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2004.11.009

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**Chapter 4**

Provisional chapter

**Mechanical Properties Comparison of Engineering**

DOI: 10.5772/intechopen.70493

Mechanical Properties Comparison of Engineering

Comparative investigations are presented into the structure and properties of selected engineering materials used for dental prosthetic restorations manufactured alternatively by the subtractive method by milling on computerised numerical control (CNC) milling machines and by the additive selective laser sintering (SLS) method of solid and porous elements using computer-aided design/computer-aided manufacturing (CAD/CAM) techniques; moreover, an original technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. Suitability was confirmed of applying the manufacturing technologies used in prosthodontics from powders by the SLS technique. The results were compared of the executed tests of strength micro-samples and of the selected prosthetic Bridges. The SLS technology for titanium, and even more for Ti6Al4V alloy, ensures the achievement of mechanical properties comparable or better than a reference Co-Cr alloy commonly used for prosthetic restorations, including prosthetic bridges fabricated by milling solid discs with CNC milling machines. For all the examined engineering materials, the milled and then sintered ZrO2 material exhibits the lowest strength properties. The results presented in

Keywords: prosthodontics, machining, additive machining, selective laser sintering,

1. Justification and scope of comparative investigations of selected alloys

Practical dental engineering encounters numerous problems with selecting the most suitable material and technology of dental manufacturing of a prosthetic restoration by a dental

> © The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and technologies of dental manufacturing of prosthetic restorations

**Materials Produced by Additive and Subtractive**

Materials Produced by Additive and Subtractive

**Technologies for Dental Prosthetic Restoration**

Technologies for Dental Prosthetic Restoration

**Application**

Application

Lech B. Dobrzański

Lech B. Dobrzański

Abstract

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

this chapter can be directly applied in dental practice.

titanium and its alloys , Co-Cr alloys, zirconium oxide

http://dx.doi.org/10.5772/intechopen.70493

Provisional chapter

**Mechanical Properties Comparison of Engineering Materials Produced by Additive and Subtractive Technologies for Dental Prosthetic Restoration Application** Mechanical Properties Comparison of Engineering Materials Produced by Additive and Subtractive Technologies for Dental Prosthetic Restoration Application

DOI: 10.5772/intechopen.70493

Lech B. Dobrzański Lech B. Dobrzański

Additional information is available at the end of the chapter Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.70493

#### Abstract

Comparative investigations are presented into the structure and properties of selected engineering materials used for dental prosthetic restorations manufactured alternatively by the subtractive method by milling on computerised numerical control (CNC) milling machines and by the additive selective laser sintering (SLS) method of solid and porous elements using computer-aided design/computer-aided manufacturing (CAD/CAM) techniques; moreover, an original technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. Suitability was confirmed of applying the manufacturing technologies used in prosthodontics from powders by the SLS technique. The results were compared of the executed tests of strength micro-samples and of the selected prosthetic Bridges. The SLS technology for titanium, and even more for Ti6Al4V alloy, ensures the achievement of mechanical properties comparable or better than a reference Co-Cr alloy commonly used for prosthetic restorations, including prosthetic bridges fabricated by milling solid discs with CNC milling machines. For all the examined engineering materials, the milled and then sintered ZrO2 material exhibits the lowest strength properties. The results presented in this chapter can be directly applied in dental practice.

Keywords: prosthodontics, machining, additive machining, selective laser sintering, titanium and its alloys , Co-Cr alloys, zirconium oxide

## 1. Justification and scope of comparative investigations of selected alloys and technologies of dental manufacturing of prosthetic restorations

Practical dental engineering encounters numerous problems with selecting the most suitable material and technology of dental manufacturing of a prosthetic restoration by a dental

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

© The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

surgeon conducting treatment. One of the important reasons is inability to compare directly the properties and structure of analogous prosthetic restorations but produced using various technologies and from different alternative engineering materials accepted for medical, especially dental, uses as compared to the restorations used most often until now. It is beyond any doubt that dental surgeons are responsible for the strategy and quality of dental treatment. The literature studies carried out as well as the Author's long-term own engineering experience in running medical practice and a dental engineering workshop show that parallel to clinical experience of a dentist conducting treatment of dental system diseases the successfulness of such activities also depends on the results of a dental engineer's activities, consisting of the appropriate execution of prosthetic restorations and implants, depending on the selection of the most appropriate engineering material and appropriate manufacturing technology, ensuring, respectively, the required biocompatibility, mechanical properties and corrosion resistance, as well as high production accuracy and aesthetic requirements. Such activities require, therefore, close cooperation of a dentist and a dental engineer having extensive and interdisciplinary engineering knowledge of the principles of construction, technology and science of materials and computer-aided methods of engineering and dental works.

materials with microporous rigid titanium and titanium alloy skeletons manufactured by selective laser sintering, whose pores are filled with living cells. The aim of such constructional and technological solutions is to ensure conditions for natural ingrowth of living tissues in the connection zone of prosthetic/implant elements, with bone or organ stumps, to eliminate the need to apply for patients the mechanical elements which are positioning and fixating the implants. In order to manufacture engineering and biological materials and implant scaffolds, it is required to seek the most advantageous proliferation conditions of living cells inside the pores of a microporous skeleton made of titanium and titanium alloys, because the adhesion and growth of living cells are dependent on the type and characteristics of the substrate. A porous zone requires surface treatment inside pores to improve proliferation conditions of living tissues, which are not advantageous for a substrate made of titanium or Ti6Al4V alloy. It appears that the improvement of proliferation conditions of cells is ensured by a substrate made of fully compatible materials, including TiO2, Al2O3 oxides and a hydroxyapatite, Ca10(PO4)6(OH)2, used as coatings of internal surfaces of pores of a microporous skeleton fabricated by SLS. Thin layers inside pores were deposited by two technologies ensuring the uniform thickness of coatings on all the walls and openings of a substrate, that is, the so-called atomic layer deposition (ALD) [19, 37–43], and the sol-gel technology of coating deposition from the liquid phase by the immersion method [37, 40, 44–52]. In order to be able to accomplish the above-mentioned general aims of the investigations, it has to be confirmed that additive technologies, especially SLS, do not lead to unacceptable deterioration of useful properties of the so-

Mechanical Properties Comparison of Engineering Materials Produced by Additive and Subtractive Technologies…

http://dx.doi.org/10.5772/intechopen.70493

113

Apart from permanent prosthetic restorations made traditionally on a metallic substructure, new technologies of introducing ceramic materials have been launched, including so-called fully porcelain systems made of zirconium oxide and aluminium oxide. For dimensional accuracy and aesthetics as well as biocompatibility and dental strength of prosthetic restorations, the constant development of manufacturing technologies of prosthetic restorations is required along with the evolvement of the engineering materials currently employed in prosthetics. Highly varied conditions of joining the artificially produced prosthetic restorations and implants with a body's natural tissues, mainly in the oral cavity, are driving especially high requirements for the engineering material applied for prosthetic restorations and implants. The reason for this is not only special conditions of using various prosthetic and implantological restorations but also different economic and technological conditions and differentiated availability and level of clinical methods and also important aesthetic reasons. Computer-aided design/computer-aided manufacturing (CAD/CAM) techniques [53–59], allowing to improve dimensional accuracy of the fabricated prosthetic restorations, have become increasingly vital over the last decade. CAD/CAM techniques find their application in dental engineering in machining, in plastic forming, in additive manufacturing technologies and in special technologies and both in metal and ceramics working. Computer-aided design enables, in particular, to produce a digital prosthetic model by scanning with an intra-cavity scanner and also to customise the designing and manufacturing of each prosthetic restoration under constant supervision of a doctor conducting the treatment. Cone beam computed tomography (CBCT), used often in dental surgeries these days as the basic diagnostic tool for comprehensive dental treatment, mainly

produced prosthetic restorations.

The dental system, especially teeth, is still subject to numerous disease processes, especially tooth decay, so widespread in many countries and practically across the globe [1–4]. This disease has a considerable social range and is a serious medical problem, not only due to local mouth cavity disorders but also due to a high risk of systemic complications stemming from the dental pulp disease, and is a direct cause of a majority of tooth losses among a wide group of people in many countries. The extraction of unhealthy teeth was a dominant practice in such cases for a long time. There may also be inborn causes of missing teeth, for example, due to missing tooth buds. The basic method of producing permanent prosthetic restorations, which is still often used, is to cast metal alloys requiring successive machining, mainly to remove elements of the gate system, and to ensure required surface roughness and dimensional accuracy. Apart from conventional technologies of manufacturing prosthetic restorations including especially powder metallurgy, casting technologies, metallic foam manufacturing, also additive technologies of fabricating solid and microporous materials in medicine and dentistry are exceptionally useful. The potential of additive technologies is much larger than of other technologies and is represented chiefly by opportunities ensured by three-dimensional designing and the related control of the structure, sizes and shape of the materials produced and the repeatability of geometric features of the elements produced, and because no wastes are produced and because the simplicity of the technology is comprised of two key stages only: designing and manufacturing.

Considering the additive technologies employed most extensively in industry, only few have found their application in prosthetics, with selective laser sintering/selective laser melting (SLS/SLM) offering the biggest opportunities [5–15]. This issue is even more important to clarify whether such additive technologies, applied in dental engineering, are competitive for conventional technologies, because original own works [5, 16–36] have been undertaken concerning constructional solutions and fabrication technologies of a new generation of custom, original, hybrid, microporous, high-strength engineering and biological surgeon conducting treatment. One of the important reasons is inability to compare directly the properties and structure of analogous prosthetic restorations but produced using various technologies and from different alternative engineering materials accepted for medical, especially dental, uses as compared to the restorations used most often until now. It is beyond any doubt that dental surgeons are responsible for the strategy and quality of dental treatment. The literature studies carried out as well as the Author's long-term own engineering experience in running medical practice and a dental engineering workshop show that parallel to clinical experience of a dentist conducting treatment of dental system diseases the successfulness of such activities also depends on the results of a dental engineer's activities, consisting of the appropriate execution of prosthetic restorations and implants, depending on the selection of the most appropriate engineering material and appropriate manufacturing technology, ensuring, respectively, the required biocompatibility, mechanical properties and corrosion resistance, as well as high production accuracy and aesthetic requirements. Such activities require, therefore, close cooperation of a dentist and a dental engineer having extensive and interdisciplinary engineering knowledge of the principles of construction, technology and science of materials and computer-aided methods of engineer-

The dental system, especially teeth, is still subject to numerous disease processes, especially tooth decay, so widespread in many countries and practically across the globe [1–4]. This disease has a considerable social range and is a serious medical problem, not only due to local mouth cavity disorders but also due to a high risk of systemic complications stemming from the dental pulp disease, and is a direct cause of a majority of tooth losses among a wide group of people in many countries. The extraction of unhealthy teeth was a dominant practice in such cases for a long time. There may also be inborn causes of missing teeth, for example, due to missing tooth buds. The basic method of producing permanent prosthetic restorations, which is still often used, is to cast metal alloys requiring successive machining, mainly to remove elements of the gate system, and to ensure required surface roughness and dimensional accuracy. Apart from conventional technologies of manufacturing prosthetic restorations including especially powder metallurgy, casting technologies, metallic foam manufacturing, also additive technologies of fabricating solid and microporous materials in medicine and dentistry are exceptionally useful. The potential of additive technologies is much larger than of other technologies and is represented chiefly by opportunities ensured by three-dimensional designing and the related control of the structure, sizes and shape of the materials produced and the repeatability of geometric features of the elements produced, and because no wastes are produced and because the simplicity of the technology is comprised of two key stages only:

Considering the additive technologies employed most extensively in industry, only few have found their application in prosthetics, with selective laser sintering/selective laser melting (SLS/SLM) offering the biggest opportunities [5–15]. This issue is even more important to clarify whether such additive technologies, applied in dental engineering, are competitive for conventional technologies, because original own works [5, 16–36] have been undertaken concerning constructional solutions and fabrication technologies of a new generation of custom, original, hybrid, microporous, high-strength engineering and biological

ing and dental works.

112 Biomaterials in Regenerative Medicine

designing and manufacturing.

materials with microporous rigid titanium and titanium alloy skeletons manufactured by selective laser sintering, whose pores are filled with living cells. The aim of such constructional and technological solutions is to ensure conditions for natural ingrowth of living tissues in the connection zone of prosthetic/implant elements, with bone or organ stumps, to eliminate the need to apply for patients the mechanical elements which are positioning and fixating the implants. In order to manufacture engineering and biological materials and implant scaffolds, it is required to seek the most advantageous proliferation conditions of living cells inside the pores of a microporous skeleton made of titanium and titanium alloys, because the adhesion and growth of living cells are dependent on the type and characteristics of the substrate. A porous zone requires surface treatment inside pores to improve proliferation conditions of living tissues, which are not advantageous for a substrate made of titanium or Ti6Al4V alloy. It appears that the improvement of proliferation conditions of cells is ensured by a substrate made of fully compatible materials, including TiO2, Al2O3 oxides and a hydroxyapatite, Ca10(PO4)6(OH)2, used as coatings of internal surfaces of pores of a microporous skeleton fabricated by SLS. Thin layers inside pores were deposited by two technologies ensuring the uniform thickness of coatings on all the walls and openings of a substrate, that is, the so-called atomic layer deposition (ALD) [19, 37–43], and the sol-gel technology of coating deposition from the liquid phase by the immersion method [37, 40, 44–52]. In order to be able to accomplish the above-mentioned general aims of the investigations, it has to be confirmed that additive technologies, especially SLS, do not lead to unacceptable deterioration of useful properties of the soproduced prosthetic restorations.

Apart from permanent prosthetic restorations made traditionally on a metallic substructure, new technologies of introducing ceramic materials have been launched, including so-called fully porcelain systems made of zirconium oxide and aluminium oxide. For dimensional accuracy and aesthetics as well as biocompatibility and dental strength of prosthetic restorations, the constant development of manufacturing technologies of prosthetic restorations is required along with the evolvement of the engineering materials currently employed in prosthetics. Highly varied conditions of joining the artificially produced prosthetic restorations and implants with a body's natural tissues, mainly in the oral cavity, are driving especially high requirements for the engineering material applied for prosthetic restorations and implants. The reason for this is not only special conditions of using various prosthetic and implantological restorations but also different economic and technological conditions and differentiated availability and level of clinical methods and also important aesthetic reasons. Computer-aided design/computer-aided manufacturing (CAD/CAM) techniques [53–59], allowing to improve dimensional accuracy of the fabricated prosthetic restorations, have become increasingly vital over the last decade. CAD/CAM techniques find their application in dental engineering in machining, in plastic forming, in additive manufacturing technologies and in special technologies and both in metal and ceramics working. Computer-aided design enables, in particular, to produce a digital prosthetic model by scanning with an intra-cavity scanner and also to customise the designing and manufacturing of each prosthetic restoration under constant supervision of a doctor conducting the treatment. Cone beam computed tomography (CBCT), used often in dental surgeries these days as the basic diagnostic tool for comprehensive dental treatment, mainly

implantological, prosthetic and endodontic treatment, creates excellent possibilities in this field. A widespread application of CBCT also permits to utilise the information acquired about the condition of a patient's tissues in planning and performing implant and prosthetic treatment [55, 60].

principally the possibility of producing solid and porous titanium parts by additive technol-

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115

This chapter depicts the outcomes of comparative investigations into the structure and properties of selected engineering materials used for dental prosthetic restorations manufactured alternatively by the subtractive method by milling on CNC milling machines and by the additive SLS method of solid and porous items using CAD/CAM techniques; moreover, an original technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. The assumptions made in this chapter are focussed on the selected engineering materials finding application in dental engineering, including mainly casting alloys of cobalt, titanium and its alloys and

Materials for milling were delivered in the form of discs with the diameter of 98.3 mm and height of 10–16 mm, and the chemical composition, according to suppliers' attestations, is

Two types of powders with a spherical shape were used, respectively, for selective laser sintering and with the composition shown in Table 2, also confirmed with spectral examina-

• Titanium powder with Grade 4 and grain size of up to 45 μm, oxygen concentration reduced to 0.14%, while average oxygen content in titanium powders is about 0.5%, the

• Ti6Al4V alloy powder with the grain diameter of 15–45 μm, for medical applications.

Co-Cr Co Cr W Mo Si Others C, Fe, Mn, N altogether 59.0 25.0 9.5 3.5 1.0 < 1.0

> 99.0 4.5-6.0

Table 1. Chemical composition of the examined materials manufactured by CAD/CAM by milling solid discs.

The results of comparative investigations are presented into the structure and properties of the mentioned engineering materials used for dental prosthetic restorations manufactured alternatively by the subtractive method by milling on CNC milling machines and by the additive SLS method of solid and porous items using CAD/CAM techniques; moreover, an original developed

ogy, for use in dentistry and reconstructive medicine.

tions with the energy dispersive spectrometry (EDS) method:

aim of which is to ensure process safety,

Material Mass concentration/fraction, %

Ti grade 2 Ti Others N, C, H, Fe, O altogether

Ti6Al4V Ti Al V Others N, C, H, Fe, O altogether

ZrO2 ZrO2 + HfO2 + Y2O3 Including Y2O3

89.4 6.2 4.0 ≤ 1.0

> 99.6 < 0.4

ceramics on the zirconium oxide matrix.

shown in Table 1.

The assumptions made require special interest in the selected engineering materials finding application in dental engineering, including mainly casting alloys of cobalt, titanium and its alloys and ceramics on the zirconium oxide matrix [61–72]. The essential aspects underlying the investigations described here were the questions whether—in each case—it is feasible to apply prosthetic restorations made on a metallic substructure of zirconium oxide, guaranteeing currently best aesthetic results, for prosthetic restorations and to explain if it is possible to use a zirconium oxide substructure in the side sections and in circular restorations together with recognising the limitations for the use of this material. Another practical aspect is to develop a methodology for application of titanium and its alloys as compared to Co-Cr alloys of the Vitalium type for implantoprosthetic restorations so as to guarantee the lowest possible mass of the completely manufactured prosthetic restoration while ensuring aesthetic requirements and the maximum, possible strength properties. For production of metallic substructures, it is necessary to ensure space for veneering porcelain, among others, in intermolecular spaces on abutments between crowns and bridge spans but also on the vestibular space, so that the dark colour of the substructure does not influence the colour of the ready veneering ceramics. Considering that disadvantages and limitations are known of casting and machining technologies such as computerised numerical control (CNC) milling currently used in dental engineering laboratories, it was decided to compare the state-of-the-art manufacturing techniques currently applied in dental engineering, including completely innovative additive techniques which—due to high implementation price, but very low price of producing a single prosthetic point—are rarely used in practice. As Co-Cr alloys are most often used in dental practice and, therefore, most often described in the literature, such were used as a reference point for the works performed already described in literature. This will allow practitioners, especially dentists, to easily transfer the results presented in this article to everyday uses and to consider such results for designing the extensiveness of prosthetic restorations and for cost valuation. The matter is very complex and requires multi-criteria optimisation, which cannot be performed individually by any prosthetic laboratory. The matter requires, therefore, methodologically planned research which can demonstrate—to the practising therapists and prosthetic workshops cooperating with them—the advantages of some material and technological solutions over those commonly in use these days. This will surely have a major impact on the outcomes and durability of dental treatment and on satisfaction, comfort and patients' improved health condition. The aspects highlighted were the reasons for outlining the topic of the chapter presenting the results of the work [18], after making a thorough literature analysis of the status of research into the dental system, its diseases and treatment methods of tooth losses as well as materials and technologies applied in prosthetics. A concept presented in this article is covered by a broader range of works followed in the recent years [5, 7, 16, 17, 21–26, 34, 55, 56, 73–79], in which the author has participated actively. The works concern principally the possibility of producing solid and porous titanium parts by additive technology, for use in dentistry and reconstructive medicine.

implantological, prosthetic and endodontic treatment, creates excellent possibilities in this field. A widespread application of CBCT also permits to utilise the information acquired about the condition of a patient's tissues in planning and performing implant and prosthetic

The assumptions made require special interest in the selected engineering materials finding application in dental engineering, including mainly casting alloys of cobalt, titanium and its alloys and ceramics on the zirconium oxide matrix [61–72]. The essential aspects underlying the investigations described here were the questions whether—in each case—it is feasible to apply prosthetic restorations made on a metallic substructure of zirconium oxide, guaranteeing currently best aesthetic results, for prosthetic restorations and to explain if it is possible to use a zirconium oxide substructure in the side sections and in circular restorations together with recognising the limitations for the use of this material. Another practical aspect is to develop a methodology for application of titanium and its alloys as compared to Co-Cr alloys of the Vitalium type for implantoprosthetic restorations so as to guarantee the lowest possible mass of the completely manufactured prosthetic restoration while ensuring aesthetic requirements and the maximum, possible strength properties. For production of metallic substructures, it is necessary to ensure space for veneering porcelain, among others, in intermolecular spaces on abutments between crowns and bridge spans but also on the vestibular space, so that the dark colour of the substructure does not influence the colour of the ready veneering ceramics. Considering that disadvantages and limitations are known of casting and machining technologies such as computerised numerical control (CNC) milling currently used in dental engineering laboratories, it was decided to compare the state-of-the-art manufacturing techniques currently applied in dental engineering, including completely innovative additive techniques which—due to high implementation price, but very low price of producing a single prosthetic point—are rarely used in practice. As Co-Cr alloys are most often used in dental practice and, therefore, most often described in the literature, such were used as a reference point for the works performed already described in literature. This will allow practitioners, especially dentists, to easily transfer the results presented in this article to everyday uses and to consider such results for designing the extensiveness of prosthetic restorations and for cost valuation. The matter is very complex and requires multi-criteria optimisation, which cannot be performed individually by any prosthetic laboratory. The matter requires, therefore, methodologically planned research which can demonstrate—to the practising therapists and prosthetic workshops cooperating with them—the advantages of some material and technological solutions over those commonly in use these days. This will surely have a major impact on the outcomes and durability of dental treatment and on satisfaction, comfort and patients' improved health condition. The aspects highlighted were the reasons for outlining the topic of the chapter presenting the results of the work [18], after making a thorough literature analysis of the status of research into the dental system, its diseases and treatment methods of tooth losses as well as materials and technologies applied in prosthetics. A concept presented in this article is covered by a broader range of works followed in the recent years [5, 7, 16, 17, 21–26, 34, 55, 56, 73–79], in which the author has participated actively. The works concern

treatment [55, 60].

114 Biomaterials in Regenerative Medicine

This chapter depicts the outcomes of comparative investigations into the structure and properties of selected engineering materials used for dental prosthetic restorations manufactured alternatively by the subtractive method by milling on CNC milling machines and by the additive SLS method of solid and porous items using CAD/CAM techniques; moreover, an original technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. The assumptions made in this chapter are focussed on the selected engineering materials finding application in dental engineering, including mainly casting alloys of cobalt, titanium and its alloys and ceramics on the zirconium oxide matrix.

Materials for milling were delivered in the form of discs with the diameter of 98.3 mm and height of 10–16 mm, and the chemical composition, according to suppliers' attestations, is shown in Table 1.

Two types of powders with a spherical shape were used, respectively, for selective laser sintering and with the composition shown in Table 2, also confirmed with spectral examinations with the energy dispersive spectrometry (EDS) method:


The results of comparative investigations are presented into the structure and properties of the mentioned engineering materials used for dental prosthetic restorations manufactured alternatively by the subtractive method by milling on CNC milling machines and by the additive SLS method of solid and porous items using CAD/CAM techniques; moreover, an original developed


Table 1. Chemical composition of the examined materials manufactured by CAD/CAM by milling solid discs.


the CAD method, and relevant data in the stereolithography (STL) format is transferred for the

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The solid elements used for the investigations, fabricated by the machining method by milling solid discs, were prepared with special CAD software, dental wings operation system (DWOS), by Dental Wings and 3D Marcam Engineering AutoFab software (software for manufacturing

The software features the width offset of the tools used for milling, to ensure the highest possible design reproduction in reality. The so-designed elements were verified by the software for surface cohesion and developed as ready designs in the stereolithography (STL) format. The design files are then implemented to a CAM module, Mayka Dental, by Picasoft. The CAM module enables to position the design prepared on a material disc according to the optimum three-axial milling axis and ensures the placement of the milled elements in a way least interfering with the design and to determine the optimum axis of precision five-axial milling to exploit optimally the capabilities of the milling machine. The so-prepared design is transformed into a machine language, and the file produced is transferred to the milling

Another group of the examined elements was manufactured by an additive technology of selective laser sintering by designing a given element virtually, in order to develop a 3D CAD model in the STL format, transferred to the machine's software and then by producing in reality the so-designed element layer by layer until producing a ready product. 3D Marcam Engineering AutoFab (software for manufacturing applications) software was used for model designing as a CAD/CAM tool. The software enables to select model dimensions, constructional features and the type of the model volume filling either as solid or porous and to choose the size of a cell unit making up the entire model. A 3D model featuring a layered structure is produced as a result of designing. The surface of the elements modelled in the STL format, which is accommodated to SLS manufacturing conditions, is shown with a net of triangles, with the accuracy of the surface being represented being inversely proportional to their size. The model size and structure have to be assumed each time and divided into layers with the expected thickness. The number of layers in a virtual model matches the actually designed number of powder layers when the given element is really produced. A set of "hexagon cross" unit cells was used, selected in a geometrical analysis and as a result of preliminary studies, from the set available in the software. By duplicating them, the entire elements were designed, consisting of nodes and single-lattice fibres, linking the particular skeleton nodes (Figure 1). Fabrication conditions were also selected, at the stage of virtual design of elements, available in the selective laser sintering system, including layer thickness, laser power, laser beam diameter, scanning rate and distance between particular remelting paths. A spatial orientation of unit cells of 45 relative to the axis x of the system of coordinates was also selected experimentally, because other orientations analysed initially, including at the angle of 0, at the angle of 45 relative to the axis y, at the angle of 45 relative to the axis x and 45 to the axis y, at the angle of 45 relative to the axis y and 45 relative to the axis x, turned out to be less advantageous due to

execution of computer-aided manufacturing in a relevant technological device.

applications) for geometrically simpler shapes of test specimens.

the mechanical properties of the elements produced (Figure 1).

machine's software.

Table 2. Chemical composition of the powders used for selective laser sintering.

technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. Experimental comparative investigations were carried out of selected engineering materials recommended for clinical use for permanent dental prostheses manufactured by CAD/CAM. The properties and structure were also compared directly of analogous prosthetic restorations but produced using the various technologies mentioned and from different alternative engineering materials already mentioned, as well, as compared to the restorations made of Co-Cr alloy by milling solid discs, used most often until now. In order to create a ranking of the materials and technologies applied, analogous elements were produced by milling from solid discs made of ZrO2 ceramics and then by sintering. Such material and technology are commonly regarded by dentist practitioners as the most avant-garde solution. Elements were produced from titanium and Ti6Al4V alloy by machining by way of milling solid discs, and analogous elements were made from such materials by the additive method in an SLS process. An original technology was established of manufacturing elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique.

## 2. Additive and machining technology of fabrication of materials and dental restorations and research methodology

The elements used for the investigations were fabricated by the machining method by milling solid discs on a CNC milling machine and by the additive SLS method. In the case of the engineering materials manufactured by the machining method by milling on CNC milling machines, technological conditions of manufacturing were chosen experimentally. The technological processes mentioned have no influence on any structural changes, except that ceramic ZrO2 material of the Cercon type has to be sintered already after milling, so the selection of manufacturing conditions does not have to be verified experimentally and structural changes do not have to be confronted with the mechanical properties obtained of the elements produced. The situation is opposite if the SLS additive technology is applied. Investigations over the selection of technological conditions of selective laser sintering involve the selection of laser power, laser beam diameter, a dependency between the laser beam diameter and the distance between laser beams and the distance between laser remelting paths to desired geometrical features of solid and porous elements produced from titanium and Ti6Al4V alloy and the structure formed depending on such technological conditions and its influence on mechanical properties. Each time before producing the examined elements, design was performed using the CAD method, and relevant data in the stereolithography (STL) format is transferred for the execution of computer-aided manufacturing in a relevant technological device.

The solid elements used for the investigations, fabricated by the machining method by milling solid discs, were prepared with special CAD software, dental wings operation system (DWOS), by Dental Wings and 3D Marcam Engineering AutoFab software (software for manufacturing applications) for geometrically simpler shapes of test specimens.

The software features the width offset of the tools used for milling, to ensure the highest possible design reproduction in reality. The so-designed elements were verified by the software for surface cohesion and developed as ready designs in the stereolithography (STL) format. The design files are then implemented to a CAM module, Mayka Dental, by Picasoft. The CAM module enables to position the design prepared on a material disc according to the optimum three-axial milling axis and ensures the placement of the milled elements in a way least interfering with the design and to determine the optimum axis of precision five-axial milling to exploit optimally the capabilities of the milling machine. The so-prepared design is transformed into a machine language, and the file produced is transferred to the milling machine's software.

technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. Experimental comparative investigations were carried out of selected engineering materials recommended for clinical use for permanent dental prostheses manufactured by CAD/CAM. The properties and structure were also compared directly of analogous prosthetic restorations but produced using the various technologies mentioned and from different alternative engineering materials already mentioned, as well, as compared to the restorations made of Co-Cr alloy by milling solid discs, used most often until now. In order to create a ranking of the materials and technologies applied, analogous elements were produced by milling from solid discs made of ZrO2 ceramics and then by sintering. Such material and technology are commonly regarded by dentist practitioners as the most avant-garde solution. Elements were produced from titanium and Ti6Al4V alloy by machining by way of milling solid discs, and analogous elements were made from such materials by the additive method in an SLS process. An original technology was established of manufacturing elements used in prosthodon-

Ti – – 0.01 0.03 0.14 0.01 0.004 < 0.4 < 0.01 remainder

Ti6Al4V 6.35 4.0 0.01 0.2 0.15 0.02 0.003 ≤ 0.4 ≤ 0.1

Al V C Fe O N H Others total Others each Ti

tics, produced with titanium and Ti6Al4V alloy powders with the SLS technique.

Powder Mass concentration of elements, %

116 Biomaterials in Regenerative Medicine

Table 2. Chemical composition of the powders used for selective laser sintering.

dental restorations and research methodology

2. Additive and machining technology of fabrication of materials and

The elements used for the investigations were fabricated by the machining method by milling solid discs on a CNC milling machine and by the additive SLS method. In the case of the engineering materials manufactured by the machining method by milling on CNC milling machines, technological conditions of manufacturing were chosen experimentally. The technological processes mentioned have no influence on any structural changes, except that ceramic ZrO2 material of the Cercon type has to be sintered already after milling, so the selection of manufacturing conditions does not have to be verified experimentally and structural changes do not have to be confronted with the mechanical properties obtained of the elements produced. The situation is opposite if the SLS additive technology is applied. Investigations over the selection of technological conditions of selective laser sintering involve the selection of laser power, laser beam diameter, a dependency between the laser beam diameter and the distance between laser beams and the distance between laser remelting paths to desired geometrical features of solid and porous elements produced from titanium and Ti6Al4V alloy and the structure formed depending on such technological conditions and its influence on mechanical properties. Each time before producing the examined elements, design was performed using Another group of the examined elements was manufactured by an additive technology of selective laser sintering by designing a given element virtually, in order to develop a 3D CAD model in the STL format, transferred to the machine's software and then by producing in reality the so-designed element layer by layer until producing a ready product. 3D Marcam Engineering AutoFab (software for manufacturing applications) software was used for model designing as a CAD/CAM tool. The software enables to select model dimensions, constructional features and the type of the model volume filling either as solid or porous and to choose the size of a cell unit making up the entire model. A 3D model featuring a layered structure is produced as a result of designing. The surface of the elements modelled in the STL format, which is accommodated to SLS manufacturing conditions, is shown with a net of triangles, with the accuracy of the surface being represented being inversely proportional to their size.

The model size and structure have to be assumed each time and divided into layers with the expected thickness. The number of layers in a virtual model matches the actually designed number of powder layers when the given element is really produced. A set of "hexagon cross" unit cells was used, selected in a geometrical analysis and as a result of preliminary studies, from the set available in the software. By duplicating them, the entire elements were designed, consisting of nodes and single-lattice fibres, linking the particular skeleton nodes (Figure 1).

Fabrication conditions were also selected, at the stage of virtual design of elements, available in the selective laser sintering system, including layer thickness, laser power, laser beam diameter, scanning rate and distance between particular remelting paths. A spatial orientation of unit cells of 45 relative to the axis x of the system of coordinates was also selected experimentally, because other orientations analysed initially, including at the angle of 0, at the angle of 45 relative to the axis y, at the angle of 45 relative to the axis x and 45 to the axis y, at the angle of 45 relative to the axis y and 45 relative to the axis x, turned out to be less advantageous due to the mechanical properties of the elements produced (Figure 1).

machine and by the additive SLS method. The bending strength was compared to dental bridges made of solid materials, that is, sintered titanium, sintered Ti6Al4V alloy, milled Ti6Al4V alloy, milled Co-Cr alloy and milled and sintered ZrO2 sinter. The results of strength tests of dental bridges selectively sintered by laser in optimised conditions were compared with the results of tests for bridges milled with CAD/CAM methods from discs of other tested materials, additionally sintered after milling in the case of ZrO2. The samples prepared correspond to the prosthetic restorations designed in the author's own clinical practice and in a dental engineering laboratory for particular patients and used by them. Geometrical features were taken into account, such as the fitting of a prosthetic restoration to a pillar, of 0.03 mm; the width of the threshold area; the space for prosthetic cement in the occlusal region, of 0.1 mm; the reduction of the crown shape considering the space for porcelain, of 1.5 mm; and the minimum thickness of the substructure according to the material manufacturer's recommendations, of 0.3 mm, for a substructure made of Co-Cr alloys, titanium and its alloys and 0.6 mm for a zirconium oxide substructure. The software features the width offset of the tools

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The so-prepared micro-samples and bridges were verified by software for surface cohesion and prepared as a ready design in the STL format (stereolithography), processed into a machine language of a CNC milling machine by Yenadent D43W or AM 125 machine by Renishaw for selective powder laser sintering. After designing the shape of the elements produced, short machining of such elements takes place with a CNC milling machine from a disc of the selected material and additional sintering in a resistance furnace for ceramic prostheses (Table 3). In the case of additive treatment, after designing the produced element the same as previously, the particular layers of metals and alloys are joined layer by layer in the

As reference results, the results were used of tests of micro-samples and prosthetic bridges fabricated by the machining method with CAD/CAM techniques by milling—with CNC milling machines—solid discs made of Co-Cr alloy of the Vitalium type, used most often in dental treatment, mainly because the alloys are commonly used in casting methods. In order to create a ranking of the materials and technologies used, analogous elements were produced by milling from solid discs made of ZrO2 ceramics and then by sintering. Titanium elements were fabricated by the machining method by milling solid discs and by the additive SLS method. Elements from Ti6Al4V powder by SLS (Table 3) were also produced for comparative purposes. Apart from examinations of solid materials, examinations of porous elements with the pore size of 200–250 μm manufactured by SLS from titanium and Ti6Al4V alloy powder were

Milling from disc on CNC milling machine Manufacturing technology Co-Cr ZrO2 Ti6Al4V Solid engineering material

Table 3. Diagram of selection of engineering materials and manufacturing technologies of the examined elements.

Sintered solid engineering material Ti6Al4V Ti

Manufacturing technology Selective laser sintering of powders

used for milling.

form of powder in an SLS process (Table 3).

also carried out for comparative purposes.

Figure 1. (a, b) "Hexagon cross unit cell"; (c) the arrangement of unit cells in the space of the system of coordinates at the angle of 45 relative to the axis x.

Figure 2. Computer model image of solid samples for static test: (a) tensile test, (b) three-point bending test and (c) compression test.

Micro-samples for static tensile, bending and compression tests were specially selected for the investigations (Figure 2), made of all the chosen materials manufactured using, respectively, the additive SLS method from powders with individual CAD/CAM and with the machining technology using CAD/CAM techniques by milling solid discs on CNC milling machines. Micro-samples for tensile strength, bending strength and compression strength tests were also designed for the conditions given above with the pore size of 200–250 μm.

In order to perform bending strength tests, models were created of three-point prosthetic bridges from three different sections of a dental arch (Figure 3), including the front, frontwing and wing section, taking into account first of all the highest possible stresses which may occur in such type of prosthetic restorations. They were made as solid from all the examined engineering materials for bending strength tests with the three-point bending method with the both analysed technologies, that is, machining method by milling solid discs on a CNC milling

Figure 3. Computer model of bridges nos. 1, 2 and 3 for bending strength tests with the three-point bending method.

machine and by the additive SLS method. The bending strength was compared to dental bridges made of solid materials, that is, sintered titanium, sintered Ti6Al4V alloy, milled Ti6Al4V alloy, milled Co-Cr alloy and milled and sintered ZrO2 sinter. The results of strength tests of dental bridges selectively sintered by laser in optimised conditions were compared with the results of tests for bridges milled with CAD/CAM methods from discs of other tested materials, additionally sintered after milling in the case of ZrO2. The samples prepared correspond to the prosthetic restorations designed in the author's own clinical practice and in a dental engineering laboratory for particular patients and used by them. Geometrical features were taken into account, such as the fitting of a prosthetic restoration to a pillar, of 0.03 mm; the width of the threshold area; the space for prosthetic cement in the occlusal region, of 0.1 mm; the reduction of the crown shape considering the space for porcelain, of 1.5 mm; and the minimum thickness of the substructure according to the material manufacturer's recommendations, of 0.3 mm, for a substructure made of Co-Cr alloys, titanium and its alloys and 0.6 mm for a zirconium oxide substructure. The software features the width offset of the tools used for milling.

The so-prepared micro-samples and bridges were verified by software for surface cohesion and prepared as a ready design in the STL format (stereolithography), processed into a machine language of a CNC milling machine by Yenadent D43W or AM 125 machine by Renishaw for selective powder laser sintering. After designing the shape of the elements produced, short machining of such elements takes place with a CNC milling machine from a disc of the selected material and additional sintering in a resistance furnace for ceramic prostheses (Table 3). In the case of additive treatment, after designing the produced element the same as previously, the particular layers of metals and alloys are joined layer by layer in the form of powder in an SLS process (Table 3).

Micro-samples for static tensile, bending and compression tests were specially selected for the investigations (Figure 2), made of all the chosen materials manufactured using, respectively, the additive SLS method from powders with individual CAD/CAM and with the machining technology using CAD/CAM techniques by milling solid discs on CNC milling machines. Micro-samples for tensile strength, bending strength and compression strength tests were also

Figure 2. Computer model image of solid samples for static test: (a) tensile test, (b) three-point bending test and

Figure 1. (a, b) "Hexagon cross unit cell"; (c) the arrangement of unit cells in the space of the system of coordinates at the

In order to perform bending strength tests, models were created of three-point prosthetic bridges from three different sections of a dental arch (Figure 3), including the front, frontwing and wing section, taking into account first of all the highest possible stresses which may occur in such type of prosthetic restorations. They were made as solid from all the examined engineering materials for bending strength tests with the three-point bending method with the both analysed technologies, that is, machining method by milling solid discs on a CNC milling

Figure 3. Computer model of bridges nos. 1, 2 and 3 for bending strength tests with the three-point bending method.

designed for the conditions given above with the pore size of 200–250 μm.

angle of 45 relative to the axis x.

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(c) compression test.

As reference results, the results were used of tests of micro-samples and prosthetic bridges fabricated by the machining method with CAD/CAM techniques by milling—with CNC milling machines—solid discs made of Co-Cr alloy of the Vitalium type, used most often in dental treatment, mainly because the alloys are commonly used in casting methods. In order to create a ranking of the materials and technologies used, analogous elements were produced by milling from solid discs made of ZrO2 ceramics and then by sintering. Titanium elements were fabricated by the machining method by milling solid discs and by the additive SLS method. Elements from Ti6Al4V powder by SLS (Table 3) were also produced for comparative purposes. Apart from examinations of solid materials, examinations of porous elements with the pore size of 200–250 μm manufactured by SLS from titanium and Ti6Al4V alloy powder were also carried out for comparative purposes.


Table 3. Diagram of selection of engineering materials and manufacturing technologies of the examined elements.

## 3. Comparison of strength properties of the examined solid engineering materials for dental prosthetic restorations manufactured by additive and machining methods

The investigations of mechanical properties of the examined engineering materials were performed in two series in each case. According to the assumed aim of the investigations, mechanical properties were compared in the first series, starting with tensile strength, each time mechanical properties of solid materials produced by the SLS technique for different laser power, that is, sintered titanium and sintered Ti6Al4V alloy. Mechanical properties were compared in the second series of all the examined solid materials, that is, sintered titanium, sintered Ti6Al4V alloy, milled Ti6Al4V alloy, milled Co-Cr alloy and milled and milled ZrO2 sinter. The results of the examinations were presented in such order for, respectively, tensile strength, bending strength and comprehensive strength.

The impact of laser power, ranging 70–110 W, was investigated in the first place on tensile strength values of sintered titanium and sintered Ti6Al4V alloy (Figure 4). The results of tensile strength tests for sintered titanium, sintered Ti6Al4V alloy, obtained in optimum conditions of production, milled Ti6Al4V alloy, milled Co-Cr alloy and milled ZrO2 sinter, are shown in Figure 5.

> The test results given in Figure 5 are an interesting reason for application of the tested materials in dental prosthetics. A result of approx. 47 MPa for tensile strength of milled and then sintered ZrO2 alloys is certainly concerning. Selectively laser sintered titanium exhibits the tensile strength of approx. 740 MPa, slightly smaller than specific for milled Co-Cr alloy samples, approx. 825 MPa, and Ti6Al4V alloy, approx. 858 MPa, and much lower than the tensile strength of approx. 1312 MPa of the selectively laser sintered Ti6Al4V alloy. It is the highest

> Figure 5. Comparison of diagrams of dependency between tensile stress and elongation for various examined materials,

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that is, for sintered Ti, sintered Ti6Al4V alloy, milled Co-Cr alloy, milled ZrO2 sinter and milled Ti6Al4V alloy.

Bending strength tests were performed the same as tensile strength tests. The investigations were performed in relation to solid materials manufactured by SLS for different laser power, that is, sintered titanium and sintered Ti6Al4V alloy (Figure 6), as well as in relation to the investigated solid materials, that is, sintered titanium, sintered Ti6Al4V alloy, milled Ti6Al4V alloy, milled Co-Cr alloy and milled ZrO2 sinter (Figure 7). It is noted that the bending strength of the sintered titanium of approx. 1682 MPA is only about 3% smaller than the bending strength of about 1740 MPa of the Co-Cr alloy but strongly smaller than the bending strength of approx. 1959 MPa of the milled Ti6Al4V alloy and much higher than the bending strength of approx. 2374 MPa of this alloy after SLS. The smallest is the bending strength of approx.

Figure 8 is the comparison of diagrams of dependency between compressive stress and

tensile strength value for all the examined engineering materials.

605 MPa of sintered ZrO2 alloys milled from discs and then sintered.

deformation for selectively laser sintered Ti samples and Ti6Al4V alloy.

Figure 4. Comparison of diagrams of dependency between tensile stress and elongation for solid samples made of Ti6Al4V alloy and pristine titanium sintered at different laser powers.

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3. Comparison of strength properties of the examined solid engineering materials for dental prosthetic restorations manufactured by additive

The investigations of mechanical properties of the examined engineering materials were performed in two series in each case. According to the assumed aim of the investigations, mechanical properties were compared in the first series, starting with tensile strength, each time mechanical properties of solid materials produced by the SLS technique for different laser power, that is, sintered titanium and sintered Ti6Al4V alloy. Mechanical properties were compared in the second series of all the examined solid materials, that is, sintered titanium, sintered Ti6Al4V alloy, milled Ti6Al4V alloy, milled Co-Cr alloy and milled and milled ZrO2 sinter. The results of the examinations were presented in such order for, respectively, tensile

The impact of laser power, ranging 70–110 W, was investigated in the first place on tensile strength values of sintered titanium and sintered Ti6Al4V alloy (Figure 4). The results of tensile strength tests for sintered titanium, sintered Ti6Al4V alloy, obtained in optimum conditions of production, milled Ti6Al4V alloy, milled Co-Cr alloy and milled ZrO2 sinter, are shown in

Figure 4. Comparison of diagrams of dependency between tensile stress and elongation for solid samples made of

and machining methods

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Figure 5.

strength, bending strength and comprehensive strength.

Ti6Al4V alloy and pristine titanium sintered at different laser powers.

Figure 5. Comparison of diagrams of dependency between tensile stress and elongation for various examined materials, that is, for sintered Ti, sintered Ti6Al4V alloy, milled Co-Cr alloy, milled ZrO2 sinter and milled Ti6Al4V alloy.

The test results given in Figure 5 are an interesting reason for application of the tested materials in dental prosthetics. A result of approx. 47 MPa for tensile strength of milled and then sintered ZrO2 alloys is certainly concerning. Selectively laser sintered titanium exhibits the tensile strength of approx. 740 MPa, slightly smaller than specific for milled Co-Cr alloy samples, approx. 825 MPa, and Ti6Al4V alloy, approx. 858 MPa, and much lower than the tensile strength of approx. 1312 MPa of the selectively laser sintered Ti6Al4V alloy. It is the highest tensile strength value for all the examined engineering materials.

Bending strength tests were performed the same as tensile strength tests. The investigations were performed in relation to solid materials manufactured by SLS for different laser power, that is, sintered titanium and sintered Ti6Al4V alloy (Figure 6), as well as in relation to the investigated solid materials, that is, sintered titanium, sintered Ti6Al4V alloy, milled Ti6Al4V alloy, milled Co-Cr alloy and milled ZrO2 sinter (Figure 7). It is noted that the bending strength of the sintered titanium of approx. 1682 MPA is only about 3% smaller than the bending strength of about 1740 MPa of the Co-Cr alloy but strongly smaller than the bending strength of approx. 1959 MPa of the milled Ti6Al4V alloy and much higher than the bending strength of approx. 2374 MPa of this alloy after SLS. The smallest is the bending strength of approx. 605 MPa of sintered ZrO2 alloys milled from discs and then sintered.

Figure 8 is the comparison of diagrams of dependency between compressive stress and deformation for selectively laser sintered Ti samples and Ti6Al4V alloy.

Figure 6. Comparison of diagrams of dependency between bending stress and bending for solid samples made of Ti6Al4V alloy and pristine titanium sintered at different laser powers.

## 4. Comparison of results of strength tests of dental bridges manufactured from the investigated materials with the additive and machining technology

tests of dental bridges selectively sintered by laser in optimised conditions were compared with the results of tests for bridges milled with CAD/CAM methods from discs of other tested materials, additionally sintered after milling in the case of ZrO2. The impact of laser power, ranging 70–110 W, was investigated in the first place on bending strength values of three bridges manufactured by SLS from sintered titanium and sintered Ti6Al4V alloy. The results of bending strength tests of a dental bridge no. 1 produced from titanium and Ti6AlV4 alloy were compared directly in Figure 9. Regardless the material produced by selective laser

Figure 7. Comparison of diagrams of dependency between tensile stress and bending for various examined materials,

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that is, for sintered Ti, sintered Ti6Al4V alloy, milled Co-Cr alloy, milled ZrO2 sinter and milled Ti6Al4V alloy.

The results of comparative investigations are presented of mechanical properties of dental bridges manufactured from the investigated engineering materials. The results of strength

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Figure 7. Comparison of diagrams of dependency between tensile stress and bending for various examined materials, that is, for sintered Ti, sintered Ti6Al4V alloy, milled Co-Cr alloy, milled ZrO2 sinter and milled Ti6Al4V alloy.

tests of dental bridges selectively sintered by laser in optimised conditions were compared with the results of tests for bridges milled with CAD/CAM methods from discs of other tested materials, additionally sintered after milling in the case of ZrO2. The impact of laser power, ranging 70–110 W, was investigated in the first place on bending strength values of three bridges manufactured by SLS from sintered titanium and sintered Ti6Al4V alloy. The results of bending strength tests of a dental bridge no. 1 produced from titanium and Ti6AlV4 alloy were compared directly in Figure 9. Regardless the material produced by selective laser

4. Comparison of results of strength tests of dental bridges manufactured

Figure 6. Comparison of diagrams of dependency between bending stress and bending for solid samples made of

The results of comparative investigations are presented of mechanical properties of dental bridges manufactured from the investigated engineering materials. The results of strength

from the investigated materials with the additive and machining

Ti6Al4V alloy and pristine titanium sintered at different laser powers.

technology

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The analogous strength properties were obtained for a dental bridge no. 2 produced from titanium and Ti6AlV4 (Figure 10). In the case of a titanium bridge, if laser power is increased from 70 to 110 W, this has a considerable effect on the value of the maximum bending strength, which increases on average from approx. 350 to approx. 1315 N. In the case of a bridge made of Ti6Al4V alloy produced by SLS, if laser power is raised in the same range, the maximum bending strength rises from approx. 525 to approx. 1560 N. In the case of Ti6Al4V alloy, each time, for each laser power, the maximum bending strength causing sample fracture is higher than for bridges produced from titanium, and this takes place in each case for relatively smaller bending. For example, selectively laser sintered solid bridges with laser power of 110 W from Ti6Al4V alloy are damaged at approx. 3% bending, when bending for titanium

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As laser power is increased from 70 to 110 W, the value of the maximum bending strength for dental bridge no. 3, produced from titanium, increases from approx. 240 to approx. 580 N, in

Figure 10. Comparison of diagrams of dependency between bending strength and bending of solid bridge no. 2

manufactured by SLS from titanium and Ti6Al4V alloy for different laser power.

exceeds 4.5% (Figure 10).

Figure 8. Comparison of diagrams of dependency between compressive stress and deformation for selectively laser sintered Ti and Ti6Al4V alloy samples.

Figure 9. Comparison of diagrams of dependency between bending strength and bending of solid bridge no. 1 manufactured by SLS from titanium and Ti6Al4V alloy for different laser power.

sintering, the maximum bending strength strongly depends on laser power. Sintering with the laser power of 70 W ensures small maximum bending force of approx. 260 and approx. 310 N, respectively, for titanium and Ti6Al4V alloy and is rising almost proportionally to, respectively, approx. 700 and approx. 730 N with laser power increase to 110 W. The maximum bending of the produced dental bridges depends very strongly on the applied laser power and a selectively sintered material. For titanium, this bending—depending on the laser power—varies between approx. 0.9 and approx. 2.4%, when it is much higher for Ti6Al4V alloy and ranges within approx. 2 to approx. 7%, and bending strength is systematically falling as bending is rising (Figure 9).

The analogous strength properties were obtained for a dental bridge no. 2 produced from titanium and Ti6AlV4 (Figure 10). In the case of a titanium bridge, if laser power is increased from 70 to 110 W, this has a considerable effect on the value of the maximum bending strength, which increases on average from approx. 350 to approx. 1315 N. In the case of a bridge made of Ti6Al4V alloy produced by SLS, if laser power is raised in the same range, the maximum bending strength rises from approx. 525 to approx. 1560 N. In the case of Ti6Al4V alloy, each time, for each laser power, the maximum bending strength causing sample fracture is higher than for bridges produced from titanium, and this takes place in each case for relatively smaller bending. For example, selectively laser sintered solid bridges with laser power of 110 W from Ti6Al4V alloy are damaged at approx. 3% bending, when bending for titanium exceeds 4.5% (Figure 10).

As laser power is increased from 70 to 110 W, the value of the maximum bending strength for dental bridge no. 3, produced from titanium, increases from approx. 240 to approx. 580 N, in

sintering, the maximum bending strength strongly depends on laser power. Sintering with the laser power of 70 W ensures small maximum bending force of approx. 260 and approx. 310 N, respectively, for titanium and Ti6Al4V alloy and is rising almost proportionally to, respectively, approx. 700 and approx. 730 N with laser power increase to 110 W. The maximum bending of the produced dental bridges depends very strongly on the applied laser power and a selectively sintered material. For titanium, this bending—depending on the laser power—varies between approx. 0.9 and approx. 2.4%, when it is much higher for Ti6Al4V alloy and ranges within approx. 2 to approx. 7%, and bending strength is systematically falling as bending is

Figure 9. Comparison of diagrams of dependency between bending strength and bending of solid bridge no. 1

manufactured by SLS from titanium and Ti6Al4V alloy for different laser power.

Figure 8. Comparison of diagrams of dependency between compressive stress and deformation for selectively laser

rising (Figure 9).

sintered Ti and Ti6Al4V alloy samples.

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Figure 10. Comparison of diagrams of dependency between bending strength and bending of solid bridge no. 2 manufactured by SLS from titanium and Ti6Al4V alloy for different laser power.

Figure 11. Comparison of diagrams of dependency between bending strength and bending of solid bridge no. 3 manufactured by SLS from titanium and Ti6Al4V alloy for different laser power.

the case of a bridge made of Ti6Al4V alloy, from approx. 340 to approx. 720 N (Figure 11). In the case of selectively laser sintered Ti6Al4V alloy, each time, for each laser power, bending strength causing sample fracture is higher than for bridges produced from titanium, and this takes place in each case for relatively smaller bending (Figure 11).

Diagrams were compared of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured from all the examined materials, that is, from selectively laser sintered titanium and Ti6Al4V alloy and milled from Ti6Al4V alloy discs, Co-Cr alloy and milled ZrO2 sinter and then sintered (Figure 12). The maximum bending strength highly relies on the dental bridge type and varies from approx. 580 N for bridge no. 3 to approx. 1320 N for bridge no. 2. On the other hand, bridge no. 3 exhibits the maximum bending of approx. 10.4% while bridge no. 2 smallest bending of approx. 4.6%. Bridge no. 1 shows indirect properties. In the case of solid bridges nos 1, 2 and 3 manufactured by SLS from Ti6Al4V alloy with laser power of 110 W (Figure 12), the maximum bending strength of approx. 1570 N occurs for bridge no. 2, and the two other bridges show the strength of approx. 720 N. Maximum bending for bridges nos 3, 2 and 1 is, respectively, approx. 3.2, approx. 3 and approx. 2.4%. In the case of bridges nos. 2, 1 and 3 manufactured by CAD/ CAM by milling discs from Ti6Al4V alloy, as presented in Figure 12, the maximum bending strength is much higher and is, respectively, approx. 2970, approx. 2030 and 1030 N, and maximum bending is, respectively, approx. 5, approx. 3.4 and 5.6%. Diagrams were also compared of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured by CAD/CAM by milling discs from Co-Cr alloy and ZrO2 sinter (Figure 12). The comparison of diagrams of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured from all the examined materials, that is, from selectively laser sintered titanium and Ti6Al4V alloy and with the CAD/CAM method by milling discs from Co-Cr alloy and ZrO2 sinter, shows that the highest maximum bending strengths are specific for bridge no. 2 regardless the material applied, and highest bending is seen for bridge no. 3.

Figure 12. Comparison of diagrams of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured from all the examined materials, that is, from selectively laser sintered titanium and Ti6Al4V alloy and with

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the CAD/CAM method by milling discs from Ti6Al4V alloy, Co-Cr alloy and sintered ZrO2 sinter.

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the case of a bridge made of Ti6Al4V alloy, from approx. 340 to approx. 720 N (Figure 11). In the case of selectively laser sintered Ti6Al4V alloy, each time, for each laser power, bending strength causing sample fracture is higher than for bridges produced from titanium, and this

Figure 11. Comparison of diagrams of dependency between bending strength and bending of solid bridge no. 3

Diagrams were compared of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured from all the examined materials, that is, from selectively laser sintered titanium and Ti6Al4V alloy and milled from Ti6Al4V alloy discs, Co-Cr alloy and milled ZrO2 sinter and then sintered (Figure 12). The maximum bending strength highly relies on the dental bridge type and varies from approx. 580 N for bridge no. 3 to approx. 1320 N for bridge no. 2. On the other hand, bridge no. 3 exhibits the maximum bending of approx. 10.4% while bridge no. 2 smallest bending of approx. 4.6%. Bridge no. 1 shows indirect properties. In the case of solid bridges nos 1, 2 and 3 manufactured by SLS from Ti6Al4V alloy with laser power of 110 W (Figure 12), the maximum bending strength of approx. 1570 N occurs for bridge no. 2, and the two other bridges show the strength of approx. 720 N. Maximum bending for bridges nos 3, 2 and 1 is, respectively, approx. 3.2, approx. 3 and approx. 2.4%. In the case of bridges nos. 2, 1 and 3 manufactured by CAD/ CAM by milling discs from Ti6Al4V alloy, as presented in Figure 12, the maximum bending strength is much higher and is, respectively, approx. 2970, approx. 2030 and 1030 N, and maximum bending is, respectively, approx. 5, approx. 3.4 and 5.6%. Diagrams were also compared of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured by CAD/CAM by milling discs from Co-Cr alloy and ZrO2 sinter (Figure 12). The comparison of diagrams of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured from all the examined materials, that is, from selectively laser sintered titanium and Ti6Al4V alloy and with the CAD/CAM method by milling discs from Co-Cr alloy and ZrO2 sinter, shows that the highest maximum bending strengths are specific for bridge no. 2 regardless the material applied, and highest

takes place in each case for relatively smaller bending (Figure 11).

manufactured by SLS from titanium and Ti6Al4V alloy for different laser power.

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bending is seen for bridge no. 3.

Figure 12. Comparison of diagrams of dependency between bending strength and bending of solid bridges nos. 1, 2 and 3 manufactured from all the examined materials, that is, from selectively laser sintered titanium and Ti6Al4V alloy and with the CAD/CAM method by milling discs from Ti6Al4V alloy, Co-Cr alloy and sintered ZrO2 sinter.

## 5. Structure of the investigated engineering materials manufactured by additive and machining technologies

It was revealed in examinations of the structure of the investigated materials in a scanning electron microscope that in the case of Co-Cr and Ti6Al4V alloy, the surface of the samples produced by CAD/CAM, by milling solid discs, exhibits roughness distinctive for precision machining (Figures 13 and 14), the same as for ceramic ZrO2 material which is sintered after performing machining (Figure 13). The examinations of chemical composition with the EDS method confirm the presence, in the first case, of Co, Cr and W (Figure 13) and in the second case of, respectively, Ti, Al and V (Figure 14) and in the third case of Zr and O (Figure 13), corresponding to the chemical composition of the materials applied. Figure 14 shows a surface structure of pristine titanium and Ti6Al4V alloy manufactured by selective laser sintering with

Figure 13. Surface structure of the samples produced by CAD/CAM by milling solid discs with the results of examinations of EDS chemical composition of (a) Co-Cr alloy, (b) sintered ZrO2 sinter after machining (SEM).

Figure 14. Surface structure of samples manufactured by selective laser sintering with the use of laser beam of 50 μm with laser power of 110 W of (a) solid titanium, (b) Ti6Al4V alloy and (c) Ti6Al4V alloy by CAD/CAM method by milling

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solid discs with results of examinations of EDS chemical composition (SEM).

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5. Structure of the investigated engineering materials manufactured by

It was revealed in examinations of the structure of the investigated materials in a scanning electron microscope that in the case of Co-Cr and Ti6Al4V alloy, the surface of the samples produced by CAD/CAM, by milling solid discs, exhibits roughness distinctive for precision machining (Figures 13 and 14), the same as for ceramic ZrO2 material which is sintered after performing machining (Figure 13). The examinations of chemical composition with the EDS method confirm the presence, in the first case, of Co, Cr and W (Figure 13) and in the second case of, respectively, Ti, Al and V (Figure 14) and in the third case of Zr and O (Figure 13), corresponding to the chemical composition of the materials applied. Figure 14 shows a surface structure of pristine titanium and Ti6Al4V alloy manufactured by selective laser sintering with

Figure 13. Surface structure of the samples produced by CAD/CAM by milling solid discs with the results of examina-

tions of EDS chemical composition of (a) Co-Cr alloy, (b) sintered ZrO2 sinter after machining (SEM).

additive and machining technologies

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Figure 14. Surface structure of samples manufactured by selective laser sintering with the use of laser beam of 50 μm with laser power of 110 W of (a) solid titanium, (b) Ti6Al4V alloy and (c) Ti6Al4V alloy by CAD/CAM method by milling solid discs with results of examinations of EDS chemical composition (SEM).

the laser beam size of 50 μm and laser power of 110 W. It was revealed in the both cases that apart from the completely sintered materials, with the shown paths of laser beam transition, fine powder particles exist, as well, which—after sintering—should be removed mechanically or by chemical etching.

X-ray tests allow to establish the phase composition of the investigated materials based on the data derived from international files. The following was identified, respectively, in particular samples: Co and Cr in Co-Cr, ZrO2 alloy, Ti in pristine titanium and Al0.3Ti0.7 phase and Ti0.8V0.2 in Ti6Al4V alloy irrespective of the manufacturing technology by milling from discs or selective laser sintering.

Investigations were carried out into the structure of the examined engineering materials subjected to the examinations of mechanical properties, and the structure was compared, especially of fractures of the examined materials produced by SLS, that is, sintered titanium and sintered Ti6Al4V alloy, and also manufactured by CAD/CAM by milling discs from solid Ti6Al4V, Co-Cr alloys and ZrO2 sinter milled and then sintered. The selectively laser sintered materials have a continuous structure without pores, as for titanium (Figure 15a), in which

Figure 15. Surface structure of solid samples manufactured by selective laser sintering (a) with the use of laser beam of 50 μm with laser power of 110 W, (b) fracture structure after a static bending test of samples of solid Ti6Al4V alloy manufactured by selective laser sintering with laser beam size of 50 μm after sintering with laser power of 70 W and (c, d) fracture structure after a static bending test of the samples produced by CAD/CAM by milling discs from solid alloys of (c) Ti6Al4V and (d) Co-Cr; SEM.

particular laser paths can be distinguished, provided that laser power for a laser beam with a diameter of 50 μm is 110 W. It is revealed by analysing a fracture of titanium, laser sintered with the laser power of 70 W, subjected to a static bending test, that—especially—finest powder particles are not sintered fully and also that a ductile fracture occurs in bridges of the sintered material. The samples sintered with the laser power of 110 W are solid and show a ductile fracture. A fracture structure of the samples from the laser sintered Ti6Al4V alloy after a static bending test is analogous to this obtained for titanium (Figure 15b) and—after

disc milling and then sintering, bridge no. 3; stereoscope microscope.

Figure 16. View of dental bridges after static bending test, manufactured from (a,b) titanium by SLS, (a) bridge no. 2 and (b) bridge no. 1; (c) from Ti6Al4V alloy by SLS bridge no. 3; and (d–f) fabricated by CAD/CAM by disc milling (d) from Ti6Al4V alloy, bridge no. 2; (e) from Co-Cr alloy, bridge no. 3; and (f) from ZrO2 sinter manufactured by CAD/CAM by

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the laser beam size of 50 μm and laser power of 110 W. It was revealed in the both cases that apart from the completely sintered materials, with the shown paths of laser beam transition, fine powder particles exist, as well, which—after sintering—should be removed mechanically

X-ray tests allow to establish the phase composition of the investigated materials based on the data derived from international files. The following was identified, respectively, in particular samples: Co and Cr in Co-Cr, ZrO2 alloy, Ti in pristine titanium and Al0.3Ti0.7 phase and Ti0.8V0.2 in Ti6Al4V alloy irrespective of the manufacturing technology by milling from discs

Investigations were carried out into the structure of the examined engineering materials subjected to the examinations of mechanical properties, and the structure was compared, especially of fractures of the examined materials produced by SLS, that is, sintered titanium and sintered Ti6Al4V alloy, and also manufactured by CAD/CAM by milling discs from solid Ti6Al4V, Co-Cr alloys and ZrO2 sinter milled and then sintered. The selectively laser sintered materials have a continuous structure without pores, as for titanium (Figure 15a), in which

Figure 15. Surface structure of solid samples manufactured by selective laser sintering (a) with the use of laser beam of 50 μm with laser power of 110 W, (b) fracture structure after a static bending test of samples of solid Ti6Al4V alloy manufactured by selective laser sintering with laser beam size of 50 μm after sintering with laser power of 70 W and (c, d) fracture structure after a static bending test of the samples produced by CAD/CAM by milling discs from solid alloys of

or by chemical etching.

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or selective laser sintering.

(c) Ti6Al4V and (d) Co-Cr; SEM.

Figure 16. View of dental bridges after static bending test, manufactured from (a,b) titanium by SLS, (a) bridge no. 2 and (b) bridge no. 1; (c) from Ti6Al4V alloy by SLS bridge no. 3; and (d–f) fabricated by CAD/CAM by disc milling (d) from Ti6Al4V alloy, bridge no. 2; (e) from Co-Cr alloy, bridge no. 3; and (f) from ZrO2 sinter manufactured by CAD/CAM by disc milling and then sintering, bridge no. 3; stereoscope microscope.

particular laser paths can be distinguished, provided that laser power for a laser beam with a diameter of 50 μm is 110 W. It is revealed by analysing a fracture of titanium, laser sintered with the laser power of 70 W, subjected to a static bending test, that—especially—finest powder particles are not sintered fully and also that a ductile fracture occurs in bridges of the sintered material. The samples sintered with the laser power of 110 W are solid and show a ductile fracture. A fracture structure of the samples from the laser sintered Ti6Al4V alloy after a static bending test is analogous to this obtained for titanium (Figure 15b) and—after sintering with the laser power of 110 W—shows a ductile fracture. A ductile fracture also exists in the case of fractographic tests of Ti6Al4V alloy but in the form of discs for sample milling (Figure 15c). The samples milled from Co-Cr alloy show, however, a brittle fracture, only with the local presence of a ductile fracture (Figure 15d).

As a zirconium oxide substructure can be completely white, but also after colouring with special paints before sintering, it can have a colour consistent with the VITA colour palette, a fully natural aesthetic effect can be achieved of a ready prosthetic restoration. It is also noteworthy that a translucence effect can be achieved, especially on secant edges of such a prosthetic restoration, responding to the most demanding patients' needs. Such effects can be attained on a substructure made of Co-Cr alloy. On the other hand, it is practically impossible to use titanium as a substructure for prosthetic crowns and bridges because the material tends to change the shade of porcelain to darker and colder; hence it is very difficult to ensure the manufacturing repeatability of prosthetic restorations and their quality. A solution ensuring the best possible aesthetic values is to use zirconium oxide as a substructure. Due to a substantial technological and therapeutic risk, a solution with the lowest strength properties is most expensive, though. Therefore, a dental engineer needs to be aware of the limitations of this material and has to design prosthetic restorations on a zirconium oxide substructure differently than when metallic materials are used. The author of this work prepares in his daily engineering practice all kinds of bridges with ¾ crowns, that is, such which are made on all the walls, except for the vestibular wall and the chewing surface, with a full structure, as so-called full contoured. Such a crown is coated from the vestibular part with veneering porcelain to ensure the best aesthetic values, and special paints and glaze are applied onto the other surfaces. The so-designed bridge, in the example shown in this work, would have walls with a thickness of approx. 2.5 mm instead of 0.6 mm, and a section area of abutments would be

Mechanical Properties Comparison of Engineering Materials Produced by Additive and Subtractive Technologies…

essentially relating to metallic materials. A much greater active section will allow to transfer higher loads, accordingly. The designer also has to keep in mind to support bridge spans on the periodontal tissue, never using the saddleback-like gum shape, but using an elliptic support with subgingival correction in each case. The so-designed prosthetic restoration will surely feature much higher strength properties than a prosthetic restoration with a reduced structure, being a reference point used in this work. Unfortunately, prosthetic restorations made on a zirconium oxide substructure with a simplified structure, as so-called caps with simple abutments, are still commonly used by dentists and dental technicians. The test results obtained seem not to confirm the applicability of the bridges, made of this material, designed this way. This puts into question a potential application of ZrO2 sinter bridges in dental practice without using a special design method of such dental restorations. This fact can be confirmed by cases, known from medical practice, when such bridges are damaged during normal use by patients. The results presented in this chapter can be directly applied in dental

The research presented in the chapter was realised in connection with the Project POIR.01.01.01-00-0485/16-00 on "IMSKA-MAT Innovative dental and maxillofacial implants

, as for the designs presented in this work and

http://dx.doi.org/10.5772/intechopen.70493

133

approx. 20–22 mm2 instead of 10–12 mm<sup>2</sup>

Acknowledgements

practice to remove consequences of teeth lost due to tooth decay.

In the case of all the examined metallic materials, and independently from the applied manufacturing technology, a fracture of samples occurs most often on a root of the abutments linking successive teeth, while in the case of bridges made of ZrO2, such bridges undergo brittle cracking in different places (Figure 16).

## 6. Suitability of application of additive manufacturing methods of prosthetic restorations of the dental system as compared to machining methods

Comparative investigations are presented in this chapter of the structure and properties of selected engineering materials used for dental prosthetic restorations manufactured alternatively by the machining method by milling on computerised numerical control (CNC) milling machines and by the additive selective laser sintering (SLS) method of solid and porous elements using CAD/CAM techniques; moreover, an original technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. The suitability was confirmed of applying the manufacturing technologies used in prosthodontics from powders by the SLS technique. The results of the executed tests of strength micro-samples and of the selected prosthetic bridges were compared. The bending strength of the titanium sintered selectively by laser, of the milled Ti6Al4V alloy and of the milled Co-Cr alloy is within the range of 1739–1959 MPa and tensile strength is within 739 and 858 MPa. It points out that the materials can be employed as prosthetic restorations, similarly as a selectively laser sintered Ti6Al4V alloy with the bending strength of approx. 2374 MPa and tensile strength of approx. 1312 MPa and pristine Ti with the bending strength of approx. 1682 MPa and tensile strength of 739 MPa. The SLS technology for titanium, and even more for Ti6Al4V alloy, ensures the achievement of mechanical properties comparable or better than a reference Co-Cr alloy which is commonly used for prosthetic restorations, including prosthetic bridges fabricated by milling solid discs with CNC milling machines. For all the examined engineering materials, the milled and then sintered ZrO2 material exhibits the lowest strength properties. The average tensile strength of this material is approx. 47 MPa, that is, it is nearly 20 times smaller than of Co-Cr alloy, and the average bending strength is smaller nearly 3 times than this of Co-Cr alloy and is approx. 605 MPa. The maximum elongation does not exceed, respectively, 6%, and maximum bending is approx. 0.3%, and the both values are much smaller than such appropriate for Co-Cr alloy. The microscopic examinations of the surface of prosthetic bridges fabricated from this material by the machining method of disc milling and then sintering show more advantageous roughness than bridges produced from Co-Cr alloy, and their fracture after bending is decisively brittle, and cracking takes place in different places, not as in the case of Co-Cr alloy. Prosthetic restorations on a substructure of sintered ZrO2 are, however, used more and more often, especially because it is necessary to ensure the best aesthetic values of the solution employed. As a zirconium oxide substructure can be completely white, but also after colouring with special paints before sintering, it can have a colour consistent with the VITA colour palette, a fully natural aesthetic effect can be achieved of a ready prosthetic restoration. It is also noteworthy that a translucence effect can be achieved, especially on secant edges of such a prosthetic restoration, responding to the most demanding patients' needs. Such effects can be attained on a substructure made of Co-Cr alloy. On the other hand, it is practically impossible to use titanium as a substructure for prosthetic crowns and bridges because the material tends to change the shade of porcelain to darker and colder; hence it is very difficult to ensure the manufacturing repeatability of prosthetic restorations and their quality. A solution ensuring the best possible aesthetic values is to use zirconium oxide as a substructure. Due to a substantial technological and therapeutic risk, a solution with the lowest strength properties is most expensive, though. Therefore, a dental engineer needs to be aware of the limitations of this material and has to design prosthetic restorations on a zirconium oxide substructure differently than when metallic materials are used. The author of this work prepares in his daily engineering practice all kinds of bridges with ¾ crowns, that is, such which are made on all the walls, except for the vestibular wall and the chewing surface, with a full structure, as so-called full contoured. Such a crown is coated from the vestibular part with veneering porcelain to ensure the best aesthetic values, and special paints and glaze are applied onto the other surfaces. The so-designed bridge, in the example shown in this work, would have walls with a thickness of approx. 2.5 mm instead of 0.6 mm, and a section area of abutments would be approx. 20–22 mm2 instead of 10–12 mm<sup>2</sup> , as for the designs presented in this work and essentially relating to metallic materials. A much greater active section will allow to transfer higher loads, accordingly. The designer also has to keep in mind to support bridge spans on the periodontal tissue, never using the saddleback-like gum shape, but using an elliptic support with subgingival correction in each case. The so-designed prosthetic restoration will surely feature much higher strength properties than a prosthetic restoration with a reduced structure, being a reference point used in this work. Unfortunately, prosthetic restorations made on a zirconium oxide substructure with a simplified structure, as so-called caps with simple abutments, are still commonly used by dentists and dental technicians. The test results obtained seem not to confirm the applicability of the bridges, made of this material, designed this way. This puts into question a potential application of ZrO2 sinter bridges in dental practice without using a special design method of such dental restorations. This fact can be confirmed by cases, known from medical practice, when such bridges are damaged during normal use by patients. The results presented in this chapter can be directly applied in dental practice to remove consequences of teeth lost due to tooth decay.

#### Acknowledgements

sintering with the laser power of 110 W—shows a ductile fracture. A ductile fracture also exists in the case of fractographic tests of Ti6Al4V alloy but in the form of discs for sample milling (Figure 15c). The samples milled from Co-Cr alloy show, however, a brittle fracture, only with

In the case of all the examined metallic materials, and independently from the applied manufacturing technology, a fracture of samples occurs most often on a root of the abutments linking successive teeth, while in the case of bridges made of ZrO2, such bridges undergo

Comparative investigations are presented in this chapter of the structure and properties of selected engineering materials used for dental prosthetic restorations manufactured alternatively by the machining method by milling on computerised numerical control (CNC) milling machines and by the additive selective laser sintering (SLS) method of solid and porous elements using CAD/CAM techniques; moreover, an original technology was presented of manufacturing the elements used in prosthodontics, produced with titanium and Ti6Al4V alloy powders with the SLS technique. The suitability was confirmed of applying the manufacturing technologies used in prosthodontics from powders by the SLS technique. The results of the executed tests of strength micro-samples and of the selected prosthetic bridges were compared. The bending strength of the titanium sintered selectively by laser, of the milled Ti6Al4V alloy and of the milled Co-Cr alloy is within the range of 1739–1959 MPa and tensile strength is within 739 and 858 MPa. It points out that the materials can be employed as prosthetic restorations, similarly as a selectively laser sintered Ti6Al4V alloy with the bending strength of approx. 2374 MPa and tensile strength of approx. 1312 MPa and pristine Ti with the bending strength of approx. 1682 MPa and tensile strength of 739 MPa. The SLS technology for titanium, and even more for Ti6Al4V alloy, ensures the achievement of mechanical properties comparable or better than a reference Co-Cr alloy which is commonly used for prosthetic restorations, including prosthetic bridges fabricated by milling solid discs with CNC milling machines. For all the examined engineering materials, the milled and then sintered ZrO2 material exhibits the lowest strength properties. The average tensile strength of this material is approx. 47 MPa, that is, it is nearly 20 times smaller than of Co-Cr alloy, and the average bending strength is smaller nearly 3 times than this of Co-Cr alloy and is approx. 605 MPa. The maximum elongation does not exceed, respectively, 6%, and maximum bending is approx. 0.3%, and the both values are much smaller than such appropriate for Co-Cr alloy. The microscopic examinations of the surface of prosthetic bridges fabricated from this material by the machining method of disc milling and then sintering show more advantageous roughness than bridges produced from Co-Cr alloy, and their fracture after bending is decisively brittle, and cracking takes place in different places, not as in the case of Co-Cr alloy. Prosthetic restorations on a substructure of sintered ZrO2 are, however, used more and more often, especially because it is necessary to ensure the best aesthetic values of the solution employed.

6. Suitability of application of additive manufacturing methods of prosthetic restorations of the dental system as compared to machining

the local presence of a ductile fracture (Figure 15d).

brittle cracking in different places (Figure 16).

132 Biomaterials in Regenerative Medicine

methods

The research presented in the chapter was realised in connection with the Project POIR.01.01.01-00-0485/16-00 on "IMSKA-MAT Innovative dental and maxillofacial implants manufactured using the innovative additive technology supported by computer-aided materials design ADDMAT" realised by the Medical and Dental Engineering Centre for Research, Design and Production ASKLEPIOS in Gliwice, Poland and co-financed by the National Centre for Research and Development in Warsaw, Poland.

[9] Kruth J-P, Mercelis P, Van Vaerenbergh J, Froyen L, Rombouts M. Binding mechanisms in selective laser sintering and selective laser melting. Rapid Prototyping Journal. 2005;

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135

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[16] Dobrzański LA, Dobrzańska-Danikiewicz AD, Malara P, Gaweł TG, Dobrzański LB, Achtelik-Franczak A. Fabrication of scaffolds from Ti6Al4V powders using the computer aided laser method. Archives of Metallurgy and Materials. 2015;60:1065-1070. DOI:

[17] Dobrzański LA, Dobrzańska-Danikiewicz AD, Achtelik-Franczak A, Dobrzański LB, Szindler M, Gaweł TG. Porous selective laser melted Ti and Ti6Al4V materials for medical applications. In: Dobrzański LA, editor. Powder Metallurgy – Fundamentals and

[18] Dobrzański LB. Struktura i własności materiałów inżynierskich na uzupełnienia protetyczne układu stomatognatycznego wytwarzane metodami przyrostowymi i ubytkowymi [Ph.D.

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## Author details

Lech B. Dobrzański

Address all correspondence to: lbdobrzanski@gmail.com

Medical and Dental Engineering Centre for Research, Design and Production ASKLEPIOS, Gliwice, Poland

## References


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manufactured using the innovative additive technology supported by computer-aided materials design ADDMAT" realised by the Medical and Dental Engineering Centre for Research, Design and Production ASKLEPIOS in Gliwice, Poland and co-financed by the National

Medical and Dental Engineering Centre for Research, Design and Production ASKLEPIOS,

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Author details

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Lech B. Dobrzański

Gliwice, Poland

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**Chapter 5**

**Provisional chapter**

**Properties of Co-Cr Dental Alloys Fabricated Using**

The aim of the present paper is to make a review of the properties of dental alloys, fabricated using Additive Technologies (AT). The microstructure and mechanical properties of Co-Cr alloys as well as the accuracy and surface roughness of dental constructions are discussed. In dentistry two different approaches can be applied for production of metal frameworks using AT. According to the first one the wax/polymeric cast patterns are fabricated by 3D printing, than the constructions are cast from dental alloy with asprinted patterns. Through the second one the metal framework is manufactured form powder alloy directly from 3D virtual model by Selective Electron Beam Melting (SEBM) or Selective Laser Melting (SLM). The microstructure and mechanical properties of Co-Cr dental alloys, cast using 3D printed patterns, are typical for cast alloys. Their dimensional and adjustment accuracy is higher comparing to constructions, produced by traditional lost-wax casting or by SLM. The surface roughness is higher than that of the samples, cast by conventional technology, but lower comparing to the SLM objects. The microstructure of SLM Co-Cr dental alloys is fine grained and more homogeneous comparing that of the cast alloys, which defines higher hardness and mechanical properties, higher wear and

**Keywords:** materials science and engineering, biomaterials, regenerative stomatology,

additive technologies, Co-Cr dental alloys, microstructure and properties

**Properties of Co-Cr Dental Alloys Fabricated Using** 

DOI: 10.5772/intechopen.69718

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Dental alloys on the basis of cobalt and chromium are one of the most preferred for production of metal frameworks of dental constructions because of their high strength, high corrosion and wear resistance, high biocompatibility, and a relatively low cost [1, 2]. The chemical composition of Co-Cr dental alloys consists of 53–67% of Co, 25–32% of Cr, 2–6% of Mo, and small

**Additive Technologies**

**Additive Technologies**

http://dx.doi.org/10.5772/intechopen.69718

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

Tsanka Dikova

**Abstract**

corrosion resistance.

**1. Introduction**

Tsanka Dikova


**Provisional chapter**

## **Properties of Co-Cr Dental Alloys Fabricated Using Additive Technologies Additive Technologies**

**Properties of Co-Cr Dental Alloys Fabricated Using** 

DOI: 10.5772/intechopen.69718

Tsanka Dikova Tsanka Dikova Additional information is available at the end of the chapter

[73] Dobrzański LA, Matula G, Dobrzańska-Danikiewicz AD, Malara P, Kremzer M, Tomiczek B, Kujawa M, Hajduczek E, Achtelik-Franczak A, Dobrzański LB, Krzysteczko J. Composite materials infiltrated by aluminium alloys based on porous skeletons from alumina, mullite and titanium produced by powder metallurgy techniques. In: Dobrzański LA, editor. Powder Metallurgy – Fundamentals and Case Studies. Rijeka: InTech; 2017.

[74] Dobrzańska J, Gołombek K, Dobrzański LB. Polymer materials used in endodontic treatment – In vitro testing. Archives of Materials Science and Engineering. 2012;58:

[75] Dobrzański LA. Applications of newly developed nanostructural and microporous materials in biomedical, tissue and mechanical engineering. Archives of Materials Science and

[76] Dobrzańska-Danikiewicz AD, Gaweł TG, Wolany W. Ti6Al4V Titanium alloy used as a modern biomimetic material. Archives of Materials Science and Engineering. 2015;76:

[77] Dobrzański LA, Dobrzańska-Danikiewicz AD, editors. Metalowe materiały mikroporowate do zastosowań medycznych. Vol. 7. Open Access Library; 2017 (in press)

[78] Malara P, Dobrzański LB. Computer aided manufacturing and design of fixed bridges restoring the lost dentition, soft tissue and the bone. Archives of Materials Science and

[79] Malara P, Dobrzański LB. Screw-retained full arch restorations – Methodology of computer aided design and manufacturing. Archives of Materials Science and Engineering.

Engineering. 2016;81:68-75. DOI: 10.5604/18972764.1230551

p. 95-137. DOI: 10.5772/65377

Engineering. 2015;76:53-114

110-115

140 Biomaterials in Regenerative Medicine

150-156

2017;83:23-29

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69718

#### **Abstract**

The aim of the present paper is to make a review of the properties of dental alloys, fabricated using Additive Technologies (AT). The microstructure and mechanical properties of Co-Cr alloys as well as the accuracy and surface roughness of dental constructions are discussed. In dentistry two different approaches can be applied for production of metal frameworks using AT. According to the first one the wax/polymeric cast patterns are fabricated by 3D printing, than the constructions are cast from dental alloy with asprinted patterns. Through the second one the metal framework is manufactured form powder alloy directly from 3D virtual model by Selective Electron Beam Melting (SEBM) or Selective Laser Melting (SLM). The microstructure and mechanical properties of Co-Cr dental alloys, cast using 3D printed patterns, are typical for cast alloys. Their dimensional and adjustment accuracy is higher comparing to constructions, produced by traditional lost-wax casting or by SLM. The surface roughness is higher than that of the samples, cast by conventional technology, but lower comparing to the SLM objects. The microstructure of SLM Co-Cr dental alloys is fine grained and more homogeneous comparing that of the cast alloys, which defines higher hardness and mechanical properties, higher wear and corrosion resistance.

**Keywords:** materials science and engineering, biomaterials, regenerative stomatology, additive technologies, Co-Cr dental alloys, microstructure and properties

## **1. Introduction**

Dental alloys on the basis of cobalt and chromium are one of the most preferred for production of metal frameworks of dental constructions because of their high strength, high corrosion and wear resistance, high biocompatibility, and a relatively low cost [1, 2]. The chemical composition of Co-Cr dental alloys consists of 53–67% of Co, 25–32% of Cr, 2–6% of Mo, and small

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

quantities of W, Si, Al, and others [3]. Cr, Mo, and W are added for strengthening of the solid solution. Due to the relatively large amount of Cr, dense passive layer of Cr2 O3 with 1–4 nm thickness on the surface as well as carbides in the microstructure of the details is formed, determining the high hardness, high corrosion, and wear resistance [4, 5]. According to the phase diagram, Co-Cr dental alloys are characterized with face-centered cubic (fcc) lattice, γ phase in high temperatures, and with hexagonal close packed (hcp) lattice, ε phase in room temperature [4, 6]. The γ phase defines the ductility, while the ε phase defines the corrosion and wear resistance of the alloy [7]. In proper alloying, the microstructure of the Co-Cr dental alloys consists mainly of γ phase and carbides of the M23C6 type [4]. Therefore, the properties of Co-Cr dental alloys depend on the ratio between γ/ε phases and the type, quantity and distribution of the carbide phase in the microstructure.

the parts are manufactured by melting metal powder layer per layer with an electron beam in a high vacuum [8, 9, 14, 16]. In SLS/SLM technology, layers of particular powder material (mainly polymers and porcelains in SLS and pure metals or alloys in SLM) are fused into a real detail by a computer-directed laser [8, 9, 12, 17]. During IJP process, an extremely small ink droplet is ejected toward the substrate. Different substances can be used as ink aqueous solution of colouring agents and binders to a ceramic suspension to produce zirconia dental restorations [8, 9, 12, 18]. Another variant of the IJP is by depositing droplets of a polymer, and each formed layer is cured by UV light, which allows polymer material with characteristics

Properties of Co-Cr Dental Alloys Fabricated Using Additive Technologies

http://dx.doi.org/10.5772/intechopen.69718

143

The main advantages of AT include production of complex objects from different materials polymers, composites, metals and alloys; manufacturing of parts with dense/porous structure and predetermined surface roughness and controllable, is an easy and relatively quick process [9]. Due to the various materials used and the great variety of additive manufacturing processes, AT can be successfully applied in many fields of dentistry for the production of different types of dental constructions. AT give even opportunities to fabricate structures of hard-to-handle materials such as cobalt and chromium alloys [19]. Polymeric study models instead of dental plaster casts, surgical guides for placement of dental implants, temporary crowns and bridges as well as resin models for lost-wax casting can be fabricated by SLA and FDM [8, 9, 12, 20–22]. SEBM and SLM are used for manufacturing of customized implants for maxillofacial surgery, dense or porous dental implants, dental crowns and bridges as well as partial denture frameworks [23–31]. IJP is suitable for printing of zirconia dental restorations, polymeric dental models, orthodontic bracket guides, wax patterns of complete dentures, or wax models for casting [8, 9, 32, 33].

There are two main approaches for the production of metal constructions from Co-Cr dental alloys using AT. The first one concerns to the fabrication of wax/polymeric cast patterns by 3D printing, and the second one is casting of the framework from dental alloy. Through the second one, the metal object is manufactured directly from the 3D virtual model by SEBM or SLM. As the additive technologies are relatively new, develop extremely fast and with their indisputable advantages enter in many areas, so their implementation into the dentistry is faster than the research and data about the quality of the details produced by them. The aim of the present chapter is to make a review of the properties of dental alloys, fabricated using additive technologies. The microstructure and mechanical properties of Co-Cr alloys as well

as the accuracy and surface roughness of dental constructions are discussed.

**2. Properties of Co-Cr dental alloys cast with 3D-printed patterns**

The wax/polymeric cast patterns can be produced from generated 3D virtual models by laserassisted or digital light projection (DLP) stereolithography, fused deposition modeling, and ink-jet printing. The type of the additive manufacturing process, the parameters of its technological regimes, and the properties of the used materials influence mainly on the geometrical characteristics, adjustment accuracy, and surface roughness of the cast patterns [33–36], thus defining the accuracy and surface quality of the cast dental constructions. As with any AT,

similar to the technical wax to be used [8, 9, 13].

The microstructure and properties of dental alloys depend on the manufacturing process and the technological regimes. Since the beginning of the last century, casting is the most common technology for production of metal constructions in dentistry. During this conventional technology, the metal frameworks are fabricated by centrifugal casting using hand-built wax patterns. The technological process is characterized with large amount of manual work, and although it is performed by a qualified dental technician, it can lead to low accuracy and satisfactory quality. The advent of the modern CAD/CAM systems and the additive technologies in the last 30 years of the last century allow to decrease the amount of manual work and production time, to improve the quality of dental constructions and as a consequence to reduce their price [8, 9].

The additive technologies (AT) are developed in the late 1980s as an alternative of the subtractive technologies. They are characterized with building of one layer at a time from a powder or liquid that is bonded by means of melting, fusing, or polymerization. The American Society for Testing and Materials (ASTM) defined additive manufacturing as 'the process of joining materials to make objects from 3D model data, usually layer upon layer, as opposed to subtractive manufacturing methodologies' [8–10]. These processes are also known as 'threedimensional printing', 'layered manufacturing', 'free-form fabrication', 'rapid prototyping' and 'rapid manufacturing' [9–11]. The ASTM international committee, intended for specification of standards for additive manufacturing—ASTM F42—created a categorization of all 3D printing technologies into seven major groups [11]. According to it, the following 3D printing technologies are used in work with biomaterials: 3D plotting/direct ink writing, laser-assisted bioprinting, selective laser sintering, stereolithography, fused deposition modeling, and robot-assisted deposition/robocasting. The stereolithography (SLA), fused deposition modeling (FDM), selective electron beam melting (SEBM), selective laser sintering (SLS), selective laser melting (SLM), and ink-jet printing (IJP) are among the AT, mostly used in the dental medicine [8, 9, 12–15].

During the SLA process, a concentrated beam of UV light is focused on the surface of a tank filled with liquid photopolymer. As the light beam draws the object on the surface of the liquid, each time a layer of resin is polymerized or cross-linked until the real object is obtained [8, 9, 12, 14, 16]. The FDM characterizes with extruding the thermoplastic materials through heated nozzle, or the material is fed from a reservoir through a syringe [9, 12, 16]. By SEBM, the parts are manufactured by melting metal powder layer per layer with an electron beam in a high vacuum [8, 9, 14, 16]. In SLS/SLM technology, layers of particular powder material (mainly polymers and porcelains in SLS and pure metals or alloys in SLM) are fused into a real detail by a computer-directed laser [8, 9, 12, 17]. During IJP process, an extremely small ink droplet is ejected toward the substrate. Different substances can be used as ink aqueous solution of colouring agents and binders to a ceramic suspension to produce zirconia dental restorations [8, 9, 12, 18]. Another variant of the IJP is by depositing droplets of a polymer, and each formed layer is cured by UV light, which allows polymer material with characteristics similar to the technical wax to be used [8, 9, 13].

quantities of W, Si, Al, and others [3]. Cr, Mo, and W are added for strengthening of the solid

thickness on the surface as well as carbides in the microstructure of the details is formed, determining the high hardness, high corrosion, and wear resistance [4, 5]. According to the phase diagram, Co-Cr dental alloys are characterized with face-centered cubic (fcc) lattice, γ phase in high temperatures, and with hexagonal close packed (hcp) lattice, ε phase in room temperature [4, 6]. The γ phase defines the ductility, while the ε phase defines the corrosion and wear resistance of the alloy [7]. In proper alloying, the microstructure of the Co-Cr dental alloys consists

alloys depend on the ratio between γ/ε phases and the type, quantity and distribution of the

The microstructure and properties of dental alloys depend on the manufacturing process and the technological regimes. Since the beginning of the last century, casting is the most common technology for production of metal constructions in dentistry. During this conventional technology, the metal frameworks are fabricated by centrifugal casting using hand-built wax patterns. The technological process is characterized with large amount of manual work, and although it is performed by a qualified dental technician, it can lead to low accuracy and satisfactory quality. The advent of the modern CAD/CAM systems and the additive technologies in the last 30 years of the last century allow to decrease the amount of manual work and production time, to improve the quality of dental constructions and as a consequence to

The additive technologies (AT) are developed in the late 1980s as an alternative of the subtractive technologies. They are characterized with building of one layer at a time from a powder or liquid that is bonded by means of melting, fusing, or polymerization. The American Society for Testing and Materials (ASTM) defined additive manufacturing as 'the process of joining materials to make objects from 3D model data, usually layer upon layer, as opposed to subtractive manufacturing methodologies' [8–10]. These processes are also known as 'threedimensional printing', 'layered manufacturing', 'free-form fabrication', 'rapid prototyping' and 'rapid manufacturing' [9–11]. The ASTM international committee, intended for specification of standards for additive manufacturing—ASTM F42—created a categorization of all 3D printing technologies into seven major groups [11]. According to it, the following 3D printing technologies are used in work with biomaterials: 3D plotting/direct ink writing, laser-assisted bioprinting, selective laser sintering, stereolithography, fused deposition modeling, and robot-assisted deposition/robocasting. The stereolithography (SLA), fused deposition modeling (FDM), selective electron beam melting (SEBM), selective laser sintering (SLS), selective laser melting (SLM), and ink-jet printing (IJP) are among the AT, mostly used in the dental

During the SLA process, a concentrated beam of UV light is focused on the surface of a tank filled with liquid photopolymer. As the light beam draws the object on the surface of the liquid, each time a layer of resin is polymerized or cross-linked until the real object is obtained [8, 9, 12, 14, 16]. The FDM characterizes with extruding the thermoplastic materials through heated nozzle, or the material is fed from a reservoir through a syringe [9, 12, 16]. By SEBM,

O3

type [4]. Therefore, the properties of Co-Cr dental

with 1–4 nm

solution. Due to the relatively large amount of Cr, dense passive layer of Cr2

mainly of γ phase and carbides of the M23C6

carbide phase in the microstructure.

142 Biomaterials in Regenerative Medicine

reduce their price [8, 9].

medicine [8, 9, 12–15].

The main advantages of AT include production of complex objects from different materials polymers, composites, metals and alloys; manufacturing of parts with dense/porous structure and predetermined surface roughness and controllable, is an easy and relatively quick process [9]. Due to the various materials used and the great variety of additive manufacturing processes, AT can be successfully applied in many fields of dentistry for the production of different types of dental constructions. AT give even opportunities to fabricate structures of hard-to-handle materials such as cobalt and chromium alloys [19]. Polymeric study models instead of dental plaster casts, surgical guides for placement of dental implants, temporary crowns and bridges as well as resin models for lost-wax casting can be fabricated by SLA and FDM [8, 9, 12, 20–22]. SEBM and SLM are used for manufacturing of customized implants for maxillofacial surgery, dense or porous dental implants, dental crowns and bridges as well as partial denture frameworks [23–31]. IJP is suitable for printing of zirconia dental restorations, polymeric dental models, orthodontic bracket guides, wax patterns of complete dentures, or wax models for casting [8, 9, 32, 33].

There are two main approaches for the production of metal constructions from Co-Cr dental alloys using AT. The first one concerns to the fabrication of wax/polymeric cast patterns by 3D printing, and the second one is casting of the framework from dental alloy. Through the second one, the metal object is manufactured directly from the 3D virtual model by SEBM or SLM. As the additive technologies are relatively new, develop extremely fast and with their indisputable advantages enter in many areas, so their implementation into the dentistry is faster than the research and data about the quality of the details produced by them. The aim of the present chapter is to make a review of the properties of dental alloys, fabricated using additive technologies. The microstructure and mechanical properties of Co-Cr alloys as well as the accuracy and surface roughness of dental constructions are discussed.

## **2. Properties of Co-Cr dental alloys cast with 3D-printed patterns**

The wax/polymeric cast patterns can be produced from generated 3D virtual models by laserassisted or digital light projection (DLP) stereolithography, fused deposition modeling, and ink-jet printing. The type of the additive manufacturing process, the parameters of its technological regimes, and the properties of the used materials influence mainly on the geometrical characteristics, adjustment accuracy, and surface roughness of the cast patterns [33–36], thus defining the accuracy and surface quality of the cast dental constructions. As with any AT, the layer's thickness, the position of the object toward the print direction, and the optical properties of the polymers (in SLA) have a decisive effect on the object's accuracy. The thinner the layer, the lower the angle to the print direction, and the lower the surface roughness, the higher the resolution and the dimensional accuracy, but the longer the production time [37].

Ishida and Miyasaka [44] investigated the accuracy of the patterns for casting of all metal crowns, manufactured by four 3D printers, using different manufacturing processes: laserassisted SLA, SLA with concentrated UV beam, FDM, and IJP. They established that in all types of manufacturing process, the outer and the inner diameters of the crowns are smaller than that of the virtual 3D model, which could be compensated with 3–5% increasing the dimensions of the virtual model. The FDM-printed crowns are characterized with the highest surface roughness, while the lowest is typical for the crowns, fabricated by laser-assisted SLA. Additionally, the difference of the roughness values along the 3D axes exists. They established that the roughness along the tooth axis of the crown created by laser-assisted SLA was the smallest, while that along the horizontal direction of the crown printed by multi-jet modeling (MJM) was the smallest. This can be explained with the specific features of the additive

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The higher roughness of the surfaces, parallel, or inclined to the print direction (*Z* axes), compared to that along *X* and *Y* axes is also established in the research of Dikova et al. [46]. Their investigations of cubic samples, printed with different polymers by DLP SLA, show that the roughness of the surfaces of the cubes with horizontal position toward the basis is less than those of the cubes, printed inclined in almost all types of polymers. All sizes of the cubes in both positions are larger than the size of the virtual model. In their subsequent investigations [47], it is observed that the decrease of the layer's thickness from 50 (recommended by the equipment manufacturer) to 35 μm leads to nearly twice decrease of the surface roughness of the samples, made of resin NextDent Cast (developed especially for manufacturing of cast patterns), in both printing positions. It also leads to the highest dimensional accuracy and the least interval of deviation. The research team has shown that the dimensions, parallel to the basis, axes *X* and *Y*, are the most precise, while those, parallel or inclined to the printing direction, axis Z, are the most deviating. The dimensional and adjustment accuracy as well as the surface roughness of four-part dental bridges, made of polymers NextDent Cast and NextDent C+B (for temporary crowns and bridges) by DLP SLA, are investigated in the work of Dikova et al. [48]. They established that the dimensions in both directions of the bridges from both polymers, fabricated with less layer's thickness of 35 μm, are smaller with 0.29–1.10%, compared to those of the virtual model, while the sizes of the bridges with larger layer's thickness of 50 μm are 1.51–3.45% more than the virtual model. Dimensions of the samples of both polymers, situated in the building direction, are larger with 1.51–3.45%, and those, which are not in the print direction, are 0.49–0.53% smaller than those of the virtual 3D model. The inaccuracy in the geometrical dimensions causes inaccuracy of adjusting and absence of gap between the bridge constructions and the gypsum model. The surface roughness of the cast patterns, manufactured with regime, recommended by the producer (50 μm layer's thickness), is relatively high with average arithmetic deviation of the surface roughness *Ra* = 3.24 μm. Decreasing the layer's thickness to 35 μm leads to the lower surface roughness of *Ra* = 2.18 μm. The surface observations with optical microscopy proved the layered structure, which is a specific feature of the objects built up via stereolithography. The higher roughness of the 3D-printed cast patterns can complicate the cast process and cause higher

manufacturing process [8, 12, 45].

roughness of the cast itself.

The accuracy of dental constructions, produced by the modern CAD/CAM systems, is defined in the standard ISO 12836:2015 'Dentistry—Digitizing devices for CAD/CAM systems for indirect dental restorations—Test methods for assessing accuracy' [38]. Three standardized geometrical figures for accuracy evaluation are described in its Annexes A, B, and C [38]. Braian et al. [39] used the figures in the Annexes A and C, specifying the measurement of an inlay-shaped object and a multiunit specimen to simulate a four-unit bridge, for determining the production tolerances of four commercially available additive manufacturing systems, working on the SLS, multi-jet, or poly-jet principles. The samples were printed with different layer's thicknesses, optimal for each system, and ensuring the highest product's accuracy (as per the equipment manufacturer). According to the ISO/IEC GUIDE 99:2007(E/F) [40], the 'accuracy' is the closeness of agreement between a measured quantity value and a true quantity value of a measured object. The authors [39] defined the terms 'resolution' and 'repeatability'. The resolution refers to the smallest feature that the system can produce, while the reproducibility is described as the system ability to produce consistent output time after time. The researchers established that the accuracy of both types of samples, produced with four printers, is different in the three directions. The multi-jet printing ensures the highest accuracy of the linear as well as angular dimensions, followed by poly-jet process and SLS. It was concluded that the suitable type of printer should be chosen according to the intended dental application.

In evaluation of restorations, fabricated by digital technologies, subtractive (milling of wax and zirconia) and additive (SLA of photopolymer and SLS of Co-Cr alloy), Bae et al. [41] found out that the dimensions of the samples of all four groups are smaller than the reference data. Concerning to the repeatability of AT, they established the smallest difference from the reference data in the specimens, fabricated using SLA, and no significant differences between the SLA and SLS methods. Therefore, they concluded that the accuracy of additive manufacturing methods is better than the subtractive ones, as the mean accuracy discrepancies are the smallest for SLA, followed by the SLS, wax, and zirconia milling. But they pointed out the following as disadvantages: the print layers, clearly seen on the surface of the SLA sample, and circular, sunken forms of approximately 80 μm in diameter on the SLS specimen surface. The investigations of Mai et al. [42] proved that the crowns fabricated by CAD/CAM systems (milling and polymer-jet 3D printing) have higher fitting accuracy than that of the molded ones. Among them, the polymer-jet 3D printing significantly enhanced the fit of the interim crowns, particularly in the occlusal region. According to Kim et al. [43], the fitting accuracy of dental crowns is affected by the number of copings fabricated by micro-stereolithography. Based on the marginal discrepancy, the most precise copings are printed when three arrays are used on a single-built platform. Therefore, there are limitations concerning to the reproducibility and accuracy of polymeric dental constructions, manufactured by 3D printing.

Ishida and Miyasaka [44] investigated the accuracy of the patterns for casting of all metal crowns, manufactured by four 3D printers, using different manufacturing processes: laserassisted SLA, SLA with concentrated UV beam, FDM, and IJP. They established that in all types of manufacturing process, the outer and the inner diameters of the crowns are smaller than that of the virtual 3D model, which could be compensated with 3–5% increasing the dimensions of the virtual model. The FDM-printed crowns are characterized with the highest surface roughness, while the lowest is typical for the crowns, fabricated by laser-assisted SLA. Additionally, the difference of the roughness values along the 3D axes exists. They established that the roughness along the tooth axis of the crown created by laser-assisted SLA was the smallest, while that along the horizontal direction of the crown printed by multi-jet modeling (MJM) was the smallest. This can be explained with the specific features of the additive manufacturing process [8, 12, 45].

the layer's thickness, the position of the object toward the print direction, and the optical properties of the polymers (in SLA) have a decisive effect on the object's accuracy. The thinner the layer, the lower the angle to the print direction, and the lower the surface roughness, the higher the resolution and the dimensional accuracy, but the longer the production time [37]. The accuracy of dental constructions, produced by the modern CAD/CAM systems, is defined in the standard ISO 12836:2015 'Dentistry—Digitizing devices for CAD/CAM systems for indirect dental restorations—Test methods for assessing accuracy' [38]. Three standardized geometrical figures for accuracy evaluation are described in its Annexes A, B, and C [38]. Braian et al. [39] used the figures in the Annexes A and C, specifying the measurement of an inlay-shaped object and a multiunit specimen to simulate a four-unit bridge, for determining the production tolerances of four commercially available additive manufacturing systems, working on the SLS, multi-jet, or poly-jet principles. The samples were printed with different layer's thicknesses, optimal for each system, and ensuring the highest product's accuracy (as per the equipment manufacturer). According to the ISO/IEC GUIDE 99:2007(E/F) [40], the 'accuracy' is the closeness of agreement between a measured quantity value and a true quantity value of a measured object. The authors [39] defined the terms 'resolution' and 'repeatability'. The resolution refers to the smallest feature that the system can produce, while the reproducibility is described as the system ability to produce consistent output time after time. The researchers established that the accuracy of both types of samples, produced with four printers, is different in the three directions. The multi-jet printing ensures the highest accuracy of the linear as well as angular dimensions, followed by poly-jet process and SLS. It was concluded that the suitable type of printer should be chosen according to the intended dental

In evaluation of restorations, fabricated by digital technologies, subtractive (milling of wax and zirconia) and additive (SLA of photopolymer and SLS of Co-Cr alloy), Bae et al. [41] found out that the dimensions of the samples of all four groups are smaller than the reference data. Concerning to the repeatability of AT, they established the smallest difference from the reference data in the specimens, fabricated using SLA, and no significant differences between the SLA and SLS methods. Therefore, they concluded that the accuracy of additive manufacturing methods is better than the subtractive ones, as the mean accuracy discrepancies are the smallest for SLA, followed by the SLS, wax, and zirconia milling. But they pointed out the following as disadvantages: the print layers, clearly seen on the surface of the SLA sample, and circular, sunken forms of approximately 80 μm in diameter on the SLS specimen surface. The investigations of Mai et al. [42] proved that the crowns fabricated by CAD/CAM systems (milling and polymer-jet 3D printing) have higher fitting accuracy than that of the molded ones. Among them, the polymer-jet 3D printing significantly enhanced the fit of the interim crowns, particularly in the occlusal region. According to Kim et al. [43], the fitting accuracy of dental crowns is affected by the number of copings fabricated by micro-stereolithography. Based on the marginal discrepancy, the most precise copings are printed when three arrays are used on a single-built platform. Therefore, there are limitations concerning to the reproducibility and accuracy of polymeric dental construc-

application.

144 Biomaterials in Regenerative Medicine

tions, manufactured by 3D printing.

The higher roughness of the surfaces, parallel, or inclined to the print direction (*Z* axes), compared to that along *X* and *Y* axes is also established in the research of Dikova et al. [46]. Their investigations of cubic samples, printed with different polymers by DLP SLA, show that the roughness of the surfaces of the cubes with horizontal position toward the basis is less than those of the cubes, printed inclined in almost all types of polymers. All sizes of the cubes in both positions are larger than the size of the virtual model. In their subsequent investigations [47], it is observed that the decrease of the layer's thickness from 50 (recommended by the equipment manufacturer) to 35 μm leads to nearly twice decrease of the surface roughness of the samples, made of resin NextDent Cast (developed especially for manufacturing of cast patterns), in both printing positions. It also leads to the highest dimensional accuracy and the least interval of deviation. The research team has shown that the dimensions, parallel to the basis, axes *X* and *Y*, are the most precise, while those, parallel or inclined to the printing direction, axis Z, are the most deviating. The dimensional and adjustment accuracy as well as the surface roughness of four-part dental bridges, made of polymers NextDent Cast and NextDent C+B (for temporary crowns and bridges) by DLP SLA, are investigated in the work of Dikova et al. [48]. They established that the dimensions in both directions of the bridges from both polymers, fabricated with less layer's thickness of 35 μm, are smaller with 0.29–1.10%, compared to those of the virtual model, while the sizes of the bridges with larger layer's thickness of 50 μm are 1.51–3.45% more than the virtual model. Dimensions of the samples of both polymers, situated in the building direction, are larger with 1.51–3.45%, and those, which are not in the print direction, are 0.49–0.53% smaller than those of the virtual 3D model. The inaccuracy in the geometrical dimensions causes inaccuracy of adjusting and absence of gap between the bridge constructions and the gypsum model. The surface roughness of the cast patterns, manufactured with regime, recommended by the producer (50 μm layer's thickness), is relatively high with average arithmetic deviation of the surface roughness *Ra* = 3.24 μm. Decreasing the layer's thickness to 35 μm leads to the lower surface roughness of *Ra* = 2.18 μm. The surface observations with optical microscopy proved the layered structure, which is a specific feature of the objects built up via stereolithography. The higher roughness of the 3D-printed cast patterns can complicate the cast process and cause higher roughness of the cast itself.

The geometrical accuracy of fixed dental prostheses, manufactured by additive technologies, is investigated in Ref. [49]. Four-part dental bridges are produced of Co-Cr alloy by three technological processes: conventional lost-wax casting with wax models, manufactured in silicon mold; lost-wax casting with 3D-printed (multi-jet modeling (MJM) cast patterns, and direct fabricating by SLM. It was established that the surface roughness of Co-Cr fixed partial dentures, cast with 3D-printed wax patterns, is more than three times higher than the roughness of dentures, cast by conventional lost-wax technology (*Ra* = 3.39 μm and *Ra* = 1.11 μm, respectively). The surface observation of Co-Cr dental bridges proves the higher roughness as well as the traces of the layered manufacturing of the cast patterns, left even after sandblasting. The fixed partial dentures, cast with 3D-printed cast patterns, possess higher accuracy of the shape, sizes, and adjustment compared to the dentures, produced by conventional lostwax casting. This is mainly due to the minimal manual work, because 3D-printed cast patterns are produced directly from the virtual model.

For fabricating the high-quality castings from Co-Cr dental alloys using 3D-printed patterns, some steps concerning to the processes of 3D printing and casting should be observed. In 3D printing (1) the printer, ensuring the cast pattern with high dimensional and fitting accuracy and satisfactory surface roughness, has to be chosen. The present review shows that the systems, working on the principle of the laser-assisted SLA, DLP SLA, MJM, and polymer-jet printing, are good candidates; (2) As each printer works with its specific materials, the right material, meeting the requirements for manufacturing the cast patterns, should be used; (3) The optimal technological regime has to be chosen, including the layer's thickness; (4) The strategy of printing has to be developed—positions, and the number of the objects; (5) Preliminary tests, calibration on the three axes *X*, *Y,* and *Z* of the printer and compensation (if needed), have to be done to guarantee dimensions with high accuracy; and (6) 3D printing, post-curing (in some processes), and final surface treatment to enhance the smoothness. In the second part, the casting processes are (1) selection of the investment material, relevant to the material of the 3D-printed cast pattern, should be done for manufacturing casting mold; (2) heating of the casting mold with temperature regime concerning to the material of the 3D-printed cast pattern and the Co-Cr alloy and (3) casting with the given Co-Cr alloy,

Properties of Co-Cr Dental Alloys Fabricated Using Additive Technologies

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147

The properties of the Co-Cr dental alloys depend on the microstructure, its morphology, and composition, which are defined by the manufacturing process and the technological regimes. The SLM is a complex thermophysical process, depending on a number of important parameters. During SLM, layers of metal powder are fused into a real object by a computer-directed laser. The process characterizes with high heating and cooling rates, leading to fine-grained microstructure of the solidified layer. As the heat is led away through the solid body, phase transformations run in the underneath layers heated above the transition temperatures [50]. Due to the high-temperature gradients during the SLM process, high residual stresses are generated in the details, which can cause subsequent deformations [17, 52–54]. These characteristics determine the specific microstructure and properties of the objects, produced by SLM as compared with that, manufactured by casting. The main technological parameters that are crucial for the production of high-quality construction are the laser power, scanning speed, laser beam diameter and distance between the traces, layer's thickness, and the working area [55–58]. In the development of any process for production of an object by SLM, it is necessary to evaluate the density, accuracy, surface roughness, hardness, and strength properties.

In SLM process, the volume of the molten metal depends on the volume energy density Ev

son for the porous structure [50]. This can be avoided by optimization of the technological parameters—increasing the laser power or decreasing the scanning speed. Varying with the input energy and scan spaces, Takaichi et al. [59] established that dense structure of SLM Co-29Cr-6Mo alloy can be obtained when the input energy of the laser scan increased more

deposited layer will occur. Therefore, the lower volume energy density Ev

,

is the main rea-

and inversely proportional to the scan-

is insufficient, the incomplete melting of the

keeping the manufacturer's requirement.

**3. Properties of Co-Cr dental alloys fabricated via SLM**

which is directly proportional to the power density Ns

ning speed V. If the volume energy density Ev

The microstructure of Co-Cr sample, cast with 3D-printed pattern (MJM), is a typical cast microstructure—inhomogeneous, consisting of large grains with dendrite morphology [50]. A large amount of lamellar and blocky carbides of different sizes are located mainly along the grain boundaries. The dendrites consist of γ phase, while the inter-dendritic regions consist of microeutectic (Co solid solution with intermetallic precipitations) and carbides of the (Cr, Mo)23C6 type. The hardness of the Co-Cr samples, cast with 3D-printed patterns, is in the range 327HV–343HV with uneven distribution due to the inhomogeneous microstructure. The dendritic grains with intermetallic phases, precipitated along the grain boundaries, were observed in the microstructure of the samples, cast with 3D-printed patterns from Co-Cr alloy remanium star CL by Kim et al. [51]. They characterized with yield strength 540 ± 20 MPa, mean percent elongation 10 ± 2 and Young's modulus 260 ± 20 GPa.

The microstructure and strength properties of Co-Cr dental alloys, cast with 3D-printed patterns, are typical for castings. Their dimensional and fitting accuracy is higher than that of the constructions, manufactured by conventional casting, while their surface roughness is more than three times higher. Therefore, the dimensional and fitting accuracy and the surface roughness of Co-Cr dental substructures are strongly influenced by the quality of the cast patterns fabricated via AT. The data about the dimensional accuracy of wax/polymeric dental constructions, produced by different additive manufacturing processes, are contrary. In some cases, the dimensions of the printed objects are smaller than the virtual model; in another case, they are larger. It depends not only on the type of the manufacturing process but also on the scale factor—the sizes of the printed object. As a consequence of the low dimensional accuracy, the fitting accuracy is lower. The strategy of printing should be chosen very carefully, because too much objects in the same built platform can cause lower dimensional and fitting accuracy. Due to the features of the 3D printing processes, the sizes of the detail can be different in the three directions *X*, *Y,* and *Z*; therefore, it should be paid attention to the position of the object, especially toward the print direction. This will influence the surface roughness, too. The parameters of the technological regimes also influence the dimensional accuracy. Decreasing the layer's thickness leads to smaller sizes and lower roughness, which in some cases defines the higher accuracy. The special requirements to the materials, used for 3D printing of cast patterns, exist—to have no burned out residue and null or minimal thermal expansion—otherwise, the casting with low quality will be the result.

For fabricating the high-quality castings from Co-Cr dental alloys using 3D-printed patterns, some steps concerning to the processes of 3D printing and casting should be observed. In 3D printing (1) the printer, ensuring the cast pattern with high dimensional and fitting accuracy and satisfactory surface roughness, has to be chosen. The present review shows that the systems, working on the principle of the laser-assisted SLA, DLP SLA, MJM, and polymer-jet printing, are good candidates; (2) As each printer works with its specific materials, the right material, meeting the requirements for manufacturing the cast patterns, should be used; (3) The optimal technological regime has to be chosen, including the layer's thickness; (4) The strategy of printing has to be developed—positions, and the number of the objects; (5) Preliminary tests, calibration on the three axes *X*, *Y,* and *Z* of the printer and compensation (if needed), have to be done to guarantee dimensions with high accuracy; and (6) 3D printing, post-curing (in some processes), and final surface treatment to enhance the smoothness. In the second part, the casting processes are (1) selection of the investment material, relevant to the material of the 3D-printed cast pattern, should be done for manufacturing casting mold; (2) heating of the casting mold with temperature regime concerning to the material of the 3D-printed cast pattern and the Co-Cr alloy and (3) casting with the given Co-Cr alloy, keeping the manufacturer's requirement.

#### **3. Properties of Co-Cr dental alloys fabricated via SLM**

The geometrical accuracy of fixed dental prostheses, manufactured by additive technologies, is investigated in Ref. [49]. Four-part dental bridges are produced of Co-Cr alloy by three technological processes: conventional lost-wax casting with wax models, manufactured in silicon mold; lost-wax casting with 3D-printed (multi-jet modeling (MJM) cast patterns, and direct fabricating by SLM. It was established that the surface roughness of Co-Cr fixed partial dentures, cast with 3D-printed wax patterns, is more than three times higher than the roughness of dentures, cast by conventional lost-wax technology (*Ra* = 3.39 μm and *Ra* = 1.11 μm, respectively). The surface observation of Co-Cr dental bridges proves the higher roughness as well as the traces of the layered manufacturing of the cast patterns, left even after sandblasting. The fixed partial dentures, cast with 3D-printed cast patterns, possess higher accuracy of the shape, sizes, and adjustment compared to the dentures, produced by conventional lostwax casting. This is mainly due to the minimal manual work, because 3D-printed cast pat-

The microstructure of Co-Cr sample, cast with 3D-printed pattern (MJM), is a typical cast microstructure—inhomogeneous, consisting of large grains with dendrite morphology [50]. A large amount of lamellar and blocky carbides of different sizes are located mainly along the grain boundaries. The dendrites consist of γ phase, while the inter-dendritic regions consist of microeutectic (Co solid solution with intermetallic precipitations) and carbides of the (Cr,

range 327HV–343HV with uneven distribution due to the inhomogeneous microstructure. The dendritic grains with intermetallic phases, precipitated along the grain boundaries, were observed in the microstructure of the samples, cast with 3D-printed patterns from Co-Cr alloy remanium star CL by Kim et al. [51]. They characterized with yield strength 540 ± 20 MPa,

The microstructure and strength properties of Co-Cr dental alloys, cast with 3D-printed patterns, are typical for castings. Their dimensional and fitting accuracy is higher than that of the constructions, manufactured by conventional casting, while their surface roughness is more than three times higher. Therefore, the dimensional and fitting accuracy and the surface roughness of Co-Cr dental substructures are strongly influenced by the quality of the cast patterns fabricated via AT. The data about the dimensional accuracy of wax/polymeric dental constructions, produced by different additive manufacturing processes, are contrary. In some cases, the dimensions of the printed objects are smaller than the virtual model; in another case, they are larger. It depends not only on the type of the manufacturing process but also on the scale factor—the sizes of the printed object. As a consequence of the low dimensional accuracy, the fitting accuracy is lower. The strategy of printing should be chosen very carefully, because too much objects in the same built platform can cause lower dimensional and fitting accuracy. Due to the features of the 3D printing processes, the sizes of the detail can be different in the three directions *X*, *Y,* and *Z*; therefore, it should be paid attention to the position of the object, especially toward the print direction. This will influence the surface roughness, too. The parameters of the technological regimes also influence the dimensional accuracy. Decreasing the layer's thickness leads to smaller sizes and lower roughness, which in some cases defines the higher accuracy. The special requirements to the materials, used for 3D printing of cast patterns, exist—to have no burned out residue and null or minimal ther-

type. The hardness of the Co-Cr samples, cast with 3D-printed patterns, is in the

terns are produced directly from the virtual model.

mean percent elongation 10 ± 2 and Young's modulus 260 ± 20 GPa.

mal expansion—otherwise, the casting with low quality will be the result.

Mo)23C6

146 Biomaterials in Regenerative Medicine

The properties of the Co-Cr dental alloys depend on the microstructure, its morphology, and composition, which are defined by the manufacturing process and the technological regimes. The SLM is a complex thermophysical process, depending on a number of important parameters. During SLM, layers of metal powder are fused into a real object by a computer-directed laser. The process characterizes with high heating and cooling rates, leading to fine-grained microstructure of the solidified layer. As the heat is led away through the solid body, phase transformations run in the underneath layers heated above the transition temperatures [50]. Due to the high-temperature gradients during the SLM process, high residual stresses are generated in the details, which can cause subsequent deformations [17, 52–54]. These characteristics determine the specific microstructure and properties of the objects, produced by SLM as compared with that, manufactured by casting. The main technological parameters that are crucial for the production of high-quality construction are the laser power, scanning speed, laser beam diameter and distance between the traces, layer's thickness, and the working area [55–58]. In the development of any process for production of an object by SLM, it is necessary to evaluate the density, accuracy, surface roughness, hardness, and strength properties.

In SLM process, the volume of the molten metal depends on the volume energy density Ev , which is directly proportional to the power density Ns and inversely proportional to the scanning speed V. If the volume energy density Ev is insufficient, the incomplete melting of the deposited layer will occur. Therefore, the lower volume energy density Ev is the main reason for the porous structure [50]. This can be avoided by optimization of the technological parameters—increasing the laser power or decreasing the scanning speed. Varying with the input energy and scan spaces, Takaichi et al. [59] established that dense structure of SLM Co-29Cr-6Mo alloy can be obtained when the input energy of the laser scan increased more than 400 J/mm<sup>3</sup> and porous, in input energy less than 150 J/mm<sup>3</sup> . Vandenbroucke and Kruth [60] ensured 99.9% density of the SLM Co-Cr-Mo alloy, working with optimized technological regime. In the investigation of the possibility for manufacturing of three-part dental bridge from Co-Cr alloy by SLM, Averyanova et al. [29] stated that all samples are densed with porosity less than 1%.

direction, while in directions, perpendicular or inclined at 45° to the building direction, fine columnar structures with diameter about 0.5 μm, elongated along the building direction, were found. In lower magnification, gradual arch-shaped molten pool boundaries, typical for the SLM process, were observed. The tensile strength of the three groups of specimens is in the range 1170–1274 MPa, which is higher than that of the cast samples. Concerning to the fatigue strength, the results confirmed the anisotropy of the SLM alloy. The samples, parallel to the building direction, exhibit significantly longer fatigue life than the cast specimens, while the

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149

The research of Kim et al. [51] shows that the SLM Co-Cr alloy remanium star CL clearly exhibits the laser scan traces, as in higher magnification, the presence of fine grains with sizes about 35 μm can be recognized. The SLM samples showed the highest mean ultimate tensile strength, followed by milled/post-sintered, cast, and milled samples. The yield strength and mean percent elongation of SLM alloy were higher than the cast alloy (R0.2 = 580 ± 50 MPa and R0.2 = 540 ± 20 MPa, respectively), while Young's modulus was lower (200 ± 10 GPa for SLM and 260 ± 20 GPa for cast). A high yield strength and relatively low but sufficient modulus of elasticity of the SLM samples allow this technology to be used for manufacturing dental constructions, such as removable partial dentures, clasps, thin-veneered crowns, and wide-

Averyanova et al. [29] investigated three-part dental bridge of Co-Cr alloy, manufactured by SLM. The hardness of the SLM Co-Cr alloy is in the range of 400 ± 14 HV10 and 462 ± 22 HV0.05, while the average tensile strength is 1157 MPa, which is equal to that of the wrought alloy and about twice higher than the cast Co-Cr alloy (655 MPa). The microstructure of SLM dental bridges is nonequilibrium, consisting mainly of 98.7 ± 1.8% fcc Co-rich solid solution and 1.3 ± 0.5% of hcp ε phase. The investigation of Dikova et al. [50] established similar hardness of Co-Cr four-part dental bridges, produced by SLM (407–460 HV), which is higher than the cast alloy and has nearly even distribution along the depth of each crown. Their subsequent investigations [66] show that the tensile strength and the yield strength of the SLM Co-Cr alloy are higher than the cast alloy (R0.2 = 720 MPa and R0.2 = 410 MPa, respectively), while the elastic modulus is comparable (213 and 209 GPa, accordingly). The higher hardness and more homogeneous microstructure of SLM Co-Cr dental alloys determine their higher

The specific features of the SLM process and the parameters of the powder materials used determine the high surface roughness of the SLM Co-Cr dental alloys. The object position and orientation to the building direction as well as the choice and modeling of the supports are also of great importance [58]. The accuracy and the relation between the surface roughness and the sloping angle were researched in Ref. [60] using special benchmark models. It was proven that the surface roughness depends on the layer's thickness and sloping angle to the basis. The average arithmetic deviation of the surface roughness *Ra* varies between 6–18 μm for 20 μm layer's thickness and 13–33 μm for 50 μm layer's thickness. The higher values concern to the lower sloping angle to the basis of 8°, while the lower values for the larger angle of 70°. The research of Kajima et al. [65] confirmed that the surface roughness depends on the building direction. It is the lowest in the samples, perpendicular to the building direction,

fatigue life of the two groups is significantly shorter than that of the cast specimens.

span bridges.

wear and corrosion resistance [67].

The specific features of the SLM process, characterizing with high heating and cooling rates, define unique microstructure of Co-Cr dental alloys and mechanical properties higher than that of the cast alloys. Meacock and Vilar [61] reported that the microstructure of biomedical Co-Cr-Mo alloy, produced by laser powder micro-deposition, is homogeneously composed of fine cellular dendrites. The average hardness was 460 HV0.2, which is higher than the values obtained by the other fabrication process. In investigation of Co-Cr-Mo parts, produced by direct metal laser sintering, Barucca et al. [62] established that the microstructure consists of γ and ε phases. The ε phase is formed by athermal martensitic transformation, and it is distributed as network of thin lamellae inside the γ phase. The higher hardness (47 HRC) is attributed to the presence of the ε-lamellae grown on the {111}γ planes that restricts the dislocation movement in the γ phase. Lu et al. [63] investigated the microstructure, hardness, mechanical properties, electrochemical behaviour, and metal release of Co-Cr-W alloy fabricated by SLM in two different scanning strategies—line and island. They established the coexistence of the γ and ε phases in the microstructure and nearly the same hardness, 570 HV for lineformed alloy and 564 HV for island-formed alloy. Their research shows that the results of tensile strength (1158.22 ± 21 MPa for line scheme and 1115.56 ± 19 MPa for island scheme), hardness, density, electrochemical, and metal release tests are independent of the scanning strategy, and the yield strength of both samples meets the ISO 22764:2006 standard for dental restorations. The tensile strength of two Co-Cr dental alloys—cast remanium GM and SLM F75— was investigated in the work of Jevremovic et al. [64]. They established more than 1.5 times higher tensile strength of the SLM samples compared to the cast ones (1363–1472 MPa and 900 MPa, respectively). Vandenbroucke and Kruth [60] did complex investigation of SLM titanium and Co-Cr-Mo alloys. The mechanical tests proved that the SLM Co-Cr details fulfil the requirements for hardness, strength, and stiffness. Concerning to the corrosion—the SLM Co-Cr samples showed lower emission than the cast ones due to the more homogeneous and finer microstructure of the laser molten material.

In investigation of the influence of the object's position to the building direction, some researchers established anisotropy of the mechanical properties and especially of fatigue strength. The microstructure and mechanical properties of SLM Co-29Cr-6Mo alloy were studied out in the work of Takaichi et al. [59]. Unique microstructure was formed, consisting of the fine cellular dendrites in the elongated grains, parallel to the building direction. The cellular boundaries were enriched with Cr and Mo, and the γ phase was dominant in the SLM building. Due to the unique microstructure, the mechanical anisotropy was confirmed in the samples, but the yield strength, Ultimate Tensile Strength (UTS), and elongation were higher than that of the cast alloy. The research of Kajima et al. [65] shows that the microstructure of SLM specimens of Co-Cr-Mo alloys is quite different from those of the cast samples, which consists of coarse dendrites with visible precipitates in the inter-dendritic regions. The fine cellular dendrites with diameter about 0.5 μm were observed in the SLM samples parallel to the building direction, while in directions, perpendicular or inclined at 45° to the building direction, fine columnar structures with diameter about 0.5 μm, elongated along the building direction, were found. In lower magnification, gradual arch-shaped molten pool boundaries, typical for the SLM process, were observed. The tensile strength of the three groups of specimens is in the range 1170–1274 MPa, which is higher than that of the cast samples. Concerning to the fatigue strength, the results confirmed the anisotropy of the SLM alloy. The samples, parallel to the building direction, exhibit significantly longer fatigue life than the cast specimens, while the fatigue life of the two groups is significantly shorter than that of the cast specimens.

than 400 J/mm<sup>3</sup>

with porosity less than 1%.

148 Biomaterials in Regenerative Medicine

finer microstructure of the laser molten material.

and porous, in input energy less than 150 J/mm<sup>3</sup>

[60] ensured 99.9% density of the SLM Co-Cr-Mo alloy, working with optimized technological regime. In the investigation of the possibility for manufacturing of three-part dental bridge from Co-Cr alloy by SLM, Averyanova et al. [29] stated that all samples are densed

The specific features of the SLM process, characterizing with high heating and cooling rates, define unique microstructure of Co-Cr dental alloys and mechanical properties higher than that of the cast alloys. Meacock and Vilar [61] reported that the microstructure of biomedical Co-Cr-Mo alloy, produced by laser powder micro-deposition, is homogeneously composed of fine cellular dendrites. The average hardness was 460 HV0.2, which is higher than the values obtained by the other fabrication process. In investigation of Co-Cr-Mo parts, produced by direct metal laser sintering, Barucca et al. [62] established that the microstructure consists of γ and ε phases. The ε phase is formed by athermal martensitic transformation, and it is distributed as network of thin lamellae inside the γ phase. The higher hardness (47 HRC) is attributed to the presence of the ε-lamellae grown on the {111}γ planes that restricts the dislocation movement in the γ phase. Lu et al. [63] investigated the microstructure, hardness, mechanical properties, electrochemical behaviour, and metal release of Co-Cr-W alloy fabricated by SLM in two different scanning strategies—line and island. They established the coexistence of the γ and ε phases in the microstructure and nearly the same hardness, 570 HV for lineformed alloy and 564 HV for island-formed alloy. Their research shows that the results of tensile strength (1158.22 ± 21 MPa for line scheme and 1115.56 ± 19 MPa for island scheme), hardness, density, electrochemical, and metal release tests are independent of the scanning strategy, and the yield strength of both samples meets the ISO 22764:2006 standard for dental restorations. The tensile strength of two Co-Cr dental alloys—cast remanium GM and SLM F75— was investigated in the work of Jevremovic et al. [64]. They established more than 1.5 times higher tensile strength of the SLM samples compared to the cast ones (1363–1472 MPa and 900 MPa, respectively). Vandenbroucke and Kruth [60] did complex investigation of SLM titanium and Co-Cr-Mo alloys. The mechanical tests proved that the SLM Co-Cr details fulfil the requirements for hardness, strength, and stiffness. Concerning to the corrosion—the SLM Co-Cr samples showed lower emission than the cast ones due to the more homogeneous and

In investigation of the influence of the object's position to the building direction, some researchers established anisotropy of the mechanical properties and especially of fatigue strength. The microstructure and mechanical properties of SLM Co-29Cr-6Mo alloy were studied out in the work of Takaichi et al. [59]. Unique microstructure was formed, consisting of the fine cellular dendrites in the elongated grains, parallel to the building direction. The cellular boundaries were enriched with Cr and Mo, and the γ phase was dominant in the SLM building. Due to the unique microstructure, the mechanical anisotropy was confirmed in the samples, but the yield strength, Ultimate Tensile Strength (UTS), and elongation were higher than that of the cast alloy. The research of Kajima et al. [65] shows that the microstructure of SLM specimens of Co-Cr-Mo alloys is quite different from those of the cast samples, which consists of coarse dendrites with visible precipitates in the inter-dendritic regions. The fine cellular dendrites with diameter about 0.5 μm were observed in the SLM samples parallel to the building

. Vandenbroucke and Kruth

The research of Kim et al. [51] shows that the SLM Co-Cr alloy remanium star CL clearly exhibits the laser scan traces, as in higher magnification, the presence of fine grains with sizes about 35 μm can be recognized. The SLM samples showed the highest mean ultimate tensile strength, followed by milled/post-sintered, cast, and milled samples. The yield strength and mean percent elongation of SLM alloy were higher than the cast alloy (R0.2 = 580 ± 50 MPa and R0.2 = 540 ± 20 MPa, respectively), while Young's modulus was lower (200 ± 10 GPa for SLM and 260 ± 20 GPa for cast). A high yield strength and relatively low but sufficient modulus of elasticity of the SLM samples allow this technology to be used for manufacturing dental constructions, such as removable partial dentures, clasps, thin-veneered crowns, and widespan bridges.

Averyanova et al. [29] investigated three-part dental bridge of Co-Cr alloy, manufactured by SLM. The hardness of the SLM Co-Cr alloy is in the range of 400 ± 14 HV10 and 462 ± 22 HV0.05, while the average tensile strength is 1157 MPa, which is equal to that of the wrought alloy and about twice higher than the cast Co-Cr alloy (655 MPa). The microstructure of SLM dental bridges is nonequilibrium, consisting mainly of 98.7 ± 1.8% fcc Co-rich solid solution and 1.3 ± 0.5% of hcp ε phase. The investigation of Dikova et al. [50] established similar hardness of Co-Cr four-part dental bridges, produced by SLM (407–460 HV), which is higher than the cast alloy and has nearly even distribution along the depth of each crown. Their subsequent investigations [66] show that the tensile strength and the yield strength of the SLM Co-Cr alloy are higher than the cast alloy (R0.2 = 720 MPa and R0.2 = 410 MPa, respectively), while the elastic modulus is comparable (213 and 209 GPa, accordingly). The higher hardness and more homogeneous microstructure of SLM Co-Cr dental alloys determine their higher wear and corrosion resistance [67].

The specific features of the SLM process and the parameters of the powder materials used determine the high surface roughness of the SLM Co-Cr dental alloys. The object position and orientation to the building direction as well as the choice and modeling of the supports are also of great importance [58]. The accuracy and the relation between the surface roughness and the sloping angle were researched in Ref. [60] using special benchmark models. It was proven that the surface roughness depends on the layer's thickness and sloping angle to the basis. The average arithmetic deviation of the surface roughness *Ra* varies between 6–18 μm for 20 μm layer's thickness and 13–33 μm for 50 μm layer's thickness. The higher values concern to the lower sloping angle to the basis of 8°, while the lower values for the larger angle of 70°. The research of Kajima et al. [65] confirmed that the surface roughness depends on the building direction. It is the lowest in the samples, perpendicular to the building direction, *Ra* = 10.22 μm, followed by that of the samples, inclined at 45<sup>o</sup> with *Ra* = 13.67 μm, and the highest in the samples, parallel to the building direction, *Ra* = 18.17 μm. It was established in Ref. [49] that the roughness of the vestibular surface of the second premolar of four-part dental bridge, manufactured by SLM, is nearly four times higher than the roughness of the conventional cast bridge (*Ra* = 4.24 μm and *Ra* = 1.11 μm, respectively) and 25% higher compared to that cast with 3D-printed patterns (*Ra* = 4.24 μm and *Ra* = 3.39 μm, respectively). It is proposed that the considerably higher roughness and partially melted powder on the surface of the SLM samples could lead to the increase of the mechanical as well as the chemical components of the adhesion of the porcelain to the Co-Cr alloys, thus promoting higher adhesion strength of the porcelain coating.

The higher roughness of SLM Co-Cr alloys can cause lower dimensional accuracy and comparatively satisfactory adjustment accuracy in comparison to the objects, cast by conventional technology or with 3D-printed patterns. In SLM of Co-Cr-Mo alloy, Vandenbroucke and Kruth [60] established the process accuracy below 40 μm, which fulfil requirements of most medical and dental applications. According to Bibb et al. [30], SLM-manufactured Co-Cr frameworks for removable partial denture possess accuracy and quality of fit comparable to the existing traditional methods used in the dental laboratories. Averyanova et al. [29] confirmed that the geometrical accuracy of three-part dental bridge of Co-Cr alloy, produced by SLM, is comparable to that of the substructures, produced by conventional technology. The good repeatability of the SLM process in manufacturing Co-Cr four-part dental bridges was reported by Dzhendov et al. [49]. The maximal deviations of the dimensions of SLM bridges were the lowest compared to the conventional casting and casting with 3D-printed patterns. But the dimensions of the SLM dentures were lower than that of the base model with −0.07/−0.23 mm. The adjustment accuracy of the SLM bridges is comparable to that of the bridges, cast with 3D-printed patterns, and is

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There is no consensus regarding the clinically acceptable limits of marginal fit of dental restorations. Most researchers agree on an acceptable marginal discrepancy (distance from the abutment margin to the metal coping in a straight line) below the range of 100–120 μm [75], as values, greater than 120 μm, are considered not clinically acceptable [76–78]. The research of Kim et al. [75] stated that the marginal fit values of the Co-Cr alloy greatly depended on the fabrication methods and, occasionally, on the alloy systems. They found out that the marginal discrepancy of the SLM crowns (98.7 μm for 20-μm-thick layer and 128.8 μm for 30-μm-thick layer) is larger than the cast crowns (65.3–70.4 μm). In SLM samples, the marginal discrepancy increases with the increase of the layer's thickness. Kaleli and Sarac [76] compared the marginal adaptation after fabrication of the framework, porcelain application, and cementation of metalceramic restorations prepared by conventional lost-wax technique, milling, DMLS, and a direct powder-bed process. They observed the lowest marginal discrepancy values in the crowns, prepared by direct process powder-bed method, followed by the DMLS, milling, and casting. The research of Pompa et al. [77] concerns to the differences of marginal fit of laser-fused and conventional technologies for production of fixed dental prostheses. They established that the copings, manufactured by SLM, have better marginal adaptation within an acceptable range. But the cement gap characterized with irregular distribution was wider in the region of the shoulder than at the point of closure. The marginal discrepancy increased after porcelain application and cementation. In comparison of the marginal fit of metal laser sintered (MLS) Co-Cr alloy copings and conventional cast Ni-Cr alloy copings, Sundar et al. [79] concluded that the MLS copings had a better marginal fit and a decrease in micro-leakage compared to the copings, manufactured by conventional lost-wax technique. Huang et al. [80] compared the marginal and internal fit of SLM metal-ceramic crowns with lost-wax cast ones. They established that the SLM Co-Cr metal-ceramic crowns were better in marginal fit, not significantly different in axial fit and less accurate in occlusal fit than that of the cast samples. Concerning to the gap distribution, Tamac et al. [81] reported nearly the same results in comparing the clinical marginal and internal adaptation of metal-ceramic crowns, fabricated by CAD/CAM milling, DMLS, and

higher than that of the conventionally cast dentures.

Concerning to the standard ISO 9693-1:2012 [68], the minimum acceptable bond strength of metal-ceramic is 25 MPa. Kaleli and Sarac [69] compared porcelain bond strength of Co-Cr frameworks manufactured by conventional lost-wax technique, milling, direct metal laser sintering (DMLS), and direct process powder-bed method. There was no significant difference between the values of porcelain bond strength to the samples, produced by different methods. The mean bond strength was 38.08 MPa for cast samples, while that of the DMLS samples was 40.73 MPa. The type of failure of the cast samples was adhesive/mixed, while that of the DMLS samples was cohesive/mixed. Li et al. [70] confirmed that there are no significant differences between the bond strength of the cast, milled, and SLM Co-Cr alloys. The milled and SLM groups showed significantly more porcelain adherence than the cast group. Akova et al. [71] also revealed no statistically significant difference of the shear bond strength of porcelain to the cast and SLM Co-Cr dental alloys (72.9 and 67.0 MPa, respectively). The failure type of the porcelain of the cast samples was of the mixed adhesioncohesion type, while the failure of the SLM samples was of the mixed/adhesive type. The similar porcelain bond strength to the cast and SLM Co-Cr dental alloys was confirmed in the work of Wu et al. [72]—54.17 and 55.78 MPa accordingly. They established that the SLM alloy had an intermediate layer with elemental interpretation between the alloy and the porcelain, resulting in an improved bonding strength. Xiang et al. [73] established no significant difference for the mean bond strength of the SLM (44 MPa) and traditional cast (43 MPa) Co-Cr samples. A mixed fracture mode on the debonding interface of both the SLM and the cast groups was observed, but the SLM group showed significantly more porcelain adherence. Only Wang et al. [74] stated that there are statistically significant differences of the porcelain bond strength to the cast, CNC milled and SLM Co-Cr samples (37.7 ± 6.5 MPa, 43.3 ± 9.2 MPa and 46.8 ± 5.1 MPa, respectively), as the debonding surface of the all samples was of the cohesive failure mode. It should be noticed out that the surface roughness of the SLM samples decreases two–three times after sandblasting before porcelain firing (from *Ra* = 15 μm of as-received SLM Co-Cr samples to *Ra* = 8 μm of glass blasted and to *Ra*-5 μm in ultrasonic ceramic field) [60]; as in the most cases, the physical appearance and the surface quality are similar to the conventionally manufactured by investment casting [58]. Current investigations show that the SLM metal-ceramic system exhibits a bonding strength that exceeds the requirement of ISO 9691: 1:2012. It even shows a better behaviour in porcelain adherence comparable to the traditional cast methods.

The higher roughness of SLM Co-Cr alloys can cause lower dimensional accuracy and comparatively satisfactory adjustment accuracy in comparison to the objects, cast by conventional technology or with 3D-printed patterns. In SLM of Co-Cr-Mo alloy, Vandenbroucke and Kruth [60] established the process accuracy below 40 μm, which fulfil requirements of most medical and dental applications. According to Bibb et al. [30], SLM-manufactured Co-Cr frameworks for removable partial denture possess accuracy and quality of fit comparable to the existing traditional methods used in the dental laboratories. Averyanova et al. [29] confirmed that the geometrical accuracy of three-part dental bridge of Co-Cr alloy, produced by SLM, is comparable to that of the substructures, produced by conventional technology. The good repeatability of the SLM process in manufacturing Co-Cr four-part dental bridges was reported by Dzhendov et al. [49]. The maximal deviations of the dimensions of SLM bridges were the lowest compared to the conventional casting and casting with 3D-printed patterns. But the dimensions of the SLM dentures were lower than that of the base model with −0.07/−0.23 mm. The adjustment accuracy of the SLM bridges is comparable to that of the bridges, cast with 3D-printed patterns, and is higher than that of the conventionally cast dentures.

*Ra* = 10.22 μm, followed by that of the samples, inclined at 45<sup>o</sup>

in porcelain adherence comparable to the traditional cast methods.

strength of the porcelain coating.

150 Biomaterials in Regenerative Medicine

highest in the samples, parallel to the building direction, *Ra* = 18.17 μm. It was established in Ref. [49] that the roughness of the vestibular surface of the second premolar of four-part dental bridge, manufactured by SLM, is nearly four times higher than the roughness of the conventional cast bridge (*Ra* = 4.24 μm and *Ra* = 1.11 μm, respectively) and 25% higher compared to that cast with 3D-printed patterns (*Ra* = 4.24 μm and *Ra* = 3.39 μm, respectively). It is proposed that the considerably higher roughness and partially melted powder on the surface of the SLM samples could lead to the increase of the mechanical as well as the chemical components of the adhesion of the porcelain to the Co-Cr alloys, thus promoting higher adhesion

Concerning to the standard ISO 9693-1:2012 [68], the minimum acceptable bond strength of metal-ceramic is 25 MPa. Kaleli and Sarac [69] compared porcelain bond strength of Co-Cr frameworks manufactured by conventional lost-wax technique, milling, direct metal laser sintering (DMLS), and direct process powder-bed method. There was no significant difference between the values of porcelain bond strength to the samples, produced by different methods. The mean bond strength was 38.08 MPa for cast samples, while that of the DMLS samples was 40.73 MPa. The type of failure of the cast samples was adhesive/mixed, while that of the DMLS samples was cohesive/mixed. Li et al. [70] confirmed that there are no significant differences between the bond strength of the cast, milled, and SLM Co-Cr alloys. The milled and SLM groups showed significantly more porcelain adherence than the cast group. Akova et al. [71] also revealed no statistically significant difference of the shear bond strength of porcelain to the cast and SLM Co-Cr dental alloys (72.9 and 67.0 MPa, respectively). The failure type of the porcelain of the cast samples was of the mixed adhesioncohesion type, while the failure of the SLM samples was of the mixed/adhesive type. The similar porcelain bond strength to the cast and SLM Co-Cr dental alloys was confirmed in the work of Wu et al. [72]—54.17 and 55.78 MPa accordingly. They established that the SLM alloy had an intermediate layer with elemental interpretation between the alloy and the porcelain, resulting in an improved bonding strength. Xiang et al. [73] established no significant difference for the mean bond strength of the SLM (44 MPa) and traditional cast (43 MPa) Co-Cr samples. A mixed fracture mode on the debonding interface of both the SLM and the cast groups was observed, but the SLM group showed significantly more porcelain adherence. Only Wang et al. [74] stated that there are statistically significant differences of the porcelain bond strength to the cast, CNC milled and SLM Co-Cr samples (37.7 ± 6.5 MPa, 43.3 ± 9.2 MPa and 46.8 ± 5.1 MPa, respectively), as the debonding surface of the all samples was of the cohesive failure mode. It should be noticed out that the surface roughness of the SLM samples decreases two–three times after sandblasting before porcelain firing (from *Ra* = 15 μm of as-received SLM Co-Cr samples to *Ra* = 8 μm of glass blasted and to *Ra*-5 μm in ultrasonic ceramic field) [60]; as in the most cases, the physical appearance and the surface quality are similar to the conventionally manufactured by investment casting [58]. Current investigations show that the SLM metal-ceramic system exhibits a bonding strength that exceeds the requirement of ISO 9691: 1:2012. It even shows a better behaviour

with *Ra* = 13.67 μm, and the

There is no consensus regarding the clinically acceptable limits of marginal fit of dental restorations. Most researchers agree on an acceptable marginal discrepancy (distance from the abutment margin to the metal coping in a straight line) below the range of 100–120 μm [75], as values, greater than 120 μm, are considered not clinically acceptable [76–78]. The research of Kim et al. [75] stated that the marginal fit values of the Co-Cr alloy greatly depended on the fabrication methods and, occasionally, on the alloy systems. They found out that the marginal discrepancy of the SLM crowns (98.7 μm for 20-μm-thick layer and 128.8 μm for 30-μm-thick layer) is larger than the cast crowns (65.3–70.4 μm). In SLM samples, the marginal discrepancy increases with the increase of the layer's thickness. Kaleli and Sarac [76] compared the marginal adaptation after fabrication of the framework, porcelain application, and cementation of metalceramic restorations prepared by conventional lost-wax technique, milling, DMLS, and a direct powder-bed process. They observed the lowest marginal discrepancy values in the crowns, prepared by direct process powder-bed method, followed by the DMLS, milling, and casting. The research of Pompa et al. [77] concerns to the differences of marginal fit of laser-fused and conventional technologies for production of fixed dental prostheses. They established that the copings, manufactured by SLM, have better marginal adaptation within an acceptable range. But the cement gap characterized with irregular distribution was wider in the region of the shoulder than at the point of closure. The marginal discrepancy increased after porcelain application and cementation. In comparison of the marginal fit of metal laser sintered (MLS) Co-Cr alloy copings and conventional cast Ni-Cr alloy copings, Sundar et al. [79] concluded that the MLS copings had a better marginal fit and a decrease in micro-leakage compared to the copings, manufactured by conventional lost-wax technique. Huang et al. [80] compared the marginal and internal fit of SLM metal-ceramic crowns with lost-wax cast ones. They established that the SLM Co-Cr metal-ceramic crowns were better in marginal fit, not significantly different in axial fit and less accurate in occlusal fit than that of the cast samples. Concerning to the gap distribution, Tamac et al. [81] reported nearly the same results in comparing the clinical marginal and internal adaptation of metal-ceramic crowns, fabricated by CAD/CAM milling, DMLS, and traditional casting. They established that mean marginal gap values were 86.64 μm for milling, 96.23 μm for DMLS, and 75.92 μm for casting. The gap values in the axial wall region were the higher for the three groups of samples, followed by the gap values of axio-occlusal and occlusal surface regions. The cement film thickness at the occlusal region and axio-occlusal region was higher for the DMLS crowns. Consequently, the laser-assisted technologies for direct production of metal dental restorations, such as SLM, SLS, and DMLS, ensure improved or at least clinically acceptable fitting values compared with that of the conventional casting.

compared to the constructions, produced by traditional lost-wax casting or by SLM. The surface roughness is higher than that of the samples, cast by conventional technology, but lower

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153

The microstructure of SLM Co-Cr dental alloys is fine grained and more homogeneous comparing to that of the cast alloys, which defines higher hardness and mechanical properties, higher wear and corrosion resistance. The surface roughness of SLM Co-Cr dental alloys is higher than that of the alloys, cast conventionally or with 3D-printed patterns. The dimensional accuracy of SLM Co-Cr details is higher than the cast samples, while the fitting accu-

[1] Youssef S, Jabbari Al. Physico-mechanical properties and prosthodontic applications of Co-Cr dental alloys: a review of the literature. Journal of Advanced Prosthodontics.

[2] Eliasson A, Arnelund CF, Johansson A. A clinical evaluation of cobalt-chromium metalceramic fixed partial dentures and crowns: A three- to seven-year retrospective study.

[3] Kissov H. Stomatologichna keramika, chast I, Osnovni principi, materiali i instrumen-

[4] Podrrez-Radziszewska M, Haimann K, Dudzinski W, Morawska-Soltysik M . Characteristic of intermetallic phases in cast dental CoCrMo alloy. Archives of Foundry Engineering.

[5] Bellefontaine G. The Corrosion of CoCrMo Alloys for Biomedical Applications [thesis].

[6] Gupta P. The Co-Cr-Mo (Cobalt-Chromium-Molybdenum) system. Journal of Phase

[7] Kurosu Sh, Nomura N, Chiba A. Effect of sigma phase in Co-29Cr-6Mo alloy on corro-

sion behavior in saline solution. Materials Transaction. 2006;**47**(8):1961-1964

[8] van Noort R. The future of dental devices is digital. Dental Materials. 2012;**28**(3-12)

compared to the SLM objects.

**Author details**

Tsanka Dikova

**References**

2014;**6**(2):138-145

2010;**10**(3):51-59

racy is improved or clinically accepted.

Address all correspondence to: tsanka\_dikova@abv.bg

The Journal of Prosthetic Dentistry. 2007;**98**:6-16

tarium [in Bulgarian]. 1st ed. Sofia: Index; 1997. 432 p

Birmingham: University of Birmingham; 2010. 88 p

Equilibria and Diffusion. 2005;**26**(1):87-92

Medical University "Prof. d-r Paraskev Stoyanov", Varna, Bulgaria

The Co-Cr dental alloys, fabricated by SLM, characterize with homogeneous fine-grained microstructure, consisting mainly of γ and ε phases in different ratios. Their unique microstructure defines higher mechanical properties,—hardness, yield and tensile strength, fatigue strength as well as higher wear and corrosion resistance—compared to the cast alloys. As a consequence of the peculiarities of the SLM manufacturing and the position of the object toward the building direction, anisotropy in mechanical properties, especially in the fatigue life, can be observed. The work with optimized technological regimes can guarantee the constructions with density higher than 99% and comparatively low roughness. But as a whole, the surface roughness of the SLM Co-Cr alloys is higher than the alloys, cast conventionally or with 3D-printed patterns, due to the specific features of the manufacturing process and the use of metal powder as raw material. It was expected that the higher surface roughness could decrease the dimensional and fitting accuracy and promote the higher adhesion strength of the porcelain coating. But the current review shows that the dimensional accuracy of the SLM details is higher than the cast samples and the fitting accuracy is improved or clinically accepted, most probably due to the CAD/CAM nature of the manufacturing process, which additionally enables high repeatability. Concerning to the adhesion strength of the porcelain to the SLM dental alloys, more authors reported that it is comparable with the adhesion strength to the cast alloys, which may be due to the decreased roughness of the SLM samples after sandblasting before porcelain firing. The Co-Cr dental alloys, fabricated by SLM, comply with the standards and requirements concerning the dimensional and fitting accuracy as well as strength properties and can be successfully used in production of dental constructions.

## **4. Conclusion**

In dentistry, two different approaches can be applied for the production of metal frameworks using AT. According to the first one, the wax/polymeric cast patterns are fabricated by 3D printing, and the constructions are cast from dental alloy with as-printed patterns. Through the second one, the metal framework is a manufactured form of powder alloy directly from the 3D virtual model by SEBM or SLM.

The specific features of the manufacturing processes, the parameters of the technological regimes, and the properties of the materials influence on the accuracy and properties of Co-Cr dental constructions.

The microstructure and mechanical properties of Co-Cr dental alloys, cast using 3D-printed patterns, are typical for cast alloys. Their dimensional and adjustment accuracy is higher compared to the constructions, produced by traditional lost-wax casting or by SLM. The surface roughness is higher than that of the samples, cast by conventional technology, but lower compared to the SLM objects.

The microstructure of SLM Co-Cr dental alloys is fine grained and more homogeneous comparing to that of the cast alloys, which defines higher hardness and mechanical properties, higher wear and corrosion resistance. The surface roughness of SLM Co-Cr dental alloys is higher than that of the alloys, cast conventionally or with 3D-printed patterns. The dimensional accuracy of SLM Co-Cr details is higher than the cast samples, while the fitting accuracy is improved or clinically accepted.

## **Author details**

Tsanka Dikova

traditional casting. They established that mean marginal gap values were 86.64 μm for milling, 96.23 μm for DMLS, and 75.92 μm for casting. The gap values in the axial wall region were the higher for the three groups of samples, followed by the gap values of axio-occlusal and occlusal surface regions. The cement film thickness at the occlusal region and axio-occlusal region was higher for the DMLS crowns. Consequently, the laser-assisted technologies for direct production of metal dental restorations, such as SLM, SLS, and DMLS, ensure improved or at least

The Co-Cr dental alloys, fabricated by SLM, characterize with homogeneous fine-grained microstructure, consisting mainly of γ and ε phases in different ratios. Their unique microstructure defines higher mechanical properties,—hardness, yield and tensile strength, fatigue strength as well as higher wear and corrosion resistance—compared to the cast alloys. As a consequence of the peculiarities of the SLM manufacturing and the position of the object toward the building direction, anisotropy in mechanical properties, especially in the fatigue life, can be observed. The work with optimized technological regimes can guarantee the constructions with density higher than 99% and comparatively low roughness. But as a whole, the surface roughness of the SLM Co-Cr alloys is higher than the alloys, cast conventionally or with 3D-printed patterns, due to the specific features of the manufacturing process and the use of metal powder as raw material. It was expected that the higher surface roughness could decrease the dimensional and fitting accuracy and promote the higher adhesion strength of the porcelain coating. But the current review shows that the dimensional accuracy of the SLM details is higher than the cast samples and the fitting accuracy is improved or clinically accepted, most probably due to the CAD/CAM nature of the manufacturing process, which additionally enables high repeatability. Concerning to the adhesion strength of the porcelain to the SLM dental alloys, more authors reported that it is comparable with the adhesion strength to the cast alloys, which may be due to the decreased roughness of the SLM samples after sandblasting before porcelain firing. The Co-Cr dental alloys, fabricated by SLM, comply with the standards and requirements concerning the dimensional and fitting accuracy as well as strength properties and can be successfully used in production of dental constructions.

In dentistry, two different approaches can be applied for the production of metal frameworks using AT. According to the first one, the wax/polymeric cast patterns are fabricated by 3D printing, and the constructions are cast from dental alloy with as-printed patterns. Through the second one, the metal framework is a manufactured form of powder alloy directly from

The specific features of the manufacturing processes, the parameters of the technological regimes, and the properties of the materials influence on the accuracy and properties of Co-Cr

The microstructure and mechanical properties of Co-Cr dental alloys, cast using 3D-printed patterns, are typical for cast alloys. Their dimensional and adjustment accuracy is higher

clinically acceptable fitting values compared with that of the conventional casting.

**4. Conclusion**

152 Biomaterials in Regenerative Medicine

dental constructions.

the 3D virtual model by SEBM or SLM.

Address all correspondence to: tsanka\_dikova@abv.bg

Medical University "Prof. d-r Paraskev Stoyanov", Varna, Bulgaria

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**Chapter 6**

**Provisional chapter**

**Perspective of Additive Manufacturing Selective Laser**

This chapter seeks to compare the properties of samples manufactured by additive manufacturing (AM) by the selective laser melting (SLM) technology and compare with the precision casting (PC) processes using the Co-Cr-Mo (ASTM F75) alloy to manufacture of dental prosthesis. This AM process can be manufactured three-dimensional models by means of a laser beam that completely melts particles of powder deposited layer by layer. However, it is still relevant to know the properties of: performance, dimensional, mechanical and microstructural of this laser melting process and compare with a convencional process. The results of mechanical evaluation showed that the SLM technique provides superior mechanical properties compared to those obtained by the PC technique. It is possible to verify that the consolidation by SLM technique results in lower presence of porosity than PC technique. In addition, PC samples presented a gross dendritic microstructure of casting process. Microstructural analysis of SLM samples results in a characteristic morphology of layer manufacturing with ultrafine grains and a high chemical homogeneity. In this way, the development of the present study evidenced to improve the manufacture of customized components (copings)

**Perspective of Additive Manufacturing Selective Laser** 

DOI: 10.5772/intechopen.69720

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

**Keywords:** Co-Cr-Mo alloy, biomaterial, additive manufacturing, selective laser melting,

**Melting in Co-Cr-Mo Alloy in the Consolidation of**

**Melting in Co-Cr-Mo Alloy in the Consolidation of** 

**Dental Prosthesis**

**Dental Prosthesis**

Marcello Vertamatti Mergulhão, Carlos Eduardo Podestá and

Marcello Vertamatti Mergulhão, Carlos Eduardo Podestá and

Maurício David Martins das Neves

Maurício David Martins das Neves

http://dx.doi.org/10.5772/intechopen.69720

using the SLM technology.

precision casting

**Abstract**

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

**Provisional chapter**

#### **Perspective of Additive Manufacturing Selective Laser Melting in Co-Cr-Mo Alloy in the Consolidation of Dental Prosthesis Melting in Co-Cr-Mo Alloy in the Consolidation of Dental Prosthesis**

**Perspective of Additive Manufacturing Selective Laser** 

DOI: 10.5772/intechopen.69720

Marcello Vertamatti Mergulhão, Carlos Eduardo Podestá and Maurício David Martins das Neves Carlos Eduardo Podestá and Maurício David Martins das Neves Additional information is available at the end of the chapter

Marcello Vertamatti Mergulhão,

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69720

#### **Abstract**

This chapter seeks to compare the properties of samples manufactured by additive manufacturing (AM) by the selective laser melting (SLM) technology and compare with the precision casting (PC) processes using the Co-Cr-Mo (ASTM F75) alloy to manufacture of dental prosthesis. This AM process can be manufactured three-dimensional models by means of a laser beam that completely melts particles of powder deposited layer by layer. However, it is still relevant to know the properties of: performance, dimensional, mechanical and microstructural of this laser melting process and compare with a convencional process. The results of mechanical evaluation showed that the SLM technique provides superior mechanical properties compared to those obtained by the PC technique. It is possible to verify that the consolidation by SLM technique results in lower presence of porosity than PC technique. In addition, PC samples presented a gross dendritic microstructure of casting process. Microstructural analysis of SLM samples results in a characteristic morphology of layer manufacturing with ultrafine grains and a high chemical homogeneity. In this way, the development of the present study evidenced to improve the manufacture of customized components (copings) using the SLM technology.

**Keywords:** Co-Cr-Mo alloy, biomaterial, additive manufacturing, selective laser melting, precision casting

Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons

## **1. Introduction**

Metal powders of cobalt-chromium (Co-Cr) alloy are widely used in various sectors of the automotive, aeronautics, and aerospace industry, because of its high wear resistance and adequate corrosion resistance also being used in surface coating to increase performance components [1, 2]. In addition, the biocompatibility properties are suitable and are being used in the manufacture of medical and dental prosthetics [3–7]. The use of Co-Cr alloys is widely discussed to manufacture medical and dental implants or prostheses [7–9] presenting positive aspects in relation to biocompatibility analysis. The necessity for characterization and biological evaluations, physical-chemical, and mechanical are basic requirements for the development of new biomaterials applied in medical devices. In general, biomaterials need to present a final clinical characteristic (bio-functionality) and biocompatibility [10, 11].

Since 1930, Cobalt-Chromium-Molybdenum (Co-Cr-Mo) alloys processed by casting were used as dental alloys and later adapted for use in orthopedic implants [12, 13]. According to Jabbari et al., Co-Cr alloys are used almost exclusively in the manufacture of metal structures prostheses and recently is replacing Ni-Cr alloy or alternatively for the production of restorations in porcelain fused to metal (PFM), because Co-Cr alloy is Ni-free and does not have allergic responses or toxic effects related to Nickel [14].

In SLM technique, the raw material is in metallic powder form and the thermal energy required for the complete melting of the powder layer comes from a laser beam, usually is used a source of Ytterbium (Yb) fiber [29]. The maximum laser power on SLM machines is approximately 400 W and the laser focus may have a diameter of approximately 100 μm [30]. In turn, the laser beam is commanded by an interface that transmits it to the optical assembly, which selectively directs (X–Y plane) the laser beam, causing the powders to melt [28–33]. The metal powder is stored in a container, which may or may not be the distributor of powder (deposited by gravity), which in turn uniformizes the powder layer (between 50 and 100 μm) on an object consolidation platform. At each consolidated layer, the platform moves on the Z–axis, according to the next layer until the component is completely consolidated. The process of component consolidation occurs in a consolidation chamber (internal environment of the SLM equipment) that is under inert atmosphere protection (argon gas) [28–33]. The basic scheme of the components present in the consolidation process by SLM can be observed in **Figure 2**, and other machine parameters and working conditions are presented

**Figure 1.** Models of dental components manufactured by AM techniques in (a) resin model by stereolithography for

Perspective of Additive Manufacturing Selective Laser Melting in Co-Cr-Mo Alloy...

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163

posterior precision casting process and (b) copings manufactured by SLM technique [17].

Details of the main parameters of the SLM technique are shown schematically in **Figure 3**.

• Laser: power "*P*", beam diameter, pulse duration, pulse frequency;

• Temperature: consolidation layer, powder feeder, uniformity;

• Compounding chamber: composition of protective atmosphere.

• Scanning: speed "*v*", track distance, strategy;

The SLM technique has several process parameters and can be grouped in five families, being these related to laser, scanning, material (powder), temperature, and consolidation chamber

• Powder (material): material properties, particle size, distribution, bed density of powder,

in **Table 1**.

[28, 29, 31].

layer thickness;

The coefficient of thermal expansion (CTE) is a thermal property of the alloy, is of great interest in cases of applications in dental components, that requires ceramic coating, such as the dental crown. In this case, it is shown by Refs. [6, 15, 16] that Co-Cr alloys should have a CTE value in the range of 14.0–14.6 × 10−6 °C−1 at temperatures from 500 to 600°C to the correct ceramic firing process, as coating of the metal component. The CTE of ceramic materials for coating applications in metallic materials should be close, providing a good adjustment due to contraction and expansion during heating, thus avoiding the possibility of voids or cracks occurring during the firing process [15, 17].

Currently, the lost wax casting method is the most widely used, but has faced competition from other manufacturing processes [17]. Several authors [18–20] describe the development of AM technologies providing the creation of final customized implants. Techniques such as stereolithography were implemented to manufacture resin models for posterior manufacture of dental prostheses (crowns and bridges) by conventional process of lost wax castings. Mechanical, chemical, and microstructural properties are evaluated in comparison to new AM technologies, for example, the selective laser melting (SLM) in relation to conventional techniques as lost wax casting [21–23]. In this way, the preparation of medical and dental components provides customized final components with high mechanical properties, compared to conventional techniques (see **Figure 1**) [24].

Notably in health area, this technology is competitive over other traditional manufacturing processes by advances occurred in the area of processes using powder metallurgy techniques [25, 26]. Selective laser melting technology is one of the innovative technologies in additive manufacturing development in the middle of the 1980s after the creation of selective laser sintering (SLS) process. SLM is a process based on the 3D construction in which it is possible by laser beam to completely melt the metallic powder particles on a previous layer [27–31].

**Figure 1.** Models of dental components manufactured by AM techniques in (a) resin model by stereolithography for posterior precision casting process and (b) copings manufactured by SLM technique [17].

In SLM technique, the raw material is in metallic powder form and the thermal energy required for the complete melting of the powder layer comes from a laser beam, usually is used a source of Ytterbium (Yb) fiber [29]. The maximum laser power on SLM machines is approximately 400 W and the laser focus may have a diameter of approximately 100 μm [30]. In turn, the laser beam is commanded by an interface that transmits it to the optical assembly, which selectively directs (X–Y plane) the laser beam, causing the powders to melt [28–33]. The metal powder is stored in a container, which may or may not be the distributor of powder (deposited by gravity), which in turn uniformizes the powder layer (between 50 and 100 μm) on an object consolidation platform. At each consolidated layer, the platform moves on the Z–axis, according to the next layer until the component is completely consolidated. The process of component consolidation occurs in a consolidation chamber (internal environment of the SLM equipment) that is under inert atmosphere protection (argon gas) [28–33]. The basic scheme of the components present in the consolidation process by SLM can be observed in **Figure 2**, and other machine parameters and working conditions are presented in **Table 1**.

Details of the main parameters of the SLM technique are shown schematically in **Figure 3**.

The SLM technique has several process parameters and can be grouped in five families, being these related to laser, scanning, material (powder), temperature, and consolidation chamber [28, 29, 31].


**1. Introduction**

162 Biomaterials in Regenerative Medicine

Metal powders of cobalt-chromium (Co-Cr) alloy are widely used in various sectors of the automotive, aeronautics, and aerospace industry, because of its high wear resistance and adequate corrosion resistance also being used in surface coating to increase performance components [1, 2]. In addition, the biocompatibility properties are suitable and are being used in the manufacture of medical and dental prosthetics [3–7]. The use of Co-Cr alloys is widely discussed to manufacture medical and dental implants or prostheses [7–9] presenting positive aspects in relation to biocompatibility analysis. The necessity for characterization and biological evaluations, physical-chemical, and mechanical are basic requirements for the development of new biomaterials applied in medical devices. In general, biomaterials need to present

Since 1930, Cobalt-Chromium-Molybdenum (Co-Cr-Mo) alloys processed by casting were used as dental alloys and later adapted for use in orthopedic implants [12, 13]. According to Jabbari et al., Co-Cr alloys are used almost exclusively in the manufacture of metal structures prostheses and recently is replacing Ni-Cr alloy or alternatively for the production of restorations in porcelain fused to metal (PFM), because Co-Cr alloy is Ni-free and does not have

The coefficient of thermal expansion (CTE) is a thermal property of the alloy, is of great interest in cases of applications in dental components, that requires ceramic coating, such as the dental crown. In this case, it is shown by Refs. [6, 15, 16] that Co-Cr alloys should have a CTE value in the range of 14.0–14.6 × 10−6 °C−1 at temperatures from 500 to 600°C to the correct ceramic firing process, as coating of the metal component. The CTE of ceramic materials for coating applications in metallic materials should be close, providing a good adjustment due to contraction and expansion during heating, thus avoiding the possibility of voids or cracks

Currently, the lost wax casting method is the most widely used, but has faced competition from other manufacturing processes [17]. Several authors [18–20] describe the development of AM technologies providing the creation of final customized implants. Techniques such as stereolithography were implemented to manufacture resin models for posterior manufacture of dental prostheses (crowns and bridges) by conventional process of lost wax castings. Mechanical, chemical, and microstructural properties are evaluated in comparison to new AM technologies, for example, the selective laser melting (SLM) in relation to conventional techniques as lost wax casting [21–23]. In this way, the preparation of medical and dental components provides customized final components with high mechanical properties, com-

Notably in health area, this technology is competitive over other traditional manufacturing processes by advances occurred in the area of processes using powder metallurgy techniques [25, 26]. Selective laser melting technology is one of the innovative technologies in additive manufacturing development in the middle of the 1980s after the creation of selective laser sintering (SLS) process. SLM is a process based on the 3D construction in which it is possible by laser beam to completely melt the metallic powder particles on a previous layer [27–31].

a final clinical characteristic (bio-functionality) and biocompatibility [10, 11].

allergic responses or toxic effects related to Nickel [14].

occurring during the firing process [15, 17].

pared to conventional techniques (see **Figure 1**) [24].


**Figure 2.** Basic schematic of components present in a SLM machine [33].

The most common parameters to be adjusted in the SLM process to optimize the manufacture of components are: laser power (*P*), scan speed (*v*), track (hatch) distance, and layer thickness. According to Refs. [27, 34, 35], the volumetric energy density of the laser (*ψ*, given in J.mm−3) relates the main parameters of consolidation, in relation to the laser as shown in Eq. (1), being: laser power (*P*), track distance (*t*), scan speed (*v*), and layer thickness (*L*).

$$
\psi = \frac{P}{t \cdot v \cdot L} \tag{1}
$$

Several authors in Refs. [27, 34, 35] report that these parameters affect the volumetric energy density, determinant in the powder melting and that in turn influences the mechanical properties and roughness of the surface of the consolidated parts. The combination of these variables can generate excess (or insufficient) energy during the consolidation process, which can lead to the balling phenomenon at consolidated specimen, which corresponds to the dissimilar or noncontinuous scan tracks [30, 36, 37]. Additionally, the balling phenomenon can generate uniform deposition of next powder layer, can cause uncontrollable porosity and

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As observed, the parameters of the SLM process involve a certain complexity, in order to obtain the fabrication of components of complete density. In order to optimize the mechanical and physical-chemical properties of the final components manufactured the consolidation strategies are the subject of discussion and study [39]. The consolidation strategies refer to the consolidation parameters already presented, as well as to the direction and orientation of laser beam scanning, angle of rotation between the layers, and the number passes of laser beam (per layer), as seen in **Figure 4a**–**h** [40–42]. Also, the physical properties of the components may be associated with the manufacturing anisotropy of the samples, see **Figure 4i**

An interesting point to consider for AM processes is the feedstock (or raw material–metal powder). These new technologies demanded a characteristic powder size distribution, format and physical properties (flowability and packing) [36]. In this case, to produce spherical metal powder the most common process is gas atomization [29, 30]. However, it is remarkable that the use of gas-atomized powders in the SLM process by the better physical properties is compared to water-atomized powders. The characteristic format of powders (gas and water

In addition, the capability to reusability of feedstock material in AM processes is a significant issue to promote economic and environmental manufacture processes [28, 46]. However, the effects and influence of the powder reuse on manufactured parts are the subject of much discussion [47–50]. The conclusions of Tang et al. [48] study, about reuse of Ti-6Al-4V powders of electron beam melting (EBM) process, appoint the increased oxygen content and particles became less spherical. Although the reuse powder improved the flowability (by little presence of satellite particles), increased the yield strength and the ultimate tensile on the AM process

delamination by the absence of inter-fusion between layers [37, 38].

**Figure 3.** Schematic of parameters involved in SLM consolidation [31].

atomized) is possible to observe in **Figure 5**.

[43–45].

of Ti-6Al-4V [48].


**Table 1.** Typical technical parameters of the SLM®280HL machine.

Perspective of Additive Manufacturing Selective Laser Melting in Co-Cr-Mo Alloy... http://dx.doi.org/10.5772/intechopen.69720 165

**Figure 3.** Schematic of parameters involved in SLM consolidation [31].

The most common parameters to be adjusted in the SLM process to optimize the manufacture of components are: laser power (*P*), scan speed (*v*), track (hatch) distance, and layer thickness. According to Refs. [27, 34, 35], the volumetric energy density of the laser (*ψ*, given in J.mm−3) relates the main parameters of consolidation, in relation to the laser as shown in Eq. (1), being:

*<sup>t</sup>* <sup>∙</sup> *<sup>v</sup>* <sup>∙</sup> *<sup>L</sup>* (1)

, 2.5–3.0 l/min

, 1700 l @ 100 l/min

laser power (*P*), track distance (*t*), scan speed (*v*), and layer thickness (*L*).

Laser power 400 W Yb-Fiber-laser

Build speed 20 ccm/h Pract. layer thickness 20–75 μm min Scan line/wall thickness 150 μm Operational beam focus 80–120 μm Scan speed 15 m/s Inert gas consumption in operation Ar/N2

Inert gas consumption venting Ar/N2

**Table 1.** Typical technical parameters of the SLM®280HL machine.

Compressed air requirement 18 l/min @ 1.5 bar

*ψ* = \_\_\_\_\_\_ *<sup>P</sup>*

**System parameters**

164 Biomaterials in Regenerative Medicine

**Figure 2.** Basic schematic of components present in a SLM machine [33].

Several authors in Refs. [27, 34, 35] report that these parameters affect the volumetric energy density, determinant in the powder melting and that in turn influences the mechanical properties and roughness of the surface of the consolidated parts. The combination of these variables can generate excess (or insufficient) energy during the consolidation process, which can lead to the balling phenomenon at consolidated specimen, which corresponds to the dissimilar or noncontinuous scan tracks [30, 36, 37]. Additionally, the balling phenomenon can generate uniform deposition of next powder layer, can cause uncontrollable porosity and delamination by the absence of inter-fusion between layers [37, 38].

As observed, the parameters of the SLM process involve a certain complexity, in order to obtain the fabrication of components of complete density. In order to optimize the mechanical and physical-chemical properties of the final components manufactured the consolidation strategies are the subject of discussion and study [39]. The consolidation strategies refer to the consolidation parameters already presented, as well as to the direction and orientation of laser beam scanning, angle of rotation between the layers, and the number passes of laser beam (per layer), as seen in **Figure 4a**–**h** [40–42]. Also, the physical properties of the components may be associated with the manufacturing anisotropy of the samples, see **Figure 4i** [43–45].

An interesting point to consider for AM processes is the feedstock (or raw material–metal powder). These new technologies demanded a characteristic powder size distribution, format and physical properties (flowability and packing) [36]. In this case, to produce spherical metal powder the most common process is gas atomization [29, 30]. However, it is remarkable that the use of gas-atomized powders in the SLM process by the better physical properties is compared to water-atomized powders. The characteristic format of powders (gas and water atomized) is possible to observe in **Figure 5**.

In addition, the capability to reusability of feedstock material in AM processes is a significant issue to promote economic and environmental manufacture processes [28, 46]. However, the effects and influence of the powder reuse on manufactured parts are the subject of much discussion [47–50]. The conclusions of Tang et al. [48] study, about reuse of Ti-6Al-4V powders of electron beam melting (EBM) process, appoint the increased oxygen content and particles became less spherical. Although the reuse powder improved the flowability (by little presence of satellite particles), increased the yield strength and the ultimate tensile on the AM process of Ti-6Al-4V [48].

**2. Experimental procedure**

ANVISA (Brazilian agency) for use in health care segment.

Micromeritics equipment (Model Accu PYC 1330 Pycnometer).

ence–empty during all tests) were composed of alumina (Al2

Co-Cr-Mo alloy gas atomized (H.C Starck®, Lübeck, Germany) was provided by the HighBond® (Indaiatuba, Brazil) in the particle size (granulometric range) of 15–45 μm. The confirmation of the chemical composition was performed by energy dispersive X-ray (Shimadzu EDX-720 equipment and by LECO). This study was based on alloy/powder with certification of

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Several physical properties of gas-atomized powder were obtained such as flow time (ASTM B212 [51]), apparent density (ASTM B213 [52]), and tap density (ASTM B527 [53]). The particle size distribution was performed using a particle analyzer by laser scattering (Cilas–Model 1064). The particle format and microstructural characterization of powders were performed via optical and scanning electron microscopy (OM–Olympus BX51M and SEM-EDS Philips

To evaluate the internal porosity of powders sample were measured by the pycnometer density in comparison to theoretical density. The density by Helium pycnometry considered only the internal porosity (excluding the open porosity) and was performed using the

Differential scanning calorimetry (DSC) analysis was performed using a sample of gas-atomized Co-Cr-Mo powder. Three runs of heating curves at rates of 10, 20, and 30°C/min were performed and under static atmosphere constituted in argon (99.999%) for minimizing the oxidation of the samples. In all experiments, both the crucible (sample holder and the refer-

mately 100 μL. The equipment used was Setsys 16/18, from Setaram with a thermocouple rod

Precision casting and selective laser melting techniques performed the manufacture of gasatomized Co-Cr-Mo powders. The tensile and three-point bending specimens were manufactured in standard dimensions according to ISO 22674-06 [54] and ASTM B528-12 [55]. **Figure 6**

The precision casting (PC) samples were performed according to ASTM F75-12 [13] by HighBond® (Indaiatuba, Brazil). The PC fabrication process method satisfied the following steps: machining of wax disks in the standard dimensions of tensile and flexural test specimens, assembly and shell building, dewaxing and pouring the Co-Cr-Mo alloy was by an

The consolidation of SLM samples was carried out by SLM Solutions™ using a selective laser melting machine SLM®280HL with a single Ytterbium laser beam (maximum power 400 W). The building consolidation of specimens was parallel to laser beam and performed using

parameters such as: layer thickness of 30 μm and diameter of laser beam of 76 μm.

**2.2. Manufacturing specimens using precision casting and selective laser melting**

O3

) with a volume of approxi-

**2.1. Powder characterization**

XL30).

of Pt/Pt Rh 10%.

shows specimens manufactured.

induction furnace at a temperature of 1489°C.

**Figure 4.** Representation of laser scanning strategies for sample consolidation via SLM, scanning in: (a) island, (b) line, (c) line at 45°, (d) line and rotation line at 45°, (e) line and rotation line at 90°, (f) line rotate at 67°, (g) internal spiral, (h) external spiral [42], and (i) building orientation of specimens [45].

Considering this important field in expansion, this chapter is part of this scenario with a focus on the dental sector, more specifically on the evaluation of mechanical properties and microstructural analysis of Co-Cr-Mo alloy to manufacture dental prostheses (copings). The aim of this chapter is to evaluate the mechanical properties and microstructures of standardized specimens made by powder metallurgy techniques using SLM from powdered gas powder of the Co-Cr-Mo alloy. The results obtained by SLM will be compared with the results of samples manufactured by precision casting.

**Figure 5.** Characteristic format of powders (316L alloy) produced by water atomization (a) and gas atomization (b). Magnitude ×500.

## **2. Experimental procedure**

#### **2.1. Powder characterization**

Considering this important field in expansion, this chapter is part of this scenario with a focus on the dental sector, more specifically on the evaluation of mechanical properties and microstructural analysis of Co-Cr-Mo alloy to manufacture dental prostheses (copings). The aim of this chapter is to evaluate the mechanical properties and microstructures of standardized specimens made by powder metallurgy techniques using SLM from powdered gas powder of the Co-Cr-Mo alloy. The results obtained by SLM will be compared with the results of

**Figure 4.** Representation of laser scanning strategies for sample consolidation via SLM, scanning in: (a) island, (b) line, (c) line at 45°, (d) line and rotation line at 45°, (e) line and rotation line at 90°, (f) line rotate at 67°, (g) internal spiral, (h)

**Figure 5.** Characteristic format of powders (316L alloy) produced by water atomization (a) and gas atomization (b).

samples manufactured by precision casting.

external spiral [42], and (i) building orientation of specimens [45].

166 Biomaterials in Regenerative Medicine

Magnitude ×500.

Co-Cr-Mo alloy gas atomized (H.C Starck®, Lübeck, Germany) was provided by the HighBond® (Indaiatuba, Brazil) in the particle size (granulometric range) of 15–45 μm. The confirmation of the chemical composition was performed by energy dispersive X-ray (Shimadzu EDX-720 equipment and by LECO). This study was based on alloy/powder with certification of ANVISA (Brazilian agency) for use in health care segment.

Several physical properties of gas-atomized powder were obtained such as flow time (ASTM B212 [51]), apparent density (ASTM B213 [52]), and tap density (ASTM B527 [53]). The particle size distribution was performed using a particle analyzer by laser scattering (Cilas–Model 1064). The particle format and microstructural characterization of powders were performed via optical and scanning electron microscopy (OM–Olympus BX51M and SEM-EDS Philips XL30).

To evaluate the internal porosity of powders sample were measured by the pycnometer density in comparison to theoretical density. The density by Helium pycnometry considered only the internal porosity (excluding the open porosity) and was performed using the Micromeritics equipment (Model Accu PYC 1330 Pycnometer).

Differential scanning calorimetry (DSC) analysis was performed using a sample of gas-atomized Co-Cr-Mo powder. Three runs of heating curves at rates of 10, 20, and 30°C/min were performed and under static atmosphere constituted in argon (99.999%) for minimizing the oxidation of the samples. In all experiments, both the crucible (sample holder and the reference–empty during all tests) were composed of alumina (Al2 O3 ) with a volume of approximately 100 μL. The equipment used was Setsys 16/18, from Setaram with a thermocouple rod of Pt/Pt Rh 10%.

#### **2.2. Manufacturing specimens using precision casting and selective laser melting**

Precision casting and selective laser melting techniques performed the manufacture of gasatomized Co-Cr-Mo powders. The tensile and three-point bending specimens were manufactured in standard dimensions according to ISO 22674-06 [54] and ASTM B528-12 [55]. **Figure 6** shows specimens manufactured.

The precision casting (PC) samples were performed according to ASTM F75-12 [13] by HighBond® (Indaiatuba, Brazil). The PC fabrication process method satisfied the following steps: machining of wax disks in the standard dimensions of tensile and flexural test specimens, assembly and shell building, dewaxing and pouring the Co-Cr-Mo alloy was by an induction furnace at a temperature of 1489°C.

The consolidation of SLM samples was carried out by SLM Solutions™ using a selective laser melting machine SLM®280HL with a single Ytterbium laser beam (maximum power 400 W). The building consolidation of specimens was parallel to laser beam and performed using parameters such as: layer thickness of 30 μm and diameter of laser beam of 76 μm.

To evaluate the internal porosity of cast, and SLM samples were measured by the pycnometry density in comparison to the theoretical density. The density by Helium pycnometry, considered only the internal porosity (excluding the open porosity) was measured using the

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The thermomechanical analysis (TMA) was performed on samples consolidated by SLM and PC of the Co-Cr-Mo alloy. The purpose of the technique was to obtain the coefficient of thermal expansion (CTE). In addition, the SLM samples were analyzed in the parallel and transversal building direction (SLM 1—parallel direction and SLM 2—transversal direction). The routine of the TMA remained the heating rate was from 10°C/min until the temperature of 1300°C. The equipment used was a Setaram–Setsys 16/18, using a thermocouple rod of Pt/Pt

Mechanical characterization of consolidated samples by PC and SLM techniques was held in five samples of each test (tensile and three-point bending), respectively, according to ISO 22674-06 [54] and ASTM B528 [55]. The three-point bending test determined the transversal rupture strength (TRS) of specimens. The TRS relates to the applied load (P) and the distance between the supports (L), over the cross area of the sample (thickness "*t*" and width "*w*"), as show in Eq. (2). Mechanical tests were performed using a universal testing machine (Instron

The microstructural characterization of consolidated Co-Cr-Mo and the fracture analysis were evaluated after tensile test. Metallography preparation consisted of mechanical grinding in SiC paper #1200 and final chemical polishing with OP-S 0.02 μm with addition of 10% HCl.

ture). The microstructural characterization was performed in both building directions using an optical microscope (OM) Olympus–BX51M and scanning electron microscope (SEM) with

The confirmation of chemical composition was performed in Co-Cr-Mo powder alloy, as also in the samples manufactured (PC and SLM). The chemical composition is presented in

<sup>2</sup> <sup>∙</sup> *<sup>t</sup>*<sup>2</sup> <sup>∙</sup> *<sup>w</sup>* (2)

(1–2 min at room tempera-

O2

Rh 10% under a static atmosphere (argon—99.999%) to exclude the sample oxidation.

Micromeritics equipment–Model Accu PYC 1330 Pycnometer.

3366) under a crosshead speed of 0.2 mm/min at room temperature.

The specimens were etching in solution: 100 ml HCl and 2 ml H2

energy dispersive X-ray (EDS) Philips XL30 and JEOL–JSM6701F.

**Table 2** comparing with the standard ASTM F75-12 [13].

*TRS* = \_\_\_\_\_\_\_ <sup>3</sup> <sup>∙</sup> *<sup>P</sup>* <sup>∙</sup> *<sup>L</sup>*

**2.4. Mechanical characterization**

**2.5. Microstructural evaluation**

**3. Results and discussion**

**3.1. Powder characterization**

**Figure 6.** Specimens manufactured of Co-Cr-Mo alloy: (a) specimens made by FP technique and (b) specimens made by SLM technique.

#### **2.3. Characterization of samples manufactured**

To evaluate the susceptibility to cell growth in the Co-Cr-Mo alloy after the consolidation processes (PC and SLM) was performed by the cytotoxicity analysis, according to ISO 10993-5 [56]. The determination of the cytotoxicity was obtained by the quantitative evaluation method, which is carried out by the measurement of cell death, cell proliferation or formation of cellular colonies.

To evaluate the internal porosity of cast, and SLM samples were measured by the pycnometry density in comparison to the theoretical density. The density by Helium pycnometry, considered only the internal porosity (excluding the open porosity) was measured using the Micromeritics equipment–Model Accu PYC 1330 Pycnometer.

The thermomechanical analysis (TMA) was performed on samples consolidated by SLM and PC of the Co-Cr-Mo alloy. The purpose of the technique was to obtain the coefficient of thermal expansion (CTE). In addition, the SLM samples were analyzed in the parallel and transversal building direction (SLM 1—parallel direction and SLM 2—transversal direction). The routine of the TMA remained the heating rate was from 10°C/min until the temperature of 1300°C. The equipment used was a Setaram–Setsys 16/18, using a thermocouple rod of Pt/Pt Rh 10% under a static atmosphere (argon—99.999%) to exclude the sample oxidation.

#### **2.4. Mechanical characterization**

Mechanical characterization of consolidated samples by PC and SLM techniques was held in five samples of each test (tensile and three-point bending), respectively, according to ISO 22674-06 [54] and ASTM B528 [55]. The three-point bending test determined the transversal rupture strength (TRS) of specimens. The TRS relates to the applied load (P) and the distance between the supports (L), over the cross area of the sample (thickness "*t*" and width "*w*"), as show in Eq. (2). Mechanical tests were performed using a universal testing machine (Instron 3366) under a crosshead speed of 0.2 mm/min at room temperature.

$$\text{TRS} = \frac{\text{3} \cdot \text{P} \cdot \text{L}}{2 \cdot \text{t} \text{2} \cdot w} \tag{2}$$

#### **2.5. Microstructural evaluation**

The microstructural characterization of consolidated Co-Cr-Mo and the fracture analysis were evaluated after tensile test. Metallography preparation consisted of mechanical grinding in SiC paper #1200 and final chemical polishing with OP-S 0.02 μm with addition of 10% HCl. The specimens were etching in solution: 100 ml HCl and 2 ml H2 O2 (1–2 min at room temperature). The microstructural characterization was performed in both building directions using an optical microscope (OM) Olympus–BX51M and scanning electron microscope (SEM) with energy dispersive X-ray (EDS) Philips XL30 and JEOL–JSM6701F.

#### **3. Results and discussion**

#### **3.1. Powder characterization**

**2.3. Characterization of samples manufactured**

of cellular colonies.

SLM technique.

168 Biomaterials in Regenerative Medicine

To evaluate the susceptibility to cell growth in the Co-Cr-Mo alloy after the consolidation processes (PC and SLM) was performed by the cytotoxicity analysis, according to ISO 10993-5 [56]. The determination of the cytotoxicity was obtained by the quantitative evaluation method, which is carried out by the measurement of cell death, cell proliferation or formation

**Figure 6.** Specimens manufactured of Co-Cr-Mo alloy: (a) specimens made by FP technique and (b) specimens made by

The confirmation of chemical composition was performed in Co-Cr-Mo powder alloy, as also in the samples manufactured (PC and SLM). The chemical composition is presented in **Table 2** comparing with the standard ASTM F75-12 [13].


**Table 2.** Chemical composition (weight %) of Co-Cr-Mo samples (powder, PC and SLM) in accordance with standard ASTM F75.

The powders to manufacture samples via SLM technique have a mean diameter less than 50 μm to improve the physical properties such as flow time, apparent density, and tap den-

According to Haan et al. [59], Co-Cr-Mo powders with diameter D90 equals to 39 μm, the flowability was 18.60 s/50 g. The results were similar to those obtained for the present study,

The result of tap density tends to be higher than the result of the apparent density, because of the particle's accommodation there is a decrease in the amount of voids between the particles [2]. Also, the smaller the apparent density, the greater the percentage of increase the tap

It is possible to verify the presence of closed porosity that is not considered as a measure of the vol-

To investigate and confirm the thermal events present in the Co-Cr-Mo alloy, the heating curves (different rates: 10, 20, and 30°C/min) of the DSC analysis obtained using the Co-Cr-Mo powder are shown in **Figure 8**. The presence of three events occurring in the heating curves of DSC is observed, being the first exothermic, the second endothermic, and the final event corresponding to the fusion of the Co-Cr-Mo alloy. It is possible to verify the temperature

In relation to the first event (exothermic), occurring around 582.81°C, it is related to the phase transformation of the alloy (precisely from Co), from the cubic face (αCo) phase to the compact hexagonal phase (εCo). In a similar analysis, Santos [16] obtained a slight peak at 600°C in the thermal analysis (DTA), however the author does not approach the occurrence. Facchini [60] describes this occurrence, the event occurs at approximately 650°C, but is described by an endothermic peak, diverging from the present analysis, in which the curve of 20°C/min occurs

The second event (endothermic) occurs around 944.52°C, may be related to allotropic transformation of element Co, by the transition of the phase of compact hexagonal structure (εCo) to the phase of cubic structure of face centered (αCo). This transformation can be confirmed

). The presence

) is approximately 1.3%.

ume of Helium and consequently reduces the value of pycnometry density (8.30 g/cm3

of internal porosity calculated in relation to theoretical density (8.38 g/cm3

variation of the events between the different temperature rates.

at approximately 600°C and which describes an exothermic peak.

sity [33]. The results of physical powder properties are summarized in **Table 3**.

Granulometric distribution (μm) D10 D50 D90 D mean

20.88 31.11 46.10 32.36

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such as 15.86 s/50 g for D90 equals to 46.10 μm.

**Physical properties Co-Cr-Mo powder**

) 4.51 ± 0.01

) 8.30 ± 0.001

**Table 3.** Results of physical properties for Co-Cr-Mo powder in the SLM range.

) 5.26 ± 0.05

) 8.38

Flow rate (s/50g) 15.85 ± 0.11

Apparent density (g/cm3

Theoretical density (g/cm3

Helium pycnometry (g/cm3

Tap density (g/cm3

density.

The characteristic format of the powder process fabrication by gas atomization is observed in **Figure 7**. The analysis in SEM shows that the powders are spherical and presented satellites (appointed by arrows—**Figure 7a**,**b**). The satellites can be formed in the surface particles during the cooling process of the spherical powder particles during gas atomization. It is noteworthy that the shape of the particle influences on packing properties, flow hate, and compressibility, as well as reports on the powder metallurgy process [2, 57, 58]. The crosssectioned powder (**Figure 7d,e**.) shows the dendritic morphology with the primarily arms and ramifications, characterizing the rapid solidification of gas atomization process.

**Figure 7.** SEM images of Co-Cr-Mo powder: (a) magnitude ×500, (b) magnitude ×1000, (c) magnitude ×2000, (d) and (e) cross-section powder after chemical etch (etch solution: HCl, H2 SO4 , and HNO3 for 60–240 s at 45°C, respectively, magnitude ×1000 and ×2000).


**Table 3.** Results of physical properties for Co-Cr-Mo powder in the SLM range.

The characteristic format of the powder process fabrication by gas atomization is observed in **Figure 7**. The analysis in SEM shows that the powders are spherical and presented satellites (appointed by arrows—**Figure 7a**,**b**). The satellites can be formed in the surface particles during the cooling process of the spherical powder particles during gas atomization. It is noteworthy that the shape of the particle influences on packing properties, flow hate, and compressibility, as well as reports on the powder metallurgy process [2, 57, 58]. The crosssectioned powder (**Figure 7d,e**.) shows the dendritic morphology with the primarily arms

**Table 2.** Chemical composition (weight %) of Co-Cr-Mo samples (powder, PC and SLM) in accordance with standard

**Figure 7.** SEM images of Co-Cr-Mo powder: (a) magnitude ×500, (b) magnitude ×1000, (c) magnitude ×2000, (d) and

SO4

, and HNO3

for 60–240 s at 45°C, respectively,

(e) cross-section powder after chemical etch (etch solution: HCl, H2

magnitude ×1000 and ×2000).

ASTM F75.

170 Biomaterials in Regenerative Medicine

and ramifications, characterizing the rapid solidification of gas atomization process.

**Elements (%) Powder PC SLM ASTM F75** Co 63.93 ± 0.16 66.38 ± 0.15 65.38 ± 0.32 Balanço

O2 0.0940 ± 0.0015 0.0187 ± 0.0016 0.0240 ± 0.0010 –

Cr 28.83 ± 0.19 26.76 ± 0.21 27.68 ± 0.13 27.00 − 30.00 ± 0.30 Mo 7.07 ± 0.31 6.68 ± 0.03 6.61 ± 0.16 5.00 − 7.00 ± 0.15 Fe 0.17 ± 0.01 0.18 ± 0.08 0.33 ± 0.06 0.75 ± 0.03 C 0.03 ± 0.01 0.02 ± 0.01 0.03 ± 0.01 0.350 ± 0.020 S 0.01 ± 0.01 0.01 ± 0.01 0.01 ± 0.01 0.010 ± 0.003 N2 0.0820 ± 0.0011 0.0416 ± 0.0015 0.1330 ± 0.0015 0.250 ± 0.020

> The powders to manufacture samples via SLM technique have a mean diameter less than 50 μm to improve the physical properties such as flow time, apparent density, and tap density [33]. The results of physical powder properties are summarized in **Table 3**.

> According to Haan et al. [59], Co-Cr-Mo powders with diameter D90 equals to 39 μm, the flowability was 18.60 s/50 g. The results were similar to those obtained for the present study, such as 15.86 s/50 g for D90 equals to 46.10 μm.

> The result of tap density tends to be higher than the result of the apparent density, because of the particle's accommodation there is a decrease in the amount of voids between the particles [2]. Also, the smaller the apparent density, the greater the percentage of increase the tap density.

> It is possible to verify the presence of closed porosity that is not considered as a measure of the volume of Helium and consequently reduces the value of pycnometry density (8.30 g/cm3 ). The presence of internal porosity calculated in relation to theoretical density (8.38 g/cm3 ) is approximately 1.3%.

> To investigate and confirm the thermal events present in the Co-Cr-Mo alloy, the heating curves (different rates: 10, 20, and 30°C/min) of the DSC analysis obtained using the Co-Cr-Mo powder are shown in **Figure 8**. The presence of three events occurring in the heating curves of DSC is observed, being the first exothermic, the second endothermic, and the final event corresponding to the fusion of the Co-Cr-Mo alloy. It is possible to verify the temperature variation of the events between the different temperature rates.

> In relation to the first event (exothermic), occurring around 582.81°C, it is related to the phase transformation of the alloy (precisely from Co), from the cubic face (αCo) phase to the compact hexagonal phase (εCo). In a similar analysis, Santos [16] obtained a slight peak at 600°C in the thermal analysis (DTA), however the author does not approach the occurrence. Facchini [60] describes this occurrence, the event occurs at approximately 650°C, but is described by an endothermic peak, diverging from the present analysis, in which the curve of 20°C/min occurs at approximately 600°C and which describes an exothermic peak.

> The second event (endothermic) occurs around 944.52°C, may be related to allotropic transformation of element Co, by the transition of the phase of compact hexagonal structure (εCo) to the phase of cubic structure of face centered (αCo). This transformation can be confirmed

The results showed the expected results for the Co-Cr-Mo alloy, because as there is a need for specific mechanical properties to be reached, the use of Co-Cr alloys for the manufacture of medical and dental implants or prostheses did not show any toxicity with the medium

**Figure 9.** Result of viability of cell growth (cytotoxicity analysis) for Co-Cr-Mo specimens manufactured by PC and SLM

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The porosity of the samples was determined and a comparison was made with average densities: theoretical, volumetric, and by helium pycnometry. The results of the mean densities

Analyzing the results of volumetric density, it is possible to identify that PC samples present a lower result than the one obtained in the SLM samples. This premise is confirmed by the result obtained, evidencing that the PC process presents superior open porosity than SLM technology. When correlating with the density by Helium pycnometry, it can be verified that

Analyzing the Helium pycnometry, it is possible to verify that both consolidation processes

the theoretical density is possible to check the internal porosity, that results, respectively, for the PC and SLM process of 2.14 and 1.80%. It can be concluded that the SLM consolidation process produces samples with lower internal porosity, and can obtain components with densities around 98.20% of theoretical, in contrast to the PC process that obtains components with

The heating curves obtained from TMA in the form of PC and SLM sample are presented in **Figure 10**. To understand the events occurred at TMA, the heating curves of DSC analysis were juxtaposed. The CTE for the consolidated samples has a different behavior between the processes of the Co-Cr-Mo alloy in a similar analysis to the present study. The CTE for the consolidated samples has a distinct behavior between the processes of the Co-Cr-Mo alloy. As it is possible to verify the CTE at temperature of 500°C of samples (SLM 1, SLM 2, and PC) is, respectively, 15.0/19.5/22.0 × 10−6 °C−1. At the temperature of 600°C, the coefficient value decreases to the values of 12.5/14.5/18.5 × 10−6 °C−1. This difference is greater for the

). Relating the Helium pycnometry density to

obtained for the samples consolidated by PC and SLM can be seen in **Table 4**.

the open porosity results in 0.24% for PC sample and 0.12% for SLM sample.

have the same theoretical density (8.24 g/cm3

density of 97.86% of theoretical.

biological [7, 8, 61].

techniques.

**Figure 8.** Heating curves of DSC analysis for Co-Cr-Mo powder in different rates (10, 20, and 30°C/min).

in the Co-Cr binary diagram occurring at about 950°C). This occurrence is similar to that described in Ref. [16, 60], which obtains endothermic peaks, respectively, at approximately 970 and 1000°C, relative to that obtained in the present study of 944.52°C. This temperature difference is associated with the chemical composition of the alloy (64Co-29Cr-7Mo of the present study), which represent alloys according to ASTM F75-12 (stoichiometry is 66Co-28Cr-6Mo) and therefore there are temperature difference of 26 and 56°C relative to the cited references. This difference can be associated to different calibrations, among the equipment used in the analysis.

In interpreting the DSC curves, it is possible to verify the melting temperature of the Co-Cr-Mo alloy. By means of the average value of the three heating rates, the melting temperature is approximately 1354.5°C, with a variation of 4°C between rates.

#### **3.2. Evaluation of samples manufactured**

The result of the cytotoxicity analysis for the processed samples is shown in **Figure 9**. According to the cytotoxicity assay with respect to the pure extract, without dilution, the samples processed by precision casting and selective laser melting showed no toxicity.

**Figure 9.** Result of viability of cell growth (cytotoxicity analysis) for Co-Cr-Mo specimens manufactured by PC and SLM techniques.

The results showed the expected results for the Co-Cr-Mo alloy, because as there is a need for specific mechanical properties to be reached, the use of Co-Cr alloys for the manufacture of medical and dental implants or prostheses did not show any toxicity with the medium biological [7, 8, 61].

The porosity of the samples was determined and a comparison was made with average densities: theoretical, volumetric, and by helium pycnometry. The results of the mean densities obtained for the samples consolidated by PC and SLM can be seen in **Table 4**.

Analyzing the results of volumetric density, it is possible to identify that PC samples present a lower result than the one obtained in the SLM samples. This premise is confirmed by the result obtained, evidencing that the PC process presents superior open porosity than SLM technology. When correlating with the density by Helium pycnometry, it can be verified that the open porosity results in 0.24% for PC sample and 0.12% for SLM sample.

in the Co-Cr binary diagram occurring at about 950°C). This occurrence is similar to that described in Ref. [16, 60], which obtains endothermic peaks, respectively, at approximately 970 and 1000°C, relative to that obtained in the present study of 944.52°C. This temperature difference is associated with the chemical composition of the alloy (64Co-29Cr-7Mo of the present study), which represent alloys according to ASTM F75-12 (stoichiometry is 66Co-28Cr-6Mo) and therefore there are temperature difference of 26 and 56°C relative to the cited references. This difference can be associated to different calibrations, among the equipment used in the

**Figure 8.** Heating curves of DSC analysis for Co-Cr-Mo powder in different rates (10, 20, and 30°C/min).

In interpreting the DSC curves, it is possible to verify the melting temperature of the Co-Cr-Mo alloy. By means of the average value of the three heating rates, the melting temperature is

The result of the cytotoxicity analysis for the processed samples is shown in **Figure 9**. According to the cytotoxicity assay with respect to the pure extract, without dilution, the samples pro-

approximately 1354.5°C, with a variation of 4°C between rates.

cessed by precision casting and selective laser melting showed no toxicity.

**3.2. Evaluation of samples manufactured**

analysis.

172 Biomaterials in Regenerative Medicine

Analyzing the Helium pycnometry, it is possible to verify that both consolidation processes have the same theoretical density (8.24 g/cm3 ). Relating the Helium pycnometry density to the theoretical density is possible to check the internal porosity, that results, respectively, for the PC and SLM process of 2.14 and 1.80%. It can be concluded that the SLM consolidation process produces samples with lower internal porosity, and can obtain components with densities around 98.20% of theoretical, in contrast to the PC process that obtains components with density of 97.86% of theoretical.

The heating curves obtained from TMA in the form of PC and SLM sample are presented in **Figure 10**. To understand the events occurred at TMA, the heating curves of DSC analysis were juxtaposed. The CTE for the consolidated samples has a different behavior between the processes of the Co-Cr-Mo alloy in a similar analysis to the present study. The CTE for the consolidated samples has a distinct behavior between the processes of the Co-Cr-Mo alloy. As it is possible to verify the CTE at temperature of 500°C of samples (SLM 1, SLM 2, and PC) is, respectively, 15.0/19.5/22.0 × 10−6 °C−1. At the temperature of 600°C, the coefficient value decreases to the values of 12.5/14.5/18.5 × 10−6 °C−1. This difference is greater for the


sample (at DSC analysis) has a dendritic microstructure, as the same of the PC sample (at TMA). TMA curves of SLM samples (SLM 1 and SLM 2) show more evident peaks, but the transition of these events set at increased temperatures. Thus, the laser fusion process has a refined and more homogeneous microstructure what hinders phase transitions and requires

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The mechanical results of the tests for the PC and SLM specimens are present in **Table 5**. Analyzing the values is possible to verify that in all properties the SLM technique results in higher properties than PC technique. According to standard ISO22674:06 [54], the SLM and

The result of TRS samples (SLM = 2501.2 ± 9.7 MPa and PC = 1072.3 ± 4.6 MPa) was satisfactory. However, there was no rupture of PC sample (test interrupted) evidenced the ductility of the precision casting process, that was confirmed by the value of higher elongation. According to Mengucci et al. [62], the TRS result for a similar composition of Co-Cr-Mo alloy, after the shoot-peened treatment followed by heat treatment for strain relief, resulted a TRS equal to 2700 ± 25 MPa. Therefore, the present results are acceptable comparing the data

The heat treatment, as hot isostatic pressing (HIP), after the additive manufacturing process (by laser melting process–SLM and EBM) of parts has been used successfully by medical and aeronautic manufactures. The HIP process is effective to obtain better results of mechanical properties (ductility and fatigue resistance) and decrease the porosity [60, 63]. Although this present study evaluated the mechanical properties of SLM and PC samples without any post-process of heat treatment is possible, check the relevant mechanical properties obtained by manufacture process. According to the results present evaluation is presented in **Table 6** to compare the mechanical properties to those presented in the literature

**Table 5.** Mechanical properties of Co-Cr-Mo alloy manufactured by PC and SLM compared to the minimum properties

**PC SLM ISO 22674 "type 5" ASTM F75 "casting"**

PC specimens satisfied the type 5 criteria in all mechanical properties.

**Mechanical properties Consolidation technique Standards**

Yield strength (MPa) 646.7 ± 44.4 731.5 ± 40.3 500 450 Rupture strength (MPa) 742.2 ± 106.8 1127.9 ± 0.1 – – Ultimate tensile strength (MPa) 771.7 ± 103.3 1136.9 ± 1.0 – 655 Elongation (%) 14.20 ± 2.8 13.7 ± 5.3 2 8 Elastic modulus "E" (GPa) 223.42 ± 15.7 225.2 ± 14.4 150 – Hardness Vickers (HV) 272.2 ± 20.5 334.8 ± 16.0 – 266−345 TRS (MPa) 1072.3 ± 4.6 2501.2 ± 9.7 – –

obtained with the study by Mengucci et al. [62].

higher temperatures.

[14, 41, 44, 62, 64].

required by standards.

**3.3. Mechanical behavior**

**Table 4.** Results of densities (medium values): theoretical, Helium pycnometry and volumetric for samples PC and SLM.

alloy processed by PC, which is associated to the dendritic microstructure formed and in relation to the casting process because it characterizes a fine grain microstructure and more packaging.

As can be seen in the TMA and DSC curves are similar in the temperature ranges of events. Facchini [60] shows the DSC and TMA curves (heating rate of 20°C/min) performed in an ASTM F75-12 composition alloy and processed by electron beam fusion (EBM). EBM has an effect similar to the SLM because both processes have a concentrated heat source and is possible to relate the DSC analysis, in which, the first peak (565–900°C) is associated with the transition from the FCC (αCo) phase to the HCP (εCo) phase, and the second peak (900–1000°C) reduces the HCP phase (εCo) and reappearance of the FCC phase (αCo).

Is possible observed two events, the first event occurs in the range of 514–614°C and the second event at 923–961°C. In the case of the PC sample (TMA curve), it is possible to verify that the events occur in a higher temperature in relation to the samples processed by SLM. This occurrence is associated to the microstructure samples of analysis performed. The powder

**Figure 10.** Heat curve of thermal analysis for Co-Cr-Mo samples: DSC curve of Co-Cr-Mo powder sample and TMA curves of consolidated samples PC and SLM (SLM 1 and SLM 2).

sample (at DSC analysis) has a dendritic microstructure, as the same of the PC sample (at TMA). TMA curves of SLM samples (SLM 1 and SLM 2) show more evident peaks, but the transition of these events set at increased temperatures. Thus, the laser fusion process has a refined and more homogeneous microstructure what hinders phase transitions and requires higher temperatures.

#### **3.3. Mechanical behavior**

alloy processed by PC, which is associated to the dendritic microstructure formed and in relation to the casting process because it characterizes a fine grain microstructure and more

**Table 4.** Results of densities (medium values): theoretical, Helium pycnometry and volumetric for samples PC and SLM.

**Theoretical Helium pycnometry Volumetric**

As can be seen in the TMA and DSC curves are similar in the temperature ranges of events. Facchini [60] shows the DSC and TMA curves (heating rate of 20°C/min) performed in an ASTM F75-12 composition alloy and processed by electron beam fusion (EBM). EBM has an effect similar to the SLM because both processes have a concentrated heat source and is possible to relate the DSC analysis, in which, the first peak (565–900°C) is associated with the transition from the FCC (αCo) phase to the HCP (εCo) phase, and the second peak (900–1000°C)

Is possible observed two events, the first event occurs in the range of 514–614°C and the second event at 923–961°C. In the case of the PC sample (TMA curve), it is possible to verify that the events occur in a higher temperature in relation to the samples processed by SLM. This occurrence is associated to the microstructure samples of analysis performed. The powder

**Figure 10.** Heat curve of thermal analysis for Co-Cr-Mo samples: DSC curve of Co-Cr-Mo powder sample and TMA

curves of consolidated samples PC and SLM (SLM 1 and SLM 2).

reduces the HCP phase (εCo) and reappearance of the FCC phase (αCo).

**)**

PC 8.42 8.24 ± 0.01 8.22 ± 0.10 SLM 8.39 8.24 ± 0.01 8.23 < 0.01

packaging.

**Sample Densities (g/cm3**

174 Biomaterials in Regenerative Medicine

The mechanical results of the tests for the PC and SLM specimens are present in **Table 5**. Analyzing the values is possible to verify that in all properties the SLM technique results in higher properties than PC technique. According to standard ISO22674:06 [54], the SLM and PC specimens satisfied the type 5 criteria in all mechanical properties.

The result of TRS samples (SLM = 2501.2 ± 9.7 MPa and PC = 1072.3 ± 4.6 MPa) was satisfactory. However, there was no rupture of PC sample (test interrupted) evidenced the ductility of the precision casting process, that was confirmed by the value of higher elongation. According to Mengucci et al. [62], the TRS result for a similar composition of Co-Cr-Mo alloy, after the shoot-peened treatment followed by heat treatment for strain relief, resulted a TRS equal to 2700 ± 25 MPa. Therefore, the present results are acceptable comparing the data obtained with the study by Mengucci et al. [62].

The heat treatment, as hot isostatic pressing (HIP), after the additive manufacturing process (by laser melting process–SLM and EBM) of parts has been used successfully by medical and aeronautic manufactures. The HIP process is effective to obtain better results of mechanical properties (ductility and fatigue resistance) and decrease the porosity [60, 63]. Although this present study evaluated the mechanical properties of SLM and PC samples without any post-process of heat treatment is possible, check the relevant mechanical properties obtained by manufacture process. According to the results present evaluation is presented in **Table 6** to compare the mechanical properties to those presented in the literature [14, 41, 44, 62, 64].


**Table 5.** Mechanical properties of Co-Cr-Mo alloy manufactured by PC and SLM compared to the minimum properties required by standards.


**Table 6.** Comparative results for mechanical properties obtained in the present study with Co-Cr alloys manufactured by selective laser melting (SLM) and casting (C) process presented in the literature.

#### **3.4. Microstructural analysis**

To understand the mechanical properties improved in the SLM specimens in relation to the casting process technique carried out the microstructural analysis by OM and SEM-EDS.

The microstructural analysis by OM of PC samples (**Figure 11a,b**) describe dendritic arms and ramifications with different solidification orientations [14]. In addition, PC sample present porous (microporous) as the SLM samples, but are uneven (a little larger but in small quantity). This occurrence is possible to form by problems of dispersing the powder in the bed layer and the presence of satellites/porous in the powder particles. SLM specimens show a characteristic morphology (weld-like structure) of laser beam melting. Is possible to check the layers formed during the manufacture process (**Figure 11c,d**)? The vertical section of SLM sample is to observe the building direction of specimen (indicated by arrow—**Figure 11d**) characterized by the overlapping of each layer and the morphology formation by the action of the laser beam such as the weld pool.

molybdenum [14]. The M23C6

(Etch: 100 ml HCl and 2 ml H<sup>2</sup>

in opposition of 338 ± 14 HV (1 mN/15 s) to the micro hardness of matrix.

. Magnitude: ×200).

better mechanical properties than the cast technique.

O2

**Figure 13** shows the microstructure of SLM specimen. It is observed that a microstructure is formed with small grains characterizing the rapid solidification during the SLM manufacturing process. The semi-quantitative analysis in the fine grains shows that it does not have different elements compositions. SLM specimen presents a homogeneous matrix with Co-Cr-Mo elements. The morphology formation of the laser melting sample was also observed after electrolytic attack [59, 61]. Also confirms that the fine grains are oriented in direction of the laser scanning. This characteristic microstructure of laser melting technique allows to achieve

**Figure 11.** MO images of Co-Cr-Mo alloy consolidated: (a–b) transversal and longitudinal view of precision casting sample and (c–d) transversal and longitudinal (arrow indicates building consolidation) view of selective laser melting

The fractures of the tensile samples were SEM analyzed (**Figure 14**) observed the formation of dimples homogeneously distributed in the microfracture of both samples. Regions with the presence of dimples are ductless and with higher toughness. However, it is apparent that the dimple formations on the SLM sample extends completely by the fracture planes and are of

carbide results in a micro hardness of 699 ± 131 HV (1 mN/15 s)

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**Figure 12** represents the SEM images of PC samples and the semi-quantitative analysis of interesting points by EDS. It is possible to identify the cast specimen with a second phase (white area) in the matrix. The semi-quantitative analysis with the EDS and the respective spectrums (**Figure 12c,d**) show that the composition of white area (point 1) is rich in Mo element, and the matrix (point 2) is composed by Co-Cr elements, with a small percentage of Mo. The phase (point 1) shows the confirmation of carbide (M23C6 ) presence, rich in chromium and

**Figure 11.** MO images of Co-Cr-Mo alloy consolidated: (a–b) transversal and longitudinal view of precision casting sample and (c–d) transversal and longitudinal (arrow indicates building consolidation) view of selective laser melting (Etch: 100 ml HCl and 2 ml H<sup>2</sup> O2 . Magnitude: ×200).

**3.4. Microstructural analysis**

C 60-65Co 26-30Cr 5-7Mo

SLM 60-65Co 26-30Cr 5-7Mo

SLM 63.8Co-24.7Cr-5.1Mo-5.4W

Present work Co-Cr-Mo

176 Biomaterials in Regenerative Medicine

Takaichi et al. [64]

Qian et al. [41]

Kajima et al. [44]

Mengucci et al. [62]

Jabbari et al. [14]

Liverani et al. [40]

the laser beam such as the weld pool.

The phase (point 1) shows the confirmation of carbide (M23C6

To understand the mechanical properties improved in the SLM specimens in relation to the casting process technique carried out the microstructural analysis by OM and SEM-EDS.

**Table 6.** Comparative results for mechanical properties obtained in the present study with Co-Cr alloys manufactured

by selective laser melting (SLM) and casting (C) process presented in the literature.

**References Alloy (wt%) σYS (MPa) El (%) σUTS (MPa) TRS (MPa) Hardness** 

C Co-28Cr-6Mo 296 ± 25 9.6 ± 2.5 912 ± 39 – – SLM 516 ± 28 10.7 ± 2.9 591 ± 37 – –

SLM 873 ± 76 – 1303 ± 73 – –

C 63Co-29Cr-6Mo 571 ± 23 11.2 ± 2 775 ± 67 – –

C 61.6Co-30Cr-6.5Mo – – – – 320 ± 12 SLM Co-29Cr-5.5Mo – – – – 371 ± 10

SLM Co 27-30Cr 5-7Mo 677 – – – 361 ± 31

PC 64Co-29Cr-7Mo 646.76 ± 44.36 14.20 ± 2.76 771.70 ± 103.32 1072.3 ± 4.6 256.7 ± 12.9 SLM 731.50 ± 40.31 13.73 ± 5.32 1136.95 ± 0.92 2501.2 ± 9.7 358.1 ± 9.8

610 – 741 – –

877 ± 37 12.3 ± 3 1170 ± 29 – –

– – 1340 ± 20 2700 ± 25 434 ± 22

The microstructural analysis by OM of PC samples (**Figure 11a,b**) describe dendritic arms and ramifications with different solidification orientations [14]. In addition, PC sample present porous (microporous) as the SLM samples, but are uneven (a little larger but in small quantity). This occurrence is possible to form by problems of dispersing the powder in the bed layer and the presence of satellites/porous in the powder particles. SLM specimens show a characteristic morphology (weld-like structure) of laser beam melting. Is possible to check the layers formed during the manufacture process (**Figure 11c,d**)? The vertical section of SLM sample is to observe the building direction of specimen (indicated by arrow—**Figure 11d**) characterized by the overlapping of each layer and the morphology formation by the action of

**Figure 12** represents the SEM images of PC samples and the semi-quantitative analysis of interesting points by EDS. It is possible to identify the cast specimen with a second phase (white area) in the matrix. The semi-quantitative analysis with the EDS and the respective spectrums (**Figure 12c,d**) show that the composition of white area (point 1) is rich in Mo element, and the matrix (point 2) is composed by Co-Cr elements, with a small percentage of Mo.

) presence, rich in chromium and

**(HV)**

molybdenum [14]. The M23C6 carbide results in a micro hardness of 699 ± 131 HV (1 mN/15 s) in opposition of 338 ± 14 HV (1 mN/15 s) to the micro hardness of matrix.

**Figure 13** shows the microstructure of SLM specimen. It is observed that a microstructure is formed with small grains characterizing the rapid solidification during the SLM manufacturing process. The semi-quantitative analysis in the fine grains shows that it does not have different elements compositions. SLM specimen presents a homogeneous matrix with Co-Cr-Mo elements. The morphology formation of the laser melting sample was also observed after electrolytic attack [59, 61]. Also confirms that the fine grains are oriented in direction of the laser scanning. This characteristic microstructure of laser melting technique allows to achieve better mechanical properties than the cast technique.

The fractures of the tensile samples were SEM analyzed (**Figure 14**) observed the formation of dimples homogeneously distributed in the microfracture of both samples. Regions with the presence of dimples are ductless and with higher toughness. However, it is apparent that the dimple formations on the SLM sample extends completely by the fracture planes and are of

**Figure 12.** SEM images of PC sample: (a) magnitude ×2000, (b) magnitude ×8000, and (c–d) EDS spectrograms of analysis at point 1 and point 2.

finer size, compared to the fracture with dimples geometrically larger of PC sample. It can be verified with SLM samples and confirmed with the mechanical results in relation to PC samples. In addition, some planar regions (indicated by arrows) show a semi-cleavage morphology. The type of fracture observed in the samples, according to Takaichi et al. [64], describes the formation of dimples along the fracture surface, as well as cracking of the wedge is appointed

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In general, the results of powders characterization showed that the granulometric range of 20–50 μm is the one that best fits in the properties of packaging, for the consolidation by SLM. The biocompatibility of the samples obtained a positive result for both processing techniques. In this way, the development of the present study evidenced to improve the manufacture of

Microstructural analysis obtained for SLM samples results in a characteristic morphology of layer manufacturing with ultrafine grains and a high chemical homogeneity. The conventional technique presented a differentiated microstructure by the gross dendritic microstruc-

The mechanical evaluation showed that the SLM technique provides superior mechanical properties (as yield strength, rupture strength, ultimate tensile strength, TRS, and hardness)

The thermal analyses showed the present phase transitions of Co-Cr-Mo alloy, as well as being possible to correlate them (TMA to DSC curves). The coefficient of thermal expansion

(CTE) resulted for both processes a similar value to alloys used in dental materials.

as a possible formation of cleavage fracture over favorable crystallographic planes.

**Figure 14.** SEM images of tensile fracture: (a) PC sample and (b) SLM sample (magnitude: ×2000).

customized dental components (copings) using the SLM technique.

compared to those obtained by the precision casting technique.

**4. Conclusions**

ture of casting process.

**Figure 13.** SEM images of SLM specimens: (a) horizontal section from backscattered electrons, and (b) from secondary electrons, (c–d) vertical section from backscattered electrons (black arrow indicates the building consolidation).

**Figure 14.** SEM images of tensile fracture: (a) PC sample and (b) SLM sample (magnitude: ×2000).

finer size, compared to the fracture with dimples geometrically larger of PC sample. It can be verified with SLM samples and confirmed with the mechanical results in relation to PC samples. In addition, some planar regions (indicated by arrows) show a semi-cleavage morphology. The type of fracture observed in the samples, according to Takaichi et al. [64], describes the formation of dimples along the fracture surface, as well as cracking of the wedge is appointed as a possible formation of cleavage fracture over favorable crystallographic planes.

#### **4. Conclusions**

**Figure 12.** SEM images of PC sample: (a) magnitude ×2000, (b) magnitude ×8000, and (c–d) EDS spectrograms of analysis

**Figure 13.** SEM images of SLM specimens: (a) horizontal section from backscattered electrons, and (b) from secondary electrons, (c–d) vertical section from backscattered electrons (black arrow indicates the building consolidation).

at point 1 and point 2.

178 Biomaterials in Regenerative Medicine

In general, the results of powders characterization showed that the granulometric range of 20–50 μm is the one that best fits in the properties of packaging, for the consolidation by SLM.

The biocompatibility of the samples obtained a positive result for both processing techniques. In this way, the development of the present study evidenced to improve the manufacture of customized dental components (copings) using the SLM technique.

Microstructural analysis obtained for SLM samples results in a characteristic morphology of layer manufacturing with ultrafine grains and a high chemical homogeneity. The conventional technique presented a differentiated microstructure by the gross dendritic microstructure of casting process.

The mechanical evaluation showed that the SLM technique provides superior mechanical properties (as yield strength, rupture strength, ultimate tensile strength, TRS, and hardness) compared to those obtained by the precision casting technique.

The thermal analyses showed the present phase transitions of Co-Cr-Mo alloy, as well as being possible to correlate them (TMA to DSC curves). The coefficient of thermal expansion (CTE) resulted for both processes a similar value to alloys used in dental materials.

The processing using laser melting proved better mechanical and thermal properties to precision casting processing technique without post-processing (thermal treatment). SLM technique evidenced a promising use to manufacture prosthetics and dental implants. Nevertheless, it is still of great concern and promising further development of laser melting process (SLM and EBM) in relation to the parameters and variables of process, as also to the post-processing method apply to AM parts. Such characteristics should be addressed to new materials and investigate in relation to the performance and bio-functionality of specific application part.

[7] Wataha JC. Biocompatibility of dental casting alloys: A review. The Journal of Prosthetic

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## **Acknowledgements**

This study was financially support by CNPq and FAPESP. The authors also thank to Ms. Amed Belaid and SLM® Solution for SLM specimens and collaboration.

## **Author details**

Marcello Vertamatti Mergulhão\*, Carlos Eduardo Podestá and Maurício David Martins das Neves

\*Address all correspondence to: marcellovertamatti@gmail.com

Nuclear and Energy Research Institute (IPEN/CNEN-SP), CCTM, São Paulo, Brazil

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The processing using laser melting proved better mechanical and thermal properties to precision casting processing technique without post-processing (thermal treatment). SLM technique evidenced a promising use to manufacture prosthetics and dental implants. Nevertheless, it is still of great concern and promising further development of laser melting process (SLM and EBM) in relation to the parameters and variables of process, as also to the post-processing method apply to AM parts. Such characteristics should be addressed to new materials and investigate in relation to the performance and bio-functionality of specific

This study was financially support by CNPq and FAPESP. The authors also thank to Ms.

Amed Belaid and SLM® Solution for SLM specimens and collaboration.

Marcello Vertamatti Mergulhão\*, Carlos Eduardo Podestá and

\*Address all correspondence to: marcellovertamatti@gmail.com

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184 Biomaterials in Regenerative Medicine


**Chapter 7**

**Provisional chapter**

**Application of 3-D Printing for Tissue Regeneration in**

The ultimate goal of any surgical procedure is to improve perioperative form and function and to minimize operative and postoperative morbidity. In recent years, many exciting and novel technological advances have been introduced in the field of oral and maxillofacial surgery. One example of such technology that is continuing to increase in prevalence is the use of 3-dimensional (3-D) printing techniques with special properties, which seems hopeful for practitioners in the field of regenerative medicine. Tissue engineering is a critical and important area in biomedical engineering for creating biological alternatives for grafts, implants, and prostheses. One of the main triad bases for tissue engineering is scaffolds, which play a great role for determining growth directions of stem cells in a 3-dimensional aspect. Mechanical strength of these scaffolds is critical as well as interconnected channels and controlled porosity or pores distribution. However, existing 3-D scaffolds proved less than ideal for actual clinical applications. In this chapter, we review the application and advancement of rapid prototyping (RP) techniques in the design and creation of synthetic scaffolds for use in tissue engineering. Also, we

**Keywords:** 3-D printing, prototyping, tissue engineering, scaffolds, bioprinting, stem cells,

**Application of 3-D Printing for Tissue Regeneration in** 

DOI: 10.5772/intechopen.70323

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Three-dimensional printing—also known as rapid prototyping—was first introduced in 1980s; during past three decades, enormous changes and development have been performed

by scientists through modifying this technology by uses, material, and also accuracy.

**Oral and Maxillofacial Surgery: What is Upcoming?**

**Oral and Maxillofacial Surgery: What is Upcoming?**

Seied Omid Keyhan, Hamidreza Fallahi,

Seied Omid Keyhan, Hamidreza Fallahi, Alireza Jahangirnia, Mohammad Taher Amirzade-Iranaq and Mohammad Hosein

Mohammad Taher Amirzade-Iranaq and Mohammad Hosein Amirzade-Iranaq

Additional information is available at the end of the chapter

survey through new and novel merging era of "bioprinting."

regenerative medicine, oral surgery, maxillofacial surgery

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.70323

Alireza Jahangirnia,

Amirzade-Iranaq

**Abstract**

**1. Introduction**

**Provisional chapter**

#### **Application of 3-D Printing for Tissue Regeneration in Oral and Maxillofacial Surgery: What is Upcoming? Application of 3-D Printing for Tissue Regeneration in Oral and Maxillofacial Surgery: What is Upcoming?**

DOI: 10.5772/intechopen.70323

Seied Omid Keyhan, Hamidreza Fallahi, Alireza Jahangirnia, Mohammad Taher Amirzade-Iranaq and Mohammad Hosein Amirzade-Iranaq Seied Omid Keyhan, Hamidreza Fallahi, Alireza Jahangirnia, Mohammad Taher Amirzade-Iranaq and Mohammad Hosein Amirzade-Iranaq

Additional information is available at the end of the chapter Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.70323

#### **Abstract**

The ultimate goal of any surgical procedure is to improve perioperative form and function and to minimize operative and postoperative morbidity. In recent years, many exciting and novel technological advances have been introduced in the field of oral and maxillofacial surgery. One example of such technology that is continuing to increase in prevalence is the use of 3-dimensional (3-D) printing techniques with special properties, which seems hopeful for practitioners in the field of regenerative medicine. Tissue engineering is a critical and important area in biomedical engineering for creating biological alternatives for grafts, implants, and prostheses. One of the main triad bases for tissue engineering is scaffolds, which play a great role for determining growth directions of stem cells in a 3-dimensional aspect. Mechanical strength of these scaffolds is critical as well as interconnected channels and controlled porosity or pores distribution. However, existing 3-D scaffolds proved less than ideal for actual clinical applications. In this chapter, we review the application and advancement of rapid prototyping (RP) techniques in the design and creation of synthetic scaffolds for use in tissue engineering. Also, we survey through new and novel merging era of "bioprinting."

**Keywords:** 3-D printing, prototyping, tissue engineering, scaffolds, bioprinting, stem cells, regenerative medicine, oral surgery, maxillofacial surgery

## **1. Introduction**

Three-dimensional printing—also known as rapid prototyping—was first introduced in 1980s; during past three decades, enormous changes and development have been performed by scientists through modifying this technology by uses, material, and also accuracy.

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

With increasing attention of scientific societies, recently, scientific literature bolded feasibility of 3-D-printed tissues and organs and its usage within laborious clinical situations. Also, this technology was used largely in accurate and highly customized devices, such as tracheobronchial splints, bionic ears, and even more. Within the field of craniofacial surgery, 3-D surgical models have been used as templates to create bone grafts, tailoring bioprosthetic implants, plate bending, cutting guides for osteotomies, and intraoperative oral splints. Using 3-D models and guides has been shown to shorten operative time and potentially reduce the complications associated with prolonged operative times.

The developing agora for 3-D desktop printers encourages wide-ranging experimentations in that subject. Generally, medical indications of these printers are such as treatment planning,

Application of 3-D Printing for Tissue Regeneration in Oral and Maxillofacial...

http://dx.doi.org/10.5772/intechopen.70323

189

Having being used in military, food industry, and art, rapid prototyping is receiving a lot of

The pioneering usage of stereolithography in oral and maxillofacial surgery was by Brix and Lambrecht in 1985. Later this technique was used by them for treatment planning in cranio-

In 1990, stereolithography was used by Mankovich et al. for treating patients having craniofacial deformities [16, 17]. They used it to simulate bony anatomy of the cranium using com-

By aiding in complex craniofacial reconstructions, 3-D printing has recently earned reputa-

Today, maxillofacial surgery can benefit from additive manufacturing in various aspects and different clinical cases [22]. This technique can help with bending plates, manufacturing templates for bone grafts, tailoring implants, osteotomy guides, and intraoperative occlusal splints [23–27]. Rapid prototyping can shorten surgery duration and simplify pre and intra-

**2. Current 3-D printing techniques used in oral and maxillofacial surgery**

From first innovation till nowadays, there are different kind of technologies introduced for 3-D printing. Binder jetting (BJ), electron beam melting (EBM), fused deposition modeling (FDM), indirect processes, laser melting (LM), laser sintering (LS), material jetting (MJ), photopolymer jetting (PJ), and stereolithography (SL) are well-known technologies of 3-D printing [14, 28, 29]. There are many different 3-D printing techniques. Benefits and disadvantages are factors to differ each technology system [14]. Among this variety of different techniques, there is a huge

discussion and usage in oral and maxillofacial region for SL, FDM, and PJ [1, 28, 30].

marizes some different three dimensional printing technologies.

**3. Biomaterials available for 3-D printing**

Each technology has its own characteristics, properties, and advantages which **Table 1** sum-

As researchers aim to investigate new materials for 3-D printing in last decade, it is obvious to see variety of biomaterials with different properties and also different applications. As **Table 2** summarizes all biomaterials used within studies all over the world for generating scaffolds for bone tissue engineering, it has to be noticed that from this large spectrum of biomaterials

operative decisions. It has enhanced efficacy and preciseness of surgeries [10].

prosthesis, implant fabrications, medical training, and other usages [4].

attention in the field of surgery in the last 10 years [6, 14].

puted tomography with complete internal components [17, 18].

tion in medicine and surgical fields [19–21].

facial surgery [15].

The goal of surgical procedures for a clinician is to improve perioperative form, recovery of function, and also minimizing operative and postoperative morbidity. Many exciting and new technological advances have ushered in a new era in the field of oral and maxillofacial surgery over the last years, which within no exaggeration 3-D printing is the novelist and controversial one.

The aim of this chapter is to introduce 3-D printing method and its role in the contemporary oral and maxillofacial surgery and to review current advantages of its application in the field of regenerative medicine.

#### **1.1. History of the technology**

Three-dimensional (3-D) printing has been utilized in diverse aspects of manufacturing to produce different objects from guns, boats, and food to models of unborn babies. From over 1450 articles related to 3-D printing listed in PubMed, nearly a third of them were solely published in the last 2 years [1].

Three-dimensional (3-D) printing is a manufacturing process that objects are fabricated in a layering method during fusing or depositing different materials such as plastic, metal, ceramics, powders, liquids, or even living cells to build a 3-D matter [2, 3]. It is a process of generating physical models from digital layouts [4, 5]. This technology demonstrates a technique that a product designed through a computer-aided scheme is manufactured in a layer-by-layer system [6]. This process is also cited as rapid prototyping (RP), solid freeform technology (SFF), or additive manufacturing (AM) [7].

3-D printing techniques are not brand new and have been existed since 30 years ago [8–10]. This technology is first introduced and invented by Charles Hull in 1986, and at first, it was utilized in the engineering and automobile industry for manufacturing polyurethane frameworks for different models, pieces, and instruments [11]. Originally, Hull employed the phrase "stereolithography" in his US Patent 4,575,330, termed "Apparatus for Production of Three—Dimensional Objects by Stereolithography" published in 1986. Stereolithography technique included subjoining layers over the top of each other, by curing photopolymers with UV lasers [12, 13].

Since then, 3-D models have been used for a diversity of different objectives. Since 1986, this process has started to accelerate and has honored recognition globally and has influenced different arenas, such as medicine.

The developing agora for 3-D desktop printers encourages wide-ranging experimentations in that subject. Generally, medical indications of these printers are such as treatment planning, prosthesis, implant fabrications, medical training, and other usages [4].

With increasing attention of scientific societies, recently, scientific literature bolded feasibility of 3-D-printed tissues and organs and its usage within laborious clinical situations. Also, this technology was used largely in accurate and highly customized devices, such as tracheobronchial splints, bionic ears, and even more. Within the field of craniofacial surgery, 3-D surgical models have been used as templates to create bone grafts, tailoring bioprosthetic implants, plate bending, cutting guides for osteotomies, and intraoperative oral splints. Using 3-D models and guides has been shown to shorten operative time and potentially reduce the complica-

The goal of surgical procedures for a clinician is to improve perioperative form, recovery of function, and also minimizing operative and postoperative morbidity. Many exciting and new technological advances have ushered in a new era in the field of oral and maxillofacial surgery over the last years, which within no exaggeration 3-D printing is the novelist and

The aim of this chapter is to introduce 3-D printing method and its role in the contemporary oral and maxillofacial surgery and to review current advantages of its application in the field

Three-dimensional (3-D) printing has been utilized in diverse aspects of manufacturing to produce different objects from guns, boats, and food to models of unborn babies. From over 1450 articles related to 3-D printing listed in PubMed, nearly a third of them were solely pub-

Three-dimensional (3-D) printing is a manufacturing process that objects are fabricated in a layering method during fusing or depositing different materials such as plastic, metal, ceramics, powders, liquids, or even living cells to build a 3-D matter [2, 3]. It is a process of generating physical models from digital layouts [4, 5]. This technology demonstrates a technique that a product designed through a computer-aided scheme is manufactured in a layer-by-layer system [6]. This process is also cited as rapid prototyping (RP), solid freeform technology

3-D printing techniques are not brand new and have been existed since 30 years ago [8–10]. This technology is first introduced and invented by Charles Hull in 1986, and at first, it was utilized in the engineering and automobile industry for manufacturing polyurethane frameworks for different models, pieces, and instruments [11]. Originally, Hull employed the phrase "stereolithography" in his US Patent 4,575,330, termed "Apparatus for Production of Three—Dimensional Objects by Stereolithography" published in 1986. Stereolithography technique included subjoining layers over the top of each other, by curing photopolymers

Since then, 3-D models have been used for a diversity of different objectives. Since 1986, this process has started to accelerate and has honored recognition globally and has influenced dif-

tions associated with prolonged operative times.

controversial one.

of regenerative medicine.

188 Biomaterials in Regenerative Medicine

**1.1. History of the technology**

lished in the last 2 years [1].

with UV lasers [12, 13].

ferent arenas, such as medicine.

(SFF), or additive manufacturing (AM) [7].

Having being used in military, food industry, and art, rapid prototyping is receiving a lot of attention in the field of surgery in the last 10 years [6, 14].

The pioneering usage of stereolithography in oral and maxillofacial surgery was by Brix and Lambrecht in 1985. Later this technique was used by them for treatment planning in craniofacial surgery [15].

In 1990, stereolithography was used by Mankovich et al. for treating patients having craniofacial deformities [16, 17]. They used it to simulate bony anatomy of the cranium using computed tomography with complete internal components [17, 18].

By aiding in complex craniofacial reconstructions, 3-D printing has recently earned reputation in medicine and surgical fields [19–21].

Today, maxillofacial surgery can benefit from additive manufacturing in various aspects and different clinical cases [22]. This technique can help with bending plates, manufacturing templates for bone grafts, tailoring implants, osteotomy guides, and intraoperative occlusal splints [23–27]. Rapid prototyping can shorten surgery duration and simplify pre and intraoperative decisions. It has enhanced efficacy and preciseness of surgeries [10].

## **2. Current 3-D printing techniques used in oral and maxillofacial surgery**

From first innovation till nowadays, there are different kind of technologies introduced for 3-D printing. Binder jetting (BJ), electron beam melting (EBM), fused deposition modeling (FDM), indirect processes, laser melting (LM), laser sintering (LS), material jetting (MJ), photopolymer jetting (PJ), and stereolithography (SL) are well-known technologies of 3-D printing [14, 28, 29].

There are many different 3-D printing techniques. Benefits and disadvantages are factors to differ each technology system [14]. Among this variety of different techniques, there is a huge discussion and usage in oral and maxillofacial region for SL, FDM, and PJ [1, 28, 30].

Each technology has its own characteristics, properties, and advantages which **Table 1** summarizes some different three dimensional printing technologies.

## **3. Biomaterials available for 3-D printing**

As researchers aim to investigate new materials for 3-D printing in last decade, it is obvious to see variety of biomaterials with different properties and also different applications. As **Table 2** summarizes all biomaterials used within studies all over the world for generating scaffolds for bone tissue engineering, it has to be noticed that from this large spectrum of biomaterials


**Techniques Advantages Disadvantages**

may be heat sterilized.

**Composed scaffolds Synthetic scaffolds Natural scaffolds**

High porosity. Variable mechanical strength. Low- to mid-range cost materials and equipment. Low accuracy in low cost equipment. Some materials

Ceramic Polymeric Inorganic Organic

βTCP PLG Coral PRP

Flurohydroxyapatite PGA Premineralized silk

coated PLA

βTCP/collagen PLA-PEG Deer horn Alginate

OsteoSet PEG-MMP Allograft

Apatite-coated PLGA Collagen gel

**Table 2.** Types of scaffolds used in bone tissue engineering in maxillo-craniofacial region [51].

TCP, tri-calcium phosphate; HA, hydroxyapatite; DBM, demineralized bone matrix; PLGA, poly(lactic-co-glycolic acid); PLA, poly(d,l-lactic acid); PGA, poly(glycolic acid); PLLA, poly(l-lactic acid); PVDC, polyvinylidene chloride; PEG, polyethylene glycol; DA, diacrylate; MMP, matrix metalloproteinases; ABB, anorganic bovine bone; Puramatrix, a self-

HA/TCP PLLA Silk fibroin protein Gelatin sponge

Low cost but limited materials—

http://dx.doi.org/10.5772/intechopen.70323

191

Gelatin Hydrogel

Partially demineralized

Hyaluronic acid based

hydrogel

bone matrix

only thermoplastics. Limited shape complexity for biological materials. Support material must be

removed.

Application of 3-D Printing for Tissue Regeneration in Oral and Maxillofacial...

PLGA Silver Collagen sponge

fibroin protein

PLA ABB PuraMatrix

PEG-DA Bio-Oss

PVDC Fibrin sealant

Polycaprolactone Gelatin foam

**Thermoplastic**

onto build platform.

Nano-hydroxyapatite/ collagen/PLLA

Octacalcium phosphate/

Nano-hydroxyapatite/ polyamide 6

Nano-hydroxyapatite/ polyamide66

Hydroxyapatite-coated

Nano-hydroxyapatite/

Octacalcium phosphate precipitated (OCP)

Demineralized bone powders/PLA

assembling peptide nanomaterial.

polyamide

alginate

HA/PLGA Ca deficient

collagen

PLGA

Fused deposition modeling (FDM) First 3-DP technology, most used in 'home' printers. Thermoplastic material extruded through nozzle

**Table 1.** 3-D printing modalities and materials [14, 31].

Calcium Magnesium Phosphate cement (CMPC)

hydroxyapatite (CDHA)

DBM/PLA Fibronectin-


**Table 1.** 3-D printing modalities and materials [14, 31].

**Techniques Advantages Disadvantages**

Rapid fabrication.

used in bulk.

technology.

Lower cost technology.

fast.

support

and accurate parts. Self-supported process.

possible.

Able to create complex shapes with high feature resolution. Lower cost materials if

Relatively fast. High resolution, high-quality finish possible. Multiple materials available various colors and physical properties including elastic materials. Lower cost

Good accuracy, smooth surfaces, relatively

Lower cost materials and technology. Can print in color. Un-set material provides

Relatively fast process. Safe materials.

Range of polymeric materials including nylon, elastomers, and composites. Strong

Polymeric materials—commonly nylon may be autoclaved. Printed object may have full mechanical functionality. Lower cost materials if used in large volume.

High strength objects can control porosity. Variety of materials including titanium, titanium alloys, cobalt chrome, stainless steel. Metal alloy may be recycled. Fine detail

High temperature process, so no support or heat treatment needed afterwards. High speed. Dense parts with controlled porosity. Only available with light curable liquid

Support materials must be removed. Resin is messy and can cause skin sensitization and may irritate by contact and inhalation. Limited shelf life and vat life. Cannot be heat sterilized. High cost technology.

Tenacious support material can be difficult to remove completely. Support material may cause skin irritation. Cannot be heat sterilized. High cost

Light curable liquid polymers and wax-like materials for casting. Support materials must be removed. Resin is messy and can cause skin sensitization, and may be irritant by contact Limited shelf life and vat life. Cannot be heat sterilized. Higher cost materials.

Low resolution. Messy powder. Low strength. Cannot be soaked or heat

Significant infrastructure required, e.g., Compressed air, climate control. Messy powders. Lower cost in bulk. Inhalation risk. High cost technology.

Elaborate infrastructure requirements. Extremely costly technology moderately costly materials. Dust and nanoparticle condensate may be hazardous to health. Explosive risk. Rough surface. Elaborate post-processing is required: Heat treatment to relieve internal stresses in printed objects. Hard to remove

polymers.

materials.

sterilized.

Rough surface.

support materials. Relatively slow process.

Extremely costly technology moderately costly materials. Dust may be hazardous to health. Explosive risk. Rough surface. Less post-processing required. Lower resolution.

**Light cured resin**

polymer.

1. Stereolithography (SLA)

Light sensitive polymer cured layer by layer by a scanning laser in a vat of liquid

190 Biomaterials in Regenerative Medicine

2. Photojet—light sensitive polymer is jetted onto a build platform from an inkjet type print head, and cured layer by layer on an incrementally descending platform.

3. DLP (digital light processing) Liquid resin is cured layer by layer by a projector light source. The object is built upside down on an incrementally

Plaster or cementaceous material set by drops of (colored) water from 'inkjet' print head. Object built layer by layer in a powder bed, on an incrementally

Selective laser sintering (SLS) for polymers. Object built layer by layer in powder bed. Heated build chamber raises temperature of material to just below melting point. Scanning laser then sinters powder layer by layer in a descending bed.

Selective laser sintering (SLS)—for metals and metal alloys. Also described as selective laser melting (SLM) or direct metal laser sintering (DMLS). Scanning laser sinters metal powder layer by layer in a cold build chamber as the build platform descends. Support structure used to tether

Electron beam melting (EBM, Arcam). Heated build chamber. Powder sintered layer by layer by scanning electron beam on descending build platform.

objects to build platform.

elevating platform.

**Powder binder**

descending platform.

**Sintered powder**


TCP, tri-calcium phosphate; HA, hydroxyapatite; DBM, demineralized bone matrix; PLGA, poly(lactic-co-glycolic acid); PLA, poly(d,l-lactic acid); PGA, poly(glycolic acid); PLLA, poly(l-lactic acid); PVDC, polyvinylidene chloride; PEG, polyethylene glycol; DA, diacrylate; MMP, matrix metalloproteinases; ABB, anorganic bovine bone; Puramatrix, a selfassembling peptide nanomaterial.

**Table 2.** Types of scaffolds used in bone tissue engineering in maxillo-craniofacial region [51].

just a whole bit of them are available for application in 3-D printing. As follows, we discuss four large categories of materials for 3-D printing of scaffolds and craniofacial tissues, which researches still aim to determine these materials complete properties and advantages.

microenvironment, also as mentioned before their ability to mechanically provide space maintenance, makes these materials interesting choice for 3-D scaffold fabrication for craniofacial applications. In contrast, ceramic scaffolds are not compatible with cell encapsulation for bioprinting. In 3-D printed ceramic scaffolds, cells quickly populate the scaffold surface, which establishing close cell-cell interactions lead to promotion of cell proliferation and differentiation. On the other hand, ceramics with properties lead to lower rates of degradation than hydrogels, which aids in prolonged guided tissue remodeling and structural support. In contrast, ceramic scaffolds are too brittle for implantation in load-bearing defect sites. Ideal scaffolds would combine the high calcium content of calcium and phosphate ceramics with the outstanding toughness of natural bone, which perhaps can only be obtained by creating scaffolds that are biomimetically mineral-

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193

Fused deposition of ceramics (FDC) in a direct printing mode generally consists of extruding a slurry including a high content (>50% w/v) of inorganic components [42]. The manufactur-

**1.** Mixture phase, which involves the preparation of the slurry. The bioceramic particles are mixed in a solvent (aqueous or nonaqueous) with a low concentration of organic poly-

**2.** Green ceramic and binder burnout phase involving the deposition of filaments of slurry following a predetermined pattern prior to drying and exposure to high temperature to

**3.** Sintering phase, which involves the exposure of the green form to elevated temperature (above 1000°C) to initiate the migration of atoms between adjacent ceramic particles, hence

It is critical for reproducible manufacturing of 3-D rapid prototyped bioceramics to have shape retention, a challenge that can be reached by adjusting the viscosity of the slurry and

Printable composites, which are usually in the form of copolymers, polymer-polymer mixtures, or polymer-ceramic mixtures [44], allow ability for the combination of variety of advantageous properties of their included components, which provide a remarkable candidate as "bioink". Considering the advantages of polymer composite hydrogels, such as interpenetrating polymer networks (IPNs) or hybrid hydrogels [45], the incorporation of synthetic fillers to printable materials recently discussed in researches [33]. The addition of silicate fillers [38] and a range of nanoparticles have been used to synthesize different types of composite scaffolds [46] to promote greater control over viscosity and stiffness of polymer hydrogels. In addition, several of silica-containing hydrogels with higher expression of genes encoding morphogenetic cytokines, such as bone morphogenetic proteins (BMPs) seems promising [47]. The combination and manufacturing mixture of hydrogels with filler materials and/or natural peptides with morphogenetic capacity demonstrate great future for application in 3-D

ized and hierarchically structured, as recent researches demonstrated that in [41].

mers/surfactants, called the binder, to obtain adequate flowability.

printing in aim to reach ultimate goal in regenerative craniofacial repair.

burn out the organic component of the mixture.

creating physical bonds called "necks."

the evaporation rate of the solvent [43].

**3.3. Composite materials**

ing of such scaffolds follows 3 steps:

#### **3.1. Polymers and hydrogels**

Polymer hydrogels are ideal candidates for the development of printable materials for tissue engineering. Hydrogels are known for remarkable tunability of rheological also presenting great mechanical, chemical, and biological properties; high biocompatibility; and similarity to native extracellular matrix (ECM) [32]. For three-dimensional printing of polymers and hydrogels, the use of materials with controlled viscosity should been noticed. This defines the range of printability of the ink. Polymer inks, which are typically printed in the prepolymer phase, need enough viscosity allowing structural support of subsequent printed layers, also enough fluidity to prevent nozzle clogging. For avoiding these difficulties, alginate hydrogels have been cross-linked with calcium ions immediately before the ink leaves the printing head or just after extrusions [33].

In recent researches, for providing suitable ink for bioprinting applications, prepolymerized cell-laden methacrylated gelatin hydrogels have been used successfully [34, 35]. Synthetic hydrogels used for cell encapsulation may limit cell-cell interactions. These interactions are critical for efficient cell proliferation, differentiation, and finally, tissue development. This can represent one of the limitations of bioprinting cell-laden hydrogels which is not present in 3-D printed scaffolds with cells seeded onto or in bioprinting of dense cell aggregates, which will discuss as follow. Hence, the requirement for the development of ECM-derived hydrogels that have tunable physical and chemical properties, are compatible with high cell viability, and provide the adequate binding sites (RGDs) for cell attachment and matrix remodeling during their early proliferative stage [32].

Synthetic polymers are most commonly used materials for 3-D printing in biomedical applications [36, 37]. However, since high temperature is usually involved during the printing of these materials, the direct incorporation of cells or growth factors in the polymer mixture is generally avoided as the cell viability or bioactivity [37] cannot be maintained throughout the manufacturing process.

Although hydrogels provide great advantages for tissue engineering applications, such as the ability of exposing cells to highly hydrated 3-D microenvironments that is similar to the natural ECM [32]. In contrast, they generally present very low stiffness (in the kPa range) compared with the majority of load-bearing tissues in the craniofacial complex (in the GPa range). Therefore, reconstruction of tissues subjected to higher mechanical loads, such as bones and teeth, usually requires the use of ceramic materials or composite scaffolds which provide more mechanical advantages, where polymers are commonly combined with inorganic fillers to increase scaffold stiffness [38].

#### **3.2. Ceramics**

Ceramic scaffolds are usually composed of calcium and phosphate mineral phases, such as hydroxyapatite [39] or b-tricalcium phosphate [40]. The noticeable ability of these scaffolds to upregulate osteogenesis due to inherent properties of the formation of a bioactive ion-rich cellular microenvironment, also as mentioned before their ability to mechanically provide space maintenance, makes these materials interesting choice for 3-D scaffold fabrication for craniofacial applications. In contrast, ceramic scaffolds are not compatible with cell encapsulation for bioprinting. In 3-D printed ceramic scaffolds, cells quickly populate the scaffold surface, which establishing close cell-cell interactions lead to promotion of cell proliferation and differentiation. On the other hand, ceramics with properties lead to lower rates of degradation than hydrogels, which aids in prolonged guided tissue remodeling and structural support. In contrast, ceramic scaffolds are too brittle for implantation in load-bearing defect sites. Ideal scaffolds would combine the high calcium content of calcium and phosphate ceramics with the outstanding toughness of natural bone, which perhaps can only be obtained by creating scaffolds that are biomimetically mineralized and hierarchically structured, as recent researches demonstrated that in [41].

Fused deposition of ceramics (FDC) in a direct printing mode generally consists of extruding a slurry including a high content (>50% w/v) of inorganic components [42]. The manufacturing of such scaffolds follows 3 steps:


It is critical for reproducible manufacturing of 3-D rapid prototyped bioceramics to have shape retention, a challenge that can be reached by adjusting the viscosity of the slurry and the evaporation rate of the solvent [43].

#### **3.3. Composite materials**

just a whole bit of them are available for application in 3-D printing. As follows, we discuss four large categories of materials for 3-D printing of scaffolds and craniofacial tissues, which

Polymer hydrogels are ideal candidates for the development of printable materials for tissue engineering. Hydrogels are known for remarkable tunability of rheological also presenting great mechanical, chemical, and biological properties; high biocompatibility; and similarity to native extracellular matrix (ECM) [32]. For three-dimensional printing of polymers and hydrogels, the use of materials with controlled viscosity should been noticed. This defines the range of printability of the ink. Polymer inks, which are typically printed in the prepolymer phase, need enough viscosity allowing structural support of subsequent printed layers, also enough fluidity to prevent nozzle clogging. For avoiding these difficulties, alginate hydrogels have been cross-linked with calcium ions immediately before the ink leaves the printing head or just after extrusions [33]. In recent researches, for providing suitable ink for bioprinting applications, prepolymerized cell-laden methacrylated gelatin hydrogels have been used successfully [34, 35]. Synthetic hydrogels used for cell encapsulation may limit cell-cell interactions. These interactions are critical for efficient cell proliferation, differentiation, and finally, tissue development. This can represent one of the limitations of bioprinting cell-laden hydrogels which is not present in 3-D printed scaffolds with cells seeded onto or in bioprinting of dense cell aggregates, which will discuss as follow. Hence, the requirement for the development of ECM-derived hydrogels that have tunable physical and chemical properties, are compatible with high cell viability, and provide the adequate binding sites (RGDs) for cell attachment and matrix remodeling

Synthetic polymers are most commonly used materials for 3-D printing in biomedical applications [36, 37]. However, since high temperature is usually involved during the printing of these materials, the direct incorporation of cells or growth factors in the polymer mixture is generally avoided as the cell viability or bioactivity [37] cannot be maintained throughout the

Although hydrogels provide great advantages for tissue engineering applications, such as the ability of exposing cells to highly hydrated 3-D microenvironments that is similar to the natural ECM [32]. In contrast, they generally present very low stiffness (in the kPa range) compared with the majority of load-bearing tissues in the craniofacial complex (in the GPa range). Therefore, reconstruction of tissues subjected to higher mechanical loads, such as bones and teeth, usually requires the use of ceramic materials or composite scaffolds which provide more mechanical advantages, where polymers are commonly combined with inorganic fillers to increase scaffold

Ceramic scaffolds are usually composed of calcium and phosphate mineral phases, such as hydroxyapatite [39] or b-tricalcium phosphate [40]. The noticeable ability of these scaffolds to upregulate osteogenesis due to inherent properties of the formation of a bioactive ion-rich cellular

researches still aim to determine these materials complete properties and advantages.

**3.1. Polymers and hydrogels**

192 Biomaterials in Regenerative Medicine

during their early proliferative stage [32].

manufacturing process.

stiffness [38].

**3.2. Ceramics**

Printable composites, which are usually in the form of copolymers, polymer-polymer mixtures, or polymer-ceramic mixtures [44], allow ability for the combination of variety of advantageous properties of their included components, which provide a remarkable candidate as "bioink". Considering the advantages of polymer composite hydrogels, such as interpenetrating polymer networks (IPNs) or hybrid hydrogels [45], the incorporation of synthetic fillers to printable materials recently discussed in researches [33]. The addition of silicate fillers [38] and a range of nanoparticles have been used to synthesize different types of composite scaffolds [46] to promote greater control over viscosity and stiffness of polymer hydrogels. In addition, several of silica-containing hydrogels with higher expression of genes encoding morphogenetic cytokines, such as bone morphogenetic proteins (BMPs) seems promising [47]. The combination and manufacturing mixture of hydrogels with filler materials and/or natural peptides with morphogenetic capacity demonstrate great future for application in 3-D printing in aim to reach ultimate goal in regenerative craniofacial repair.

#### **3.4. Cell aggregates and spheroids**

Over recent years, many of researches aimed to evaluate and study cell aggregates and spheroids for use in tissue engineering and regenerative medicine [48]. As this method cited correctly and appropriately as "scaffold-free printing," in fact small quantities of hydrogel are used to facilitate cell aggregation. In this method for 3-D printing, or in an appropriate way called "bioprinting," multicellular spheroids are deposited using extrusion printers and allowed to self-assemble into the desired 3-D structure (**Figure 1**). As it is clear, these systems allow direct fabrication of tissue constructs which in contrast to other methods have extremely high cell densities. Although in load-bearing tissues with high amount of mineral components and noticeable mechanical properties use of this methods still looks uncertain, the ability to position aggregates of heterotypic cells with microscale precision (**Figure 2**) seems promising

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195

Researches aimed to investigate novel technologies for 3-D printing and introduced some novel methods including phase-separation, self-assembly, electrospinning, freeze drying, solvent casting/particulate leaching, gas foaming, and melt molding [52]. Using scaffolds, the architecture of native extracellular matrices can be mimicked at the nanoscale level and therefore provide the primary base for the regeneration of new tissue [53]. Originally, a "top-down" approach was used as a tissue engineering method for scaffold fabrication. In this method, cells are seeded onto a biodegradable and biocompatible scaffold and are predicted to migrate and fill the scaffold hence creating their own matrix. By using this technique, several avascular tissues such as bladder [54] and skin [55] have been engineered effectively. However, due to the limited diffusion properties of these scaffolds, this technique faces several challenges for fabrication of more complex tissues such as heart and liver [56]. Therefore, "bottom-up" methods have been developed to overcome this problem [57]. Bottom-up approaches include cell-encapsulation with microscale hydrogels, cell aggregation by self-assembly, generation of cell sheets, and direct printing of cells [58]. These complex tissue blocks can be assembled using various methods including microfluidics [59], magnetic fields [60], acoustic fields [61], and surface tension [62]. These methods are relatively easy and have provided a solid foundation for the fabrication of scaffolds. However, as mentioned previously, these conventional methods suffer from several limitations including inadequate control over scaffold properties such as pore size, pore geometry, distribution of high levels of interconnectivity, and mechanical strength. As such, it is necessary to develop technologies with sufficient control so as to design more intricate tissue-specific scaffolds. In addition, scaffolds can be coated using surface modification techniques (such as introducing functional groups) to enhance cell migration, attachment and proliferation. Three-dimensional printing allows scaffolds to become more precisely fabricated (similar to that of the computer-aided design (CAD)) with higher flexibility in the type of materials used to make such scaffolds. Three-dimensional printing uses an additive manufacturing process where a structure is fabricated using a layerby-layer process. Materials deposited for the formation of the scaffold may be cross-linked or polymerized through heat, ultraviolet light, or binder solutions. Using this technology, 3-D

as an excellent alternative to bioprint complex tissues consisting variety of cells [49].

**4. Manufacturing of scaffolds with 3-D printing technology**

printed scaffolds can be prepared for optimized tissue engineering [52].

For appropriate formation of tissue architecture, the seeding cells (often stem cells) require a 3-D environment/matrix similar to that of the ECM. The ECM acts as a medium to provide proteins and proteoglycans among other nutrients for cellular growth. The

**Figure 1.** SEM view of multicellular spheroids of HUSMCs (A), CHO cells (B) and HFBs (C*)* (adapted from Norotte et al. [50]).

**Figure 2.** Principles of spheroids bioprinting technology: (a) bioprinter (general view); (b) multiple bioprinter nozzles; (c) tissue spheroids before dispensing; (d) tissue spheroids during dispensing; (e) schematic view of continuous dispensing in air; (f) schematic view of continuous dispensing in fluid; (g) schematic view of digital dispensing in air; (h) schematic view of digital dispensing in fluid; (i) schematic view of bioassembly of tubular tissue construct using bioprinting of selfassembled tissue spheroids illustrating sequential steps of layer-by-layer tissue spheroid deposition and tissue fusion process (adopted from Mironov et al. [48]).

way called "bioprinting," multicellular spheroids are deposited using extrusion printers and allowed to self-assemble into the desired 3-D structure (**Figure 1**). As it is clear, these systems allow direct fabrication of tissue constructs which in contrast to other methods have extremely high cell densities. Although in load-bearing tissues with high amount of mineral components and noticeable mechanical properties use of this methods still looks uncertain, the ability to position aggregates of heterotypic cells with microscale precision (**Figure 2**) seems promising as an excellent alternative to bioprint complex tissues consisting variety of cells [49].

## **4. Manufacturing of scaffolds with 3-D printing technology**

**Figure 1.** SEM view of multicellular spheroids of HUSMCs (A), CHO cells (B) and HFBs (C*)* (adapted from Norotte et al. [50]).

Over recent years, many of researches aimed to evaluate and study cell aggregates and spheroids for use in tissue engineering and regenerative medicine [48]. As this method cited correctly and appropriately as "scaffold-free printing," in fact small quantities of hydrogel are used to facilitate cell aggregation. In this method for 3-D printing, or in an appropriate

**Figure 2.** Principles of spheroids bioprinting technology: (a) bioprinter (general view); (b) multiple bioprinter nozzles; (c) tissue spheroids before dispensing; (d) tissue spheroids during dispensing; (e) schematic view of continuous dispensing in air; (f) schematic view of continuous dispensing in fluid; (g) schematic view of digital dispensing in air; (h) schematic view of digital dispensing in fluid; (i) schematic view of bioassembly of tubular tissue construct using bioprinting of selfassembled tissue spheroids illustrating sequential steps of layer-by-layer tissue spheroid deposition and tissue fusion

process (adopted from Mironov et al. [48]).

**3.4. Cell aggregates and spheroids**

194 Biomaterials in Regenerative Medicine

Researches aimed to investigate novel technologies for 3-D printing and introduced some novel methods including phase-separation, self-assembly, electrospinning, freeze drying, solvent casting/particulate leaching, gas foaming, and melt molding [52]. Using scaffolds, the architecture of native extracellular matrices can be mimicked at the nanoscale level and therefore provide the primary base for the regeneration of new tissue [53]. Originally, a "top-down" approach was used as a tissue engineering method for scaffold fabrication. In this method, cells are seeded onto a biodegradable and biocompatible scaffold and are predicted to migrate and fill the scaffold hence creating their own matrix. By using this technique, several avascular tissues such as bladder [54] and skin [55] have been engineered effectively. However, due to the limited diffusion properties of these scaffolds, this technique faces several challenges for fabrication of more complex tissues such as heart and liver [56]. Therefore, "bottom-up" methods have been developed to overcome this problem [57]. Bottom-up approaches include cell-encapsulation with microscale hydrogels, cell aggregation by self-assembly, generation of cell sheets, and direct printing of cells [58]. These complex tissue blocks can be assembled using various methods including microfluidics [59], magnetic fields [60], acoustic fields [61], and surface tension [62]. These methods are relatively easy and have provided a solid foundation for the fabrication of scaffolds. However, as mentioned previously, these conventional methods suffer from several limitations including inadequate control over scaffold properties such as pore size, pore geometry, distribution of high levels of interconnectivity, and mechanical strength. As such, it is necessary to develop technologies with sufficient control so as to design more intricate tissue-specific scaffolds. In addition, scaffolds can be coated using surface modification techniques (such as introducing functional groups) to enhance cell migration, attachment and proliferation. Three-dimensional printing allows scaffolds to become more precisely fabricated (similar to that of the computer-aided design (CAD)) with higher flexibility in the type of materials used to make such scaffolds. Three-dimensional printing uses an additive manufacturing process where a structure is fabricated using a layerby-layer process. Materials deposited for the formation of the scaffold may be cross-linked or polymerized through heat, ultraviolet light, or binder solutions. Using this technology, 3-D printed scaffolds can be prepared for optimized tissue engineering [52].

For appropriate formation of tissue architecture, the seeding cells (often stem cells) require a 3-D environment/matrix similar to that of the ECM. The ECM acts as a medium to provide proteins and proteoglycans among other nutrients for cellular growth. The ECM also provides structural support to allow for cellular functionality such as regulating cellular communication, growth, and assembly [63]. With this in mind, scientists and engineers originally attempted to replicate the ECM through conventional techniques, which consequently established a framework for using more advanced techniques, such as 3-D printing, to yield higher quality scaffolds. The 3-D printing technique can create defined scaffold structures with controlled pore size and interconnectivity and the ability to support cell growth and tissue formation [64–66]. The current methods for 3-D printing involve a CAD, which is then relayed to each 3-D printing system to "print" the desired scaffold structure. Through various 3-D printing technologies, discussed below, researchers are trying to fabricate biocompatible scaffolds that efficiently support tissue formation (**Table 3**).

## **5. Bioprinting advantages aiming for clinical use**

The goal of tissue engineering is to create functional tissues and organs for regenerative therapies and ultimately organ transplantation/replacement. Trial and error was the long and tedious process mainly used to advance the field of regenerative medicine by clarifying the success of techniques.

Researchers needed to come up with a list of requirements in order to measure their successes or failures in tissue fabrication [48, 67]. This list was generated from the observations of natural human tissue.

As gold standard of fabricated tissues is to be as similar as possible to natural tissues in the human body in different parameters, then these fabricated tissues must:


Also, the printers used for tissue fabrication required standardization as well [67, 69].


**Printing method**

Direct 3-D printing/

• • plex structures

Versatile in terms of usable materials

No support is necessary for overhang or com-

inkjet

W/electrospinning

Bioplotting

• •

Fused deposition

• •

Low cytotoxicity vs direct 3-D printing

• • • shapes

• • • resistant)

• • processing)

Very high temp required (up to 1400°C)

Expensive and time consuming (processing and post

Limitation on materials (must be shrinkage and heat

• • • • •

(Mouse/heart)

197

(Mouse/skin)

(Rat/bone)

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(Rat/heart)

(Mouse/bone)

Application of 3-D Printing for Tissue Regeneration in Oral and Maxillofacial...

Selective laser sintering

 • strength

• structures

•

Fine resolution

Powder bed provides support for complex

Provides scaffolds with high mechanical

Low Resolution

Post-processing may be necessary

Materials used are nonbiodegradable

Requires support structure for overhangs and complex

Limitation on materials (often requires thermoplastics)

• •

(Rat/bone)

(Swine/bone)

Relatively inexpensive (printers and materials)

modeling

Soft tissue applications

Prints viable cells

• ing processing)

•

Requires support structure for printing complex shapes

Limitation on nozzle size\* (\*Must not be cytotoxic dur-

• • • • •

(Mouse/tooth

regeneration)

•

(Mouse/skin)

(Mouse/cartilage)

(Rat/cartilage)

(Rabbit/cartilage)

(Rabbit/trachea)

**Advantages**

**Disadvantages**

• • sintering

•

Time consuming (post-processing)

Potential toxicity (incompletely removed binders)

Low mechanical strength prints compared to laser

**Preclinical progress**

• • • •

(Mouse/cartilage)

(Mouse/bone)

(Rabbit/bone)

(Rat/bone)


ECM also provides structural support to allow for cellular functionality such as regulating cellular communication, growth, and assembly [63]. With this in mind, scientists and engineers originally attempted to replicate the ECM through conventional techniques, which consequently established a framework for using more advanced techniques, such as 3-D printing, to yield higher quality scaffolds. The 3-D printing technique can create defined scaffold structures with controlled pore size and interconnectivity and the ability to support cell growth and tissue formation [64–66]. The current methods for 3-D printing involve a CAD, which is then relayed to each 3-D printing system to "print" the desired scaffold structure. Through various 3-D printing technologies, discussed below, researchers are trying to fabricate biocompatible scaffolds that efficiently support tissue formation

The goal of tissue engineering is to create functional tissues and organs for regenerative therapies and ultimately organ transplantation/replacement. Trial and error was the long and tedious process mainly used to advance the field of regenerative medicine by clarifying the

Researchers needed to come up with a list of requirements in order to measure their successes or failures in tissue fabrication [48, 67]. This list was generated from the observations of natu-

As gold standard of fabricated tissues is to be as similar as possible to natural tissues in the

**1.** Be able to integrate with naturally occurring tissue, and attach via microsutures, glues [68],

human body in different parameters, then these fabricated tissues must:

**3.** Become fully vascularized in order to sustain its functionality [68, 71].

Also, the printers used for tissue fabrication required standardization as well [67, 69].

**1.** The bioprinting machines required set extreme sterilization methods to eliminate unwarranted contamination with previously used materials or foreign matter from the

**2.** The conditions for printing must be ideal for tissue fabrication, so factors such as humidity

**3.** Nozzle size and methods of delivery affect the viability of the materials being printed; therefore, there must be set ideals for delivery methods in relation to the various printing

**5. Bioprinting advantages aiming for clinical use**

(**Table 3**).

success of techniques.

196 Biomaterials in Regenerative Medicine

or through cell adhesion [69–71].

**2.** Be capable of essential functions in vivo [48].

and temperature must be closely monitored.

ral human tissue.

environment.

materials.


**Table 3.** Preclinical researches on various 3-D printing techniques for manufacturing scaffolds for tissue engineering [52]. As a result, researchers created a few methods of printing with the goal of finding a solution to the given problems for optimal tissue biofabrication [48, 68, 69]. Thermal inkjet bioprinting with bioink and direct-write bioprinting both make use of modified inkjet printers but with varied application techniques. Organ printing with tissue spheroids is the recent achievement

Application of 3-D Printing for Tissue Regeneration in Oral and Maxillofacial...

and disadvantages of all three common methods "Thermal Inkjet Bioprinting," "Direct-Write Bioprinting," and "Spheroid Organ Printing." Organ printing, otherwise known as the biomedical

> • Avascular • Aneural • Alymphatic • Thin

• Only nourishable via diffusion

• Self-organization is defined as, "a process in which patterning at the global level of a system emerges solely from numerous interactions among the lower-level components of the system."

• Self-assembly is defined to be, "the autonomous organization of com

> ponents into patterns or structures without human intervention."


Possibility of printing tissues with different compositions.

**4** review advantages

"Bioink," which is a water-based liquid consisting of proteins, enzymes, and cells suspended in

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• The bioink of direct-write printers may consist of hydro

• This method categorized in pneumatic, mechanical, and a pneumatic-mechanical hybrid. It was concluded that the pneumatic systems work better with high viscosity materials, while mechanical systems are better suited in working with materials of

Pluronic F127.

low viscosity.

spheroids.

Researchers fabricated three types of spheroids to create a vascular tree: solid or nonlumenized spheroids, spheroids with one big lumen (mono-lumenized spheroids), and microvascularized tissue

gels of varying consistencies that are composed of agarose, alginate, collagen type I, and


199

a media or saline.

of researches which seems promising to fabricate tissues directly. **Table**

**Method Tissue characteristics Note**

**Type of bioprinting**

Thermal inkjet bioprinting

Direct-write bioprinting

Spheroid organ printing • Bottom up • Layer-by-layer

• Digital control of print. • Several printing units simultaneously. • Application of variety of materials simultaneously. • Faster turnaround time for printed products.

Spheroids are punched into "biopaper" which is a sprayed layer of hydrogel. Each spheroid is made of living cells, thereby creating a ball of "living materials" capable of selfassembly and self-fusion. Alternatively, the spheroids can be digitally placed, undergo selfassembly, and fuse without the

**Table 4.** 3-D bioprinting technique advantages and properties[67].

use of hydrogel.

As a result, researchers created a few methods of printing with the goal of finding a solution to the given problems for optimal tissue biofabrication [48, 68, 69]. Thermal inkjet bioprinting with bioink and direct-write bioprinting both make use of modified inkjet printers but with varied application techniques. Organ printing with tissue spheroids is the recent achievement of researches which seems promising to fabricate tissues directly. **Table 4** review advantages and disadvantages of all three common methods "Thermal Inkjet Bioprinting," "Direct-Write Bioprinting," and "Spheroid Organ Printing." Organ printing, otherwise known as the biomedical


**Printing method**

Stereolithography

• • •

Electrospinning

• • • •

Indirect 3-D printing

• • obtained

Good for prototyping/preproduction

• printing)

• • • product)

**Table 3.**

Preclinical researches on various 3-D printing techniques for manufacturing scaffolds for tissue engineering [52].

Long production times (mold

→ cast

→ processing

→

Mold required for casting

Low accuracies/resolution

Requires proprietary waxes for biocompatibility (wax

• •

(Mouse/tooth

regeneration)

(Rat/bone)

Material versatility casting once mold is

Low shear stress (bioelectrospraying)

Soft tissue engineering

Cell printing

Speed of fabrication

Smooth surface finish

Speed of fabrication

Very high resolution

**Advantages**

**Disadvantages**

• • • objects

• • •

High voltage (1–30 kV) requirements

Nonuniform pore sizes

Random orientation of fibers

Support system is necessary for overhang and intricate

Expensive (two photon printers)

Materials must be photopolymers

**Preclinical progress**

• • • •

(Mouse/

biocompatibility)

• •

(Rabbit/vascular

tissue)

(Rat/bone)

(Pig/tendon)

198 Biomaterials in Regenerative Medicine

(Rabbit/trachea)

(Rat/bone)

application of rapid prototyping, may be defined as additive layer-by-layer biomanufacturing of cells. Advantages of organ printing include its automated approach offering a pathway for a scalable and reproducible mass production of tissue-engineered products. This also allows the precise simultaneous 3-D positioning of several cell types, hence enabling the creation of tissue with a high level of cell density. Organ printing may be used to solve the problem of vascularization in thick tissue constructs, and moreover, this technology may be done in situ. Therefore, this emerging transforming technology has potential for surpassing traditional solid scaffold-based tissue engineering [72].

**7. Future aspects of 3-D printing for regenerative medicine**

tissue constructs continues to be a challenge.

exciting area.

In this chapter, we have illustrated current guiding principles for 3-D bioprinting in tissue fabrication, as well as recent advances and technological developments. The speed at which our knowledge has advanced with additive manufacturing and automated printing systems shows a promise to expand our basic science and engineering capabilities toward addressing health care problems. One of the significant developments in 3-D bioprinting is to manufacture cell microenvironments from molecular to macroscopic scales, which are requested and suitable for tissue engineering and regenerative medicine. As novel methods and technologies introduced in recent years for 3-D printing of biomaterials, promising overview of future appears to manufacture scaffolds for tissue engineering that reach the gold standards and also better comprehensions of stem cells microenvironments and interactions. By aid of various novel technologies, such as microfluidic systems [75, 76], biopatterning [77], and layer-by-layer assembly [76, 78], researchers are now able to biomanufacture microtissue constructs within scaffolds and even also within scaffold-free environments. Considering the great and enormous improvements of biomaterial for tissue engineering, in contrast, there are still certain challenges and difficulties that need more attention. Vascularization is one of the limitations which receive most of attentions [79, 80] due to the fact that this challenge leads to hypoxia, apoptosis, and immediate cell death. For resolving this issue and providing sufficient space for vascularization, researchers attempts to fabricate porous scaffolds [81], to provide sufficient space for vascularization. However, this approach cannot overcome the vascularization challenge completely due to the diffusion of cells and other materials into these porous structures [82]. Forming interconnected, well-defined vascular structures during biomanufacturing process seems to lead to resolving this difficulty and providing better results during process. Other issues that have to be noticed are mechanical strength and stability in 3-D tissue engineering which is one of the key requirements [83]. To be clear in regeneration of hard (e.g., bone) and soft (e.g., vascular grafts) tissues, modulus of elasticity is a crucial parameter that desires improvement [84–86]. Furthermore, the development of a totally closed bioprinting system that integrates printing and post-printing processes such as in-vitro culture and maturation of

Application of 3-D Printing for Tissue Regeneration in Oral and Maxillofacial...

http://dx.doi.org/10.5772/intechopen.70323

201

With advances in near future, which help finding solutions for the challenges mentioned above, bioprinting technologies will potentially help improvements of rapid clinical solutions and advances in medical implants. Further, we envision that the integration of cells and biomaterials through bioprinting with microfluidic technologies are likely to create unique microenvironments for various applications in cancer biology, tissue engineering, and regenerative medicine [87–91]. Additionally, developments on high-throughput biomanufacturing of 3-D architectures will pave the way for further advancements of in vitro screening and diagnostic applications, potentially enabling complex organ constructs. In the meantime, it is only the effective interplay of engineering concepts in combination with the well-established fundamentals of biology that will realize the true potential of this

## **6. Current limitations**

#### **6.1. Vascularization**

In order to create a complete and functional organ, the researchers must be able to create thick complex tissues with full vascularization containing lumens of various sizes, large vascular structures to microstructures, in order to sustain the surrounding organ tissue. The best way to achieve this type of vascularization is to fabricate the vascular system and tissue simultaneously, of which is easier said than done [48]. Thorough vascularization remains a common theme for current bioprinting limitations. Without a functional circulatory system, tissue constructs are limited to a means of diffusion for nutrition, which in itself is limited to just a few hundred microns [69].

Current methods of vascularization call for the infiltration of host microvessels into an implanted construct [67, 73, 74].

Yet, this strategy is lacking in control and specificity for the developing microvessels. The invading microvessels have a limited penetration depth which prevents the successful incorporation of the microvessels into larger layered constructs. Additionally, the penetration of the vascular system itself may result in a distortion of the region penetrated or in the destruction of the fabricated tissue altogether. For these reasons, it would be ideal to construct tissues with direct vessel in-growth, or vascularization created within the tissue itself, all before implantation.

#### **6.2. Tissue components and costs**

In addition to vascularization, native tissues contain unique cellular combinations and organizations. There is a need to develop techniques that mimic the complexity of native tissues in order to drive tissue recovery and replacement for medical applications [69]. With the production of organs such as kidneys, for example, at least one million glomeruli and nephrons would need to be generated. Not only would the fabrication be a massive undertaking but also the fabricated tissue would need to be scalable. Scalability of biofabricated tissues is not presently a reality. Yet, spheroids have shown promise toward being scalable with further development. Finally, another major limitation for the development of natural-like, fully functioning fabricated human tissue is economic [68]. This challenge must definitely be overcome if biofabrication technology is to allow the creation of a functional living human organ.

## **7. Future aspects of 3-D printing for regenerative medicine**

application of rapid prototyping, may be defined as additive layer-by-layer biomanufacturing of cells. Advantages of organ printing include its automated approach offering a pathway for a scalable and reproducible mass production of tissue-engineered products. This also allows the precise simultaneous 3-D positioning of several cell types, hence enabling the creation of tissue with a high level of cell density. Organ printing may be used to solve the problem of vascularization in thick tissue constructs, and moreover, this technology may be done in situ. Therefore, this emerging transforming technology has potential for surpassing traditional solid scaffold-based

In order to create a complete and functional organ, the researchers must be able to create thick complex tissues with full vascularization containing lumens of various sizes, large vascular structures to microstructures, in order to sustain the surrounding organ tissue. The best way to achieve this type of vascularization is to fabricate the vascular system and tissue simultaneously, of which is easier said than done [48]. Thorough vascularization remains a common theme for current bioprinting limitations. Without a functional circulatory system, tissue constructs are limited to a means of diffusion for nutrition, which in itself is limited to just a few hundred microns [69].

Current methods of vascularization call for the infiltration of host microvessels into an

Yet, this strategy is lacking in control and specificity for the developing microvessels. The invading microvessels have a limited penetration depth which prevents the successful incorporation of the microvessels into larger layered constructs. Additionally, the penetration of the vascular system itself may result in a distortion of the region penetrated or in the destruction of the fabricated tissue altogether. For these reasons, it would be ideal to construct tissues with direct vessel in-growth, or vascularization created within the tissue itself, all before implantation.

In addition to vascularization, native tissues contain unique cellular combinations and organizations. There is a need to develop techniques that mimic the complexity of native tissues in order to drive tissue recovery and replacement for medical applications [69]. With the production of organs such as kidneys, for example, at least one million glomeruli and nephrons would need to be generated. Not only would the fabrication be a massive undertaking but also the fabricated tissue would need to be scalable. Scalability of biofabricated tissues is not presently a reality. Yet, spheroids have shown promise toward being scalable with further development. Finally, another major limitation for the development of natural-like, fully functioning fabricated human tissue is economic [68]. This challenge must definitely be overcome if biofabrication technology is to allow the creation of a functional living human organ.

tissue engineering [72].

200 Biomaterials in Regenerative Medicine

**6.1. Vascularization**

**6. Current limitations**

implanted construct [67, 73, 74].

**6.2. Tissue components and costs**

In this chapter, we have illustrated current guiding principles for 3-D bioprinting in tissue fabrication, as well as recent advances and technological developments. The speed at which our knowledge has advanced with additive manufacturing and automated printing systems shows a promise to expand our basic science and engineering capabilities toward addressing health care problems. One of the significant developments in 3-D bioprinting is to manufacture cell microenvironments from molecular to macroscopic scales, which are requested and suitable for tissue engineering and regenerative medicine. As novel methods and technologies introduced in recent years for 3-D printing of biomaterials, promising overview of future appears to manufacture scaffolds for tissue engineering that reach the gold standards and also better comprehensions of stem cells microenvironments and interactions. By aid of various novel technologies, such as microfluidic systems [75, 76], biopatterning [77], and layer-by-layer assembly [76, 78], researchers are now able to biomanufacture microtissue constructs within scaffolds and even also within scaffold-free environments. Considering the great and enormous improvements of biomaterial for tissue engineering, in contrast, there are still certain challenges and difficulties that need more attention. Vascularization is one of the limitations which receive most of attentions [79, 80] due to the fact that this challenge leads to hypoxia, apoptosis, and immediate cell death. For resolving this issue and providing sufficient space for vascularization, researchers attempts to fabricate porous scaffolds [81], to provide sufficient space for vascularization. However, this approach cannot overcome the vascularization challenge completely due to the diffusion of cells and other materials into these porous structures [82]. Forming interconnected, well-defined vascular structures during biomanufacturing process seems to lead to resolving this difficulty and providing better results during process. Other issues that have to be noticed are mechanical strength and stability in 3-D tissue engineering which is one of the key requirements [83]. To be clear in regeneration of hard (e.g., bone) and soft (e.g., vascular grafts) tissues, modulus of elasticity is a crucial parameter that desires improvement [84–86]. Furthermore, the development of a totally closed bioprinting system that integrates printing and post-printing processes such as in-vitro culture and maturation of tissue constructs continues to be a challenge.

With advances in near future, which help finding solutions for the challenges mentioned above, bioprinting technologies will potentially help improvements of rapid clinical solutions and advances in medical implants. Further, we envision that the integration of cells and biomaterials through bioprinting with microfluidic technologies are likely to create unique microenvironments for various applications in cancer biology, tissue engineering, and regenerative medicine [87–91]. Additionally, developments on high-throughput biomanufacturing of 3-D architectures will pave the way for further advancements of in vitro screening and diagnostic applications, potentially enabling complex organ constructs. In the meantime, it is only the effective interplay of engineering concepts in combination with the well-established fundamentals of biology that will realize the true potential of this exciting area.

## **Author details**

Seied Omid Keyhan1,2,3,4, Hamidreza Fallahi<sup>5</sup> , Alireza Jahangirnia<sup>6</sup> , Mohammad Taher Amirzade-Iranaq<sup>7</sup> and Mohammad Hosein Amirzade-Iranaq8,9,10,11\*

\*Address all correspondence to: h.amirzade@gmail.com

1 Department of Oral & Maxillofacial Surgery, Faculty of Dentistry, Birjand University of Medical Science, Birjand, Iran

[4] Hoy MB. 3D printing: Making things at the library. Medical Reference Services Quarterly.

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2 Vice Presidential Organization of Technology of the Islamic Republic of Iran, Iran

3 Stem Cell & Regenerative Medicine Network, Shahid Beheshti University of Medical Sciences, Tehran, Iran

4 Cranio Maxillofacial Research Center, Tehran Dental Branch, Islamic Azad University, Tehran, Iran

5 Oral and Maxillofacial Surgery, Jundishapur University of Medical Sciences, Ahvaz, Iran

6 Oral and Maxillofacial Surgeon, Private Practice, Tehran, Iran

7 Biomaterial Medical Engineer, Department of Materials, Faculty of Materials Engineering, Islamic Azad University Najaf Abad Branch, Isfahan, Iran

8 Student Research Committee (SRC), Baqiyatallah University of Medical Sciences, Tehran, Iran

9 Universal Network of Interdisciplinary Research in Oral and Maxillofacial Surgery (UNIROMS), Universal Scientific Education and Research Network (USERN), Tehran, Iran

10 Student Research Committee, Shahid Sadoughi University of Medical Sciences, Yazd, Iran

11 International Otorhinolaryngology Research Association (IORA), Universal Scientific Education and Research Network (USERN), Tehran, Iran

## **References**


[4] Hoy MB. 3D printing: Making things at the library. Medical Reference Services Quarterly. 2013;**32**(1):93-99

**Author details**

202 Biomaterials in Regenerative Medicine

Tehran, Iran

Tehran, Iran

Tehran, Iran

Yazd, Iran

**References**

2014;**9**(4):535-549

Seied Omid Keyhan1,2,3,4, Hamidreza Fallahi<sup>5</sup>

\*Address all correspondence to: h.amirzade@gmail.com

6 Oral and Maxillofacial Surgeon, Private Practice, Tehran, Iran

Islamic Azad University Najaf Abad Branch, Isfahan, Iran

Education and Research Network (USERN), Tehran, Iran

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Mohammad Taher Amirzade-Iranaq<sup>7</sup>

Medical Science, Birjand, Iran

, Alireza Jahangirnia<sup>6</sup>

1 Department of Oral & Maxillofacial Surgery, Faculty of Dentistry, Birjand University of

3 Stem Cell & Regenerative Medicine Network, Shahid Beheshti University of Medical Sciences,

4 Cranio Maxillofacial Research Center, Tehran Dental Branch, Islamic Azad University,

5 Oral and Maxillofacial Surgery, Jundishapur University of Medical Sciences, Ahvaz, Iran

7 Biomaterial Medical Engineer, Department of Materials, Faculty of Materials Engineering,

8 Student Research Committee (SRC), Baqiyatallah University of Medical Sciences,

11 International Otorhinolaryngology Research Association (IORA), Universal Scientific

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[61] Xu F, Finley TD, Turkaydin M, Sung Y, Gurkan UA, Yavuz AS, et al. The assembly of cell-encapsulating microscale hydrogels using acoustic waves. Biomaterials.

[62] Kachouie NN, Du Y, Bae H, Khabiry M, Ahari AF, Zamanian B, et al. Directed assembly of cell-laden hydrogels for engineering functional tissues. Organogenesis. 2010;**6**(4):234-244

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[79] Rupnick MA, Panigrahy D, Zhang C-Y, Dallabrida SM, Lowell BB, Langer R, et al. Adipose tissue mass can be regulated through the vasculature. Proceedings of the National Academy of Sciences. 2002;**99**(16):10730-10735

**Chapter 8**

**Provisional chapter**

**Tissue Engineering: Use of Growth Factors in Bone**

Tissue healing is a complex process involving a cascade of cellular and molecular events that are mostly shared by the different tissues of the body. Interestingly, the tissue repair process initiates immediately after a traumatic injury and is mediated and controlled by a wide range of cytokines, proteins, and growth factors released from platelets upon activation. Consequently, many growth factors have been considered as therapeutic molecules for the repair or regeneration of a wide range of tissues. Although their role has been only partially elucidated, the potential benefit of most growth factors has been demonstrated. In the last few years, the development of platelet-rich preparations has revolutionized the field of regenerative medicine, due to the repair capacities of the platelet-released growth factors that stimulate and accelerate both soft and hard tissue healing and regeneration. Today, autologous platelet concentrates (APCs) are used in a wide range of disciplines such as dentistry, oral surgery, orthopedics, sport medicine, dermatology, and plastic and reconstructive surgery. The purpose of this chapter is to describe the current evidence regarding the benefits of using autologous platelet concentrates in various oral surgery proce-

**Keywords:** platelet concentrates, growth factors, tissue regeneration, oral surgery

**Tissue Engineering: Use of Growth Factors in Bone** 

DOI: 10.5772/intechopen.69875

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Autologous platelet concentrates (APCs) have been widely used in many different clinical situations that require a rapid tissue healing and regeneration as it is especially the case in oral and maxillofacial surgery, orthopedics, sports medicine, ophthalmology, and in the treatment of skin ulcers. APCs are hemocomponents, obtained through centrifugation of patient's own blood, in order to collect the most active blood components: platelets, fibrin,

Carmen Mortellaro and Massimo Del Fabbro

Additional information is available at the end of the chapter

Carmen Mortellaro and Massimo Del Fabbro

Additional information is available at the end of the chapter

dures, using a systematic review approach.

http://dx.doi.org/10.5772/intechopen.69875

**Regeneration**

**Abstract**

**1. Introduction**

**Regeneration**


**Provisional chapter**

## **Tissue Engineering: Use of Growth Factors in Bone Regeneration Regeneration**

**Tissue Engineering: Use of Growth Factors in Bone** 

DOI: 10.5772/intechopen.69875

Carmen Mortellaro and Massimo Del Fabbro Carmen Mortellaro and Massimo Del Fabbro

Additional information is available at the end of the chapter Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69875

#### **Abstract**

[79] Rupnick MA, Panigrahy D, Zhang C-Y, Dallabrida SM, Lowell BB, Langer R, et al. Adipose tissue mass can be regulated through the vasculature. Proceedings of the

[80] Langer RS, Vacanti JP. Tissue engineering: The challenges ahead. Scientific American.

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[83] Hockaday L, Kang K, Colangelo N, Cheung P, Duan B, Malone E, et al. Rapid 3D printing of anatomically accurate and mechanically heterogeneous aortic valve hydrogel

[84] Mironov V, Boland T, Trusk T, Forgacs G, Markwald RR. Organ printing: Computeraided jet-based 3D tissue engineering. TRENDS in Biotechnology. 2003;**21**(4):157-161 [85] Prestwich GD. Hyaluronic acid-based clinical biomaterials derived for cell and molecule delivery in regenerative medicine. Journal of Controlled Release. 2011;**155**(2):193-199 [86] Guven S, Lindsey JS, Poudel I, Chinthala S, Nickerson MD, Gerami-Naini B, et al. Functional maintenance of differentiated embryoid bodies in microfluidic systems: A platform for personalized medicine. Stem Cells Translational Medicine. 2015;**4**(3):261-268

[87] Rizvi I, Gurkan UA, Tasoglu S, Alagic N, Celli JP, Mensah LB, et al. Flow induces epithelial-mesenchymal transition, cellular heterogeneity and biomarker modulation in 3D ovarian cancer nodules. Proceedings of the National Academy of Sciences.

[88] Luo Z, Güven S, Gozen I, Chen P, Tasoglu S, Anchan RM, et al. Deformation of a single mouse oocyte in a constricted microfluidic channel. Microfluidics and Nanofluidics.

[89] Anchan R, Lindsey J, Ng N, Parasar P, Guven S, El Assal R, et al. Human IPSC-derived steroidogenic cells maintain endocrine function with extended culture in a microfluidic

[90] Gurkan UA, Tasoglu S, Akkaynak D, Avci O, Unluisler S, Canikyan S, et al. Smart interface materials integrated with microfluidics for on-demand local capture and release of

[91] Gurkan UA, Anand T, Tas H, Elkan D, Akay A, Keles HO, et al. Controlled viable release of selectively captured label-free cells in microchannels. Lab On a Chip. 2011;**11**(23):

National Academy of Sciences. 2002;**99**(16):10730-10735

scaffolds. Biofabrication. 2012;**4**(3):035005

1999;**280**:86-89

208 Biomaterials in Regenerative Medicine

2001;**12**(1):107-124

2013;**110**(22):E1974-E1983

chip system. Fertility and Sterility. 2015;**104**(3):e73

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2015;**19**(4):883-890

3979-3989

Tissue healing is a complex process involving a cascade of cellular and molecular events that are mostly shared by the different tissues of the body. Interestingly, the tissue repair process initiates immediately after a traumatic injury and is mediated and controlled by a wide range of cytokines, proteins, and growth factors released from platelets upon activation. Consequently, many growth factors have been considered as therapeutic molecules for the repair or regeneration of a wide range of tissues. Although their role has been only partially elucidated, the potential benefit of most growth factors has been demonstrated. In the last few years, the development of platelet-rich preparations has revolutionized the field of regenerative medicine, due to the repair capacities of the platelet-released growth factors that stimulate and accelerate both soft and hard tissue healing and regeneration. Today, autologous platelet concentrates (APCs) are used in a wide range of disciplines such as dentistry, oral surgery, orthopedics, sport medicine, dermatology, and plastic and reconstructive surgery. The purpose of this chapter is to describe the current evidence regarding the benefits of using autologous platelet concentrates in various oral surgery procedures, using a systematic review approach.

**Keywords:** platelet concentrates, growth factors, tissue regeneration, oral surgery

## **1. Introduction**

Autologous platelet concentrates (APCs) have been widely used in many different clinical situations that require a rapid tissue healing and regeneration as it is especially the case in oral and maxillofacial surgery, orthopedics, sports medicine, ophthalmology, and in the treatment of skin ulcers. APCs are hemocomponents, obtained through centrifugation of patient's own blood, in order to collect the most active blood components: platelets, fibrin,

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and in certain cases also leukocytes. The final product has a platelet concentration higher than the basal level, consequently has an increased number of platelet-derived growth factors [1]. The rationale of the clinical use of such platelet-rich preparations is based upon the concept of exploiting their contents enriched of numerous mitogenic platelet-derived growth factors (including platelet-derived growth factor (PDGF), transforming growth factor-β (TGF-β), endothelial growth factor (EGF), vascular endothelial growth factor (VEGF), insulin-like growth factor-1 (IGF-1), basic fibroblast growth factor (FGF), and hepatocyte growth factor (HGF)) as well as other key molecules in promoting tissue healing (as adhesive proteins, pro-coagulant factors, cytokines, chemokines, and anti-microbial proteins [2–5]) to stimulate many biological functions, such as chemotaxis, angiogenesis, proliferation, and differentiation (**Table 1**) in order to enhance hard and soft tissue healing.

**2. Brief history of platelet concentrates**

More than 40 years ago, these technologies were originally used as sealant-adhesive agents in the treatment of hemorrhage with the aim of blocking the blood leakage [6]. Subsequently, other molecules involved in the coagulation process were combined to such fibrin preparations to improve their adhesive properties. These preparations were referred as "plateletfibrinogen-thrombin mixtures" and were successfully used in ophthalmology [7, 8], general

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It is noteworthy that the application of these preparations was essentially related to their adhesive properties and the platelets were served only to reinforce the fibrin matrix architecture. A few years later, it was developed the concept that these preparations could have healing and regenerative properties. In the late 1980s, Knighton et al. [12–14] used the autologous "Platelet-Derived Wound Healing Factors (PDWHF)," which were prepared through twostep centrifugation process, in the treatment of chronic nonhealing cutaneous ulcers. In 1997, Whitman et al. used platelet concentrate referred as "platelet gel" in oral and maxillofacial

The term "Platelet-Rich Plasma" (PRP) was, for the first time, introduced by Kingsley et al. to describe a thrombocyte concentrate [16] used for the treatment of severe thrombopenia. However, the use of PRP term really started with Marx in 1998 [1] when he published a comparative clinical study in which the PRP regenerative potential was demonstrated in a series of patients undergoing mandibular reconstruction. Afterward, the PRP product was associated with the concept of platelet growth factors and their potential contribution to the

According to the PRP protocol, the blood is collected in tubes containing anticoagulants and processed by two centrifugation steps. **Figure 1** illustrates schematically the specific protocol [17]. The final PRP product can be applied to the surgical site with a syringe or be activated by thrombin and/or calcium chloride to trigger platelet activation and to stimulate the fibrin

After blood collection into tubes with anticoagulant, the first centrifugation at low force (soft spin) allows the separation of blood into three distinct layers: red blood cells at the bottom, a cellular plasma (platelet-poor plasma (PPP)) in the upper portion, and a whitish layer called buffy coat located between them containing the highest concentration of platelets and leukocytes. For the production of Pure-PRP (P-PRP), PPP and the superficial buffy coat layer are transferred into another tube and centrifuged at high forces (hard spin), after which most of the PPP and leukocytes are discarded and the final P-PRP can be collected. For the production of PRP rich in leukocytes (L-PRP), PPP, the entire buffy coat layer and some residual red blood cells are collected and transferred in another tube to be hard spin centrifuged. To obtain

surgery [9], and neurosurgery [10]. Other authors termed them as "gel foam" [11].

**2.1. Fibrin adhesives**

surgery [15].

**2.2. Platelet-rich plasma**

enhancement of tissue healing.

polymerization.


Extracted and readapted from Anitua et al. [2].

**Table 1.** Platelet α-granule contents and their functional categories.

## **2. Brief history of platelet concentrates**

#### **2.1. Fibrin adhesives**

and in certain cases also leukocytes. The final product has a platelet concentration higher than the basal level, consequently has an increased number of platelet-derived growth factors [1]. The rationale of the clinical use of such platelet-rich preparations is based upon the concept of exploiting their contents enriched of numerous mitogenic platelet-derived growth factors (including platelet-derived growth factor (PDGF), transforming growth factor-β (TGF-β), endothelial growth factor (EGF), vascular endothelial growth factor (VEGF), insulin-like growth factor-1 (IGF-1), basic fibroblast growth factor (FGF), and hepatocyte growth factor (HGF)) as well as other key molecules in promoting tissue healing (as adhesive proteins, pro-coagulant factors, cytokines, chemokines, and anti-microbial proteins [2–5]) to stimulate many biological functions, such as chemotaxis, angiogenesis, proliferation, and differentia-

tion (**Table 1**) in order to enhance hard and soft tissue healing.

Adhesive proteins VWF + propeptide, Fg, Fn, Vn, TSP-1,

Proteases and antiproteases TIMP-4, MMP-4, inhibitor of FIX,

Basic proteins and others PF4, β-TG, PBP, CTAP III, NAP-2,

Membrane glycoproteins αIIbβ3, αvβ3, GPIb, PECAM-1, most

**Table 1.** Platelet α-granule contents and their functional categories.

Extracted and readapted from Anitua et al. [2].

protein

and ES

Others CS-4, AB, and Ig Diverse

Clotting factors and associated

210 Biomaterials in Regenerative Medicine

Fibrinolytic factors and associated

Growth factors, cytokines, and

proteins

proteins

chemokines

**Category Term Biological activities**

F V/Va, F XI, Gas6, protein S, HMWK, AT, and TFPI

Pgn, PAI-I, u-PAm, OSN, α2-AP, HRGP, TAFI, and α2-M

PDGF, TGFβ-1 and -2, EGF, IGF-1,VEGF, bFGF, FGF-2, HGF, CCL5, IL-8, MIP-1α, CXCL5, MCP-3, ANG-1, and IL-1β, neutrophil chemotactic

Antimicrobial proteins TC Bactericidal and fungicidal properties

plasma membrane constituents, receptors for primary agonists, CD40L, TF, and P-selection

PN-II, C1-INH, and A1AT

Cell contact interactions, clotting, and extracellular matrix composition

Thrombin production and regulation

Plasmin production and vascular

Angiogenesis, vascular modeling, regulation of coagulation, and regulation of cellular behavior

Chemotaxis, cell proliferation and differentiation, and angiogenesis

Regulation of angiogenesis, vascular modeling, and cellular interactions

Platelet aggregation and adhesion, endocytosis of proteins, inflammation, thrombin generation, and platelet-leukocyte interactions

and angiogenesis

modeling

and laminin-8

More than 40 years ago, these technologies were originally used as sealant-adhesive agents in the treatment of hemorrhage with the aim of blocking the blood leakage [6]. Subsequently, other molecules involved in the coagulation process were combined to such fibrin preparations to improve their adhesive properties. These preparations were referred as "plateletfibrinogen-thrombin mixtures" and were successfully used in ophthalmology [7, 8], general surgery [9], and neurosurgery [10]. Other authors termed them as "gel foam" [11].

It is noteworthy that the application of these preparations was essentially related to their adhesive properties and the platelets were served only to reinforce the fibrin matrix architecture. A few years later, it was developed the concept that these preparations could have healing and regenerative properties. In the late 1980s, Knighton et al. [12–14] used the autologous "Platelet-Derived Wound Healing Factors (PDWHF)," which were prepared through twostep centrifugation process, in the treatment of chronic nonhealing cutaneous ulcers. In 1997, Whitman et al. used platelet concentrate referred as "platelet gel" in oral and maxillofacial surgery [15].

#### **2.2. Platelet-rich plasma**

The term "Platelet-Rich Plasma" (PRP) was, for the first time, introduced by Kingsley et al. to describe a thrombocyte concentrate [16] used for the treatment of severe thrombopenia. However, the use of PRP term really started with Marx in 1998 [1] when he published a comparative clinical study in which the PRP regenerative potential was demonstrated in a series of patients undergoing mandibular reconstruction. Afterward, the PRP product was associated with the concept of platelet growth factors and their potential contribution to the enhancement of tissue healing.

According to the PRP protocol, the blood is collected in tubes containing anticoagulants and processed by two centrifugation steps. **Figure 1** illustrates schematically the specific protocol [17]. The final PRP product can be applied to the surgical site with a syringe or be activated by thrombin and/or calcium chloride to trigger platelet activation and to stimulate the fibrin polymerization.

After blood collection into tubes with anticoagulant, the first centrifugation at low force (soft spin) allows the separation of blood into three distinct layers: red blood cells at the bottom, a cellular plasma (platelet-poor plasma (PPP)) in the upper portion, and a whitish layer called buffy coat located between them containing the highest concentration of platelets and leukocytes. For the production of Pure-PRP (P-PRP), PPP and the superficial buffy coat layer are transferred into another tube and centrifuged at high forces (hard spin), after which most of the PPP and leukocytes are discarded and the final P-PRP can be collected. For the production of PRP rich in leukocytes (L-PRP), PPP, the entire buffy coat layer and some residual red blood cells are collected and transferred in another tube to be hard spin centrifuged. To obtain

**2.4. Plasma rich in growth factors**

target site or may be preactivated by adding 0.2 ml of 10% CaCl<sup>2</sup>

**Figure 2.** Protocol for PRF production. (Adapted from Dohan Ehrenfest et al. [19]).

**2.5. Technical differences between PRGF, PRP, and PRF**

PRGF differs from PRP for the following technical aspects:

**2.** Requires a single centrifugation for the preparation.

In addition, PRGF also differs from PRF for these features:

**5.** Platelet concentration is reduced (2–3 fold the baseline, respect to 5–8× for PRP).

**2.** PRGF liquid can be combined with bone graft materials for bone regeneration procedure.

**3.** PRP liquid can be mixed with bone graft materials for bone regeneration procedures.

**1.** Different products can be obtained (liquid, gel, membrane, and fibrin clot).

**1.** Blood volume drawn is minimal (5–40 ml).

**4.** Does not contain proinflammatory cytokines.

Differences between PRP and PRF are as follows:

**1.** PRP preparation requires two centrifugations.

**2.** Different products can be obtained (liquid and fibrin clot).

**3.** Does not contain leukocytes.

In parallel to the introduction of PRP and PRF, Anitua in 1999 proposed another platelet concentrate protocol, denominated plasma rich in growth factors (PRGF) [21]. Briefly, blood is collected in tubes containing 0.2 ml of 3.8% trisodium citrate as anticoagulant. After a centrifugation at 580 g for 8 min, red blood cells and buffy coat layer are deposited at the bottom of the tube and the plasma component above. The latter is then manually separated into two fractions. The lower portion of about 2 ml, above the buffy coat, is the PRGF, whereas the upper portion is the plasma poor in growth factors (PPGF) (**Figure 3**). The final PRGF product may be applied as a liquid fraction to the

to induce the clot formation [22].

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**Figure 1.** Protocol for PRP production.

the final L-PRP, PPP is discarded leading to an L-PRP that contains the buffy coat with most of the platelets and leukocytes, some residual red blood cells and PPP (adapted from Dohan Ehrenfest et al. [17]).

Currently, more than 20 different commercial systems for PRP preparation exist, which may lead to products with different features, especially regarding the composition and the cell concentration rate respect to baseline. On average, a 5–8× concentration is achieved though a ratio of up to 11× has been reported with PRP.

#### **2.3. Platelet-rich fibrin**

In 2001, Choukroun et al. developed a protocol for producing a hemocomponent named platelet-rich fibrin (PRF) [18]. Here, the blood is collected in tubes in the absence of anticoagulant and centrifuged with moderate forces (3000 rpm in a dedicated centrifuge) for 12 min. Afterward, three layers are formed: red blood cells and acellular plasma are located, respectively, at the bottom and at the top of the tube, and the fibrin clot, positioned between them, is PRF (**Figure 2**). Since the formation of the PRF clot naturally occurs within the tube, it has a strong fibrin matrix in which most of the platelets and leukocytes are embedded [19]. Since its introduction, PRF has undergone some development: the advanced PRF (a-PRF) was launched a few years ago, characterized by a reduced centrifugation speed and time (2700 rpm, 8–10 min), which allows a more even cell distribution within the clot [20]. Recently, the injectable PRF (i-PRF) has also been developed, which may be obtained with a further softer spin (1500 rpm, 3 min), consisting of a liquid form, very rich in white cells that can be used by infiltration into tissues and joints.

**Figure 2.** Protocol for PRF production. (Adapted from Dohan Ehrenfest et al. [19]).

#### **2.4. Plasma rich in growth factors**

In parallel to the introduction of PRP and PRF, Anitua in 1999 proposed another platelet concentrate protocol, denominated plasma rich in growth factors (PRGF) [21]. Briefly, blood is collected in tubes containing 0.2 ml of 3.8% trisodium citrate as anticoagulant. After a centrifugation at 580 g for 8 min, red blood cells and buffy coat layer are deposited at the bottom of the tube and the plasma component above. The latter is then manually separated into two fractions. The lower portion of about 2 ml, above the buffy coat, is the PRGF, whereas the upper portion is the plasma poor in growth factors (PPGF) (**Figure 3**). The final PRGF product may be applied as a liquid fraction to the target site or may be preactivated by adding 0.2 ml of 10% CaCl<sup>2</sup> to induce the clot formation [22].

#### **2.5. Technical differences between PRGF, PRP, and PRF**

PRGF differs from PRP for the following technical aspects:


the final L-PRP, PPP is discarded leading to an L-PRP that contains the buffy coat with most of the platelets and leukocytes, some residual red blood cells and PPP (adapted from Dohan

Currently, more than 20 different commercial systems for PRP preparation exist, which may lead to products with different features, especially regarding the composition and the cell concentration rate respect to baseline. On average, a 5–8× concentration is achieved though a

In 2001, Choukroun et al. developed a protocol for producing a hemocomponent named platelet-rich fibrin (PRF) [18]. Here, the blood is collected in tubes in the absence of anticoagulant and centrifuged with moderate forces (3000 rpm in a dedicated centrifuge) for 12 min. Afterward, three layers are formed: red blood cells and acellular plasma are located, respectively, at the bottom and at the top of the tube, and the fibrin clot, positioned between them, is PRF (**Figure 2**). Since the formation of the PRF clot naturally occurs within the tube, it has a strong fibrin matrix in which most of the platelets and leukocytes are embedded [19]. Since its introduction, PRF has undergone some development: the advanced PRF (a-PRF) was launched a few years ago, characterized by a reduced centrifugation speed and time (2700 rpm, 8–10 min), which allows a more even cell distribution within the clot [20]. Recently, the injectable PRF (i-PRF) has also been developed, which may be obtained with a further softer spin (1500 rpm, 3 min), consisting of a liquid form, very rich in white cells that can be used by infiltration into tissues and joints.

Ehrenfest et al. [17]).

**Figure 1.** Protocol for PRP production.

212 Biomaterials in Regenerative Medicine

**2.3. Platelet-rich fibrin**

ratio of up to 11× has been reported with PRP.


In addition, PRGF also differs from PRF for these features:


Differences between PRP and PRF are as follows:


could not allow sound meta-analyses. Regarding the bone formation, the qualitative synthesis of the histological analyses reported a better bone quality in biopsies retrieved from sites treated with platelet concentrate. Furthermore, the meta-analysis of the histomorphometric evaluation of the bone formation (including only few of the included studies) revealed that sites treated with platelet concentrate showed a statistically significant greater proportion of new bone than the controls, at 3 months of follow-up [23]. Even though the results of the metaanalysis suggest a beneficial effect of autologous platelet concentrates on bone formation, caution should be paid on interpreting such results, since the available evidence is scarce and of limited quality [23]. In spite of the relatively numerous randomized clinical studies assessing the value of APCs for enhancing postextraction socket healing, the main reason that prevents a wide meta-analysis is that the methods used for assessing bone regeneration and socket preservation are very different, providing different information that cannot be aggregated. In fact, different studies used different techniques like histological and immunohistological analysis, histomorphometric evaluation, scintigraphy, micro-TC, intraoral radiography, cone beam computed tomography, and clinical assessment, for the evaluation of different variables as percentage of new bone formation, osteoblasts activity, bone density, crestal bone level changes, ridge width and height, and soft tissues health status [23, 24]. All studies that evaluated postoperative symptoms like pain, swelling, trismus, and adverse events like dry socket, alveolitis, and acute inflammation concluded that APCs are effective in reducing symptoms and the incidence of adverse events, with an overall improvement of patients' quality of life [24]. The most recent systematic review on this topic concludes that APCs should be used in postextraction sites in order to improve clinical and radiographic outcomes such as bone density and soft tissue healing, as well as to reduce postoperative symptoms [24]. The actual benefit of APCs on decreasing of pain in extraction sockets, however, though consistently reported, is still not quantifiable [24]. In **Figure 4**, a brief sequence of pictures documenting a

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**Figure 4.** Double tooth extraction in the upper jaw and placement of PRGF in the extraction socket. (A) Fresh sockets after atraumatic extraction. (B) Sockets filled with PRGF clot; it can be sutured for a better stabilization. (C) 14 days after

surgery: excellent healing of soft tissues is shown.

**Figure 3.** Process of PRGF production. Details are given in the text. (Adapted from Anitua [22]).

## **3. Clinical evidence of the efficacy of autologous platelet concentrates in oral regenerative surgical procedures**

#### **3.1. General aspects**

In the last couple of decades, it has been observed a growing interest in the use of autologous platelet concentrates during oral regenerative surgical procedures as an adjunctive tool to enhance the hard and soft tissue healing. The following sections will summarize the recent evidence on the efficacy of autologous platelet concentrates in the dental field. The recent evidence has been obtained using a systematic review approach. The focused question was: "Does the adjunct of autologous platelet concentrate produce benefits to hard and soft tissue healing in oral surgery procedures in terms of tissue parameters, postoperative complications as well as patient's postoperative quality of life?" In order to address the aim of this chapter, electronic searches were performed on the main scientific databases (MEDLINE, Scopus, and Cochrane Central Register of Controlled Trials). Proper search terms were used, combined by Boolean operators. Only controlled clinical trials, randomized clinical trials, and existing systematic reviews and meta-analyses of the literature were included. The outcomes were complications and adverse events, treatment success, discomfort/quality of life, bone healing and remodeling assessed by histological and radiographic techniques, and soft tissue healing. The surgical procedures taken into consideration were: tooth extraction, periodontal surgery, endodontic surgery and treatment of immature necrotic teeth, maxillary sinus augmentation, and implant treatment. When possible, a quantitative analysis was undertaken by meta-analysis approach, using the software RevMan (RevMan, Version 5.3, The Nordic Cochrane Center, The Cochrane Collaboration, Copenhagen, Denmark, 2014).

#### **3.2. Alveolar postextraction healing**

Several recent systematic reviews evaluated the efficacy of autologous platelet concentrates in enhancing alveolar socket healing after tooth extraction [23, 24]. Beneficial effects were generally reported in terms of better soft tissue healing, better clinical and histological epithelialization of wound margins, and a faster wound closure, although the heterogeneity of the data could not allow sound meta-analyses. Regarding the bone formation, the qualitative synthesis of the histological analyses reported a better bone quality in biopsies retrieved from sites treated with platelet concentrate. Furthermore, the meta-analysis of the histomorphometric evaluation of the bone formation (including only few of the included studies) revealed that sites treated with platelet concentrate showed a statistically significant greater proportion of new bone than the controls, at 3 months of follow-up [23]. Even though the results of the metaanalysis suggest a beneficial effect of autologous platelet concentrates on bone formation, caution should be paid on interpreting such results, since the available evidence is scarce and of limited quality [23]. In spite of the relatively numerous randomized clinical studies assessing the value of APCs for enhancing postextraction socket healing, the main reason that prevents a wide meta-analysis is that the methods used for assessing bone regeneration and socket preservation are very different, providing different information that cannot be aggregated. In fact, different studies used different techniques like histological and immunohistological analysis, histomorphometric evaluation, scintigraphy, micro-TC, intraoral radiography, cone beam computed tomography, and clinical assessment, for the evaluation of different variables as percentage of new bone formation, osteoblasts activity, bone density, crestal bone level changes, ridge width and height, and soft tissues health status [23, 24]. All studies that evaluated postoperative symptoms like pain, swelling, trismus, and adverse events like dry socket, alveolitis, and acute inflammation concluded that APCs are effective in reducing symptoms and the incidence of adverse events, with an overall improvement of patients' quality of life [24]. The most recent systematic review on this topic concludes that APCs should be used in postextraction sites in order to improve clinical and radiographic outcomes such as bone density and soft tissue healing, as well as to reduce postoperative symptoms [24]. The actual benefit of APCs on decreasing of pain in extraction sockets, however, though consistently reported, is still not quantifiable [24]. In **Figure 4**, a brief sequence of pictures documenting a

**3. Clinical evidence of the efficacy of autologous platelet concentrates in** 

**Figure 3.** Process of PRGF production. Details are given in the text. (Adapted from Anitua [22]).

In the last couple of decades, it has been observed a growing interest in the use of autologous platelet concentrates during oral regenerative surgical procedures as an adjunctive tool to enhance the hard and soft tissue healing. The following sections will summarize the recent evidence on the efficacy of autologous platelet concentrates in the dental field. The recent evidence has been obtained using a systematic review approach. The focused question was: "Does the adjunct of autologous platelet concentrate produce benefits to hard and soft tissue healing in oral surgery procedures in terms of tissue parameters, postoperative complications as well as patient's postoperative quality of life?" In order to address the aim of this chapter, electronic searches were performed on the main scientific databases (MEDLINE, Scopus, and Cochrane Central Register of Controlled Trials). Proper search terms were used, combined by Boolean operators. Only controlled clinical trials, randomized clinical trials, and existing systematic reviews and meta-analyses of the literature were included. The outcomes were complications and adverse events, treatment success, discomfort/quality of life, bone healing and remodeling assessed by histological and radiographic techniques, and soft tissue healing. The surgical procedures taken into consideration were: tooth extraction, periodontal surgery, endodontic surgery and treatment of immature necrotic teeth, maxillary sinus augmentation, and implant treatment. When possible, a quantitative analysis was undertaken by meta-analysis approach, using the software RevMan (RevMan, Version 5.3, The Nordic Cochrane Center,

Several recent systematic reviews evaluated the efficacy of autologous platelet concentrates in enhancing alveolar socket healing after tooth extraction [23, 24]. Beneficial effects were generally reported in terms of better soft tissue healing, better clinical and histological epithelialization of wound margins, and a faster wound closure, although the heterogeneity of the data

**oral regenerative surgical procedures**

The Cochrane Collaboration, Copenhagen, Denmark, 2014).

**3.2. Alveolar postextraction healing**

**3.1. General aspects**

214 Biomaterials in Regenerative Medicine

**Figure 4.** Double tooth extraction in the upper jaw and placement of PRGF in the extraction socket. (A) Fresh sockets after atraumatic extraction. (B) Sockets filled with PRGF clot; it can be sutured for a better stabilization. (C) 14 days after surgery: excellent healing of soft tissues is shown.

case of double extraction in the posterior upper jaw, subsequent positioning of platelet concentrate (PRGF) into extraction sockets, and the postsurgical clinical healing is shown.

the control group [37]. This suggests that the adjunct of P-PRP to the endodontic surgery procedure may produce significant beneficial effect to patients' quality of life during the early

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Another pilot clinical study on endodontic surgery compared a group of seven patients treated with the adjunct of PRF versus four control patients [38]. In addition to confirming the beneficial effects in the early postoperative period regarding significant reduction of pain and swelling, a significantly better healing of the lesion was observed after 2 and 3 months but not after 1-year follow-up [38]. This preliminary study had a very small sample size, so results should be interpreted cautiously. The latter two studies have been included in a recently published Cochrane systematic review on endodontic procedures for the retreatment of periapical defects, which concluded that there is evidence that adjunctive use of a gel of plasma rich in growth factors reduces postoperative pain compared with no grafting [39]. Regarding other possible beneficial effects of APCs in endodontic surgery,

The use of platelet concentrates in association with grafting material during maxillary sinus augmentation procedure provided conflicting results in both preclinical and clinical studies [40, 41]. A recent meta-analysis documented that PRP combined to graft materials, in this type of surgical procedure, had no adjunctive effect on bone formation, on implant survival and implant stability as well as it did not show any statistically significant differences on marginal bone loss or alveolar bone height, compared to the bone graft alone [42]. Similar conclusions were also reported in other systematic reviews [43, 44]. However, another metaanalysis reported opposite conclusions concerning the bone formation supporting the use of PRP for sinus bone graft [45]. Furthermore, beneficial effects on soft tissue healing as well as reduction of postoperative discomfort were often reported [43]. Such variability in results could be ascribed to a number of factors. First of all, different techniques have been adopted for the preparation of platelet concentrates, leading to the products with different characteristics (final concentration of platelets and white cells, presence or absence of leukocytes, use of anticoagulants and activators, different mechanical consistence of the product, and association with different graft types) and, presumably, different biological activities. Secondly, different studies may also differ in experimental design, objectives, outcome variables, inclusion criteria, and follow-up duration. Furthermore, the sinus augmentation technique, though representing a very popular model for the assessment of bone substitutes for bone regeneration, suffers from a number of confounding factors that make standardization difficult, like patient age, residual bone quality and quantity, smoking habits, volume of graft used, porosity and general features of the graft material (e.g., intraoral or extraoral autografts, allografts, xenografts, and alloplasts), graft resorbability over time, graft healing time, and intra- and postsurgical complications like membrane perforation, infection, expertise of the clinician, techniques adopted for evaluating bone formation, including the position of the biopsy (crestal or lateral). Finally, if one aims at evaluating the effect of APCs on implant survival and success in the maxillary sinus augmentation, a number of additional factors concerning the implants and the prosthetic reconstruction must be considered, e.g., implant length

postoperative period.

further evidence is needed.

**3.5. Maxillary sinus augmentation**

#### **3.3. Periodontal defects**

Several systematic reviews and some meta-analyses evaluated the efficacy of autologous platelet concentrates in the treatment of periodontal defects, including intrabony defects, gingival recessions, and furcation defects [25–33]. Beneficial effects on clinical and radiographic outcomes in the treatment of infrabony defects were reported, although a high heterogeneity emerged among the clinical studies in terms of outcomes evaluated and bioactive agents/ procedures combined with autologous platelet concentrates [25, 28, 30–32]. Furthermore, two meta-analyses [28, 30] concluded that PRP might exert positive adjunctive effects in the surgical treatment of such defects when combined with grafting materials, but no adjunctive effects were found in association with the guided tissue regeneration technique. Indeed, the latter is considered the gold standard for the treatment of periodontal intrabony defects and its use could probably mask the PRP effect. Regarding the autologous platelet concentrates' effect on gingival recessions, very few systematic reviews have been conducted, presumably due to the limited data about it. PRP or PRF did not show any clinical improvements in the treatment of gingival recessions or furcation defects [28, 33]. A Cochrane systematic review on this topic is still ongoing [34] and its results will certainly shed light on the actual evidence level regarding this topic, possibly confirming early indications of previous systematic reviews.

#### **3.4. Endodontics and endodontic surgery**

Platelet concentrates have been recently used in the clinical treatment of immature necrotic teeth, with the aim of regenerating the intracanal pulp and stimulating tooth development, as well as in the surgical treatment of teeth with apical periodontitis to enhance healing of periapical tissues. Clinical evidence on the benefits of the use of platelet concentrates in these pathologies exists but is still scarce. A recent systematic review concluded that periapical healing and apical closure were improved in those immature necrotic teeth treated with PRP compared with the control group without PRP, even though not statistically significant, and a significant better thickening of dentinal walls and root lengthening were also reported [35]. However, from the histological point of view, it seems that a true regeneration of necrotic pulp tissue of either mature or immature teeth was not achieved after using platelet concentrates. In fact, the neoformed intracanal tissues were mainly cemento-like, bone-like, and connective tissue. Root canals were repopulated with living tissue that only marginally resembled the original pulp. Despite this, the root maturation may be achieved and teeth function is not compromised [36].

Though the use of APCs in the management of periapical lesions could be considered a proper indication, very scarce studies are present regarding this topic. One randomized clinical study evaluated the postoperative quality of life in patients undergoing endodontic surgery [37]. The test group of 18 patients was treated with the adjunct of P-PRP and the control group was treated conventionally, without P-PRP. The test group showed significantly less pain and swelling, fewer analgesics taken, and improved functional activities as compared with the control group [37]. This suggests that the adjunct of P-PRP to the endodontic surgery procedure may produce significant beneficial effect to patients' quality of life during the early postoperative period.

Another pilot clinical study on endodontic surgery compared a group of seven patients treated with the adjunct of PRF versus four control patients [38]. In addition to confirming the beneficial effects in the early postoperative period regarding significant reduction of pain and swelling, a significantly better healing of the lesion was observed after 2 and 3 months but not after 1-year follow-up [38]. This preliminary study had a very small sample size, so results should be interpreted cautiously. The latter two studies have been included in a recently published Cochrane systematic review on endodontic procedures for the retreatment of periapical defects, which concluded that there is evidence that adjunctive use of a gel of plasma rich in growth factors reduces postoperative pain compared with no grafting [39]. Regarding other possible beneficial effects of APCs in endodontic surgery, further evidence is needed.

#### **3.5. Maxillary sinus augmentation**

case of double extraction in the posterior upper jaw, subsequent positioning of platelet concentrate (PRGF) into extraction sockets, and the postsurgical clinical healing is shown.

Several systematic reviews and some meta-analyses evaluated the efficacy of autologous platelet concentrates in the treatment of periodontal defects, including intrabony defects, gingival recessions, and furcation defects [25–33]. Beneficial effects on clinical and radiographic outcomes in the treatment of infrabony defects were reported, although a high heterogeneity emerged among the clinical studies in terms of outcomes evaluated and bioactive agents/ procedures combined with autologous platelet concentrates [25, 28, 30–32]. Furthermore, two meta-analyses [28, 30] concluded that PRP might exert positive adjunctive effects in the surgical treatment of such defects when combined with grafting materials, but no adjunctive effects were found in association with the guided tissue regeneration technique. Indeed, the latter is considered the gold standard for the treatment of periodontal intrabony defects and its use could probably mask the PRP effect. Regarding the autologous platelet concentrates' effect on gingival recessions, very few systematic reviews have been conducted, presumably due to the limited data about it. PRP or PRF did not show any clinical improvements in the treatment of gingival recessions or furcation defects [28, 33]. A Cochrane systematic review on this topic is still ongoing [34] and its results will certainly shed light on the actual evidence level regarding

this topic, possibly confirming early indications of previous systematic reviews.

Platelet concentrates have been recently used in the clinical treatment of immature necrotic teeth, with the aim of regenerating the intracanal pulp and stimulating tooth development, as well as in the surgical treatment of teeth with apical periodontitis to enhance healing of periapical tissues. Clinical evidence on the benefits of the use of platelet concentrates in these pathologies exists but is still scarce. A recent systematic review concluded that periapical healing and apical closure were improved in those immature necrotic teeth treated with PRP compared with the control group without PRP, even though not statistically significant, and a significant better thickening of dentinal walls and root lengthening were also reported [35]. However, from the histological point of view, it seems that a true regeneration of necrotic pulp tissue of either mature or immature teeth was not achieved after using platelet concentrates. In fact, the neoformed intracanal tissues were mainly cemento-like, bone-like, and connective tissue. Root canals were repopulated with living tissue that only marginally resembled the original pulp. Despite this, the root maturation may be achieved and teeth function is not

Though the use of APCs in the management of periapical lesions could be considered a proper indication, very scarce studies are present regarding this topic. One randomized clinical study evaluated the postoperative quality of life in patients undergoing endodontic surgery [37]. The test group of 18 patients was treated with the adjunct of P-PRP and the control group was treated conventionally, without P-PRP. The test group showed significantly less pain and swelling, fewer analgesics taken, and improved functional activities as compared with

**3.4. Endodontics and endodontic surgery**

compromised [36].

**3.3. Periodontal defects**

216 Biomaterials in Regenerative Medicine

The use of platelet concentrates in association with grafting material during maxillary sinus augmentation procedure provided conflicting results in both preclinical and clinical studies [40, 41]. A recent meta-analysis documented that PRP combined to graft materials, in this type of surgical procedure, had no adjunctive effect on bone formation, on implant survival and implant stability as well as it did not show any statistically significant differences on marginal bone loss or alveolar bone height, compared to the bone graft alone [42]. Similar conclusions were also reported in other systematic reviews [43, 44]. However, another metaanalysis reported opposite conclusions concerning the bone formation supporting the use of PRP for sinus bone graft [45]. Furthermore, beneficial effects on soft tissue healing as well as reduction of postoperative discomfort were often reported [43]. Such variability in results could be ascribed to a number of factors. First of all, different techniques have been adopted for the preparation of platelet concentrates, leading to the products with different characteristics (final concentration of platelets and white cells, presence or absence of leukocytes, use of anticoagulants and activators, different mechanical consistence of the product, and association with different graft types) and, presumably, different biological activities. Secondly, different studies may also differ in experimental design, objectives, outcome variables, inclusion criteria, and follow-up duration. Furthermore, the sinus augmentation technique, though representing a very popular model for the assessment of bone substitutes for bone regeneration, suffers from a number of confounding factors that make standardization difficult, like patient age, residual bone quality and quantity, smoking habits, volume of graft used, porosity and general features of the graft material (e.g., intraoral or extraoral autografts, allografts, xenografts, and alloplasts), graft resorbability over time, graft healing time, and intra- and postsurgical complications like membrane perforation, infection, expertise of the clinician, techniques adopted for evaluating bone formation, including the position of the biopsy (crestal or lateral). Finally, if one aims at evaluating the effect of APCs on implant survival and success in the maxillary sinus augmentation, a number of additional factors concerning the implants and the prosthetic reconstruction must be considered, e.g., implant length and width, shape, surface micro- and nano-geometry, type of implant-abutment connection, implant primary stability, and number and position of the implants. In fact, it is well known that implant survival in the augmented maxillary sinus is more variable than that of implants placed in the posterior maxilla.

**3.8. Diabetes**

**3.9. Irradiated patients**

mon in diabetic patients following tooth extraction.

with P-PRP and subsequently healed [64].

in patients with a history of head and neck radiotherapy.

**3.10. Osteoradionecrosis (ORN)**

**3.11. Coagulation defects**

In a split-mouth study in 34 patients with diabetes mellitus candidate to bilateral tooth extraction of a total of 127 teeth, alveolar sockets on the test side were treated with P-PRP and on the control side they were left to heal in a natural way [63]. The P-PRP group showed a significantly better healing (evaluated through the Healing Index) and a faster closure of the alveolus (by evaluation of the residual socket volume) at 3 and 7 days after extraction, as compared to the control group [63]. The authors concluded that platelet growth factors stimulate a faster epithelialization, protecting the alveolus in early healing steps. Hence, it is avoided occurrence of alveolitis, very com-

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In a split-mouth study on 20 patients that underwent radiotherapy for head and neck cancer, and candidate to bilateral tooth extraction, alveolar sockets on the directly irradiated side were treated with P-PRP (test group) and on the untreated side they were left to heal in a natural way (control group) [64]. Twenty-four bilateral extractions were performed in the mandible and 33 in the maxilla for a total of 114 extractions. The P-PRP group showed a significantly better healing, in terms of Healing Index and residual socket volume at each follow-up (7, 14, and 21 days), as compared to the control group. Patients were followed up to 24 months after surgery. In the control group, two cases with bone exposure were retreated

In a series of 10 patients who developed osteoradionecrosis, debridement of necrotic bone (performed with ultrasonic instruments) was associated with the adjunct of P-PRP. All patients successfully healed, with no intraoperative or postoperative complications up to 12 months of follow-up [65]. Tissue regeneration and closure was excellent, and postoperative pain, assessed through visual analogue score (VAS) scores, was low. In spite of the absence of a control group, this study suggested that P-PRP may be beneficial as an adjunct to surgical treatment of ORN, for predictable enhancement of tissue vascularization and epithelialization

A case-control study was performed on 66 patients affected by severe thrombocytopenia (<50,000 platelets/μL) and in need for at least two tooth extraction each [66]. For these patients, postoperative bleeding represents an important issue. Teeth were extracted in two consecutive interventions. In one intervention, the patients received a platelet transfusion before extraction (systemic treatment) and in the other session, the postalveolar sockets were filled with P-PRP (local treatment). Patients were evaluated frequently in the first 7 days after extraction. The group treated locally with P-PRP showed a statistically significant reduction in postoperative bleeding, hematoma, and need for reintervention, as compared to the group receiving a systemic infusion of platelets [66]. Therefore, P-PRP in postalveolar sockets may strongly

A randomized clinical trial evaluating the effect of P-PRP adjunct on postoperative quality of life of patients undergoing maxillary sinus augmentation procedure, found a beneficial effect of P-PRP regarding pain, swelling, hematoma, and other postoperative symptoms, improving patient's acceptance of this often demanding procedure [46].

## **3.6. Implant dentistry**

As it is claimed that platelet concentrates may promote bone regeneration, several animal studies have been conducted to assess the PRP effect on the osseointegration process, through histological and histomorphometrical evaluation, but controversial results have been reported. In fact, some studies did not demonstrate any advantages of PRP over non-PRP control groups at stimulating faster bone formation or higher bone-implant contact [47–49]. By contrast, histomorphometric analyses of the bone-implant interface in the early healing phase after implantation (6 or 8 weeks) revealed a significantly higher percentage of boneimplant contact in implants coated with liquid PRP formulation compared to those not PRPbioactivated [50–52]. In addition to being time dependent, PRP effect is also site dependent since its effect has been reported to decrease with increasing distance from the site of application [52]. Similarly, liquid-PRP showed a tendency to increase the bone apposition to roughened titanium implants during early healing phase [53, 54].

Clinical studies reported a higher bone formation around the implants [55] and a good preservation of the alveolar crest around postextraction implants [56, 57] when APCs were used.

APCs have been also combined with several different types of grafting materials during regenerative procedures associated with implantoprosthesis rehabilitations, showing satisfying results and positive patient-related outcomes [58–61]. A long-term clinical study (10–12 years) on short implant placement in association with PRGF reported an implant survival rate of 98.9% and marginal bone loss inferior to 1 mm [62].

#### **3.7. Systemically compromised patients**

Patients affected by chronic systemic conditions like osteoporosis, cancer, diabetes type I or II, immunodeficiency, hematological/coagulation defects ,and other conditions, often present difficulty in healing even after simple surgical procedures, like the extraction of a tooth. Therefore, the use of a safe tool that may enhance the healing process in a natural way may represent a remarkable benefit for these patients.

A few studies have been published on the use of platelet concentrates in systemically compromised patients, most of them are regarding postextraction healing. Some examples will be reported.

#### **3.8. Diabetes**

and width, shape, surface micro- and nano-geometry, type of implant-abutment connection, implant primary stability, and number and position of the implants. In fact, it is well known that implant survival in the augmented maxillary sinus is more variable than that of implants

A randomized clinical trial evaluating the effect of P-PRP adjunct on postoperative quality of life of patients undergoing maxillary sinus augmentation procedure, found a beneficial effect of P-PRP regarding pain, swelling, hematoma, and other postoperative symptoms, improving

As it is claimed that platelet concentrates may promote bone regeneration, several animal studies have been conducted to assess the PRP effect on the osseointegration process, through histological and histomorphometrical evaluation, but controversial results have been reported. In fact, some studies did not demonstrate any advantages of PRP over non-PRP control groups at stimulating faster bone formation or higher bone-implant contact [47–49]. By contrast, histomorphometric analyses of the bone-implant interface in the early healing phase after implantation (6 or 8 weeks) revealed a significantly higher percentage of boneimplant contact in implants coated with liquid PRP formulation compared to those not PRPbioactivated [50–52]. In addition to being time dependent, PRP effect is also site dependent since its effect has been reported to decrease with increasing distance from the site of application [52]. Similarly, liquid-PRP showed a tendency to increase the bone apposition to rough-

Clinical studies reported a higher bone formation around the implants [55] and a good preservation of the alveolar crest around postextraction implants [56, 57] when APCs

APCs have been also combined with several different types of grafting materials during regenerative procedures associated with implantoprosthesis rehabilitations, showing satisfying results and positive patient-related outcomes [58–61]. A long-term clinical study (10–12 years) on short implant placement in association with PRGF reported an implant survival rate

Patients affected by chronic systemic conditions like osteoporosis, cancer, diabetes type I or II, immunodeficiency, hematological/coagulation defects ,and other conditions, often present difficulty in healing even after simple surgical procedures, like the extraction of a tooth. Therefore, the use of a safe tool that may enhance the healing process in a natural way may

A few studies have been published on the use of platelet concentrates in systemically compromised patients, most of them are regarding postextraction healing. Some examples will

placed in the posterior maxilla.

218 Biomaterials in Regenerative Medicine

**3.6. Implant dentistry**

were used.

be reported.

patient's acceptance of this often demanding procedure [46].

ened titanium implants during early healing phase [53, 54].

of 98.9% and marginal bone loss inferior to 1 mm [62].

represent a remarkable benefit for these patients.

**3.7. Systemically compromised patients**

In a split-mouth study in 34 patients with diabetes mellitus candidate to bilateral tooth extraction of a total of 127 teeth, alveolar sockets on the test side were treated with P-PRP and on the control side they were left to heal in a natural way [63]. The P-PRP group showed a significantly better healing (evaluated through the Healing Index) and a faster closure of the alveolus (by evaluation of the residual socket volume) at 3 and 7 days after extraction, as compared to the control group [63]. The authors concluded that platelet growth factors stimulate a faster epithelialization, protecting the alveolus in early healing steps. Hence, it is avoided occurrence of alveolitis, very common in diabetic patients following tooth extraction.

#### **3.9. Irradiated patients**

In a split-mouth study on 20 patients that underwent radiotherapy for head and neck cancer, and candidate to bilateral tooth extraction, alveolar sockets on the directly irradiated side were treated with P-PRP (test group) and on the untreated side they were left to heal in a natural way (control group) [64]. Twenty-four bilateral extractions were performed in the mandible and 33 in the maxilla for a total of 114 extractions. The P-PRP group showed a significantly better healing, in terms of Healing Index and residual socket volume at each follow-up (7, 14, and 21 days), as compared to the control group. Patients were followed up to 24 months after surgery. In the control group, two cases with bone exposure were retreated with P-PRP and subsequently healed [64].

#### **3.10. Osteoradionecrosis (ORN)**

In a series of 10 patients who developed osteoradionecrosis, debridement of necrotic bone (performed with ultrasonic instruments) was associated with the adjunct of P-PRP. All patients successfully healed, with no intraoperative or postoperative complications up to 12 months of follow-up [65]. Tissue regeneration and closure was excellent, and postoperative pain, assessed through visual analogue score (VAS) scores, was low. In spite of the absence of a control group, this study suggested that P-PRP may be beneficial as an adjunct to surgical treatment of ORN, for predictable enhancement of tissue vascularization and epithelialization in patients with a history of head and neck radiotherapy.

#### **3.11. Coagulation defects**

A case-control study was performed on 66 patients affected by severe thrombocytopenia (<50,000 platelets/μL) and in need for at least two tooth extraction each [66]. For these patients, postoperative bleeding represents an important issue. Teeth were extracted in two consecutive interventions. In one intervention, the patients received a platelet transfusion before extraction (systemic treatment) and in the other session, the postalveolar sockets were filled with P-PRP (local treatment). Patients were evaluated frequently in the first 7 days after extraction. The group treated locally with P-PRP showed a statistically significant reduction in postoperative bleeding, hematoma, and need for reintervention, as compared to the group receiving a systemic infusion of platelets [66]. Therefore, P-PRP in postalveolar sockets may strongly reduce the risk of hemorrhage and related complications in thrombocytopenic patients, due to the ability of stimulating healing and its hemostatic properties.

#### **3.12. Bisphosphonate-related osteonecrosis of the jaws**

Recently renamed medication-related osteonecrosis of the jaw (MRONJ), it is an adverse drug reaction consisting of progressive bone destruction in the maxillofacial region of patients under current or previous treatment with a bisphosphonate or another antiresorptive or antiangiogenic drug. Since APCs demonstrated to enhance bone and soft tissue healing in oral surgery procedures, it is reasonable to believe that they might provide benefits to these patients. A recent systematic review included 18 studies, reporting on 362 patients undergoing oral surgery in combination with APCs [67]. The adjunct of APC in the surgical treatment of necrotic bone removal significantly reduced osteonecrosis recurrence as compared to control (**Figure 5**). APC was associated with a reduced, though not significant, incidence of BRONJ after tooth extraction. Heterogeneity among studies was found regarding bisphosphonate type, clinical indication, administration route, treatment duration, triggering factors, study design, follow-up duration, type of APC, and outcomes adopted to evaluate treatment success [67]. Though the results of this review must be cautiously interpreted, since they are based on low evidence level studies, with limited sample size, they are suggestive of possible beneficial effects of APC when associated with surgical procedures for treatment or prevention of BRONJ. To confirm such indication, prospective comparative studies with a large sample size are urgently needed. Another subsequent systematic review on the same topic substantially confirmed these results, highlighting the need for well-done evidence-based comparative studies [68].

#### **3.13. Implants in patients assuming bisphosphonates**

In a multicenter study on a cohort of 235 middle-aged osteoporotic women under bisphosphonates therapy, the outcome of a total of 1267 implants was evaluated after a minimum follow-up of 24 months [69]. The implants were always placed in combination with P-PRP, used as a coating over implant surface at insertion (**Figure 6**). The main outcomes (adverse events) were implant failure and incidence of osteonecrosis of the jaw (ONJ). Sixteen implants were lost in 16 patients up to 120 months after placement, representing a survival of 98.7% on an implant basis and 93.2% on a patient basis [69] No cases of ONJ occurred. The results are

**Figure 6.** Multiple implant placements in an osteoporotic woman under bisphosphonates. PRGF membrane was used to

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Overall, the above results suggest that the use of procedures aimed at enhancing tissue healing, such as autologous platelet concentrates, may produce relevant benefits in patients at risk

The use of autologous platelet concentrates generally produced beneficial effects, though the level of evidence differs among various surgical procedures. In postextraction sockets and periodontal intrabony defects, the advantage of using APCs, both alone and combined with bone substitutes, is well documented by a number of randomized clinical studies. In maxillary sinus augmentation, controversial outcomes exist, due to few published controlled studies. Also, in endodontic surgery and implant treatment, there is a paucity of evidencebased studies, even though all show beneficial effects of APC. The variability among protocols and outcomes in different studies often prevents the possibility of performing meta-analysis and is thought to be related to the controversial results sometimes observed. Better soft tissue healing, improved patients' quality of life and reduced incidence of adverse events and

due to their compromised systemic condition and should be recommended.

perfectly in line with those of healthy patients.

cover the region involved from implant surgery.

**4. Conclusion**

**Figure 5.** (A) A case of an oncologic patient under bisphosphonate affected by osteonecrosis of the upper jaw, undergoing respective surgery for removal of the necrotic bone. (B) PRGF is placed within the region involved from resection. (C) One year after surgery, the region is completely healed.

**Figure 6.** Multiple implant placements in an osteoporotic woman under bisphosphonates. PRGF membrane was used to cover the region involved from implant surgery.

events) were implant failure and incidence of osteonecrosis of the jaw (ONJ). Sixteen implants were lost in 16 patients up to 120 months after placement, representing a survival of 98.7% on an implant basis and 93.2% on a patient basis [69] No cases of ONJ occurred. The results are perfectly in line with those of healthy patients.

Overall, the above results suggest that the use of procedures aimed at enhancing tissue healing, such as autologous platelet concentrates, may produce relevant benefits in patients at risk due to their compromised systemic condition and should be recommended.

## **4. Conclusion**

reduce the risk of hemorrhage and related complications in thrombocytopenic patients, due

Recently renamed medication-related osteonecrosis of the jaw (MRONJ), it is an adverse drug reaction consisting of progressive bone destruction in the maxillofacial region of patients under current or previous treatment with a bisphosphonate or another antiresorptive or antiangiogenic drug. Since APCs demonstrated to enhance bone and soft tissue healing in oral surgery procedures, it is reasonable to believe that they might provide benefits to these patients. A recent systematic review included 18 studies, reporting on 362 patients undergoing oral surgery in combination with APCs [67]. The adjunct of APC in the surgical treatment of necrotic bone removal significantly reduced osteonecrosis recurrence as compared to control (**Figure 5**). APC was associated with a reduced, though not significant, incidence of BRONJ after tooth extraction. Heterogeneity among studies was found regarding bisphosphonate type, clinical indication, administration route, treatment duration, triggering factors, study design, follow-up duration, type of APC, and outcomes adopted to evaluate treatment success [67]. Though the results of this review must be cautiously interpreted, since they are based on low evidence level studies, with limited sample size, they are suggestive of possible beneficial effects of APC when associated with surgical procedures for treatment or prevention of BRONJ. To confirm such indication, prospective comparative studies with a large sample size are urgently needed. Another subsequent systematic review on the same topic substantially confirmed these results, highlighting the need for well-done evidence-based

In a multicenter study on a cohort of 235 middle-aged osteoporotic women under bisphosphonates therapy, the outcome of a total of 1267 implants was evaluated after a minimum follow-up of 24 months [69]. The implants were always placed in combination with P-PRP, used as a coating over implant surface at insertion (**Figure 6**). The main outcomes (adverse

**Figure 5.** (A) A case of an oncologic patient under bisphosphonate affected by osteonecrosis of the upper jaw, undergoing respective surgery for removal of the necrotic bone. (B) PRGF is placed within the region involved from resection.

to the ability of stimulating healing and its hemostatic properties.

**3.12. Bisphosphonate-related osteonecrosis of the jaws**

**3.13. Implants in patients assuming bisphosphonates**

(C) One year after surgery, the region is completely healed.

comparative studies [68].

220 Biomaterials in Regenerative Medicine

The use of autologous platelet concentrates generally produced beneficial effects, though the level of evidence differs among various surgical procedures. In postextraction sockets and periodontal intrabony defects, the advantage of using APCs, both alone and combined with bone substitutes, is well documented by a number of randomized clinical studies. In maxillary sinus augmentation, controversial outcomes exist, due to few published controlled studies. Also, in endodontic surgery and implant treatment, there is a paucity of evidencebased studies, even though all show beneficial effects of APC. The variability among protocols and outcomes in different studies often prevents the possibility of performing meta-analysis and is thought to be related to the controversial results sometimes observed. Better soft tissue healing, improved patients' quality of life and reduced incidence of adverse events and complications are systematically reported when using APC. In conclusion, the use of such autologous products is recommended for improving predictability and patients' acceptance of treatment in oral surgery procedures.

[9] Pearl RM, Wustrack KO, Harbury C, Rubenstein E, Kaplan EN. Microvascular anastomosis using a blood product sealant-adhesive. Surgery, Gynecology & Obstetrics. 1977;**144**:

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[10] Silverberg GD, Harbury CB, Rubenstein E. A physiological sealant for cerebrospinal

[11] Fischer H. A method of suture-free anastomosis of nerve transplantation is being reported, using facial nerve as the example (author's transl). Laryngologie, Rhinologie,

[12] Knighton DR, Ciresi KF, Fiegel VD, Austin LL, Butler EL. Classification and treatment of chronic nonhealing wounds. Successful treatment with autologous platelet-derived

[13] Knighton DR, Doucette M, Fiegel VD, Ciresi K, Butler E, Austin L. The use of platelet derived wound healing formula in human clinical trials. Progress in Clinical and

[14] Knighton DR, Ciresi K, Fiegel VD, Schumerth S, Butler E, Cerra F. Stimulation of repair in chronic, nonhealing, cutaneous ulcers using platelet-derived wound healing formula.

[15] Whitman DH, Berry RL, Green DM. Platelet gel: An autologous alternative to fibrin glue with applications in oral and maxillofacial surgery. Journal of Oral and Maxillofacial

[16] Kingsley CS. Blood coagulation; evidence of an antagonist to factor VI in platelet-rich

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## **Author details**

Carmen Mortellaro<sup>1</sup> \* and Massimo Del Fabbro2,3

\*Address all correspondence to: carmen.mortellaro@med.uniupo.it

1 Department of Health Sciences, Avogadro University of Eastern Piedmont, Novara, Italy

2 Department of Biomedical, Surgical and Dental Sciences, Università degli Studi di Milano, Milano, Italy

3 IRCCS Istituto Ortopedico Galeazzi, Milano, Italy

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complications are systematically reported when using APC. In conclusion, the use of such autologous products is recommended for improving predictability and patients' acceptance

1 Department of Health Sciences, Avogadro University of Eastern Piedmont, Novara, Italy

2 Department of Biomedical, Surgical and Dental Sciences, Università degli Studi di Milano,

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**Chapter 9**

Provisional chapter

**Laser Processing of Silicon for Synthesis of Better**

DOI: 10.5772/intechopen.69856

Laser Processing of Silicon for Synthesis of Better

The increasing demand for new biomaterials and fabrication methods provides an opportunity for silicon to solve current challenges in the field. Laser processing is becoming more common as the public begins to understand its simplicity and value. When an abundant material is paired with a reliable and economic fabrication method, biomedical devices can be created and improved. In this chapter, different laser parameters of the Nd:YAG laser are investigated and the topographic and physical trends are analyzed. The biocompatibility is assessed for scanning speed, line spacing, overlap number, pulse frequency, and laser power with the use of simulated body fluid (SBF) and fibroblast culturing (NIH 3T3). Not only can nanosecond pulses increase the biocompatibility of silicon by generating silicon oxide nanofibers, but the substrate

Science fiction has motivated intelligent minds to enhance the quality of living for the last century. A well-known example in fantasy is bionic limbs controlled by the mind. Individuals who have lost or permanently damaged limbs can benefit from procuring an aesthetically pleasing and fully functioning bionic replacement to restore or improve their quality of life. The field of biomaterials engineering has been making monumental advances by producing devices such as biosensors, bioMEMS, and artificial hearts [1–3]. There is a continuous growth in population today, demanding the attention from biomedical fields to improve lifestyles and create better body functionality. Although current devices' interfaces with the human body have come a long way, there is still a long way to go in the fabrication methods of the required scaffolds.

> © The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

**Biomaterials**

Biomaterials

Abstract

1. Introduction

Candace Colpitts and Amirkianoosh Kiani

Candace Colpitts and Amirkianoosh Kiani

http://dx.doi.org/10.5772/intechopen.69856

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

becomes bioactive with the manipulation of cell interactions.

Keywords: silicon, nanofibers, Nd:YAG laser, fibroblasts

Provisional chapter

## **Laser Processing of Silicon for Synthesis of Better Biomaterials** Laser Processing of Silicon for Synthesis of Better

DOI: 10.5772/intechopen.69856

Biomaterials

Candace Colpitts and Amirkianoosh Kiani

Additional information is available at the end of the chapter Candace Colpitts and Amirkianoosh Kiani

http://dx.doi.org/10.5772/intechopen.69856 Additional information is available at the end of the chapter

#### Abstract

The increasing demand for new biomaterials and fabrication methods provides an opportunity for silicon to solve current challenges in the field. Laser processing is becoming more common as the public begins to understand its simplicity and value. When an abundant material is paired with a reliable and economic fabrication method, biomedical devices can be created and improved. In this chapter, different laser parameters of the Nd:YAG laser are investigated and the topographic and physical trends are analyzed. The biocompatibility is assessed for scanning speed, line spacing, overlap number, pulse frequency, and laser power with the use of simulated body fluid (SBF) and fibroblast culturing (NIH 3T3). Not only can nanosecond pulses increase the biocompatibility of silicon by generating silicon oxide nanofibers, but the substrate becomes bioactive with the manipulation of cell interactions.

Keywords: silicon, nanofibers, Nd:YAG laser, fibroblasts

## 1. Introduction

Science fiction has motivated intelligent minds to enhance the quality of living for the last century. A well-known example in fantasy is bionic limbs controlled by the mind. Individuals who have lost or permanently damaged limbs can benefit from procuring an aesthetically pleasing and fully functioning bionic replacement to restore or improve their quality of life. The field of biomaterials engineering has been making monumental advances by producing devices such as biosensors, bioMEMS, and artificial hearts [1–3]. There is a continuous growth in population today, demanding the attention from biomedical fields to improve lifestyles and create better body functionality. Although current devices' interfaces with the human body have come a long way, there is still a long way to go in the fabrication methods of the required scaffolds.

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

© The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

This chapter outlines one single pathway of research done to broaden the opportunities in the biomaterials industry. The bioactivity of laser-treated silicon is investigated through the use of in vitro testing. This research investigates of the trends of different laser parameters, including power, frequency, scanning speed, line spacing, and overlap number.

2. Laser processing and surface characterization

microscopy (FESEM) and 3D optical microscopy.

2.1. Laser system

laser-operating software.

2.2. Biocompatibility evaluation

2.3. Temperature evaluation

kept in an incubator at 36.5C for up to 6 weeks.

draq5 (1:10,000 dilution) overnight for fluorescence imaging.

Microfabrication with lasers is becoming increasingly popular in many industries including biomaterials [21]. Laser irradiation introduces surface irregularities and chemical changes to the silicon surface. The laser irradiates a simple line pattern onto pure silicon with <100> orientation. There are a number of methods used to analyze the condition of the laserprocessed substrat. Images of the samples are taken with field emission scanning electron

Laser Processing of Silicon for Synthesis of Better Biomaterials

http://dx.doi.org/10.5772/intechopen.69856

231

The laser used in this research is a SOL-20 1064 nm Nd:YAG nanosecond laser by Bright Solutions Inc. The JD2204 Sino Galvo two-axis Galvo scanner has an input of 10 mm and a beam displacement of 13.4 mm. The theoretically determined spot size diameter is 19 μm. The laser pulses can range from lengths of 6 to 35 ns, frequencies of 10 to 100 kHz, powers up to 20 W, and scanning speeds up to 3000 mm/s. For this research, scanning speeds of 100–1000 mm/s, powers of 7, 10, and 15 W, and frequencies of 25, 70, and 100 kHz are used. Line spacing varies from 0.025, 0.05, and 0.10 mm. Overlap number, or number of times the laser repeats the same pattern, varies from 1, 2, and 3. The manipulation of these parameters is easily executed through the

The biocompatibility of a material is influenced by surface roughness, reflectivity, and chemical content of the substrate. The chemical content is assessed using micro-Raman and energydispersive X-ray (EDX) analysis. The surface roughness is determined with the use of 3D optical microscopy, and the reflectivity is determined with light spectroscopy. The biocompatibility is also determined with the use of simulated body fluid, which is a form of in vitro assessment—testing done outside of the body. Simulated body fluid is a solution that mimics the ion concentration of human blood plasma. When a biocompatible material is submerged in the liquid, hydroxyapatite forms on the surface [22, 23]. The submerged samples in the SBF are

Cell interactions with the laser-processed silicon substrate are also evaluated with cultured mouse embryonic fibroblasts (NIH 3T3). Cells are seeded at 2400 cells/cm<sup>2</sup> in triplicate and incubated for 72 h at 37C. The samples are incubated under 5% CO2 in Dulbecco's modified Eagle medium (DMEM) supplemented with 10% heat-inactivated calf serum. Phosphate-buffered saline (PBS) was then used to rinse the nonadherent cells from the samples overnight at 4C. Staining was then done to the samples with phalloidin (1:2000 dilution) and

Different frequencies, powers, and pulse widths change the behavior of the laser pulses. Determining the temperature will help investigate the pulse energy and how it affects the

#### 1.1. Challenges in the biomaterials industry

The current challenges in the biomedical field include finding biocompatible and bioactive organic or inorganic materials and simplifying manufacturing processes. Devices that are implanted inside the body require materials that are biocompatible; the behavior of this material when interacting with the human body must not have any toxic effects and must perform a specific task. For a material to be bioactive, it requires to be biocompatible and have a biological effect and provoke a positive and controlled biological response. Current biocompatible materials that are commonly used are gold, titanium, polymers, bioceramics, and composites [4–8].

Silicon was chosen as the material for this research due to its abundance and semiconductor abilities. Microfabricated silicon is widely used today in the microelectronics and photovoltaics industry [9, 10]. Silicon in its pure form is not biocompatible [11, 12]. However, silicon can be packaged in a biocompatible material such as titanium [13, 14]. It has been found that porous silicon is biocompatible [11]. The current method used for creating a porous layer is etching with hydrochloric acid. Acid etching is a long process that requires the use of dangerous chemicals and is consequently environmentally friendly. The challenge that silicon faces in the biomaterials industry is to find a superior surface alteration method.

#### 1.2. Laser processing

Technology that easily controls and creates an accurate pattern on a microscale is required in the microelectronics industry. A good solution to this criterion is a laser, which has been commonly used for surface texturing of steel [15, 16]. It was found that this method of surface treatment allowed the generation of micropores with different characteristics. Unlike acid etching, a laser is great for the modification of silicon since it is very clean, high resolution, and controllability of intensity and depth of penetration. The Nd:YAG pulsed laser is a particularly good solution since it is cost-effective, stable, and has the required high power range. Another advantage to this method is that there are no chemicals involved, which eliminates the complex processes of preparation and environmental concerns. Above all, using a laser is a single-step process. The economic and simplistic benefits that are associated with this approach are valuable to the biomedical industry.

High-end picosecond and femtosecond pulsed laser systems have also been used for generation of porous silicon particles [17, 18]. In this research, it is found that a Nd:YAG nanosecond pulse laser can achieve the desired biocompatible silicon. The nanosecond laser is much more economical and commercially available than the faster pulse lasers. The nanosecond laser is also currently used in the medical industry for procedures such as eye and dental surgeries [19, 20].

## 2. Laser processing and surface characterization

Microfabrication with lasers is becoming increasingly popular in many industries including biomaterials [21]. Laser irradiation introduces surface irregularities and chemical changes to the silicon surface. The laser irradiates a simple line pattern onto pure silicon with <100> orientation. There are a number of methods used to analyze the condition of the laserprocessed substrat. Images of the samples are taken with field emission scanning electron microscopy (FESEM) and 3D optical microscopy.

#### 2.1. Laser system

This chapter outlines one single pathway of research done to broaden the opportunities in the biomaterials industry. The bioactivity of laser-treated silicon is investigated through the use of in vitro testing. This research investigates of the trends of different laser parameters, including

The current challenges in the biomedical field include finding biocompatible and bioactive organic or inorganic materials and simplifying manufacturing processes. Devices that are implanted inside the body require materials that are biocompatible; the behavior of this material when interacting with the human body must not have any toxic effects and must perform a specific task. For a material to be bioactive, it requires to be biocompatible and have a biological effect and provoke a positive and controlled biological response. Current biocompatible materials that are commonly used are gold, titanium, polymers, bioceramics, and

Silicon was chosen as the material for this research due to its abundance and semiconductor abilities. Microfabricated silicon is widely used today in the microelectronics and photovoltaics industry [9, 10]. Silicon in its pure form is not biocompatible [11, 12]. However, silicon can be packaged in a biocompatible material such as titanium [13, 14]. It has been found that porous silicon is biocompatible [11]. The current method used for creating a porous layer is etching with hydrochloric acid. Acid etching is a long process that requires the use of dangerous chemicals and is consequently environmentally friendly. The challenge that silicon faces in the

Technology that easily controls and creates an accurate pattern on a microscale is required in the microelectronics industry. A good solution to this criterion is a laser, which has been commonly used for surface texturing of steel [15, 16]. It was found that this method of surface treatment allowed the generation of micropores with different characteristics. Unlike acid etching, a laser is great for the modification of silicon since it is very clean, high resolution, and controllability of intensity and depth of penetration. The Nd:YAG pulsed laser is a particularly good solution since it is cost-effective, stable, and has the required high power range. Another advantage to this method is that there are no chemicals involved, which eliminates the complex processes of preparation and environmental concerns. Above all, using a laser is a single-step process. The economic and simplistic benefits that are associated with this

High-end picosecond and femtosecond pulsed laser systems have also been used for generation of porous silicon particles [17, 18]. In this research, it is found that a Nd:YAG nanosecond pulse laser can achieve the desired biocompatible silicon. The nanosecond laser is much more economical and commercially available than the faster pulse lasers. The nanosecond laser is also currently used in the medical industry for procedures such as eye and dental surger-

power, frequency, scanning speed, line spacing, and overlap number.

biomaterials industry is to find a superior surface alteration method.

approach are valuable to the biomedical industry.

1.1. Challenges in the biomaterials industry

composites [4–8].

230 Biomaterials in Regenerative Medicine

1.2. Laser processing

ies [19, 20].

The laser used in this research is a SOL-20 1064 nm Nd:YAG nanosecond laser by Bright Solutions Inc. The JD2204 Sino Galvo two-axis Galvo scanner has an input of 10 mm and a beam displacement of 13.4 mm. The theoretically determined spot size diameter is 19 μm. The laser pulses can range from lengths of 6 to 35 ns, frequencies of 10 to 100 kHz, powers up to 20 W, and scanning speeds up to 3000 mm/s. For this research, scanning speeds of 100–1000 mm/s, powers of 7, 10, and 15 W, and frequencies of 25, 70, and 100 kHz are used. Line spacing varies from 0.025, 0.05, and 0.10 mm. Overlap number, or number of times the laser repeats the same pattern, varies from 1, 2, and 3. The manipulation of these parameters is easily executed through the laser-operating software.

#### 2.2. Biocompatibility evaluation

The biocompatibility of a material is influenced by surface roughness, reflectivity, and chemical content of the substrate. The chemical content is assessed using micro-Raman and energydispersive X-ray (EDX) analysis. The surface roughness is determined with the use of 3D optical microscopy, and the reflectivity is determined with light spectroscopy. The biocompatibility is also determined with the use of simulated body fluid, which is a form of in vitro assessment—testing done outside of the body. Simulated body fluid is a solution that mimics the ion concentration of human blood plasma. When a biocompatible material is submerged in the liquid, hydroxyapatite forms on the surface [22, 23]. The submerged samples in the SBF are kept in an incubator at 36.5C for up to 6 weeks.

Cell interactions with the laser-processed silicon substrate are also evaluated with cultured mouse embryonic fibroblasts (NIH 3T3). Cells are seeded at 2400 cells/cm<sup>2</sup> in triplicate and incubated for 72 h at 37C. The samples are incubated under 5% CO2 in Dulbecco's modified Eagle medium (DMEM) supplemented with 10% heat-inactivated calf serum. Phosphate-buffered saline (PBS) was then used to rinse the nonadherent cells from the samples overnight at 4C. Staining was then done to the samples with phalloidin (1:2000 dilution) and draq5 (1:10,000 dilution) overnight for fluorescence imaging.

#### 2.3. Temperature evaluation

Different frequencies, powers, and pulse widths change the behavior of the laser pulses. Determining the temperature will help investigate the pulse energy and how it affects the surface topographic properties and chemical structure. The surface temperature is modeled using the two-dimensional heat equation in cylindrical coordinates. The heat equation in cylindrical coordinates is found in Eq. (1).

$$\frac{1}{r}\frac{\partial}{\partial r}\left(\kappa r \frac{\partial T}{\partial r}\right) + \frac{1}{r^2} \frac{\partial}{\partial \phi}\left(\kappa \frac{\partial T}{\partial \phi}\right) + \frac{\partial}{\partial z}\left(\kappa \frac{\partial T}{\partial z}\right) + \dot{q} = \rho c\_p \frac{\partial T}{\partial t} \tag{1}$$

where κ is the heat conduction coefficient, ρ is the material density (kg/m<sup>3</sup> ), cp is the specific heat (J/kgK), and q\_ is the rate at which energy is generated per unit volume of the medium (W/m<sup>3</sup> ) [24]. Since it is assumed that κ, ρ, and cp are constant, and there is no energy generation within the silicon (q\_ ¼ 0), Eq. (1) can be simplified into Eq. (2).

$$\frac{1}{r}\frac{\partial}{\partial r}\left(r\frac{\partial T}{\partial r}\right) + \frac{\partial}{\partial z}\left(\frac{\partial T}{\partial z}\right) = \frac{1}{a}\frac{\partial T}{\partial t} \tag{2}$$

Eqs. (6) and (7) will determine the ablation depth after a train of pulses. Each consecutive pulse

<sup>S</sup> <sup>¼</sup> v\_scan

where S is the spatial separation between each consecutive pulse, v\_scan is the scanning speed

The three scanning parameters discussed in this section are overlap number, line spacing, and scanning speed. Each of these is easily set through the laser-operating software. These parameters have a direct effect of the surface topography and oxidation levels of the silicon sub-

The field emission scanning electron microscope images are an effective way of investigating the physical results from the laser ablation on the silicon samples. Figure 1 shows the effect of different overlap numbers. At 1 overlap (OL), the line pattern is distinct and relatively clean. When the OL number is increased to 2, the line pattern is less definite and contains more irregularities. Finally, increasing the overlap number once again to 3, the line pattern is almost

The effect of different line spacing was then investigated with FESEM. The overlap number was kept constant at 1. At the largest line spacing of 0.10 mm, the line pattern is discrete. At the smallest line spacing of 0.025 mm, the amount of imperfections is high with no distinctive line

Knowing that a higher overlap number and a smaller line spacing made for the highest level of laser-ablated substrate, a sample with 0.025 mm line spacing and three overlaps was created to observe the surface characteristics. The high magnification FESEM image in Figure 3 of this sample shows a nanofibrous substrate. These nanofibers are the result of a high-energy reactive plume that forms on the surface during laser ablation [26]. The plume generates a heataffected zone that causes the silicon ions to react with the oxygen ions, creating these nanoscale

The effect of scanning speed was then investigated with 3D optical microscopy. The data of the results from scanning speeds of 100, 200, 500, 800, and 1000 mm/s are mapped and compared in Figure 4. It is expected that the lower scanning speeds have larger depths due to a higher number of pulses ablating the surface area. However, with a closer look at Figure 4, the lower scanning speeds have shallow depths and a relatively high wall of built-up material along the

hscanðrÞ ¼ hðrÞ þ hðS � rÞ (6)

<sup>f</sup> (7)

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adds to the penetration of the preceding pulse, S, resulting in a deeper groove.

in m/s, and f is the pulse frequency.

strates.

3.1. Topography analysis

SiO2 fibers [27–29].

3. Effect of scanning parameters

unrecognizable with a substantial amount of imperfections.

separation. The effect of line spacing can be seen in Figure 2.

where a = κ/ρcp, which is the thermal diffusivity (m2 /s). The boundary conditions for this equation include the initial temperature being room temperature, the pulse intensity is at its maximum during the pulse at z = 0, the intensity is zero between pulses, and the temperature change is zero at r = ∞, and z = ∞. With these conditions, the single-pulse temperature change of a high absorption material, ΔT, for a square pulse, can be obtained (Eq. (3)) [25].

$$
\Delta T(r, z, \tau) = \frac{I\_{\max} \mathcal{V} \sqrt{\kappa}}{\sqrt{\pi} K} \int\_0^t \frac{1}{\sqrt{t} \left[1 + \frac{8\kappa t}{\mathcal{W}^2}\right]} e^{-\left[\frac{z^2}{4\kappa t} + \frac{r^2}{4\kappa t + \frac{1}{\pi} \mathcal{W}^2}\right]} dt \tag{3}
$$

$$\text{where } I\_{\text{max}} = \frac{P\_{\text{peak}}}{A} = \frac{4P\_{\text{measured}}}{\pi d^2 f \tau} \tag{4}$$

where Imax (Eq. (4)) is the peak intensity which is the peak power divided by the spot area, Pmeasured is the experimental measured power, f is the pulse frequency, τ is the pulse duration, γ is equal to the Fresnel energy reflectivity (R) subtracted from 1 (1–R) with a R value of 0.325 and a <sup>γ</sup> value of 0.675, <sup>κ</sup> is Silicon's diffusivity with a value of 9.07�10�<sup>3</sup> m2 /s, K is Silicon's conductivity with a value of 155 W/mK, τ is the laser pulse duration, W is the beam's filed radius (1/e) with a value of 1.94�10�<sup>5</sup> m, <sup>z</sup> is the ablation pit depth, and <sup>r</sup> is the ablation pit radius. Using this equation with the assumption that the pulse is square-shaped, the temperature can be determined at the center of ablation (r = 0) and at the surface (z = 0). From Eq. (3), an analytical expression can be made to determine the depth of the ablated groove at the center of the ablation with respect to radius by using the mean value theorem.

$$h(r) = \sqrt{-4\kappa\tau \ln\left\{\frac{\beta \mathcal{K} \Delta T\_B}{\mathcal{Y} I\_{\text{max}}} \sqrt{\frac{\pi}{\kappa \beta \tau}} \left(1 + \frac{8\beta \kappa \tau}{\mathcal{W}^2}\right)\right\} - \frac{r^2}{1 + \frac{\mathcal{W}^2}{8\beta \kappa \tau}}} \le h(0) \tag{5}$$

where β is an experimentally determined correction factor of 0.5 and ΔTB is silicon's boiling temperature of 2972 K. A more accurate representation of the experimental results found in Eqs. (6) and (7) will determine the ablation depth after a train of pulses. Each consecutive pulse adds to the penetration of the preceding pulse, S, resulting in a deeper groove.

$$h\_{\text{scan}}(r) = h(r) + h(\mathbb{S} - r) \tag{6}$$

$$S = \frac{v\_- \text{scam}}{f} \tag{7}$$

where S is the spatial separation between each consecutive pulse, v\_scan is the scanning speed in m/s, and f is the pulse frequency.

## 3. Effect of scanning parameters

The three scanning parameters discussed in this section are overlap number, line spacing, and scanning speed. Each of these is easily set through the laser-operating software. These parameters have a direct effect of the surface topography and oxidation levels of the silicon substrates.

#### 3.1. Topography analysis

surface topographic properties and chemical structure. The surface temperature is modeled using the two-dimensional heat equation in cylindrical coordinates. The heat equation in

> ∂T ∂φ � �

þ ∂ ∂z κ ∂T ∂z � �

) [24]. Since it is assumed that κ, ρ, and cp are constant, and there is no energy generation

∂T ∂z � �

1 ffiffi <sup>t</sup> <sup>p</sup> <sup>1</sup> <sup>þ</sup> <sup>8</sup>κ<sup>t</sup> W<sup>2</sup> h i <sup>e</sup>

<sup>A</sup> <sup>¼</sup> <sup>4</sup>Pmeasured πd<sup>2</sup> f τ

¼ 1 a ∂T

> � <sup>z</sup><sup>2</sup> <sup>4</sup>κ<sup>t</sup> <sup>þ</sup> <sup>r</sup><sup>2</sup> <sup>4</sup>κtþ<sup>1</sup> 2 W2 h i

heat (J/kgK), and q\_ is the rate at which energy is generated per unit volume of the medium

þ ∂ ∂z

a high absorption material, ΔT, for a square pulse, can be obtained (Eq. (3)) [25].

κ p ffiffiffi <sup>π</sup> <sup>p</sup> <sup>K</sup>

where Imax <sup>¼</sup> Ppeak

and a <sup>γ</sup> value of 0.675, <sup>κ</sup> is Silicon's diffusivity with a value of 9.07�10�<sup>3</sup> m2

of the ablation with respect to radius by using the mean value theorem.

γImax

�4κτln <sup>β</sup>KΔTB

hðrÞ ¼

ðτ 0

where Imax (Eq. (4)) is the peak intensity which is the peak power divided by the spot area, Pmeasured is the experimental measured power, f is the pulse frequency, τ is the pulse duration, γ is equal to the Fresnel energy reflectivity (R) subtracted from 1 (1–R) with a R value of 0.325

conductivity with a value of 155 W/mK, τ is the laser pulse duration, W is the beam's filed radius (1/e) with a value of 1.94�10�<sup>5</sup> m, <sup>z</sup> is the ablation pit depth, and <sup>r</sup> is the ablation pit radius. Using this equation with the assumption that the pulse is square-shaped, the temperature can be determined at the center of ablation (r = 0) and at the surface (z = 0). From Eq. (3), an analytical expression can be made to determine the depth of the ablated groove at the center

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

1 þ

8βκτ W<sup>2</sup>

vuut <sup>≤</sup> <sup>h</sup>ð0<sup>Þ</sup> (5)

� <sup>r</sup><sup>2</sup> <sup>1</sup> <sup>þ</sup> <sup>W</sup><sup>2</sup> 8βκτ

ffiffiffiffiffiffiffiffi π κβτ r

� � � �

where β is an experimentally determined correction factor of 0.5 and ΔTB is silicon's boiling temperature of 2972 K. A more accurate representation of the experimental results found in

equation include the initial temperature being room temperature, the pulse intensity is at its maximum during the pulse at z = 0, the intensity is zero between pulses, and the temperature change is zero at r = ∞, and z = ∞. With these conditions, the single-pulse temperature change of

þ q\_ ¼ ρcp

∂T

<sup>∂</sup><sup>t</sup> (2)

/s). The boundary conditions for this

<sup>∂</sup><sup>t</sup> (1)

dt (3)

/s, K is Silicon's

(4)

), cp is the specific

cylindrical coordinates is found in Eq. (1).

232 Biomaterials in Regenerative Medicine

1 r ∂ ∂r κr ∂T ∂r � �

(W/m<sup>3</sup>

þ 1 r2 ∂ <sup>∂</sup><sup>φ</sup> <sup>κ</sup>

within the silicon (q\_ ¼ 0), Eq. (1) can be simplified into Eq. (2).

where a = κ/ρcp, which is the thermal diffusivity (m2

1 r ∂ ∂r r ∂T ∂r � �

<sup>Δ</sup>Tðr, z, <sup>τ</sup>Þ ¼ Imax<sup>γ</sup> ffiffiffi

where κ is the heat conduction coefficient, ρ is the material density (kg/m<sup>3</sup>

The field emission scanning electron microscope images are an effective way of investigating the physical results from the laser ablation on the silicon samples. Figure 1 shows the effect of different overlap numbers. At 1 overlap (OL), the line pattern is distinct and relatively clean. When the OL number is increased to 2, the line pattern is less definite and contains more irregularities. Finally, increasing the overlap number once again to 3, the line pattern is almost unrecognizable with a substantial amount of imperfections.

The effect of different line spacing was then investigated with FESEM. The overlap number was kept constant at 1. At the largest line spacing of 0.10 mm, the line pattern is discrete. At the smallest line spacing of 0.025 mm, the amount of imperfections is high with no distinctive line separation. The effect of line spacing can be seen in Figure 2.

Knowing that a higher overlap number and a smaller line spacing made for the highest level of laser-ablated substrate, a sample with 0.025 mm line spacing and three overlaps was created to observe the surface characteristics. The high magnification FESEM image in Figure 3 of this sample shows a nanofibrous substrate. These nanofibers are the result of a high-energy reactive plume that forms on the surface during laser ablation [26]. The plume generates a heataffected zone that causes the silicon ions to react with the oxygen ions, creating these nanoscale SiO2 fibers [27–29].

The effect of scanning speed was then investigated with 3D optical microscopy. The data of the results from scanning speeds of 100, 200, 500, 800, and 1000 mm/s are mapped and compared in Figure 4. It is expected that the lower scanning speeds have larger depths due to a higher number of pulses ablating the surface area. However, with a closer look at Figure 4, the lower scanning speeds have shallow depths and a relatively high wall of built-up material along the

Figure 1. FESEM images of pattern overlaps of 1, 2, and 3 with a line spacing of 0.10 mm, a laser power of 10.5 W, a frequency of 100 kHz, and a scanning speed of 400 mm/s.

Figure 2. FESEM images of line spacing of 0.025, 0.05, and 0.10 mm with an overlap number of 1, a laser power of 13.3 W, a frequency of 100 kHz, and a scanning speed of 400 mm/s.

sides of the groove. At higher scanning speeds, when the pulses are farther apart, there is less penetration on the surface, leading to a shallower groove. At lower scanning speeds, the hightemperature ablated material from the walls caves back into the deep groove and solidifies into a much shallower groove than initially dredged. Scanning speeds of 200 and 800 mm/s show a

Figure 5. Ablation depth after a train of pulses at difference scanning speeds at a power of 15 W and a frequency of 100 kHz.

Figure 4. Profile data from the 3D optical microscope for scanning speeds of 100, 200, 500, 800, and 1000 mm/s at an

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overlap number of 1, a power of 15 W, and a frequency of 100 kHz.

Eqs. (6) and (7) are then used to find the theoretical ablation depths at various scanning speeds. Observing the trend in Figure 5, it is clear that the ablation depth decreases with increasing scanning speed. Both the experimental data and theoretical results are in close agreement.

Each sample was submerged in simulated body fluid (SBF) for 6 weeks and kept at a constant temperature of 36.5C. The samples were then emerged from the fluid and assessed with energy-dispersive X-ray (EDX). The SBF-soaked samples were found to contain a traces of phosphorous and calcium, which is indicative of the presence of a bone-like apatite. Hydroxyapatite is formed by the nucleation of calcium phosphate ions [5, 27]. The silicon oxide layer created by the laser plume has a negative charge, which attracts the positively charged calcium phosphate. The resulting substrate contains this bone-like apatite, which was seen on the lasertreated silicon samples. The EDX results from the sample with three overlap and 0.025 mm line

deep groove with a smaller amount of material built up than the 100 mm/s sample.

3.2. Bioactivity assessment

spacing can be seen in Figure 6.

Figure 3. Presence of nanofibers detected on FESEM image of sample with three overlaps and 0.025 mm line spacing with power of 13.3 W, frequency of 100 kHz, and a scanning speed of 400 mm/s.

Figure 4. Profile data from the 3D optical microscope for scanning speeds of 100, 200, 500, 800, and 1000 mm/s at an overlap number of 1, a power of 15 W, and a frequency of 100 kHz.

Figure 5. Ablation depth after a train of pulses at difference scanning speeds at a power of 15 W and a frequency of 100 kHz.

sides of the groove. At higher scanning speeds, when the pulses are farther apart, there is less penetration on the surface, leading to a shallower groove. At lower scanning speeds, the hightemperature ablated material from the walls caves back into the deep groove and solidifies into a much shallower groove than initially dredged. Scanning speeds of 200 and 800 mm/s show a deep groove with a smaller amount of material built up than the 100 mm/s sample.

Eqs. (6) and (7) are then used to find the theoretical ablation depths at various scanning speeds. Observing the trend in Figure 5, it is clear that the ablation depth decreases with increasing scanning speed. Both the experimental data and theoretical results are in close agreement.

#### 3.2. Bioactivity assessment

Figure 1. FESEM images of pattern overlaps of 1, 2, and 3 with a line spacing of 0.10 mm, a laser power of 10.5 W, a

Figure 2. FESEM images of line spacing of 0.025, 0.05, and 0.10 mm with an overlap number of 1, a laser power of 13.3 W,

Figure 3. Presence of nanofibers detected on FESEM image of sample with three overlaps and 0.025 mm line spacing with

frequency of 100 kHz, and a scanning speed of 400 mm/s.

234 Biomaterials in Regenerative Medicine

a frequency of 100 kHz, and a scanning speed of 400 mm/s.

power of 13.3 W, frequency of 100 kHz, and a scanning speed of 400 mm/s.

Each sample was submerged in simulated body fluid (SBF) for 6 weeks and kept at a constant temperature of 36.5C. The samples were then emerged from the fluid and assessed with energy-dispersive X-ray (EDX). The SBF-soaked samples were found to contain a traces of phosphorous and calcium, which is indicative of the presence of a bone-like apatite. Hydroxyapatite is formed by the nucleation of calcium phosphate ions [5, 27]. The silicon oxide layer created by the laser plume has a negative charge, which attracts the positively charged calcium phosphate. The resulting substrate contains this bone-like apatite, which was seen on the lasertreated silicon samples. The EDX results from the sample with three overlap and 0.025 mm line spacing can be seen in Figure 6.

Figure 6. EDX results of sample with three overlaps, 0.025 mm line spacing, 400 mm/s, a power of 13.3 W, and a frequency of 400 mm/s.

This presence of bone-like apatite confirms that the biocompatibility of pure silicon was enhanced with nanosecond laser pulses. A smaller line spacing and higher overlap number generates more SiO2 nanofibers, which provides a favorable site for the nucleation of apatite.

## 4. Effect of frequency

The range of frequencies used in this section is 25, 70, and 100 kHz. For these experiments, the scanning speed was kept constant at 100 mm/s, the power at 15W, and the overlap number at 1.

#### 4.1. Topography analysis

Figures 7 and 8 show the topography changes in the frequency samples. A lower frequency produces a wider and shallower groove, while the higher frequencies yield a thinner yet deeper ablated groove. The theoretical results in for a single pulse in Figure 9 show that the groove increases in depth and decreases in width as frequency increases, which is in close agreement with the experimental results.

frequencies. Each recurring pulse adds to the penetration of preceding pulse. Higher frequencies have more pulses, resulting in a deeper penetration of the surface, which develops a trench

Figure 9. Theoretical profile data for a single pulse for frequencies of 25, 70, and 100 kHz at an overlap number of 1, a

Figure 7. 3D optical microscopy images of samples with frequencies of 25, 70, and 100 kHz at a power of 15 W, a scanning

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Figure 8. Experimental profile data from the 3D optical microscope for frequencies of 25, 70, and 100 kHz at an overlap

Mouse embryonic fibroblast cell interactions were examined for each frequency. The cell count in Figure 11 establishes that there are a higher number of cells in the higher frequency grooves.

with a larger depth.

4.3. Bioactivity assessment

power of 15 W, and a scanning speed of 100 mm/s.

speed of 100 mm/s, and 1 overlap.

number of 1, a power of 15 W, and a scanning speed of 100 mm/s.

#### 4.2. Temperature analysis

The temperature is determined for each frequency with Eq. (3) and can be found in Figure 10. The higher temperatures are associated with the lower frequencies on both z-axis and r-axis. By increasing the frequency, the pulse energy decreases which results in lower temperatures and a smaller heat-affected zone [30]. Due to reduced size of the heat-affected zone, the shape of the groove consequently changes size as well. This results in the thinner grooves at higher

Figure 7. 3D optical microscopy images of samples with frequencies of 25, 70, and 100 kHz at a power of 15 W, a scanning speed of 100 mm/s, and 1 overlap.

Figure 8. Experimental profile data from the 3D optical microscope for frequencies of 25, 70, and 100 kHz at an overlap number of 1, a power of 15 W, and a scanning speed of 100 mm/s.

Figure 9. Theoretical profile data for a single pulse for frequencies of 25, 70, and 100 kHz at an overlap number of 1, a power of 15 W, and a scanning speed of 100 mm/s.

frequencies. Each recurring pulse adds to the penetration of preceding pulse. Higher frequencies have more pulses, resulting in a deeper penetration of the surface, which develops a trench with a larger depth.

#### 4.3. Bioactivity assessment

This presence of bone-like apatite confirms that the biocompatibility of pure silicon was enhanced with nanosecond laser pulses. A smaller line spacing and higher overlap number generates more SiO2 nanofibers, which provides a favorable site for the nucleation of apatite.

Figure 6. EDX results of sample with three overlaps, 0.025 mm line spacing, 400 mm/s, a power of 13.3 W, and a

The range of frequencies used in this section is 25, 70, and 100 kHz. For these experiments, the scanning speed was kept constant at 100 mm/s, the power at 15W, and the overlap number at 1.

Figures 7 and 8 show the topography changes in the frequency samples. A lower frequency produces a wider and shallower groove, while the higher frequencies yield a thinner yet deeper ablated groove. The theoretical results in for a single pulse in Figure 9 show that the groove increases in depth and decreases in width as frequency increases, which is in close

The temperature is determined for each frequency with Eq. (3) and can be found in Figure 10. The higher temperatures are associated with the lower frequencies on both z-axis and r-axis. By increasing the frequency, the pulse energy decreases which results in lower temperatures and a smaller heat-affected zone [30]. Due to reduced size of the heat-affected zone, the shape of the groove consequently changes size as well. This results in the thinner grooves at higher

4. Effect of frequency

frequency of 400 mm/s.

236 Biomaterials in Regenerative Medicine

4.1. Topography analysis

4.2. Temperature analysis

agreement with the experimental results.

Mouse embryonic fibroblast cell interactions were examined for each frequency. The cell count in Figure 11 establishes that there are a higher number of cells in the higher frequency grooves.

5. Effect of power

5.1. Topography analysis

mm/s, and 1 overlap.

was 1.

Laser power immensely influences the surface properties when treatment is done to a material. The power for this section varies from 7, 10 to 15 W. For these experiments, the frequency was kept constant at 100 kHz, the scanning speed was set to 400 mm/s, and the overlap number

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The FESEM images of each power sample are shown in Figure 12. The experimental 3D optical microscopy profile data is shown in Figure 13. These results show that at higher powers, the groove will increase in both width and depth. Unlike the frequency trends, the size of the heataffected zone increases with power. The theoretical single-pulse depths found with Eq. (5) are shown in Figure 14. These results also show that the groove width and depth increase with

Figure 12. FESEM images of samples with powers of 7, 10, and 15 W at a frequency of 100 kHz, a scanning speed of 400

Figure 13. Experimental profile data from the 3D optical microscope for powers of 7, 10, and 15 W at an overlap number

of 1, a frequency of 100 kHz, and a scanning speed of 400 mm/s.

power and are in close agreement with the experimental results.

Figure 10. The single-pulse temperature on the surface of the silicon with respect to radius.

Figure 11. The number of cells within the laser-treated groove for each frequency as well as the amount of cells outside the groove within 100 μm from the edge of the groove [30].

The cells show a strong preference for the treated areas and show avoidance in the untreated areas. There is also a presence of fibronectin within the cells, which is an ECM protein secreted during embryonic development and wound healing, potentially leading to collagen deposition and tissue morphogenesis [30, 31]. These results confirm that the biocompatibility is enhanced with higher frequencies.

## 5. Effect of power

Laser power immensely influences the surface properties when treatment is done to a material. The power for this section varies from 7, 10 to 15 W. For these experiments, the frequency was kept constant at 100 kHz, the scanning speed was set to 400 mm/s, and the overlap number was 1.

#### 5.1. Topography analysis

The FESEM images of each power sample are shown in Figure 12. The experimental 3D optical microscopy profile data is shown in Figure 13. These results show that at higher powers, the groove will increase in both width and depth. Unlike the frequency trends, the size of the heataffected zone increases with power. The theoretical single-pulse depths found with Eq. (5) are shown in Figure 14. These results also show that the groove width and depth increase with power and are in close agreement with the experimental results.

The cells show a strong preference for the treated areas and show avoidance in the untreated areas. There is also a presence of fibronectin within the cells, which is an ECM protein secreted during embryonic development and wound healing, potentially leading to collagen deposition and tissue morphogenesis [30, 31]. These results confirm that the biocompatibility is enhanced

Figure 11. The number of cells within the laser-treated groove for each frequency as well as the amount of cells outside

Figure 10. The single-pulse temperature on the surface of the silicon with respect to radius.

with higher frequencies.

238 Biomaterials in Regenerative Medicine

the groove within 100 μm from the edge of the groove [30].

Figure 12. FESEM images of samples with powers of 7, 10, and 15 W at a frequency of 100 kHz, a scanning speed of 400 mm/s, and 1 overlap.

Figure 13. Experimental profile data from the 3D optical microscope for powers of 7, 10, and 15 W at an overlap number of 1, a frequency of 100 kHz, and a scanning speed of 400 mm/s.

Figure 14. Theoretical profile data for a single pulse for powers of 7, 10, and 15 W with an overlap number of 1, a frequency of 100 kHz, and a scanning speed of 400 mm/s.

ablated areas, which contains residual stress. When the power is increased, there is a larger transfer energy, resulting in a larger stress zone. These residual stress zones contain mismatched

Laser Processing of Silicon for Synthesis of Better Biomaterials

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241

Knowing that the power level of the laser pulse can control the residual stress zone size and cell behavior, this research can provide opportunities in cell manipulation and cell programming.

This chapter aims to introduce nanosecond laser processing for the enhancement of biocompatibility of pure silicon for various biomedical technologies. These results can contribute to the design of manufacturing processes of innovative biomedical devices to enhance the quality of living for a number of individuals. This research investigates the trends of various laser parameters including three scanning parameters (line spacing, overlap number, and scanning speed), pulse frequency, and laser power. Biocompatible in vitro assessment was conducted through the use of simulated body fluid (SBF) and cell culturing with NIH 3T3 fibroblasts. Samples with smaller line spacing and higher overlap numbers showed more generation of SiO2 nanofibers, which were shown to be biocompatible under SBF assessment. Biocompatibility increased with frequency due to the SiO2 being more prominent on high frequency samples and containing more fibroblast cell proliferation. Fibroblasts also showed preference to higher powers. However, the heat-affected zone immediately outside the ablated areas showed a mismatch of crystal orientations causing residual stress. These stress zones were avoided by cells, which led

to promising results for the potential in cell programming and manipulation.

Silicon Hall: Laser Micro/Nanofabrication Facility, Department of Mechanical Engineering,

crystal orientations due to tensile stresses causing crystal distortion.

Figure 16. Fibroblast cells avoiding the zones beside the grooves for powers of 7, 10, and 15 W.

6. Summary

Author details

Candace Colpitts and Amirkianoosh Kiani\*

University of New Brunswick, NB, Canada

\*Address all correspondence to: a.kiani@unb.ca

#### 5.2. Temperature analysis

The temperature is determined for each power with Eq. (3) and can be found in Figure 15. As expected, the higher temperatures are found with higher powers. At lower powers, the heataffected zone is smaller, allowing for both a thinner and shallower groove. When the pulse power is increased, the temperatures in the high-density plume are increased, causing more generation of the SiO2 nanofibers [26].

#### 5.3. Bioactivity assessment

Samples were once again assessed with fibroblast culturing for each power. When viewing the cell interactions under the microscope, cells were accumulated inside the laser-treated area as expected. Interestingly, the cell count was low directly beside the grooves and began to become more concentrated farther away from the edge of the groove. The fibroblasts avoided the zones immediately beside the grooves on each side. This can be seen in Figure 16.

This phenomenon is a result from the shockwave that is generated from the high-energy plume during laser ablation [26]. The shockwave transfers energy to the surface with results in high intensity thermal stress. The thermal shock causes a small zone directly beside the

Figure 15. The single-pulse temperature on the surface of the silicon with respect to radius. Tb is the boiling temperature of silicon at 3538 K.

Figure 16. Fibroblast cells avoiding the zones beside the grooves for powers of 7, 10, and 15 W.

ablated areas, which contains residual stress. When the power is increased, there is a larger transfer energy, resulting in a larger stress zone. These residual stress zones contain mismatched crystal orientations due to tensile stresses causing crystal distortion.

Knowing that the power level of the laser pulse can control the residual stress zone size and cell behavior, this research can provide opportunities in cell manipulation and cell programming.

#### 6. Summary

5.2. Temperature analysis

240 Biomaterials in Regenerative Medicine

5.3. Bioactivity assessment

of silicon at 3538 K.

generation of the SiO2 nanofibers [26].

frequency of 100 kHz, and a scanning speed of 400 mm/s.

The temperature is determined for each power with Eq. (3) and can be found in Figure 15. As expected, the higher temperatures are found with higher powers. At lower powers, the heataffected zone is smaller, allowing for both a thinner and shallower groove. When the pulse power is increased, the temperatures in the high-density plume are increased, causing more

Figure 14. Theoretical profile data for a single pulse for powers of 7, 10, and 15 W with an overlap number of 1, a

Samples were once again assessed with fibroblast culturing for each power. When viewing the cell interactions under the microscope, cells were accumulated inside the laser-treated area as expected. Interestingly, the cell count was low directly beside the grooves and began to become more concentrated farther away from the edge of the groove. The fibroblasts avoided the zones

This phenomenon is a result from the shockwave that is generated from the high-energy plume during laser ablation [26]. The shockwave transfers energy to the surface with results in high intensity thermal stress. The thermal shock causes a small zone directly beside the

Figure 15. The single-pulse temperature on the surface of the silicon with respect to radius. Tb is the boiling temperature

immediately beside the grooves on each side. This can be seen in Figure 16.

This chapter aims to introduce nanosecond laser processing for the enhancement of biocompatibility of pure silicon for various biomedical technologies. These results can contribute to the design of manufacturing processes of innovative biomedical devices to enhance the quality of living for a number of individuals. This research investigates the trends of various laser parameters including three scanning parameters (line spacing, overlap number, and scanning speed), pulse frequency, and laser power. Biocompatible in vitro assessment was conducted through the use of simulated body fluid (SBF) and cell culturing with NIH 3T3 fibroblasts. Samples with smaller line spacing and higher overlap numbers showed more generation of SiO2 nanofibers, which were shown to be biocompatible under SBF assessment. Biocompatibility increased with frequency due to the SiO2 being more prominent on high frequency samples and containing more fibroblast cell proliferation. Fibroblasts also showed preference to higher powers. However, the heat-affected zone immediately outside the ablated areas showed a mismatch of crystal orientations causing residual stress. These stress zones were avoided by cells, which led to promising results for the potential in cell programming and manipulation.

#### Author details

Candace Colpitts and Amirkianoosh Kiani\*

\*Address all correspondence to: a.kiani@unb.ca

Silicon Hall: Laser Micro/Nanofabrication Facility, Department of Mechanical Engineering, University of New Brunswick, NB, Canada

## References

[1] Turner AP, Pickup JC. Diabetes mellitus: Biosensors for research and management. Biosensors. 1985;1(1):85-115

[17] Bonse J, et al. Femtosecond laser ablation of silicon–modification thresholds and mor-

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[18] Kiani A, et al. Leaf-like nanotips synthesized on femtosecond laser-irradiated dielectric

[19] Wood JP, et al. Nanosecond pulse lasers for retinal applications. Lasers in Surgery and

[20] Al-Hadeethi Y, et al. Data fitting to study ablated hard dental tissues by nanosecond laser

[21] Khademhosseini A, et al. Microscale technologies for tissue engineering and biology. Proceedings of the National Academy of Sciences of the United States of America.

[22] Canham LT. Bioactive silicon structure fabrication through nanoetching techniques.

[23] Kamitakahara M, et al. Bioactivity and mechanical properties of polydimethylsiloxane (PDMS)-CaO-SiO2 hybrids with different PDMS contents. The Journal of Sol-Gel Science

[24] Incropera FP, et al. Introduction to convection. In: Anonymous, editors. Introduction to

[25] Hendow ST, Shakir SA. Structuring materials with nanosecond laser pulses. Optics

[26] Tavangar A, Tan B, Venkatakrishnan K. Study of the formation of 3-D titania nanofibrous structure by MHz femtosecond laser in ambient air. The Journal of Applied Physics.

[27] Tavangar A, Tan B, Venkatakrishnan K. The influence of laser-induced 3-D titania nanofibrous platforms on cell behavior. Journal of biomedical nanotechnology. 2013;9(11):1837-1846 [28] Kiani A, Venkatakrishnan K, Tan B. Optical absorption enhancement in 3D nanofibers coated on polymer substrate for photovoltaic devices. Optics Express. 2015;23(11):A569-A575 [29] Colpitts C, Kiani A. Synthesis of bioactive three-dimensional silicon-oxide nanofibrous structures on the silicon substrate for bionic devices' fabrication. Nanotechnology and

[30] Colpitts C, et al. Mammalian fibroblast cells show strong preference for laser-generated hybrid amorphous silicon-SiO2 textures. Journal of Applied Biomaterials and Functional

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[1] Turner AP, Pickup JC. Diabetes mellitus: Biosensors for research and management. Bio-

[2] Grayson AR, et al. A BioMEMS review: MEMS technology for physiologically integrated

[3] Copeland JG, et al. Cardiac replacement with a total artificial heart as a bridge to trans-

[4] Li X, et al. The use of nanoscaled fibers or tubes to improve biocompatibility and bioac-

[5] Radmanesh M, Kiani A. Effects of laser pulse numbers on surface biocompatibility of titanium for implant fabrication. Journal of Biomaterials and Nanobiotechnology. 2015;

[6] Swetha M, et al. Biocomposites containing natural polymers and hydroxyapatite for bone tissue engineering. International Journal of Biological Macromolecules. 2010;47(1):1-4 [7] Yue Z, et al. Controlled delivery for neuro-bionic devices. Advanced Drug Delivery

[8] Vallet-Regí M, Colilla M, González B. Medical applications of organic–inorganic hybrid materials within the field of silica-based bioceramics. Chemical Society Reviews. 2011;40(2):

[9] Green ML, et al. Ultrathin (<4 nm) SiO2 and Si–O–N gate dielectric layers for silicon microelectronics: Understanding the processing, structure, and physical and electrical limits. The Journal of Applied Physics. 2001;90(5):2057-2121. DOI: http://dx.doi.org/10.1063/1.1385803

[10] Jeong S, et al. Hybrid silicon nanocone-polymer solar cells. Nano Letters. 2012;12(6): 2971-2976. Available: http://dx.doi.org/10.1021/nl300713x. DOI: 10.1021/nl300713x [11] Buckberry L, Baylis S. Porous silicon as a biomaterial. Materials World. 1999;7:213-215 [12] Shaoqiang C, et al. Hydroxyapatite coating on porous silicon substrate obtained by

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tivity of biomedical materials. Journal of Nanomaterials. 2013;2013:14


**Chapter 10**

**Provisional chapter**

**Measurement and Simulation of Permeation and**

**Measurement and Simulation of Permeation and** 

Hao-Hsiang Hsu, Katharina Schimek, Uwe Marx and

Hao-Hsiang Hsu, Katharina Schimek, Uwe Marx and

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.72904

Ralf Pörtner

**Abstract**

and discussed.

**1. Introduction**

an interesting tool.

Ralf Pörtner

**Diffusion in Native and Cultivated Tissue Constructs**

Characterization of native skin or cultured 3D skin models with respect to permeability plays an important role for the development and testing of pharmaceuticals and cosmetics. Extensive efforts have been dedicated to determining the key parameters describing permeability and diffusion. Whereas respective methods are well established for native skin biopsies, only few are available for 3D skin models, as these have usually much lower dimensions. In this chapter, some fundamentals about permeation and diffusion as well as state of the art of measurement methods used for skin biopsies are summarized. An alternative method for the determination of the permeation in a membrane insert system and the use of a modular simulation to support permeability studies is presented

**Keywords:** skin models, permeation, diffusion, membrane insert system

**Diffusion in Native and Cultivated Tissue Constructs**

DOI: 10.5772/intechopen.72904

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Animal experiments are still a common method of testing drugs and also for skin, which is ethically controversial, cost-intensive and time-consuming. The fact is that drug testing on

Permeability studies are indispensable to characterize the transport of substances through the skin, either natural skin or cultivated 3D skin models. This is evident for dermal drug delivery systems, where drugs can be applied in the form of creams or patches on the skin. Furthermore, permeability studies play an important role in toxicity tests applied in drug development as well as substance testing. In this respect, human three-dimensional skin models have become

**Provisional chapter**

## **Measurement and Simulation of Permeation and Diffusion in Native and Cultivated Tissue Constructs Diffusion in Native and Cultivated Tissue Constructs**

**Measurement and Simulation of Permeation and** 

DOI: 10.5772/intechopen.72904

Hao-Hsiang Hsu, Katharina Schimek, Uwe Marx and Ralf Pörtner Ralf Pörtner Additional information is available at the end of the chapter

Hao-Hsiang Hsu, Katharina Schimek, Uwe Marx and

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.72904

#### **Abstract**

Characterization of native skin or cultured 3D skin models with respect to permeability plays an important role for the development and testing of pharmaceuticals and cosmetics. Extensive efforts have been dedicated to determining the key parameters describing permeability and diffusion. Whereas respective methods are well established for native skin biopsies, only few are available for 3D skin models, as these have usually much lower dimensions. In this chapter, some fundamentals about permeation and diffusion as well as state of the art of measurement methods used for skin biopsies are summarized. An alternative method for the determination of the permeation in a membrane insert system and the use of a modular simulation to support permeability studies is presented and discussed.

**Keywords:** skin models, permeation, diffusion, membrane insert system

## **1. Introduction**

Permeability studies are indispensable to characterize the transport of substances through the skin, either natural skin or cultivated 3D skin models. This is evident for dermal drug delivery systems, where drugs can be applied in the form of creams or patches on the skin. Furthermore, permeability studies play an important role in toxicity tests applied in drug development as well as substance testing. In this respect, human three-dimensional skin models have become an interesting tool.

Animal experiments are still a common method of testing drugs and also for skin, which is ethically controversial, cost-intensive and time-consuming. The fact is that drug testing on

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons

human skin is more efficient in comparison to animal skin just like rat, mouse and guinea pig [1, 2]. Furthermore, in 2013 the European Regulation (EC) No 1223/2009 entered into force, which prohibits animal experiments for cosmetics products. Artificial skin models based on human cells are intended to replace animal experiments. Especially the barrier function between artificial and human skin can differ, so permeation and diffusion investigation in this area is necessary.

most difficult way for substances, because they should have lipophilic and hydrophilic prop-

Measurement and Simulation of Permeation and Diffusion in Native and Cultivated Tissue…

http://dx.doi.org/10.5772/intechopen.72904

247

Alternatively, there is a way through the intercellular spaces between the cells. This is called the intercellular route. The intermediate space consists of cholesterols, ceramides and free fatty acids [7]. Because of the fatty acids, lipophilic substances can pass easier through the intercellular route in contrast to lipophobic substances [8]. Another barrier in this intercellular space is the tight junctions [9]. These are networks of strands which are formed by membrane proteins connecting cells. They are located between the keratinocytes in the stratum granulosum of the epidermis. An important function of the tight junctions, which are formed during

The appendageal route describes the penetration of a substance through skin appendages like hair follicles and glands. Since hair follicles and glands build only a small part of the human skin, their relevance for skin permeation was neglected for a long time. Its importance was shown recently by a researcher as permeation is better in a skin area containing hair follicles and glands in comparison to an area without them. A specific characteristic of the hair follicle is it reservoir function. In follicles, substances can be stored up to 10 days and can penetrate gradually into the skin. This aspect is interesting for drug delivery over the skin. For example, alcohol prefers the appendageal penetration route, as it would otherwise evaporate quite fast

So far the penetration and permeation in the skin was described. To get more in detail, the physical aspect of diffusion and permeation will be explained. Diffusion is a transport process where molecules move via Brownian motion in a volume or area. It is driven by the concentration gradient in the direction from higher to lower concentration [15]. Adolf Fick (1829–1901) verified the coherence between heat transfer and diffusion, which led to the Fick's first law of diffusion:

According to Eq. (1), the flux *F* in the one partial direction *x* is proportional to the gradient of

Permeation is an aspect of diffusion. Whereas diffusion is related to the movement of molecules in a system, permeation describes how fast molecules move through a system. An

through a membrane with a surface *A*. The acceptor concentration *cA* of the permeating substance on the other side can be detected over the time *t*. Eq. (2) for the permeation coefficient *P*

*dt* <sup>=</sup> *<sup>P</sup>* <sup>∙</sup> *<sup>A</sup> <sup>c</sup>*

the preferred parameter, as it is easier to measure compared to the diffusion coefficient. Due to the different layers of the skin, permeation and diffusion coefficient changes all the time

\_\_\_*D VA*

\_\_\_*<sup>A</sup>*

*<sup>D</sup>* ≫ *c A* <sup>∂</sup>*<sup>x</sup>* (1)

. With respect to skin, the permeation coefficient is

and a donor concentration *cD*

(2)

erties. Until now, it is not clear if hydrophilic substances choose this pathway [6].

the differentiation of keratinocytes, is to protect the skin from water loss [10].

on the surface of the skin. [11–14]

*F* = −*D* \_\_\_ <sup>∂</sup>*<sup>C</sup>*

concentration *C*. *D* is the diffusion coefficient or diffusivity. [16, 17]

example is the permeation of a substance within a volume *VA*

can be derived from the Fick's law of diffusion [18, 19]:

*dc*

This equation can only be used for *c*

from layer to layer.

Most methods for determination of diffusion and permeability parameters have been developed for large biopsies and can hardly be applied for small-scale 3D skin tissue cultures. But this is indispensable, if multi-well test systems with several samples run in parallel or multiorgan-chips are applied. Therefore, this chapter will first give a brief overview of standard methods used to investigate permeation and diffusion on the skin. Then an alternative method based on skin tissue cultures in a membrane insert system is introduced. By this, the permeation coefficient of substances through a skin-tissue barrier can be determined. The diffusion coefficient is estimated via parameter optimization performed in COMSOL Multiphysics. This software tool helps to describe the physical effects of the experimental set-ups more precisely and can further be used to reduce the required amount of experiments significantly.

## **2. Penetration, diffusion and permeation through the skin**

Penetration describes the entering of a substance into the skin. The entering process and depth of substance penetration through the skin depends on the physical and chemical character of the substance and the skin. The permeation of the skin is the pathway of a substance from the surface to the blood vessel. From a scientific point of view, it is the permeation of a substance through the skin layers. Diffusion is the physical process of randomized particle movement. If a concentration gradient exists, the particles move in the direction of lower concentration.

The human skin consists of three layers, the epidermis, dermis and subcutis. The stratum corneum (horny layer) is the upper layer of the epidermis and forms the main barrier of the skin. Substances such as drugs and chemicals penetrate through the skin barrier in three possible routes: the transcellular, the intercellular and the appendageal route (see **Figure 1**) [3–5].

The transcellular route leads the permeating substances directly through the cells. Here, the substances have to pass alternating lipophilic and hydrophilic layers. This is probably the

**Figure 1.** Schematic illustration of the three possible pathways of a permeating substance through the stratum corneum.

most difficult way for substances, because they should have lipophilic and hydrophilic properties. Until now, it is not clear if hydrophilic substances choose this pathway [6].

human skin is more efficient in comparison to animal skin just like rat, mouse and guinea pig [1, 2]. Furthermore, in 2013 the European Regulation (EC) No 1223/2009 entered into force, which prohibits animal experiments for cosmetics products. Artificial skin models based on human cells are intended to replace animal experiments. Especially the barrier function between artificial and human skin can differ, so permeation and diffusion investigation in

Most methods for determination of diffusion and permeability parameters have been developed for large biopsies and can hardly be applied for small-scale 3D skin tissue cultures. But this is indispensable, if multi-well test systems with several samples run in parallel or multiorgan-chips are applied. Therefore, this chapter will first give a brief overview of standard methods used to investigate permeation and diffusion on the skin. Then an alternative method based on skin tissue cultures in a membrane insert system is introduced. By this, the permeation coefficient of substances through a skin-tissue barrier can be determined. The diffusion coefficient is estimated via parameter optimization performed in COMSOL Multiphysics. This software tool helps to describe the physical effects of the experimental set-ups more precisely

and can further be used to reduce the required amount of experiments significantly.

Penetration describes the entering of a substance into the skin. The entering process and depth of substance penetration through the skin depends on the physical and chemical character of the substance and the skin. The permeation of the skin is the pathway of a substance from the surface to the blood vessel. From a scientific point of view, it is the permeation of a substance through the skin layers. Diffusion is the physical process of randomized particle movement. If a concentration gradient exists, the particles move in the direction of lower concentration. The human skin consists of three layers, the epidermis, dermis and subcutis. The stratum corneum (horny layer) is the upper layer of the epidermis and forms the main barrier of the skin. Substances such as drugs and chemicals penetrate through the skin barrier in three possible routes: the transcellular, the intercellular and the appendageal route (see **Figure 1**) [3–5]. The transcellular route leads the permeating substances directly through the cells. Here, the substances have to pass alternating lipophilic and hydrophilic layers. This is probably the

**Figure 1.** Schematic illustration of the three possible pathways of a permeating substance through the stratum corneum.

**2. Penetration, diffusion and permeation through the skin**

this area is necessary.

246 Biomaterials in Regenerative Medicine

Alternatively, there is a way through the intercellular spaces between the cells. This is called the intercellular route. The intermediate space consists of cholesterols, ceramides and free fatty acids [7]. Because of the fatty acids, lipophilic substances can pass easier through the intercellular route in contrast to lipophobic substances [8]. Another barrier in this intercellular space is the tight junctions [9]. These are networks of strands which are formed by membrane proteins connecting cells. They are located between the keratinocytes in the stratum granulosum of the epidermis. An important function of the tight junctions, which are formed during the differentiation of keratinocytes, is to protect the skin from water loss [10].

The appendageal route describes the penetration of a substance through skin appendages like hair follicles and glands. Since hair follicles and glands build only a small part of the human skin, their relevance for skin permeation was neglected for a long time. Its importance was shown recently by a researcher as permeation is better in a skin area containing hair follicles and glands in comparison to an area without them. A specific characteristic of the hair follicle is it reservoir function. In follicles, substances can be stored up to 10 days and can penetrate gradually into the skin. This aspect is interesting for drug delivery over the skin. For example, alcohol prefers the appendageal penetration route, as it would otherwise evaporate quite fast on the surface of the skin. [11–14]

So far the penetration and permeation in the skin was described. To get more in detail, the physical aspect of diffusion and permeation will be explained. Diffusion is a transport process where molecules move via Brownian motion in a volume or area. It is driven by the concentration gradient in the direction from higher to lower concentration [15]. Adolf Fick (1829–1901) verified the coherence between heat transfer and diffusion, which led to the Fick's first law of diffusion:

$$F = -D\frac{\partial \mathbb{C}}{\partial \mathbf{x}}\tag{1}$$

According to Eq. (1), the flux *F* in the one partial direction *x* is proportional to the gradient of concentration *C*. *D* is the diffusion coefficient or diffusivity. [16, 17]

Permeation is an aspect of diffusion. Whereas diffusion is related to the movement of molecules in a system, permeation describes how fast molecules move through a system. An example is the permeation of a substance within a volume *VA* and a donor concentration *cD* through a membrane with a surface *A*. The acceptor concentration *cA* of the permeating substance on the other side can be detected over the time *t*. Eq. (2) for the permeation coefficient *P* can be derived from the Fick's law of diffusion [18, 19]:

$$\frac{dc\_A}{dt} = P \cdot A \frac{c\_D}{V\_A} \tag{2}$$

This equation can only be used for *c <sup>D</sup>* ≫ *c A* . With respect to skin, the permeation coefficient is the preferred parameter, as it is easier to measure compared to the diffusion coefficient. Due to the different layers of the skin, permeation and diffusion coefficient changes all the time from layer to layer.

## **3. State of the art for investigation of penetration, diffusion and permeation within the skin**

To understand and investigate diffusion and permeation of the skin biopsies, several methods have been established. Some are summarized in the following.

they can penetrate into the skin. It is minimal invasive and possible to investigate the penetration directly on human skin. A disadvantage of this method is the undefined thicknesses of the stripped layer. It varies from experiment to experiment and differs with the skin model or

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249

As mentioned before, the above methods all together provide a detailed characterization of diffusion and permeation effects within the skin. But most of them can hardly be adapted to skin tissue models used in drug and substance testing. Here usually small culture devices, e.g. 12- or 96-well plates are preferred, as they allow for handling of a large number of samples in parallel. Furthermore, most methods require treatment of the sample in one or the other way. Therefore, it is quite difficult to determine the time-depending changes of diffusion and permeation.

The need to evaluate skin permeation, test cosmetic products and toxicologically screen topically applied compounds is evident. Historically, several millions of animal experiments have been performed worldwide to address this purpose [40]. Since animal experiments are under massive debate, ethical and regulatory issues, but also severe differences between animal and human data pushed the development and commercialization of diverse *in vitro* skin models [41, 42]. Human skin equivalents (HSEs) can be categorized into two main groups: the epidermis-only and full-thickness models. For both, the differentiation of keratinocytes and hence development of the various layers of the epidermis is important to model actual skin barrier properties more closely. In this context, the direct exposure to air as well as the culture media that supply nutrients for cell growth and differentiation from below has been found to be beneficial [43]. Growing cells on a porous membrane is one of the most commonly used ways to accomplish this air-liquid interface culture. According to the Organization for Economic Co-operation and Development (OECD) test guideline 431 (skin corrosion) and 439 (skin irritation), currently validated skin models include EpiSkin™ (L'oreal, France), EpiDerm™ SIT (MatTek Corporation, USA), SkinEthic™ RHE (SkinEthic laboratories, France), EpiCS® (CellSystems, Germany) and LabCyte EPI-MODEL24 SIT (Japan Tissue Engineering Co., Japan). These 3D skin models are all composed of one cell type only, the keratinocytes, mimicking the epidermis of native human skin and are especially advantageous with respect to high reproducibility [44]. However still not validated, there are also commercially available full thickness skin models composed of an additional dermal layer (e.g., GraftSkin®, EpiDermFT®, and Pheninon®). The models described are nowadays widely used for animal-

free tests in drug development as well as, the chemical and cosmetics industries.

A common method for *in vitro* cultivation of skin models is the use of a membrane insert system. This system consists of a plastic vessel with a permeable membrane at the bottom. It enables the cultivation of skin tissue models in airlift on the membrane and guarantees the

**5. Determination of permeation and diffusion coefficients in membrane insert systems via measurement and simulation**

skin type. The thickness can be estimated by weighing. [36–39]

**4. Skin tissue models**

The Franz diffusion cell is a well-known device to measure the permeation of a substance through a skin biopsy. This device consists of two chambers where the skin biopsy (or any other barrier) is fixed in between. The test substance can be applied to the top chamber (donor) of the skin; it permeates through the barrier into the bottom chamber (acceptor). The fluid in the bottom chamber is mixed by means of a magnetic stirrer. On this side, samples can be taken and the concentration can be analyzed. The concentration of the acceptor is plotted over the time and the permeation coefficient can be calculated according to Eq. (2). The whole system can be temperature-controlled. The usual size (height) of a Franz diffusion cell is in the range of 19–179 mm. Besides permeation investigation, this system is also used to test the quality of skin models and the effects of pharmaceutical substances on the skin. [18, 20–26]

Fluorescence recovery after photobleaching (FRAP) is a method to measure molecular diffusion in tissues or gels, mainly for high molecular weight compounds. For this, the substance to be analyzed must be labeled with a fluorochrome. Mostly fluorescence-labeled proteins or FITC-dextranes (fluorescein isothiocyanate-dextranes) with different molecular sizes are used. The tissues or gels have to be soaked with this substance. This can be realized by storing the material in the fluorescence substance for some days or in case of a gel, to directly polymerize in the fluorescent substance. Then, a confocal laser is used to bleach out a certain area (mostly a line or a point) in the material. Because of diffusion, bleached molecules will move and change their position with fluorescent particles and the fluorescence recovers. After the bleaching process, the area will be scanned several times. The recovery time of fluorescence intensity is used to determine the diffusion coefficient of the substance in the material. For this, software for image analysis is used. [27–30]

Further examples for imaging methods for the determination of diffusion of molecules in skin are Fourier-transform-infrared (FTIR) spectroscopy [31, 32], two-photon fluorescence correlation spectroscopy in combination with fluorescence correlation spectroscopy (FCS) [33, 34] and optical coherence tomography [35]. These methods are noninvasive and nondestructive. Furthermore, some of them can detect molecules without fluorescence labeling. One big disadvantage is the equipment. For these imaging methods, special microscopes or also costintensive tomographs are needed.

A method to investigate the penetration process of substances into the skin is tape stripping. After treatment of the skin with the substance of interest the stratum corneum is ripped of layer by layer with an adhesive film. Then, the amount of the substance can be analyzed. For this, there are different methods to determine the concentration of the substance. One method is to detect the substances directly on the film, for example titanium dioxide can be analyzed with X-ray fluorescent measurement and fluorescent-labeled substances can be detected via laser scanning microscopy. Another possibility is to remove the skin layer from the film and apply standardized analytical methods to determine the substance concentration. With tape stripping it is possible to observe where the substance of interest is localized and how deep they can penetrate into the skin. It is minimal invasive and possible to investigate the penetration directly on human skin. A disadvantage of this method is the undefined thicknesses of the stripped layer. It varies from experiment to experiment and differs with the skin model or skin type. The thickness can be estimated by weighing. [36–39]

As mentioned before, the above methods all together provide a detailed characterization of diffusion and permeation effects within the skin. But most of them can hardly be adapted to skin tissue models used in drug and substance testing. Here usually small culture devices, e.g. 12- or 96-well plates are preferred, as they allow for handling of a large number of samples in parallel. Furthermore, most methods require treatment of the sample in one or the other way. Therefore, it is quite difficult to determine the time-depending changes of diffusion and permeation.

## **4. Skin tissue models**

**3. State of the art for investigation of penetration, diffusion and** 

have been established. Some are summarized in the following.

this, software for image analysis is used. [27–30]

intensive tomographs are needed.

To understand and investigate diffusion and permeation of the skin biopsies, several methods

The Franz diffusion cell is a well-known device to measure the permeation of a substance through a skin biopsy. This device consists of two chambers where the skin biopsy (or any other barrier) is fixed in between. The test substance can be applied to the top chamber (donor) of the skin; it permeates through the barrier into the bottom chamber (acceptor). The fluid in the bottom chamber is mixed by means of a magnetic stirrer. On this side, samples can be taken and the concentration can be analyzed. The concentration of the acceptor is plotted over the time and the permeation coefficient can be calculated according to Eq. (2). The whole system can be temperature-controlled. The usual size (height) of a Franz diffusion cell is in the range of 19–179 mm. Besides permeation investigation, this system is also used to test the quality of skin models and the effects of pharmaceutical substances on the skin. [18, 20–26]

Fluorescence recovery after photobleaching (FRAP) is a method to measure molecular diffusion in tissues or gels, mainly for high molecular weight compounds. For this, the substance to be analyzed must be labeled with a fluorochrome. Mostly fluorescence-labeled proteins or FITC-dextranes (fluorescein isothiocyanate-dextranes) with different molecular sizes are used. The tissues or gels have to be soaked with this substance. This can be realized by storing the material in the fluorescence substance for some days or in case of a gel, to directly polymerize in the fluorescent substance. Then, a confocal laser is used to bleach out a certain area (mostly a line or a point) in the material. Because of diffusion, bleached molecules will move and change their position with fluorescent particles and the fluorescence recovers. After the bleaching process, the area will be scanned several times. The recovery time of fluorescence intensity is used to determine the diffusion coefficient of the substance in the material. For

Further examples for imaging methods for the determination of diffusion of molecules in skin are Fourier-transform-infrared (FTIR) spectroscopy [31, 32], two-photon fluorescence correlation spectroscopy in combination with fluorescence correlation spectroscopy (FCS) [33, 34] and optical coherence tomography [35]. These methods are noninvasive and nondestructive. Furthermore, some of them can detect molecules without fluorescence labeling. One big disadvantage is the equipment. For these imaging methods, special microscopes or also cost-

A method to investigate the penetration process of substances into the skin is tape stripping. After treatment of the skin with the substance of interest the stratum corneum is ripped of layer by layer with an adhesive film. Then, the amount of the substance can be analyzed. For this, there are different methods to determine the concentration of the substance. One method is to detect the substances directly on the film, for example titanium dioxide can be analyzed with X-ray fluorescent measurement and fluorescent-labeled substances can be detected via laser scanning microscopy. Another possibility is to remove the skin layer from the film and apply standardized analytical methods to determine the substance concentration. With tape stripping it is possible to observe where the substance of interest is localized and how deep

**permeation within the skin**

248 Biomaterials in Regenerative Medicine

The need to evaluate skin permeation, test cosmetic products and toxicologically screen topically applied compounds is evident. Historically, several millions of animal experiments have been performed worldwide to address this purpose [40]. Since animal experiments are under massive debate, ethical and regulatory issues, but also severe differences between animal and human data pushed the development and commercialization of diverse *in vitro* skin models [41, 42]. Human skin equivalents (HSEs) can be categorized into two main groups: the epidermis-only and full-thickness models. For both, the differentiation of keratinocytes and hence development of the various layers of the epidermis is important to model actual skin barrier properties more closely. In this context, the direct exposure to air as well as the culture media that supply nutrients for cell growth and differentiation from below has been found to be beneficial [43]. Growing cells on a porous membrane is one of the most commonly used ways to accomplish this air-liquid interface culture. According to the Organization for Economic Co-operation and Development (OECD) test guideline 431 (skin corrosion) and 439 (skin irritation), currently validated skin models include EpiSkin™ (L'oreal, France), EpiDerm™ SIT (MatTek Corporation, USA), SkinEthic™ RHE (SkinEthic laboratories, France), EpiCS® (CellSystems, Germany) and LabCyte EPI-MODEL24 SIT (Japan Tissue Engineering Co., Japan). These 3D skin models are all composed of one cell type only, the keratinocytes, mimicking the epidermis of native human skin and are especially advantageous with respect to high reproducibility [44]. However still not validated, there are also commercially available full thickness skin models composed of an additional dermal layer (e.g., GraftSkin®, EpiDermFT®, and Pheninon®). The models described are nowadays widely used for animalfree tests in drug development as well as, the chemical and cosmetics industries.

## **5. Determination of permeation and diffusion coefficients in membrane insert systems via measurement and simulation**

A common method for *in vitro* cultivation of skin models is the use of a membrane insert system. This system consists of a plastic vessel with a permeable membrane at the bottom. It enables the cultivation of skin tissue models in airlift on the membrane and guarantees the supply of nutrients from below. As the membrane insert system has two separate chambers, it can be used for permeation studies similar to the Franz diffusion cell. In the following, an experimental procedure for the determination of permeation coefficients in Transwell® systems (12 and 96 well) and simulations with COMSOL Multiphysics for estimation of diffusion coefficients will be discussed. Details can be found in [53].

structure enables the computation of different physical problems in one simulation. The permeation experiment was simulated with the module "transport diluted species", which uses the Fick's law to simulate diffusion processes. In order to determine the diffusion coefficient, a parameter optimization was performed with the "optimization module". Some simplifications have to be done in order to simulate the experiment. In the experiment, the permeating substance passes through a barrier and a membrane. For the simulation, these two phases were resumed as one homogenous material. It is not possible to resolve the different phases, as investigations that are more detailed would be necessary. The diffusion coefficient of the permeating substance in the liquid phases (in donor and acceptor) was

Measurement and Simulation of Permeation and Diffusion in Native and Cultivated Tissue…

represents the molecular distribution in the mixing process. Furthermore, the geometry of the membrane insert system was simplified. In reality, the system is slightly conical. The simplified geometry is a cylinder. All boundaries of the geometry were set as "no slip". The concave surface of the agarose gel was approximated with a spherical shape. In **Figure 3**,

Different membrane insert sizes influence the permeation within the system. Investigations with fluorescein sodium salt and 2% agarose gel in 96- and 12-well membrane insert systems are shown in **Figure 4a**. The time course of the acceptor concentration in the 12-well system was steeper in comparison to the 96-well system. Therefore, the fluorescein sodium salt permeates faster through the barrier in the 12-well system compared to the 96-well system. A reason for this is the concentration gradient in the gap between the membrane and the bottom of the receiver vessel. Because of the ratio "volume to permeation surface" and the gap size, the concentration will be balanced faster in case of the 12-well system. The gradient can be reduced by increasing the mixing frequency and amplitude. However, this is limited due to spillover of the liquid. By simulating these experiments, the different concentration distributions below the membrane can be visualized. In **Figure 4c** and **d** the concentration at different time points is plotted over the length below the membrane from the middle point to the edge of the receiver vessel (see red line on **Figure 4b**). The concentration difference between the middle and edge in 12-well systems is higher than that in 96-well systems. This indicates a better and faster concentration balance in the larger system, which explains the

The reproducibility of the suggested method is an important aspect. The permeation coefficient determined in different permeation experiments with fluorescein sodium salt through 2% agarose gel differed up to 40.9%. Although the value seems to be quite high, it is still within the range of deviation of permeability experiments with Franz diffusion cells reported in the literature [24]. It was found that small concentration variations during the preparation of the donor substance can cause large variation. By using one stock solution for different experiments, the deviation can be reduced to 29%. Further reasons for the deviation can also be small pipetting errors and variations of the barrier, for example, variation of the gel con-

/s. This parameter

251

http://dx.doi.org/10.5772/intechopen.72904

determined in preliminary mixing tests and was found to be 1 × 10−9 m2

**5.2. Influence of different settings and validation of the system.**

accelerated permeation.

centration or volume.

the geometry and the mesh of the 96- and 12-well Transwell® systems are shown.

#### **5.1. Measurement and simulation.**

The scheme of a permeation experiment in a membrane insert system is shown in **Figure 2**. The tissue barrier is established on the membrane. It is composed of agarose gel or 3D tissue to validate the method. The 3D tissues consist of a collagen matrix with human fibroblasts within and HaCaT cells on the top of the matrix. On top of the barrier, the donor is applied. The donor contains the substance to be analyzed, which permeates through the barrier. The acceptor, which collects the permeating substance, is located on the other side of the membrane in the receiver vessel. Temperature, humidity and mixing are parameters that influence the permeation and should be kept constant. It is recommended to perform the experiment on a shaker in an incubator with 37°C and ≥90% humidity (conditions for human cell culture). To avoid hydrodynamic pressure, the fluid surface in the insert (donor) and in the receiver vessel (acceptor) should be on the same level. The used volume for the experiment in 12 and 96 Transwell® systems is shown in **Table 1**. Sampling, like in the Franz diffusion cell, is difficult because of the small volume in the acceptor. A solution is the use of fluorescence-labeled substances, by which the permeate concentration can be detected via fluorescence measurement in the receiver vessel. A more elaborate possibility is to run several permeation experiments in parallel and take a sample from one vessel per time point. Then, the concentration of the substance can be measured analytically.

The permeation experiment was simulated with COMSOL Multiphysics. This program is based on the finite element method and offers different physic simulation modules. This

**Figure 2.** Schematic sketch of a membrane insert system.


**Table 1.** Liquid volume in the acceptor and the donor in different membrane insert systems with a barrier of 2 mm thickness.

structure enables the computation of different physical problems in one simulation. The permeation experiment was simulated with the module "transport diluted species", which uses the Fick's law to simulate diffusion processes. In order to determine the diffusion coefficient, a parameter optimization was performed with the "optimization module". Some simplifications have to be done in order to simulate the experiment. In the experiment, the permeating substance passes through a barrier and a membrane. For the simulation, these two phases were resumed as one homogenous material. It is not possible to resolve the different phases, as investigations that are more detailed would be necessary. The diffusion coefficient of the permeating substance in the liquid phases (in donor and acceptor) was determined in preliminary mixing tests and was found to be 1 × 10−9 m2 /s. This parameter represents the molecular distribution in the mixing process. Furthermore, the geometry of the membrane insert system was simplified. In reality, the system is slightly conical. The simplified geometry is a cylinder. All boundaries of the geometry were set as "no slip". The concave surface of the agarose gel was approximated with a spherical shape. In **Figure 3**, the geometry and the mesh of the 96- and 12-well Transwell® systems are shown.

#### **5.2. Influence of different settings and validation of the system.**

supply of nutrients from below. As the membrane insert system has two separate chambers, it can be used for permeation studies similar to the Franz diffusion cell. In the following, an experimental procedure for the determination of permeation coefficients in Transwell® systems (12 and 96 well) and simulations with COMSOL Multiphysics for estimation of diffusion

The scheme of a permeation experiment in a membrane insert system is shown in **Figure 2**. The tissue barrier is established on the membrane. It is composed of agarose gel or 3D tissue to validate the method. The 3D tissues consist of a collagen matrix with human fibroblasts within and HaCaT cells on the top of the matrix. On top of the barrier, the donor is applied. The donor contains the substance to be analyzed, which permeates through the barrier. The acceptor, which collects the permeating substance, is located on the other side of the membrane in the receiver vessel. Temperature, humidity and mixing are parameters that influence the permeation and should be kept constant. It is recommended to perform the experiment on a shaker in an incubator with 37°C and ≥90% humidity (conditions for human cell culture). To avoid hydrodynamic pressure, the fluid surface in the insert (donor) and in the receiver vessel (acceptor) should be on the same level. The used volume for the experiment in 12 and 96 Transwell® systems is shown in **Table 1**. Sampling, like in the Franz diffusion cell, is difficult because of the small volume in the acceptor. A solution is the use of fluorescence-labeled substances, by which the permeate concentration can be detected via fluorescence measurement in the receiver vessel. A more elaborate possibility is to run several permeation experiments in parallel and take a sample from one vessel per time point. Then, the concentration of the

The permeation experiment was simulated with COMSOL Multiphysics. This program is based on the finite element method and offers different physic simulation modules. This

**Table 1.** Liquid volume in the acceptor and the donor in different membrane insert systems with a barrier of 2 mm

coefficients will be discussed. Details can be found in [53].

**5.1. Measurement and simulation.**

250 Biomaterials in Regenerative Medicine

substance can be measured analytically.

**Figure 2.** Schematic sketch of a membrane insert system.

thickness.

Different membrane insert sizes influence the permeation within the system. Investigations with fluorescein sodium salt and 2% agarose gel in 96- and 12-well membrane insert systems are shown in **Figure 4a**. The time course of the acceptor concentration in the 12-well system was steeper in comparison to the 96-well system. Therefore, the fluorescein sodium salt permeates faster through the barrier in the 12-well system compared to the 96-well system. A reason for this is the concentration gradient in the gap between the membrane and the bottom of the receiver vessel. Because of the ratio "volume to permeation surface" and the gap size, the concentration will be balanced faster in case of the 12-well system. The gradient can be reduced by increasing the mixing frequency and amplitude. However, this is limited due to spillover of the liquid. By simulating these experiments, the different concentration distributions below the membrane can be visualized. In **Figure 4c** and **d** the concentration at different time points is plotted over the length below the membrane from the middle point to the edge of the receiver vessel (see red line on **Figure 4b**). The concentration difference between the middle and edge in 12-well systems is higher than that in 96-well systems. This indicates a better and faster concentration balance in the larger system, which explains the accelerated permeation.

The reproducibility of the suggested method is an important aspect. The permeation coefficient determined in different permeation experiments with fluorescein sodium salt through 2% agarose gel differed up to 40.9%. Although the value seems to be quite high, it is still within the range of deviation of permeability experiments with Franz diffusion cells reported in the literature [24]. It was found that small concentration variations during the preparation of the donor substance can cause large variation. By using one stock solution for different experiments, the deviation can be reduced to 29%. Further reasons for the deviation can also be small pipetting errors and variations of the barrier, for example, variation of the gel concentration or volume.

**Figure 3.** Geometry of the 96 a) and 12 b) Transwell® system and the used mesh c) and d) implemented in COMSOL **Mult**i**physics**.

a 12 Transwell® system. The 3D tissues consisted of a collagen matrix with different constellations of cell layers. Human primary fibroblasts were integrated into the collagen matrix and HaCaT cells were seeded on the top. A representative example of the 3D tissue is shown in

**Figure 4.** Permeation in 12- and 96-well membrane insert systems. (a) Experimental results of permeation of fluorescein sodium salt through 2% agarose. (b) Geometry of the 12 and 96 Transwell® system for the simulation and the position of the concentration measurement. (c) and (d) the concentration distribution at different time over the length at the red

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253

The results show that the permeation coefficient decreases when additional cell layers are added in the tissue model. The permeation coefficient of fluorescein sodium salt through collagen matrix (without cells) is 2.18 × 10−8 m/s, 1.85 × 10−8 m/s in tissue models with fibroblast and 1.67 × 10−8 m/s in models with HaCaT cells (see **Figure 6b)**. These results represent very

The particle size influences the permeation behavior through gels and biomaterials. It is well known that smaller molecules permeate faster through a matrix mesh than larger particles. This was already observed for permeation experiments through sclera [19], human epidermal membrane [45], human skin [33] and rat skin [3]. Fluorescein sodium salt and fluorescein isothiocyanate-dextrans (FITC-dextran) were used to vary the molecular size from 376 g/mol up

well the barrier function of keratinocytes of the skin [3].

lined position on (b) ((c) 96 and (d) 12 Transwell® system).

**Figure 5**.

The membrane itself also has an influence on the permeation. Same experiments as described above were carried out with different membranes in a 12 Transwell® system. Membranes consisting of polycarbonate (PC) with 0.4 and 3.0 μm pore size showed a quite similar permeation coefficient of 8.03 and 8.1 × 10−8 m/s. For polyethylene (PE) membranes with 0.4 μm pore size the mean value of the permeation coefficient was 5.94 × 10−8 m/s and for PE with 3.0 μm pore size it was 8.59 × 10−8 m/s (see **Figure 6a**). Except for the PE membrane with pore size of 3.0 μm there is no significant difference. The reason for this is the pore density of the membrane. In total, the pore surface of PE membranes with 0.4 μm pore size is 0.25 mm2 per 1 cm2 and for the other membranes 6.3–7.05 mm2 per 1 cm2 . The material of the membrane seems not to influence the permeation.

The suggested method is sensitive enough to determine a cover layer of HaCaT cells of 3D skin tissue. To prove this, in a permeation experiment different 3D tissue model was tested in

Measurement and Simulation of Permeation and Diffusion in Native and Cultivated Tissue… http://dx.doi.org/10.5772/intechopen.72904 253

**Figure 4.** Permeation in 12- and 96-well membrane insert systems. (a) Experimental results of permeation of fluorescein sodium salt through 2% agarose. (b) Geometry of the 12 and 96 Transwell® system for the simulation and the position of the concentration measurement. (c) and (d) the concentration distribution at different time over the length at the red lined position on (b) ((c) 96 and (d) 12 Transwell® system).

a 12 Transwell® system. The 3D tissues consisted of a collagen matrix with different constellations of cell layers. Human primary fibroblasts were integrated into the collagen matrix and HaCaT cells were seeded on the top. A representative example of the 3D tissue is shown in **Figure 5**.

The membrane itself also has an influence on the permeation. Same experiments as described above were carried out with different membranes in a 12 Transwell® system. Membranes consisting of polycarbonate (PC) with 0.4 and 3.0 μm pore size showed a quite similar permeation coefficient of 8.03 and 8.1 × 10−8 m/s. For polyethylene (PE) membranes with 0.4 μm pore size the mean value of the permeation coefficient was 5.94 × 10−8 m/s and for PE with 3.0 μm pore size it was 8.59 × 10−8 m/s (see **Figure 6a**). Except for the PE membrane with pore size of 3.0 μm there is no significant difference. The reason for this is the pore density of the membrane. In

**Figure 3.** Geometry of the 96 a) and 12 b) Transwell® system and the used mesh c) and d) implemented in COMSOL

The suggested method is sensitive enough to determine a cover layer of HaCaT cells of 3D skin tissue. To prove this, in a permeation experiment different 3D tissue model was tested in

per 1 cm2

. The material of the membrane seems not to

and for

total, the pore surface of PE membranes with 0.4 μm pore size is 0.25 mm2

per 1 cm2

the other membranes 6.3–7.05 mm2

influence the permeation.

**Mult**i**physics**.

252 Biomaterials in Regenerative Medicine

The results show that the permeation coefficient decreases when additional cell layers are added in the tissue model. The permeation coefficient of fluorescein sodium salt through collagen matrix (without cells) is 2.18 × 10−8 m/s, 1.85 × 10−8 m/s in tissue models with fibroblast and 1.67 × 10−8 m/s in models with HaCaT cells (see **Figure 6b)**. These results represent very well the barrier function of keratinocytes of the skin [3].

The particle size influences the permeation behavior through gels and biomaterials. It is well known that smaller molecules permeate faster through a matrix mesh than larger particles. This was already observed for permeation experiments through sclera [19], human epidermal membrane [45], human skin [33] and rat skin [3]. Fluorescein sodium salt and fluorescein isothiocyanate-dextrans (FITC-dextran) were used to vary the molecular size from 376 g/mol up

in **Figure 8a**. In the case of larger molecular size, the simulation is different from experimental results. A closer look shows that the experimental data increased and flattened faster compared to the simulation (see **Figure 8b)**. A possible reason for this effect could be the presence of particle size distribution in the substance. The migration of smaller particles reduces the lag time in the beginning of the permeation process, where large particles can increase the friction and slows down the diffusion. Especially the second effect leads to abnormal diffusion [46, 47].

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255

**Figure 7.** Permeability coefficient plotted over stokes radius. Results of permeation experiment in 96 Transwell® system

**Figure 8.** Simulation of permeation experiment with different molecular sizes. Results of permeation experiments and simulations with COMSOL **Multiphysics** of (a) fluorescein sodium salt and (b) FITC dextran 10,000 mg/mol in 96

This cannot be simulated with equations based on Fick's law.

with fluorescein sodium salt and FITC-dextran through 2% agarose gel.

Transwell systems through 2% agarose gel.

**Figure 5.** Hematoxylin and eosin stain of 3D tissue model consisting of collagen matrix with fibroblast and HaCaT seeded on the top.

**Figure 6.** Permeation through different membranes and 3D skin tissues. (a) Results of permeation experiment with fluorescein sodium salt through 2% agarose gel on different membranes (n = 3). The 12 Transwell® system was used and membranes consisting of polycarbonate (PC) and polyethylene (PE) with pore sizes of 3.0 and 0.4 μm were tested. (b) Results of permeation experiments with fluorescein sodium salt through different 3D tissues in a 12 Transwell® system (n = 6). The 3D skin tissue consisted of collagen (Col.), collagen with primary human fibroblast (Col. + F.), collagen with HaCaTs (Col. + H.) and collagen with primary human fibroblast and HaCaTs (Col. F. + H.).

to 40,000 g/mol for permeation experiments in a 96 Transwell® system. The results show similar correlations between permeation coefficient and molecular size as the studies mentioned above. There is almost a linear relationship between these two factors, which is well described by the Navier–Stokes equation (see **Figure 7**). An exception is FITC-dextran 40,000 g/mol, which deviates from the linearity.

These experiments were simulated with COMSOL Multiphysics, where the diffusion coefficient is fitted on the experimental data. The simulation based on Fick's law is quite accurate for the permeation of substances with a small molecular size from 376 g/mol up to 4000 g/ mol. The simulation shows good agreement with the experiment which is exemplary shown in **Figure 8a**. In the case of larger molecular size, the simulation is different from experimental results. A closer look shows that the experimental data increased and flattened faster compared to the simulation (see **Figure 8b)**. A possible reason for this effect could be the presence of particle size distribution in the substance. The migration of smaller particles reduces the lag time in the beginning of the permeation process, where large particles can increase the friction and slows down the diffusion. Especially the second effect leads to abnormal diffusion [46, 47]. This cannot be simulated with equations based on Fick's law.

**Figure 7.** Permeability coefficient plotted over stokes radius. Results of permeation experiment in 96 Transwell® system with fluorescein sodium salt and FITC-dextran through 2% agarose gel.

to 40,000 g/mol for permeation experiments in a 96 Transwell® system. The results show similar correlations between permeation coefficient and molecular size as the studies mentioned above. There is almost a linear relationship between these two factors, which is well described by the Navier–Stokes equation (see **Figure 7**). An exception is FITC-dextran 40,000 g/mol,

HaCaTs (Col. + H.) and collagen with primary human fibroblast and HaCaTs (Col. F. + H.).

**Figure 6.** Permeation through different membranes and 3D skin tissues. (a) Results of permeation experiment with fluorescein sodium salt through 2% agarose gel on different membranes (n = 3). The 12 Transwell® system was used and membranes consisting of polycarbonate (PC) and polyethylene (PE) with pore sizes of 3.0 and 0.4 μm were tested. (b) Results of permeation experiments with fluorescein sodium salt through different 3D tissues in a 12 Transwell® system (n = 6). The 3D skin tissue consisted of collagen (Col.), collagen with primary human fibroblast (Col. + F.), collagen with

**Figure 5.** Hematoxylin and eosin stain of 3D tissue model consisting of collagen matrix with fibroblast and HaCaT

These experiments were simulated with COMSOL Multiphysics, where the diffusion coefficient is fitted on the experimental data. The simulation based on Fick's law is quite accurate for the permeation of substances with a small molecular size from 376 g/mol up to 4000 g/ mol. The simulation shows good agreement with the experiment which is exemplary shown

which deviates from the linearity.

seeded on the top.

254 Biomaterials in Regenerative Medicine

**Figure 8.** Simulation of permeation experiment with different molecular sizes. Results of permeation experiments and simulations with COMSOL **Multiphysics** of (a) fluorescein sodium salt and (b) FITC dextran 10,000 mg/mol in 96 Transwell systems through 2% agarose gel.

## **6. Conclusion and prospective**

The studies have shown that the membrane insert system is a possible alternative for permeation studies. An advantage of the system is the small size. The membrane insert system of 96-well plates from CORNING has a cultivation surface of 0.143 cm<sup>2</sup> and a height of 2 cm. This reduces significantly the number of cells, materials and substances needed for the cultivation. In comparison to the Franz diffusion cell, the handling of such a system is easier and one experiment can be run with a large specimen number. A time-intensive mounting process of the samples (skin) is no longer required and the experiment can directly execute in the system. The sensitivity of this system is good enough to differ between 3D tissues as well as different cell layers and to detect different molecular sizes of the substance.

FRAP fluorescence recovery after photobleaching

HaCaT human adult low calcium high temperature

, Katharina Schimek<sup>2</sup>

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1-18. DOI: 10.1016/S0378-5173(01)00731-1

2194. DOI: 10.1002/jps.20695

**208**:7-11. PubMed ID: 10884933

ID: 11509908

\*Address all correspondence to: poertner@tuhh.de

OECD organization for economic co-operation and development

, Uwe Marx2

1 Institute of Bioprocess and Biosystems Engineering, Hamburg University of Technology,

[1] Ackermann K, Lombardi Borgia S, Korting H, Mewes K, Schäfer-Korting M.The Phenion® full-thickness skin model for percutaneous absorption testing. Skin Pharmacology and

[2] Asbill C. Evaluation of a human bio-engineered skin equivalent for drug permeation studies. Pharmaceutical Research. 2000;**17**(9):1092-1097. DOI: 10.1023/A:1026405712870

[3] Hadgraft J. Skin, the final frontier. International Journal of Pharmaceutics. 2001;**224**(1-2):

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[7] Bouwstra J, Pilgram G, Gooris G, Koerten H, Ponec M. New aspects of the skin barrier organization. Skin Pharmacology and Applied Skin Physiology. 2001;**14**:52-62. PubMed

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and Ralf Pörtner1

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\*

http://dx.doi.org/10.5772/intechopen.72904

257

FTIR Fourier-transform-infrared

HSE human skin equivalents

**Author details**

Hao-Hsiang Hsu1

Hamburg, Germany

**References**

2 TissUse GmbH, Berlin, Germany

It should be considered that the permeation is detected through a membrane and the size of the system influences the permeation properties. The specimen has to cover up the whole membrane, otherwise the substance will pass by. Unlike the imaging and stripping method, it is not possible to measure the diffusion and penetration inside the membrane insert system. Alternatively, the diffusion can be calculated or estimated by simulation. Furthermore, this method can be used to investigate changes in the permeation behavior of the skin model during the cultivation or it can also be adapted for other systems, which use membrane insert systems. An example is the Two-Organ-on-a-Chip, a variant of TissUse's Multi-Organ-Chip platform [48–52]. This device enables the integration of skin models in a membrane insert system. Therefore this method can be used to investigate the permeation process into an organon-a-chip system in order to understand the substance distribution.

With the help of the simulation in COMSOL Multiphysics, it is possible to calculate the diffusion process in the membrane insert system. It is limited to small particle sizes and normal diffusion described by Fick's law. Otherwise, it is possible to optimize the simulation by integration of abnormal diffusion. Furthermore, the simulation is an attractive tool to support the experiments. On the one hand, it can be used to understand physical phenomena and to reduce experimental effort. On the other hand, it is modular and can be integrated into a more complex system to support permeation studies.

## **Acknowledgements**

This work was created with financial support from Deutsche Forschungsgemeinschaft (DFG) under grant No. PO413/12-1 and LA 1028/7-1.

## **Appendices and nomenclatures**

EC European regulation

FITC dextran fluorescein isothiocyanate-dextrans


## **Author details**

**6. Conclusion and prospective**

256 Biomaterials in Regenerative Medicine

The studies have shown that the membrane insert system is a possible alternative for permeation studies. An advantage of the system is the small size. The membrane insert system

This reduces significantly the number of cells, materials and substances needed for the cultivation. In comparison to the Franz diffusion cell, the handling of such a system is easier and one experiment can be run with a large specimen number. A time-intensive mounting process of the samples (skin) is no longer required and the experiment can directly execute in the system. The sensitivity of this system is good enough to differ between 3D tissues as well as

It should be considered that the permeation is detected through a membrane and the size of the system influences the permeation properties. The specimen has to cover up the whole membrane, otherwise the substance will pass by. Unlike the imaging and stripping method, it is not possible to measure the diffusion and penetration inside the membrane insert system. Alternatively, the diffusion can be calculated or estimated by simulation. Furthermore, this method can be used to investigate changes in the permeation behavior of the skin model during the cultivation or it can also be adapted for other systems, which use membrane insert systems. An example is the Two-Organ-on-a-Chip, a variant of TissUse's Multi-Organ-Chip platform [48–52]. This device enables the integration of skin models in a membrane insert system. Therefore this method can be used to investigate the permeation process into an organ-

With the help of the simulation in COMSOL Multiphysics, it is possible to calculate the diffusion process in the membrane insert system. It is limited to small particle sizes and normal diffusion described by Fick's law. Otherwise, it is possible to optimize the simulation by integration of abnormal diffusion. Furthermore, the simulation is an attractive tool to support the experiments. On the one hand, it can be used to understand physical phenomena and to reduce experimental effort. On the other hand, it is modular and can be integrated into a more

This work was created with financial support from Deutsche Forschungsgemeinschaft (DFG)

and a height of 2 cm.

of 96-well plates from CORNING has a cultivation surface of 0.143 cm<sup>2</sup>

different cell layers and to detect different molecular sizes of the substance.

on-a-chip system in order to understand the substance distribution.

complex system to support permeation studies.

under grant No. PO413/12-1 and LA 1028/7-1.

**Appendices and nomenclatures**

EC European regulation

FITC dextran fluorescein isothiocyanate-dextrans

**Acknowledgements**

Hao-Hsiang Hsu1 , Katharina Schimek<sup>2</sup> , Uwe Marx2 and Ralf Pörtner1 \*

\*Address all correspondence to: poertner@tuhh.de

1 Institute of Bioprocess and Biosystems Engineering, Hamburg University of Technology, Hamburg, Germany

2 TissUse GmbH, Berlin, Germany

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260 Biomaterials in Regenerative Medicine


**Chapter 11**

**Provisional chapter**

**Biomaterials for Tendon/Ligament and Skin**

**Biomaterials for Tendon/Ligament and Skin** 

DOI: 10.5772/intechopen.69716

Tendon/ligament injury or skin injuries due to diseases, trauma, and surgery are common. Timely functional repair and tissue regeneration is a key to improve the quality of life of the patient while reducing health care cost. Tendon/ligament/skin is also enriched in a common extracellular matrix (ECM), collagen I, III, and elastin. Tissue engineering and regenerative medicine, the combination of (stem) cells, growth factors, and biomaterial scaffolds, is an emergent field, which has attracted substantial attention over the years. Biomaterials are considered the foundation of regenerative medicine. A key to find a new solution to tendon/ligament/skin healing is to synthesize new functional biomaterials, which have better biomechanical properties, biodegradability, and cell supporting prop‐ erties. This chapter will review existing FDA‐approved biomaterial‐based therapy, as well

> © 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

From material science point of view, tendon/ligament and skin tissue are similar in that they are mainly composed of collagen and elastin. Up to 80% of the dry weight is collagen type I. However, their micro and hierarchical structures and functions are very different from each other. Tendon/ligament is composed of densely packed, aligned collagen fiber bundles, whereas skin is composed of a layered structure of collagen random naonfibril network. **Table 1**

The regeneration of a large size, lost tendon/ligament, and skin tissue normally involves a type of stem cells/progenitor cells of endogenous origin or exogenous origin combined with a

summarizes the key properties of tendon/ligament and skin tissue in human.

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69716

as those in development.

**Keywords:** biomaterials, tendon, skin, regeneration

**Regeneration**

**Regeneration**

Xingguo Cheng

**Abstract**

**1. Introduction**

Xingguo Cheng

**Provisional chapter**

## **Biomaterials for Tendon/Ligament and Skin Regeneration Regeneration**

**Biomaterials for Tendon/Ligament and Skin** 

DOI: 10.5772/intechopen.69716

#### Xingguo Cheng Xingguo Cheng Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69716

#### **Abstract**

Tendon/ligament injury or skin injuries due to diseases, trauma, and surgery are common. Timely functional repair and tissue regeneration is a key to improve the quality of life of the patient while reducing health care cost. Tendon/ligament/skin is also enriched in a common extracellular matrix (ECM), collagen I, III, and elastin. Tissue engineering and regenerative medicine, the combination of (stem) cells, growth factors, and biomaterial scaffolds, is an emergent field, which has attracted substantial attention over the years. Biomaterials are considered the foundation of regenerative medicine. A key to find a new solution to tendon/ligament/skin healing is to synthesize new functional biomaterials, which have better biomechanical properties, biodegradability, and cell supporting prop‐ erties. This chapter will review existing FDA‐approved biomaterial‐based therapy, as well as those in development.

**Keywords:** biomaterials, tendon, skin, regeneration

## **1. Introduction**

From material science point of view, tendon/ligament and skin tissue are similar in that they are mainly composed of collagen and elastin. Up to 80% of the dry weight is collagen type I. However, their micro and hierarchical structures and functions are very different from each other. Tendon/ligament is composed of densely packed, aligned collagen fiber bundles, whereas skin is composed of a layered structure of collagen random naonfibril network. **Table 1** summarizes the key properties of tendon/ligament and skin tissue in human.

The regeneration of a large size, lost tendon/ligament, and skin tissue normally involves a type of stem cells/progenitor cells of endogenous origin or exogenous origin combined with a


effect, whereas pure PRP (P‐PRP) without leucocytes have anabolic effects and results in over‐ proliferation and scar tissue formation. The complex interaction between PRP and stem cells may explain the different preclinical outcome and warrant a large random clinical trial.

Biomaterials for Tendon/Ligament and Skin Regeneration

http://dx.doi.org/10.5772/intechopen.69716

265

Fibrin can be derived from human plasma. After addition of thrombin, it will form a gel, a process used for blood clotting. Stem cells can be added together with fibrin and thrombin and sprayed onto the wound for the promotion of wound closure and healing. The fibrin combined with bone marrow stem cells spray was successfully tested to prevent ulceration and acceler‐ ate wound closure in mice [4]. A small human clinical trial showed this approach accelerates wound closure and resurfacing without adverse effects [5]. Poly(ethylene glycol) PEG‐modified fibrin combined with adipose‐derived stem cells (ADSCs) also showed promising results in

Human placenta‐derived biomaterial is unique in that it has immune privilege while it con‐ tains multiple growth factors. Human ADSCs (hADSCs) seeded onto radiosterilized human amnion are viable and can proliferate. These cells are able to migrate over these scaffolds as demonstrated by using time‐lapse microscopy. In addition, the scaffolds induce hADSCs to secrete interleukin‐10, an important negative regulator of inflammation [7]. This suggests that placenta‐derived biomaterial may be a good substrate for stem cells and used for skin/tendon

3D micronized (300–600 μm) amniotic membrane (mAM) was made by means of repeated freeze‐thawing cycles to deplete cell components and homogenized with a macrohomoge‐ nizer in liquid nitrogen. These mAM loaded with epidermal stem cells (ESCs) (ESC‐mAM) was further transplanted to full‐thickness skin defects in nude mice. ESCs survived well and formed a new epidermis. Four weeks after transplantation, papilla‐like structures were observed, and collagen fibers were well and regularly arranged in the newly formed dermal layer. In conclusion, the mAM as a novel natural microcarrier possesses an intact basement

A recent study showed that decellularized matrix from different tissues (tendon, bone, and skin) affect the differentiation of stem cell in a different way. Decellularized bone matrix may induce the undesirable osteogenic differentiation of stem cells, while tendon or skin matrix does not have such an effect [9]. The ECM components provide a niche for proper differentia‐ tion of stem cells. For example, ECM without biglycan (Bgn) and fibromodulin (Fmod) will affect the differentiation of tendon stem cells by modulating bone morphogenetic protein signaling and impairs tendon formation *in vivo* [10]. Autologous origin, but decelluarized der‐ mal matrix using trypsin and Trion X‐100, after combined with ADSCs, was found to enhance

**2.2. FDA‐approved ECM grafts for tendon augmentation and skin regeneration**

wound healing in a murine model of a full‐thickness skin defect [11].

*2.1.2. Fibrin*

a pig burn model [6].

applications.

*2.1.3. Amniotic membrane*

membrane structure and bioactivities [8].

**Table 1.** A general comparison of two different connective tissues: tendon/ligament and skin.

biomaterial. Stem cells can be derived from tendon/ligament or skin, or from embryo, pla‐ centa, adipose tissue, bone marrow, umbilical cord, etc. This review focuses only on regenera‐ tion using exogenous stem cells coupled with a biomaterial matrix/carrier implanted to the wound site. Biomaterial can play a key role in protection of the cells from dehydration while it serves as a temporary substrate for stem cells to proliferate, or differentiate, and synthesize tissue‐specific matrix. The morphology, topography, composition, stiffness of biomaterials may play a key role in controlling the differentiation of stem cells, in addition to biochemical factors, mechanical cues, or genetic/cellular cues. This chapter focuses on biomaterial explored for tendon/ligament/skin tissue regeneration applications.

#### **1.1. Research methods**

We performed a comprehensive search of PubMed using keywords "tendon," "ligament," "skin," "regeneration," "scaffold", over the years 1970–2016. All articles relevant to the subject were retrieved, and their bibliographies hand searched for further references in the context to biomaterials for tendon/ligament/skin regeneration

## **2. Results**

#### **2.1. Biomaterial directly derived from patients**

#### *2.1.1. PRP*

Platelet‐rich plasma (PRP) is derived from blood and PRP gel is widely used for tendon/ ligament repair. Recently, PRP was combined with adipose‐derived stem cells (ADSCs) and it was found that PRP combined with stem cells resulted in improved mechanical strength in a rabbit tendon model compared to PRP gel alone [1]. Similar results were also observed using PRP with tendon‐derived stem cells (TDSCs) [2]. However, in a sheep model, no differ‐ ences were observed between the PRP group and PRP‐stem cell group [3]. This approach is highly translational, since both autologous stem cells and PRP can be obtained from the same patients. The concern may be the leucocytes‐containing PRP (L‐PRP) that have a catabolic effect, whereas pure PRP (P‐PRP) without leucocytes have anabolic effects and results in over‐ proliferation and scar tissue formation. The complex interaction between PRP and stem cells may explain the different preclinical outcome and warrant a large random clinical trial.

#### *2.1.2. Fibrin*

biomaterial. Stem cells can be derived from tendon/ligament or skin, or from embryo, pla‐ centa, adipose tissue, bone marrow, umbilical cord, etc. This review focuses only on regenera‐ tion using exogenous stem cells coupled with a biomaterial matrix/carrier implanted to the wound site. Biomaterial can play a key role in protection of the cells from dehydration while it serves as a temporary substrate for stem cells to proliferate, or differentiate, and synthesize tissue‐specific matrix. The morphology, topography, composition, stiffness of biomaterials may play a key role in controlling the differentiation of stem cells, in addition to biochemical factors, mechanical cues, or genetic/cellular cues. This chapter focuses on biomaterial explored

**Tendon/ligament Skin**

Structure 1D Aligned fiber bundles (nano‐macro) 2D random nanofibrils (weave basket

pattern)

Main composition >80% Collagen, ∼5% elastin elastin >80% Collagen, ∼5% elastin Minor composition Fibronectin, proteoglycan Fibronectin, proteoglycan Main cells Fibroblast Fibroblast and keratinocytes

Vasculature Few Abundant

Mechanical properties Strong and elastic Weak and elastic

**Table 1.** A general comparison of two different connective tissues: tendon/ligament and skin.

We performed a comprehensive search of PubMed using keywords "tendon," "ligament," "skin," "regeneration," "scaffold", over the years 1970–2016. All articles relevant to the subject were retrieved, and their bibliographies hand searched for further references in the context to

Platelet‐rich plasma (PRP) is derived from blood and PRP gel is widely used for tendon/ ligament repair. Recently, PRP was combined with adipose‐derived stem cells (ADSCs) and it was found that PRP combined with stem cells resulted in improved mechanical strength in a rabbit tendon model compared to PRP gel alone [1]. Similar results were also observed using PRP with tendon‐derived stem cells (TDSCs) [2]. However, in a sheep model, no differ‐ ences were observed between the PRP group and PRP‐stem cell group [3]. This approach is highly translational, since both autologous stem cells and PRP can be obtained from the same patients. The concern may be the leucocytes‐containing PRP (L‐PRP) that have a catabolic

for tendon/ligament/skin tissue regeneration applications.

biomaterials for tendon/ligament/skin regeneration

**2.1. Biomaterial directly derived from patients**

**1.1. Research methods**

264 Biomaterials in Regenerative Medicine

**2. Results**

*2.1.1. PRP*

Fibrin can be derived from human plasma. After addition of thrombin, it will form a gel, a process used for blood clotting. Stem cells can be added together with fibrin and thrombin and sprayed onto the wound for the promotion of wound closure and healing. The fibrin combined with bone marrow stem cells spray was successfully tested to prevent ulceration and acceler‐ ate wound closure in mice [4]. A small human clinical trial showed this approach accelerates wound closure and resurfacing without adverse effects [5]. Poly(ethylene glycol) PEG‐modified fibrin combined with adipose‐derived stem cells (ADSCs) also showed promising results in a pig burn model [6].

#### *2.1.3. Amniotic membrane*

Human placenta‐derived biomaterial is unique in that it has immune privilege while it con‐ tains multiple growth factors. Human ADSCs (hADSCs) seeded onto radiosterilized human amnion are viable and can proliferate. These cells are able to migrate over these scaffolds as demonstrated by using time‐lapse microscopy. In addition, the scaffolds induce hADSCs to secrete interleukin‐10, an important negative regulator of inflammation [7]. This suggests that placenta‐derived biomaterial may be a good substrate for stem cells and used for skin/tendon applications.

3D micronized (300–600 μm) amniotic membrane (mAM) was made by means of repeated freeze‐thawing cycles to deplete cell components and homogenized with a macrohomoge‐ nizer in liquid nitrogen. These mAM loaded with epidermal stem cells (ESCs) (ESC‐mAM) was further transplanted to full‐thickness skin defects in nude mice. ESCs survived well and formed a new epidermis. Four weeks after transplantation, papilla‐like structures were observed, and collagen fibers were well and regularly arranged in the newly formed dermal layer. In conclusion, the mAM as a novel natural microcarrier possesses an intact basement membrane structure and bioactivities [8].

#### **2.2. FDA‐approved ECM grafts for tendon augmentation and skin regeneration**

A recent study showed that decellularized matrix from different tissues (tendon, bone, and skin) affect the differentiation of stem cell in a different way. Decellularized bone matrix may induce the undesirable osteogenic differentiation of stem cells, while tendon or skin matrix does not have such an effect [9]. The ECM components provide a niche for proper differentia‐ tion of stem cells. For example, ECM without biglycan (Bgn) and fibromodulin (Fmod) will affect the differentiation of tendon stem cells by modulating bone morphogenetic protein signaling and impairs tendon formation *in vivo* [10]. Autologous origin, but decelluarized der‐ mal matrix using trypsin and Trion X‐100, after combined with ADSCs, was found to enhance wound healing in a murine model of a full‐thickness skin defect [11].

Many FDA‐approved human or animal decelluarized tissue matrixes have been approved for direct tendon/ligament/skin repair without any (stem) cells (**Table 2**). The key advantages of decellularized tissue grafts are that they largely maintain the main architecture, composi‐ tion, and mechanical properties of native tissues. These allografts/xenografts are processed to remove immunogenic cells, DNAs, and certain immunogenic molecules. Typical problems are that these grafts are slower to repair the tissue and some fail to restore the proper func‐ tions (e.g., scarring). For tissue regeneration using these ECM biological grafts, stem cells may need to be reseeded onto the grafts for recellularization. For a dense tendon allograft, direct cell seeding may be difficult. The recellularization onto the surface may be achieved using a cell‐loaded gel coating [12]. However, it is highly desirable to get cells inside the grafts as well. Thus, ECM grafts may be processed to have a much higher porosity than the original tissue. Proration/incision into the ECM grafts may help the penetration of cells and nutrients. Instead of being coated with cell‐laden gel, an interesting approach is to use a stem cell‐sheet to wrap around a frozen tendon graft for implantation [13]. Interesting, a case report showed that a dermal allograft combined with PRP and autologous mesenchymal stem cells (MSCs) derived from peripheral blood (PB‐MSCs) resulted in enhanced healing of human rotator cuff [14].

patch (SIS) as carrier of ADSCs significantly increases the survival of stem cells as compared to direct injection of ADSCs into the skin wounds [18]. There may be a synergistic angiogen‐ esis promoting effects of biologic graft with stem cells, which may be important for tendon/

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Another interesting approach is to use ECM directly secreted by the cells as a carrier for stem cells. For example, stem cells can be incubated at 37°C in a temperature‐responsive flask (e.g., upcellTM, Cellseed, Japan) and they will product ECM after addition of ascorbic acid. The cell sheets can be lifted at room temperature since the Poly‐n‐isopropylacrylamide (Poly‐NIPAM)‐ based substrate will become soluble at lower temperature. The cell‐ECM sheet has been explored for promotion of tendon/ligament healing [20] as well as diabetic skin wound healing [21].

(5 mm × 2 mm) collagen sponge scaffolds (Zimmer Dental) were used for the culture of (BM) MSCs. Cell‐seeded scaffolds were placed in culture dishes and incubated for 2 hours in a minimum volume of growth medium, after which more medium was applied to submerge the scaffolds. After an additional 24‐hour culture, cells seeded in scaffolds were treated with 10 ng/mL of recombinant (BMP) 12 for 12 hours. The medium was then replaced with fresh growth medium and scaffolds were either cultured for an additional 7 days or immediately implanted into partial calcaneal tendon defects in rats. It was shown that after 21 days, the BMP12‐treated, collagen‐cell scaffold results in robust formation of tendon‐like tissue [22]. Similarly, a collagen carrier combined with ASDCs showed they did not improve the biome‐ chanical properties of the tendon‐to‐bone healing. However, the ADSCs group showed less inflammation, which may lead to a more elastic repair and less scarred healing in a rat model.

Integra® bilayer wound Matrix (LifeSciences Corp., Plainsboro, New Jersey) is a dermal acellular analog composed of bovine collagen type I crosslinked with glycosaminoglycans. Importantly, inclusion of WJ‐MSC into Integra induced significant up‐regulation of prototypical angiogenic and healing factors, stimulating pleiotropic aspects of neovascularization in experimental set‐ tings of angiogenesis, without altering the inflammatory response in the animals, thus dem‐ onstrating their potential benefit in therapeutic care of wounds and skin grafts [23]. Recently, Integra dermal matrix scaffold engineered with adult mesenchymal stem cells and platelet‐rich

plasma was investigated *in vitro* and demonstrated promising preliminary results [24].

Biopolymer matrix such as collagen, gelatin, hyaluronic acid, chitosan, silk, cellulose are the essential ECM components of human, animal, or plant cells. They have showed great promise for attachment, proliferation, and differentiation of stem cells and have been explored for

skin wound healing [19].

**2.3. FDA‐approved biomaterial**

*2.3.1. Collagen sponge scaffold*

*2.3.2. Integra bilayer*

**2.4. Synthetic biopolymer matrix**

tendon and skin regeneration.

Some of the FDA‐approved ECM biomatrices were combined with stem cells, and investigated for tendon/ligament/skin regeneration applications. Human acellular allograft (Alloderm) was investigated for direct cell seeding using ADSCs. It was concluded that human ADSCs can attach to Alloderm with the dermis side up in a petri dish [15]. ADSC seeded onto Alloderm was also implanted *in vivo* for skin regenerations with promising results [16]. Strattice was evaluated for seeding with rat MSCs [17]. Thus, the stem cell‐seeded biologic graft can be used as a tendon wrap or a skin regeneration material. A study was performed to compare the survival and proliferation of stem cells via bioluminescent imaging. The use of biologic graft


**Table 2.** An example list of biologic ECM grafts with potential for tendon/skin regeneration.

patch (SIS) as carrier of ADSCs significantly increases the survival of stem cells as compared to direct injection of ADSCs into the skin wounds [18]. There may be a synergistic angiogen‐ esis promoting effects of biologic graft with stem cells, which may be important for tendon/ skin wound healing [19].

Another interesting approach is to use ECM directly secreted by the cells as a carrier for stem cells. For example, stem cells can be incubated at 37°C in a temperature‐responsive flask (e.g., upcellTM, Cellseed, Japan) and they will product ECM after addition of ascorbic acid. The cell sheets can be lifted at room temperature since the Poly‐n‐isopropylacrylamide (Poly‐NIPAM)‐ based substrate will become soluble at lower temperature. The cell‐ECM sheet has been explored for promotion of tendon/ligament healing [20] as well as diabetic skin wound healing [21].

#### **2.3. FDA‐approved biomaterial**

#### *2.3.1. Collagen sponge scaffold*

Many FDA‐approved human or animal decelluarized tissue matrixes have been approved for direct tendon/ligament/skin repair without any (stem) cells (**Table 2**). The key advantages of decellularized tissue grafts are that they largely maintain the main architecture, composi‐ tion, and mechanical properties of native tissues. These allografts/xenografts are processed to remove immunogenic cells, DNAs, and certain immunogenic molecules. Typical problems are that these grafts are slower to repair the tissue and some fail to restore the proper func‐ tions (e.g., scarring). For tissue regeneration using these ECM biological grafts, stem cells may need to be reseeded onto the grafts for recellularization. For a dense tendon allograft, direct cell seeding may be difficult. The recellularization onto the surface may be achieved using a cell‐loaded gel coating [12]. However, it is highly desirable to get cells inside the grafts as well. Thus, ECM grafts may be processed to have a much higher porosity than the original tissue. Proration/incision into the ECM grafts may help the penetration of cells and nutrients. Instead of being coated with cell‐laden gel, an interesting approach is to use a stem cell‐sheet to wrap around a frozen tendon graft for implantation [13]. Interesting, a case report showed that a dermal allograft combined with PRP and autologous mesenchymal stem cells (MSCs) derived from peripheral blood (PB‐MSCs) resulted in enhanced healing of

Some of the FDA‐approved ECM biomatrices were combined with stem cells, and investigated for tendon/ligament/skin regeneration applications. Human acellular allograft (Alloderm) was investigated for direct cell seeding using ADSCs. It was concluded that human ADSCs can attach to Alloderm with the dermis side up in a petri dish [15]. ADSC seeded onto Alloderm was also implanted *in vivo* for skin regenerations with promising results [16]. Strattice was evaluated for seeding with rat MSCs [17]. Thus, the stem cell‐seeded biologic graft can be used as a tendon wrap or a skin regeneration material. A study was performed to compare the survival and proliferation of stem cells via bioluminescent imaging. The use of biologic graft

submucosa (SIS) treated with peracetic acid/ethanol (90%

crosslinked with hexamethylene diisocyanate (HMDI)

collagen, 5–10% lipids)

(97% collagen, 2% elastin)

collagen II)

Alloderm Human skin (decelluarized)

Restore 10‐layered porcine small intestine

BioBlanket Crosslinked porous bovine dermis

Strattice Porcine acelluar dermal matrix

**Table 2.** An example list of biologic ECM grafts with potential for tendon/skin regeneration.

TissueMend Noncrosslinked fetal biovine dermal matrix [19]

Permacol or Zimmer collagen repair patch Porcine acelluar dermis treated with trypsin, solvent, and

Cuffpatch EDC‐crosslinked matrix from eight layers of treated SIS

OrthoADAPT Crosslinked equine pericardium (90% collagen I, 10%

GraftJacket Human skin

human rotator cuff [14].

266 Biomaterials in Regenerative Medicine

(5 mm × 2 mm) collagen sponge scaffolds (Zimmer Dental) were used for the culture of (BM) MSCs. Cell‐seeded scaffolds were placed in culture dishes and incubated for 2 hours in a minimum volume of growth medium, after which more medium was applied to submerge the scaffolds. After an additional 24‐hour culture, cells seeded in scaffolds were treated with 10 ng/mL of recombinant (BMP) 12 for 12 hours. The medium was then replaced with fresh growth medium and scaffolds were either cultured for an additional 7 days or immediately implanted into partial calcaneal tendon defects in rats. It was shown that after 21 days, the BMP12‐treated, collagen‐cell scaffold results in robust formation of tendon‐like tissue [22]. Similarly, a collagen carrier combined with ASDCs showed they did not improve the biome‐ chanical properties of the tendon‐to‐bone healing. However, the ADSCs group showed less inflammation, which may lead to a more elastic repair and less scarred healing in a rat model.

#### *2.3.2. Integra bilayer*

Integra® bilayer wound Matrix (LifeSciences Corp., Plainsboro, New Jersey) is a dermal acellular analog composed of bovine collagen type I crosslinked with glycosaminoglycans. Importantly, inclusion of WJ‐MSC into Integra induced significant up‐regulation of prototypical angiogenic and healing factors, stimulating pleiotropic aspects of neovascularization in experimental set‐ tings of angiogenesis, without altering the inflammatory response in the animals, thus dem‐ onstrating their potential benefit in therapeutic care of wounds and skin grafts [23]. Recently, Integra dermal matrix scaffold engineered with adult mesenchymal stem cells and platelet‐rich plasma was investigated *in vitro* and demonstrated promising preliminary results [24].

#### **2.4. Synthetic biopolymer matrix**

Biopolymer matrix such as collagen, gelatin, hyaluronic acid, chitosan, silk, cellulose are the essential ECM components of human, animal, or plant cells. They have showed great promise for attachment, proliferation, and differentiation of stem cells and have been explored for tendon and skin regeneration.

#### *2.4.1. Collagen*

Collagen monomer solution can be extracted from skin of fetal bovine calf in close herd. Collagen‐based biomaterial has been widely used for tendon/skin regeneration. For tendon application, it is highly desirable to prepare aligned collagen fiber scaffolds. Anisotropic col‐ lagen biomaterial can be prepared by directional freeze drying [25], electrospinning, or a novel process called electrochemical process. Electrochemically aligned collagen fiber and skin substrate have been coupled with stem cells for both tendon [26] and skin regeneration [27], respectively. Collagen carrier/gel [28] or collagen combined with stem cell accelerated the wound healing in healing‐impaired db/db mice [29].

and proliferation of the mesenchymal stromal cells *in vitro*. *In vivo* assays performed with mesenchymal stromal cells grown on the surface of the dense membranes showed accelera‐

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A thiol‐ene Michael‐type addition was utilized for rapid encapsulation of MSCs within a gela‐ tin/PEG biomatrix according to Eq. (1). The MSCs/gelatin/PEG biomatrix was applied as a provisional dressing to full‐thickness wounds in Sprague‐Dawley rats. Biomatrix resulted in attenuated immune cell infiltration, lack of foreign giant cell (FBGC) formation, accelerated wound closure and re‐epithelialization, as well as enhanced neovascularization and granula‐

Gelatin ‐ PEG ‐ SH + PEG‐diacrylate → Gelatin ‐ PEG hydrogel (1)

Electrospun nanofiberous scaffolds prepared from silk fibroin protein were seeded with bone marrow‐derived mesenchymal stem cells (MSCs) and epidermal stem cells (ESCs). The con‐ structs were evaluated for wound re‐epithelization, collagen synthesis, as well as the skin appendages regeneration. It was shown that both the transplantation of MSCs and ESCs could significantly accelerate the skin re‐epithelization, stimulate the collagen synthesis. Furthermore, the regenerative features of MSCs and ESCs in activating the blood vessels and hair follicles for‐ mation, respectively, were suggested [35]. Combination of silk with collagen or poly(lactic‐co‐ glycolic acid) (PLGA) [18] and stem cells were evaluated in a rabbit tendon defect model [36].

A stroma skin substitute was first generated by using a mixture of human fibrin obtained from frozen human plasma and 0.1% agarose. An average of 250,000 cultured skin fibro‐ blasts were added to 5 mL of the mixture immediately before inducing the polymerization of the artificial stroma on Transwell (Corning Enterprises, Corning, NY, http://www.corning. com) porous inserts. Once the stromas jellified, human umbilical cord Wharton's jelly stem cells (HWJSCs) were seeded on top of the skin artificial stromas and cultured. It was demon‐ strated that this 3D bioactive scaffold can stratify and form epithelial cell‐like layers and well‐ formed cell‐cell junctions [37]. The authors also demonstrated similar strategy can be used for regeneration of oral mucosa using human Wharton's jelly stem cells (HWJSCs)‐seeded

Sodium CMC was evaluated as a substrate for ADSCs and implanted in adult male Wistar rats. CMC at 10 mg/mL associated with ADSCs increased the rate of cell proliferation of the granulation tissue and epithelium thickness when compared to untreated lesions (Sham).

tion of wound healing in Wistar rats [33].

tion tissue formation by 7 days [34].

*2.4.6. Gelatin/PEG*

*2.4.7. Silk*

*2.4.8. Fibrin‐agarose*

fibrin‐agarose‐mucosal fibroblasts.

*2.4.9. Sodium carboxymethylcellulose (CMC)*

#### *2.4.2. Pullulan‐collagen hydrogel*

A biomimetic pullulan‐collagen hydrogel was used to create a functional biomaterial‐based stem cell niche for the delivery of MSCs into wounds. Murine bone marrow‐derived MSCs were seeded into hydrogels and compared to MSCs grown in standard culture conditions. Hydrogels induced MSC secretion of angiogenic cytokines and expression of transcription factors associated with maintenance of pluripotency and self‐renewal. MSC‐seeded hydro‐ gels showed significantly accelerated healing and a return of skin appendages [30].

#### *2.4.3. Gelatin*

Gelatin is denatured collagen. Human ADSCs laden gelatin microcryogels (GMs) were evalu‐ ated *in vitro* as a stem cell carrier. The cell phenotype markers, stemness genes, differenti‐ ation, secretion of growth factors, cell apoptosis, and cell memory were compared against cells without a carrier. The priming effects of GMs on upregulation of stemness genes and improved secretion of growth factors of hASCs were demonstrated.

#### *2.4.4. Hyaluronic acid (HA)*

Hyaluronic acid (HA) is a nonsulfated, linear polysaccharide with the repeating disaccharide, β‐1,4‐D‐glucuronic acid ‐ β‐1,3‐N‐acetyl‐D‐glucosamine (Mw: 100–8000 KDa). It is an ECM component. ADSCs combined with hyaluronic acid (ADSC‐HA) dermal filler were implanted in rats and compared with HA alone. It was demonstrated that ADSC‐HA has better filling effects than HA alone. A total of 70% of stem cells remain in the injection site after 3 months. These suggested stem cells have the potential to improve the esthetic effects and longevity of dermal fillers [31].

#### *2.4.5. Chitin/Chitosan*

Chitosan‐hyaluron membrane: Hsu et al. investigated adult ADSCs spheroids combined with a chitosan‐hyaluron membrane and showed biomaterial combined with stem cells promoted wound healing in a rat skin repair model [32]. Dense and porous chitosan‐xanthan mem‐ branes seeded with multipotent mesenchymal stromal cells were evaluated for the treatment of skin wounds. The membranes showed to be nonmutagenic and allowed efficient adhesion and proliferation of the mesenchymal stromal cells *in vitro*. *In vivo* assays performed with mesenchymal stromal cells grown on the surface of the dense membranes showed accelera‐ tion of wound healing in Wistar rats [33].

#### *2.4.6. Gelatin/PEG*

*2.4.1. Collagen*

268 Biomaterials in Regenerative Medicine

*2.4.3. Gelatin*

*2.4.4. Hyaluronic acid (HA)*

dermal fillers [31].

*2.4.5. Chitin/Chitosan*

Collagen monomer solution can be extracted from skin of fetal bovine calf in close herd. Collagen‐based biomaterial has been widely used for tendon/skin regeneration. For tendon application, it is highly desirable to prepare aligned collagen fiber scaffolds. Anisotropic col‐ lagen biomaterial can be prepared by directional freeze drying [25], electrospinning, or a novel process called electrochemical process. Electrochemically aligned collagen fiber and skin substrate have been coupled with stem cells for both tendon [26] and skin regeneration [27], respectively. Collagen carrier/gel [28] or collagen combined with stem cell accelerated

A biomimetic pullulan‐collagen hydrogel was used to create a functional biomaterial‐based stem cell niche for the delivery of MSCs into wounds. Murine bone marrow‐derived MSCs were seeded into hydrogels and compared to MSCs grown in standard culture conditions. Hydrogels induced MSC secretion of angiogenic cytokines and expression of transcription factors associated with maintenance of pluripotency and self‐renewal. MSC‐seeded hydro‐

Gelatin is denatured collagen. Human ADSCs laden gelatin microcryogels (GMs) were evalu‐ ated *in vitro* as a stem cell carrier. The cell phenotype markers, stemness genes, differenti‐ ation, secretion of growth factors, cell apoptosis, and cell memory were compared against cells without a carrier. The priming effects of GMs on upregulation of stemness genes and

Hyaluronic acid (HA) is a nonsulfated, linear polysaccharide with the repeating disaccharide, β‐1,4‐D‐glucuronic acid ‐ β‐1,3‐N‐acetyl‐D‐glucosamine (Mw: 100–8000 KDa). It is an ECM component. ADSCs combined with hyaluronic acid (ADSC‐HA) dermal filler were implanted in rats and compared with HA alone. It was demonstrated that ADSC‐HA has better filling effects than HA alone. A total of 70% of stem cells remain in the injection site after 3 months. These suggested stem cells have the potential to improve the esthetic effects and longevity of

Chitosan‐hyaluron membrane: Hsu et al. investigated adult ADSCs spheroids combined with a chitosan‐hyaluron membrane and showed biomaterial combined with stem cells promoted wound healing in a rat skin repair model [32]. Dense and porous chitosan‐xanthan mem‐ branes seeded with multipotent mesenchymal stromal cells were evaluated for the treatment of skin wounds. The membranes showed to be nonmutagenic and allowed efficient adhesion

gels showed significantly accelerated healing and a return of skin appendages [30].

improved secretion of growth factors of hASCs were demonstrated.

the wound healing in healing‐impaired db/db mice [29].

*2.4.2. Pullulan‐collagen hydrogel*

A thiol‐ene Michael‐type addition was utilized for rapid encapsulation of MSCs within a gela‐ tin/PEG biomatrix according to Eq. (1). The MSCs/gelatin/PEG biomatrix was applied as a provisional dressing to full‐thickness wounds in Sprague‐Dawley rats. Biomatrix resulted in attenuated immune cell infiltration, lack of foreign giant cell (FBGC) formation, accelerated wound closure and re‐epithelialization, as well as enhanced neovascularization and granula‐ tion tissue formation by 7 days [34].

$$\text{Gelatin -PEG} \cdot \text{SH} + \text{PEC-diacrylate} \rightarrow \text{Gelatin -PEC hydrogen} \tag{1}$$

#### *2.4.7. Silk*

Electrospun nanofiberous scaffolds prepared from silk fibroin protein were seeded with bone marrow‐derived mesenchymal stem cells (MSCs) and epidermal stem cells (ESCs). The con‐ structs were evaluated for wound re‐epithelization, collagen synthesis, as well as the skin appendages regeneration. It was shown that both the transplantation of MSCs and ESCs could significantly accelerate the skin re‐epithelization, stimulate the collagen synthesis. Furthermore, the regenerative features of MSCs and ESCs in activating the blood vessels and hair follicles for‐ mation, respectively, were suggested [35]. Combination of silk with collagen or poly(lactic‐co‐ glycolic acid) (PLGA) [18] and stem cells were evaluated in a rabbit tendon defect model [36].

#### *2.4.8. Fibrin‐agarose*

A stroma skin substitute was first generated by using a mixture of human fibrin obtained from frozen human plasma and 0.1% agarose. An average of 250,000 cultured skin fibro‐ blasts were added to 5 mL of the mixture immediately before inducing the polymerization of the artificial stroma on Transwell (Corning Enterprises, Corning, NY, http://www.corning. com) porous inserts. Once the stromas jellified, human umbilical cord Wharton's jelly stem cells (HWJSCs) were seeded on top of the skin artificial stromas and cultured. It was demon‐ strated that this 3D bioactive scaffold can stratify and form epithelial cell‐like layers and well‐ formed cell‐cell junctions [37]. The authors also demonstrated similar strategy can be used for regeneration of oral mucosa using human Wharton's jelly stem cells (HWJSCs)‐seeded fibrin‐agarose‐mucosal fibroblasts.

#### *2.4.9. Sodium carboxymethylcellulose (CMC)*

Sodium CMC was evaluated as a substrate for ADSCs and implanted in adult male Wistar rats. CMC at 10 mg/mL associated with ADSCs increased the rate of cell proliferation of the granulation tissue and epithelium thickness when compared to untreated lesions (Sham). CMC is capable to allow the growth of ADSCs and is safe for this biological application up to the concentration of 20 mg/mL. These findings suggest that CMC is a promising biomaterial to be used in cell therapy [38].

#### **2.5. Synthetic nondegradable polymer‐based biomaterial**

**Table 3** summarizes common nondegradable polymer used as a biomaterial or as a modifier of ECM‐based biopolymer for skin and tendon tissue engineering applications. The incorporation of such polymer allows biopolymer (collagen, gelatin, HA) to be able to gelled at physiological condition (e.g., under UV/blue light, ambient temperature, or room temperature aqueous free radical initiator. Stem cells can be directly encapsulated inside the substrate during gel formation.

angiogenesis. Moreover, USCs could secrete vascular endothelial growth factor(VEGF) and transforming growth factor (TGF)‐β1, and USC‐conditioned medium enhanced the migra‐ tion, proliferation, and tube formation of endothelial cells. This data suggested that USCs in combination with PCL/GT significantly prompted the healing of full‐thickness skin wounds in rabbits. Similarly, electrospun poly (L‐lactide‐co‐ε‐caprolactone)/poloxamer (PLCL/P123) scaffolds combined with ADSCs enhanced skin wound healing in a rat model [47]. Chitosan‐ crosslinked poly (3‐hydroxybutyrate‐co‐3‐hydroxyvalerate) (PHBV) was used to load unre‐ stricted somatic stem cells. The cell‐laden scaffold showed better results during the healing

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**PLGA PCL PHBV**

**Table 4.** Common synthetic, degradable polymer used for tendon/skin regeneration applications.

Biomaterials play an important role for attachment, survival, and function of stem cells. Many biomaterials are either used alone or as one component of the product for the regeneration of tendon/ligament/skin. Despite abundance of biomaterial developed, the optimal biomaterials that meet the structural, mechanical, functional requirement of tendon/ligament tissues to be regenerated remain a challenge. Novel biomaterial fabrication process, biomaterial design, and biomaterial synthesis toward tendon/ligament and skin regeneration are urgently needed.

[1] Uysal CA, et al. Adipose‐derived stem cells enhance primary tendon repair: Biomechanical and immunohistochemical evaluation. Journal of Plastic Reconstructive & Aesthetic

process of skin defects in rat models [48, 49].

Address all correspondence to: xcheng@swri.org

Surgery. 2012:**65**(12):1712‐1719

Southwest Research Institute, San Antonio, TX, USA

**3. Conclusion**

**Author details**

Xingguo Cheng

**References**

#### *2.5.1. Poly(NIPAM)‐based*

A biodegradable, multifunctional crosslinker and an n‐isopropylacrylamide (NIPAM)‐based, thermosensitive hydrogel was synthesized to carry BMSCs to treat diabetic skin ulcers. The crosslinker contains an arginylglycylaspartic acid(GRD)‐like motif that promotes cell attach‐ ment and differentiation of BMSCs. After hydrogel association with BMSCs treated the diabetic skin wound in mice, significantly greater wound contraction was observed in the hydrogel + BMSCs group. Histology and immunohistochemistry results confirmed that this treatment contributed to the rapid healing of diabetic skin wounds by promoting granulation tissue formation, angiogenesis, extracellular matrix (ECM) secretion, wound contraction, and re‐epithelialization. These results show that a hydrogel laden with BMSCs may be a promis‐ ing therapeutic strategy for the management of diabetic ulcers [39].

**Table 3.** Common synthetic, nondegradable polymer used as hydrogel or gel components.

#### **2.6. Polyester‐based degradable polymer biomaterial**

The structure of common polyester is shown in **Table 4**. Stem cell‐coated polyester suture was evaluated for tendon applications in a rat model [40]. Electrospun PLGA fiber may be more suit‐ able for tendon regeneration than film [41]. (PGA/PLA) fiber combined with ADSCs improve tendon in a rabbit tendon model [42]. Open cell PLGA seeded with stem cells produced more collagen type I [43]. Knited PLGA encapsulated with stem cell/alginate gel [44]. PGA sheet with MSCs were able to regenerate tendon‐bone insertions and the tendon in rabbit [45]. Electrospun polycaprolactone/gelatin (PCL/GT) membrane and human urine‐derived stem cells (USCs) were evaluated for skin wound healing in a rabbit model [46]. USCs‐PCL/GT‐treated wounds closed much faster, with increased re‐epithelialization, collagen formation, and

**Table 4.** Common synthetic, degradable polymer used for tendon/skin regeneration applications.

angiogenesis. Moreover, USCs could secrete vascular endothelial growth factor(VEGF) and transforming growth factor (TGF)‐β1, and USC‐conditioned medium enhanced the migra‐ tion, proliferation, and tube formation of endothelial cells. This data suggested that USCs in combination with PCL/GT significantly prompted the healing of full‐thickness skin wounds in rabbits. Similarly, electrospun poly (L‐lactide‐co‐ε‐caprolactone)/poloxamer (PLCL/P123) scaffolds combined with ADSCs enhanced skin wound healing in a rat model [47]. Chitosan‐ crosslinked poly (3‐hydroxybutyrate‐co‐3‐hydroxyvalerate) (PHBV) was used to load unre‐ stricted somatic stem cells. The cell‐laden scaffold showed better results during the healing process of skin defects in rat models [48, 49].

## **3. Conclusion**

CMC is capable to allow the growth of ADSCs and is safe for this biological application up to the concentration of 20 mg/mL. These findings suggest that CMC is a promising biomaterial

**Table 3** summarizes common nondegradable polymer used as a biomaterial or as a modifier of ECM‐based biopolymer for skin and tendon tissue engineering applications. The incorporation of such polymer allows biopolymer (collagen, gelatin, HA) to be able to gelled at physiological condition (e.g., under UV/blue light, ambient temperature, or room temperature aqueous free radical initiator. Stem cells can be directly encapsulated inside the substrate during gel formation.

A biodegradable, multifunctional crosslinker and an n‐isopropylacrylamide (NIPAM)‐based, thermosensitive hydrogel was synthesized to carry BMSCs to treat diabetic skin ulcers. The crosslinker contains an arginylglycylaspartic acid(GRD)‐like motif that promotes cell attach‐ ment and differentiation of BMSCs. After hydrogel association with BMSCs treated the diabetic skin wound in mice, significantly greater wound contraction was observed in the hydrogel + BMSCs group. Histology and immunohistochemistry results confirmed that this treatment contributed to the rapid healing of diabetic skin wounds by promoting granulation tissue formation, angiogenesis, extracellular matrix (ECM) secretion, wound contraction, and re‐epithelialization. These results show that a hydrogel laden with BMSCs may be a promis‐

The structure of common polyester is shown in **Table 4**. Stem cell‐coated polyester suture was evaluated for tendon applications in a rat model [40]. Electrospun PLGA fiber may be more suit‐ able for tendon regeneration than film [41]. (PGA/PLA) fiber combined with ADSCs improve tendon in a rabbit tendon model [42]. Open cell PLGA seeded with stem cells produced more collagen type I [43]. Knited PLGA encapsulated with stem cell/alginate gel [44]. PGA sheet with MSCs were able to regenerate tendon‐bone insertions and the tendon in rabbit [45]. Electrospun polycaprolactone/gelatin (PCL/GT) membrane and human urine‐derived stem cells (USCs) were evaluated for skin wound healing in a rabbit model [46]. USCs‐PCL/GT‐treated wounds closed much faster, with increased re‐epithelialization, collagen formation, and

to be used in cell therapy [38].

270 Biomaterials in Regenerative Medicine

*2.5.1. Poly(NIPAM)‐based*

**2.5. Synthetic nondegradable polymer‐based biomaterial**

ing therapeutic strategy for the management of diabetic ulcers [39].

**Table 3.** Common synthetic, nondegradable polymer used as hydrogel or gel components.

**Poly(NIPAM) PEGDA Poly(methyl acrylate) (PMA)**

**2.6. Polyester‐based degradable polymer biomaterial**

Biomaterials play an important role for attachment, survival, and function of stem cells. Many biomaterials are either used alone or as one component of the product for the regeneration of tendon/ligament/skin. Despite abundance of biomaterial developed, the optimal biomaterials that meet the structural, mechanical, functional requirement of tendon/ligament tissues to be regenerated remain a challenge. Novel biomaterial fabrication process, biomaterial design, and biomaterial synthesis toward tendon/ligament and skin regeneration are urgently needed.

## **Author details**

Xingguo Cheng Address all correspondence to: xcheng@swri.org Southwest Research Institute, San Antonio, TX, USA

## **References**

[1] Uysal CA, et al. Adipose‐derived stem cells enhance primary tendon repair: Biomechanical and immunohistochemical evaluation. Journal of Plastic Reconstructive & Aesthetic Surgery. 2012:**65**(12):1712‐1719

[2] Xu K, et al. Platelet‐rich plasma activates tendon‐derived stem cells to promote regenera‐ tion of Achilles tendon rupture in rats. Journal of Tissue Engineering and Regenerative Medicine. 2017;**11**(4):1173‐1184

[17] Gao Y, et al. Methodology of fibroblast and mesenchymal stem cell coating of surgi‐ cal meshes: A pilot analysis. Journal of Biomedical Materials Research Part B Applied

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[18] Lam MT, et al. Effective delivery of stem cells using an extracellular matrix patch results in increased cell survival and proliferation and reduced scarring in skin wound healing.

[19] Liu S, et al. Synergistic angiogenesis promoting effects of extracellular matrix scaffolds and adipose‐derived stem cells during wound repair. Tissue Engineering Part A. 2011;

[20] Mifune Y, et al. Tendon graft revitalization using adult anterior cruciate ligament (ACL)‐ derived CD34+ cell sheets for ACL reconstruction. Biomaterials. 2013;**34**(22):5476‐5487

[21] Kato Y, et al. Allogeneic transplantation of an adipose‐derived stem cell sheet combined with artificial skin accelerates wound healing in a rat wound model of type 2 diabetes

[22] Lee JY, et al. BMP‐12 treatment of adult mesenchymal stem cells *in vitro* augments ten‐ don‐like tissue formation and defect repair *in vivo*. PLoS One. 2011;**6**(3):e17531

[23] Edwards SS, et al. Functional analysis reveals angiogenic potential of human mesenchy‐ mal stem cells from Wharton's jelly in dermal regeneration. Angiogenesis. 2014;**17**(4):

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**Chapter 12**

**Provisional chapter**

**Hydrogels in Regenerative Medicine**

**Hydrogels in Regenerative Medicine**

DOI: 10.5772/intechopen.70409

Polymer scaffolds have many various applications in the field of tissue engineering, drug delivery, and implantation. They are applied as dispensing devices for bioactive molecules and as three-dimensional (3D) structures that provide stimulants that organize cells and direct desired original tissue formation. Hydrogels are preferred scaffolding material because they are structurally similar to the extracellular matrix of many tissues, often processed under mild conditions, and can be delivered in a minimally invasive manner. Hydrogel materials formed a group of polymeric materials. The hydrophilic structure allows them to hold large amounts of water in their three-dimensional backbone. As a result, hydrogels are used as scaffolding material for drug and growth factor transmission, tissue engineering modifications, and many other applications. In this chapter, we describe the physical and chemical structure of hydrogels, side groups, cross-linkings, swelling properties, types of polymers and fabrication methods, and application fields.

**Keywords:** hydrogel, polymers, drug delivery, tissue engineering

preferred as they promote cell attachment and proliferation.

hydrogels in regenerative medicine is aimed.

Yasemin Budama-Kilinc, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and

Yasemin Budama-Kilinc, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Hydrogels are polymer-based substances that contain high levels of water and show the physical characteristic of the extracellular matrix. In regenerative medicine, in terms of the material requirements for tissue scaffold or therapeutic transfer systems, the cell matrix is preferred due to its structural and compound-based similarities. Moreover, hydrogels are

In this chapter, compiling the studies about the preparation of hydrogels and application of

http://dx.doi.org/10.5772/intechopen.70409

Eslin Üstün

**Abstract**

**1. Introduction**

Eslin Üstün

**Provisional chapter**

## **Hydrogels in Regenerative Medicine**

**Hydrogels in Regenerative Medicine**

Yasemin Budama-Kilinc, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and Eslin Üstün Bahar Aslan, Burcu Özkan, Hande Mutlu and Eslin Üstün Additional information is available at the end of the chapter

Yasemin Budama-Kilinc, Rabia Cakir-Koc,

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.70409

#### **Abstract**

Polymer scaffolds have many various applications in the field of tissue engineering, drug delivery, and implantation. They are applied as dispensing devices for bioactive molecules and as three-dimensional (3D) structures that provide stimulants that organize cells and direct desired original tissue formation. Hydrogels are preferred scaffolding material because they are structurally similar to the extracellular matrix of many tissues, often processed under mild conditions, and can be delivered in a minimally invasive manner. Hydrogel materials formed a group of polymeric materials. The hydrophilic structure allows them to hold large amounts of water in their three-dimensional backbone. As a result, hydrogels are used as scaffolding material for drug and growth factor transmission, tissue engineering modifications, and many other applications. In this chapter, we describe the physical and chemical structure of hydrogels, side groups, cross-linkings, swelling properties, types of polymers and fabrication methods, and application fields.

DOI: 10.5772/intechopen.70409

**Keywords:** hydrogel, polymers, drug delivery, tissue engineering

## **1. Introduction**

Hydrogels are polymer-based substances that contain high levels of water and show the physical characteristic of the extracellular matrix. In regenerative medicine, in terms of the material requirements for tissue scaffold or therapeutic transfer systems, the cell matrix is preferred due to its structural and compound-based similarities. Moreover, hydrogels are preferred as they promote cell attachment and proliferation.

In this chapter, compiling the studies about the preparation of hydrogels and application of hydrogels in regenerative medicine is aimed.

## **2. Physical and chemical properties of hydrogels**

#### **2.1. Crosslinking and general characteristic of hydrogels**

Hydrogels are a three-dimensional (3D) polymeric network structure with the ability to absorb water and other biological fluids. The water content of hydrogels determines its specific physicochemical character: mildness, density, and low surface tension of hydrogels [1–4].

hydrophilicity of the polymer and thus increases the water absorption capacity. As a result of the electrostatic repulsive forces of these loads, the solubility and swelling ratios in the web structure are increased. In ionic hydrogels containing acid groups, swelling increases in basic medium, while hydrogels containing basic groups increase swelling in acidic media [9, 10].

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 279

Hydrogels are polymeric webs that absorb and retain water in large quantities. In the polymeric cannula, there are hydrophilic groups or areas that provide a hydrated hydrogel structure in an aqueous environment. As is the case with the term "network," cross-links must be

Physical gel called hydrogels is the result of cross-linking of ionic interaction and physical interactions such as hydrogen bonds, coordination bonds, and hydrophobic interactions [11]. These hydrogels alter the temperature, pH, or solvent composition to form a homogeneous solution and re-gel when they return to their initial conditions. Because of these behaviors, physical gel is also known as reversible gel [12]. They are generally unstable and mechanically weak [13].

Generally, ionic interactions are used to produce physically cross-linked hydrogels. For example, alginate hydrogels are formed with calcium ions. The physical cross-links are found between ionic interactions [14], crystallization [14], amphiphilic block and graft copolymers

The hydrogels in this group are called as irreversible because they are cross-linked by strong chemical bonds between their chains and are not redissolved by changing temperature, pH, or solvent composition. Chemical cross-linked nets have permanent junctions [18, 19]. Chemical cross-linking involves permanent covalent bonds that result in better mechanical strength and more stable hydrogels than physical cross-linking. Incorporation of chemical cross-linkers, however, can cause toxicity problems. Cross-linking of hydrogel can occur with some other method such as radical polymerization, chemical reaction, and high-energy irradiation.

Chemically cross-linked gels can be obtained by radical polymerization of polymers of low molecular weight drugs in the presence of cross-linking agents; poly(2-hydroxyethyl methacrylate) (pHEMA) is a well-known and frequently studied hydrogel system. This hydrogel was first described by Wichterle and Lim [19] and was obtained by polymerization of HEMA in the presence of a suitable cross-linking agent (e.g., ethylene glycol dimethacrylate). A wide

The covalent bonds between the polymer chains can be formed by reaction with complementary reactivity of the functional groups, formation of an amine carboxylic acid, or OH/

 reaction such as isocyanate or Schiff base formation. Chemical cross-linking bonds can occur with aldehydes, with additional reactions and condensation reactions. Water-soluble

variety of hydrogel systems have been synthesized using similar procedures [20].

present to prevent the hydrophilic polymer chains from dissolving in aqueous phase.

**2.3. Cross-linking methods in hydrogel fabrication**

[15], hydrogen bonds [16], and protein interactions [17].

*2.3.1. Physically cross-linked hydrogel*

*2.3.2. Chemical cross-linked hydrogels*

NH3

Hydrogels can be classified according to their preparation methods, ionic charges, physical structures, and cross-linking conditions. According to physical structures, hydrogels can be divided into some groups that are called amorphous hydrogels, semi-crystalline hydrogels, hydrogen-bonded hydrogels, physical hydrogels, and chemical hydrogels.

Various cross-linking approaches, including chemical and physical, have been used to create polymer networks and to protect 3D structures in aqueous media. Physical interactions between the polymer chains in the physically cross-linked gels prevent hydrogel breakdown, while the covalent bonds between the polymer chains in the chemically cross-linked gel form a constant hydrogel.

Physically cross-linked hydrogels are formed by changes in environmental conditions (pH, temperature, and ionic interactions), hydrogen bonds, and protein interactions. Chemically cross-linked gels are obtained by radical polymerization, chemical reactions, energy irradiation, and enzymatic cross-linking [5].

Hydrogels can also have many physical forms: solid mold forms (such as soft contact lenses), pressed powder matrices (pills or capsules taken orally), microparticles (bio-structured carriers or wound treatments), coatings (in implants or in catheters, in pills or capsules, or in capillary electrophoresis), membranes or disks (such as two-dimensional electrophoresis gels), encapsulated solute (in osmotic pumps), and liquids (gel forms that are heated or cooled) [6].

#### **2.2. Hydrogels side groups**

Hydrogels are also called "hydrophilic polymers" because of their water-like properties. Hydrogels have functional groups, such as **─**SO<sup>3</sup> H, **─**COOH, **─**CONH<sup>2</sup> , **─**OH, and **─**NH<sup>2</sup> , which provide hydrophilic character to the polymer chains in their networks. Some of these functional groups (**─**SO<sup>3</sup> and **─**COO hydrogel nets) are negatively charged in aqueous media and can adsorb metal ions through electrostatic interactions [7].

Depending on their chemical and physical structure, there are three types of hydrogels: neutral hydrogels, ionic hydrogels, and interconnected network structures. Neutral (nonionic) hydrogels are homopolymeric or copolymeric hydrogels which do not have charged groups in their structure. Ionic hydrogels, also known as polyelectrolytes, are prepared from ionically charged monomers. These hydrogels are termed cationic and anionic hydrogels, as their monomeric charge is positive and negative [8]. The presence of charged groups in the main chain of ionic hydrogels enhances their susceptibility to stimuli. The presence of ionizable functional groups such as carboxyl and sulfonic acid or ammonium salt increases the hydrophilicity of the polymer and thus increases the water absorption capacity. As a result of the electrostatic repulsive forces of these loads, the solubility and swelling ratios in the web structure are increased. In ionic hydrogels containing acid groups, swelling increases in basic medium, while hydrogels containing basic groups increase swelling in acidic media [9, 10].

#### **2.3. Cross-linking methods in hydrogel fabrication**

Hydrogels are polymeric webs that absorb and retain water in large quantities. In the polymeric cannula, there are hydrophilic groups or areas that provide a hydrated hydrogel structure in an aqueous environment. As is the case with the term "network," cross-links must be present to prevent the hydrophilic polymer chains from dissolving in aqueous phase.

#### *2.3.1. Physically cross-linked hydrogel*

**2. Physical and chemical properties of hydrogels**

hydrogen-bonded hydrogels, physical hydrogels, and chemical hydrogels.

Hydrogels are a three-dimensional (3D) polymeric network structure with the ability to absorb water and other biological fluids. The water content of hydrogels determines its specific physicochemical character: mildness, density, and low surface tension of hydrogels [1–4]. Hydrogels can be classified according to their preparation methods, ionic charges, physical structures, and cross-linking conditions. According to physical structures, hydrogels can be divided into some groups that are called amorphous hydrogels, semi-crystalline hydrogels,

Various cross-linking approaches, including chemical and physical, have been used to create polymer networks and to protect 3D structures in aqueous media. Physical interactions between the polymer chains in the physically cross-linked gels prevent hydrogel breakdown, while the covalent bonds between the polymer chains in the chemically cross-linked gel form

Physically cross-linked hydrogels are formed by changes in environmental conditions (pH, temperature, and ionic interactions), hydrogen bonds, and protein interactions. Chemically cross-linked gels are obtained by radical polymerization, chemical reactions, energy irradia-

Hydrogels can also have many physical forms: solid mold forms (such as soft contact lenses), pressed powder matrices (pills or capsules taken orally), microparticles (bio-structured carriers or wound treatments), coatings (in implants or in catheters, in pills or capsules, or in capillary electrophoresis), membranes or disks (such as two-dimensional electrophoresis gels), encapsulated solute (in osmotic pumps), and liquids (gel forms that are heated or cooled) [6].

Hydrogels are also called "hydrophilic polymers" because of their water-like properties.

which provide hydrophilic character to the polymer chains in their networks. Some of these

Depending on their chemical and physical structure, there are three types of hydrogels: neutral hydrogels, ionic hydrogels, and interconnected network structures. Neutral (nonionic) hydrogels are homopolymeric or copolymeric hydrogels which do not have charged groups in their structure. Ionic hydrogels, also known as polyelectrolytes, are prepared from ionically charged monomers. These hydrogels are termed cationic and anionic hydrogels, as their monomeric charge is positive and negative [8]. The presence of charged groups in the main chain of ionic hydrogels enhances their susceptibility to stimuli. The presence of ionizable functional groups such as carboxyl and sulfonic acid or ammonium salt increases the

H, **─**COOH, **─**CONH<sup>2</sup>

and **─**COO hydrogel nets) are negatively charged in aqueous media

, **─**OH, and **─**NH<sup>2</sup>

,

**2.1. Crosslinking and general characteristic of hydrogels**

a constant hydrogel.

278 Biomaterials in Regenerative Medicine

tion, and enzymatic cross-linking [5].

Hydrogels have functional groups, such as **─**SO<sup>3</sup>

and can adsorb metal ions through electrostatic interactions [7].

**2.2. Hydrogels side groups**

functional groups (**─**SO<sup>3</sup>

Physical gel called hydrogels is the result of cross-linking of ionic interaction and physical interactions such as hydrogen bonds, coordination bonds, and hydrophobic interactions [11]. These hydrogels alter the temperature, pH, or solvent composition to form a homogeneous solution and re-gel when they return to their initial conditions. Because of these behaviors, physical gel is also known as reversible gel [12]. They are generally unstable and mechanically weak [13].

Generally, ionic interactions are used to produce physically cross-linked hydrogels. For example, alginate hydrogels are formed with calcium ions. The physical cross-links are found between ionic interactions [14], crystallization [14], amphiphilic block and graft copolymers [15], hydrogen bonds [16], and protein interactions [17].

#### *2.3.2. Chemical cross-linked hydrogels*

The hydrogels in this group are called as irreversible because they are cross-linked by strong chemical bonds between their chains and are not redissolved by changing temperature, pH, or solvent composition. Chemical cross-linked nets have permanent junctions [18, 19]. Chemical cross-linking involves permanent covalent bonds that result in better mechanical strength and more stable hydrogels than physical cross-linking. Incorporation of chemical cross-linkers, however, can cause toxicity problems. Cross-linking of hydrogel can occur with some other method such as radical polymerization, chemical reaction, and high-energy irradiation.

Chemically cross-linked gels can be obtained by radical polymerization of polymers of low molecular weight drugs in the presence of cross-linking agents; poly(2-hydroxyethyl methacrylate) (pHEMA) is a well-known and frequently studied hydrogel system. This hydrogel was first described by Wichterle and Lim [19] and was obtained by polymerization of HEMA in the presence of a suitable cross-linking agent (e.g., ethylene glycol dimethacrylate). A wide variety of hydrogel systems have been synthesized using similar procedures [20].

The covalent bonds between the polymer chains can be formed by reaction with complementary reactivity of the functional groups, formation of an amine carboxylic acid, or OH/ NH3 reaction such as isocyanate or Schiff base formation. Chemical cross-linking bonds can occur with aldehydes, with additional reactions and condensation reactions. Water-soluble polymers owe their solubility properties to the presence of functional groups (mainly **─**OH, COOH, NH) which can be used for hydrogel formation.

*2.5.1. Natural and synthetic polymers*

tides, and leucine [37] zipper;

lose, and gelatin/alginate [39]; and

among these three main types.

*2.5.2. Biodegradable synthetic polymers*

gels by chemical or physical cross-linking [51, 52].

• DNA [40].

Natural polymers have been used as scaffolds for tissue engineering as natural hydrogels, regarding their biological compatibility, natural biodegradability, and critical biological func-

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 281

• proteins such as collagen, silk, gelatin, fibrin, lysozyme, Matrigel™, and genetically engineered proteins [34–36] like calmodulin (a calcium-binding protein), elastin-like polypep-

• protein/polysaccharide hybrid polymers like collagen/HA, laminin/chitosan, fibrin/cellu-

Compared to natural polymers, synthetic polymers have more reproducible physical and chemical properties that are critical for the construction of scaffolding. At present, synthetic polymers have emerged as an important alternative option in the manufacture of hydrogel tissue-engineered skeletons, since they can be molecularly modified by block structures, molecular weights, mechanical strength, and biodegradability [41–43]. Synthetic polymers used in the preparation of synthetic hydrogels can be classified into three main types: nonbiodegradable [44, 45], biodegradable [46], and bioactive polymers [47]. In the following section, the biodegradable synthetic polymers are disclosed because of its widespread use from

Biodegradability is one of the most important aspects of scaffolds for tissue engineering. It is highly desirable to ensure that the biodegradation rate coincides with new tissue regeneration at the defect site [48, 49]. Many polymers, proteins, cellulose, starch, and chitin created in the environment are limited to making hydrogel skeletons that are biodegradable but have their own biodegradability and mechanical properties. Most common biodegradable synthetic polymer groups contain poly (lactic acid) (PLA), poly(ε-caprolactone) (PCL), poly(glycolic acid) (PGA) and copolymers [50]. They can be used to modify hydrophilic polymers such as PEG to form acrylate macromers or amphiphilic polymers to produce biodegradable hydro-

The mechanical stability of the gel is an important feature when a scaffold is designed. The strength of the hydrogels can be increased by incorporating cross-linking agents, comonomers, because the higher cross-linking degree provides more fragility and less elasticity.

Biodegradable synthetic hydrogels include various vinylic monomers such as 2-hydroxyethyl methacrylate (HEMA), N-isopropylacrylamide (NIPAM), 2-hydroxypropyl methacrylate (HPMA), acrylamide (AAm), acrylic acid (AAc) or macromers [53–55] N,N′-methylenebis and methoxyl poly(ethylene glycol) (PEG), monoacrylate (mPEGMA or PEGMA), and diacrylate (PEGDA) with cross-linkers such as ethylene glycol diacrylate (EGDA), and these

• polysaccharides like hyaluronic acid (HA), dextran, agarose, and chitosan [38].

tions. There are four main types of natural polymers, including:

High-energy radiation, especially gamma and electron beams, can be used to polymerize unsaturated compounds. This means that water-soluble polymers derivatized with vinyl groups can be converted into hydrogels by high-energy irradiation [21]. Hydrogels may also be obtained by radiation-induced polymerization of a mixture consisting of a monofunctional acrylate (e.g., acryloyl-L-proline methyl ester) and a suitable cross-linker [22]. Moreover, highenergy irradiation cross-links water-soluble polymers without additional groups. During irradiation of aqueous solutions of polymers (gamma or electron transfer), the radicals are generally present in the polymer chain, e.g., homolytic cleavage of C–H bonds.

#### **2.4. Water content and resources**

Hydrogels can generally hold 20–100 times water and their own weight depends on their structures [23]. If hydrogels can hold higher than 100%, they are called superabsorbent hydrogels. According to the water contents, hydrogels can be categorized as low swelling grade hydrogels (20–50%), medium swelling grade hydrogels (50–90%), high swelling grade hydrogels (90–99.5%), and superabsorbent hydrogels (>99.5%) [24, 25].

High water content hydrogels are usually more useful in the medical field because of permeability and biocompatibility [26]. Synthesis of biodegradable hydrogels with both good mechanical properties and high water content is still a major challenge, because as the amount of water increases, the mechanical properties of hydrogel become weak [27].

Because water content of hydrogels is high, these gels resemble natural living tissues more than synthetic biological materials, so the biocompatibility is higher. When the polymer network contacts the aqueous solutions, swelling occurs to achieve thermodynamic stability due to the water concentration gradient [28].

The hydrophilicity of the polymer chain and the cross-link density are parameters effective in hydrogel swelling. Hydrophilic groups of polymer networks in hydrogels are groups that like water. Therefore, they have a high affinity for water. These groups allow hydrogels to swell in water, become soft, and gain elastic properties [29]. Hydrophilic group-containing hydrogels swell at a higher rate than hydrophobic group-containing hydrogels. Because hydrophobic groups do not like water, they prevent water molecules from penetrating into the structure [30].

The structures of the high cross-linked hydrogels are more rigid, and cross-linking inhibits the movement of the polymer chain. When the amount of cross-linking agent increases, the hydrogel becomes less flexible, and the swelling ability of the hydrogel is reduced [31, 32].

#### **2.5. Polymers used in hydrogel fabrication**

Polymer materials have many uses such as catheters, heart valves, contact and intraocular lenses, dental prostheses, ophthalmic applications, urology, and gastroenterology applications [33]. In tissue engineering different polymers are used for different tissues. It is possible to classify them in four different groups.

#### *2.5.1. Natural and synthetic polymers*

Natural polymers have been used as scaffolds for tissue engineering as natural hydrogels, regarding their biological compatibility, natural biodegradability, and critical biological functions. There are four main types of natural polymers, including:


polymers owe their solubility properties to the presence of functional groups (mainly **─**OH,

High-energy radiation, especially gamma and electron beams, can be used to polymerize unsaturated compounds. This means that water-soluble polymers derivatized with vinyl groups can be converted into hydrogels by high-energy irradiation [21]. Hydrogels may also be obtained by radiation-induced polymerization of a mixture consisting of a monofunctional acrylate (e.g., acryloyl-L-proline methyl ester) and a suitable cross-linker [22]. Moreover, highenergy irradiation cross-links water-soluble polymers without additional groups. During irradiation of aqueous solutions of polymers (gamma or electron transfer), the radicals are

Hydrogels can generally hold 20–100 times water and their own weight depends on their structures [23]. If hydrogels can hold higher than 100%, they are called superabsorbent hydrogels. According to the water contents, hydrogels can be categorized as low swelling grade hydrogels (20–50%), medium swelling grade hydrogels (50–90%), high swelling grade hydro-

High water content hydrogels are usually more useful in the medical field because of permeability and biocompatibility [26]. Synthesis of biodegradable hydrogels with both good mechanical properties and high water content is still a major challenge, because as the amount

Because water content of hydrogels is high, these gels resemble natural living tissues more than synthetic biological materials, so the biocompatibility is higher. When the polymer network contacts the aqueous solutions, swelling occurs to achieve thermodynamic stability due

The hydrophilicity of the polymer chain and the cross-link density are parameters effective in hydrogel swelling. Hydrophilic groups of polymer networks in hydrogels are groups that like water. Therefore, they have a high affinity for water. These groups allow hydrogels to swell in water, become soft, and gain elastic properties [29]. Hydrophilic group-containing hydrogels swell at a higher rate than hydrophobic group-containing hydrogels. Because hydrophobic groups do not like water, they prevent water molecules from penetrating into the structure [30]. The structures of the high cross-linked hydrogels are more rigid, and cross-linking inhibits the movement of the polymer chain. When the amount of cross-linking agent increases, the hydrogel becomes less flexible, and the swelling ability of the hydrogel is reduced [31, 32].

Polymer materials have many uses such as catheters, heart valves, contact and intraocular lenses, dental prostheses, ophthalmic applications, urology, and gastroenterology applications [33]. In tissue engineering different polymers are used for different tissues. It is possible

generally present in the polymer chain, e.g., homolytic cleavage of C–H bonds.

gels (90–99.5%), and superabsorbent hydrogels (>99.5%) [24, 25].

of water increases, the mechanical properties of hydrogel become weak [27].

COOH, NH) which can be used for hydrogel formation.

**2.4. Water content and resources**

280 Biomaterials in Regenerative Medicine

to the water concentration gradient [28].

**2.5. Polymers used in hydrogel fabrication**

to classify them in four different groups.

Compared to natural polymers, synthetic polymers have more reproducible physical and chemical properties that are critical for the construction of scaffolding. At present, synthetic polymers have emerged as an important alternative option in the manufacture of hydrogel tissue-engineered skeletons, since they can be molecularly modified by block structures, molecular weights, mechanical strength, and biodegradability [41–43]. Synthetic polymers used in the preparation of synthetic hydrogels can be classified into three main types: nonbiodegradable [44, 45], biodegradable [46], and bioactive polymers [47]. In the following section, the biodegradable synthetic polymers are disclosed because of its widespread use from among these three main types.

#### *2.5.2. Biodegradable synthetic polymers*

Biodegradability is one of the most important aspects of scaffolds for tissue engineering. It is highly desirable to ensure that the biodegradation rate coincides with new tissue regeneration at the defect site [48, 49]. Many polymers, proteins, cellulose, starch, and chitin created in the environment are limited to making hydrogel skeletons that are biodegradable but have their own biodegradability and mechanical properties. Most common biodegradable synthetic polymer groups contain poly (lactic acid) (PLA), poly(ε-caprolactone) (PCL), poly(glycolic acid) (PGA) and copolymers [50]. They can be used to modify hydrophilic polymers such as PEG to form acrylate macromers or amphiphilic polymers to produce biodegradable hydrogels by chemical or physical cross-linking [51, 52].

The mechanical stability of the gel is an important feature when a scaffold is designed. The strength of the hydrogels can be increased by incorporating cross-linking agents, comonomers, because the higher cross-linking degree provides more fragility and less elasticity.

Biodegradable synthetic hydrogels include various vinylic monomers such as 2-hydroxyethyl methacrylate (HEMA), N-isopropylacrylamide (NIPAM), 2-hydroxypropyl methacrylate (HPMA), acrylamide (AAm), acrylic acid (AAc) or macromers [53–55] N,N′-methylenebis and methoxyl poly(ethylene glycol) (PEG), monoacrylate (mPEGMA or PEGMA), and diacrylate (PEGDA) with cross-linkers such as ethylene glycol diacrylate (EGDA), and these nonbiodegradable polymers such as self-assembly of Pluronic® polymers that have poly(ethylene oxide) (PEO)-poly(propylene oxide) (PPO)-PEO structure [53–55] are modified with poly(vinyl alcohol) (PVA) via chemical cross-linking.

Hydrogels are administered as low-viscosity liquids that penetrate the dentin tubules when exposed to capillary action. Treatment with a gel that forms a long-term block is provided when it is swollen and wet; at this point it is also permeable to oxygen and electrolytes [67]. The use of hydrogels in dental equipment offers several advantages over other biomaterials such as (1) long-term grip despite the effects of surface abrasion, (2) its softness removes the pressure (3) brushing and abrasion only cleanse the exposed parts of the gel and leave the remainder in the tubules for a long time (4), the hardened gel is colorless and therefore maintains the normal appearance of the tooth, (5) the gel structure continues to be permeable to oxygen and electrolytes, and (6) there are no hydraulic pressures that can cause pain without

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 283

Oral hydrogels are drug-free, are completely natural as wound dressings, and are approved by the FDA (Food and Drug Administration) for all oral injuries. They can be used in tooth

Depending on the hydrostatic pressure, the sharp fluid movements in these micro-channels can cause dentin pain. The microorganisms present in the oral cavity ferment carbohydrates to form lactic acid, which causes the pH to drop. The acidic medium then causes the hydroxyapatite to dissolve in the hard walls of the micro-channels. This increases the permeability of the channels and facilitates the penetration of microorganisms. The spread of bacterial toxins causes an inflammatory reaction in the pulp, which causes symptoms such as hypersensitivity and pain. For this reason, the development of new prophylactic agents and methods to reduce tooth decay and an overly sensitive tooth neck is of great importance [70]. Ionic hydrogels, especially nanogels, have started being used in toothpastes. Micro-channels filled with a fibrous protein and a bio-hydrogel composed of a liquid phase enter the human teeth. The gel network is bound to the organic collagen matrix and the hydroxyapatite crystals on

Conventional drug administration can result in toxicity and side effects because of requirement of high dosages and repeated administration to stimulate a therapeutic effect, and these can reduce efficacy and patient compliance [72–74]. Researchers have focused on controlled drug delivery systems during the recent years for reducing toxicity and by enhancing efficacy of therapeutics [75]. Drug release is controlled by drug delivery systems in terms of transpor-

Hydrogels have been considerably researched as the carrier for drug delivery systems [78]. They have low interfacial tension; thus, they exhibit minimum tendency to adsorb proteins from body fluids [79, 80]. These biomaterials have paid attention due to their peculiar properties such as temperature and pH sensitivity, swelling in aqueous medium, and sensitivity toward other stimuli [78]. Their highly porous structure can be modified through controlling the density of cross-links in the gel matrix and the affinity of the hydrogels for the aqueous environment in which they are swollen. Drug release at a rate dependent on the diffusion coefficient of the small molecule or macromolecule through the gel network and loading of

drugs into the gel matrix can be provided due to their porous structure [80, 81].

extractions, grafts, implants, lesions, or for wounds in the mouth [69].

fluid flow [68].

the walls of the dentin tubules [71].

**3.2. Applications of hydrogels in drug delivery**

tation to tissues and cells in time and in space [76, 77].

Synthetic polymers can be easily synthesized on a large scale and manipulated at the molecular level by polymerization, cross-linking, and functionalization. However, most synthetic hydrogels alone generally function as passive scaffolds only for cells and do not increase active cellular interactions [56]. As already mentioned, natural polymers such as proteins exhibit different tertiary structures and regulate active cellular response, biological identification, and cell-triggered remodeling. Thus, combining the properties of synthetic and natural polymers to form hybrid hydrogels has become a direct approach to creating bioactive hydrogel scaffolds for tissue engineering. These same examples of hybrid hydrogel polymers include:


The hydrogel is obtained in different forms depending on the application area by selecting the appropriate polymer. It is possible to produce hydrogels which can be heat-adjusted, injectable, or in the form of film, foam, or gel. These can be controlled by different parameters such as cross-linking of structures and synthesis of polymers.

## **3. Application of hydrogels**

The hydrogels with various features are widely used as wound dressing, drug delivery, dental equipment, pharmaceutic industry, injectable polymer systems, implants, and tissue engineering.

#### **3.1. Applications of hydrogels in dentistry**

In recent years, injectable hydrogels have emerged as a promising biomaterial for the therapeutic delivery of cells and bioactive molecules for tissue regeneration in dentistry. The controllability of deterioration and release behavior, its adaptability in the clinical setting for minimally invasive surgical procedures, and its ability to remain three-dimensional (3D) after gelling mean that these hydrogels can be used in medical practice [64]. In dental tissue engineering, injectable hydrogels are used to replace countless tissues, especially cartilage, bones, nerves, blood vessels, and soft tissues. Furthermore, it has been demonstrated that the feasibility of introducing dental pulp stem cells with hydrogels has supported the matrix and growth factors [65].

Hydrogel is biologically compatible and absorbed into the body without any side effects. The gel is completely disintegrated so no by-products or new chemicals are formed. For these reasons, they are preferred as dental fillers, as dental equipment, and even as coating for dental implants [66].

Hydrogels are administered as low-viscosity liquids that penetrate the dentin tubules when exposed to capillary action. Treatment with a gel that forms a long-term block is provided when it is swollen and wet; at this point it is also permeable to oxygen and electrolytes [67]. The use of hydrogels in dental equipment offers several advantages over other biomaterials such as (1) long-term grip despite the effects of surface abrasion, (2) its softness removes the pressure (3) brushing and abrasion only cleanse the exposed parts of the gel and leave the remainder in the tubules for a long time (4), the hardened gel is colorless and therefore maintains the normal appearance of the tooth, (5) the gel structure continues to be permeable to oxygen and electrolytes, and (6) there are no hydraulic pressures that can cause pain without fluid flow [68].

Oral hydrogels are drug-free, are completely natural as wound dressings, and are approved by the FDA (Food and Drug Administration) for all oral injuries. They can be used in tooth extractions, grafts, implants, lesions, or for wounds in the mouth [69].

Depending on the hydrostatic pressure, the sharp fluid movements in these micro-channels can cause dentin pain. The microorganisms present in the oral cavity ferment carbohydrates to form lactic acid, which causes the pH to drop. The acidic medium then causes the hydroxyapatite to dissolve in the hard walls of the micro-channels. This increases the permeability of the channels and facilitates the penetration of microorganisms. The spread of bacterial toxins causes an inflammatory reaction in the pulp, which causes symptoms such as hypersensitivity and pain. For this reason, the development of new prophylactic agents and methods to reduce tooth decay and an overly sensitive tooth neck is of great importance [70]. Ionic hydrogels, especially nanogels, have started being used in toothpastes. Micro-channels filled with a fibrous protein and a bio-hydrogel composed of a liquid phase enter the human teeth. The gel network is bound to the organic collagen matrix and the hydroxyapatite crystals on the walls of the dentin tubules [71].

#### **3.2. Applications of hydrogels in drug delivery**

nonbiodegradable polymers such as self-assembly of Pluronic® polymers that have poly(ethylene oxide) (PEO)-poly(propylene oxide) (PPO)-PEO structure [53–55] are modified

Synthetic polymers can be easily synthesized on a large scale and manipulated at the molecular level by polymerization, cross-linking, and functionalization. However, most synthetic hydrogels alone generally function as passive scaffolds only for cells and do not increase active cellular interactions [56]. As already mentioned, natural polymers such as proteins exhibit different tertiary structures and regulate active cellular response, biological identification, and cell-triggered remodeling. Thus, combining the properties of synthetic and natural polymers to form hybrid hydrogels has become a direct approach to creating bioactive hydrogel scaffolds for tissue engineering. These same examples of hybrid hydrogel polymers include:

• PEG-modified natural polymers [57–60], like fibrinogen, heparin, dextran, HA, and albumin;

The hydrogel is obtained in different forms depending on the application area by selecting the appropriate polymer. It is possible to produce hydrogels which can be heat-adjusted, injectable, or in the form of film, foam, or gel. These can be controlled by different parameters such

The hydrogels with various features are widely used as wound dressing, drug delivery, dental equipment, pharmaceutic industry, injectable polymer systems, implants, and tissue

In recent years, injectable hydrogels have emerged as a promising biomaterial for the therapeutic delivery of cells and bioactive molecules for tissue regeneration in dentistry. The controllability of deterioration and release behavior, its adaptability in the clinical setting for minimally invasive surgical procedures, and its ability to remain three-dimensional (3D) after gelling mean that these hydrogels can be used in medical practice [64]. In dental tissue engineering, injectable hydrogels are used to replace countless tissues, especially cartilage, bones, nerves, blood vessels, and soft tissues. Furthermore, it has been demonstrated that the feasibility of introducing dental pulp stem cells with hydrogels has supported the matrix and growth factors [65].

Hydrogel is biologically compatible and absorbed into the body without any side effects. The gel is completely disintegrated so no by-products or new chemicals are formed. For these reasons, they are preferred as dental fillers, as dental equipment, and even as coating for dental

• PNIPAm-modified natural polymers, like collagen, chitosan, and alginate [61];

• PVA and other synthetic polymer (e.g., pluronic)-modified natural polymers [63].

• Synthetic peptide-modified proteins or polysaccharides [62]; and

as cross-linking of structures and synthesis of polymers.

**3. Application of hydrogels**

**3.1. Applications of hydrogels in dentistry**

engineering.

implants [66].

with poly(vinyl alcohol) (PVA) via chemical cross-linking.

282 Biomaterials in Regenerative Medicine

Conventional drug administration can result in toxicity and side effects because of requirement of high dosages and repeated administration to stimulate a therapeutic effect, and these can reduce efficacy and patient compliance [72–74]. Researchers have focused on controlled drug delivery systems during the recent years for reducing toxicity and by enhancing efficacy of therapeutics [75]. Drug release is controlled by drug delivery systems in terms of transportation to tissues and cells in time and in space [76, 77].

Hydrogels have been considerably researched as the carrier for drug delivery systems [78]. They have low interfacial tension; thus, they exhibit minimum tendency to adsorb proteins from body fluids [79, 80]. These biomaterials have paid attention due to their peculiar properties such as temperature and pH sensitivity, swelling in aqueous medium, and sensitivity toward other stimuli [78]. Their highly porous structure can be modified through controlling the density of cross-links in the gel matrix and the affinity of the hydrogels for the aqueous environment in which they are swollen. Drug release at a rate dependent on the diffusion coefficient of the small molecule or macromolecule through the gel network and loading of drugs into the gel matrix can be provided due to their porous structure [80, 81].

Hydrogels have also been served as drug protectors particularly for proteins and peptides from in vivo environment. Furthermore, these swollen polymers are beneficial as targetable carriers for bioactive drugs with tissue specificity [78]. In general hydrogels are divided for their morphology, swelling property, and elasticity. The release mechanism of the drug from the swollen polymeric material is determined by swelling, and also elasticity affects both the stability of these drug carriers and the mechanical strength of the network [82].

devices can be designed to administer drugs locally to the specific areas in the GI tract. Several hydrogels are currently being investigated as potential devices for colon-specific drug delivery. These include chemically or physically cross-linked polysaccharides, such as dextran, guar gum, azo cross-linked poly(acrylic acid), amidated pectin, and inulin [91–93]. They are designed to be highly swollen or degraded in the presence of colonic enzymes or microflora,

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 285

The eye has shielding mechanisms at work, like tear drainage, the power of blinking, and low permeability of the cornea. These are some of the many physiological constraints preventing

The usual eye drops that contain a drug solution are most likely to be eliminated quickly from the eye; thus, medication administered will have limited absorption. This in turn leads to poor ophthalmic bioavailability (the degree to which a drug or other substances is absorbed at the

Researchers have been motivated to come up with a system that delivers the drug for a longer stay time in the eye. Suspension and ointments which are some of the dosage forms can remain in the eye, but they also can be uncomfortable at times due to the unpleasant feeling

Hydrogels have elastic properties, and this is why they can also represent an ocular drainageresistant device. Additionally, they could give a better feeling to patients, with less gritty sensations. In particular, in situ forming hydrogels are attractive as an ocular drug delivery system because of their facility in dosing as a liquid, and their long-term retention property as

Drug delivery to the skin has been traditionally conducted for a specific area use of dermatological drugs to treat skin diseases or for disinfection of the skin itself. Longer duration and a constant rate are among the benefits of transdermal drug delivery that drug delivery can be easily canceled on demand by simply removing the devices. Thus, since hydrogels are high in water content, swollen hydrogels can provide a better feeling for the skin when compared to

Subcutaneously inserted exogenous materials may more or less evoke potentially undesirable body responses, such as inflammation, carcinogenicity, and immunogenicity. Therefore,

Due to their high water content, hydrogels are generally considered as biocompatible materials. Hydrogels have been used to improve cellular adhesion [99]. They also provide several promising properties: minimal mechanical irritation upon living organism implantation, due

providing colon specificity in drug delivery [94].

site of physiological activity after being given) [96].

given to the eye from characteristics of solids and semisolids [97].

biocompatibility is a prerequisite that makes materials implantable.

a successful drug delivery to the eye [95].

a gel after dosage has been given [98].

traditional ointments and patches [96].

*3.2.5. Transdermal delivery*

*3.2.6. Subcutaneous delivery*

*3.2.4. Ocular delivery*

The body has environmental variables, known as low pH and high temperatures. In this endeavor, it either includes physical, such as temperature, or chemical, such as in pH, ions which can be used for site-specific controlled drug delivery [80, 83]. Due to their ability to "sense" changes in environmental properties, stimuli-sensitive hydrogels, also known as "smart" hydrogels, are very different from inactive hydrogels. Their response will be to increase or decrease their degree of swelling [84, 85]. "Smart" hydrogels have a property of changing the intensity which is particularly useful in the application of drug delivery as drug release can be set off upon environmental changes.

Besides, molecules of different sizes have the ability to go into and out of hydrogels (going in is called drug loading and going out of is called drug release) which makes them permissible for dry or swollen polymeric networks as drug delivery systems for nasal, oral, ocular, buccal, parenteral, epidermal, and subcutaneous routes of administration [86].

#### *3.2.1. Peroral drug delivery*

Up to now, pharmaceutical application of hydrogels has been most commonly done orally. Taken orally, hydrogels are able to deliver drugs to four major areas: the mouth, stomach, small intestine, and colon. If their swelling properties or bioadhesive abilities are controlled, hydrogels can be useful in the release of drugs in a controlled manner at these desired areas. Additionally, they can also stick to certain specific regions in the oral pathway, leading to a locally enhanced drug concentration and, thus, promoting the drug absorption at the release site [87].

#### *3.2.2. Drug delivery in the oral cavity*

Drug delivery to the oral cavity can have many functions. Hydrogel are used ın the local treatment of diseases pertaining to the mouth (i.e., periodontal, stomatitis, viral and fungal infections) and oral cavity cancers. With this in mind, many types of bioadhesive hydrogel systems have been devised since the early 1980s. To name one, a hydrogel-based ointment can also be used for the topical treatment of certain diseases in the oral cavity. It can be used not only as a drug delivery device but also as a liposome delivery medium [88, 89].

#### *3.2.3. Drug delivery in the gastrointestinal (GI) tract*

The GI tract is undoubtedly the most popular route of drug delivery because of the easiness of administration of drugs for a manageable therapy and its large surface area for absorption to the entire body. It is, however, the most complex route, so that changeable approaches are needed to give drugs for effective therapy [90]. Like buccal delivery, hydrogel-based devices can be designed to administer drugs locally to the specific areas in the GI tract. Several hydrogels are currently being investigated as potential devices for colon-specific drug delivery. These include chemically or physically cross-linked polysaccharides, such as dextran, guar gum, azo cross-linked poly(acrylic acid), amidated pectin, and inulin [91–93]. They are designed to be highly swollen or degraded in the presence of colonic enzymes or microflora, providing colon specificity in drug delivery [94].

#### *3.2.4. Ocular delivery*

Hydrogels have also been served as drug protectors particularly for proteins and peptides from in vivo environment. Furthermore, these swollen polymers are beneficial as targetable carriers for bioactive drugs with tissue specificity [78]. In general hydrogels are divided for their morphology, swelling property, and elasticity. The release mechanism of the drug from the swollen polymeric material is determined by swelling, and also elasticity affects both the

The body has environmental variables, known as low pH and high temperatures. In this endeavor, it either includes physical, such as temperature, or chemical, such as in pH, ions which can be used for site-specific controlled drug delivery [80, 83]. Due to their ability to "sense" changes in environmental properties, stimuli-sensitive hydrogels, also known as "smart" hydrogels, are very different from inactive hydrogels. Their response will be to increase or decrease their degree of swelling [84, 85]. "Smart" hydrogels have a property of changing the intensity which is particularly useful in the application of drug delivery as drug

Besides, molecules of different sizes have the ability to go into and out of hydrogels (going in is called drug loading and going out of is called drug release) which makes them permissible for dry or swollen polymeric networks as drug delivery systems for nasal, oral, ocular, buccal,

Up to now, pharmaceutical application of hydrogels has been most commonly done orally. Taken orally, hydrogels are able to deliver drugs to four major areas: the mouth, stomach, small intestine, and colon. If their swelling properties or bioadhesive abilities are controlled, hydrogels can be useful in the release of drugs in a controlled manner at these desired areas. Additionally, they can also stick to certain specific regions in the oral pathway, leading to a locally enhanced drug concentration and, thus, promoting the drug absorption at the release

Drug delivery to the oral cavity can have many functions. Hydrogel are used ın the local treatment of diseases pertaining to the mouth (i.e., periodontal, stomatitis, viral and fungal infections) and oral cavity cancers. With this in mind, many types of bioadhesive hydrogel systems have been devised since the early 1980s. To name one, a hydrogel-based ointment can also be used for the topical treatment of certain diseases in the oral cavity. It can be used not only as

The GI tract is undoubtedly the most popular route of drug delivery because of the easiness of administration of drugs for a manageable therapy and its large surface area for absorption to the entire body. It is, however, the most complex route, so that changeable approaches are needed to give drugs for effective therapy [90]. Like buccal delivery, hydrogel-based

stability of these drug carriers and the mechanical strength of the network [82].

parenteral, epidermal, and subcutaneous routes of administration [86].

a drug delivery device but also as a liposome delivery medium [88, 89].

release can be set off upon environmental changes.

*3.2.1. Peroral drug delivery*

284 Biomaterials in Regenerative Medicine

*3.2.2. Drug delivery in the oral cavity*

*3.2.3. Drug delivery in the gastrointestinal (GI) tract*

site [87].

The eye has shielding mechanisms at work, like tear drainage, the power of blinking, and low permeability of the cornea. These are some of the many physiological constraints preventing a successful drug delivery to the eye [95].

The usual eye drops that contain a drug solution are most likely to be eliminated quickly from the eye; thus, medication administered will have limited absorption. This in turn leads to poor ophthalmic bioavailability (the degree to which a drug or other substances is absorbed at the site of physiological activity after being given) [96].

Researchers have been motivated to come up with a system that delivers the drug for a longer stay time in the eye. Suspension and ointments which are some of the dosage forms can remain in the eye, but they also can be uncomfortable at times due to the unpleasant feeling given to the eye from characteristics of solids and semisolids [97].

Hydrogels have elastic properties, and this is why they can also represent an ocular drainageresistant device. Additionally, they could give a better feeling to patients, with less gritty sensations. In particular, in situ forming hydrogels are attractive as an ocular drug delivery system because of their facility in dosing as a liquid, and their long-term retention property as a gel after dosage has been given [98].

#### *3.2.5. Transdermal delivery*

Drug delivery to the skin has been traditionally conducted for a specific area use of dermatological drugs to treat skin diseases or for disinfection of the skin itself. Longer duration and a constant rate are among the benefits of transdermal drug delivery that drug delivery can be easily canceled on demand by simply removing the devices. Thus, since hydrogels are high in water content, swollen hydrogels can provide a better feeling for the skin when compared to traditional ointments and patches [96].

#### *3.2.6. Subcutaneous delivery*

Subcutaneously inserted exogenous materials may more or less evoke potentially undesirable body responses, such as inflammation, carcinogenicity, and immunogenicity. Therefore, biocompatibility is a prerequisite that makes materials implantable.

Due to their high water content, hydrogels are generally considered as biocompatible materials. Hydrogels have been used to improve cellular adhesion [99]. They also provide several promising properties: minimal mechanical irritation upon living organism implantation, due to their soft, elastic properties; prevention of protein adsorption and cell adhesion arising from the low interfacial tension between water and hydrogels; broad acceptability for individual drugs with different hydrophilicities and molecular sizes; and unique possibilities (cross-linking density and swelling) to manipulate the release of incorporated drugs. Some of these may offer an advantage for the delivery of certain delicate drugs, such as peptides and proteins [73, 80].

largest region—the cartilaginous tissue has a low density of round chondrocytes surrounded by an ECM consisting of hard, avascular glycosaminoglycan (GAG) and type II collagen [108, 109]. Hydrogel design to improve cartilage repair has progressed in recent years. Some using hydrogels include the development of improved network cross-linking (e.g., double networks), new techniques for processing hydrogels (e.g., 3D printing), and better incorporation of biologi-

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 287

Hydrogel has emerged as a promising port because of its potential features in a wide range

The meniscus plays an important role in maintaining homeostasis of the knee joint [110]. Meniscus lesions are tight, have poor healing ability, and can produce tibiofemoral osteoarthritis. Current options for reconstructive treatment focus primarily on the treatment of lesions in the peripheral vascularized area [111]. On the contrary, a small number of approaches can stimulate the repair of damaged, meniscal tissue in the middle, avascular part. Tissue engineering approaches are of great interest to repair or replace damaged meniscus tissue in this

Hydrogel-based biomaterials show a special interest for meniscus repair, as their cores contain relatively high rates of proteoglycans responsible for viscoelastic compaction properties and degree of hydration. Hydrogels that exhibit high water content and provide a specific three-dimensional (3D) microenvironment can be designed such that the meniscus tissue can

Different polymers of both natural and synthetic origin have been manipulated to produce hydrogels that host the respective cell populations for meniscus regeneration and to provide platforms for meniscus tissue replacement. To date, these compounds have been used to design controlled delivery systems of bioactive molecules involved in the meniscus repair

Significant advances have been made in repair techniques for tendon injuries in recent years, but the treatment of finger flexor tendon injury is still one of the most difficult and important

The main problem in tendon repair is adhesion between the tendon and surrounding tissue. We note that polyvinyl alcohol hydrogel (PVA-H) has a high water content with both adhesion-inhibiting function and synovial perfusion function. PVA-H has been extensively studied as a biomaterial [113, 114], and mechanical strength and biocompatibility have been

It has been reported that high water content PVA-H has selective permeability penetrating into low molecular weight nutrients such as glucose and protein, but not substances with cell-size levels such as fibroblasts and leukocytes [114, 115]. In addition, the gel prevents

process or to host genetically modified cells to improve meniscus repair.

cal markers (e.g., controlled release).

*3.3.3. Meniscus tissue engineering*

*3.3.4. Tendon tissue engineering*

problems in the orthopedic area.

excellent.

area.

and its ability to trap cells in the material.

exactly resemble this topographic composition [112].

#### **3.3. Applications of hydrogels in tissue engineering**

Tissue engineering is an interdisciplinary field that uses the principles of engineering and life sciences that aims to improving, maintaining, and renewing the functions of tissues and organs. It was emerged by Y. C. Fung in 1985 as a technical science department by using medical and material engineering [100].

Hydrogels are frequently used in biomedical and bioengineering fields due to their controllable different properties, similarity to tissues, and their ability to form scaffolds for different tissues due to their tunable physical and mechanical properties [101, 102]. Hydrogels used in tissue engineering have low viscosity before injection and require rapid gel formation in the physiological environment of the tissue [99]. It is frequently used in bone tissue, cartilage tissue, vascular tissue, meniscus, tendon, skin, cornea, and soft tissues. Hydrogel film can be used to fabricate or repair tissue in a loss of function. In the tissue engineering studies, synthetic or natural polymers are used for hydrogel production [103–105]. In tissue studies, adhesion of cells to artificial tissues in the course of cell culture, spreading, and population is an important feature.

#### *3.3.1. Bone tissue engineering (BTE)*

The bone is a dynamic and vascularized tissue with the ability to heal naturally on injury. However, repair mechanism cannot be sufficient in large bone defects. Current approaches often impose limitations that consider autografts and various allografts. BTE is based on the use of 3D matrices that both promote cellular growth and divergence to promote bone regeneration. The hydrogel matrices may contain biological materials such as cells and growth factors [106]. One of the studies is about that to assess the effect of different natural skeletons on this coculture system, cells were encapsulated in alginate and/or collagen hydrogel scaffolds. A researcher has discovered that in addition to the cell-cell proximity between the two cell types, the natural cell-binding capabilities of hydrogels such as collagen are preferred. Researchers have discovered increased osteogenic and angiogenic potential, as evidenced by increased protein and protein expression of ALP, BMP-2, VEGF, and PECAM [107].

#### *3.3.2. Cartilage tissue engineering*

Simply, articular cartilage is a tissue composed of a single cell type (chondrocytes) embedded in an extracellular matrix (ECM). However, the structure is more complex, and, depending on depth, the three ECM contents include three depth-related layers with varying behavior of structure and chondrocyte: superficial zone, middle zone, and deep zone. In the mid-region—the largest region—the cartilaginous tissue has a low density of round chondrocytes surrounded by an ECM consisting of hard, avascular glycosaminoglycan (GAG) and type II collagen [108, 109]. Hydrogel design to improve cartilage repair has progressed in recent years. Some using hydrogels include the development of improved network cross-linking (e.g., double networks), new techniques for processing hydrogels (e.g., 3D printing), and better incorporation of biological markers (e.g., controlled release).

Hydrogel has emerged as a promising port because of its potential features in a wide range and its ability to trap cells in the material.

#### *3.3.3. Meniscus tissue engineering*

to their soft, elastic properties; prevention of protein adsorption and cell adhesion arising from the low interfacial tension between water and hydrogels; broad acceptability for individual drugs with different hydrophilicities and molecular sizes; and unique possibilities (cross-linking density and swelling) to manipulate the release of incorporated drugs. Some of these may offer an advantage for the delivery of certain delicate drugs, such as peptides and

Tissue engineering is an interdisciplinary field that uses the principles of engineering and life sciences that aims to improving, maintaining, and renewing the functions of tissues and organs. It was emerged by Y. C. Fung in 1985 as a technical science department by using medi-

Hydrogels are frequently used in biomedical and bioengineering fields due to their controllable different properties, similarity to tissues, and their ability to form scaffolds for different tissues due to their tunable physical and mechanical properties [101, 102]. Hydrogels used in tissue engineering have low viscosity before injection and require rapid gel formation in the physiological environment of the tissue [99]. It is frequently used in bone tissue, cartilage tissue, vascular tissue, meniscus, tendon, skin, cornea, and soft tissues. Hydrogel film can be used to fabricate or repair tissue in a loss of function. In the tissue engineering studies, synthetic or natural polymers are used for hydrogel production [103–105]. In tissue studies, adhesion of cells to artificial tissues in the course of cell culture, spreading, and population is

The bone is a dynamic and vascularized tissue with the ability to heal naturally on injury. However, repair mechanism cannot be sufficient in large bone defects. Current approaches often impose limitations that consider autografts and various allografts. BTE is based on the use of 3D matrices that both promote cellular growth and divergence to promote bone regeneration. The hydrogel matrices may contain biological materials such as cells and growth factors [106]. One of the studies is about that to assess the effect of different natural skeletons on this coculture system, cells were encapsulated in alginate and/or collagen hydrogel scaffolds. A researcher has discovered that in addition to the cell-cell proximity between the two cell types, the natural cell-binding capabilities of hydrogels such as collagen are preferred. Researchers have discovered increased osteogenic and angiogenic potential, as evidenced by

increased protein and protein expression of ALP, BMP-2, VEGF, and PECAM [107].

Simply, articular cartilage is a tissue composed of a single cell type (chondrocytes) embedded in an extracellular matrix (ECM). However, the structure is more complex, and, depending on depth, the three ECM contents include three depth-related layers with varying behavior of structure and chondrocyte: superficial zone, middle zone, and deep zone. In the mid-region—the

proteins [73, 80].

286 Biomaterials in Regenerative Medicine

an important feature.

*3.3.1. Bone tissue engineering (BTE)*

*3.3.2. Cartilage tissue engineering*

**3.3. Applications of hydrogels in tissue engineering**

cal and material engineering [100].

The meniscus plays an important role in maintaining homeostasis of the knee joint [110]. Meniscus lesions are tight, have poor healing ability, and can produce tibiofemoral osteoarthritis. Current options for reconstructive treatment focus primarily on the treatment of lesions in the peripheral vascularized area [111]. On the contrary, a small number of approaches can stimulate the repair of damaged, meniscal tissue in the middle, avascular part. Tissue engineering approaches are of great interest to repair or replace damaged meniscus tissue in this area.

Hydrogel-based biomaterials show a special interest for meniscus repair, as their cores contain relatively high rates of proteoglycans responsible for viscoelastic compaction properties and degree of hydration. Hydrogels that exhibit high water content and provide a specific three-dimensional (3D) microenvironment can be designed such that the meniscus tissue can exactly resemble this topographic composition [112].

Different polymers of both natural and synthetic origin have been manipulated to produce hydrogels that host the respective cell populations for meniscus regeneration and to provide platforms for meniscus tissue replacement. To date, these compounds have been used to design controlled delivery systems of bioactive molecules involved in the meniscus repair process or to host genetically modified cells to improve meniscus repair.

#### *3.3.4. Tendon tissue engineering*

Significant advances have been made in repair techniques for tendon injuries in recent years, but the treatment of finger flexor tendon injury is still one of the most difficult and important problems in the orthopedic area.

The main problem in tendon repair is adhesion between the tendon and surrounding tissue. We note that polyvinyl alcohol hydrogel (PVA-H) has a high water content with both adhesion-inhibiting function and synovial perfusion function. PVA-H has been extensively studied as a biomaterial [113, 114], and mechanical strength and biocompatibility have been excellent.

It has been reported that high water content PVA-H has selective permeability penetrating into low molecular weight nutrients such as glucose and protein, but not substances with cell-size levels such as fibroblasts and leukocytes [114, 115]. In addition, the gel prevents adhesion on the surface by preventing cell viability. Successful tendon tissue engineering requires the use of three-dimensional (3D) biomimetic scaffold molds with physical and biochemical properties of native tendon tissue. Here, the development and characterization of a new composite scaffold fabricated by co-electrospinning of poly-ε-caprolactone (PCL) and methacrylated gelatin (mGLT) are reported.

wound. In addition, hydrogel dressings, when placed on a dry surface of the wound, provide a moist wound environment that enables the rapid healing of the wound moistening of

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 289

Hydrogel wound dressings can easily be shaped and leave no residue. Since the dressing does not stick to the wound surface, it is painless to remove from the wound surface and can be easily cleaned. Compared to semipermeable membranes, they have water vapor permeability. In addition, using hydrogel dressing, the drug can be applied locally to the wound, and the diffusion of the drug into the wound can be controlled by controlling the gel cross-linking grade. Hydrogels are used in dry necrotic wounds, superficial wounds, non-exudate, nonin-

PVA has excellent and easy film-forming properties and can be mixed with synthetic and natural polymers due to its water-soluble, biodegradable, noncarcinogenic, and biocompatible character. The final properties of the mixed material depend on the properties of the materials being assembled, and the PVA properties change after mixing. Sodium alginate is one of the popular natural polymers used in wound dressing applications by combining PVA polymer

Gelatin is a protein produced by partial hydrolysis of collagen extracted from the skin (44%, usually cow and pig), bone (27%), and connective tissues or organs (28%) of animals. Gelatin has biological activities due to its natural origin, which is suitable to be used as a wound dressing material, a drug delivery carrier, and a scaffold for tissue engineering. Gelatin has the ability to form strong hydrogels and transparent films that are easily designed as insolu-

Corneal disease is a leading cause of blindness all over the world [130]. Donor corneal transplantation is preferred in corneal disease. However, they are limited in supply, especially in

For this reason, alternative treatment methods are needed to meet the demand for increased corneal replacement. Efforts have focused on the development of tissue engineering structures that use hydrogels and biopolymer materials to replace native corneas [131–135].

An ideal tissue engineering structure should imitate the native cornea closely in order to fulfill its critical functions. Ruberti and Zieske have described three design considerations that must be met for an artificial corneal formation to be functional: (i) protection, (ii) transmis-

In order to meet the abovementioned design criteria, the tissue engineering structure must have the following essential characteristics [137]: (i) to stimulate the migration and proliferation of corneal epithelial cells for the formation of a functional corneal epithelium, (ii) to resist the intraocular pressure, and thus to prevent rupture, (iii) simulating the nanoscale fibrillary order of the corneal stroma to show a high degree of transparency (>90°C) and a refractive index comparable to that of the native cornea (3.8 MPa) [138] and the refractive

as the main or additional component due to its high water swelling ability [128].

ble hydrophilic polymers for strong regeneration and tissue implantation [129].

wounds [125].

fected wounds [126, 127].

*3.3.7. Cornea tissue engineering*

sion, and (iii) refraction [136].

developing countries.

#### *3.3.5. Skin tissue engineering*

Various types of tissue engineering scaffolds have been developed and used for epidermic engineering [116]. Ideally, this scaffold should have some biologic properties (i.e., to support keratinocyte adhesion, proliferation, and differentiation) and appropriate mechanical properties [117]. The mechanical properties of the skeletons are defined as a key modulator of keratinocyte behavior with increased cell adhesion and proliferation on hard surfaces [118].

The scaffold should also be strong enough and flexible to support easy handling during surgery and to support the natural movements of the tissues [119]. In addition, these scaffolds should only ideally deteriorate after adequate recovery, and this process can take up to 8 weeks [120]. In addition, for some clinical applications, the scaffold molds need to be crosslinked quickly in place for the contour of the wound [121].

Based on these requirements, natural hydrogels are considered attractive candidates for epidermis engineering due to their unique combination of biological and physical properties, including biocompatibility, as well as on-site cross-linking capacities, mimicking extracellular matrix (ECM), adjustable mechanical, swelling, and degradation properties [122].

#### *3.3.6. Hydrogel wound dressing*

An ideal wound dressing should maintain a moist environment at the wound interface, allow gas exchange, create microorganism barrier, and remove extreme exudates. It should also be made from a readily available biomaterial that is nontoxic and nonallergic, unattached, and easily removed without trauma, requires minimal processing, has antimicrobial properties, and accelerates wound healing [123].

Hydrogels are promising materials for wound dressing and for treatment of severe burns. In these applications, the use of hydrogel dressing offers several advantages. These are light materials and contain water in significant quantities. For this reason, hydrogels imitate some of the important qualities of people's skin and texture. Hydrogels can absorb large amounts of water or biological fluid content. They also protect the three-dimensional structure. Its advantages over other wound dressings are its good mechanical properties, oxygen permeability, liquid absorption, hydration of the wound bed, shape stability, softness similar to soft peripheral tissue, and cooling effect on the wound surface. Depending on the hydration state of the tissues, the hydrogel may give or absorb water to the wound periphery [124].

Hydrogel dressings' elastic sheet, film, or gel (amorphous) can be found in form. The hydrogel in leaf form has many of the ideal wound dressing properties. The hydrogels in this form have convenient constructions and can be cut and prepared in a manner suitable for the wound. In addition, hydrogel dressings, when placed on a dry surface of the wound, provide a moist wound environment that enables the rapid healing of the wound moistening of wounds [125].

Hydrogel wound dressings can easily be shaped and leave no residue. Since the dressing does not stick to the wound surface, it is painless to remove from the wound surface and can be easily cleaned. Compared to semipermeable membranes, they have water vapor permeability. In addition, using hydrogel dressing, the drug can be applied locally to the wound, and the diffusion of the drug into the wound can be controlled by controlling the gel cross-linking grade. Hydrogels are used in dry necrotic wounds, superficial wounds, non-exudate, noninfected wounds [126, 127].

PVA has excellent and easy film-forming properties and can be mixed with synthetic and natural polymers due to its water-soluble, biodegradable, noncarcinogenic, and biocompatible character. The final properties of the mixed material depend on the properties of the materials being assembled, and the PVA properties change after mixing. Sodium alginate is one of the popular natural polymers used in wound dressing applications by combining PVA polymer as the main or additional component due to its high water swelling ability [128].

Gelatin is a protein produced by partial hydrolysis of collagen extracted from the skin (44%, usually cow and pig), bone (27%), and connective tissues or organs (28%) of animals. Gelatin has biological activities due to its natural origin, which is suitable to be used as a wound dressing material, a drug delivery carrier, and a scaffold for tissue engineering. Gelatin has the ability to form strong hydrogels and transparent films that are easily designed as insoluble hydrophilic polymers for strong regeneration and tissue implantation [129].

#### *3.3.7. Cornea tissue engineering*

adhesion on the surface by preventing cell viability. Successful tendon tissue engineering requires the use of three-dimensional (3D) biomimetic scaffold molds with physical and biochemical properties of native tendon tissue. Here, the development and characterization of a new composite scaffold fabricated by co-electrospinning of poly-ε-caprolactone (PCL) and

Various types of tissue engineering scaffolds have been developed and used for epidermic engineering [116]. Ideally, this scaffold should have some biologic properties (i.e., to support keratinocyte adhesion, proliferation, and differentiation) and appropriate mechanical properties [117]. The mechanical properties of the skeletons are defined as a key modulator of keratinocyte behavior with increased cell adhesion and proliferation on hard surfaces [118]. The scaffold should also be strong enough and flexible to support easy handling during surgery and to support the natural movements of the tissues [119]. In addition, these scaffolds should only ideally deteriorate after adequate recovery, and this process can take up to 8 weeks [120]. In addition, for some clinical applications, the scaffold molds need to be cross-

Based on these requirements, natural hydrogels are considered attractive candidates for epidermis engineering due to their unique combination of biological and physical properties, including biocompatibility, as well as on-site cross-linking capacities, mimicking extracellular

An ideal wound dressing should maintain a moist environment at the wound interface, allow gas exchange, create microorganism barrier, and remove extreme exudates. It should also be made from a readily available biomaterial that is nontoxic and nonallergic, unattached, and easily removed without trauma, requires minimal processing, has antimicrobial properties,

Hydrogels are promising materials for wound dressing and for treatment of severe burns. In these applications, the use of hydrogel dressing offers several advantages. These are light materials and contain water in significant quantities. For this reason, hydrogels imitate some of the important qualities of people's skin and texture. Hydrogels can absorb large amounts of water or biological fluid content. They also protect the three-dimensional structure. Its advantages over other wound dressings are its good mechanical properties, oxygen permeability, liquid absorption, hydration of the wound bed, shape stability, softness similar to soft peripheral tissue, and cooling effect on the wound surface. Depending on the hydration state

of the tissues, the hydrogel may give or absorb water to the wound periphery [124].

Hydrogel dressings' elastic sheet, film, or gel (amorphous) can be found in form. The hydrogel in leaf form has many of the ideal wound dressing properties. The hydrogels in this form have convenient constructions and can be cut and prepared in a manner suitable for the

matrix (ECM), adjustable mechanical, swelling, and degradation properties [122].

methacrylated gelatin (mGLT) are reported.

linked quickly in place for the contour of the wound [121].

*3.3.5. Skin tissue engineering*

288 Biomaterials in Regenerative Medicine

*3.3.6. Hydrogel wound dressing*

and accelerates wound healing [123].

Corneal disease is a leading cause of blindness all over the world [130]. Donor corneal transplantation is preferred in corneal disease. However, they are limited in supply, especially in developing countries.

For this reason, alternative treatment methods are needed to meet the demand for increased corneal replacement. Efforts have focused on the development of tissue engineering structures that use hydrogels and biopolymer materials to replace native corneas [131–135].

An ideal tissue engineering structure should imitate the native cornea closely in order to fulfill its critical functions. Ruberti and Zieske have described three design considerations that must be met for an artificial corneal formation to be functional: (i) protection, (ii) transmission, and (iii) refraction [136].

In order to meet the abovementioned design criteria, the tissue engineering structure must have the following essential characteristics [137]: (i) to stimulate the migration and proliferation of corneal epithelial cells for the formation of a functional corneal epithelium, (ii) to resist the intraocular pressure, and thus to prevent rupture, (iii) simulating the nanoscale fibrillary order of the corneal stroma to show a high degree of transparency (>90°C) and a refractive index comparable to that of the native cornea (3.8 MPa) [138] and the refractive index, (iv) transparency to support the maintenance of functional keratocytes to preserve and de novo collagen tissue synthesis, and (v) to compensate for and thereby prevent the swelling rate of the native cornea to preserve and distribute the water content [139].

Sterilization of high water content lenses is not possible by heat and enzymatic way for a long time. They have high protein accumulation and oxygen permeability, and they easily lose

Hydrogels in Regenerative Medicine http://dx.doi.org/10.5772/intechopen.70409 291

The materials of soft contact lenses that contain polymers are not natural. They can be split into two main groups. One of these groups is hydrogel materials: hydrogel materials and silicone. The second most important material used in soft contact lenses is silicone. Elastic silicone is soft in the elastomer form. In this form, soft silicone has high oxygen permeability but also has low surface wettability. For this reason, it is coated with hydrophilic materials. The most important development in contact lenses in the recent years has been on the combination of silicone and hydrogel materials. The high oxygen permeability of silicone when added to other materials lowers its disadvantages due to the high water maintenance of hydrogels [145]. Silicone hydrogel lenses contain less water than traditional hydrogel lenses. Transparent silicone hydrogel lenses with high oxygen permeability cause less dryness in the eye and are much more comfortable to use throughout the day. With the development of silicone hydrogel lenses, the contact lens industry has experienced a revolution. With these new "super permeable" transparent contact lenses, much more oxygen can reach the cornea layer. A high oxygen supply is vital for the maintenance of eye health. This development has lead to the production of "long lasting" contact lenses, which can be used 24/7. The silicone hydrogel material is so effective that it is used to make more than 50% of contact lenses that are manufactured. Compared to traditional and hydrogel contact lenses, silicone hydrogel contact lenses have been found to have about 40 times less risk of producing infective keratitis. The fact that silicone hydrogel contact lenses are adapted to the cornea physiology, have high oxygen permeability, and have advanced surface and design features should not overshadow

that they too have mechanical and inflammatory effects on the cornea [146].

Hydrogels have a very important role in the field of biomedicine and nanotechnology. The future success of hydrogels is based on the synthesis of new polymers or on the modification of natural polymers to solve certain biological and medical difficulties. Most scientific research shows that the studies of hydrogels have a bright future. New approaches to hydrogel design are increased in the investigation of these biomaterials. Fast response, self-assembly, high and good mechanical properties, and super-porous hydrogel are just a few examples of biomateri-

Yasemin Budama-Kilinc\*, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and

Yildiz Technical University, Faculty of Chemical and Metallurgical Engineering, Department

\*Address all correspondence to: yaseminbudama@gmail.com

of Bioengineering, Esenler, Istanbul, Turkey

**3.4. Future aspect**

**Author details**

Eslin Üstün

als with an intelligent future.

water; thus, their tensile strength durability is low [144].

#### *3.3.8. Hydrogel prosthesis*

In general, medical bionic implants include innumerable solutions that remove injuries and post-resection losses, as well as various dysfunctions resulting from the operative treatment of tumors or inflammatory processes [140]. Swelling ability of hydrogel is used for various body regions in prostheses. A simple example of this is the design of an endoprosthesis which is not removed from the site for use in the treatment of patients with biliary obstruction. Here, the hydrogel ring located in the grooves around the endoprosthesis is placed on both sides of the occlusion. The swelling of the rings prevents the device from coming out of its place, thus protecting the internal drainage [141].

Carboxymethyl cellulose (CMC) hydrogel is a biodegradable, nontoxic, and non-mutagenic viscoelastic gel. It has been in clinical use since 1984 and has been used as a monobloc breast prosthesis since 1994. In the study by Brunner et al., 122 patients with hydrogel implants were evaluated between February 2000 and February 2005. It has been found that CMC hydrogel implants have higher radiotranslucency than silicone gel and that the integrity of this material is easy to prove by clinical examination. In the event of a rupture, the implant can be replaced immediately. It can be easily attached and placed by means of a small cut due to its high elastic shell. Studies conducted in patients with CMC hydrogel-filled breast prostheses have determined that a negligible complication rate and a high satisfaction rate in patients are seen. Another advantage of the CMC hydrogel-filled implant is that it is a natural feel comparable to that of silicone gel-filled breast implants [142].

Another usage area of hydrogels is contact lenses. The use of contact lenses to treat visual impairment began in the nineteenth century. The materials used in contact lenses for visual impairment must be pure. They must possess the necessary physical and mechanical properties such as malleability, transparency, wettability, mechanical stability, and oxygen permeability. They should be able to be sterilized and must be noncarcinogenic.

Contact lenses are generally classified as "hard" or "soft" depending on their elasticity. Although hard contact lenses are more durable, users do not prefer them, since they require an adaptation period. Hard contact lenses are resistant to hydrophobic substances such as poly(methyl methacrylate) (PMMA) or poly(hexafluoroisopropyl methacrylate) (HFIM). Despite this, soft contact lenses are resistant to hydrogels [143].

FDA-approved hydrogel contact lenses are divided into four groups depending on their water contents and the reactive groups on their polymer surface. These four groups are low water content, high water content, ionize (reactive chemical groups on the polymer surface), nonionize (those who do not carry reactive chemical group on the polymer surface). Low water content lenses should be compatible with all lens care systems. Contact lenses have some properties which are suitable for implantation. These properties are about protein accumulation, tensile strength, and color changing. Protein accumulation and color changing should be low, and lens strength has to be durable.

Sterilization of high water content lenses is not possible by heat and enzymatic way for a long time. They have high protein accumulation and oxygen permeability, and they easily lose water; thus, their tensile strength durability is low [144].

The materials of soft contact lenses that contain polymers are not natural. They can be split into two main groups. One of these groups is hydrogel materials: hydrogel materials and silicone. The second most important material used in soft contact lenses is silicone. Elastic silicone is soft in the elastomer form. In this form, soft silicone has high oxygen permeability but also has low surface wettability. For this reason, it is coated with hydrophilic materials. The most important development in contact lenses in the recent years has been on the combination of silicone and hydrogel materials. The high oxygen permeability of silicone when added to other materials lowers its disadvantages due to the high water maintenance of hydrogels [145].

Silicone hydrogel lenses contain less water than traditional hydrogel lenses. Transparent silicone hydrogel lenses with high oxygen permeability cause less dryness in the eye and are much more comfortable to use throughout the day. With the development of silicone hydrogel lenses, the contact lens industry has experienced a revolution. With these new "super permeable" transparent contact lenses, much more oxygen can reach the cornea layer. A high oxygen supply is vital for the maintenance of eye health. This development has lead to the production of "long lasting" contact lenses, which can be used 24/7. The silicone hydrogel material is so effective that it is used to make more than 50% of contact lenses that are manufactured. Compared to traditional and hydrogel contact lenses, silicone hydrogel contact lenses have been found to have about 40 times less risk of producing infective keratitis. The fact that silicone hydrogel contact lenses are adapted to the cornea physiology, have high oxygen permeability, and have advanced surface and design features should not overshadow that they too have mechanical and inflammatory effects on the cornea [146].

#### **3.4. Future aspect**

index, (iv) transparency to support the maintenance of functional keratocytes to preserve and de novo collagen tissue synthesis, and (v) to compensate for and thereby prevent the

In general, medical bionic implants include innumerable solutions that remove injuries and post-resection losses, as well as various dysfunctions resulting from the operative treatment of tumors or inflammatory processes [140]. Swelling ability of hydrogel is used for various body regions in prostheses. A simple example of this is the design of an endoprosthesis which is not removed from the site for use in the treatment of patients with biliary obstruction. Here, the hydrogel ring located in the grooves around the endoprosthesis is placed on both sides of the occlusion. The swelling of the rings prevents the device from coming out of its place, thus

Carboxymethyl cellulose (CMC) hydrogel is a biodegradable, nontoxic, and non-mutagenic viscoelastic gel. It has been in clinical use since 1984 and has been used as a monobloc breast prosthesis since 1994. In the study by Brunner et al., 122 patients with hydrogel implants were evaluated between February 2000 and February 2005. It has been found that CMC hydrogel implants have higher radiotranslucency than silicone gel and that the integrity of this material is easy to prove by clinical examination. In the event of a rupture, the implant can be replaced immediately. It can be easily attached and placed by means of a small cut due to its high elastic shell. Studies conducted in patients with CMC hydrogel-filled breast prostheses have determined that a negligible complication rate and a high satisfaction rate in patients are seen. Another advantage of the CMC hydrogel-filled implant is that it is a natural feel comparable

Another usage area of hydrogels is contact lenses. The use of contact lenses to treat visual impairment began in the nineteenth century. The materials used in contact lenses for visual impairment must be pure. They must possess the necessary physical and mechanical properties such as malleability, transparency, wettability, mechanical stability, and oxygen perme-

Contact lenses are generally classified as "hard" or "soft" depending on their elasticity. Although hard contact lenses are more durable, users do not prefer them, since they require an adaptation period. Hard contact lenses are resistant to hydrophobic substances such as poly(methyl methacrylate) (PMMA) or poly(hexafluoroisopropyl methacrylate) (HFIM). Despite this, soft con-

FDA-approved hydrogel contact lenses are divided into four groups depending on their water contents and the reactive groups on their polymer surface. These four groups are low water content, high water content, ionize (reactive chemical groups on the polymer surface), nonionize (those who do not carry reactive chemical group on the polymer surface). Low water content lenses should be compatible with all lens care systems. Contact lenses have some properties which are suitable for implantation. These properties are about protein accumulation, tensile strength, and color changing. Protein accumulation and color changing should be

ability. They should be able to be sterilized and must be noncarcinogenic.

swelling rate of the native cornea to preserve and distribute the water content [139].

*3.3.8. Hydrogel prosthesis*

290 Biomaterials in Regenerative Medicine

protecting the internal drainage [141].

to that of silicone gel-filled breast implants [142].

tact lenses are resistant to hydrogels [143].

low, and lens strength has to be durable.

Hydrogels have a very important role in the field of biomedicine and nanotechnology. The future success of hydrogels is based on the synthesis of new polymers or on the modification of natural polymers to solve certain biological and medical difficulties. Most scientific research shows that the studies of hydrogels have a bright future. New approaches to hydrogel design are increased in the investigation of these biomaterials. Fast response, self-assembly, high and good mechanical properties, and super-porous hydrogel are just a few examples of biomaterials with an intelligent future.

## **Author details**

Yasemin Budama-Kilinc\*, Rabia Cakir-Koc, Bahar Aslan, Burcu Özkan, Hande Mutlu and Eslin Üstün

\*Address all correspondence to: yaseminbudama@gmail.com

Yildiz Technical University, Faculty of Chemical and Metallurgical Engineering, Department of Bioengineering, Esenler, Istanbul, Turkey

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**Chapter 13**

**Provisional chapter**

**Natural Rubber Latex Biomaterials in Bone**

**Natural Rubber Latex Biomaterials in Bone** 

DOI: 10.5772/intechopen.69855

Natural rubber latex (NRL) is a white and milky solution that exudes from *Hevea brasiliensis* bark when perforated, and it has been appointed as a new promising biomaterial. NRL has been proven to be a very biocompatible material, and several new biomedical applications have been proposed. NRL has been proven to stimulate angiogenesis, cellular adhesion and formation of extracellular matrix, besides promoting replacement and regeneration of tissue. NRL also can be used as an occlusive membrane for guided bone regeneration (GBR) with promising results. Therefore, the aim of this chapter is to review NRL studies and to present NRL membrane as a promising biomembrane for use in bone trauma and injury.

**Keywords:** natural rubber latex, biomembranes, biomaterial, bone regeneration

*Hevea brasiliensis*, popularly known as rubber tree, is a plant species that belongs to the Euphorbiaceae family. This plant synthesizes latex by a system of laticiferous rings, organized as paracirculatory vessel systems, in the inner bark of the plant. Latex is the cytoplasm of the laticiferous cells, and its composition resembles the composition of common cells, except for having 30–45% of natural rubber [1]. The latex of *H. brasiliensis* exudes from the bark when it

Most of the harvested latex is coagulated for the manufacture of "dry rubber" products, including automotive tires. The latex of *H. brasiliensis* can be stabilized in an uncoagulated form with the use of ammonia, which allows the latex to be used for the manufacture of other

> © 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

**Regenerative Medicine**

**Regenerative Medicine**

http://dx.doi.org/10.5772/intechopen.69855

Dalita G.S.M. Cavalcante and Aldo Eloizo Job

Additional information is available at the end of the chapter

Leandra E. Kerche-Silva, Dalita G.S.M. Cavalcante

Additional information is available at the end of the chapter

Leandra E. Kerche-Silva,

and Aldo Eloizo Job

**Abstract**

**1. Introduction**

is perforated (**Figure 1**).

products, such as surgical gloves (**Figure 2**) [2].

## **Chapter 13**

**Provisional chapter**

## **Natural Rubber Latex Biomaterials in Bone Regenerative Medicine Regenerative Medicine**

**Natural Rubber Latex Biomaterials in Bone** 

DOI: 10.5772/intechopen.69855

Leandra E. Kerche-Silva, Dalita G.S.M. Cavalcante and Aldo Eloizo Job and Aldo Eloizo Job Additional information is available at the end of the chapter

Leandra E. Kerche-Silva, Dalita G.S.M. Cavalcante

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69855

**Abstract**

Natural rubber latex (NRL) is a white and milky solution that exudes from *Hevea brasiliensis* bark when perforated, and it has been appointed as a new promising biomaterial. NRL has been proven to be a very biocompatible material, and several new biomedical applications have been proposed. NRL has been proven to stimulate angiogenesis, cellular adhesion and formation of extracellular matrix, besides promoting replacement and regeneration of tissue. NRL also can be used as an occlusive membrane for guided bone regeneration (GBR) with promising results. Therefore, the aim of this chapter is to review NRL studies and to present NRL membrane as a promising biomembrane for use in bone trauma and injury.

**Keywords:** natural rubber latex, biomembranes, biomaterial, bone regeneration

#### **1. Introduction**

*Hevea brasiliensis*, popularly known as rubber tree, is a plant species that belongs to the Euphorbiaceae family. This plant synthesizes latex by a system of laticiferous rings, organized as paracirculatory vessel systems, in the inner bark of the plant. Latex is the cytoplasm of the laticiferous cells, and its composition resembles the composition of common cells, except for having 30–45% of natural rubber [1]. The latex of *H. brasiliensis* exudes from the bark when it is perforated (**Figure 1**).

Most of the harvested latex is coagulated for the manufacture of "dry rubber" products, including automotive tires. The latex of *H. brasiliensis* can be stabilized in an uncoagulated form with the use of ammonia, which allows the latex to be used for the manufacture of other products, such as surgical gloves (**Figure 2**) [2].

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

**Figure 1.** Latex of *Hevea brasiliensis* on exuding of bark after drilling.

**Figure 3.** Rubber particles surrounded by a layer of protein-phospholipid (adapted from Ref. [3]).

In the last years, researchers have been publishing about NRL membranes (**Figure 4**) that has been proven to be an important inductor of wound healing, inductor of esophagic wall regeneration and tympanic membrane regeneration, by mechanisms involved with angiogenesis [7–9]. NRL has also been studied for reconstructing temporal muscle fascia [9] and as arterial prosthesis in animal models, healing of ocular conjunctiva and neoangiogenesis in rabbits [10]. Besides forming biomembranes, which represent a complex colloid made of rubber particles and lutoids bodies suspended in a protein-rich media, NRL can be centrifuged in high speed and separated in fractions that mainly consist of a superior phase of rubber particles, an aqueous phase called centrifuged serum (C-serum) and an inferior phase called bottom serum

Natural Rubber Latex Biomaterials in Bone Regenerative Medicine

http://dx.doi.org/10.5772/intechopen.69855

305

The rubber particles represent 25–45% of the fresh NRL and they have a medium diameter of 50 Å to 3 μm [15, 16]. The main proteins found in the rubber particles surface are Hev b 1 and Hev b 3 [17]. Centrifuged serum is the solution composed by carbohydrates, electrolytes, proteins and amino acids. This solution is composed by the cytoplasm of the laticiferous cells and contains a great amount of proteins related to the cell metabolism [18]. This phase is implicated with most of the biological properties of NRL [19]. Bottom serum is mainly constituted of lutoids, spherical

**Figure 4.** NRL membrane.

(B-serum) (**Figure 5**) [12–14].

**Figure 2.** Products made by the latex of *Hevea brasiliensis*.

Natural rubber latex (NRL) from *H. brasiliensis* is a colloidal anionic system formed by rubber particles (1,4-*cis*-polyisoprene) stabilized by phospholipids and proteins molecules (**Figure 3**) [4, 5]. One-third of the weight of *H. brasiliensis* latex is made of natural rubber, but 1–2% of its weight consists of hundreds of proteins [5]. Other constituents such as lipids, Quebrachitol, ribonucleic acids and organic salts are also present [6].

In the last years, researchers have been publishing about NRL membranes (**Figure 4**) that has been proven to be an important inductor of wound healing, inductor of esophagic wall regeneration and tympanic membrane regeneration, by mechanisms involved with angiogenesis [7–9]. NRL has also been studied for reconstructing temporal muscle fascia [9] and as arterial prosthesis in animal models, healing of ocular conjunctiva and neoangiogenesis in rabbits [10].

Besides forming biomembranes, which represent a complex colloid made of rubber particles and lutoids bodies suspended in a protein-rich media, NRL can be centrifuged in high speed and separated in fractions that mainly consist of a superior phase of rubber particles, an aqueous phase called centrifuged serum (C-serum) and an inferior phase called bottom serum (B-serum) (**Figure 5**) [12–14].

The rubber particles represent 25–45% of the fresh NRL and they have a medium diameter of 50 Å to 3 μm [15, 16]. The main proteins found in the rubber particles surface are Hev b 1 and Hev b 3 [17]. Centrifuged serum is the solution composed by carbohydrates, electrolytes, proteins and amino acids. This solution is composed by the cytoplasm of the laticiferous cells and contains a great amount of proteins related to the cell metabolism [18]. This phase is implicated with most of the biological properties of NRL [19]. Bottom serum is mainly constituted of lutoids, spherical

**Figure 3.** Rubber particles surrounded by a layer of protein-phospholipid (adapted from Ref. [3]).

**Figure 4.** NRL membrane.

Natural rubber latex (NRL) from *H. brasiliensis* is a colloidal anionic system formed by rubber particles (1,4-*cis*-polyisoprene) stabilized by phospholipids and proteins molecules (**Figure 3**) [4, 5]. One-third of the weight of *H. brasiliensis* latex is made of natural rubber, but 1–2% of its weight consists of hundreds of proteins [5]. Other constituents such as lipids, Quebrachitol,

ribonucleic acids and organic salts are also present [6].

**Figure 2.** Products made by the latex of *Hevea brasiliensis*.

**Figure 1.** Latex of *Hevea brasiliensis* on exuding of bark after drilling.

304 Biomaterials in Regenerative Medicine

Bones protect vital organs and provide storage for calcium and phosphate. Bone compartments designated for mechanical functions are called cortical bone, and the bone compartments designed for metabolic functions are called trabecular bone. Bones can also be distinguished by their formation. When the formation occurs in a direct way, it is called intramembranous ossification and it is characterized by the condensation of mesenchymal stem cells that become osteoblasts. This type of process occurs in the flat bones of neuro- and viscerocranium and in part of the clavicle [22]. When bone formation occurs in an indirect way, it is called endochondral ossification, and it is characterized by the differentiation of mesenchymal stem cells into cartilage first and this cartilage is later replaced by bone [23]. Endochondral ossification occurs in long bones, in vertebrae and in the skull base and the posterior part of the skull [22].

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Bone modeling occurs during the growth process, and bone remodeling occurs during lifetime. These two processes take place under the control of various substances such as parathyroid hormone (PTH), calcitonin, vitamin D, growth hormone (GH), steroids, soluble cytokines and growth factors (i.e., macrophage colony-stimulating factor (M-CSF), receptor activator of nuclear κB ligand (RANKL), vascular endothelial growth factor (VEGF) and interleukin-6 (IL-6) family) [11]. Microfractures or factors related to bone microenvironment generate different stimuli that induce osteoblast to produce RANKL, which interacts with its receptor expressed by osteoclasts. This interaction activates the polarization of the osteoclasts that secrets enzymes required for bone resorption. Osteoblasts synthesize type I collagen (that represents 90% of the proteins in bone matrix) and procollagen I N-terminal peptide (PINP) that is considered a

Type I collagen together with other fibrillar collagens, bone proteins (osteopontin, bone sialoprotein and osteocalcin), proteoglycans, fibronectin and glycosaminoglycans compose unmineralized osteoid. The development of the osteoblast along with the stimulation of the osteogenic genes and mineral deposition are dependent on the pigment epithelium-derived factor (PEDF) [25]. Osteoblast phosphatases are responsible for the mineralization process, since they release

osteoblasts that remain trapped inside the mineralized matrix acquire a stellar shape after morphologic change and form a network that produces signaling through bone tissue. These

During osteogenesis, bone morphogenetics proteins (BMPs) and WNT signaling pathways are very important. The BMP pathways activate intracellular proteins called SMAD that control the expression of the gene RUNX2 (runt-related transcriptional factor 2), and this gene codifies a transcriptional factor that stimulates the mesenchymal stem cells to differentiate into osteogenic lineage [27]. WNT pathway is formed by proteins that are involved in many other

Depending on the level of differentiation of progenitor cells, WNT classic pathway induces or inhibits osteoblast formation. It also controls bone resorption when increasing the ratio of osteoprotegerin (OPG)/RANKL proteins [28]. These proteins are specifically produced by osteoblast

) 6 (OH)2

] [26]. The

phosphates that along with calcium form hydroxyapatite crystals [Ca10(PO4

biological processes, and it also regulates the expression of RUNX2 gene [28].

**2.1. Bone formation and bone remodeling molecular pathways**

marker of bone formation [24].

cells are called osteocytes.

**Figure 5.** NRL fractions after high speed centrifugation: (A) rubber particles, (B) centrifuged serum and (C) bottom serum.

vacuoles that are osmotically sensitive and induce the latex flow to stop [20, 21]. Hevein is the main protein found in bottom serum, and it is implicated with allergenic reactions [18].

Since many biological and biomedical properties can be implicated with NRL from *H. brasiliensis*, the aim of this chapter is to review and explore the studies showing the applications of this biomaterial in bone regenerative medicine.

## **2. Bone remodeling**

Bone is an organ capable of replacing old and disrupted tissue through a remodeling process. Mechanical changes required by skeletal functions make the remodeling process indispensable for the bones. The cells responsible for this process are osteoblasts and osteoclasts, and the first promote bone formation and the latter bone resorption. Osteoblasts are derived of the osteogenic differentiation of mesenchymal stem cells. Other important skeletal cells that have their origin in mature osteoblasts are the osteocytes, and these cells in particular are surrounded by extracellular matrix and have the ability to regulate osteoblast and osteoclast activities to maintain bone homeostasis [11].

Bones protect vital organs and provide storage for calcium and phosphate. Bone compartments designated for mechanical functions are called cortical bone, and the bone compartments designed for metabolic functions are called trabecular bone. Bones can also be distinguished by their formation. When the formation occurs in a direct way, it is called intramembranous ossification and it is characterized by the condensation of mesenchymal stem cells that become osteoblasts. This type of process occurs in the flat bones of neuro- and viscerocranium and in part of the clavicle [22]. When bone formation occurs in an indirect way, it is called endochondral ossification, and it is characterized by the differentiation of mesenchymal stem cells into cartilage first and this cartilage is later replaced by bone [23]. Endochondral ossification occurs in long bones, in vertebrae and in the skull base and the posterior part of the skull [22].

Bone modeling occurs during the growth process, and bone remodeling occurs during lifetime. These two processes take place under the control of various substances such as parathyroid hormone (PTH), calcitonin, vitamin D, growth hormone (GH), steroids, soluble cytokines and growth factors (i.e., macrophage colony-stimulating factor (M-CSF), receptor activator of nuclear κB ligand (RANKL), vascular endothelial growth factor (VEGF) and interleukin-6 (IL-6) family) [11].

Microfractures or factors related to bone microenvironment generate different stimuli that induce osteoblast to produce RANKL, which interacts with its receptor expressed by osteoclasts. This interaction activates the polarization of the osteoclasts that secrets enzymes required for bone resorption. Osteoblasts synthesize type I collagen (that represents 90% of the proteins in bone matrix) and procollagen I N-terminal peptide (PINP) that is considered a marker of bone formation [24].

Type I collagen together with other fibrillar collagens, bone proteins (osteopontin, bone sialoprotein and osteocalcin), proteoglycans, fibronectin and glycosaminoglycans compose unmineralized osteoid. The development of the osteoblast along with the stimulation of the osteogenic genes and mineral deposition are dependent on the pigment epithelium-derived factor (PEDF) [25]. Osteoblast phosphatases are responsible for the mineralization process, since they release phosphates that along with calcium form hydroxyapatite crystals [Ca10(PO4 ) 6 (OH)2 ] [26]. The osteoblasts that remain trapped inside the mineralized matrix acquire a stellar shape after morphologic change and form a network that produces signaling through bone tissue. These cells are called osteocytes.

#### **2.1. Bone formation and bone remodeling molecular pathways**

vacuoles that are osmotically sensitive and induce the latex flow to stop [20, 21]. Hevein is the

**Figure 5.** NRL fractions after high speed centrifugation: (A) rubber particles, (B) centrifuged serum and (C) bottom serum.

Since many biological and biomedical properties can be implicated with NRL from *H. brasiliensis*, the aim of this chapter is to review and explore the studies showing the applications of

Bone is an organ capable of replacing old and disrupted tissue through a remodeling process. Mechanical changes required by skeletal functions make the remodeling process indispensable for the bones. The cells responsible for this process are osteoblasts and osteoclasts, and the first promote bone formation and the latter bone resorption. Osteoblasts are derived of the osteogenic differentiation of mesenchymal stem cells. Other important skeletal cells that have their origin in mature osteoblasts are the osteocytes, and these cells in particular are surrounded by extracellular matrix and have the ability to regulate osteoblast and osteoclast

main protein found in bottom serum, and it is implicated with allergenic reactions [18].

this biomaterial in bone regenerative medicine.

activities to maintain bone homeostasis [11].

**2. Bone remodeling**

306 Biomaterials in Regenerative Medicine

During osteogenesis, bone morphogenetics proteins (BMPs) and WNT signaling pathways are very important. The BMP pathways activate intracellular proteins called SMAD that control the expression of the gene RUNX2 (runt-related transcriptional factor 2), and this gene codifies a transcriptional factor that stimulates the mesenchymal stem cells to differentiate into osteogenic lineage [27]. WNT pathway is formed by proteins that are involved in many other biological processes, and it also regulates the expression of RUNX2 gene [28].

Depending on the level of differentiation of progenitor cells, WNT classic pathway induces or inhibits osteoblast formation. It also controls bone resorption when increasing the ratio of osteoprotegerin (OPG)/RANKL proteins [28]. These proteins are specifically produced by osteoblast to either inhibit (OPG) or enhance (RANKL) osteoclasts activity [29]. **Figure 6** schematically represents the complex network of molecular signaling pathways during osteogenesis.

by macrophages, inflammatory cells and mesenchymal cells [35]. TNF-α stimulates osteoclast cells and promotes mesenchymal stem cells to start endochondral bone formation. TNF-α also stimulates apoptosis in hypertrophic chondrocytes, and a delay in the reabsorption of mineralized cartilage stops bone formation. When TNF-α is overproduced, as in diabetics, the cartilage is removed in a premature way, delaying bone formation and healing [36]. Right after the injury, during osteoclastogenesis, RANKL and OPG expressions are elevated. In

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During bone remodeling, IL-1 and IL-6 are the most important IL. IL-1 is produced by macrophages and promotes the formation of the primary cartilaginous callus and also promotes angiogenesis in the site [38]. IL-1 also stimulates the production of IL-6 by the osteoblasts, and IL-6 also stimulates angiogenesis by the production of VEGF. It induces the differentiation of

In bone remodeling, mesenchymal, osteoblasts and chondrocytes cells produce BMPs that can work independently or in collaboration with each other and with other members of the TGF-β family to start osteoclastogenesis [40]. BMPs are structurally and functionally related, but they exhibit different patterns of expression during bone remodeling. Murine studies have shown that BMP-2 expression is more elevated in the first 24 h after the fracture, suggesting its primary function in the beginning of the bone repair. This BMP is also related to the main-

BMP-3, BMP-4, BMP-7 and BMP-8 are expressed during the healing time (14–21 days after the injury) when there is reabsorption of the calcified cartilage and osteogenesis [43]. It has been proposed that BMP-7 is the most potent inductor of differentiation in mesenchymal stem cells

Bone is a connective tissue that possesses high vascularization. During bone endochondral and intramembranous ossification, regeneration and remodeling, the vasculature plays an important role [45]. Ten to fifteen percent of total cardiac output goes to skeletal system [46]. The role of the blood vessels in the bone is not only to supply the bones with oxygen and nutrients but also to provide them with growth factor, hormones and neurotransmitters (e.g., brain-derived serotonin), maintaining the bone activity and cell

Osteogenesis can occur through endochondral or intramembranous ossification; in both ways, angiogenesis is a critical stage and it is associated with the production of VEGF by hypertrophic chondrocytes or mesenchymal stem cells, respectively [49]. VEGF attracts endothelial cells acting as a chemotactic molecule that controls differentiation and function of osteoblasts and osteoclasts, participating in bone modeling and bone remodeling. A loss of the VEGF leads to incomplete bone vascularization and automatic disturbed endochondral ossification [50]. On the other hand, overexpression of VEGF can lead to osteosclerosis, highly increased bone formation that is a result of intense osteoblast differentiation, ending in altered bone

bone remodeling, the expression of these proteins is diminished [37].

**2.2. Vasculature involved in bone formation and bone remodeling**

osteoblasts and osteoclasts [39].

to osteoblasts [44].

survivor [47, 48].

morphology [51].

tenance of the normal bone mass [41, 42].

Besides the pathways described above, systemic hormones also regulate osteogenic commitment or differentiation of mesenchymal cells. Examples of these hormones are PTH, glucocorticoids and estrogens. Local growth factor signaling, such as bone transforming growth factor-β (TGF-β1/2), insulin-like growth factor (IGF), fibroblast growth factor 2 (FGF-2), VEGF, cytokine modulators (prostaglandins) and mitogen-activated protein kinases (MAPK), also regulates the differentiation of mesenchymal stem cells [30]. Other enzyme that plays an important role in bone regulation is Peptidyl-prolyl *cis-trans* isomerase NIMA-interacting 1 (PIN1), which interacts with RUNX2, SMAD1/5 and ß-catenin proteins [31].

Osteogenesis is also regulated by epigenetic factors such as DNA methylation, microRNAs (miRNAs), histone acetylation and deacetylation, and chromatin structure modification [32, 33]. Especially short noncoding miRNAs have been proven to affect both osteoblast (bone formation) and osteoclast (bone resorption) lineage. Some miRNAs regulate osteoblastogenesis by posttranscription regulation of RUNX2 (e.g., miR-34c, miR-133a, miR135a, miR-137, miR-205, miR-217, etc.), Osterix (OSX) (e.g., miR-31, miR-93, miR-143, miR-145, etc.) and type I collagen (e.g., miR-29, miR-Let7) proteins [34].

After a bone fracture, an inflammatory response is activated and it is crucial for the process of bone regeneration, bone remodeling and healing. This inflammatory response involves the secretion of tumor necrosis factor-α (TNF-α) and interleukins (IL) IL-1, IL-6, IL-11 and IL-18

**Figure 6.** Schematic representation of cells and regulatory molecules involved in osteoblastogenesis and bone formation (adapted from [11]).

by macrophages, inflammatory cells and mesenchymal cells [35]. TNF-α stimulates osteoclast cells and promotes mesenchymal stem cells to start endochondral bone formation. TNF-α also stimulates apoptosis in hypertrophic chondrocytes, and a delay in the reabsorption of mineralized cartilage stops bone formation. When TNF-α is overproduced, as in diabetics, the cartilage is removed in a premature way, delaying bone formation and healing [36]. Right after the injury, during osteoclastogenesis, RANKL and OPG expressions are elevated. In bone remodeling, the expression of these proteins is diminished [37].

to either inhibit (OPG) or enhance (RANKL) osteoclasts activity [29]. **Figure 6** schematically

Besides the pathways described above, systemic hormones also regulate osteogenic commitment or differentiation of mesenchymal cells. Examples of these hormones are PTH, glucocorticoids and estrogens. Local growth factor signaling, such as bone transforming growth factor-β (TGF-β1/2), insulin-like growth factor (IGF), fibroblast growth factor 2 (FGF-2), VEGF, cytokine modulators (prostaglandins) and mitogen-activated protein kinases (MAPK), also regulates the differentiation of mesenchymal stem cells [30]. Other enzyme that plays an important role in bone regulation is Peptidyl-prolyl *cis-trans* isomerase NIMA-interacting 1 (PIN1), which

Osteogenesis is also regulated by epigenetic factors such as DNA methylation, microRNAs (miRNAs), histone acetylation and deacetylation, and chromatin structure modification [32, 33]. Especially short noncoding miRNAs have been proven to affect both osteoblast (bone formation) and osteoclast (bone resorption) lineage. Some miRNAs regulate osteoblastogenesis by posttranscription regulation of RUNX2 (e.g., miR-34c, miR-133a, miR135a, miR-137, miR-205, miR-217, etc.), Osterix (OSX) (e.g., miR-31, miR-93, miR-143, miR-145, etc.) and type I collagen

After a bone fracture, an inflammatory response is activated and it is crucial for the process of bone regeneration, bone remodeling and healing. This inflammatory response involves the secretion of tumor necrosis factor-α (TNF-α) and interleukins (IL) IL-1, IL-6, IL-11 and IL-18

**Figure 6.** Schematic representation of cells and regulatory molecules involved in osteoblastogenesis and bone formation

represents the complex network of molecular signaling pathways during osteogenesis.

interacts with RUNX2, SMAD1/5 and ß-catenin proteins [31].

(e.g., miR-29, miR-Let7) proteins [34].

308 Biomaterials in Regenerative Medicine

(adapted from [11]).

During bone remodeling, IL-1 and IL-6 are the most important IL. IL-1 is produced by macrophages and promotes the formation of the primary cartilaginous callus and also promotes angiogenesis in the site [38]. IL-1 also stimulates the production of IL-6 by the osteoblasts, and IL-6 also stimulates angiogenesis by the production of VEGF. It induces the differentiation of osteoblasts and osteoclasts [39].

In bone remodeling, mesenchymal, osteoblasts and chondrocytes cells produce BMPs that can work independently or in collaboration with each other and with other members of the TGF-β family to start osteoclastogenesis [40]. BMPs are structurally and functionally related, but they exhibit different patterns of expression during bone remodeling. Murine studies have shown that BMP-2 expression is more elevated in the first 24 h after the fracture, suggesting its primary function in the beginning of the bone repair. This BMP is also related to the maintenance of the normal bone mass [41, 42].

BMP-3, BMP-4, BMP-7 and BMP-8 are expressed during the healing time (14–21 days after the injury) when there is reabsorption of the calcified cartilage and osteogenesis [43]. It has been proposed that BMP-7 is the most potent inductor of differentiation in mesenchymal stem cells to osteoblasts [44].

#### **2.2. Vasculature involved in bone formation and bone remodeling**

Bone is a connective tissue that possesses high vascularization. During bone endochondral and intramembranous ossification, regeneration and remodeling, the vasculature plays an important role [45]. Ten to fifteen percent of total cardiac output goes to skeletal system [46]. The role of the blood vessels in the bone is not only to supply the bones with oxygen and nutrients but also to provide them with growth factor, hormones and neurotransmitters (e.g., brain-derived serotonin), maintaining the bone activity and cell survivor [47, 48].

Osteogenesis can occur through endochondral or intramembranous ossification; in both ways, angiogenesis is a critical stage and it is associated with the production of VEGF by hypertrophic chondrocytes or mesenchymal stem cells, respectively [49]. VEGF attracts endothelial cells acting as a chemotactic molecule that controls differentiation and function of osteoblasts and osteoclasts, participating in bone modeling and bone remodeling. A loss of the VEGF leads to incomplete bone vascularization and automatic disturbed endochondral ossification [50]. On the other hand, overexpression of VEGF can lead to osteosclerosis, highly increased bone formation that is a result of intense osteoblast differentiation, ending in altered bone morphology [51].

In endochondral ossification, cartilage vascularization begins with the formation of a primary vessel that is projected from the perichondrial vascular network to the adjacent cartilage. Then, a capillary glomerulus is formed at the leading edge and the entire vascular unit grows. Following elongation, a backward expansion of the capillary network occurs and it tightly surrounds a pair of main vessels composed by an arteriole and a venule [52].

inflammatory reaction, and after 42 days, the NRL implants were making close contact to the

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The induction of angiogenesis is crucial for bone regeneration and remodeling, and a study using chick embryo chorioallantoic membrane (CAM) showed that NRL membranes possess angiogenic properties. The experiment also showed that the angiogenic capacity of the NRL membranes remains active and increases in temperatures ranging from 65 to 85°C. These results showed that the heating used to prepare NRL membrane does not affect its biological properties. The same study used centrifuged latex to demonstrate that the poly-isoprene is

Large fractures can mean significant reconstructive problems and sometimes require special procedures for regeneration. And this regeneration is dependent on blood clot stability, local vascularization, defect size and protection against invasion of competitive nonosteogenic tissues [58, 59]. Guided bone regeneration (GBR) was a technique developed to enhance bone repair, and in this procedure, an occlusive membrane is used to provide to the osteogenic cells

Different types of membranes have been tested for use in GBR, but resorbable membranes represent the most interesting alternative since they avoid removal surgery [61]. NRL membranes were tested as an occlusive barrier in GBR of large defects in rabbit calvaria. NRL membranes successfully enhanced bone regeneration process in the group of the treated animals, and it was shown by a statistically higher volume of mature bone in all periods of study that was up to 120 days. The NRL membranes worked as a passive barrier membrane that prevented epithelial and connective tissue migration, thus facilitating the proliferation and migration of regenerative bone cells into the wound. The NRL membranes used for GBR, as well as the ones used to treat dental alveoli defects, did not induce the formation of foreign

Another way to accelerate bone regeneration and remodeling would be incorporating BMPs to NRL membrane. As shown above, BMPs induce bone remodeling and many studies have tried to develop a BMP delivery system that could sustain gradual release of BMPs for dental and orthopedic use [63–66]. A study using bovine serum albumin (BSA) in the place of BMP (same molecular weight) and NRL membranes prepared at different polymerization temperatures showed that NRL membrane was able to release BSA for 18 days. This indicates a promising future of these membranes as active occlusive membrane in GBR and they could

Based on previous studies, a protein was isolated from NRL (P-1) and its inductive bone repair properties were compared to recombinant human bone morphogenetic protein-2 (rhBMP-2), a commercial available human protein with good osteoinductive capabilities. To compare these two proteins, a carrier made of monoolein gel was used. P-1 is still being characterized to better understand its biological properties and its function in the laticifer cells of *Hevea brasiliensis*. However, associated with monoolein gel, P-1 was able to induce new bone

not the part of latex responsible for the biological properties [4].

better conditions to perform bone remodeling [60].

release BMPs for 18 days accelerating bone healing [55].

formation on bone defect of rat calvaria [67].

body inflammatory reaction [62].

bone at many sites.

In intramembranous ossification, capillaries of a small diameter move into thin avascular layer of mesenchyme that surrounds the mesenchymal condensation center. In this center, mesenchymal cells secrete VEGF that attracts endothelial cells. At the start of mineralization of the bone, the first blood vessels associate with an extensive external network of blood vessels [53].

During bone remodeling and bone regeneration, the blood vessels play an important role. Osteoclasts form a cutting cone that moves forward resorbing dead bone or damaged/old bone matrix. A blood vessel follows this cutting cone, delivering nutrients, hormones and growth factors to osteoblast that will produce new bone [48]. OSX-expressing osteoblast precursors are involved in the process of remodeling, accompanying the blood vessels, positioning themselves in a perivascular localization, showing the tight relationship between osteogenesis and angiogenesis [54].

## **3. Bone remodeling and latex**

To clinically manage situations such as bone loss, injury or disease, researchers have been engaged for a long time to find a biocompatible material that is innocuous, promotes osseointegration, is manageable and has low cost for the people who need it. Natural rubber latex (NRL) extracted from *Hevea brasiliensis* has been used in the industrial manufacturing of several products, such as gloves, condoms, balloons and parts of medical and dental equipment [55].

Based on a new manufacturing process, several new biomedical applications have been proposed since NRL has been shown to be very biocompatible, stimulating cellular adhesion, the formation of extracellular matrix, and promoting the replacement and regeneration of tissue [8]. This new manufacturing process is based on the production of a biomembrane of NRL that has been used to replace vessels, esophagus, pericardium and abdominal wall [7]. In all of these experiments, the biomembrane promoted rapid tissue repair and elicited an inflammatory response that resembled the inflammatory response of normal healing process and not one of rejection process [8].

The membrane of NRL was also tested for the repair of bone defects in dental alveoli of rats [56], and the histological examination of the extraction sockets revealed a pattern of normal repair with characteristics similar to those reported for other materials [57]. Histometric evaluation of the areas close to the implant during the initial 7 days demonstrated progressive and accelerated osteogenesis by a decrease in the thickness of the fibrous capsule. At long-term, NRL implants did not induce the formation of foreign-body reaction nor persistent inflammatory reaction, and after 42 days, the NRL implants were making close contact to the bone at many sites.

In endochondral ossification, cartilage vascularization begins with the formation of a primary vessel that is projected from the perichondrial vascular network to the adjacent cartilage. Then, a capillary glomerulus is formed at the leading edge and the entire vascular unit grows. Following elongation, a backward expansion of the capillary network occurs and it tightly

In intramembranous ossification, capillaries of a small diameter move into thin avascular layer of mesenchyme that surrounds the mesenchymal condensation center. In this center, mesenchymal cells secrete VEGF that attracts endothelial cells. At the start of mineralization of the bone, the first blood vessels associate with an extensive external network of blood ves-

During bone remodeling and bone regeneration, the blood vessels play an important role. Osteoclasts form a cutting cone that moves forward resorbing dead bone or damaged/old bone matrix. A blood vessel follows this cutting cone, delivering nutrients, hormones and growth factors to osteoblast that will produce new bone [48]. OSX-expressing osteoblast precursors are involved in the process of remodeling, accompanying the blood vessels, positioning themselves in a perivascular localization, showing the tight relationship between

To clinically manage situations such as bone loss, injury or disease, researchers have been engaged for a long time to find a biocompatible material that is innocuous, promotes osseointegration, is manageable and has low cost for the people who need it. Natural rubber latex (NRL) extracted from *Hevea brasiliensis* has been used in the industrial manufacturing of several products, such as gloves, condoms, balloons and parts of medical and dental equipment [55]. Based on a new manufacturing process, several new biomedical applications have been proposed since NRL has been shown to be very biocompatible, stimulating cellular adhesion, the formation of extracellular matrix, and promoting the replacement and regeneration of tissue [8]. This new manufacturing process is based on the production of a biomembrane of NRL that has been used to replace vessels, esophagus, pericardium and abdominal wall [7]. In all of these experiments, the biomembrane promoted rapid tissue repair and elicited an inflammatory response that resembled the inflammatory response of normal healing process and

The membrane of NRL was also tested for the repair of bone defects in dental alveoli of rats [56], and the histological examination of the extraction sockets revealed a pattern of normal repair with characteristics similar to those reported for other materials [57]. Histometric evaluation of the areas close to the implant during the initial 7 days demonstrated progressive and accelerated osteogenesis by a decrease in the thickness of the fibrous capsule. At long-term, NRL implants did not induce the formation of foreign-body reaction nor persistent

surrounds a pair of main vessels composed by an arteriole and a venule [52].

sels [53].

310 Biomaterials in Regenerative Medicine

osteogenesis and angiogenesis [54].

**3. Bone remodeling and latex**

not one of rejection process [8].

The induction of angiogenesis is crucial for bone regeneration and remodeling, and a study using chick embryo chorioallantoic membrane (CAM) showed that NRL membranes possess angiogenic properties. The experiment also showed that the angiogenic capacity of the NRL membranes remains active and increases in temperatures ranging from 65 to 85°C. These results showed that the heating used to prepare NRL membrane does not affect its biological properties. The same study used centrifuged latex to demonstrate that the poly-isoprene is not the part of latex responsible for the biological properties [4].

Large fractures can mean significant reconstructive problems and sometimes require special procedures for regeneration. And this regeneration is dependent on blood clot stability, local vascularization, defect size and protection against invasion of competitive nonosteogenic tissues [58, 59]. Guided bone regeneration (GBR) was a technique developed to enhance bone repair, and in this procedure, an occlusive membrane is used to provide to the osteogenic cells better conditions to perform bone remodeling [60].

Different types of membranes have been tested for use in GBR, but resorbable membranes represent the most interesting alternative since they avoid removal surgery [61]. NRL membranes were tested as an occlusive barrier in GBR of large defects in rabbit calvaria. NRL membranes successfully enhanced bone regeneration process in the group of the treated animals, and it was shown by a statistically higher volume of mature bone in all periods of study that was up to 120 days. The NRL membranes worked as a passive barrier membrane that prevented epithelial and connective tissue migration, thus facilitating the proliferation and migration of regenerative bone cells into the wound. The NRL membranes used for GBR, as well as the ones used to treat dental alveoli defects, did not induce the formation of foreign body inflammatory reaction [62].

Another way to accelerate bone regeneration and remodeling would be incorporating BMPs to NRL membrane. As shown above, BMPs induce bone remodeling and many studies have tried to develop a BMP delivery system that could sustain gradual release of BMPs for dental and orthopedic use [63–66]. A study using bovine serum albumin (BSA) in the place of BMP (same molecular weight) and NRL membranes prepared at different polymerization temperatures showed that NRL membrane was able to release BSA for 18 days. This indicates a promising future of these membranes as active occlusive membrane in GBR and they could release BMPs for 18 days accelerating bone healing [55].

Based on previous studies, a protein was isolated from NRL (P-1) and its inductive bone repair properties were compared to recombinant human bone morphogenetic protein-2 (rhBMP-2), a commercial available human protein with good osteoinductive capabilities. To compare these two proteins, a carrier made of monoolein gel was used. P-1 is still being characterized to better understand its biological properties and its function in the laticifer cells of *Hevea brasiliensis*. However, associated with monoolein gel, P-1 was able to induce new bone formation on bone defect of rat calvaria [67].

This protein extracted from NRL (P-1) was used in combination with a fibrin sealant in the repair of rat tibial bone defects. This combination was successful, being immunoidentified by the presence of osteoblasts in the area, showing high osteogenic and osteoconductive capacity for bone healing [68].

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der Rubberc. 1956;**33**:195

Allergy. 2008;**38**:989-912

Moura et al. [69] used different polymerized NRL membranes in rabbit calvaria with bone defect. These membranes were compared to polytetrafluoroethylene (PTFE) membranes, which are an extensively studied material considered gold standard for occlusive membranes [70]. In this study, NRL membranes performed their role as biological barriers and achieved a similar performance to the PTFE membrane. One of the polymerized NRL membranes was ammonia free, and these were the membranes that significantly improved the bone repair process producing higher bone formation, being more effective than PTFE membrane. These results were achieved since the method of preparation of these NRL ammonia-free membranes preserved the angiogenic stimulus of the membranes. These membranes also did not lead to bone tissue hypersensitization.

NRL was also coated with calcium phosphate (Ca/P) and tested for biomedical application. Biomaterials added with Ca/P present biological, chemical and mechanical properties very similar to the mineral phase of the bone besides the ability to bond to the host tissue. A hemolytic test was performed, and this material did not affect the blood cells, being so ready for animal tests [71].

These results showed above indicate NRL membranes as a promising future biomembrane that could be used to accelerate bone healing. More experiments are being done already in humans. Since NRL membranes present intense angiogenic activity and wound healing activity, NRL membranes are being commercialized in Brazil and other 60 countries around the world as a band-aid curative (BIOCURE®) for the treatment of ulcers in diabetic patients.

## **Author details**

Leandra E. Kerche-Silva\*, Dalita G.S.M. Cavalcante and Aldo Eloizo Job

\*Address all correspondence to: leakerche@gmail.com

Department of Physics, Chemistry and Biology, Universidade Estadual Paulista "Júlio de Mesquita Filho", Presidente Prudente, São Paulo, Brazil

## **References**


[3] Nawamawat K, Sakdapipanich JT, Ho CC, Ma Y, Vancso JG. Surface nanostructure of *Hevea brasiliensis* natural rubber latex particles. Colloids and Surfaces A. 2011;**390**:157-166

This protein extracted from NRL (P-1) was used in combination with a fibrin sealant in the repair of rat tibial bone defects. This combination was successful, being immunoidentified by the presence of osteoblasts in the area, showing high osteogenic and osteoconductive capacity

Moura et al. [69] used different polymerized NRL membranes in rabbit calvaria with bone defect. These membranes were compared to polytetrafluoroethylene (PTFE) membranes, which are an extensively studied material considered gold standard for occlusive membranes [70]. In this study, NRL membranes performed their role as biological barriers and achieved a similar performance to the PTFE membrane. One of the polymerized NRL membranes was ammonia free, and these were the membranes that significantly improved the bone repair process producing higher bone formation, being more effective than PTFE membrane. These results were achieved since the method of preparation of these NRL ammonia-free membranes preserved the angiogenic stimulus of the membranes. These membranes also did not lead to bone tissue hypersensitization. NRL was also coated with calcium phosphate (Ca/P) and tested for biomedical application. Biomaterials added with Ca/P present biological, chemical and mechanical properties very similar to the mineral phase of the bone besides the ability to bond to the host tissue. A hemolytic test was performed, and this material did not affect the blood cells, being so ready for

These results showed above indicate NRL membranes as a promising future biomembrane that could be used to accelerate bone healing. More experiments are being done already in humans. Since NRL membranes present intense angiogenic activity and wound healing activity, NRL membranes are being commercialized in Brazil and other 60 countries around the world as a band-aid curative (BIOCURE®) for the treatment of ulcers in diabetic patients.

Department of Physics, Chemistry and Biology, Universidade Estadual Paulista "Júlio de

[1] Jacob JL, Auzac J, Prevôt JL. The composition of natural latex from *Hevea brasiliensis*.

[2] Arif SAM, Hamilton RG, Yusof F, Chew NP, Loke YH, Nimkar S. Isolation and characterization of the early nodule-specific protein homologue (Hev b 13), an allergenic lipolytic esterase from *Hevea brasiliensis* latex. Journal of Biological Chemistry. 2004;**279**:23933-23941

Leandra E. Kerche-Silva\*, Dalita G.S.M. Cavalcante and Aldo Eloizo Job

Clinical Reviews in Allergy & Immunology. 1993;**11**:325-337

\*Address all correspondence to: leakerche@gmail.com

Mesquita Filho", Presidente Prudente, São Paulo, Brazil

for bone healing [68].

312 Biomaterials in Regenerative Medicine

animal tests [71].

**Author details**

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**Chapter 14**

**Provisional chapter**

**Systematic Study of Ethylene-Vinyl Acetate (EVA) in**

**Manufacturing of Protector Devices for the Orofacial** 

DOI: 10.5772/intechopen.69969

**Systematic Study of Ethylene-Vinyl Acetate (EVA) in the** 

Fracture of facial bones and dental elements, and laceration of soft tissue, have increased in sports over recent years. Dentist is the only professional responsible for the mouth protection design, the knowledge about suitable materials is essential. EVA is a thermoplastic material, available in the market, easy of handling and processing, and low-cost. However, it is important to understand the mechanical properties and ability to absorb and to dissipate the impact energy, when this material is submitted to different environments, such as oral cavity with saliva and different temperatures. This chapter show provides a systematic evaluation of the EVA application in orofacial protectors while focusing on sports. The research comprises two aspects: experimental tests and numerical analyses. During experimental tests, EVA was analyzed in special buccal conditions, concerning temperature and presence of saliva. Regarding the presence of saliva, more specific studies about its influence on the mechanical behavior of EVA were performed. In the numerical analyses of the EVA orofacial protector, the studies focused on its effect on the nasal bone integrity, and in the zygomatic bone protection. However, life cycle should be analyzed, since its performance deteriorates over time. Mainly due to the saliva-originated changes to the EVA mechanical characteristics, it can behave as a rigid material. For facial protection, a better performance is obtained with a combination of rigid and soft EVA material. According to the experimental and numerical results from a systematic study of EVA, its application to orofacial protection can be considered satisfactory. **Keywords:** material tests, orofacial protection, trauma in sports, protection in sports

> © 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

**the Manufacturing of Protector Devices for the**

**Orofacial System**

Larissa Driemeier

**Abstract**

Gilmar Ferreira Batalha and

**System**

Reinaldo Brito e Dias, Neide Pena Coto,

Reinaldo Brito e Dias, Neide Pena Coto,

http://dx.doi.org/10.5772/intechopen.69969

Gilmar Ferreira Batalha and Larissa Driemeier

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

**Provisional chapter**

## **Systematic Study of Ethylene-Vinyl Acetate (EVA) in the Manufacturing of Protector Devices for the Orofacial System Manufacturing of Protector Devices for the Orofacial System**

DOI: 10.5772/intechopen.69969

**Systematic Study of Ethylene-Vinyl Acetate (EVA) in the** 

Reinaldo Brito e Dias, Neide Pena Coto, Gilmar Ferreira Batalha and Larissa Driemeier Larissa Driemeier

Reinaldo Brito e Dias, Neide Pena Coto,

Additional information is available at the end of the chapter Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69969

Gilmar Ferreira Batalha and

#### **Abstract**

Fracture of facial bones and dental elements, and laceration of soft tissue, have increased in sports over recent years. Dentist is the only professional responsible for the mouth protection design, the knowledge about suitable materials is essential. EVA is a thermoplastic material, available in the market, easy of handling and processing, and low-cost. However, it is important to understand the mechanical properties and ability to absorb and to dissipate the impact energy, when this material is submitted to different environments, such as oral cavity with saliva and different temperatures. This chapter show provides a systematic evaluation of the EVA application in orofacial protectors while focusing on sports. The research comprises two aspects: experimental tests and numerical analyses. During experimental tests, EVA was analyzed in special buccal conditions, concerning temperature and presence of saliva. Regarding the presence of saliva, more specific studies about its influence on the mechanical behavior of EVA were performed. In the numerical analyses of the EVA orofacial protector, the studies focused on its effect on the nasal bone integrity, and in the zygomatic bone protection. However, life cycle should be analyzed, since its performance deteriorates over time. Mainly due to the saliva-originated changes to the EVA mechanical characteristics, it can behave as a rigid material. For facial protection, a better performance is obtained with a combination of rigid and soft EVA material. According to the experimental and numerical results from a systematic study of EVA, its application to orofacial protection can be considered satisfactory.

**Keywords:** material tests, orofacial protection, trauma in sports, protection in sports

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

## **1. Introduction**

According to the World Health Organization (WHO) [1], "Health is a state of complete physical, mental and social well-being and not merely the absence of disease or infirmity."

Regarding mechanical behavior, polymers are classified as *elastomers*, *plastics* and *fibers* [4].

cost [5].

mass.

During dentists' day-to-day operations, *resin* is the most commonly used polymer. Most of these resins are based on methacrylate, with methyl methacrylate as the main ingredient. Resins are easy to manipulate, without demanding elaborate techniques; the final resin products are esthetically acceptable and offer excellent balance when used in the oral environment, in the presence of saliva and chewing conditions, besides being low

Systematic Study of Ethylene-Vinyl Acetate (EVA) in the Manufacturing of Protector Devices...

http://dx.doi.org/10.5772/intechopen.69969

321

Ethylene-vinyl acetate (EVA), object of this study, is a thermoplastic copolymer derived from petroleum. For dental use, it is in the form of rigid or flexible flat plates, in thicknesses of 1–5 mm, without the presence of blowing agents, differently from EVA plates available in the common market. It is indicated as a shock absorber material, for producing mouth and facial protectors for sports practice, as pointed out by Coto et al. [6], as well as dental bleaching trays,

The copolymer of ethylene and vinyl acetate is a thermoplastic polymer, formed by different monomers: ethylene and vinyl acetate (**Figure 1**). Monomers merge through a polymerization process—a set of reactions among simple molecules to form a macromolecule of high molar

The EVA presents semi-crystalline structure; its geometry is composed of an amorphous and a crystalline part. The damping capacity of EVA increases as the percentage of vinyl acetate decreases. As already mentioned, EVA is a macromolecule composed of repeated units linked by covalent bonds and its primary units of constitution are two monomers whose physical properties depend on their size and molecular weight. Polymeric materials generally exhibit

, temperature resistance (glass transition temperatures

orthodontic restraints and as the base for facial prostheses.

**2. Ethylene-vinyl acetate**

density ranging from 0.926 to 0.950 g/cm<sup>3</sup>

**Figure 1.** Polymerization reaction between ethylene and vinyl acetate, resulting in EVA.

Areas of study related to life, health and disease are called human health sciences. Medicine, Biology, Biomedicine, Nursing, Speech Therapy, Pharmacy and Biochemistry, Sports Science, Physical Education, Psychology, Occupational Therapy, Nutrition, Physiotherapy, Bioengineering and Dentistry are part of this program. All these research areas focus on improving or maintaining the patient quality of life, in accordance with the conditions dictated by WHO.

In dentistry, a particularly important area is related to the endless search for materials that can more efficiently help the maintenance and/or return of the individual's well-being. Researchers in dentistry seek and study materials that may replace dental organs, may be accepted in the alveolar and dentofacial complex, or may protect the orofacial complex from injuries.

Therefore, due to the technical-scientific excellence required in its attributions, dentistry is a science that requires constant updating of materials science and applications. It is worth highlighting that, to indicate a safe and efficient clinical application for a particular material, mechanical, physical, chemical and biological properties must be known.

According to Anusavice et al. [2], four groups of materials are used and studied in dentistry: metals, ceramics, composites and polymers. These materials are separated into modalities, according to their application: preventive, restorative or auxiliary materials.

*Auxiliary* are materials with recognized importance and application but which do not fit into the first two modalities. It is the best option for describing the function of polymers.

Polymers are an important category of materials for dentistry. They are versatile, since they can be combined in order to improve mechanical properties, and moreover, they are reproducible and homogeneous [3].

The term polymer derives from the Greek words: poly-many and mer-unit; or, more specifically, it is a macromolecule composed of repeating units linked by a covalent bond. Its physical properties depend on the length of the molecule and its molar mass. When the polymer is formed by a single type of *mer*, it is called **homopolymer**; otherwise, it is called a **copolymer**.

According to their malleability, polymers are classified into thermoplastic and thermosetting. When the temperature is raised above its melt point, the *thermoplastic polymer* becomes softer and more fluid, allowing it to be molded. When the heat source is removed, the thermoplastic hardens in the molded shape. Since it occurs without chemical curing, it is a reversible physical transformation. In turn, with the addition of a second material and/or heat, *thermosetting polymers* soften and cure, forming cross-links that prevent the material from returning to the primary form. This process cannot be repeated.

Regarding mechanical behavior, polymers are classified as *elastomers*, *plastics* and *fibers* [4].

During dentists' day-to-day operations, *resin* is the most commonly used polymer. Most of these resins are based on methacrylate, with methyl methacrylate as the main ingredient. Resins are easy to manipulate, without demanding elaborate techniques; the final resin products are esthetically acceptable and offer excellent balance when used in the oral environment, in the presence of saliva and chewing conditions, besides being low cost [5].

Ethylene-vinyl acetate (EVA), object of this study, is a thermoplastic copolymer derived from petroleum. For dental use, it is in the form of rigid or flexible flat plates, in thicknesses of 1–5 mm, without the presence of blowing agents, differently from EVA plates available in the common market. It is indicated as a shock absorber material, for producing mouth and facial protectors for sports practice, as pointed out by Coto et al. [6], as well as dental bleaching trays, orthodontic restraints and as the base for facial prostheses.

## **2. Ethylene-vinyl acetate**

**1. Introduction**

320 Biomaterials in Regenerative Medicine

tated by WHO.

ducible and homogeneous [3].

primary form. This process cannot be repeated.

a **copolymer**.

According to the World Health Organization (WHO) [1], "Health is a state of complete physi-

Areas of study related to life, health and disease are called human health sciences. Medicine, Biology, Biomedicine, Nursing, Speech Therapy, Pharmacy and Biochemistry, Sports Science, Physical Education, Psychology, Occupational Therapy, Nutrition, Physiotherapy, Bioengineering and Dentistry are part of this program. All these research areas focus on improving or maintaining the patient quality of life, in accordance with the conditions dic-

In dentistry, a particularly important area is related to the endless search for materials that can more efficiently help the maintenance and/or return of the individual's well-being. Researchers in dentistry seek and study materials that may replace dental organs, may be accepted in the

Therefore, due to the technical-scientific excellence required in its attributions, dentistry is a science that requires constant updating of materials science and applications. It is worth highlighting that, to indicate a safe and efficient clinical application for a particular material,

According to Anusavice et al. [2], four groups of materials are used and studied in dentistry: metals, ceramics, composites and polymers. These materials are separated into modalities,

*Auxiliary* are materials with recognized importance and application but which do not fit into

Polymers are an important category of materials for dentistry. They are versatile, since they can be combined in order to improve mechanical properties, and moreover, they are repro-

The term polymer derives from the Greek words: poly-many and mer-unit; or, more specifically, it is a macromolecule composed of repeating units linked by a covalent bond. Its physical properties depend on the length of the molecule and its molar mass. When the polymer is formed by a single type of *mer*, it is called **homopolymer**; otherwise, it is called

According to their malleability, polymers are classified into thermoplastic and thermosetting. When the temperature is raised above its melt point, the *thermoplastic polymer* becomes softer and more fluid, allowing it to be molded. When the heat source is removed, the thermoplastic hardens in the molded shape. Since it occurs without chemical curing, it is a reversible physical transformation. In turn, with the addition of a second material and/or heat, *thermosetting polymers* soften and cure, forming cross-links that prevent the material from returning to the

the first two modalities. It is the best option for describing the function of polymers.

alveolar and dentofacial complex, or may protect the orofacial complex from injuries.

mechanical, physical, chemical and biological properties must be known.

according to their application: preventive, restorative or auxiliary materials.

cal, mental and social well-being and not merely the absence of disease or infirmity."

The copolymer of ethylene and vinyl acetate is a thermoplastic polymer, formed by different monomers: ethylene and vinyl acetate (**Figure 1**). Monomers merge through a polymerization process—a set of reactions among simple molecules to form a macromolecule of high molar mass.

The EVA presents semi-crystalline structure; its geometry is composed of an amorphous and a crystalline part. The damping capacity of EVA increases as the percentage of vinyl acetate decreases. As already mentioned, EVA is a macromolecule composed of repeated units linked by covalent bonds and its primary units of constitution are two monomers whose physical properties depend on their size and molecular weight. Polymeric materials generally exhibit density ranging from 0.926 to 0.950 g/cm<sup>3</sup> , temperature resistance (glass transition temperatures

**Figure 1.** Polymerization reaction between ethylene and vinyl acetate, resulting in EVA.

close to 0 to −7°C). Among the main characteristics of EVA, its elastic behavior characterized by the Young Modulus ranging from 15 to 80 MPa can be highlighted. Flexible EVA, for example, behaves similarly to elastomers, and its elasticity is considerable.

**2.2. Hardness**

proportionally.

**2.3. Transparency**

**2.4. Damping**

**2.5. Viscoelasticity**

BR 20 023048 9).

pyrolysis.

material becomes more transparent.

The hardness of a polymer is determined by the penetration of the Durometer indenter foot into a small sample (Shore Hardness). The increase in vinyl acetate content reduces the hardness of EVA, mainly due to the decrease in its degree of crystallization. Although hardness and stiffness are different properties, in some cases, it is possible to establish an empirical correlation between them for a given family of polymers. In some cases, as the degree of crystallization of EVA increases, the stiffness and hardness increase

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The polymer crystals of EVA act as physical obstacles to the passage of light. Accordingly, as the polymer crystals concentration decreases, increasing the content of vinyl acetate, the

It is the ability of the material to absorb the mechanical energy to mainly overcome internal friction. The damping capacity of EVA increases as the vinyl acetate content reduces. Damping capacity is sometimes unduly related to hardness. However, a hard polymer can be

In many of the practical conditions in which polymers are requested or tested, their mechanical response is found to be time-dependent, which characterizes these materials as viscoelastic, as already mentioned. This absorption may occur due to the internal friction between the macromolecules, by shape changes (rotation of the carbon–carbon bonds around its own axis) or by flow. Furthermore, in case of impact, the viscous portion is responsible, for the delay in the elastic response, which will depend on the stimulus and on the time necessary to coil and

The study of the mechanical properties of EVA focused on mouth guards and facial protectors. Particularly for the facial protector designed in this study, a patent was applied (number

Several experimental tests made with the material, available in the literature [10–12], confirm this percentage, which is proportionally inverse to the EVA damping capacity. Moreover, the EVA was carefully characterized in POLITENO—Brazil (now BRASKEN) for analysis. The analyses of vinyl acetate percentage were performed by means of

designed to have the same damping capacity of a soft polymer.

**3. Experimental study of EVA applied to oral protection**

to uncoil the polymer macromolecule [4, 7–9].

In most practical situations in which EVA is applied or mechanically tested, it is possible to observe that the material's mechanical response is time dependent, that is, it is a viscoelastic material. This characteristic of viscosity is important for energy dissipation.

In the chemical industry, EVA is presented in grain form as shown in **Figure 2**.

Some mechanical properties of EVA are discussed as follows.

#### **2.1. Stiffness**

The initial stiffness of the EVA can be measured by its modulus of elasticity, that is, angle of inclination of the approximated straight line that relates stresses as a function of the strains, in elastic regime. In the elastic regime, the energy absorbed by the deformed material is totally restored by removing the stress. The higher the vinyl acetate concentration, the more flexible the EVA material is, due to the reduction in the degree of crystallization.

The degree of EVA crystallization is proportional to the latent heat of fusion (Δ*Hf* ), and its value increases as the concentration of crystals present in EVA increases. However, EVA is not a totally crystalline polymer because, in the solid state, it contains two phases: amorphous and crystalline. In fact, the presence of a glass transition temperature (*T<sup>g</sup>* ) means that it contains an amorphous phase, since *T<sup>g</sup>* is a thermal transition exclusive of the amorphous phase, that is, it is the temperature at which the macromolecules of the amorphous phase acquire rotational mobility. The amorphous phase of EVA is represented by a macromolecule entanglement which lacks an ordered and periodic three-dimensional structure. The crystalline phase, on the other hand, is characterized by a three-dimensional ordered and periodic structure of macromolecules folded one on the other, assuming the lamellar format. The melting temperature (*Tf* ) is also a thermal transition, in which the crystalline phase disintegrates and the polymer becomes a viscous liquid.

**Figure 2.** EVA in granules.

#### **2.2. Hardness**

close to 0 to −7°C). Among the main characteristics of EVA, its elastic behavior characterized by the Young Modulus ranging from 15 to 80 MPa can be highlighted. Flexible EVA, for example,

In most practical situations in which EVA is applied or mechanically tested, it is possible to observe that the material's mechanical response is time dependent, that is, it is a viscoelastic

The initial stiffness of the EVA can be measured by its modulus of elasticity, that is, angle of inclination of the approximated straight line that relates stresses as a function of the strains, in elastic regime. In the elastic regime, the energy absorbed by the deformed material is totally restored by removing the stress. The higher the vinyl acetate concentration, the more flexible

increases as the concentration of crystals present in EVA increases. However, EVA is not a totally crystalline polymer because, in the solid state, it contains two phases: amorphous and crystal-

temperature at which the macromolecules of the amorphous phase acquire rotational mobility. The amorphous phase of EVA is represented by a macromolecule entanglement which lacks an ordered and periodic three-dimensional structure. The crystalline phase, on the other hand, is characterized by a three-dimensional ordered and periodic structure of macromolecules folded

transition, in which the crystalline phase disintegrates and the polymer becomes a viscous liquid.

is a thermal transition exclusive of the amorphous phase, that is, it is the

), and its value

) is also a thermal

) means that it contains an amor-

behaves similarly to elastomers, and its elasticity is considerable.

Some mechanical properties of EVA are discussed as follows.

line. In fact, the presence of a glass transition temperature (*T<sup>g</sup>*

**2.1. Stiffness**

322 Biomaterials in Regenerative Medicine

phous phase, since *T<sup>g</sup>*

**Figure 2.** EVA in granules.

material. This characteristic of viscosity is important for energy dissipation.

the EVA material is, due to the reduction in the degree of crystallization.

The degree of EVA crystallization is proportional to the latent heat of fusion (Δ*Hf*

one on the other, assuming the lamellar format. The melting temperature (*Tf*

In the chemical industry, EVA is presented in grain form as shown in **Figure 2**.

The hardness of a polymer is determined by the penetration of the Durometer indenter foot into a small sample (Shore Hardness). The increase in vinyl acetate content reduces the hardness of EVA, mainly due to the decrease in its degree of crystallization. Although hardness and stiffness are different properties, in some cases, it is possible to establish an empirical correlation between them for a given family of polymers. In some cases, as the degree of crystallization of EVA increases, the stiffness and hardness increase proportionally.

#### **2.3. Transparency**

The polymer crystals of EVA act as physical obstacles to the passage of light. Accordingly, as the polymer crystals concentration decreases, increasing the content of vinyl acetate, the material becomes more transparent.

#### **2.4. Damping**

It is the ability of the material to absorb the mechanical energy to mainly overcome internal friction. The damping capacity of EVA increases as the vinyl acetate content reduces. Damping capacity is sometimes unduly related to hardness. However, a hard polymer can be designed to have the same damping capacity of a soft polymer.

#### **2.5. Viscoelasticity**

In many of the practical conditions in which polymers are requested or tested, their mechanical response is found to be time-dependent, which characterizes these materials as viscoelastic, as already mentioned. This absorption may occur due to the internal friction between the macromolecules, by shape changes (rotation of the carbon–carbon bonds around its own axis) or by flow. Furthermore, in case of impact, the viscous portion is responsible, for the delay in the elastic response, which will depend on the stimulus and on the time necessary to coil and to uncoil the polymer macromolecule [4, 7–9].

## **3. Experimental study of EVA applied to oral protection**

The study of the mechanical properties of EVA focused on mouth guards and facial protectors. Particularly for the facial protector designed in this study, a patent was applied (number BR 20 023048 9).

Several experimental tests made with the material, available in the literature [10–12], confirm this percentage, which is proportionally inverse to the EVA damping capacity. Moreover, the EVA was carefully characterized in POLITENO—Brazil (now BRASKEN) for analysis. The analyses of vinyl acetate percentage were performed by means of pyrolysis.

**Figure 3.** Compression test of the EVA specimens—ABNT NBR 8690—with Schob Pendulum.

In **Figure 3**, the Schob pendulum was used to measure the resilience of the EVA. Six experiments were performed, three used to calibrate the system and three to measure the property. The EVA was observed to have a great damping potential, since it absorbed 50% of the applied energy.

Experimental compression tests were performed to the mechanical characterization of the EVA.

**Figure 4** shows the Instron® machine and the recording of the compression tests, performed by a Photron Ultima APX-RS high-speed camera (3000 frames per second). The record helped the study of the nonlinear material behavior of the EVA.

Particularly, **Figure 5** shows that EVA undergoes considerable plastic deformation before failure.

The test was controlled by optical pyrometer, maintaining the temperature around 37–39°C,

**Figure 4.** (A) EVA compression test in an Instron machine. (B) Detail of the geometry of the specimen: flat discs with 30

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Five EVA mouth guards of each thickness (3 and 4 mm) were made for each test group, using models of a superior dental arch in stone gypsum and metalvander® vacuum-form machine. The geometry respected the recommendation of American Academy for Sports Dentistry [13], that is, 2 mm below the bottom of the vestibular groove, 10 mm beyond the palatine gingival

The heating time for both thicknesses was 4 min, approximately; aspiration time was 45 s

(**Figure 9**). All the protectors were immersed in cold water for 10 min.

**Figure 5.** (A) EVA specimens for tensile test. (B) Detail of the specimen after failure.

close to the mouth temperature (**Figure 8**).

mm in diameter. (C) Compression test recording.

and extension up to the second upper molars.

#### **3.1. Mechanical study of the operation of a mouth guard**

To reproduce conditions as close as possible to a real situation, models in epoxy resin were manufactured from a patient model (**Figure 6**).

As illustrated in **Figure 7**, the models of the upper and lower arches were fixed in a compression device that allows the lower arch to move while maintaining the upper arch fixed. The compression device was coupled to a Universal Kratos Test Machine, data acquisition system, 20 kN load cell. The aperture, initially in occlusion, was controlled by the extensometer, with a maximum opening of 18 mm. Compression tests were performed, at a velocity of 42.86 mm/min.

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**Figure 4.** (A) EVA compression test in an Instron machine. (B) Detail of the geometry of the specimen: flat discs with 30 mm in diameter. (C) Compression test recording.

**Figure 5.** (A) EVA specimens for tensile test. (B) Detail of the specimen after failure.

In **Figure 3**, the Schob pendulum was used to measure the resilience of the EVA. Six experiments were performed, three used to calibrate the system and three to measure the property. The EVA was observed to have a great damping potential, since it absorbed 50% of the applied energy.

**Figure 3.** Compression test of the EVA specimens—ABNT NBR 8690—with Schob Pendulum.

Experimental compression tests were performed to the mechanical characterization of the

**Figure 4** shows the Instron® machine and the recording of the compression tests, performed by a Photron Ultima APX-RS high-speed camera (3000 frames per second). The record helped

Particularly, **Figure 5** shows that EVA undergoes considerable plastic deformation before

To reproduce conditions as close as possible to a real situation, models in epoxy resin were

As illustrated in **Figure 7**, the models of the upper and lower arches were fixed in a compression device that allows the lower arch to move while maintaining the upper arch fixed. The compression device was coupled to a Universal Kratos Test Machine, data acquisition system, 20 kN load cell. The aperture, initially in occlusion, was controlled by the extensometer, with a maximum opening of 18 mm. Compression tests were performed, at a velocity of 42.86 mm/min.

the study of the nonlinear material behavior of the EVA.

**3.1. Mechanical study of the operation of a mouth guard**

manufactured from a patient model (**Figure 6**).

EVA.

324 Biomaterials in Regenerative Medicine

failure.

The test was controlled by optical pyrometer, maintaining the temperature around 37–39°C, close to the mouth temperature (**Figure 8**).

Five EVA mouth guards of each thickness (3 and 4 mm) were made for each test group, using models of a superior dental arch in stone gypsum and metalvander® vacuum-form machine. The geometry respected the recommendation of American Academy for Sports Dentistry [13], that is, 2 mm below the bottom of the vestibular groove, 10 mm beyond the palatine gingival and extension up to the second upper molars.

The heating time for both thicknesses was 4 min, approximately; aspiration time was 45 s (**Figure 9**). All the protectors were immersed in cold water for 10 min.

**Figure 6.** Model of dental arches made of epoxy resin.

The groups are divided as follows:

**A.** Mouth guard—3-mm-thick blade

**B.** Mouth guard—4-mm-thick blade

named **protB**).

• Room temperature/without saliva

• Oral temperature/without saliva

• Room temperature/saturated in saliva

**Figure 9.** Manufacture of mouth guards in vacuum form Metalvander® machine.

• Oral temperature/saturated in saliva

• Room temperature/without saliva

• Oral temperature/without saliva

• Room temperature/saturated in saliva

• Oral temperature/saturated in saliva

**Table 1** shows a variation in the compression maximum load in Newtons (N) when evaluating the 3-mm-thick mouth guard (here named **protA**) and the 4-mm-thick mouth guard (here

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**Figure 7.** Test set: the Kratos® Universal Testing Machine, dental arch models and extensometer.

**Figure 8.** Maintenance of the system temperature with an optical pyrometer.

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**Figure 9.** Manufacture of mouth guards in vacuum form Metalvander® machine.

The groups are divided as follows:

**A.** Mouth guard—3-mm-thick blade


**Figure 8.** Maintenance of the system temperature with an optical pyrometer.

**Figure 7.** Test set: the Kratos® Universal Testing Machine, dental arch models and extensometer.

**Figure 6.** Model of dental arches made of epoxy resin.

326 Biomaterials in Regenerative Medicine


**Table 1** shows a variation in the compression maximum load in Newtons (N) when evaluating the 3-mm-thick mouth guard (here named **protA**) and the 4-mm-thick mouth guard (here named **protB**).


When a facial bone is fractured, undergoing or not surgery reduction, it should not be exposed to any trauma during the bone healing process, which lasts about 30 days [23–27]. If surgical reduction is required, it should occur within the first 2–3 h after the injury occurs

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Cases of surgical reduction may disrupt the performance of athletes. In these cases, the use of the facial protector can allow an early and safe return of the athlete to training and competitions [22, 27, 28, 30]. In general, 4–7 days are required for the face molding and for manufac-

For the present FEM analysis, the cortical bone is represented as a linear elastic, homogeneous and isotropic material. The mechanical properties—density, Young's modulus, Poisson coefficient and maximum strength—of each bone depend on its composition, as reported by Lotti

**Table 2** presents the maximum compressive load of each bone portion of interest for dentistry. Particularly for the cortical bone, the elastic material parameters are listed in

The soft tissue named here is composed of the skin and the muscular portion of the studied

The soft tissue is a hyperelastic nonlinear material [33–37] here represented by the well-

**Table 4** lists the parameters used for soft tissue in the FEA. The elastic parameters are the same as those adopted by several car manufacturers to simulate pedestrian—car impact—and Ogden parameters were obtained by Coto et al. [6], according to the definition in the finite-

**Bone Max. load (N) Max compressive stress (N/mm2**

**Structure Young Modulus (MPa) Poisson coefficient Density (t/mm3**

Cortical bone 13,700 0.32 2.28

Frontal 1000–6494 ≥7.58

Zygomatic 489–2401 1.38–4.17 Mandible 668–1801 1.03–2.07 Nasal 342–450 0.13–0.34 **)**

**)**

[28, 29].

**Table 3**.

region.

turing/producing the protector.

et al., Handbook in 2006 [31].

known Ogden model [35, 37].

element software LS-Dyna.

**Table 2.** Face bone resistance [32].

**Table 3.** Elastic parameters for the cortical bone.

*4.1.2. Human soft tissue*

**Table 1.** Maximum load variation (N) as a function of the thickness variable, with their respective standard deviations (SD) and significance levels (*p* ≤ 0.05), for protectors A and B.

When comparing **protA** and **protB** protectors in **Table 1**, **protB** was observed to require an additional force of 173 N. It agrees with Craig and Godwin [14]: "The energy absorbed in the cyclic moment of compressive deformation should reduce the locally transmitted energy" and thus avoid the rupture of the polymer layer between the teeth. The EVA material acts as a shock absorber, guaranteeing a low energy transmission to the teeth of the dental arch [15].

These data become more relevant when the final measurements of the guard thicknesses are observed. At the end, they presented mean differences in thickness of approximately 0.55 mm, instead of the nominal 1 mm difference. This is already expected, since during the manufacture of the individualized buccal protector there is a loss of thickness between 25 and 50%, also observed in the literature [14, 16]. Analyzing **Table 1** again, one can conclude that a small difference, 0.55 mm, increased the force around 173 N.

## **4. Numerical analysis of EVA applied to facial protection**

Studies in the biological area involving impact have become impossible to perform *in vivo* due to ethical awareness. On the other hand, engineering presented rapid technological development of tools that allow for more detailed analyses, using complex geometries and offering refined results of behavior of virtually modeled bodies [17]. Particularly, the finite element method (FEM) is a powerful tool, able to virtually mimic different complex phenomena, including the impact of an object on a human face.

However, to analyze the performance of different EVA geometries and properties (flexible and rigid forms) via FEM, it is necessary to determine the parameters and constitutive laws for the materials (tissue, bone, EVA), geometry of the studied problem (face and projectile) and boundary conditions (initial velocity of the projectile, displacement restrictions in the system).

#### **4.1. Material parameters and constitutive laws**

#### *4.1.1. Face bones*

Most of the bony framework of the face has high-level resistance, since it protects vital elements, such as the brain, the eyes and the neuromuscular structures. Yet, it is also composed of very fragile bones, such as the maxilla, nasal bones and the malar portion of the zygomatic bone [18–22].

When a facial bone is fractured, undergoing or not surgery reduction, it should not be exposed to any trauma during the bone healing process, which lasts about 30 days [23–27]. If surgical reduction is required, it should occur within the first 2–3 h after the injury occurs [28, 29].

Cases of surgical reduction may disrupt the performance of athletes. In these cases, the use of the facial protector can allow an early and safe return of the athlete to training and competitions [22, 27, 28, 30]. In general, 4–7 days are required for the face molding and for manufacturing/producing the protector.

For the present FEM analysis, the cortical bone is represented as a linear elastic, homogeneous and isotropic material. The mechanical properties—density, Young's modulus, Poisson coefficient and maximum strength—of each bone depend on its composition, as reported by Lotti et al., Handbook in 2006 [31].

**Table 2** presents the maximum compressive load of each bone portion of interest for dentistry. Particularly for the cortical bone, the elastic material parameters are listed in **Table 3**.

#### *4.1.2. Human soft tissue*

When comparing **protA** and **protB** protectors in **Table 1**, **protB** was observed to require an additional force of 173 N. It agrees with Craig and Godwin [14]: "The energy absorbed in the cyclic moment of compressive deformation should reduce the locally transmitted energy" and thus avoid the rupture of the polymer layer between the teeth. The EVA material acts as a shock absorber, guaranteeing a low energy transmission to the teeth of the

**Table 1.** Maximum load variation (N) as a function of the thickness variable, with their respective standard deviations

**Coefficient Max. load (N) Standard deviation Significance** *p*

**protA** 2046 20 0.00 **protB** 2219 20 0.00

These data become more relevant when the final measurements of the guard thicknesses are observed. At the end, they presented mean differences in thickness of approximately 0.55 mm, instead of the nominal 1 mm difference. This is already expected, since during the manufacture of the individualized buccal protector there is a loss of thickness between 25 and 50%, also observed in the literature [14, 16]. Analyzing **Table 1** again, one can conclude that a small

Studies in the biological area involving impact have become impossible to perform *in vivo* due to ethical awareness. On the other hand, engineering presented rapid technological development of tools that allow for more detailed analyses, using complex geometries and offering refined results of behavior of virtually modeled bodies [17]. Particularly, the finite element method (FEM) is a powerful tool, able to virtually mimic different complex phenomena,

However, to analyze the performance of different EVA geometries and properties (flexible and rigid forms) via FEM, it is necessary to determine the parameters and constitutive laws for the materials (tissue, bone, EVA), geometry of the studied problem (face and projectile) and boundary conditions (initial velocity of the projectile, displacement restrictions in the system).

Most of the bony framework of the face has high-level resistance, since it protects vital elements, such as the brain, the eyes and the neuromuscular structures. Yet, it is also composed of very fragile bones, such as the maxilla, nasal bones and the malar portion of the zygomatic

difference, 0.55 mm, increased the force around 173 N.

(SD) and significance levels (*p* ≤ 0.05), for protectors A and B.

including the impact of an object on a human face.

**4.1. Material parameters and constitutive laws**

**4. Numerical analysis of EVA applied to facial protection**

dental arch [15].

328 Biomaterials in Regenerative Medicine

*4.1.1. Face bones*

bone [18–22].

The soft tissue named here is composed of the skin and the muscular portion of the studied region.

The soft tissue is a hyperelastic nonlinear material [33–37] here represented by the wellknown Ogden model [35, 37].

**Table 4** lists the parameters used for soft tissue in the FEA. The elastic parameters are the same as those adopted by several car manufacturers to simulate pedestrian—car impact—and Ogden parameters were obtained by Coto et al. [6], according to the definition in the finiteelement software LS-Dyna.


**Table 2.** Face bone resistance [32].


**Table 3.** Elastic parameters for the cortical bone.


**Table 4.** Material model for human tissue.

#### *4.1.3. Flexible EVA*

Flexible EVA has high elasticity and low mechanical resistance. A reverse analysis method was adopted to extract the material properties from the experimental tests described. In the reverse method, material parameters are tuned such that numerical predictions match the experimental curves (**Figure 10**).

**Table 5** summarizes the material parameters, used to characterize flexible EVA, according to the Ogden hyperelastic model, available in the commercial software LS-Dyna® and adopted in this study.

#### *4.1.4. Rigid EVA*

The inverse methodology was again adopted here, to characterize rigid EVA. **Figure 11** shows the similarity between the experimental and the numerical compression tests.

**Table 6** shows the material parameters used for the rigid EVA, according to the Ogden model

**Figure 11.** (A) Experimental and numerical curve for compression tests of rigid EVA. (B) Specimen configuration at

**/α<sup>2</sup> Poisson coefficient Shear Modulus (MPa) Density (t/mm3**

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Flexible EVA 7.0/0.8 2.6/2.6 0.48 10.0 2.0 E − 9

**)**

331

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**)**

As for numerical simulations of the human face, the geometry is a challenge, due to the great

To overcome this problem, a scientific partnership was established with the Renato Archer Information Technology Center (CTI Renato Archer). They provided the face images (**Figure 12**), obtained by computerized tomography (CT) and using in-house

Rigid EVA 1.0/0.05 10.0/−4.0 0.49 175 1.2 E − 9

**/α<sup>2</sup> Poisson coefficient Shear Modulus (MPa) Density (t/mm3**

available in the software LS-Dyna®.

different instants of the numerical analysis [37].

**μ1**

**Table 5.** The material model for flexible EVA.

**/α<sup>1</sup> μ2**

*4.1.5. Geometry*

software.

number of particularities.

**μ1**

**Table 6.** Material model for rigid EVA.

**/α<sup>1</sup> μ2**

**Figure 10.** (A) Experimental and numerical curve for compression test. (B) Specimen configuration at different instants of the numerical analysis [6].


**Table 5.** The material model for flexible EVA.

*4.1.3. Flexible EVA*

in this study.

*4.1.4. Rigid EVA*

of the numerical analysis [6].

experimental curves (**Figure 10**).

**Table 4.** Material model for human tissue.

330 Biomaterials in Regenerative Medicine

Flexible EVA has high elasticity and low mechanical resistance. A reverse analysis method was adopted to extract the material properties from the experimental tests described. In the reverse method, material parameters are tuned such that numerical predictions match the

**Elastic parameters Ogden parameters**

**) μ1**

**/α<sup>1</sup> μ2**

**/α<sup>2</sup>**

**Shear Modulus (MPa) Poisson coefficient Density (t/mm3**

Tissue 0.69 0.495 1.438 E – 9 7.0/0.8 2.6/2.6

**Table 5** summarizes the material parameters, used to characterize flexible EVA, according to the Ogden hyperelastic model, available in the commercial software LS-Dyna® and adopted

The inverse methodology was again adopted here, to characterize rigid EVA. **Figure 11** shows

**Figure 10.** (A) Experimental and numerical curve for compression test. (B) Specimen configuration at different instants

the similarity between the experimental and the numerical compression tests.

**Figure 11.** (A) Experimental and numerical curve for compression tests of rigid EVA. (B) Specimen configuration at different instants of the numerical analysis [37].

**Table 6** shows the material parameters used for the rigid EVA, according to the Ogden model available in the software LS-Dyna®.

#### *4.1.5. Geometry*

As for numerical simulations of the human face, the geometry is a challenge, due to the great number of particularities.

To overcome this problem, a scientific partnership was established with the Renato Archer Information Technology Center (CTI Renato Archer). They provided the face images (**Figure 12**), obtained by computerized tomography (CT) and using in-house software.


**Table 6.** Material model for rigid EVA.

**Figure 12.** (A) The bone and (B) soft tissue configuration of the face, obtained from CT [37].

#### **4.2. Numerical analyses**

Using the material and geometrical parameters defined so far, it is possible to perform complex numerical analyses of the face, with different load conditions. Here, software LS-Dyna was used. The mesh generation and data analyses were performed with the pre- and postprocessors HyperMesh and HyperView, respectively [38, 39].

#### *4.2.1. EVA as nasal protector for sport*

Coto et al. [6] studied the performance of EVA nasal protectors undergoing the impact of a rigid ball in the face with a 3D FE model (**Figure 13**). The material used was a combination of 1 mm of rigid EVA with 2 mm of flexible EVA. The author concluded that the proposed protector could absorb and dissipate the energy from the impact of a ball with mass of 0.025 kg at initial velocity of 20 m/s. The energy is high enough to fracture the nasal bone if there is no protector (**Figure 14**).

**4.3. Study of EVA to protect the zygomatic bone**

protector and (D) with protector. Figure is extracted from Coto et al. [37].

G1 2 1 1 G2 3 1 – G3 2 1 –

exclusively formed by the cortical bone [41].

**Table 7.** Configurations analyzed for rigid and flexible EVA.

to **Table 7**.

The zygomatic bone forms the prominence of the cheek, part of the lateral wall and floor of the orbit. Due to its location and prominence, it presents a high risk of fracture [39–41]. The thickness is not constant in its extension. The zygomatic bone is composed of cortical and spongy bone in the thicker portion, and in the region near the frontonasal suture, it is almost

**Figure 14.** (A, B) Normal compressive stress in the bones of the frontal region, after impact, (A) without and (B) with nasal protector. (C, D) Normal compressive stress in the soft tissue of the frontal region, after the impact, (C) without

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A simplified geometry of overlapping discs with a 100-mm radius was considered. The layers were composed of bone tissue (zygomatic bone portion, lower malar portion, near the nasal front suture), soft tissue and three proposed rigid and flexible EVA combinations, according

**Flexible EVA (thickness, mm) Rigid EVA (thickness, mm) Flexible EVA (thickness, mm)**

According to Coto et al. [37], rigid EVA reduced the velocity of impact and the flexible EVA increased the time interval of the impulse, thus decreasing the peak load transmitted to the bone.

**Figure 13.** FE model. (A) Without the protector. (B) With the protector. Figure is extracted from Coto et al. [37].

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**Figure 14.** (A, B) Normal compressive stress in the bones of the frontal region, after impact, (A) without and (B) with nasal protector. (C, D) Normal compressive stress in the soft tissue of the frontal region, after the impact, (C) without protector and (D) with protector. Figure is extracted from Coto et al. [37].

#### **4.3. Study of EVA to protect the zygomatic bone**

**4.2. Numerical analyses**

332 Biomaterials in Regenerative Medicine

no protector (**Figure 14**).

*4.2.1. EVA as nasal protector for sport*

Using the material and geometrical parameters defined so far, it is possible to perform complex numerical analyses of the face, with different load conditions. Here, software LS-Dyna was used. The mesh generation and data analyses were performed with the pre- and post-

Coto et al. [6] studied the performance of EVA nasal protectors undergoing the impact of a rigid ball in the face with a 3D FE model (**Figure 13**). The material used was a combination of 1 mm of rigid EVA with 2 mm of flexible EVA. The author concluded that the proposed protector could absorb and dissipate the energy from the impact of a ball with mass of 0.025 kg at initial velocity of 20 m/s. The energy is high enough to fracture the nasal bone if there is

According to Coto et al. [37], rigid EVA reduced the velocity of impact and the flexible EVA increased the time interval of the impulse, thus decreasing the peak load transmitted to the bone.

**Figure 13.** FE model. (A) Without the protector. (B) With the protector. Figure is extracted from Coto et al. [37].

processors HyperMesh and HyperView, respectively [38, 39].

**Figure 12.** (A) The bone and (B) soft tissue configuration of the face, obtained from CT [37].

The zygomatic bone forms the prominence of the cheek, part of the lateral wall and floor of the orbit. Due to its location and prominence, it presents a high risk of fracture [39–41]. The thickness is not constant in its extension. The zygomatic bone is composed of cortical and spongy bone in the thicker portion, and in the region near the frontonasal suture, it is almost exclusively formed by the cortical bone [41].

A simplified geometry of overlapping discs with a 100-mm radius was considered. The layers were composed of bone tissue (zygomatic bone portion, lower malar portion, near the nasal front suture), soft tissue and three proposed rigid and flexible EVA combinations, according to **Table 7**.


**Table 7.** Configurations analyzed for rigid and flexible EVA.

**Figure 15** shows the geometry for G1. An extra configuration formed only by the cortical bone and soft tissue was also included in the analyses as a control group (CG) as shown in **Figure 16**.

**Figure 16** also shows the projectile, here represented by a golf ball, with parameters obtained in Bartlett et al. [43] (**Table 8**). The ball had a velocity of 10 m/s at the instant of impact.

The parameters used for the cortical bone, soft tissue, flexible and rigid EVA are in **Tables 3**–**6**, respectively. The thicknesses of bone and soft tissue were 10.3 mm [42] and 12.3 mm [43], respectively.

> The analyses were performed by the LS-Dyna software. The minimum compression stress was controlled. The maximum pressure allowed for the bone and the EVA (rigid or flexible) was of 2.7 MPa and 5.0 MPa, respectively. The friction value considered was 0.5 between ball

**Density (t/mm3**

**) Radius (mm) Velocity (m/s)**

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335

Systematic Study of Ethylene-Vinyl Acetate (EVA) in the Manufacturing of Protector Devices...

**Figure 17** shows the pressure for the CG. The figure shows the high level of pressure at the

According to the analyses, the results showed that in the three models proposed, there was the maximum performance of EVA, but the best protection to the studied bone is given by the

G2 model. **Figure 18** shows the pressure profile in the EVA for G1 and G3.

zygomatic bone, exceeding the failure limit of 2.7 MPa.

**Figure 19** shows the energy conversion during impact in G2.

and soft tissue.

**Young Modulus** 

**Poisson coefficient**

**Table 8.** Geometric and material characteristics of the projectile.

Golf ball 392 0.45 1.15 E − 9 21 10

**(MPa)**

**5. FEA results**

**Figure 17.** Pressure for the CG.

**Figure 15.** Simplified geometry (Group G1), soft EVA, rigid EVA, soft tissue amd bone.

**Figure 16.** Control Group (CG), soft tissue and zigomatic bone.


**Table 8.** Geometric and material characteristics of the projectile.

The analyses were performed by the LS-Dyna software. The minimum compression stress was controlled. The maximum pressure allowed for the bone and the EVA (rigid or flexible) was of 2.7 MPa and 5.0 MPa, respectively. The friction value considered was 0.5 between ball and soft tissue.

#### **5. FEA results**

**Figure 15** shows the geometry for G1. An extra configuration formed only by the cortical bone and soft tissue was also included in the analyses as a control group (CG) as shown in **Figure 16**. **Figure 16** also shows the projectile, here represented by a golf ball, with parameters obtained in Bartlett et al. [43] (**Table 8**). The ball had a velocity of 10 m/s at the instant of impact.

334 Biomaterials in Regenerative Medicine

The parameters used for the cortical bone, soft tissue, flexible and rigid EVA are in **Tables 3**–**6**, respectively. The thicknesses of bone and soft tissue were 10.3 mm [42] and 12.3 mm [43], respectively.

**Figure 15.** Simplified geometry (Group G1), soft EVA, rigid EVA, soft tissue amd bone.

**Figure 16.** Control Group (CG), soft tissue and zigomatic bone.

**Figure 17** shows the pressure for the CG. The figure shows the high level of pressure at the zygomatic bone, exceeding the failure limit of 2.7 MPa.

According to the analyses, the results showed that in the three models proposed, there was the maximum performance of EVA, but the best protection to the studied bone is given by the G2 model. **Figure 18** shows the pressure profile in the EVA for G1 and G3.

**Figure 19** shows the energy conversion during impact in G2.

**Figure 17.** Pressure for the CG.

**6. Conclusions**

**Author details**

**References**

[Accessed: June 10, 2014]

1979;**7**(2):283-290

Reinaldo Brito e Dias1

of any material used in the area must be known.

In human health science research, the study of materials that may replace organs is in constant evolution. Particularly in dentistry, the material should be easy to manipulate, esthetically acceptable, stable to use in the oral environment, in the presence of saliva and chewing conditions and low cost. Moreover, mechanical, physical, chemical and biological properties

Systematic Study of Ethylene-Vinyl Acetate (EVA) in the Manufacturing of Protector Devices...

http://dx.doi.org/10.5772/intechopen.69969

337

EVA was the object of this study. It is a thermoplastic copolymer derived from petroleum.

(initial velocity of the projectile, displacement restrictions in the system).

, Neide Pena Coto<sup>1</sup>

1 School of Dentistry, University of São Paulo, Sao Paulo, Brazil

2 School of Engineering, University of São Paulo, Sao Paulo, Brazil

\*Address all correspondence to: neidecoto@gmail.com

lator)]. 12th ed. Rio de Janeiro: Elsevier; 2013

Initially, the material was studied in the mouth environment, and it was mechanically and chemically characterized. Finally, the material is molded and applied to facial protection. The application is numerical, since studies in the biological area involving impact have become impossible to perform *in vivo*. FEM is a powerful tool, able to virtually mimic different complex phenomena. The quality of the results strongly depends on the correct material characterization, precise geometry of the analyzed structure and real boundary conditions

The facial protector was tested during the impact of a golf ball in the nasal bone and, through a simplified model, in the zygomatic bone. The proposed protector is able to amortize the impact, and its configuration does not compromise peripheral view and does not cause discomfort to the athlete.

[1] WHO. Available from: http://www.direitoshumanos.usp.br/index.php/OMS-Organização-Mundial-da-Saúde/constituicao-da-organizacao-mundial-da-saude-omswho.html. 2014.

[2] Anusavice KJ, Shen C, Rawls R. Phillips materiais dentários. [Roberto Braga, et al. (trans-

[3] Wong EW, White RC. Development of a shock absorbing biomedical elastomer for a new total elbow replacement design. Biomaterials, Medical Devices, and Artificial Organs.

\*, Gilmar Ferreira Batalha<sup>2</sup>

and Larissa Driemeier<sup>2</sup>

**Figure 18.** Results of pressure profile in the protector for G1 and G3, respectively.

**Figure 19.** (a) Energy conversion during impact for G2; pressure in the (b) semirigid and (c) flexible EVA.

## **6. Conclusions**

In human health science research, the study of materials that may replace organs is in constant evolution. Particularly in dentistry, the material should be easy to manipulate, esthetically acceptable, stable to use in the oral environment, in the presence of saliva and chewing conditions and low cost. Moreover, mechanical, physical, chemical and biological properties of any material used in the area must be known.

EVA was the object of this study. It is a thermoplastic copolymer derived from petroleum.

Initially, the material was studied in the mouth environment, and it was mechanically and chemically characterized. Finally, the material is molded and applied to facial protection.

The application is numerical, since studies in the biological area involving impact have become impossible to perform *in vivo*. FEM is a powerful tool, able to virtually mimic different complex phenomena. The quality of the results strongly depends on the correct material characterization, precise geometry of the analyzed structure and real boundary conditions (initial velocity of the projectile, displacement restrictions in the system).

The facial protector was tested during the impact of a golf ball in the nasal bone and, through a simplified model, in the zygomatic bone. The proposed protector is able to amortize the impact, and its configuration does not compromise peripheral view and does not cause discomfort to the athlete.

## **Author details**

**Figure 18.** Results of pressure profile in the protector for G1 and G3, respectively.

336 Biomaterials in Regenerative Medicine

**Figure 19.** (a) Energy conversion during impact for G2; pressure in the (b) semirigid and (c) flexible EVA.

Reinaldo Brito e Dias1 , Neide Pena Coto<sup>1</sup> \*, Gilmar Ferreira Batalha<sup>2</sup> and Larissa Driemeier<sup>2</sup>


## **References**


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**Chapter 15**

**Provisional chapter**

**Tailoring Bioengineered Scaffolds for Regenerative**

The vision to unravel and develop biological healing mechanisms based on evolving molecular and cellular technologies has led to a worldwide scientific endeavor to establish regenerative medicine. This is a multidisciplinary field that involves basic and preclinical research and development on the repair, replacement, and regrowth or regeneration of cells, tissues, or organs in both diseases (congenital or acquired) and traumas. A total of over 63,000 patients were officially placed on organs' waiting lists on 31 December 2013 in the European Union (European Commission, 2014). Tissue engineering and regenerative medicine have emerged as promising fields to achieve proper solutions for these concerns. However, we are far from having patient-specific tissue engineering scaffolds that mimic the native tissue regarding both structure and function. The proposed chapter is a qualitative review over the biomaterials, processes, and scaffold designs for tailored bioprinting. Relevant literature on bioengineered scaffolds for regenerative medicine will be updated. It is well known that mechanical properties play significant effects on biologic behavior which highlight the importance of an extensively discussion on tailoring biomechanical properties for bioengineered scaffolds. The following topics will be discussed: scaffold design, biomaterials and scaffolds bioactivity, biofabrication processes, scaffolds biodegradability, and cell viability. Moreover, new insights will be pointed out. **Keywords:** tailored scaffold, biomaterials, bioprinting, biomechanics, regenerative medicine

**Tailoring Bioengineered Scaffolds for Regenerative** 

DOI: 10.5772/intechopen.69857

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

In a society that is in constant development, the discovery of "new" scientific and technological knowledge must (i) progress at an incredibly fast pace, (ii) target a wide audience, and (iii) have a practical impact in the society. The health sciences are naturally a priority area of

**Medicine**

**Medicine**

Sandra Amado, Pedro Morouço, Paula Pascoal-Faria and Nuno Alves

Sandra Amado, Pedro Morouço, Paula Pascoal-Faria and Nuno Alves

http://dx.doi.org/10.5772/intechopen.69857

**Abstract**

**1. Introduction**

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter


**Provisional chapter**

## **Tailoring Bioengineered Scaffolds for Regenerative Medicine Medicine**

**Tailoring Bioengineered Scaffolds for Regenerative** 

DOI: 10.5772/intechopen.69857

Sandra Amado, Pedro Morouço, Paula Pascoal-Faria and Nuno Alves Paula Pascoal-Faria and Nuno Alves Additional information is available at the end of the chapter

Sandra Amado, Pedro Morouço,

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69857

#### **Abstract**

vivo by indentation and static friction tests. Skin Research and Technology. 2009 Feb;

[37] Coto NP, Driemeier L, Roveri GO, Meira JBC, Dias RB, Noritomi PY. Numerical study of the face bone behaviour when impacted by rigid ball. Journal of Biomechanics. 2012

[38] Withnall C, Shewchenko N, Gittens R, Dvorak J. Biomechanical investigation of head impacts in football. British Journal of Sports Medicine. 2005 Aug;**39**(Suppl 1):i49-i57 [39] Verschueren P, Delye H, Depreitere B, Van Lierde C, Haex B, Berckmans D. A new test set-up for skull fracture characterisation. Journal of Biomechanics. 2007;**40**(15):3389-3396

[40] Gialain IO, Dias RB, Coto NP. Mouthguard: A new technique for the partially edentu-

[41] Del Neri NB, Araujo-Pires AC, Andreo JC, Rubira-Bullen IR, Ferreira Júnior O. Zygomaticofacial foramen location accuracy and reliability in cone-beam computed

[42] de Almeida NH, Michel-Crosato E, de Paiva LA, Biazevic MG. Facial soft tissue thickness in the Brazilian population: New reference data and anatomical landmarks. Forensic

[43] Bartlett R, Gratton C, Rolf CG, editors. Encyclopedia of International Sports Studies.

tomography. Acta Odontologica Scandinavica. 2014 Feb;**72**(2):157-160

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Abington: Routledge; 1987

**15**(1):68-76

340 Biomaterials in Regenerative Medicine

Jul;**45**:1121

The vision to unravel and develop biological healing mechanisms based on evolving molecular and cellular technologies has led to a worldwide scientific endeavor to establish regenerative medicine. This is a multidisciplinary field that involves basic and preclinical research and development on the repair, replacement, and regrowth or regeneration of cells, tissues, or organs in both diseases (congenital or acquired) and traumas. A total of over 63,000 patients were officially placed on organs' waiting lists on 31 December 2013 in the European Union (European Commission, 2014). Tissue engineering and regenerative medicine have emerged as promising fields to achieve proper solutions for these concerns. However, we are far from having patient-specific tissue engineering scaffolds that mimic the native tissue regarding both structure and function. The proposed chapter is a qualitative review over the biomaterials, processes, and scaffold designs for tailored bioprinting. Relevant literature on bioengineered scaffolds for regenerative medicine will be updated. It is well known that mechanical properties play significant effects on biologic behavior which highlight the importance of an extensively discussion on tailoring biomechanical properties for bioengineered scaffolds. The following topics will be discussed: scaffold design, biomaterials and scaffolds bioactivity, biofabrication processes, scaffolds biodegradability, and cell viability. Moreover, new insights will be pointed out.

**Keywords:** tailored scaffold, biomaterials, bioprinting, biomechanics, regenerative medicine

## **1. Introduction**

In a society that is in constant development, the discovery of "new" scientific and technological knowledge must (i) progress at an incredibly fast pace, (ii) target a wide audience, and (iii) have a practical impact in the society. The health sciences are naturally a priority area of

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

research, mostly because of the impact they have on the augment human life expectancy, by developing advanced and patient-specific therapies.

providing an adequate biomechanical environment for tissue regeneration [4]. However, adapting the adequate technology with enhanced biomaterials, in order to obtain customized

Tailoring Bioengineered Scaffolds for Regenerative Medicine

http://dx.doi.org/10.5772/intechopen.69857

343

This chapter intends to provide a synopses in patient-specific engineering scaffolds. A revision of the scaffold design, biomaterials, and advanced manufacturing processes will help to establish new research paradigms on tailoring bioengineered scaffolds for regenerative medicine. Recent advances will be highlighted to stimulate the readers for future insights and

Scaffold modeling plays a key role in tissue engineering and regenerative medicine. A welldesigned 3D scaffold is a fundamental tool to guide tissue formation both in vitro and in vivo. Properties such as high surface-area-to-volume ratio, porosity, pore size, pore design, pore interconnectivity, permeability, and degradation should be taken into account when designing scaffold for different and tailored applications. These will allow a desirable biological network for cell migration, nutrient transportation, and the mechanical stiffness, and strength can be therefore obtained [5, 6]. Growth factors (GFs) and drug release (DR) should also be considered to achieve an optimized tissue growth as scaffold degraded. Moreover, some authors have shown the benefits for tissue generation of using curvature and concave

To address and fulfill aforementioned requirements, two scaffold design approaches can be used according to the flowchart presented in **Figure 2**. The first one is based on the native tissue, whereas the second one is based on the unit digital cell model, both addressing tailored scaffold geometry. The geometry obtained can then be used on computer-aided engineering studies to optimize the performance of the tailored bioengineering scaffold. Finally, a physical optimized scaffold can be produced using 3D printing or AM technology before in vitro and/or in vivo implantation of the scaffold. Accordingly, several research works have been developed concerning tailored scaffold geometry and its fabrication. In these studies, physical scaffolds have been used directly for in vitro and/or in vivo studies. Nevertheless, the link between computer modeling and computer-aided engineering to tailor bioengineering scaffold remains a paramount challenge. When solved, it can significantly reduce animal

In the computer modeling based on native tissue, different noninvasive 3D scanning techniques can be considered to obtain the 3D anatomical geometric model. The most used are computed tomography (CT), μCT (micro-CT), and nCT (nano-CT), which considered different scale levels [8–10], as well as magnetic resonance image (MRI) and 3D optical techniques. All these techniques used different physical principles to obtain a series of two-dimensional (2D) images or a 3D point cloud of the sample of the native tissue studied. CT requires the exposition of the sample of the native tissue to ionizing radiation, whereas MRI uses a magnetic field and pulses of radio wave energy (avoiding radiation) both obtaining a series of 2D images. In

implants that mimic the native tissue, is nowadays a challenge with a huge potential.

possibilities.

**2. Scaffold design**

experimental studies.

surfaces compared to convex and planar ones [7].

Only the complexity of human tissues could justify that in the 1980s tissue engineering emerged as a scientific field with an enormous potential. Targeting to regenerate the bone, cartilage, skin, or other tissues and organs, bridging the anatomy with its physiology/function is a paramount challenge to be solved. Several efforts have been made, by research groups spread worldwide, to tailor bioengineering scaffolds (sometimes denominated by tissue constructs) that could mimic native tissues. However, the achievement of three-dimensional (3D) complex organ structures is far from being tangible. Due to its nature, tissue engineering gathers scientists, engineers, and physicians in multidisciplinary teams using a variety of methods to construct biological substitutes [1]. Indeed, significant efforts are being developed worldwide in the fields of tissue engineering and regenerative medicine, but full tissue or organ regeneration remains a paramount challenge. Therefore, these multidisciplinary scientific fields apply a wide variety of methodologies, where multidisciplinary research teams can provide suitable inputs for its development [2].

One of the major goals is to produce biological substitutes to restore, maintain, or improve tissue function, using biocompatible and biodegradable support structures, i.e., scaffolds, in conjunction with human cells (**Figure 1**) [3]. Gathering tissue engineering and regenerative medicine, researchers have been interested on developing alternative approaches for restoring functionality. To do so, one of the most promising methodologies involves the use of additive manufacturing (AM) processes. AM technologies allow the production of complex 3D structures concerning mainly a high level of control, predefined geometry, size, and interconnected pores in a reproducible way. This optimized controlled organization enhances the vascularization and, thus, transports oxygen and nutrients throughout the whole structure,

**Figure 1.** Relevant steps based on the use of scaffolds for tissue engineering.

providing an adequate biomechanical environment for tissue regeneration [4]. However, adapting the adequate technology with enhanced biomaterials, in order to obtain customized implants that mimic the native tissue, is nowadays a challenge with a huge potential.

This chapter intends to provide a synopses in patient-specific engineering scaffolds. A revision of the scaffold design, biomaterials, and advanced manufacturing processes will help to establish new research paradigms on tailoring bioengineered scaffolds for regenerative medicine. Recent advances will be highlighted to stimulate the readers for future insights and possibilities.

## **2. Scaffold design**

research, mostly because of the impact they have on the augment human life expectancy, by

Only the complexity of human tissues could justify that in the 1980s tissue engineering emerged as a scientific field with an enormous potential. Targeting to regenerate the bone, cartilage, skin, or other tissues and organs, bridging the anatomy with its physiology/function is a paramount challenge to be solved. Several efforts have been made, by research groups spread worldwide, to tailor bioengineering scaffolds (sometimes denominated by tissue constructs) that could mimic native tissues. However, the achievement of three-dimensional (3D) complex organ structures is far from being tangible. Due to its nature, tissue engineering gathers scientists, engineers, and physicians in multidisciplinary teams using a variety of methods to construct biological substitutes [1]. Indeed, significant efforts are being developed worldwide in the fields of tissue engineering and regenerative medicine, but full tissue or organ regeneration remains a paramount challenge. Therefore, these multidisciplinary scientific fields apply a wide variety of methodologies, where multidisciplinary research teams can

One of the major goals is to produce biological substitutes to restore, maintain, or improve tissue function, using biocompatible and biodegradable support structures, i.e., scaffolds, in conjunction with human cells (**Figure 1**) [3]. Gathering tissue engineering and regenerative medicine, researchers have been interested on developing alternative approaches for restoring functionality. To do so, one of the most promising methodologies involves the use of additive manufacturing (AM) processes. AM technologies allow the production of complex 3D structures concerning mainly a high level of control, predefined geometry, size, and interconnected pores in a reproducible way. This optimized controlled organization enhances the vascularization and, thus, transports oxygen and nutrients throughout the whole structure,

developing advanced and patient-specific therapies.

342 Biomaterials in Regenerative Medicine

provide suitable inputs for its development [2].

**Figure 1.** Relevant steps based on the use of scaffolds for tissue engineering.

Scaffold modeling plays a key role in tissue engineering and regenerative medicine. A welldesigned 3D scaffold is a fundamental tool to guide tissue formation both in vitro and in vivo. Properties such as high surface-area-to-volume ratio, porosity, pore size, pore design, pore interconnectivity, permeability, and degradation should be taken into account when designing scaffold for different and tailored applications. These will allow a desirable biological network for cell migration, nutrient transportation, and the mechanical stiffness, and strength can be therefore obtained [5, 6]. Growth factors (GFs) and drug release (DR) should also be considered to achieve an optimized tissue growth as scaffold degraded. Moreover, some authors have shown the benefits for tissue generation of using curvature and concave surfaces compared to convex and planar ones [7].

To address and fulfill aforementioned requirements, two scaffold design approaches can be used according to the flowchart presented in **Figure 2**. The first one is based on the native tissue, whereas the second one is based on the unit digital cell model, both addressing tailored scaffold geometry. The geometry obtained can then be used on computer-aided engineering studies to optimize the performance of the tailored bioengineering scaffold. Finally, a physical optimized scaffold can be produced using 3D printing or AM technology before in vitro and/or in vivo implantation of the scaffold. Accordingly, several research works have been developed concerning tailored scaffold geometry and its fabrication. In these studies, physical scaffolds have been used directly for in vitro and/or in vivo studies. Nevertheless, the link between computer modeling and computer-aided engineering to tailor bioengineering scaffold remains a paramount challenge. When solved, it can significantly reduce animal experimental studies.

In the computer modeling based on native tissue, different noninvasive 3D scanning techniques can be considered to obtain the 3D anatomical geometric model. The most used are computed tomography (CT), μCT (micro-CT), and nCT (nano-CT), which considered different scale levels [8–10], as well as magnetic resonance image (MRI) and 3D optical techniques. All these techniques used different physical principles to obtain a series of two-dimensional (2D) images or a 3D point cloud of the sample of the native tissue studied. CT requires the exposition of the sample of the native tissue to ionizing radiation, whereas MRI uses a magnetic field and pulses of radio wave energy (avoiding radiation) both obtaining a series of 2D images. In

internal structures, such as the central nervous system, heart, and kidneys of a rat [9]. The 2D individual images obtained using the previous techniques described are then assembled and realigned, and therefore a 3D geometric anatomical model distinguishing different types of

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Microscopy optical technique is also used to obtain the 3D anatomical geometric model of the native tissue. However, it can only differentiate every type of tissue down to the level of the individual cell at the cost of a huge computational effort. Other 3D optical techniques, such as 3D structured light, cannot differentiate the types of tissues presented and only allow the generation of the outer 3D geometric model of the native sample. In the case of the native tissue sample (temporomandibular joint disc (TMJ disc)) presented in the flowchart of **Figure 2**, a 3D point cloud was obtained using a white light 3D scanning system (Steinbichler—COMET 5®), and then an appropriate software was applied to replicate the 3D geometric anatomical

Hybrid modalities can also be used to construct the 3D model of the same specimen in order to take the advantage of each technique to differentiate the different types of tissues [9].

The second approach uses computer-aided design (CAD) techniques to create a 3D unit cell which is used as pattern. Then, a desired number of patterns are automatically generated and combined until a complete 3D geometric model of the scaffold is obtained with controlled architecture (**Figure 2**). Following this approach, the main scientific achievements reported

Permanent tailored engineering scaffolds have been designed mainly for bone repair of large segmental defects caused by fracture, tumor, or infection. In 2013, Wieding et al. [13] reported a numerical study which is used to determine the suitability of open porous of titanium scaffolds to act as bone scaffolds under physiological loading conditions. Uniaxial compression structural modulus of the titanium scaffolds was tailored ranging from 3.5 to 19.1 GPa as a function of the scaffold porosity from 64 to 80%. Results revealed that minimizing the amount of material of the inner core had a smaller influence than increasing the porosity when the scaffolds were under biomechanical loading. It was also noted that the scaffold design could act similarly to the intact bone. In order to tailor the mechanical properties of cellular structured scaffolds, [14] designed metal scaffolds with high porosity (62–92%) to tailor both compressive strength (4.0–113.0 MPa) and elastic modulus (0.2–6.3 GPa), respectively, were comparable to trabecular and cortical bone. Porous titanium scaffolds were also investigated by van der Stok et al. [15, 16] for grafting large bone defects. Mechanical properties were tailored, whereas high porosity of the scaffold allowed the incorporation of colloidal gelatin gels for time- and dose-controlled delivery of dual growth factors (bone morphogenetic protein-2 (BMP-2) and/or fibroblast growth factor-2 (FGF-2)), promoting a quasi-full bone regeneration. The scaffold was designed based on a decahedron pattern and composed by 120-μm-thick titanium struts with porous size ranging from 240 to 730 μm. Porous size, porosity, porous volume, compression strength,

BMP-2/FGF-2. In 2012, Van Bael et al. [17] developed six distinct geometries of Ti6Al4V scaffolds in three different pore shapes (triangular, hexagonal, and rectangular) and two different

, 14 MPa, and 0.4 GPa, respectively, allowing

) for a composite scaffold with

tissue is obtained.

model of the TMJ disc [12].

are based on permanent or temporary scaffolds.

and Young's modulus were 490 μm, 88%, 55 mm<sup>3</sup>

to achieve an optimized bone volume regeneration (~50 mm<sup>3</sup>

**Figure 2.** Computer modeling and simulation to tailor bioengineering scaffolds.

CT, these images are displayed by density, while in MRI they are compiled and segmented by its signal intensity. Additionally, both techniques can be differentiated by its resolution. The high resolution of CT allows the characterization of the micro-architecture and the mechanical properties of the tissue scaffolds [11]; however, this technique has a drawback regarding soft tissues of similar density. It is more efficient in differentiating hard tissues with sharply defined density changes, such as the interface between bone and soft tissues. To overcome this problem, contrast agents can be added [6]. Although the resolution of MRI is inferior to CT scans, with the advance of technology, it is improving, allowing the 3D representation of internal structures, such as the central nervous system, heart, and kidneys of a rat [9]. The 2D individual images obtained using the previous techniques described are then assembled and realigned, and therefore a 3D geometric anatomical model distinguishing different types of tissue is obtained.

Microscopy optical technique is also used to obtain the 3D anatomical geometric model of the native tissue. However, it can only differentiate every type of tissue down to the level of the individual cell at the cost of a huge computational effort. Other 3D optical techniques, such as 3D structured light, cannot differentiate the types of tissues presented and only allow the generation of the outer 3D geometric model of the native sample. In the case of the native tissue sample (temporomandibular joint disc (TMJ disc)) presented in the flowchart of **Figure 2**, a 3D point cloud was obtained using a white light 3D scanning system (Steinbichler—COMET 5®), and then an appropriate software was applied to replicate the 3D geometric anatomical model of the TMJ disc [12].

Hybrid modalities can also be used to construct the 3D model of the same specimen in order to take the advantage of each technique to differentiate the different types of tissues [9].

The second approach uses computer-aided design (CAD) techniques to create a 3D unit cell which is used as pattern. Then, a desired number of patterns are automatically generated and combined until a complete 3D geometric model of the scaffold is obtained with controlled architecture (**Figure 2**). Following this approach, the main scientific achievements reported are based on permanent or temporary scaffolds.

Permanent tailored engineering scaffolds have been designed mainly for bone repair of large segmental defects caused by fracture, tumor, or infection. In 2013, Wieding et al. [13] reported a numerical study which is used to determine the suitability of open porous of titanium scaffolds to act as bone scaffolds under physiological loading conditions. Uniaxial compression structural modulus of the titanium scaffolds was tailored ranging from 3.5 to 19.1 GPa as a function of the scaffold porosity from 64 to 80%. Results revealed that minimizing the amount of material of the inner core had a smaller influence than increasing the porosity when the scaffolds were under biomechanical loading. It was also noted that the scaffold design could act similarly to the intact bone. In order to tailor the mechanical properties of cellular structured scaffolds, [14] designed metal scaffolds with high porosity (62–92%) to tailor both compressive strength (4.0–113.0 MPa) and elastic modulus (0.2–6.3 GPa), respectively, were comparable to trabecular and cortical bone. Porous titanium scaffolds were also investigated by van der Stok et al. [15, 16] for grafting large bone defects. Mechanical properties were tailored, whereas high porosity of the scaffold allowed the incorporation of colloidal gelatin gels for time- and dose-controlled delivery of dual growth factors (bone morphogenetic protein-2 (BMP-2) and/or fibroblast growth factor-2 (FGF-2)), promoting a quasi-full bone regeneration. The scaffold was designed based on a decahedron pattern and composed by 120-μm-thick titanium struts with porous size ranging from 240 to 730 μm. Porous size, porosity, porous volume, compression strength, and Young's modulus were 490 μm, 88%, 55 mm<sup>3</sup> , 14 MPa, and 0.4 GPa, respectively, allowing to achieve an optimized bone volume regeneration (~50 mm<sup>3</sup> ) for a composite scaffold with BMP-2/FGF-2. In 2012, Van Bael et al. [17] developed six distinct geometries of Ti6Al4V scaffolds in three different pore shapes (triangular, hexagonal, and rectangular) and two different

CT, these images are displayed by density, while in MRI they are compiled and segmented by its signal intensity. Additionally, both techniques can be differentiated by its resolution. The high resolution of CT allows the characterization of the micro-architecture and the mechanical properties of the tissue scaffolds [11]; however, this technique has a drawback regarding soft tissues of similar density. It is more efficient in differentiating hard tissues with sharply defined density changes, such as the interface between bone and soft tissues. To overcome this problem, contrast agents can be added [6]. Although the resolution of MRI is inferior to CT scans, with the advance of technology, it is improving, allowing the 3D representation of

**Figure 2.** Computer modeling and simulation to tailor bioengineering scaffolds.

344 Biomaterials in Regenerative Medicine

pore sizes (500 and 1000 μm) aiming to understand the effect of pore geometry of Ti6Al4V bone scaffolds on the in vitro biological behavior of human periosteum-derived cells. The main result showed that a functional Ti6Al4V-graded scaffold, with specific morphological and mechanical properties, will contribute to enhance cell seeding and at the same time can maintain nutrient transport throughout the whole scaffold during in vitro culturing by avoiding pore occlusion.

In 2012, Melchels et al. [28] reviewed additive manufacturing of tissues and organs. Authors also addressed tailored engineering scaffolds for breast reconstruction, focusing pore size and porosity for the generation of three scaffold models. Cipitria et al. [29] developed a poly (ε-caprolactone) (PCL) scaffold incorporating recombinant human bone morphogenetic protein 7 (rhBMP-7) for the regeneration of critical-sized defects in sheep tibiae. PCL scaffold with b-tricalcium phosphate (mPCL-TCP) to promote bone regeneration was designed based on a honeycomb structure with large interconnected pores to facilitate cellular bridging, ingrowth of bone tissue, and efficient mass transport and vascular infiltration. Moreover, Domingos et al. [30] developed PCL scaffolds for tissue engineering purposes. Authors addressed internal/ external scaffold geometry, different material deposition strategies, and the biocompatibility of the material used. 3D PCL porous scaffolds (rectangular porous prisms) were designed with an average porosity of ~76% using commercial computer-aided design software. These structures were then produced via bioextrusion in a 0/90 lay-down pattern trying to reproduce a honeycomb-like pattern of fully interconnected square pores. Similar bioextruded scaffolds were designed (regular dimensions of 600 × 600 mm) to have a well-defined internal geometry with square interconnected pores and uniform distribution. The overall porosity of the structures was found to be ~76%. In vitro degradation of the scaffold was studied as a function of the degradation environment, pore size, and geometry [30, 31]. Scaffold degradation plays a key role when tailoring scaffold properties. In 2016, Morouço et al. [32] developed three types of PCL scaffolds reinforced with cellulose nanofibers (CNF), with and without the addition of hydroxyapatite nanoparticles (HANP), aiming to tailor scaffold properties for tissue engineering applications. The authors studied scaffold porosity, mechanical properties, and biocompatibility as a function of three material combinations. PCL, PCL/CNF, and PCL/CNF/HANP scaffolds were described with porous fully interconnected and porosity (%) of 49.0, 49.5, and 50.0; compressive modulus (MPa) of 54.42, 64.58, and 70.88; and maximum compressive stress (MPa) of 10.96, 11.35, and 12.12, respectively. These structures were then produced via bioextrusion in a 0/90 lay-down pattern. Some authors [33] studied hybrid hierarchical 3D scaffolds with well-controlled architecture for both macro- and microscale. Hybrid and hierarchical 3D structures include thick filaments with the diameter of hundreds of microns, and thin filaments with sub-10 μm dimensions were developed. The microscale features can help in cell seeding, alignment, and guidance. Trying to mimic morphological and mechanical behavior of a blood vessel, Vaz et al. [34] proposed a tailored tissue engineering scaffold. Design parameters such as bilayered tubular scaffold, stiff and oriented outside fibrous layer, and a pliable and randomly oriented fibrous inner layer were considered, combining two biomaterials (PLA/PCL). Structural and mechanical properties of the scaffolds were examined using scanning electron microscopy (SEM) and tensile testing. Cell viability was investigated using 3T3 mouse fibroblasts and the tubular scaffold in an appropriated in vitro environment. The proposed scaffold presented appropriate characteristics to be considered a candidate for blood vessel tissue engineering. Other authors also proposed to tailor tissue engineering scaffolds trying to mimic

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extracellular matrix morphology of natural tissue for blood vessel applications [35, 36].

In 2012, Chantarapanich et al. [37] developed a computer-aided design library based on polyhedrons for tissue engineering applications. Close-cellular scaffold included truncated octahedron, rhombicuboctahedron, and rhombitruncated cuboctahedron, while open-cellular

Temporary (or biodegradable) tailored engineering scaffolds have been designed as a tissue engineering approach that uses degradable porous biomaterial incorporating biological cells and/or molecules to regenerate tissues such as the bone, cartilage, skeletal muscle, nerve, and blood vessels. Scaffold design must be able to create hierarchical porous structures to fulfill all mechanical and biological requirements. In 2005, Hollister [18] introduced the concept of hierarchical scaffold design as geometric features at scales from the nanometer to millimeter level that will determine how well the scaffold meets conflicting mechanical function and mass transport needs. In 2011, Khoda et al. [19] developed a functionally gradient variational porosity architecture (hierarchical design) for hollowed scaffolds. In 2014, Giannitelli et al. [20] reviewed tailored scaffold architecture with microstructural features. Authors highlighted the growing interest in the development of innovative scaffold designs to overcome often conflict requirements (such as biological and mechanical ones). Considering different pore size gradients, Sobral et al. [21] designed and manufactured. The goal was to enhance cell seeding efficiency and control the spatial organization of cells within the scaffold. Some authors [22] also emphasized the importance of scaffold pore size gradients in osteogenic differentiation of human mesenchymal stromal cells. In 2010, Puppi et al. [23] in deep reviewed the design of biodegradable and bioactive polymeric scaffolds, with properly suited architecture and tailored properties for bone and cartilage tissue regeneration. According to the authors, a good scaffold design must account that macro- and microstructural properties affect cells survival, signaling, growth, propagation, and reorganization and play also a major role in modeling cell shape and gene expressions, both related to cell growth and preservation of native phenotypes [24, 25]. In addition, several scaffold designs were developed and then manufactured using different AM processes. For example, Fierz et al. [26] designed three labeled anisotropic 3D hydroxyapatite scaffolds (pixel-wise and labeled layer-wise) with tailored pores ranging from the nanometer to millimeter scale for the reconstruction of centimeter-sized osseous defects. Seventy percent micrometer-wide pores were successfully interconnected, and virtual spheres (diameter of up to 350 ± 35 μm) were used to simulate cell migration along the pores linked with central channel. Melchels et al. [27] designed poly-dl-lactic acid (PDLLA) porous scaffolds with a gyroid architecture. This architecture was mathematically defined, allowing a precise control of porosity and pore size of a fully interconnected pore network. As noted by the authors, cell seeding of porous structures prepared from hydrophobic polymers, such as PDLLA, was difficult. Moreover, the penetration of a cell suspension was further hindered by the high tortuosity and poor interconnectivity of pore networks when manufactured by salt-leaching or freeze-drying conventional methods. Therefore, very open scaffold structure of the gyroid architecture that facilitates the penetration of water into PDLLA scaffold was manufactured by stereolithography. It was highlighted that the cells were well attached and homogeneously distributed throughout the porous scaffold. Good mechanical properties can be tailored in predesigned (porous) architectures from PDLLA based on gyroid architecture. In 2012, Melchels et al. [28] reviewed additive manufacturing of tissues and organs. Authors also addressed tailored engineering scaffolds for breast reconstruction, focusing pore size and porosity for the generation of three scaffold models. Cipitria et al. [29] developed a poly (ε-caprolactone) (PCL) scaffold incorporating recombinant human bone morphogenetic protein 7 (rhBMP-7) for the regeneration of critical-sized defects in sheep tibiae. PCL scaffold with b-tricalcium phosphate (mPCL-TCP) to promote bone regeneration was designed based on a honeycomb structure with large interconnected pores to facilitate cellular bridging, ingrowth of bone tissue, and efficient mass transport and vascular infiltration. Moreover, Domingos et al. [30] developed PCL scaffolds for tissue engineering purposes. Authors addressed internal/ external scaffold geometry, different material deposition strategies, and the biocompatibility of the material used. 3D PCL porous scaffolds (rectangular porous prisms) were designed with an average porosity of ~76% using commercial computer-aided design software. These structures were then produced via bioextrusion in a 0/90 lay-down pattern trying to reproduce a honeycomb-like pattern of fully interconnected square pores. Similar bioextruded scaffolds were designed (regular dimensions of 600 × 600 mm) to have a well-defined internal geometry with square interconnected pores and uniform distribution. The overall porosity of the structures was found to be ~76%. In vitro degradation of the scaffold was studied as a function of the degradation environment, pore size, and geometry [30, 31]. Scaffold degradation plays a key role when tailoring scaffold properties. In 2016, Morouço et al. [32] developed three types of PCL scaffolds reinforced with cellulose nanofibers (CNF), with and without the addition of hydroxyapatite nanoparticles (HANP), aiming to tailor scaffold properties for tissue engineering applications. The authors studied scaffold porosity, mechanical properties, and biocompatibility as a function of three material combinations. PCL, PCL/CNF, and PCL/CNF/HANP scaffolds were described with porous fully interconnected and porosity (%) of 49.0, 49.5, and 50.0; compressive modulus (MPa) of 54.42, 64.58, and 70.88; and maximum compressive stress (MPa) of 10.96, 11.35, and 12.12, respectively. These structures were then produced via bioextrusion in a 0/90 lay-down pattern. Some authors [33] studied hybrid hierarchical 3D scaffolds with well-controlled architecture for both macro- and microscale. Hybrid and hierarchical 3D structures include thick filaments with the diameter of hundreds of microns, and thin filaments with sub-10 μm dimensions were developed. The microscale features can help in cell seeding, alignment, and guidance. Trying to mimic morphological and mechanical behavior of a blood vessel, Vaz et al. [34] proposed a tailored tissue engineering scaffold. Design parameters such as bilayered tubular scaffold, stiff and oriented outside fibrous layer, and a pliable and randomly oriented fibrous inner layer were considered, combining two biomaterials (PLA/PCL). Structural and mechanical properties of the scaffolds were examined using scanning electron microscopy (SEM) and tensile testing. Cell viability was investigated using 3T3 mouse fibroblasts and the tubular scaffold in an appropriated in vitro environment. The proposed scaffold presented appropriate characteristics to be considered a candidate for blood vessel tissue engineering. Other authors also proposed to tailor tissue engineering scaffolds trying to mimic extracellular matrix morphology of natural tissue for blood vessel applications [35, 36].

pore sizes (500 and 1000 μm) aiming to understand the effect of pore geometry of Ti6Al4V bone scaffolds on the in vitro biological behavior of human periosteum-derived cells. The main result showed that a functional Ti6Al4V-graded scaffold, with specific morphological and mechanical properties, will contribute to enhance cell seeding and at the same time can maintain nutrient transport throughout the whole scaffold during in vitro culturing by avoiding pore occlusion. Temporary (or biodegradable) tailored engineering scaffolds have been designed as a tissue engineering approach that uses degradable porous biomaterial incorporating biological cells and/or molecules to regenerate tissues such as the bone, cartilage, skeletal muscle, nerve, and blood vessels. Scaffold design must be able to create hierarchical porous structures to fulfill all mechanical and biological requirements. In 2005, Hollister [18] introduced the concept of hierarchical scaffold design as geometric features at scales from the nanometer to millimeter level that will determine how well the scaffold meets conflicting mechanical function and mass transport needs. In 2011, Khoda et al. [19] developed a functionally gradient variational porosity architecture (hierarchical design) for hollowed scaffolds. In 2014, Giannitelli et al. [20] reviewed tailored scaffold architecture with microstructural features. Authors highlighted the growing interest in the development of innovative scaffold designs to overcome often conflict requirements (such as biological and mechanical ones). Considering different pore size gradients, Sobral et al. [21] designed and manufactured. The goal was to enhance cell seeding efficiency and control the spatial organization of cells within the scaffold. Some authors [22] also emphasized the importance of scaffold pore size gradients in osteogenic differentiation of human mesenchymal stromal cells. In 2010, Puppi et al. [23] in deep reviewed the design of biodegradable and bioactive polymeric scaffolds, with properly suited architecture and tailored properties for bone and cartilage tissue regeneration. According to the authors, a good scaffold design must account that macro- and microstructural properties affect cells survival, signaling, growth, propagation, and reorganization and play also a major role in modeling cell shape and gene expressions, both related to cell growth and preservation of native phenotypes [24, 25]. In addition, several scaffold designs were developed and then manufactured using different AM processes. For example, Fierz et al. [26] designed three labeled anisotropic 3D hydroxyapatite scaffolds (pixel-wise and labeled layer-wise) with tailored pores ranging from the nanometer to millimeter scale for the reconstruction of centimeter-sized osseous defects. Seventy percent micrometer-wide pores were successfully interconnected, and virtual spheres (diameter of up to 350 ± 35 μm) were used to simulate cell migration along the pores linked with central channel. Melchels et al. [27] designed poly-dl-lactic acid (PDLLA) porous scaffolds with a gyroid architecture. This architecture was mathematically defined, allowing a precise control of porosity and pore size of a fully interconnected pore network. As noted by the authors, cell seeding of porous structures prepared from hydrophobic polymers, such as PDLLA, was difficult. Moreover, the penetration of a cell suspension was further hindered by the high tortuosity and poor interconnectivity of pore networks when manufactured by salt-leaching or freeze-drying conventional methods. Therefore, very open scaffold structure of the gyroid architecture that facilitates the penetration of water into PDLLA scaffold was manufactured by stereolithography. It was highlighted that the cells were well attached and homogeneously distributed throughout the porous scaffold. Good mechanical properties can be tailored in predesigned (porous) architectures from PDLLA based on gyroid architecture.

346 Biomaterials in Regenerative Medicine

In 2012, Chantarapanich et al. [37] developed a computer-aided design library based on polyhedrons for tissue engineering applications. Close-cellular scaffold included truncated octahedron, rhombicuboctahedron, and rhombitruncated cuboctahedron, while open-cellular scaffold included hexahedron, truncated octahedron, truncated hexahedron, cuboctahedron, rhombicuboctahedron, and rhombitruncated cuboctahedron. Both relationship between pore size and porosity of close-cellular scaffolds and relationship between pore size/beam thickness and porosity of open-cellular scaffolds were studied. The study concluded that some design combinations were not good for making the open-cellular scaffold, generating enclosed pores inside the scaffold, and, therefore, they were excluded from the digital library. Compressive stresses were computed as a function of polyhedron-based geometries which can also be helpful for tailoring mechanical properties of the scaffolds.

synthetic materials can be used for the fabrication of scaffolds including polymers, ceramics, bioactive glass, calcium phosphates, and biometals. For example, scaffolds fabricated from bioactive ceramic materials such as hydroxyapatite and tricalcium phosphate show promise because of their biological ability to support bone tissue regeneration. However, the use of ceramics as scaffold materials is limited because of their inherent brittleness and difficult processability [39]. In 2006, Rezwan et al.'s [41] review showed that conventional material processing methods have been adapted and extended for incorporation of inorganic bioactive phases into porous and interconnected 3D polymer networks. The biomaterials were extended from purely synthetic materials to material/biologic hybrids, engineering at the same time bioactivity and biodegradability [41]. Addressing this issues, in 2015, Fiedler et al. [42] focused on the mechanical characterization of PCL-bioglass composites and concluded that the addition of bioglass was found to decrease the elastic gradient and yield stress if two scaffolds of the same density are compared and the highest bioglass content (35%) seems beneficial as it (i) does not significantly deteriorate the scaffold mechanical properties and

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The next generation of synthetic biodegradable, bioactive, living composite biomaterials that feature high adaptiveness to the biological environment [41] considers the incorporation of biomolecules as promising and is currently under extensive research. Incorporating biomolecules such as growth factors during scaffold processing with the aim to accelerate local tissue healing however are not simple as biomolecules are sensitive to elevated temperatures and extreme chemical conditions. A promising strategy is the immobilization of proteins and growth factors

"Soft" material routes like sol-gel processing might be a strategy to incorporate biomolecules during scaffold fabrication. To the authors' knowledge, however, sol-gel-derived bioactive organic/inorganic hybrids have not yet been formed into highly interconnected porous structures, which would be essential for application of these composites as scaffolds. Another related challenge was the elucidation of the local impact of growth factors on the cell and tissue systems, including long-term effects [41]. As pointed out in Section 2, mechanical property is one of the most critical parameters that determine the performance of a designed implant. It mainly depends on the process and structural properties of the biomaterials. Therefore, it is possible to achieve desired mechanical properties through modifying the structural characteristics of a biomaterial. Biological behavior of cell assessment after surface modifications is required to check its biocompatibility and bioactivity [38]. The study of the interactions of biochemical and geometrical cues on stem cell differentiation and alignment should be also considered. The capability to spatially control stem cell orientation and differentiation toward multiple phenotypes simultaneously, i.e., myocyte, tenocyte, and osteoblast, allows cells grown in vitro to more closely mimic aspects of native tissue organization and structure [44]. Although the precise mechanism behind geometry-induced cell alignment is presently unknown, it is likely that the alignment of cells observed on fibers may be attributed to a combination of factors including physical space constraint and relative stiffness of the underlying substrate (fiber); ultimately affecting changes in both cell spreading and cell stiffness, cells may be predisposed toward a specific orientation through the modulation of mechanotrans-

in the post-processing phase via surface functionalization of the scaffold [43].

duction pathways via cytoskeletal rearrangements [9].

(ii) promotes bioactivity.

In the computer-aided engineering based on tailored scaffold geometry, several digital features should be taken into account to obtain computer-tailored bioengineering scaffolds. Such features encompass cell and growth factor encapsulating, cell aggregation, cell-cell and cell-tissue interaction, vascularization, scaffold degradation (or not if permanent) and tissue growth, drug release, and scaffold mechanical behavior (**Figure 2**). To help digital prediction of cell/tissue phenomena, several automated methods exist, namely, cell counting, cell geometry determination, chromosomal counting, correlation of DNA expression determined through microarrays, interpreting fluorescence data, determining cell's lineage, and cross correlating gene expression with predicted in vivo pathology. All of these features have predictive value for determination of tissue viability and the differentiative rate of cells seeded with the goal of tissue culture. A detailed description about both accumulation of the expression data and large-scale computer cross correlation (between this expression and expressions commonly used in pathology) is provided in Ref. [9], as well as a number of specific tools for tissue analysis/identification.

Despite of the aforementioned research works, new scaffold designs integrating cell/GF/tissue phenomena and scaffold mechanical behavior (geometric characteristics and materials) are needed for regenerative medicine. These complex hierarchical 3D structures must be designed according to the structural heterogeneity of the host tissue and/or scaffold environment.

## **3. Biomaterials and scaffold bioactivity**

Tissues possess different structures and properties that a tissue engineering scaffold should be tailored to. A general requirement for all biomaterial scaffolds is to reproduce an extracellular matrix (ECM) environment for supporting cell growth outside of the body. Moreover, scaffold should host cell adhesion, proliferation, and ECM production. Hence, the scaffold should surrogate the missing ECM. Tissue engineering products can be designed to conduct, induct, or block tissue responses and architectures [38]. Besides providing the three-dimensional growth of cells in an organized way, an ideal scaffold should be characterized by biocompatibility, biodegradability, appropriate mechanical properties, interconnectivity of pores with appropriate size to retain cells, and low exchanges of nutrients and waste products [39]. Tailoring biomaterials for enhanced biofunctionality can be achieved using a variety of approaches that involve the introduction of chemical, topographical, or mechanical cues via top-down or bottom-up approaches [40]. Therefore, the selection of the starting materials and of the fabrication techniques is of paramount importance. Numerous natural and synthetic materials can be used for the fabrication of scaffolds including polymers, ceramics, bioactive glass, calcium phosphates, and biometals. For example, scaffolds fabricated from bioactive ceramic materials such as hydroxyapatite and tricalcium phosphate show promise because of their biological ability to support bone tissue regeneration. However, the use of ceramics as scaffold materials is limited because of their inherent brittleness and difficult processability [39]. In 2006, Rezwan et al.'s [41] review showed that conventional material processing methods have been adapted and extended for incorporation of inorganic bioactive phases into porous and interconnected 3D polymer networks. The biomaterials were extended from purely synthetic materials to material/biologic hybrids, engineering at the same time bioactivity and biodegradability [41]. Addressing this issues, in 2015, Fiedler et al. [42] focused on the mechanical characterization of PCL-bioglass composites and concluded that the addition of bioglass was found to decrease the elastic gradient and yield stress if two scaffolds of the same density are compared and the highest bioglass content (35%) seems beneficial as it (i) does not significantly deteriorate the scaffold mechanical properties and (ii) promotes bioactivity.

scaffold included hexahedron, truncated octahedron, truncated hexahedron, cuboctahedron, rhombicuboctahedron, and rhombitruncated cuboctahedron. Both relationship between pore size and porosity of close-cellular scaffolds and relationship between pore size/beam thickness and porosity of open-cellular scaffolds were studied. The study concluded that some design combinations were not good for making the open-cellular scaffold, generating enclosed pores inside the scaffold, and, therefore, they were excluded from the digital library. Compressive stresses were computed as a function of polyhedron-based geometries which can also be help-

In the computer-aided engineering based on tailored scaffold geometry, several digital features should be taken into account to obtain computer-tailored bioengineering scaffolds. Such features encompass cell and growth factor encapsulating, cell aggregation, cell-cell and cell-tissue interaction, vascularization, scaffold degradation (or not if permanent) and tissue growth, drug release, and scaffold mechanical behavior (**Figure 2**). To help digital prediction of cell/tissue phenomena, several automated methods exist, namely, cell counting, cell geometry determination, chromosomal counting, correlation of DNA expression determined through microarrays, interpreting fluorescence data, determining cell's lineage, and cross correlating gene expression with predicted in vivo pathology. All of these features have predictive value for determination of tissue viability and the differentiative rate of cells seeded with the goal of tissue culture. A detailed description about both accumulation of the expression data and large-scale computer cross correlation (between this expression and expressions commonly used in pathology) is provided in Ref. [9], as well as a number of specific tools for

Despite of the aforementioned research works, new scaffold designs integrating cell/GF/tissue phenomena and scaffold mechanical behavior (geometric characteristics and materials) are needed for regenerative medicine. These complex hierarchical 3D structures must be designed according to the structural heterogeneity of the host tissue and/or scaffold environment.

Tissues possess different structures and properties that a tissue engineering scaffold should be tailored to. A general requirement for all biomaterial scaffolds is to reproduce an extracellular matrix (ECM) environment for supporting cell growth outside of the body. Moreover, scaffold should host cell adhesion, proliferation, and ECM production. Hence, the scaffold should surrogate the missing ECM. Tissue engineering products can be designed to conduct, induct, or block tissue responses and architectures [38]. Besides providing the three-dimensional growth of cells in an organized way, an ideal scaffold should be characterized by biocompatibility, biodegradability, appropriate mechanical properties, interconnectivity of pores with appropriate size to retain cells, and low exchanges of nutrients and waste products [39]. Tailoring biomaterials for enhanced biofunctionality can be achieved using a variety of approaches that involve the introduction of chemical, topographical, or mechanical cues via top-down or bottom-up approaches [40]. Therefore, the selection of the starting materials and of the fabrication techniques is of paramount importance. Numerous natural and

ful for tailoring mechanical properties of the scaffolds.

tissue analysis/identification.

348 Biomaterials in Regenerative Medicine

**3. Biomaterials and scaffold bioactivity**

The next generation of synthetic biodegradable, bioactive, living composite biomaterials that feature high adaptiveness to the biological environment [41] considers the incorporation of biomolecules as promising and is currently under extensive research. Incorporating biomolecules such as growth factors during scaffold processing with the aim to accelerate local tissue healing however are not simple as biomolecules are sensitive to elevated temperatures and extreme chemical conditions. A promising strategy is the immobilization of proteins and growth factors in the post-processing phase via surface functionalization of the scaffold [43].

"Soft" material routes like sol-gel processing might be a strategy to incorporate biomolecules during scaffold fabrication. To the authors' knowledge, however, sol-gel-derived bioactive organic/inorganic hybrids have not yet been formed into highly interconnected porous structures, which would be essential for application of these composites as scaffolds. Another related challenge was the elucidation of the local impact of growth factors on the cell and tissue systems, including long-term effects [41]. As pointed out in Section 2, mechanical property is one of the most critical parameters that determine the performance of a designed implant. It mainly depends on the process and structural properties of the biomaterials. Therefore, it is possible to achieve desired mechanical properties through modifying the structural characteristics of a biomaterial. Biological behavior of cell assessment after surface modifications is required to check its biocompatibility and bioactivity [38]. The study of the interactions of biochemical and geometrical cues on stem cell differentiation and alignment should be also considered. The capability to spatially control stem cell orientation and differentiation toward multiple phenotypes simultaneously, i.e., myocyte, tenocyte, and osteoblast, allows cells grown in vitro to more closely mimic aspects of native tissue organization and structure [44]. Although the precise mechanism behind geometry-induced cell alignment is presently unknown, it is likely that the alignment of cells observed on fibers may be attributed to a combination of factors including physical space constraint and relative stiffness of the underlying substrate (fiber); ultimately affecting changes in both cell spreading and cell stiffness, cells may be predisposed toward a specific orientation through the modulation of mechanotransduction pathways via cytoskeletal rearrangements [9].

Multilayer scaffolds and combinations of several biomaterials are a better option to create graded structures that resemble the biological interface. The development of multilayer scaffolds and the controlled release of bioactive molecules to promote in situ regeneration of biological tissues are some of the latest technologies that intended to improve on the available traditional treatments. To confirm the potential of these novel approaches, long-term evaluation is necessary with special focus on studying the biological and mechanical properties of the synthesized tissues [45].

Regarding the jet-based techniques, with a common resolution of 10–50 μm, it is difficult to obtain an adequate structural support. It consists of dispensing a jet of small droplets of liquid material, also called as bioink, in a spatially controlled manner. There are two different approaches, thermal inkjet printing and piezoelectric-actuated inkjet printing, having the former lower suitability for 3D bioengineering scaffolds. Using a piezoelectric actuator, research has been able to suppress some of the thermal constraints [51]. For instance, good viability of printed cell populations was obtained for human fibroblast cell line [52], and recently a silk-based ink eliminated the usage of any cytotoxic organic or inorganic solvents [53]. Even though jet-based techniques are the pioneer techniques used for tissue engineering, translating it to the construct of large 3D structures is a challenge to overcome, mostly because of the

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351

The most successful attempts to engineer cell-containing bioengineering scaffolds have been achieved through robotic dispensing systems. These technologies are based on a controlled extrusion of a material in a continuous fashion, instead of liquid droplets (**Figure 3**) and are developed at the Center for Rapid and Sustainable Development (CDRSP) of the Polytechnic Institute of Leiria, Portugal. Therefore, this approach enables the printing of hydrogels encapsulating cells in a very controlled architecture [28]. The most common methods are the pneumatic [54] or mechanical [55] dispensing systems, comprising (i) a dispensing system and a stage with the capability of moving along the *x*, *y*, and *z* axes; (ii) a light source to illuminate the working area and/or for photoinitiator activation; and (iii) a piezoelectric humidifier [56], with some of them using multiple printing heads to permit the dispensing of various materials without retooling [57]. However, researcher should bear in mind that optimal balance

low-viscous solutions that do not provide strong and complex 3D structures.

should be aimed between pressure and nozzle size (to obtain higher cell viability).

**Figure 3.** Equipments developed at the CDRSP, Portugal: (a) one-head extrusion system, (b) dual-head extrusion system, (c) system combining an extrusion head and a syringe for hydrogel deposition, and (d) hybrid system combining

extrusion with up to three hydrogels.

Scaffolds should be designed more as a bioactive system rather than just passive cell carriers. Thus, integration of fabrication techniques with surface modification may also act as route to obtain nanofibrous scaffolds with better understanding of cell scaffolds both in vivo and in vitro. Similarly and significantly, the biomaterial as design strategy can be used in a better way to relate science and engineering, and use this advanced knowledge to engineer more advanced tissue scaffolds [46, 47].

## **4. Biofabrication processes**

Aiming to tailor bioengineering scaffolds that closely mimic the native tissues, AM technologies are suitable to dispense biomaterials (with live cells or cell aggregates) at specific, and desired, locations [48]. The usage of these technologies has been commonly divided in three categories: (i) the jet-based techniques, (ii) robotic dispensing techniques, and (iii) laserinduced forward transfer [49, 50]. Each of these techniques has advantages and drawbacks (**Table 1**). Thus, understanding its limitations and potentials is a must-do to choose the right approach for the specific tissue that is aimed to regenerate. Furthermore, some advancements have been recently achieved with integrated/hybrid systems. These systems combine different techniques within the same equipment aiming to generate a multifunctional graded construct with tailored properties similar to the native tissue.

In the available literature, it is possible to find investigations using different approaches, for the same type of tissue. In fact, there seems to be a trend to some research groups get specialized in some type of technology and use it for various goals.


**Table 1.** Comparison of the three AM approaches for tissue engineering.

Regarding the jet-based techniques, with a common resolution of 10–50 μm, it is difficult to obtain an adequate structural support. It consists of dispensing a jet of small droplets of liquid material, also called as bioink, in a spatially controlled manner. There are two different approaches, thermal inkjet printing and piezoelectric-actuated inkjet printing, having the former lower suitability for 3D bioengineering scaffolds. Using a piezoelectric actuator, research has been able to suppress some of the thermal constraints [51]. For instance, good viability of printed cell populations was obtained for human fibroblast cell line [52], and recently a silk-based ink eliminated the usage of any cytotoxic organic or inorganic solvents [53]. Even though jet-based techniques are the pioneer techniques used for tissue engineering, translating it to the construct of large 3D structures is a challenge to overcome, mostly because of the low-viscous solutions that do not provide strong and complex 3D structures.

Multilayer scaffolds and combinations of several biomaterials are a better option to create graded structures that resemble the biological interface. The development of multilayer scaffolds and the controlled release of bioactive molecules to promote in situ regeneration of biological tissues are some of the latest technologies that intended to improve on the available traditional treatments. To confirm the potential of these novel approaches, long-term evaluation is necessary with special focus on studying the biological and mechanical properties of

Scaffolds should be designed more as a bioactive system rather than just passive cell carriers. Thus, integration of fabrication techniques with surface modification may also act as route to obtain nanofibrous scaffolds with better understanding of cell scaffolds both in vivo and in vitro. Similarly and significantly, the biomaterial as design strategy can be used in a better way to relate science and engineering, and use this advanced knowledge to engineer more

Aiming to tailor bioengineering scaffolds that closely mimic the native tissues, AM technologies are suitable to dispense biomaterials (with live cells or cell aggregates) at specific, and desired, locations [48]. The usage of these technologies has been commonly divided in three categories: (i) the jet-based techniques, (ii) robotic dispensing techniques, and (iii) laserinduced forward transfer [49, 50]. Each of these techniques has advantages and drawbacks (**Table 1**). Thus, understanding its limitations and potentials is a must-do to choose the right approach for the specific tissue that is aimed to regenerate. Furthermore, some advancements have been recently achieved with integrated/hybrid systems. These systems combine different techniques within the same equipment aiming to generate a multifunctional graded con-

In the available literature, it is possible to find investigations using different approaches, for the same type of tissue. In fact, there seems to be a trend to some research groups get special-

**Jet based Robotic dispensing Laser-induced forward** 

**transfer**

the synthesized tissues [45].

350 Biomaterials in Regenerative Medicine

advanced tissue scaffolds [46, 47].

**4. Biofabrication processes**

Adapted from Ref. [51].

struct with tailored properties similar to the native tissue.

ized in some type of technology and use it for various goals.

**Table 1.** Comparison of the three AM approaches for tissue engineering.

Resolution + +/– ++ Fabrication speed +/– ++ − Hydrogel viscosity – + +/− Gelation speed ++ +/– ++ Cell density – + +/– The most successful attempts to engineer cell-containing bioengineering scaffolds have been achieved through robotic dispensing systems. These technologies are based on a controlled extrusion of a material in a continuous fashion, instead of liquid droplets (**Figure 3**) and are developed at the Center for Rapid and Sustainable Development (CDRSP) of the Polytechnic Institute of Leiria, Portugal. Therefore, this approach enables the printing of hydrogels encapsulating cells in a very controlled architecture [28]. The most common methods are the pneumatic [54] or mechanical [55] dispensing systems, comprising (i) a dispensing system and a stage with the capability of moving along the *x*, *y*, and *z* axes; (ii) a light source to illuminate the working area and/or for photoinitiator activation; and (iii) a piezoelectric humidifier [56], with some of them using multiple printing heads to permit the dispensing of various materials without retooling [57]. However, researcher should bear in mind that optimal balance should be aimed between pressure and nozzle size (to obtain higher cell viability).

**Figure 3.** Equipments developed at the CDRSP, Portugal: (a) one-head extrusion system, (b) dual-head extrusion system, (c) system combining an extrusion head and a syringe for hydrogel deposition, and (d) hybrid system combining extrusion with up to three hydrogels.

Recently, a US research group presented an integrated tissue-organ printer (ITOP) for the production of human-scale tissue bioengineering scaffolds of any shape, in a single structure [50]. Combining different procedures, it was possible to successfully engineer (i) a mandible bone, (ii) an ear-shaped cartilage, and (iii) a skeletal muscle. Furthermore, some institutions are now combining different technologies, for instance, merging a robotic dispensing system with a jet-based printing for muscle-tendon unit repair [58].

tissue construct fabrication process. These biomaterials can encompass the physical environment inside of which the cells will reside, as well as the biochemical signal cells need to func-

Tailoring Bioengineered Scaffolds for Regenerative Medicine

http://dx.doi.org/10.5772/intechopen.69857

353

Scaffolds represent the space available for the tissue to develop and the physical support for cell growth. Scaffold mechanical properties should allow shape maintenance during tissue regeneration and enable stress transfer and load bearing. Moreover, during the first stage of tissue reconstruction, wound contraction forces act against the process, and enough mechanical strength and stiffness of the scaffold is required. Scaffold porosity is a fundamental characteristic for providing available space for cells to migrate and for vascularization of the tissue. Furthermore, the larger the surface available, the more cell interactions will arise. In general, the biological activity of a scaffold is determined by ligand density. Scaffold composition and porous fraction, that is, the total surface of the structure exposed to cells, determine the ligand density. Highly specific surface areas allow for cell attachment and anchorage, and a high pore volume fraction enables cell growth, migration, and effective transportation of fluids and nutrients. In particular, microporosity is important for capillary ingrowth and interactions between cells and matrix, while macroporosity is relevant to nutrient supply and waste removal of cell metabolism. The rate of degradation is also strictly connected to the degree of

As in the development of the tissue-engineered organs, regeneration of functional tissue requires maintenance of cell viability and differentiated function, encouragement of cell proliferation, modulation of the direction and speed of cell migration, and regulation of cellular adhesion [61]. Cell viability may be judged by morphological changes or by changes in membrane permeability and/or physiological state inferred from the exclusion of certain dyes or the uptake and retention of others. Cultured cells are seeded onto a three-dimensional biocompatible scaffold that will slowly degrade and resorb as the soft and hard structures grow and assimilate in vitro and/or in vivo [2]. Cell viability during 3D bioprinting is dependent on the shear stress experienced during extrusion, which in turn is dependent on the viscosity of the solution, the applied pressure, and the needle diameter. In addition, any post-printing bioink cross-linking may also impact on cell viability [62]. Cell viability can be measured with Live/Dead Viability/ Cytotoxicity assay after printing [63] and could vary with dispensing pressure and nozzle diameter. It decreases as the pressure increases and the nozzle diameter decreases, and it is seen that the effect of pressure is significantly larger than the effect of the nozzle diameter. At higher pres-

sures, there is an increase in the number of apoptotic cells as well as necrotic cells [64].

demonstrate that novel biofabrication techniques may resolve this problem [65].

Tissue bioengineered scaffolds targeted for in vivo applications are typically restricted to a thickness of only a few hundreds of microns, owing to the diffusion limitations of oxygen and nutrients [43]. One of the major challenges in tissue engineering for translation in clinical applications is the vascularization of bioengineering scaffolds of clinically relevant size. Insufficient vascularization inhibits nutrient and host cell delivery or migrations and leads to improper cell integration or cell death. While vascularization remains a challenge to maintain viability of large biofabricated tissue bioengineering scaffolds, recent advances in the field

tion as they would in the body [46].

porosity [38].

Lastly, the laser-induced forward transfer (LIFT) is a not-so-common technology for tissue engineering, but is gaining significant importance in this domain [56]. It is based on using three layers of different components: the first layer based on a donor slide, covered by a laser energy-absorbing layer and completed with a cell-bioink component [51]. There are three main advantages for this technology: it is suitable for using (i) a wide range of materials, (ii) a very high precise deposition (but in small 3D structures), and (iii) a clog-free process without the use of nozzles. However, it requires a rapid gelation process, and researchers should bear in mind that several factors should be considered (e.g., laser wavelength, bioink viscosity; for more info read Ref. [59]). Apart from these constraints, successful cell viability (>90%) has been reported for printing skin cell lines and human mesenchymal stem cells and to prepare a cardiac patch [60].

## **5. Scaffold biodegradability and cell viability**

The main objective of tissue engineering is to allow the cells of the body to replace the implanted scaffold over a period. Because bioengineering scaffolds are not intended as permanent implants (besides some of them have shown good results mainly in bone regeneration), they must therefore be biodegradable, so that the need of surgical removal can be avoided. Furthermore, the degradation products should be nontoxic and should be able to swiftly exit from the body without interference with other organs. In addition to this, the intermediate product, the timing of the degradation process, and the route and mechanism of degradation are equally important aspects that need to be taken care [47]. Scaffold materials should fulfill several requirements. A scaffold is not just a passive support for cell growth, but a device whose properties affects the regeneration cascade. Mechanical properties, surface properties, and morphology are in turn relevant to the specific application. Degradation kinetics and the rate at which scaffold properties change with degradation should always be predictable. In particular, the degradation behavior of biomaterials can follow several mechanisms and is controlled by different factors. Understanding the degradation kinetics and mechanism of biomaterials is necessary to optimize their possible usage. The rate of degradation is also strictly connected to the degree of porosity [38].

One of the general variables that need to be thoroughly considered to successfully bioprint viable and functional tissue bioengineering scaffolds is the inclusion of supportive biomaterials, generally in the form of proteins and polymers, which (1) facilitate the deposition method by mechanical means and (2) provide support and protection to the cells during and after the tissue construct fabrication process. These biomaterials can encompass the physical environment inside of which the cells will reside, as well as the biochemical signal cells need to function as they would in the body [46].

Recently, a US research group presented an integrated tissue-organ printer (ITOP) for the production of human-scale tissue bioengineering scaffolds of any shape, in a single structure [50]. Combining different procedures, it was possible to successfully engineer (i) a mandible bone, (ii) an ear-shaped cartilage, and (iii) a skeletal muscle. Furthermore, some institutions are now combining different technologies, for instance, merging a robotic dispensing system

Lastly, the laser-induced forward transfer (LIFT) is a not-so-common technology for tissue engineering, but is gaining significant importance in this domain [56]. It is based on using three layers of different components: the first layer based on a donor slide, covered by a laser energy-absorbing layer and completed with a cell-bioink component [51]. There are three main advantages for this technology: it is suitable for using (i) a wide range of materials, (ii) a very high precise deposition (but in small 3D structures), and (iii) a clog-free process without the use of nozzles. However, it requires a rapid gelation process, and researchers should bear in mind that several factors should be considered (e.g., laser wavelength, bioink viscosity; for more info read Ref. [59]). Apart from these constraints, successful cell viability (>90%) has been reported for printing skin cell lines and human mesenchymal stem cells and to prepare

The main objective of tissue engineering is to allow the cells of the body to replace the implanted scaffold over a period. Because bioengineering scaffolds are not intended as permanent implants (besides some of them have shown good results mainly in bone regeneration), they must therefore be biodegradable, so that the need of surgical removal can be avoided. Furthermore, the degradation products should be nontoxic and should be able to swiftly exit from the body without interference with other organs. In addition to this, the intermediate product, the timing of the degradation process, and the route and mechanism of degradation are equally important aspects that need to be taken care [47]. Scaffold materials should fulfill several requirements. A scaffold is not just a passive support for cell growth, but a device whose properties affects the regeneration cascade. Mechanical properties, surface properties, and morphology are in turn relevant to the specific application. Degradation kinetics and the rate at which scaffold properties change with degradation should always be predictable. In particular, the degradation behavior of biomaterials can follow several mechanisms and is controlled by different factors. Understanding the degradation kinetics and mechanism of biomaterials is necessary to optimize their possible usage. The rate of degradation is also

One of the general variables that need to be thoroughly considered to successfully bioprint viable and functional tissue bioengineering scaffolds is the inclusion of supportive biomaterials, generally in the form of proteins and polymers, which (1) facilitate the deposition method by mechanical means and (2) provide support and protection to the cells during and after the

with a jet-based printing for muscle-tendon unit repair [58].

**5. Scaffold biodegradability and cell viability**

strictly connected to the degree of porosity [38].

a cardiac patch [60].

352 Biomaterials in Regenerative Medicine

Scaffolds represent the space available for the tissue to develop and the physical support for cell growth. Scaffold mechanical properties should allow shape maintenance during tissue regeneration and enable stress transfer and load bearing. Moreover, during the first stage of tissue reconstruction, wound contraction forces act against the process, and enough mechanical strength and stiffness of the scaffold is required. Scaffold porosity is a fundamental characteristic for providing available space for cells to migrate and for vascularization of the tissue. Furthermore, the larger the surface available, the more cell interactions will arise. In general, the biological activity of a scaffold is determined by ligand density. Scaffold composition and porous fraction, that is, the total surface of the structure exposed to cells, determine the ligand density. Highly specific surface areas allow for cell attachment and anchorage, and a high pore volume fraction enables cell growth, migration, and effective transportation of fluids and nutrients. In particular, microporosity is important for capillary ingrowth and interactions between cells and matrix, while macroporosity is relevant to nutrient supply and waste removal of cell metabolism. The rate of degradation is also strictly connected to the degree of porosity [38].

As in the development of the tissue-engineered organs, regeneration of functional tissue requires maintenance of cell viability and differentiated function, encouragement of cell proliferation, modulation of the direction and speed of cell migration, and regulation of cellular adhesion [61]. Cell viability may be judged by morphological changes or by changes in membrane permeability and/or physiological state inferred from the exclusion of certain dyes or the uptake and retention of others. Cultured cells are seeded onto a three-dimensional biocompatible scaffold that will slowly degrade and resorb as the soft and hard structures grow and assimilate in vitro and/or in vivo [2]. Cell viability during 3D bioprinting is dependent on the shear stress experienced during extrusion, which in turn is dependent on the viscosity of the solution, the applied pressure, and the needle diameter. In addition, any post-printing bioink cross-linking may also impact on cell viability [62]. Cell viability can be measured with Live/Dead Viability/ Cytotoxicity assay after printing [63] and could vary with dispensing pressure and nozzle diameter. It decreases as the pressure increases and the nozzle diameter decreases, and it is seen that the effect of pressure is significantly larger than the effect of the nozzle diameter. At higher pressures, there is an increase in the number of apoptotic cells as well as necrotic cells [64].

Tissue bioengineered scaffolds targeted for in vivo applications are typically restricted to a thickness of only a few hundreds of microns, owing to the diffusion limitations of oxygen and nutrients [43]. One of the major challenges in tissue engineering for translation in clinical applications is the vascularization of bioengineering scaffolds of clinically relevant size. Insufficient vascularization inhibits nutrient and host cell delivery or migrations and leads to improper cell integration or cell death. While vascularization remains a challenge to maintain viability of large biofabricated tissue bioengineering scaffolds, recent advances in the field demonstrate that novel biofabrication techniques may resolve this problem [65].

## **6. New insights: 3D to 4D**

Doing a survey on the Web of Science®, it is noticeable that the number of original articles on tissue engineering and regenerative medicine has experienced a tremendous increase over the past 10 years (**Figure 4**; review papers and proceedings not included). Likewise, bioprinting is attracting a lot of researchers presenting an exponential increase in the last 3 years. Meanwhile, 3D bioprinting market was valued at \$98.6 million in 2015, and an annual growth of 36% for the next 6 years is expected [66].

automation of stimulus, and standardizing the assessment procedures to evaluate the result

This publication is supported by the Portuguese Foundation for Science and Technology (FCT) through the following projects: UID/Multi/04044/2013 and PTDC/EMS-SIS/7032/2014.

, Paula Pascoal-Faria<sup>4</sup>

1 Centre for Rapid and Sustainable Product Development, Polytechnic Institute of Leiria,

2 CIPER-FMH, Centro Interdisciplinar de Estudo de Performance Humana, Faculdade de

4 Centre for Rapid and Sustainable Product Development & Mathematics Department, School of Technology and Management, Polytechnic Institute of Leiria, Leiria, Portugal

[1] Morouço P, Alves N, Amado S. The role of biomechanics in tissue engineering. Austin

[2] Hutmacher DW. Scaffold design and fabrication technologies for engineering tissues state of the art and future perspectives. Journal of Biomaterials Science Polymer Edition.

[3] Malda J, Groll J. A step towards clinical translation of biofabrication. Trends in

[4] Zadpoor AA, Malda J. Additive manufacturing of biomaterials, tissues, and organs.

[5] Parker GJM, Wheeler-Kingshott CAM, Barker GJ. Diffusion tensor imaging. IEEE

Motricidade Humana (FMH), Universidade de Lisboa (UL), Lisboa, Portugal

3 School of Health Sciences at the Polytechnic Institute of Leiria, Leiria, Portugal

5 Mechanical Engineering Department, School of Technology and Management,

and Nuno Alves1,5

Tailoring Bioengineered Scaffolds for Regenerative Medicine

http://dx.doi.org/10.5772/intechopen.69857

355

is crucial for enhanced regenerative medicine approaches [68].

\*Address all correspondence to: sandra.amado@ipleiria.pt

Polytechnic Institute of Leiria, Leiria, Portugal

Biotechnology. 2016;**34**(5):356-357

Journal of Biomedical Engineering. 2016;**3**(1):1-2

Annals of Biomedical Engineering. 2017;**45**(1):1-11

Transactions on Medical Imaging. 2002;**21**(5):505-512

**Acknowledgements**

**Author details**

Leiria, Portugal

**References**

2001;**12**(1):107-124

Sandra Amado1,2,3\*, Pedro Morouço1

Nevertheless, 3D bioprinting has been focused on the development of bioengineering scaffolds that lack a crucial element for mimicking native live tissues: its ability to acutely change according to its function. That is why leading research groups have recently proposed the four-dimensional (4D) bioprinting (time is integrated with 3D bioprinting) as an enhanced approach for tissue engineering and regenerative medicine: the development of stimuliresponsive biomaterials that can be printed and dynamic to intended stimulation. However, several challenges arise, namely, (i) bioinks have to be optimized to achieve successful bioprinting; (ii) processes must be mechanically designed to obtain robust shape-changing capability of the bioengineering scaffolds [67]; (iii) specific bioreactors for complex tissue function maturation need to be invented; and (iv) evaluation procedures should be defined to examine the functionality response.

Therefore, the most promising approach is to optimize the cell-bioengineering scaffold interactions, becoming feasible to explore the usage of computer modeling to examine the further responses. Developing "smart" biomaterials (also referred as "intelligent," "stimuli responsive," "stimuli sensitive," or "environmentally sensitive") to allow the dynamic changes of the structure, upgrade of the printing processes into defined architecture for targeting tissues,

**Figure 4.** Number of original articles on tissue engineering and regenerative medicine 2006–2016.

automation of stimulus, and standardizing the assessment procedures to evaluate the result is crucial for enhanced regenerative medicine approaches [68].

## **Acknowledgements**

**6. New insights: 3D to 4D**

354 Biomaterials in Regenerative Medicine

the functionality response.

of 36% for the next 6 years is expected [66].

Doing a survey on the Web of Science®, it is noticeable that the number of original articles on tissue engineering and regenerative medicine has experienced a tremendous increase over the past 10 years (**Figure 4**; review papers and proceedings not included). Likewise, bioprinting is attracting a lot of researchers presenting an exponential increase in the last 3 years. Meanwhile, 3D bioprinting market was valued at \$98.6 million in 2015, and an annual growth

Nevertheless, 3D bioprinting has been focused on the development of bioengineering scaffolds that lack a crucial element for mimicking native live tissues: its ability to acutely change according to its function. That is why leading research groups have recently proposed the four-dimensional (4D) bioprinting (time is integrated with 3D bioprinting) as an enhanced approach for tissue engineering and regenerative medicine: the development of stimuliresponsive biomaterials that can be printed and dynamic to intended stimulation. However, several challenges arise, namely, (i) bioinks have to be optimized to achieve successful bioprinting; (ii) processes must be mechanically designed to obtain robust shape-changing capability of the bioengineering scaffolds [67]; (iii) specific bioreactors for complex tissue function maturation need to be invented; and (iv) evaluation procedures should be defined to examine

Therefore, the most promising approach is to optimize the cell-bioengineering scaffold interactions, becoming feasible to explore the usage of computer modeling to examine the further responses. Developing "smart" biomaterials (also referred as "intelligent," "stimuli responsive," "stimuli sensitive," or "environmentally sensitive") to allow the dynamic changes of the structure, upgrade of the printing processes into defined architecture for targeting tissues,

**Figure 4.** Number of original articles on tissue engineering and regenerative medicine 2006–2016.

This publication is supported by the Portuguese Foundation for Science and Technology (FCT) through the following projects: UID/Multi/04044/2013 and PTDC/EMS-SIS/7032/2014.

## **Author details**

Sandra Amado1,2,3\*, Pedro Morouço1 , Paula Pascoal-Faria<sup>4</sup> and Nuno Alves1,5

\*Address all correspondence to: sandra.amado@ipleiria.pt

1 Centre for Rapid and Sustainable Product Development, Polytechnic Institute of Leiria, Leiria, Portugal

2 CIPER-FMH, Centro Interdisciplinar de Estudo de Performance Humana, Faculdade de Motricidade Humana (FMH), Universidade de Lisboa (UL), Lisboa, Portugal

3 School of Health Sciences at the Polytechnic Institute of Leiria, Leiria, Portugal

4 Centre for Rapid and Sustainable Product Development & Mathematics Department, School of Technology and Management, Polytechnic Institute of Leiria, Leiria, Portugal

5 Mechanical Engineering Department, School of Technology and Management, Polytechnic Institute of Leiria, Leiria, Portugal

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**Chapter 16**

**Provisional chapter**

**Biomaterials and Stem Cells: Promising Tools in Tissue**

Biomaterial sciences and tissue engineering approaches are currently fundamental strat‐ egies for the development of regenerative medicine. Stem cells (SCs) are a unique cell type capable of self‐renewal and reconstructing damaged tissues. At the present time, adult SCs isolated from postnatal tissues are widely used in clinical applications. Their characteristics such as a multipotent differentiation capacity and immunomodulatory activity make them a promising tool to use in patients. Modern material technologies allow for the development of innovative biomaterials that closely correspond to require‐ ments of the current biomedical application. Biomaterials, such as ceramics and metals, are already used as implants to replace or improve the functionality of the damaged tissue or organ. However, the continuous development of modern technology opens new insights of polymeric and smart material applications. Moreover, biomaterials may enhance the SCs biological activity and their implementation by establishing a specific microenvironment mimicking natural cell niche. Thus, the synergistic advancement in the fields of biomaterial and medical sciences constitutes a challenge for the development of effective therapies in humans including combined applications of novel biomaterials

**Keywords:** adult stem cells, biomaterials, regenerative medicine, tissue engineering

**Biomaterials and Stem Cells: Promising Tools in Tissue** 

DOI: 10.5772/intechopen.70122

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Regenerative medicine represents a new interdisciplinary field of clinical science focused on the development and implementation of novel strategies to enhance the process of regenera‐ tion of impaired cells, tissues and organs as well as replacing damaged cells with new, fully

**Engineering and Biomedical Applications**

**Engineering and Biomedical Applications**

Małgorzata Sekuła and Ewa K. Zuba‐Surma

Additional information is available at the end of the chapter

Małgorzata Sekuła and Ewa K. Zuba-Surma

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.70122

**Abstract**

and SCs populations.

functional cells of the required phenotype [1, 2].

**1. Introduction**


**Provisional chapter**

## **Biomaterials and Stem Cells: Promising Tools in Tissue Engineering and Biomedical Applications Engineering and Biomedical Applications**

**Biomaterials and Stem Cells: Promising Tools in Tissue** 

DOI: 10.5772/intechopen.70122

Małgorzata Sekuła and Ewa K. Zuba‐Surma Małgorzata Sekuła and Ewa K. Zuba-Surma Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.70122

#### **Abstract**

[64] Nair K, Gandhi M, Khalil S, Yan KC, Marcolongo M, Barbee K, et al. Characterization of cell viability during bioprinting processes. Biotechnology Journal. 2009;**4**(8):1168-1177

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Biomaterials II; 2017;**5**:236-266

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Biofabrication. 2016 2;**9**(1):012001

com/press-release/3d-bioprinting-market)

regenerative medicine. Biomaterials. 2007;**28**(34):5068-5073

Biomaterial sciences and tissue engineering approaches are currently fundamental strat‐ egies for the development of regenerative medicine. Stem cells (SCs) are a unique cell type capable of self‐renewal and reconstructing damaged tissues. At the present time, adult SCs isolated from postnatal tissues are widely used in clinical applications. Their characteristics such as a multipotent differentiation capacity and immunomodulatory activity make them a promising tool to use in patients. Modern material technologies allow for the development of innovative biomaterials that closely correspond to require‐ ments of the current biomedical application. Biomaterials, such as ceramics and metals, are already used as implants to replace or improve the functionality of the damaged tissue or organ. However, the continuous development of modern technology opens new insights of polymeric and smart material applications. Moreover, biomaterials may enhance the SCs biological activity and their implementation by establishing a specific microenvironment mimicking natural cell niche. Thus, the synergistic advancement in the fields of biomaterial and medical sciences constitutes a challenge for the development of effective therapies in humans including combined applications of novel biomaterials and SCs populations.

**Keywords:** adult stem cells, biomaterials, regenerative medicine, tissue engineering

## **1. Introduction**

Regenerative medicine represents a new interdisciplinary field of clinical science focused on the development and implementation of novel strategies to enhance the process of regenera‐ tion of impaired cells, tissues and organs as well as replacing damaged cells with new, fully functional cells of the required phenotype [1, 2].

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

To improve the effectiveness of such regeneration processes, one of the potential approaches is application of stem cell (SC)‐based therapy. In addition, the combination of stem cells with biocompatible materials that may constitute a scaffold for the seeded cells may lead to enforcement of biological activity of stem cells and as such accelerate the process of regenera‐ tion or restoration of impaired tissue [3, 4].

• Unipotent SCs (progenitors)—The cells capable to differentiate into one or two particular types of specialized cell present in particular tissue type; this group includes several popu‐ lation of tissue‐committed progenitors such as endothelial progenitor cells, cardiac SCs,

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363

Recently, more attention has been directed to potential utilization of SCs in clinical applica‐ tions in patients. Due to the legal and ethical restrictions, the employment of embryonic SCs, which possess the largest spectrum of differentiation capacity, is controversial and prohib‐ ited in many countries. Moreover, SCs with pluripotent characteristics may lead to adverse side effects following their injection including teratoma formation [19, 20]. Therefore, the therapies employing adult SCs play a major role in human treatment as safe and effective approaches. Transplantations of autologous SCs isolated from mobilized peripheral blood or bone marrow are currently widely used in haematological patients with malignancies such as leukaemia and lymphoma [21]. Moreover, haematopoietic stem cells (HSCs) residing pre‐ dominantly in adult bone marrow are widely used for bone marrow reconstitution in patients suffering from several genetic and autoimmune diseases, blood cancers and haematopoietic

Current growing expectations for further advancements in regenerative medicine are highly focused on mesenchymal stem/stromal cells (MSCs) belonging to adult SCs isolated from sev‐ eral tissue types [16, 18, 22, 23]. MSCs are multipotent, non‐haematopoietic cells, which can be isolated from various sources including bone marrow, adipose tissue, cord blood, umbilical cord, Wharton's jelly and other tissues of the adult organism [22, 24]. Isolation of this type of cells does not raise any ethical concerns and is a relatively easy procedure. Moreover, MSCs are characterized by low immunogenicity with simultaneous immunomodulatory effect, and after transplantation, teratoma formation does not occur in the recipient organism [16, 18, 22, 23, 25]. Furthermore, potential regenerative applicability of MSCs is also enhanced by their paracrine activity related to several molecules released to their environment that may impact on other neighbouring cells affecting their functions [22, 25]. MSCs produce and release bioac‐ tive molecules, including multitude growth factors (e.g. TGF‐β1, bFGF, BMP‐4), anti‐inflam‐ matory factors (e.g. IL‐10, PGE2, HGF) and cytoprotective agents (e.g. IL‐6, MCP‐1, IGF‐1), which promote resident cells to divide and remodel the damaged tissue [26]. All these listed

MSCs possess a robust proliferation capacity as well as a potential to differentiate into several lineages of mesodermal origin, including bone, cartilage and adipose tissue [23]. Moreover, they have been also shown to give rise to other cell types, such as endothelial, cardiac or liver cells, which may also be utilized in tissue regeneration [16, 27]. These unique biological values may be utilized for the development of personalized treatment strategy for several diseases and provide the progress in establishing modern cell‐based therapy. Cell‐based therapy rep‐ resents a promising perspective of treatment directed at the regeneration of damaged tissues or organs using stem cells or progenitor cells both in the autologous and allogeneic system

satellite cells of skeletal muscles, neural progenitors and others [5, 9, 10].

**2.2. Types of SCs with potential clinical application**

features make MSCs as promising tool for biomedical research.

defects [5, 7].

[16, 28–30].

## **2. Therapeutic applications of stem cells**

#### **2.1. Classification and characteristics of stem cells (SCs)**

Stem cells (SCs) are a unique type of cells characterized by the ability to (i) self‐renewal through unlimited cell divisions and (ii) differentiate into other types of specialized cells, including epithelial, muscle, neuronal cells and others [5, 6].

Based on the origin and source of isolation, SCs may be included in two main groups: (i) embryonic SCs (ESCs)—derived from embryos at different stages of development and (ii) adult SCs (ASCs)—isolated from several postnatal and adult tissue sources, including the umbilical cord, cord blood, bone marrow, adipose tissue, central nervous system, retina, skel‐ etal muscle and other mature tissues [7, 8]. The differentiation capacity of embryonic SCs allows them to form any individual organs and fully differentiated cells of the whole body, which corresponds to pluripotency of these SCs. In opposite, most of adult SCs are multi‐ potent and lineage‐restricted (monopotent) and generally give rise to certain cell types of one germ layer or cell lineage. They residue in several niches, including bone marrow, liver, muscle, brain and others, where they may be activate towards tissue‐ or organ‐specific cells under certain physiologic or experimental conditions. Moreover, it has been shown that adult SCs may provide efficient regeneration of impaired organs in both preclinical and clinical conditions [7].

Based on the differentiation capacity, SCs populations belong to the following types [5, 9–11]:


• Unipotent SCs (progenitors)—The cells capable to differentiate into one or two particular types of specialized cell present in particular tissue type; this group includes several popu‐ lation of tissue‐committed progenitors such as endothelial progenitor cells, cardiac SCs, satellite cells of skeletal muscles, neural progenitors and others [5, 9, 10].

#### **2.2. Types of SCs with potential clinical application**

To improve the effectiveness of such regeneration processes, one of the potential approaches is application of stem cell (SC)‐based therapy. In addition, the combination of stem cells with biocompatible materials that may constitute a scaffold for the seeded cells may lead to enforcement of biological activity of stem cells and as such accelerate the process of regenera‐

Stem cells (SCs) are a unique type of cells characterized by the ability to (i) self‐renewal through unlimited cell divisions and (ii) differentiate into other types of specialized cells,

Based on the origin and source of isolation, SCs may be included in two main groups: (i) embryonic SCs (ESCs)—derived from embryos at different stages of development and (ii) adult SCs (ASCs)—isolated from several postnatal and adult tissue sources, including the umbilical cord, cord blood, bone marrow, adipose tissue, central nervous system, retina, skel‐ etal muscle and other mature tissues [7, 8]. The differentiation capacity of embryonic SCs allows them to form any individual organs and fully differentiated cells of the whole body, which corresponds to pluripotency of these SCs. In opposite, most of adult SCs are multi‐ potent and lineage‐restricted (monopotent) and generally give rise to certain cell types of one germ layer or cell lineage. They residue in several niches, including bone marrow, liver, muscle, brain and others, where they may be activate towards tissue‐ or organ‐specific cells under certain physiologic or experimental conditions. Moreover, it has been shown that adult SCs may provide efficient regeneration of impaired organs in both preclinical and clinical

Based on the differentiation capacity, SCs populations belong to the following types [5, 9–11]: • Totipotent SCs (TSCs)—The most developmentally primitive and potent SCs are capable to differentiate into any cell type from three germ layers (mesoderm, ectoderm and endo‐ derm) forming whole organism as well as into extra‐embryonic tissues such as placenta;

• Pluripotent SCs (PSCs)—The cells sustaining the capacity to differentiate into all cell types from three germ layers, but they are not able to give rise to placenta; PSCs are naturally present in developing embryo in stage of morula, in inner cell mass (ICM) of developing blastocyst and in the epiblast of gastrula and in limited number may also be found in adult tissues as remnants from embryonic development [9, 14]; PSCs may also be de novo created via genetic reprogramming of somatic cells and are called 'induced PSCs' (iPS cells) [11, 15]. • Multipotent SCs—The SCs typically capable to give rise to all cell types within one germ layer; the best described examples are mesenchymal SCs (MSCs) isolated from several

the best examples of TSCs are zygote and first blastomeres [10–13].

adult and postnatal tissues [10, 16–18];

tion or restoration of impaired tissue [3, 4].

362 Biomaterials in Regenerative Medicine

conditions [7].

**2. Therapeutic applications of stem cells**

**2.1. Classification and characteristics of stem cells (SCs)**

including epithelial, muscle, neuronal cells and others [5, 6].

Recently, more attention has been directed to potential utilization of SCs in clinical applica‐ tions in patients. Due to the legal and ethical restrictions, the employment of embryonic SCs, which possess the largest spectrum of differentiation capacity, is controversial and prohib‐ ited in many countries. Moreover, SCs with pluripotent characteristics may lead to adverse side effects following their injection including teratoma formation [19, 20]. Therefore, the therapies employing adult SCs play a major role in human treatment as safe and effective approaches. Transplantations of autologous SCs isolated from mobilized peripheral blood or bone marrow are currently widely used in haematological patients with malignancies such as leukaemia and lymphoma [21]. Moreover, haematopoietic stem cells (HSCs) residing pre‐ dominantly in adult bone marrow are widely used for bone marrow reconstitution in patients suffering from several genetic and autoimmune diseases, blood cancers and haematopoietic defects [5, 7].

Current growing expectations for further advancements in regenerative medicine are highly focused on mesenchymal stem/stromal cells (MSCs) belonging to adult SCs isolated from sev‐ eral tissue types [16, 18, 22, 23]. MSCs are multipotent, non‐haematopoietic cells, which can be isolated from various sources including bone marrow, adipose tissue, cord blood, umbilical cord, Wharton's jelly and other tissues of the adult organism [22, 24]. Isolation of this type of cells does not raise any ethical concerns and is a relatively easy procedure. Moreover, MSCs are characterized by low immunogenicity with simultaneous immunomodulatory effect, and after transplantation, teratoma formation does not occur in the recipient organism [16, 18, 22, 23, 25]. Furthermore, potential regenerative applicability of MSCs is also enhanced by their paracrine activity related to several molecules released to their environment that may impact on other neighbouring cells affecting their functions [22, 25]. MSCs produce and release bioac‐ tive molecules, including multitude growth factors (e.g. TGF‐β1, bFGF, BMP‐4), anti‐inflam‐ matory factors (e.g. IL‐10, PGE2, HGF) and cytoprotective agents (e.g. IL‐6, MCP‐1, IGF‐1), which promote resident cells to divide and remodel the damaged tissue [26]. All these listed features make MSCs as promising tool for biomedical research.

MSCs possess a robust proliferation capacity as well as a potential to differentiate into several lineages of mesodermal origin, including bone, cartilage and adipose tissue [23]. Moreover, they have been also shown to give rise to other cell types, such as endothelial, cardiac or liver cells, which may also be utilized in tissue regeneration [16, 27]. These unique biological values may be utilized for the development of personalized treatment strategy for several diseases and provide the progress in establishing modern cell‐based therapy. Cell‐based therapy rep‐ resents a promising perspective of treatment directed at the regeneration of damaged tissues or organs using stem cells or progenitor cells both in the autologous and allogeneic system [16, 28–30].

According to the current U.S. National Institutes of Health database including clinical trials conducted worldwide, there are currently more than 240 clinical trials being conducted in the world employing MSCs in patients [31]. Examples of the application of MSCs isolated from different sources in the treatment of selected diseases are shown in **Table 1**.

**2.3. Utilization of SC derivatives as a potential alternative to cell‐based therapy**

exosomes (30–100 nm) fractions [36, 37].

cell‐based therapy.

including in heart repair [38, 40].

**3.1. Material requirements for biomaterials**

Immunomodulatory properties of SCs are important features involved in tissue repair, which are directly related to their paracrine activity. Despite the directly released molecules, mam‐ malian cells, including SCs, are able to produce extracellular vesicles (EVs) carrying bioac‐ tive factors, which may additionally be involved in the modulation of the repair process of damaged tissues [35]. EVs represent heterogeneous population of small, circular struc‐ tures surrounded with the protein‐lipid membrane that are released by cells including SCs. Importantly, the size and molecular composition of EVs are different and unique depending on the cell type of origin and the mechanism of their biogenesis. Depending on the size of EVs, they may be distinguished in apoptotic bodies (1–5 µm), microparticles (100 nm–1 µm) and

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Several recent scientific reports indicate that EVs express surface markers characterizing the cells from which they are released, along with EV‐specific antigens including tetraspanins (CD9, CD63 and CD81), endosome or membrane‐binding proteins (TG101), signal transduc‐ tion or scaffolding proteins (syntenin) [36, 37]. Importantly, EVs may also include various types of bioactive components (e.g. mRNA, miRNA and enzymes), as well as receptors, adhe‐ sion or signalling proteins [38, 39]. Importantly, the contents of EVs can be effectively trans‐ ferred to the target cells, change their function and impact in the regeneration of impaired tissues. Moreover, the presence of protein‐lipid membrane on the surface of EVs can protect their bioactive content from extracellular enzymes and therefore the cargo may be delivered in a fully functional form into targeted cells [38, 40]. Thus, EVs are recognized as mediators of intercellular communication and constitute an alternative or reinforcement of a standard

The biological relevance of EVs has been established in different experimental settings. Depending on the origin and content of EVs, they may enhance immune system, endorse anti‐ tumour responses and thus may provide important tools for novel anti‐tumour therapies, such as melanoma treatment [41]. EVs may also be utilized as drug delivery vehicles [42], in regenerative medicine [43] and immune therapy [44]. Recently, our study also indicated that SC‐derived EVs may be utilized as a novel tool for regenerative therapies of ischemic tissue

However, further studies are required for comprehensive analysis of the mechanisms of EVs

Biomaterial by definition is a 'substance (other than a drug), synthetic or natural, that can be used as a system or part of a system that treats, augments, or replaces any tissue, organ, or function of the body' [45]. Thus, according to the definition biomaterials are progressively

**3. Medical application of selected natural and synthetic biomaterials**

action and potential clinical applications of these promising SC derivatives.

One of the great new opportunities in medical science is the possibility of obtaining the induced pluripotent SCs (iPS cells) by genetic reprogramming of mature cells into the stage of pluripotency [15]. Due to the discovery of this phenomenon, professor Shinya Yamanaka was honoured with the Nobel Prize in medicine and physiology in 2012. Since then, iPS cells consti‐ tute an excellent model for *in vitro* studies of molecular mechanisms associated with the devel‐ opment and progression of several diseases, including Parkinson's disease [32], Huntington disease [33], Down syndrome [34] and others. Moreover, dozens of laboratories are questing for optimal utilization of these cells in tissue regeneration. However, due to the possibility of teratoma formation after iPS transplantation, their applications in medicine are still limited.

Thus, despite the fact that several new rising SCs types are being examined and optimized for future applications, the most commonly applicable SCs in cell therapies of distinct human dis‐ eases are adult stem cells including predominantly MSCs derived from bone marrow, adipose tissue and umbilical cord as well as HSCs harvested from bone marrow, mobilized peripheral blood and cord blood [16, 18, 28–30].


Source: Ref. [31].

**Table 1.** The application of MSCs in selected clinical trials.

#### **2.3. Utilization of SC derivatives as a potential alternative to cell‐based therapy**

According to the current U.S. National Institutes of Health database including clinical trials conducted worldwide, there are currently more than 240 clinical trials being conducted in the world employing MSCs in patients [31]. Examples of the application of MSCs isolated from

One of the great new opportunities in medical science is the possibility of obtaining the induced pluripotent SCs (iPS cells) by genetic reprogramming of mature cells into the stage of pluripotency [15]. Due to the discovery of this phenomenon, professor Shinya Yamanaka was honoured with the Nobel Prize in medicine and physiology in 2012. Since then, iPS cells consti‐ tute an excellent model for *in vitro* studies of molecular mechanisms associated with the devel‐ opment and progression of several diseases, including Parkinson's disease [32], Huntington disease [33], Down syndrome [34] and others. Moreover, dozens of laboratories are questing for optimal utilization of these cells in tissue regeneration. However, due to the possibility of teratoma formation after iPS transplantation, their applications in medicine are still limited.

Thus, despite the fact that several new rising SCs types are being examined and optimized for future applications, the most commonly applicable SCs in cell therapies of distinct human dis‐ eases are adult stem cells including predominantly MSCs derived from bone marrow, adipose tissue and umbilical cord as well as HSCs harvested from bone marrow, mobilized peripheral

**Type of MSCs Condition ClinicalTrials.gov identifier**

Aplastic anaemia NCT03055078 Stroke NCT02580019 Pneumoconiosis NCT02668068 Rheumatoid arthritis NCT02643823 Sweat gland diseases NCT02304562

Chronic myocardial ischaemia NCT02460770 Acute respiratory distress syndrome NCT02097641 Middle cerebral artery infarction NCT01461720 Prostate cancer NCT01983709 Stroke NCT02564328

Multiple sclerosis NCT02326935 Hair restoration NCT02865421

Umbilical cord‐derived MSCs Hepatic cirrhosis NCT02652351

Bone marrow‐derived MSCs Acute myocardial infarction NCT01652209

Adipose‐derived MSCs Infantile spinal muscular atrophy NCT02855112

different sources in the treatment of selected diseases are shown in **Table 1**.

blood and cord blood [16, 18, 28–30].

364 Biomaterials in Regenerative Medicine

Source: Ref. [31].

**Table 1.** The application of MSCs in selected clinical trials.

Immunomodulatory properties of SCs are important features involved in tissue repair, which are directly related to their paracrine activity. Despite the directly released molecules, mam‐ malian cells, including SCs, are able to produce extracellular vesicles (EVs) carrying bioac‐ tive factors, which may additionally be involved in the modulation of the repair process of damaged tissues [35]. EVs represent heterogeneous population of small, circular struc‐ tures surrounded with the protein‐lipid membrane that are released by cells including SCs. Importantly, the size and molecular composition of EVs are different and unique depending on the cell type of origin and the mechanism of their biogenesis. Depending on the size of EVs, they may be distinguished in apoptotic bodies (1–5 µm), microparticles (100 nm–1 µm) and exosomes (30–100 nm) fractions [36, 37].

Several recent scientific reports indicate that EVs express surface markers characterizing the cells from which they are released, along with EV‐specific antigens including tetraspanins (CD9, CD63 and CD81), endosome or membrane‐binding proteins (TG101), signal transduc‐ tion or scaffolding proteins (syntenin) [36, 37]. Importantly, EVs may also include various types of bioactive components (e.g. mRNA, miRNA and enzymes), as well as receptors, adhe‐ sion or signalling proteins [38, 39]. Importantly, the contents of EVs can be effectively trans‐ ferred to the target cells, change their function and impact in the regeneration of impaired tissues. Moreover, the presence of protein‐lipid membrane on the surface of EVs can protect their bioactive content from extracellular enzymes and therefore the cargo may be delivered in a fully functional form into targeted cells [38, 40]. Thus, EVs are recognized as mediators of intercellular communication and constitute an alternative or reinforcement of a standard cell‐based therapy.

The biological relevance of EVs has been established in different experimental settings. Depending on the origin and content of EVs, they may enhance immune system, endorse anti‐ tumour responses and thus may provide important tools for novel anti‐tumour therapies, such as melanoma treatment [41]. EVs may also be utilized as drug delivery vehicles [42], in regenerative medicine [43] and immune therapy [44]. Recently, our study also indicated that SC‐derived EVs may be utilized as a novel tool for regenerative therapies of ischemic tissue including in heart repair [38, 40].

However, further studies are required for comprehensive analysis of the mechanisms of EVs action and potential clinical applications of these promising SC derivatives.

## **3. Medical application of selected natural and synthetic biomaterials**

#### **3.1. Material requirements for biomaterials**

Biomaterial by definition is a 'substance (other than a drug), synthetic or natural, that can be used as a system or part of a system that treats, augments, or replaces any tissue, organ, or function of the body' [45]. Thus, according to the definition biomaterials are progressively used in tissue engineering. They may be utilized for the construction of implants to replace lost or damaged organs or tissues and may also constitute a scaffold for enhanced stem cells to reconstruct not fully functional tissue [45, 46].

cytotoxicity [47, 48]. Moreover, they may exhibit specific protein‐binding sites that improve integration with cells after transplantation [48]. Thus, they are considered predominantly

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367

One of the most common natural biomaterials is collagen—an important component of con‐ nective tissue, including bones, tendons, ligaments and skin [46, 50]. Collagen is simply absorbed into the body, is non‐toxic and exhibits a low immune response and as such is a perfect biocompatible material with an adequate mechanical strength and flexibility for sev‐ eral applications. Moreover, collagen enhances cell adhesion to such surface, stimulates also biological interactions between cells and facilitates restoration of the natural microenviron‐ ment of cell niche and thereby may support the reconstruction of several damaged tissues

Collagen may be employed for tissue engineering in the form of sponges, gels, hydrogels and sheets. It may also be chemically crosslinked in order to enhance or alter the rate of degradation of the fibres [51]. Currently, collagen preparations are used predominantly in wound healing and cartilage regeneration. Injectable form of collagen is used for cosmetic and aesthetic medicine as a tissue filler. In addition, collagen‐based membranes are used in the periodontal treatment as a barrier preventing the migration of epithelial cells. It also forms a favourable microenvironment for stem cells to facilitate reconstruction of the damaged area

Synthetic materials are considered as an alternative to natural materials. Due to their defined chemical composition and the ability to control the mechanical and physical properties, they are extensively used in therapeutic applications and basic biological studies [48, 52–55].

Due to distinct variants of polymerization reaction and formation of co‐polymers, multiple synthetic polymers with wide range of physical and chemical properties may be achieved in chemical laboratories. Moreover, novel technologies in the synthesis and formation of more complex structures allow for the production of advanced composites [54]. Synthetic polymers, such as poly(ethylene) (PE), polyurethanes (PUR), polylactides (PLA) and poly(glycolide) (PGA), are widely employed as implants and components of medical devices [56]. Moreover, polymers may constitute suitable scaffold for cell propagation and enhance their biological activity, including neural stem cells, retinal progenitor cells or smooth muscle cells [55, 57, 58]. Thus, this group of biomaterials is currently in a special focus of scientists working on com‐ bined approaches using biocompatible scaffolds and stem cells for tissue repair [55, 57, 58]. Biodegradable polymers, including polyhydroxycarboxylic acids, such as PGA, PLA, poly(3‐ hydroxybutyrate), poly(4‐hydroxybutyrate) and poly(∈‐caprolactone) (PCL) are of wide interest in the development of novel technologies [56]. One of their potential applications is utilization in the treatment of cardiovascular diseases. Our recent studies have shown the positive impact of both PCL and PLA scaffolds on proliferation, migration and proangiogenic potential of mesenchymal SCs derived from umbilical cord tissue *in vitro*, suggesting the pos‐ sible applications of these materials in cardiovascular repair *in vivo* (unpublished data) [59].

interesting for tissue engineering applications.

[46, 48, 50].

[50, 51].

*3.2.2. Synthetic biomaterials*

Due to the wide range of potential applications of biomaterials in regenerative medicine, their physical and chemical properties may be different [45, 47]. However, in order to use a bioma‐ terial in medical application, it should follow relevant requirements such as biocompatibility and biofunctionality [45, 47]:


Other important properties of biomaterials, which are affecting the potential application in medicine, include [45, 48, 49] the following:


#### **3.2. Applications of biomaterials**

Several biomaterials useful for distinct applications in medical sciences, including in tissue repair and organ reconstruction, have already been developed over the last few decades [45, 47]. The biomaterial sciences are currently one of the highly advancing fields, which also closely cooperate with biotechnological and medical studies. Recent advancement in regen‐ erative medicine strongly requires such strong support from biomaterial sciences, which may provide novel solutions for tissue repair [4, 49].

Among the biomaterials recognized and developed for potential medical purposes, here are multitude materials commonly present in natural sources or *de novo* designed and created for such purposes.

#### *3.2.1. Naturally derived biomaterials*

Natural materials commonly present in nature such as agarose, collagen, alginate, chitosan, hyaluronate or fibrin fully cooperate with living tissues of the recipient and possess low cytotoxicity [47, 48]. Moreover, they may exhibit specific protein‐binding sites that improve integration with cells after transplantation [48]. Thus, they are considered predominantly interesting for tissue engineering applications.

One of the most common natural biomaterials is collagen—an important component of con‐ nective tissue, including bones, tendons, ligaments and skin [46, 50]. Collagen is simply absorbed into the body, is non‐toxic and exhibits a low immune response and as such is a perfect biocompatible material with an adequate mechanical strength and flexibility for sev‐ eral applications. Moreover, collagen enhances cell adhesion to such surface, stimulates also biological interactions between cells and facilitates restoration of the natural microenviron‐ ment of cell niche and thereby may support the reconstruction of several damaged tissues [46, 48, 50].

Collagen may be employed for tissue engineering in the form of sponges, gels, hydrogels and sheets. It may also be chemically crosslinked in order to enhance or alter the rate of degradation of the fibres [51]. Currently, collagen preparations are used predominantly in wound healing and cartilage regeneration. Injectable form of collagen is used for cosmetic and aesthetic medicine as a tissue filler. In addition, collagen‐based membranes are used in the periodontal treatment as a barrier preventing the migration of epithelial cells. It also forms a favourable microenvironment for stem cells to facilitate reconstruction of the damaged area [50, 51].

#### *3.2.2. Synthetic biomaterials*

used in tissue engineering. They may be utilized for the construction of implants to replace lost or damaged organs or tissues and may also constitute a scaffold for enhanced stem cells

Due to the wide range of potential applications of biomaterials in regenerative medicine, their physical and chemical properties may be different [45, 47]. However, in order to use a bioma‐ terial in medical application, it should follow relevant requirements such as biocompatibility

• Biocompatibility is the ability to integrate with the recipient's cells in a safe manner and

• Biofunctionality is the ability to perform a specific biological function, based on the rel‐

Other important properties of biomaterials, which are affecting the potential application in

• Biodegradation—Decomposition of the material in a natural way, when degradation prod‐

• Bioresorbability—Decomposition of the material in a natural way at a certain period of time after implantation. Non‐toxic‐degraded products are removed from the body *via* met‐

• Non‐toxicity—From the surface or porous of the material does not elute any toxic compo‐ nents, such as surfactants, stabilizers, catalysts, pigments and UV absorbents, which were

• Mechanical properties—Biomaterial should possess particular mechanical properties con‐

Several biomaterials useful for distinct applications in medical sciences, including in tissue repair and organ reconstruction, have already been developed over the last few decades [45, 47]. The biomaterial sciences are currently one of the highly advancing fields, which also closely cooperate with biotechnological and medical studies. Recent advancement in regen‐ erative medicine strongly requires such strong support from biomaterial sciences, which may

Among the biomaterials recognized and developed for potential medical purposes, here are multitude materials commonly present in natural sources or *de novo* designed and created for

Natural materials commonly present in nature such as agarose, collagen, alginate, chitosan, hyaluronate or fibrin fully cooperate with living tissues of the recipient and possess low

used during production and that are incompatible with living organisms.

evant parameters of the physical and mechanical properties.

ucts remain in the human body but without adverse side effects.

sistent with the anatomical site into which it will be implanted.

abolic pathways (hydrolytic or enzymatic degradation).

to reconstruct not fully functional tissue [45, 46].

and biofunctionality [45, 47]:

366 Biomaterials in Regenerative Medicine

without adverse side effects.

**3.2. Applications of biomaterials**

*3.2.1. Naturally derived biomaterials*

such purposes.

provide novel solutions for tissue repair [4, 49].

medicine, include [45, 48, 49] the following:

Synthetic materials are considered as an alternative to natural materials. Due to their defined chemical composition and the ability to control the mechanical and physical properties, they are extensively used in therapeutic applications and basic biological studies [48, 52–55].

Due to distinct variants of polymerization reaction and formation of co‐polymers, multiple synthetic polymers with wide range of physical and chemical properties may be achieved in chemical laboratories. Moreover, novel technologies in the synthesis and formation of more complex structures allow for the production of advanced composites [54]. Synthetic polymers, such as poly(ethylene) (PE), polyurethanes (PUR), polylactides (PLA) and poly(glycolide) (PGA), are widely employed as implants and components of medical devices [56]. Moreover, polymers may constitute suitable scaffold for cell propagation and enhance their biological activity, including neural stem cells, retinal progenitor cells or smooth muscle cells [55, 57, 58]. Thus, this group of biomaterials is currently in a special focus of scientists working on com‐ bined approaches using biocompatible scaffolds and stem cells for tissue repair [55, 57, 58].

Biodegradable polymers, including polyhydroxycarboxylic acids, such as PGA, PLA, poly(3‐ hydroxybutyrate), poly(4‐hydroxybutyrate) and poly(∈‐caprolactone) (PCL) are of wide interest in the development of novel technologies [56]. One of their potential applications is utilization in the treatment of cardiovascular diseases. Our recent studies have shown the positive impact of both PCL and PLA scaffolds on proliferation, migration and proangiogenic potential of mesenchymal SCs derived from umbilical cord tissue *in vitro*, suggesting the pos‐ sible applications of these materials in cardiovascular repair *in vivo* (unpublished data) [59].

Synthetic polymers may also be used in biodegradable stents implanted after a heart attack and greatly contribute to patient recovery [56]. Importantly, the material should have suitable decomposition kinetics. Too long decomposition time (i.e. in the case of PLA or PGA) may lead to late stent thrombosis or blockages [56, 60]. One of a possible solution of this problem is to use rapidly biodegradable polymer stents coated with SCs to help rebuild damaged tissue and additionally stimulate resident cells to grow.

Hydrogels demonstrate transparency and bioadhesive properties and they are widely used in the pharmaceutical and dermatological industries by local administration or filling the defects caused by injury [77]. They may also be utilized as an injectable material for bone and cartilage tissue engineering, which may be combined with appropriate cell injection [53, 78, 79]. It has been shown that *in situ* implementation of hydrogels promotes osteoblast differen‐ tiation [53, 79]. Therefore, injectable therapy constitutes a promising approach for non‐inva‐ sive technique of transplantation, where also cell‐based component may be added to enhance

Biomaterials and Stem Cells: Promising Tools in Tissue Engineering and Biomedical Applications

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369

Smart materials represent a new generation of biomaterials, exceeding the functionality of the currently widely used construction materials. Smart materials are characterized by the ability to alter their physical characteristics in a controlled manner including changing the shape, colour, stiffness or stickiness in response to several external stimuli, such as temperature, hydrostatic pressure, electric and magnetic field or radiation [80]. These changes are related to the revealing or eliciting the new functionality of the material and may be utilized in biomedi‐ cal applications. Through the common connection between the internal sensor, the activator and a specific control mechanism, smart materials are able to respond to external stimuli. Importantly, these mechanisms are also responsible for the return to the original state, when

• **Colour changing materials**—Materials that change colour in a reversible manner, depend‐ ing on electrical, optical or thermal changes. These types of materials are exploited, for example, in optoelectronic components, lenses, lithium batteries, ferroelectric memory,

• **Light‐emitting materials**—Materials emitting visible or invisible light, as a result of exter‐ nal stimuli such as short wavelength radiation (e.g. X‐rays, ultraviolet light), temperature and electric voltage. They are utilized in electronics, filters for glasses, devices that detect UV rays, in criminology and in geology to identify minerals and rocks. They may also be exploited as a component of protective clothing, safety elements and warning materials

• **Shape memory materials**—Metal alloys that change shape as a result of temperature increase or decrease, respectively, to the set value. The reversibility of the process is to return to its original shape by changing the temperature or under the influence of the applied motion (the effect of pseudoelasticity). These materials are used in temperature sensors, electronics, robot‐ ics, telecommunications and production of medical devices (micro‐pump, surgical clamps, orthodontic wire, long‐ and short‐term implants, suture tightening on a stiffen wound, ortho‐

• **Self‐assembling materials**—Materials that exhibit the intrinsic ability to spontaneously connect individual elements into an ordered 2D or 3D structure. In addition, they can also

Smart materials include several types such as listed below [52, 80, 82–84]:

temperature sensors or as the indicators of battery consumption [80, 81].

paedic devices, bone nails, clamps, surgical instruments and others) [83, 84].

tissue repair.

[80, 81].

*3.2.4. Smart materials*

a stimulant disappeared [80, 81].

Other types of common synthetic materials useful for biomedical applications are ceramics. It has been well described that ceramic scaffolds, such as, for example, hydroxyapatite (HA) and tri‐calcium phosphate (TCP), are characterized by biocompatibility, high mechanical stiffness (Young's modulus), very low elasticity and a hard brittle surface [49]. Due to their chemical and structural similarity to the mineral phase of native bone, these materials may enhance osteoblast proliferation and therefore they are widely utilized in bone regeneration [61, 62]. Moreover, ceramics may be exploited in dental and orthopaedic procedures to fill bone defects or as a bioactive coating material for implants to increase their integration after transplantation [63, 64]. However, their clinical applications are still limited due to the dif‐ ficulties with the ability to change the shape of the material dedicated for transplantation and controlling time of their degradation rate [49, 65].

Similarly, titanium (Ti)‐based metallic materials have been widely optimized for bone repair due to their mechanical properties and resistance to corrosion following the transplantation [66–68]. It has been shown that titanium scaffolds are effectively colonized by osteoblasts responsible for bone formation and this process may be enhanced *via* additional modifica‐ tions of the scaffold surface by its roughening, coating with HA or graphene oxide (GO), as well as its biofunctionalization with bioactive molecules such as heparin and bone morpho‐ genetic protein 2 (BMP‐2) [69–72].

Importantly, graphene in its different forms is currently being considered as a potential new promising material for biomedical applications including tissue repair [73, 74]. This 2D carbon biocompatible material exhibits great electrical, conductive and physical properties, which make it interesting for potential applications for drug delivery and scaffold coating in regen‐ erative therapies [74, 75]. It has been shown that graphene may enhance osteogenic differenti‐ ation of SCs [72, 73]. Moreover, our recent data also suggest the beneficial impact of graphene oxide (GO) on proliferative capacity, viability and differentiation potential of umbilical cord tissue‐derived MSCs, which confirms the possibility of future graphene employment in tissue repair [76].

#### *3.2.3. Hydrogels*

Hydrogels are frequently used biomaterials in the biomedical applications and represent sys‐ tems consisting of two or more compartments comprising a three‐dimensional (3D) network of polymer chains and water that fills the spaces between the macromolecules [77, 78]. The main characteristics of hydrogels include the biocompatibility and ability to swell in solution until they reach a state of equilibrium. These allow them to be injected into the body in a non‐invasive manner [77, 78].

Hydrogels demonstrate transparency and bioadhesive properties and they are widely used in the pharmaceutical and dermatological industries by local administration or filling the defects caused by injury [77]. They may also be utilized as an injectable material for bone and cartilage tissue engineering, which may be combined with appropriate cell injection [53, 78, 79]. It has been shown that *in situ* implementation of hydrogels promotes osteoblast differen‐ tiation [53, 79]. Therefore, injectable therapy constitutes a promising approach for non‐inva‐ sive technique of transplantation, where also cell‐based component may be added to enhance tissue repair.

#### *3.2.4. Smart materials*

Synthetic polymers may also be used in biodegradable stents implanted after a heart attack and greatly contribute to patient recovery [56]. Importantly, the material should have suitable decomposition kinetics. Too long decomposition time (i.e. in the case of PLA or PGA) may lead to late stent thrombosis or blockages [56, 60]. One of a possible solution of this problem is to use rapidly biodegradable polymer stents coated with SCs to help rebuild damaged tissue

Other types of common synthetic materials useful for biomedical applications are ceramics. It has been well described that ceramic scaffolds, such as, for example, hydroxyapatite (HA) and tri‐calcium phosphate (TCP), are characterized by biocompatibility, high mechanical stiffness (Young's modulus), very low elasticity and a hard brittle surface [49]. Due to their chemical and structural similarity to the mineral phase of native bone, these materials may enhance osteoblast proliferation and therefore they are widely utilized in bone regeneration [61, 62]. Moreover, ceramics may be exploited in dental and orthopaedic procedures to fill bone defects or as a bioactive coating material for implants to increase their integration after transplantation [63, 64]. However, their clinical applications are still limited due to the dif‐ ficulties with the ability to change the shape of the material dedicated for transplantation and

Similarly, titanium (Ti)‐based metallic materials have been widely optimized for bone repair due to their mechanical properties and resistance to corrosion following the transplantation [66–68]. It has been shown that titanium scaffolds are effectively colonized by osteoblasts responsible for bone formation and this process may be enhanced *via* additional modifica‐ tions of the scaffold surface by its roughening, coating with HA or graphene oxide (GO), as well as its biofunctionalization with bioactive molecules such as heparin and bone morpho‐

Importantly, graphene in its different forms is currently being considered as a potential new promising material for biomedical applications including tissue repair [73, 74]. This 2D carbon biocompatible material exhibits great electrical, conductive and physical properties, which make it interesting for potential applications for drug delivery and scaffold coating in regen‐ erative therapies [74, 75]. It has been shown that graphene may enhance osteogenic differenti‐ ation of SCs [72, 73]. Moreover, our recent data also suggest the beneficial impact of graphene oxide (GO) on proliferative capacity, viability and differentiation potential of umbilical cord tissue‐derived MSCs, which confirms the possibility of future graphene employment in tissue

Hydrogels are frequently used biomaterials in the biomedical applications and represent sys‐ tems consisting of two or more compartments comprising a three‐dimensional (3D) network of polymer chains and water that fills the spaces between the macromolecules [77, 78]. The main characteristics of hydrogels include the biocompatibility and ability to swell in solution until they reach a state of equilibrium. These allow them to be injected into the body in a

and additionally stimulate resident cells to grow.

368 Biomaterials in Regenerative Medicine

controlling time of their degradation rate [49, 65].

genetic protein 2 (BMP‐2) [69–72].

repair [76].

*3.2.3. Hydrogels*

non‐invasive manner [77, 78].

Smart materials represent a new generation of biomaterials, exceeding the functionality of the currently widely used construction materials. Smart materials are characterized by the ability to alter their physical characteristics in a controlled manner including changing the shape, colour, stiffness or stickiness in response to several external stimuli, such as temperature, hydrostatic pressure, electric and magnetic field or radiation [80]. These changes are related to the revealing or eliciting the new functionality of the material and may be utilized in biomedi‐ cal applications. Through the common connection between the internal sensor, the activator and a specific control mechanism, smart materials are able to respond to external stimuli. Importantly, these mechanisms are also responsible for the return to the original state, when a stimulant disappeared [80, 81].

Smart materials include several types such as listed below [52, 80, 82–84]:


have the ability to bind metal atoms, ions, molecules or semiconductors. They are widely used in biological research and nanotechnology, that is, in the tissue regeneration, as com‐ ponents for the storage of drugs, crystal engineering, as artificial proteins with pH‐sensitive structure, as semi‐permeable membrane as well as for the production of electronic proces‐ sors and displays [52, 82].

such as biocompatibility, mechanical stability, permeability, appropriate size, strength and durability [47]. One of the most common encapsulation materials is alginate. Due to the fact that the procedure for cell encapsulation using alginate can be performed under physiologi‐ cal conditions (physiological temperature and pH) and using isotonic solutions, it is widely distributed through clinical and industrial applications. Moreover, this natural biodegradable polymer that mimics the extracellular matrix and promotes cell functions and metabolism has been established in cartilage regenerative approaches [86, 87]. Microencapsulation technol‐ ogy represents a novel cell culture system that allows maintaining cell viability and differen‐ tiation of interested cell lines. It also may support the extracellular matrix production and cell

Biomaterials and Stem Cells: Promising Tools in Tissue Engineering and Biomedical Applications

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371

Significant advancement of regenerative medicine, nanomedicine and biomaterials engi‐ neering offers extended possibilities to obtain novel, effective achievements, which may be utilized in biomedical applications. The effect of interdisciplinary activity resulted in the development of bioactive scaffolds that promote cell propagation and enhance their biologi‐ cal activity. However, some difficulties in biomaterial‐ and cell‐based therapy are still unclear and need to be addressed for widespread investigations. Nevertheless, integrative research in biomaterials and medicine fields is a challenge to develop effective therapies for cancer,

This work is supported by grants from the National Science Centre (NCN): SONATA BIS‐3 (UMO‐2013/10/E/NZ3/007500), SYMFONIA 3 (UMO‐2015/16/W/NZ4/00071) and the National Centre for Research and Development (NCBR): STRATEGMED III (BioMiStem project; ID 303570) to EZS. The Faculty of Biochemistry, Biophysics and Biotechnology at the Jagiellonian University, Krakow, Poland, is a partner of the Leading National Research Center (KNOW)

\*

1 Malopolska Centre of Biotechnology, Jagiellonian University, Krakow, Poland

2 Department of Cell Biology, Faculty of Biochemistry, Biophysics and Biotechnology,

civilization diseases and provide further development of tissue engineering.

supported by the Ministry of Science and Higher Education.

and Ewa K. Zuba‐Surma2

\*Address all correspondence to: ewa.zuba‐surma@uj.edu.pl

Jagiellonian University, Krakow, Poland

organization in reconstructed tissue [86].

**5. Conclusions**

**Acknowledgements**

**Author details**

Małgorzata Sekuła1

• **Self‐repairing materials**—Structural damage of this type of material is automatically and autonomously recovered by inducing a change in the shape or the self‐assembly of the mol‐ ecules. This process is not a method of complete repair of the impaired material; however, it may be used in the military, automotive, aviation and electronics industries [52, 82].

## **4. Novel aspects of the application of stem cells and biomaterials in tissue engineering and regenerative medicine**

#### **4.1. Biomaterials approaches for enhancement of SCs‐based therapy**

Modern approaches in current regenerative medicine include developing biocompatible scaf‐ folds and combining them with living cell of selected type and bioactive molecules, in order to enhance the regeneration process of damaged tissues and organs [47].

Growing evidence indicate different populations of stem cells as a promising tool that may be utilized in tissue engineering and repair. Importantly, despite the regenerative prop‐ erties of SCs, the restoration processes in damaged tissue are long and may not often be fully effective for functional recovery of damaged tissue. On the other hand, appropriate stimulation of reparative capacity of SCs may be achieved by modulation of chemical and physical properties of optimized biomaterials [47, 70, 77]. Therefore, simultaneous applica‐ tion of optimized and well‐combined SCs and biomaterials may open new perspectives for the synergistic effective cooperation of both such components to improve the efficiency of the regeneration process [77]. Biomaterials may enhance the biological activity of SCs by establishing a specific niche related to their native microenvironment. This type of cell‐bio‐ material interactions leads to stimulation of cell adhesion, proliferation and directed differ‐ entiation of the cells implemented at the injured site [47, 70, 77]. Therefore, therapy based on biomaterials and SCs opens new possibilities for the development of innovative medicine [47, 77].

Currently, growing evidence is focused on encapsulation of native SCs prior to their trans‐ plantation [47, 85]. Cells encapsulation technique is based on the immobilization of cells in a semi‐permeable membrane, which protects cells against mechanical damage and immune system response. Notably, the construction of the microcapsules allows bidirectional diffu‐ sion of nutrients, oxygen and wastes and therefore provides appropriate conditions for cell development [47, 85].

Encapsulated cells may be subjected to transplantation and directed differentiation. The material used to construct the microcapsules should possess particular physical properties, such as biocompatibility, mechanical stability, permeability, appropriate size, strength and durability [47]. One of the most common encapsulation materials is alginate. Due to the fact that the procedure for cell encapsulation using alginate can be performed under physiologi‐ cal conditions (physiological temperature and pH) and using isotonic solutions, it is widely distributed through clinical and industrial applications. Moreover, this natural biodegradable polymer that mimics the extracellular matrix and promotes cell functions and metabolism has been established in cartilage regenerative approaches [86, 87]. Microencapsulation technol‐ ogy represents a novel cell culture system that allows maintaining cell viability and differen‐ tiation of interested cell lines. It also may support the extracellular matrix production and cell organization in reconstructed tissue [86].

## **5. Conclusions**

have the ability to bind metal atoms, ions, molecules or semiconductors. They are widely used in biological research and nanotechnology, that is, in the tissue regeneration, as com‐ ponents for the storage of drugs, crystal engineering, as artificial proteins with pH‐sensitive structure, as semi‐permeable membrane as well as for the production of electronic proces‐

• **Self‐repairing materials**—Structural damage of this type of material is automatically and autonomously recovered by inducing a change in the shape or the self‐assembly of the mol‐ ecules. This process is not a method of complete repair of the impaired material; however, it may be used in the military, automotive, aviation and electronics industries [52, 82].

**4. Novel aspects of the application of stem cells and biomaterials in** 

Modern approaches in current regenerative medicine include developing biocompatible scaf‐ folds and combining them with living cell of selected type and bioactive molecules, in order

Growing evidence indicate different populations of stem cells as a promising tool that may be utilized in tissue engineering and repair. Importantly, despite the regenerative prop‐ erties of SCs, the restoration processes in damaged tissue are long and may not often be fully effective for functional recovery of damaged tissue. On the other hand, appropriate stimulation of reparative capacity of SCs may be achieved by modulation of chemical and physical properties of optimized biomaterials [47, 70, 77]. Therefore, simultaneous applica‐ tion of optimized and well‐combined SCs and biomaterials may open new perspectives for the synergistic effective cooperation of both such components to improve the efficiency of the regeneration process [77]. Biomaterials may enhance the biological activity of SCs by establishing a specific niche related to their native microenvironment. This type of cell‐bio‐ material interactions leads to stimulation of cell adhesion, proliferation and directed differ‐ entiation of the cells implemented at the injured site [47, 70, 77]. Therefore, therapy based on biomaterials and SCs opens new possibilities for the development of innovative medicine

Currently, growing evidence is focused on encapsulation of native SCs prior to their trans‐ plantation [47, 85]. Cells encapsulation technique is based on the immobilization of cells in a semi‐permeable membrane, which protects cells against mechanical damage and immune system response. Notably, the construction of the microcapsules allows bidirectional diffu‐ sion of nutrients, oxygen and wastes and therefore provides appropriate conditions for cell

Encapsulated cells may be subjected to transplantation and directed differentiation. The material used to construct the microcapsules should possess particular physical properties,

**tissue engineering and regenerative medicine**

**4.1. Biomaterials approaches for enhancement of SCs‐based therapy**

to enhance the regeneration process of damaged tissues and organs [47].

sors and displays [52, 82].

370 Biomaterials in Regenerative Medicine

[47, 77].

development [47, 85].

Significant advancement of regenerative medicine, nanomedicine and biomaterials engi‐ neering offers extended possibilities to obtain novel, effective achievements, which may be utilized in biomedical applications. The effect of interdisciplinary activity resulted in the development of bioactive scaffolds that promote cell propagation and enhance their biologi‐ cal activity. However, some difficulties in biomaterial‐ and cell‐based therapy are still unclear and need to be addressed for widespread investigations. Nevertheless, integrative research in biomaterials and medicine fields is a challenge to develop effective therapies for cancer, civilization diseases and provide further development of tissue engineering.

## **Acknowledgements**

This work is supported by grants from the National Science Centre (NCN): SONATA BIS‐3 (UMO‐2013/10/E/NZ3/007500), SYMFONIA 3 (UMO‐2015/16/W/NZ4/00071) and the National Centre for Research and Development (NCBR): STRATEGMED III (BioMiStem project; ID 303570) to EZS. The Faculty of Biochemistry, Biophysics and Biotechnology at the Jagiellonian University, Krakow, Poland, is a partner of the Leading National Research Center (KNOW) supported by the Ministry of Science and Higher Education.

## **Author details**

Małgorzata Sekuła1 and Ewa K. Zuba‐Surma2 \*

\*Address all correspondence to: ewa.zuba‐surma@uj.edu.pl

1 Malopolska Centre of Biotechnology, Jagiellonian University, Krakow, Poland

2 Department of Cell Biology, Faculty of Biochemistry, Biophysics and Biotechnology, Jagiellonian University, Krakow, Poland

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**Chapter 17**

**Provisional chapter**

**Identification of Fe3O4 Nanoparticles Biomedical**

DOI: 10.5772/intechopen.69717

The application of magnetic nanoparticles for biomedical research is an interdisciplinary problem. The use of nano‐ and microsized powder materials as developed technology for obtaining bionanomaterials with magnetocatalytic properties has been investigated. Control over immobilization can be carried by means of magnetic properties. Synthesis of superparamagnetic nanoparticles is developed not only for the benefit of fundamental science, but also for many technologies, such as technologies of magnetic storage media, magnetic ink for printers, but mainly for biosensors and medical applications. All the biomedical applications require that the nanoparticles have high enough levels of satu‐ ration of magnetization; their size should be less than 100 nm with a small deviation in size. Appropriate coating of the surface of magnetic nanoparticles should be nontoxic, biocompatible with the target of bioorganic compound. The techniques of measurement of magnetic nanoparticle properties by means of vibrational magnetometers, as well as by means of a set of smart sensor devices in accordance with new concept of Internet of Things (IoTh), were described. The first method is based on vibrating sample magnetom‐ eter technique. The second method is based on direct measurement of three dimensions

**O4 Nanoparticles Biomedical** 

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

Use of nanomaterials has become one of the innovation research directions in materials sci‐ ence, biochemistry and medicine. Small sizes of nanoparticles (NPs) lead to the appearance

**Keywords:** magnetic nanoparticles, drug transportation, magnetic field sensor

**Purpose by Magnetometric Methods**

**Purpose by Magnetometric Methods**

Zoia Duriagina, Roman Holyaka, Tetiana Tepla, Volodymyr Kulyk, Peter Arras and Elena Eyngorn

Additional information is available at the end of the chapter

(3D) of nanoparticles' magnetic field components.

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69717

**Abstract**

**1. Introduction**

Zoia Duriagina, Roman Holyaka, Tetiana Tepla, Volodymyr Kulyk, Peter Arras and Elena Eyngorn

**Identification of Fe3**


**Provisional chapter**

## **Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods Purpose by Magnetometric Methods**

DOI: 10.5772/intechopen.69717

**O4 Nanoparticles Biomedical** 

Zoia Duriagina, Roman Holyaka, Tetiana Tepla, Volodymyr Kulyk, Peter Arras and Elena Eyngorn Tetiana Tepla, Volodymyr Kulyk, Peter Arras and Elena Eyngorn

Additional information is available at the end of the chapter Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/intechopen.69717

Zoia Duriagina, Roman Holyaka,

**Identification of Fe3**

#### **Abstract**

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B: Engineering. 2016;**87**:92‐119. DOI: 10.1016/j.compositesb.2015.09.057

DOI: 10.1155/2015/948040

378 Biomaterials in Regenerative Medicine

ISBN‐13: 978‐1‐4613‐7449‐7

2013;**54**:24‐29. DOI: 10.1016/j.ijbiomac.2012.11.026

1527‐2648(200203)4:3<91::AID‐ADEM91>3.0.CO;2‐B

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& Interfaces. 2016;**8**:10070‐10087. DOI: 10.1021/acsami.6b01295

Cytotechnology. 2008;**58**:49‐56. DOI: 10.1007/s10616‐008‐9161‐0

The application of magnetic nanoparticles for biomedical research is an interdisciplinary problem. The use of nano‐ and microsized powder materials as developed technology for obtaining bionanomaterials with magnetocatalytic properties has been investigated. Control over immobilization can be carried by means of magnetic properties. Synthesis of superparamagnetic nanoparticles is developed not only for the benefit of fundamental science, but also for many technologies, such as technologies of magnetic storage media, magnetic ink for printers, but mainly for biosensors and medical applications. All the biomedical applications require that the nanoparticles have high enough levels of satu‐ ration of magnetization; their size should be less than 100 nm with a small deviation in size. Appropriate coating of the surface of magnetic nanoparticles should be nontoxic, biocompatible with the target of bioorganic compound. The techniques of measurement of magnetic nanoparticle properties by means of vibrational magnetometers, as well as by means of a set of smart sensor devices in accordance with new concept of Internet of Things (IoTh), were described. The first method is based on vibrating sample magnetom‐ eter technique. The second method is based on direct measurement of three dimensions (3D) of nanoparticles' magnetic field components.

**Keywords:** magnetic nanoparticles, drug transportation, magnetic field sensor

## **1. Introduction**

Use of nanomaterials has become one of the innovation research directions in materials sci‐ ence, biochemistry and medicine. Small sizes of nanoparticles (NPs) lead to the appearance

of new unique functional properties. The methods of such material preparation are improv‐ ing every year and become more accessible. The technology of material structure "design" by combining different types of materials, starting from metals, their compounds (oxides, nitrides, borides and hydrides) and ending with organic and inorganic polymers [1], has recently been widely used. Nanomaterials based on metals operate in power generating, opti‐ cal industry and biomedicine. The unique superparamagnetic, optical and electrochemical properties are accentuated first of all. Use of magnetic nanoparticles (MNPs) for biomedi‐ cal researches is an interdisciplinary problem and to solve it experts in different areas are required—medical workers, biologists, specialists in the field of materials science and electro‐ dynamics. Application of nanomaterials with coating for biomedical purposes foresees such functional properties of the given system as adsorption, adhesion strength, biocompatibil‐ ity, and certain magnetic and optoelectronic properties that will identify the stability of the obtained structure‐energy state. The task of the materials science researchers is to develop complicated nanoobjects with optimal properties and morphological structure and to create the system of monitoring (Internet of Things (IoTh)) by their life activity cycle. The perfect biomedical systems can be considered such that the stable functional properties are possessed.

In particular, magnetism of nanoparticles is an important information carrier and what is especially valuable can be implemented through industrial means. This action at a distance in combination with a typical magnetic field penetration into human body tissues opens many new applications, including transport and purposeful delivery of biomagnets to a cor‐ responding biological object [2]. Magnetism can be exhibited in a greater number of new nanomaterials. It is known that, under transition to nanosizes in metals and their compounds, new specific characteristics appear. Thus, magnetic properties of bulk gold and platinum are nonmagnetic, but at the nanosize they are magnetic. Surface atoms are not only different to bulk atoms, but they can also be modified by interaction with other chemical species, that is, by capping the nanoparticles. This phenomenon opens the possibility to modify the physi‐ cal properties of the nanoparticles by capping them with appropriate molecules. Actually, it should be possible that nonferromagnetic bulk materials exhibit ferromagnetic‐like behavior when prepared in a nanorange. One can obtain magnetic nanoparticles of Pd, Pt and the surprising case of Au (that is diamagnetic in bulk) from nonmagnetic bulk materials. In the case of Pt and Pd we can obtain the ferromagnetism. However, gold nanoparticles become ferromagnetic when they are capped with appropriate molecules: the charge localized at the particle surface gives rise to ferromagnetic‐like behavior. This observation suggested that modification of the d band structure by chemical bonding can induce ferromagnetic‐like char‐ acter in metallic clusters [3].

perspective in case the clearly defined and checked interaction between magnetic nanopar‐

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

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381

It is known that living organisms are built of cells with usual sizes of about 10 μm. The objects from which a cell consists of are much less than 1 μm. It should be considered that the syn‐ thesized nanoparticles are of sizes 2–450 nm, commensurable with the sizes of intercellular biological objects, viruses, proteins and genes. Thus, the nanoparticles by the sizes and mass occupy the intermediate place between single molecule and living cells. Still the main advan‐ tage of these materials is their ability to perform the preset functions under the effect of exter‐

This occurs due to the formation of potential barriers at their boundary that limit the charge carrier movement in different directions and give the electron processes a mainly quantum character with dominating role of the interface. It should be noted that in this case a non‐ linear dependence of equilibrium concentration of defects at the interface is observed that increases the dependence of properties on nanostructure sizes [10]. This is especially impor‐ tant when delivering drugs and means of diagnostics. Depending on the sizes, other proper‐ ties of nanoparticles as toxicity, adsorption ability and magnetism also change. In particular, when the size of nanoparticles is less than 10 nm, they pass into a superparamagnetic state that, when adsorbing the energy of the external high‐frequency electromagnetic field, pro‐ motes conversion of the energy state into the thermal one. This enables a tumor heating to a

ticles and living cells is provided.

**Figure 1.** Use of magnetic nanoparticles in medicine.

nal magnetic field (**Figure 2**).

temperature of 43°С and thus its destruction [11, 12].

#### **2. Use of magnetic nanoparticles in medicine**

Magnetic biomaterials are widely used in medicine, especially in cardiology, neurosurgery, oncology, radiology and cellular biology for cell separation, to perform immunological analysis, magnetic resonance spectroscopy, information preservation and so on [1–8]. They are also used as the X‐ray contrast means and magnetosensitive composites for drugs and genes delivery, in radionuclide therapy and hyperthermia (**Figure 1**). Such applications are Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods http://dx.doi.org/10.5772/intechopen.69717 381

**Figure 1.** Use of magnetic nanoparticles in medicine.

of new unique functional properties. The methods of such material preparation are improv‐ ing every year and become more accessible. The technology of material structure "design" by combining different types of materials, starting from metals, their compounds (oxides, nitrides, borides and hydrides) and ending with organic and inorganic polymers [1], has recently been widely used. Nanomaterials based on metals operate in power generating, opti‐ cal industry and biomedicine. The unique superparamagnetic, optical and electrochemical properties are accentuated first of all. Use of magnetic nanoparticles (MNPs) for biomedi‐ cal researches is an interdisciplinary problem and to solve it experts in different areas are required—medical workers, biologists, specialists in the field of materials science and electro‐ dynamics. Application of nanomaterials with coating for biomedical purposes foresees such functional properties of the given system as adsorption, adhesion strength, biocompatibil‐ ity, and certain magnetic and optoelectronic properties that will identify the stability of the obtained structure‐energy state. The task of the materials science researchers is to develop complicated nanoobjects with optimal properties and morphological structure and to create the system of monitoring (Internet of Things (IoTh)) by their life activity cycle. The perfect biomedical systems can be considered such that the stable functional properties are possessed. In particular, magnetism of nanoparticles is an important information carrier and what is especially valuable can be implemented through industrial means. This action at a distance in combination with a typical magnetic field penetration into human body tissues opens many new applications, including transport and purposeful delivery of biomagnets to a cor‐ responding biological object [2]. Magnetism can be exhibited in a greater number of new nanomaterials. It is known that, under transition to nanosizes in metals and their compounds, new specific characteristics appear. Thus, magnetic properties of bulk gold and platinum are nonmagnetic, but at the nanosize they are magnetic. Surface atoms are not only different to bulk atoms, but they can also be modified by interaction with other chemical species, that is, by capping the nanoparticles. This phenomenon opens the possibility to modify the physi‐ cal properties of the nanoparticles by capping them with appropriate molecules. Actually, it should be possible that nonferromagnetic bulk materials exhibit ferromagnetic‐like behavior when prepared in a nanorange. One can obtain magnetic nanoparticles of Pd, Pt and the surprising case of Au (that is diamagnetic in bulk) from nonmagnetic bulk materials. In the case of Pt and Pd we can obtain the ferromagnetism. However, gold nanoparticles become ferromagnetic when they are capped with appropriate molecules: the charge localized at the particle surface gives rise to ferromagnetic‐like behavior. This observation suggested that modification of the d band structure by chemical bonding can induce ferromagnetic‐like char‐

acter in metallic clusters [3].

380 Biomaterials in Regenerative Medicine

**2. Use of magnetic nanoparticles in medicine**

Magnetic biomaterials are widely used in medicine, especially in cardiology, neurosurgery, oncology, radiology and cellular biology for cell separation, to perform immunological analysis, magnetic resonance spectroscopy, information preservation and so on [1–8]. They are also used as the X‐ray contrast means and magnetosensitive composites for drugs and genes delivery, in radionuclide therapy and hyperthermia (**Figure 1**). Such applications are perspective in case the clearly defined and checked interaction between magnetic nanopar‐ ticles and living cells is provided.

It is known that living organisms are built of cells with usual sizes of about 10 μm. The objects from which a cell consists of are much less than 1 μm. It should be considered that the syn‐ thesized nanoparticles are of sizes 2–450 nm, commensurable with the sizes of intercellular biological objects, viruses, proteins and genes. Thus, the nanoparticles by the sizes and mass occupy the intermediate place between single molecule and living cells. Still the main advan‐ tage of these materials is their ability to perform the preset functions under the effect of exter‐ nal magnetic field (**Figure 2**).

This occurs due to the formation of potential barriers at their boundary that limit the charge carrier movement in different directions and give the electron processes a mainly quantum character with dominating role of the interface. It should be noted that in this case a non‐ linear dependence of equilibrium concentration of defects at the interface is observed that increases the dependence of properties on nanostructure sizes [10]. This is especially impor‐ tant when delivering drugs and means of diagnostics. Depending on the sizes, other proper‐ ties of nanoparticles as toxicity, adsorption ability and magnetism also change. In particular, when the size of nanoparticles is less than 10 nm, they pass into a superparamagnetic state that, when adsorbing the energy of the external high‐frequency electromagnetic field, pro‐ motes conversion of the energy state into the thermal one. This enables a tumor heating to a temperature of 43°С and thus its destruction [11, 12].

**Figure 2.** Several classes of NP used for the design of smart micro/nanocarriers, including micelles, dendrimers, nanocages, CNT, polymeric conjugates, nanogels, magnetic NP, quantum dots (QD), nanospheres and gold (Au) NP [9].

form magnetite or maghemite and a shell from dextran or silicon. Colloid solutions based on biomolecule‐sorbed magnetic iron oxides, surrounded by a polymer layer, can be used for interaction with biological targets and direct them into a tumor with the aim of treatment or diagnostics. In this case, the controlled movement of magnetic particles in a body occurs because of the use of magnetic field gradient, regulated such that particles penetrate directly

**Figure 3.** Graphic illustration of the structure of a multifunctional/multimodality MNP and different types of magnetic

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

http://dx.doi.org/10.5772/intechopen.69717

383

Magnetic nanoparticle‐based drugs can be divided into two groups: for external use (oint‐ ments, applicators, etc.) and for internal use (liquids, suspensions and magnetic containers). To improve physicochemical and medical biological properties when delivering drugs, the sizes of magnetic material particles in the blood stream should not exceed 100 nm. Under such conditions, they can be also used for hypoferritic anemia treatment. Today, the drugs obtained by adding precious metals to iron are paid more and more attention to. They are used as anti‐inflammatory, antibacterial, anticancer and anesthetic compounds of the oint‐

Magnetic carriers of medical purpose should be biocompatible with a human organism, non‐ toxic and nonallergenic. To improve biocompatibility, magnetic carriers are coated with chi‐ tosan, dextran, starch, carbon, gelatin, polymer starch coatings and so on (**Figure 4**). It should be noted that silicon oxide also increases biocompatibility of nanoparticles, and in this case,

conditions can be regulated by changing the size and porosity of the obtained systems [4].

nation with bioselective elements—ferments have been widely used. This is important for developing technologies of manufacturing bionanomaterials with catalytic properties. High

adsorption and formation of surface ligaments with other adsorbed particles. This guarantees

Recently, nano‐ and microsize powder materials are based on Fe3

particles as in microcapsule. Release rate of drugs in these

particles is caused by their higher ability to ion or atomic exchange,

O4

, especially in combi‐

into the affected area [4].

cores.

ments for external use.

chemical activity of Fe3

iron is localized inside the SiO2

O4

It has recently been proven that the effects of heat are more cytotoxic for cancer cells than they are for surrounding normal cells. Because of this, researchers have focused on hyper‐ thermia as a method to selectively treat cancer cells. In Gilchrist's invaluable study, the cancer cells were heated locally using magnetic NP with a 1.2‐MHz magnetic field. Since then, other observations have demonstrated that magnetic‐induced hyperthermia in animals that have been injected directly with MNP can produce tumor regression through the application of magnetic fields to the solid tumors [9].

For therapeutic aims, magnetic nanoparticles are rarely used in a pure form. Usually, they are encapsulated or situated in bioinert matrices made of different organic compounds (**Figure 3**). This allows decreasing of the possible toxic effect of magnetic phase and simul‐ taneous increasing of its stability due to immobilization on the surface of such capsules or matrices of medical aids. Encapsulation is usually carried out in suspensions of ultradisper‐ sive ferromagnetic or superparamagnetic particles containing stabilizing reagents—so called magnetic liquids [2, 5].

The structure of a multifunctional/multimodality MNP with a magnetic core, a polymeric coating and targeting ligands extended from the surface of MNP with the aid of polymeric spacers. Therapeutic payloads (drugs and genes) and imaging reporters (fluorophores and radionuclides) can be either embedded in the coating, or conjugated on the surface [5].

Magnetic contrast agents using which an image of a human body is obtained with a nuclear magnetic resonance tomography are widely known. Particles of these agents consist of a core

**Figure 3.** Graphic illustration of the structure of a multifunctional/multimodality MNP and different types of magnetic cores.

form magnetite or maghemite and a shell from dextran or silicon. Colloid solutions based on biomolecule‐sorbed magnetic iron oxides, surrounded by a polymer layer, can be used for interaction with biological targets and direct them into a tumor with the aim of treatment or diagnostics. In this case, the controlled movement of magnetic particles in a body occurs because of the use of magnetic field gradient, regulated such that particles penetrate directly into the affected area [4].

It has recently been proven that the effects of heat are more cytotoxic for cancer cells than they are for surrounding normal cells. Because of this, researchers have focused on hyper‐ thermia as a method to selectively treat cancer cells. In Gilchrist's invaluable study, the cancer cells were heated locally using magnetic NP with a 1.2‐MHz magnetic field. Since then, other observations have demonstrated that magnetic‐induced hyperthermia in animals that have been injected directly with MNP can produce tumor regression through the application of

**Figure 2.** Several classes of NP used for the design of smart micro/nanocarriers, including micelles, dendrimers, nanocages, CNT, polymeric conjugates, nanogels, magnetic NP, quantum dots (QD), nanospheres and gold (Au) NP [9].

For therapeutic aims, magnetic nanoparticles are rarely used in a pure form. Usually, they are encapsulated or situated in bioinert matrices made of different organic compounds (**Figure 3**). This allows decreasing of the possible toxic effect of magnetic phase and simul‐ taneous increasing of its stability due to immobilization on the surface of such capsules or matrices of medical aids. Encapsulation is usually carried out in suspensions of ultradisper‐ sive ferromagnetic or superparamagnetic particles containing stabilizing reagents—so called

The structure of a multifunctional/multimodality MNP with a magnetic core, a polymeric coating and targeting ligands extended from the surface of MNP with the aid of polymeric spacers. Therapeutic payloads (drugs and genes) and imaging reporters (fluorophores and radionuclides) can be either embedded in the coating, or conjugated on the surface [5].

Magnetic contrast agents using which an image of a human body is obtained with a nuclear magnetic resonance tomography are widely known. Particles of these agents consist of a core

magnetic fields to the solid tumors [9].

magnetic liquids [2, 5].

382 Biomaterials in Regenerative Medicine

Magnetic nanoparticle‐based drugs can be divided into two groups: for external use (oint‐ ments, applicators, etc.) and for internal use (liquids, suspensions and magnetic containers). To improve physicochemical and medical biological properties when delivering drugs, the sizes of magnetic material particles in the blood stream should not exceed 100 nm. Under such conditions, they can be also used for hypoferritic anemia treatment. Today, the drugs obtained by adding precious metals to iron are paid more and more attention to. They are used as anti‐inflammatory, antibacterial, anticancer and anesthetic compounds of the oint‐ ments for external use.

Magnetic carriers of medical purpose should be biocompatible with a human organism, non‐ toxic and nonallergenic. To improve biocompatibility, magnetic carriers are coated with chi‐ tosan, dextran, starch, carbon, gelatin, polymer starch coatings and so on (**Figure 4**). It should be noted that silicon oxide also increases biocompatibility of nanoparticles, and in this case, iron is localized inside the SiO2 particles as in microcapsule. Release rate of drugs in these conditions can be regulated by changing the size and porosity of the obtained systems [4].

Recently, nano‐ and microsize powder materials are based on Fe3 O4 , especially in combi‐ nation with bioselective elements—ferments have been widely used. This is important for developing technologies of manufacturing bionanomaterials with catalytic properties. High chemical activity of Fe3 O4 particles is caused by their higher ability to ion or atomic exchange, adsorption and formation of surface ligaments with other adsorbed particles. This guarantees

• Precipitation in the conditions of oxidizing hydrolysis of iron sulfate Fe2+ in acid

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

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385

• Precipitate deposition from the solutions of iron chlorides Fe2+ and Fe3+ under alkali hydro‐

When modifying the conditions of magnetite core synthesis and forming the core shell, it is possible to obtain magnetic nanocomposites that differ not only by size but also by the ratio of "core–shell" sizes, magnetic susceptibility and complexity of surface microrelief that will

However, nanoparticle synthesis has remained until now a complicated task. This is related with the difficulty of formation of homodispersive population of checked‐size magnetic par‐ ticles. It should be noted that there is no still a complete understanding of the mechanisms of core formation and growth. Besides the technology of obtaining crystalline nanoparticles with high saturation, magnetization needs further improvement. Moreover, nanoparticles lose their stability with time. This occurs due to the decrease of their free surface energy as a result of agglomeration. It is necessary to develop strategies of the surface coating preparation to simplify the process and effectively prevent agglomeration and segmentation of superpara‐ magnetic particles. As a result, a stable solution for injections or frozen‐dried (lyophilized)

be aimed at establishing the interrelations between their structure and pharmacokinetics. Nature of the coating on the iron oxide surface will determine not only the size of colloids but also will affect the pharmacokinetics and metabolic properties. This will enable modeling of their capture by reticuloendothelial system (RES) and facilitate diffusion to the tumor tissue. Processes of modification of magnetic nanoparticle surface, which are used for connection of biovectors, must be also improved. This is decisive in optimization of superparamagnetic

Iron oxides (in particular magnetite) are among the most investigated materials in human

the sake of fundamental science and technological applications, as for example magnetic data carriers, for biosensor, medical applications and magnetic inks [19–21]. Superparamagnetic nanoparticles of iron oxide with corresponding surface microgeometry can be used for increasing the image contrast, restoration of tissues, detoxification of biological fluids, hyper‐

Detailed characterization of magnetite NPs is therefore necessary in order to obtain an accurate relationship between their electronic, magnetic and structural properties. Nominally claimed magnetite NPs are often (and to various extents) composed of nonstoichiometric oxide phases, and their instability in air ultimately causes oxidation to maghemite. The oxidization rate of

O4

O4

nanoparticles, should

is developing for

• Precipitation in the conditions of alkali hydrolysis of iron chlorides Fe2+ and Fe3+.

environment.

lysis of urea.

specify the absorption properties.

powder, that is easily dissolved, can be obtained [17].

nanoparticles likeness with biological objects [18].

 **magnetite**

thermia, directed delivery of drugs and cell separation.

history. Synthesis of superparamagnetic nanoparticles of iron oxide Fe3

**O4**

**2.2. Properties of Fe3**

Investigation of superparamagnetic nanoparticles, especially Fe3

**Figure 4.** Schematic representation of the proposed magnetic sensitive carrier and triggered drug release mechanism [9].

the creation of bioparticles (immobilized ferments on the particle surface) with their further use in biosensorics and in enzymatic reactions. It is known that powder particles of Fe3 O4 retain magnetization even if no external magnetic field is present, possessing own magnetic moment. However, remaining magnetization has a negative effect—a tendency of nanopar‐ ticles to agglomeration. Therefore, for the effective use of such magnetic particles, their chemi‐ cal stability should be ensured by applying the corresponding coating on their surface.

#### **2.1. Methods of obtaining magnetic nanoparticles**

Physical and chemical properties of magnetic nanoparticles are determined by their structure and method of preparation. In most cases, these are particles of size from 1 to 100 nm with clearly expressed superparamagnetism. The most widely used methods of magnetic nanopar‐ ticle synthesis are as follows:


To synthesize magnetite‐based nanocomposites, the following methodical approaches are mainly used [13–16].


When modifying the conditions of magnetite core synthesis and forming the core shell, it is possible to obtain magnetic nanocomposites that differ not only by size but also by the ratio of "core–shell" sizes, magnetic susceptibility and complexity of surface microrelief that will specify the absorption properties.

However, nanoparticle synthesis has remained until now a complicated task. This is related with the difficulty of formation of homodispersive population of checked‐size magnetic par‐ ticles. It should be noted that there is no still a complete understanding of the mechanisms of core formation and growth. Besides the technology of obtaining crystalline nanoparticles with high saturation, magnetization needs further improvement. Moreover, nanoparticles lose their stability with time. This occurs due to the decrease of their free surface energy as a result of agglomeration. It is necessary to develop strategies of the surface coating preparation to simplify the process and effectively prevent agglomeration and segmentation of superpara‐ magnetic particles. As a result, a stable solution for injections or frozen‐dried (lyophilized) powder, that is easily dissolved, can be obtained [17].

Investigation of superparamagnetic nanoparticles, especially Fe3 O4 nanoparticles, should be aimed at establishing the interrelations between their structure and pharmacokinetics. Nature of the coating on the iron oxide surface will determine not only the size of colloids but also will affect the pharmacokinetics and metabolic properties. This will enable modeling of their capture by reticuloendothelial system (RES) and facilitate diffusion to the tumor tissue. Processes of modification of magnetic nanoparticle surface, which are used for connection of biovectors, must be also improved. This is decisive in optimization of superparamagnetic nanoparticles likeness with biological objects [18].

#### **2.2. Properties of Fe3 O4 magnetite**

the creation of bioparticles (immobilized ferments on the particle surface) with their further use in biosensorics and in enzymatic reactions. It is known that powder particles of Fe3

**Figure 4.** Schematic representation of the proposed magnetic sensitive carrier and triggered drug release mechanism [9].

retain magnetization even if no external magnetic field is present, possessing own magnetic moment. However, remaining magnetization has a negative effect—a tendency of nanopar‐ ticles to agglomeration. Therefore, for the effective use of such magnetic particles, their chemi‐ cal stability should be ensured by applying the corresponding coating on their surface.

Physical and chemical properties of magnetic nanoparticles are determined by their structure and method of preparation. In most cases, these are particles of size from 1 to 100 nm with clearly expressed superparamagnetism. The most widely used methods of magnetic nanopar‐

• Deposition—synthesis of iron oxides from water solutions of their salts in inert gas atmo‐

O4

To synthesize magnetite‐based nanocomposites, the following methodical approaches are

ganic metal compounds in boiling organic solvents that contain stabilizing surface active

nanoparticles during decomposition of or‐

**2.1. Methods of obtaining magnetic nanoparticles**

sphere at ambient or elevated temperature.

• Thermal decomposition—synthesis of Fe3

ticle synthesis are as follows:

384 Biomaterials in Regenerative Medicine

• Microemulsion technology.

agents (SAA).

mainly used [13–16].

O4

Iron oxides (in particular magnetite) are among the most investigated materials in human history. Synthesis of superparamagnetic nanoparticles of iron oxide Fe3 O4 is developing for the sake of fundamental science and technological applications, as for example magnetic data carriers, for biosensor, medical applications and magnetic inks [19–21]. Superparamagnetic nanoparticles of iron oxide with corresponding surface microgeometry can be used for increasing the image contrast, restoration of tissues, detoxification of biological fluids, hyper‐ thermia, directed delivery of drugs and cell separation.

Detailed characterization of magnetite NPs is therefore necessary in order to obtain an accurate relationship between their electronic, magnetic and structural properties. Nominally claimed magnetite NPs are often (and to various extents) composed of nonstoichiometric oxide phases, and their instability in air ultimately causes oxidation to maghemite. The oxidization rate of magnetite NPs in ambient conditions is size‐dependent and can range from several months for NPs <10 nm, up to several years for NPs ∼100 nm. However, there is no general con‐ sensus regarding the complex oxidation mechanism, which presumably proceeds through a continual set of intermediate phases accompanied by cations and vacancies (C–V) reordering. Magnetite and maghemite both crystallize in the face‐centered cubic (fcc) spinel structure, whose unit cell is composed of 32 O2− ions placed at the 32e crystallographic position, and 24 Fe ions distributed over the 64 tetrahedral 8a (A) and 32 octahedral 16d (B) crystallographic positions (**Figure 5**). Magnetite and maghemite can be represented with a single formula: (Fe3+) A[(Fe2+) 1–3δ(Fe3+) 1+2δ δ ] BO4 , where stands for vacancies and 0 ≤ δ ≤1/3. In pure magnetite (δ = 0), all A sites are occupied by Fe3+ ions, while B sites are equally occupied with Fe2+ and Fe3+. In pure maghemite (δ = 1/3), all Fe ions are in 3+ state, with a tendency to regular arrangement at B sites (two occupied followed by one vacant site). Vacancies occur preferentially at octahe‐ dral sites, but they can also mix over octahedral and tetrahedral sites. The degree of vacancy ordering in maghemite decreases with a particle size, and it is believed that for NPs smaller than 20 nm vacancy ordering vanishes. Magnetite and maghemite are both ferrimagnetic (the two uneven ferromagnetic sublattices FeA and FeB are antiferromagnetically aligned), with comparable saturation magnetizations (MS = 90 emu/g for magnetite; MS = 83.5 emu/g for well‐ordered crystalline maghemite samples) and extraordinary high Curie temperatures (TC = 858 K magnetite; TC = 790–893 K maghemite, depending on the degree of C–V ordering).

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387

The important difference between magnetite and maghemite is that maghemite is an insulator with energy gap Eg ≈ 2 eV, while magnetite is half‐metal with much narrower Eg = 0.1–0.5 eV (depending on the sample quality). Furthermore, bulk magnetite undergoes so‐called Verwey order‐disorder phase transition to the insulating state at temperatures 120–125 K, accompa‐ nied with structural change from cubic to monoclinic lattice symmetry and various anomalies in the physical properties. The decisive influence on this complex transformation is ascribed to charge and orbital ordering involved in the three‐site distortions. The exact structural parameters of the low temperature (LT) crystal structure, which is thought to have at least

Things are even more elusive when the sample size is in the nanometer range. According to some recent reports for NPs with the mean size ~ 50 nm, the Verwey temperature (TV) shifts down to 20 K and it cannot be observed for smaller particles. The Verwey transition is weakly size‐dependent in magnetite NPs larger than 20 nm, slightly suppressed in NPs smaller than 20 nm, and completely vanishes for NPs smaller than 6 nm. These inconsistencies are often ascribed to the fact that final properties of magnetite NPs strongly depend also on structural

using spectrophotometer Bruker Vertex 70 with attachment attenuated total reflectance (ATR). Powder samples were dried on the microscope slide surface and irradiated by a laser beam.

from the value of 1200 сm−1, and in the range of 600–1200 cm−1 a significant background of the wave number area of the substances of glass on which the samples were dried is observed. In this range, only a signal from Si‐O (for a wave number 1053 cm−1) is clearly seen. The blind area should contain signals from C‐O (for a wave number of about 1100 cm−1) and signals of

Obtained data can be used as an introductory information for processing monitoring of

particles strongly affects its microstructure and properties. **Figure 7** presents a

micro‐ and nanoparticles.

O4

**O4**

–NP characterization was performed by infrared Fourier spectroscopy

‐NP at 600 cm−1, similarly to micro X‐ray spectral analysis, indi‐

 **nanoparticles**

–NP (**Figure 6**, **Table 1**). Peak intensity increases

four inequivalent octahedral Fe sites, are extremely difficult to determine.

**2.3. Investigation of the structure and properties of Fe3**

O4

group Si–O–C (for a wave number of about 1108 cm−1).

O4

dynamics in situ of nanoparticle modification.

O4

cated the presence of Fe‐O bonds in Fe3

order [23].

Size of Fe3

O4

difference in microstructure of Fe<sup>3</sup>

In this research, the Fe3

The intensive peak from Fe3

**Figure 5.** Left: monoclinic crystal structure of Fe3 O4 in the low‐temperature phase corresponding to a subcell of a2×a2×2a with P2/c symmetry. Right: 3D isosurfaces of unoccupied density of states of the O 2p integrated between the fermi level and 1 eV above within the corresponding Fe3 O4 cubes highlighted in the crystal structure. For simplicity, we denote the octahedral Fe site with nominal 2+ and 3+ valences as Fe2+ and Fe3+, respectively [22].

at B sites (two occupied followed by one vacant site). Vacancies occur preferentially at octahe‐ dral sites, but they can also mix over octahedral and tetrahedral sites. The degree of vacancy ordering in maghemite decreases with a particle size, and it is believed that for NPs smaller than 20 nm vacancy ordering vanishes. Magnetite and maghemite are both ferrimagnetic (the two uneven ferromagnetic sublattices FeA and FeB are antiferromagnetically aligned), with comparable saturation magnetizations (MS = 90 emu/g for magnetite; MS = 83.5 emu/g for well‐ordered crystalline maghemite samples) and extraordinary high Curie temperatures (TC = 858 K magnetite; TC = 790–893 K maghemite, depending on the degree of C–V ordering).

magnetite NPs in ambient conditions is size‐dependent and can range from several months for NPs <10 nm, up to several years for NPs ∼100 nm. However, there is no general con‐ sensus regarding the complex oxidation mechanism, which presumably proceeds through a continual set of intermediate phases accompanied by cations and vacancies (C–V) reordering. Magnetite and maghemite both crystallize in the face‐centered cubic (fcc) spinel structure, whose unit cell is composed of 32 O2− ions placed at the 32e crystallographic position, and 24 Fe ions distributed over the 64 tetrahedral 8a (A) and 32 octahedral 16d (B) crystallographic positions (**Figure 5**). Magnetite and maghemite can be represented with a single formula:

(δ = 0), all A sites are occupied by Fe3+ ions, while B sites are equally occupied with Fe2+ and Fe3+. In pure maghemite (δ = 1/3), all Fe ions are in 3+ state, with a tendency to regular arrangement

, where stands for vacancies and 0 ≤ δ ≤1/3. In pure magnetite

in the low‐temperature phase corresponding to a subcell of a2×a2×2a

cubes highlighted in the crystal structure. For simplicity, we denote the

(Fe3+)

A[(Fe2+)

1–3δ(Fe3+)

386 Biomaterials in Regenerative Medicine

**Figure 5.** Left: monoclinic crystal structure of Fe3

and 1 eV above within the corresponding Fe3

O4

O4

octahedral Fe site with nominal 2+ and 3+ valences as Fe2+ and Fe3+, respectively [22].

with P2/c symmetry. Right: 3D isosurfaces of unoccupied density of states of the O 2p integrated between the fermi level

1+2δ δ ] BO4

The important difference between magnetite and maghemite is that maghemite is an insulator with energy gap Eg ≈ 2 eV, while magnetite is half‐metal with much narrower Eg = 0.1–0.5 eV (depending on the sample quality). Furthermore, bulk magnetite undergoes so‐called Verwey order‐disorder phase transition to the insulating state at temperatures 120–125 K, accompa‐ nied with structural change from cubic to monoclinic lattice symmetry and various anomalies in the physical properties. The decisive influence on this complex transformation is ascribed to charge and orbital ordering involved in the three‐site distortions. The exact structural parameters of the low temperature (LT) crystal structure, which is thought to have at least four inequivalent octahedral Fe sites, are extremely difficult to determine.

Things are even more elusive when the sample size is in the nanometer range. According to some recent reports for NPs with the mean size ~ 50 nm, the Verwey temperature (TV) shifts down to 20 K and it cannot be observed for smaller particles. The Verwey transition is weakly size‐dependent in magnetite NPs larger than 20 nm, slightly suppressed in NPs smaller than 20 nm, and completely vanishes for NPs smaller than 6 nm. These inconsistencies are often ascribed to the fact that final properties of magnetite NPs strongly depend also on structural order [23].

#### **2.3. Investigation of the structure and properties of Fe3 O4 nanoparticles**

In this research, the Fe3 O4 –NP characterization was performed by infrared Fourier spectroscopy using spectrophotometer Bruker Vertex 70 with attachment attenuated total reflectance (ATR). Powder samples were dried on the microscope slide surface and irradiated by a laser beam.

The intensive peak from Fe3 O4 ‐NP at 600 cm−1, similarly to micro X‐ray spectral analysis, indi‐ cated the presence of Fe‐O bonds in Fe3 O4 –NP (**Figure 6**, **Table 1**). Peak intensity increases from the value of 1200 сm−1, and in the range of 600–1200 cm−1 a significant background of the wave number area of the substances of glass on which the samples were dried is observed. In this range, only a signal from Si‐O (for a wave number 1053 cm−1) is clearly seen. The blind area should contain signals from C‐O (for a wave number of about 1100 cm−1) and signals of group Si–O–C (for a wave number of about 1108 cm−1).

Obtained data can be used as an introductory information for processing monitoring of dynamics in situ of nanoparticle modification.

Size of Fe3 O4 particles strongly affects its microstructure and properties. **Figure 7** presents a difference in microstructure of Fe<sup>3</sup> O4 micro‐ and nanoparticles.

**Figure 6.** ІR‐Fourier spectra of Fe<sup>3</sup> O4 nanoparticles.

To estimate the sizes of Fe3 O4 ‐NP atomic‐force microscopy was used, which software prod‐ ucts allowed the establishment of the scanned particle structure (**Figure 8**). Investigation data prove the results of microscopic investigations on the spherical nanoparticle accumulation. Two types of structural inhomogeneity can be distinguished on the investigated sample sur‐ face: conglomerates of conical‐like nanoparticles and granular texture of substrate. Surface topography is characterized by a rough relief with morphological regions of blocked struc‐ ture. Blocks are characterized by nonisometric round form with no surface faceting. Their height above the substrate surface is in the range from 5 to 10 nm and base diameter from 30 to 50 nm. In this case, the dominant size of nanoparticles is 8 nm.

by technical characteristics of the equipment. The quantitative phase magnetic analysis uses properties of ferromagnetics, which they acquire in strong magnetic fields—in the state of technical saturation. Primary magnetic properties that are structurally insensitive are obtained from the curve of temperature dependence of saturation magnetization. Such characteristics are magnetization and Curie point. These values give the information about phase composi‐ tion of material and its changes in the process of certain thermal operations and also in the process of deformation. Curie point and saturation magnetization are called primary mag‐ netic properties because their values are determined by the nature of the ferromagnetic phase

nanoparticles; (b) Fe3

O4

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

‐NP on the surface (a) and histogram of their distribution by

microparticles. Microstructures were

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389

Changes in magnetization and Curie point of separate phases observed in the process of inves‐ tigation of a certain material are not caused by their particle growth but are conditioned by the change of chemical composition of phases and their crystalline composition only. Based on these data, it is possible to study the kinetics of phase formation at the very early stages of the process (for particles sizes up to 10−6 cm) by the growth of magnetization and Curie point to their normal values. Thus, a weak dependence of saturation magnetization and Curie point

O4

size (b). Lateral dimension of particles in nanometers is given on X‐axis, and a number of scanned particles on Y‐axis.

(crystal lattice, electron structure of atoms and chemical phase composition) [25].

O4 : (а) Fe<sup>3</sup> O4

**Figure 7.** Microstructure of synthesized Fe<sup>3</sup>

obtained according to the scanning microscopy data.

**Figure 8.** Character of distribution of synthesized Fe3

It is established at the same time that the size of Fe3 O4 microparticles is about 1μm (**Figure 9(b)**).

Use of Fe3 O4 nanoparticles with the aim of their functionalization (application of shells, medi‐ cal aids and markers on them) or introduction in a living organism for hyperthermia foresees the application of surface coating. In this case, the analysis of the dimensions and properties of nanoparticles using a simple method becomes more complex [24]. In such cases, magnetic methods become one of the methods of particle categorization.

#### **2.4. Methods of magnetic properties investigation**

Among the magnetic research methods in materials science, the magnetic phase analysis is especially widely used. Its possibilities and effectiveness to a great extent are determined


**Table 1.** Correspondence of IR‐Fourier spectra peaks of Fe3 O4 –NP to chemical bonds [13–16]. Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods http://dx.doi.org/10.5772/intechopen.69717 389

**Figure 7.** Microstructure of synthesized Fe<sup>3</sup> O4 : (а) Fe<sup>3</sup> O4 nanoparticles; (b) Fe3 O4 microparticles. Microstructures were obtained according to the scanning microscopy data.

To estimate the sizes of Fe3

**Figure 6.** ІR‐Fourier spectra of Fe<sup>3</sup>

388 Biomaterials in Regenerative Medicine

Use of Fe3

O4

O4

nanoparticles.

O4

to 50 nm. In this case, the dominant size of nanoparticles is 8 nm.

methods become one of the methods of particle categorization.

**Wave number (cm−1) Chemical bond** 3230 ‐OH valence vibrations

1421 Not described

600 Fe‐O

**Table 1.** Correspondence of IR‐Fourier spectra peaks of Fe3

1617 ‐OH deformation ("wagging")

It is established at the same time that the size of Fe3

**2.4. Methods of magnetic properties investigation**

‐NP atomic‐force microscopy was used, which software prod‐

microparticles is about 1μm (**Figure 9(b)**).

ucts allowed the establishment of the scanned particle structure (**Figure 8**). Investigation data prove the results of microscopic investigations on the spherical nanoparticle accumulation. Two types of structural inhomogeneity can be distinguished on the investigated sample sur‐ face: conglomerates of conical‐like nanoparticles and granular texture of substrate. Surface topography is characterized by a rough relief with morphological regions of blocked struc‐ ture. Blocks are characterized by nonisometric round form with no surface faceting. Their height above the substrate surface is in the range from 5 to 10 nm and base diameter from 30

O4

cal aids and markers on them) or introduction in a living organism for hyperthermia foresees the application of surface coating. In this case, the analysis of the dimensions and properties of nanoparticles using a simple method becomes more complex [24]. In such cases, magnetic

Among the magnetic research methods in materials science, the magnetic phase analysis is especially widely used. Its possibilities and effectiveness to a great extent are determined

O4

–NP to chemical bonds [13–16].

nanoparticles with the aim of their functionalization (application of shells, medi‐

by technical characteristics of the equipment. The quantitative phase magnetic analysis uses properties of ferromagnetics, which they acquire in strong magnetic fields—in the state of technical saturation. Primary magnetic properties that are structurally insensitive are obtained from the curve of temperature dependence of saturation magnetization. Such characteristics are magnetization and Curie point. These values give the information about phase composi‐ tion of material and its changes in the process of certain thermal operations and also in the process of deformation. Curie point and saturation magnetization are called primary mag‐ netic properties because their values are determined by the nature of the ferromagnetic phase (crystal lattice, electron structure of atoms and chemical phase composition) [25].

Changes in magnetization and Curie point of separate phases observed in the process of inves‐ tigation of a certain material are not caused by their particle growth but are conditioned by the change of chemical composition of phases and their crystalline composition only. Based on these data, it is possible to study the kinetics of phase formation at the very early stages of the process (for particles sizes up to 10−6 cm) by the growth of magnetization and Curie point to their normal values. Thus, a weak dependence of saturation magnetization and Curie point

**Figure 8.** Character of distribution of synthesized Fe3 O4 ‐NP on the surface (a) and histogram of their distribution by size (b). Lateral dimension of particles in nanometers is given on X‐axis, and a number of scanned particles on Y‐axis.

Two measuring methods are used in vibration magnetometers—direct or compensation. The last one, in its turn, is of two types: the method of current shell and differential method. Compensation method allows avoiding the dependence of measuring results on the values of amplitude and frequency of vibrations. It is optimal to perform high‐temperature mea‐ suring preliminary assuming the measures of stabilization of sample mechanical vibration. Constant amplitude and frequency of sample vibrations is a required condition of providing

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

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391

The main cause of vibration amplitude instability is the changes in the moving parts of the magnetometer. Stability of measuring parameters is determined, in addition, by the stability of the alternating current generator which feeds the vibrator. Vibration amplitude stabiliza‐

Such a device for investigation of the magnetic properties possesses a number of unique char‐ acteristics. At room temperature, it is possible to build a hysteresis loop, partial hysteresis loops, initial curves of magnetization and demagnetization, and dependence of a magnetic moment on the sample orientation. At elevated temperatures, thermomagnetic measure‐ ments, high‐temperature hysteresis measurements, and time dependences of magnetic moment at different temperatures can be carried out. Temperature measuring interval ranges from 80 to 1100 K at power 3.5 кWt. Magnetic measurement in this work was performed using a vibration magnetometer [15]. Curves of the investigated samples over magnetization were recorded in magnetic fields from −200 кА/m to +200 кА/m by measuring the dependence of the given magnetization І/І200on the magnetic field strength Н, І200—magnetization at mag‐

a zero coercive force Нс and residual magnetization Іr (**Figure 10**(а)). Probably particles of

 due to a high degree of dispersivity are in a superparamagnetic state—state that is typical of microscopic and nanoscopic particles of ferromagnetic materials. In this case, a magnetic moment of the like particle changes its orientation spontaneously and randomly or due to thermal fluctuations. When external magnetic field is absent, superparamagnetics

‐NP sample remagnetization have a nonhysteresis form with

an adequate accuracy of measuring by the direct method.

netic field strength 200 кА/m.

As one can see, curves of Fe3

**Figure 10.** Curves of remagnetization of Fe3

Fe3 O4

tion was realized by providing a negative feedback of the generator.

O4

O4

‐NP (а) and Fe<sup>3</sup>

O4

‐MP (b) samples.

**Figure 9.** Character of distribution of synthesized microparticles of Fe3 O4 (а) and histogram of their distribution by size (b). Transverse dimension of particles in nanometers is given on X‐axis, and a number of scanned particles on Y‐axis (b).

on the stress state, form and degree of ferromagnetic dispersion allows us to choose these physical parameters as the quality characteristics of phases. Just these factors that strongly affect the results of the quantitative X‐ray analysis, in the magnetic phase analysis, practically have no effect on the investigation results.

To carry out magnetic phase analysis, it is necessary to build experimentally a temperature dependence of saturation magnetization of the tested material sample. Equipment for the analy‐ sis should provide measuring of magnetic properties in a wide range of temperatures and heat‐ ing and cooling rates. Peculiarity of the determination of the transition temperature of a number of dispersive magnetic systems to the paramagnetic state is that their magnetic phase analysis must be performed at high heating rates. Since in a number of cases the mass of the investigated samples can be very small (nanomaterials, fine‐dispersive powders, amorphous films, surface layers after different types of heat treatment) and as a result the value of the magnetic moment is relatively small (from 10−2 to 10−3 А·м<sup>2</sup> ), the equipment must possess a very high sensitivity.

For such investigations, it is most reasonable to use the method of vibrating specimen that is realized through vibration magnetometers [26].

Measurement accuracy with vibration magnetometer depends on the calibration accuracy. We have proposed the method of comparison when magnetometer is calibrated using a stan‐ dard with a known magnetic moment. As a standard, carbonyl iron of certain batches is used.

Application of the comparative method for calibration requires ensuring of the following conditions:


Two measuring methods are used in vibration magnetometers—direct or compensation. The last one, in its turn, is of two types: the method of current shell and differential method. Compensation method allows avoiding the dependence of measuring results on the values of amplitude and frequency of vibrations. It is optimal to perform high‐temperature mea‐ suring preliminary assuming the measures of stabilization of sample mechanical vibration. Constant amplitude and frequency of sample vibrations is a required condition of providing an adequate accuracy of measuring by the direct method.

The main cause of vibration amplitude instability is the changes in the moving parts of the magnetometer. Stability of measuring parameters is determined, in addition, by the stability of the alternating current generator which feeds the vibrator. Vibration amplitude stabiliza‐ tion was realized by providing a negative feedback of the generator.

Such a device for investigation of the magnetic properties possesses a number of unique char‐ acteristics. At room temperature, it is possible to build a hysteresis loop, partial hysteresis loops, initial curves of magnetization and demagnetization, and dependence of a magnetic moment on the sample orientation. At elevated temperatures, thermomagnetic measure‐ ments, high‐temperature hysteresis measurements, and time dependences of magnetic moment at different temperatures can be carried out. Temperature measuring interval ranges from 80 to 1100 K at power 3.5 кWt. Magnetic measurement in this work was performed using a vibration magnetometer [15]. Curves of the investigated samples over magnetization were recorded in magnetic fields from −200 кА/m to +200 кА/m by measuring the dependence of the given magnetization І/І200on the magnetic field strength Н, І200—magnetization at mag‐ netic field strength 200 кА/m.

on the stress state, form and degree of ferromagnetic dispersion allows us to choose these physical parameters as the quality characteristics of phases. Just these factors that strongly affect the results of the quantitative X‐ray analysis, in the magnetic phase analysis, practically

(b). Transverse dimension of particles in nanometers is given on X‐axis, and a number of scanned particles on Y‐axis (b).

O4

(а) and histogram of their distribution by size

To carry out magnetic phase analysis, it is necessary to build experimentally a temperature dependence of saturation magnetization of the tested material sample. Equipment for the analy‐ sis should provide measuring of magnetic properties in a wide range of temperatures and heat‐ ing and cooling rates. Peculiarity of the determination of the transition temperature of a number of dispersive magnetic systems to the paramagnetic state is that their magnetic phase analysis must be performed at high heating rates. Since in a number of cases the mass of the investigated samples can be very small (nanomaterials, fine‐dispersive powders, amorphous films, surface layers after different types of heat treatment) and as a result the value of the magnetic moment

For such investigations, it is most reasonable to use the method of vibrating specimen that is

Measurement accuracy with vibration magnetometer depends on the calibration accuracy. We have proposed the method of comparison when magnetometer is calibrated using a stan‐ dard with a known magnetic moment. As a standard, carbonyl iron of certain batches is used. Application of the comparative method for calibration requires ensuring of the following

• insignificant changes in the position and deviation from the perfect sphere‐like symmetry

• sample holder (container) introduces only an insignificant value to the signal value and can

), the equipment must possess a very high sensitivity.

have no effect on the investigation results.

390 Biomaterials in Regenerative Medicine

**Figure 9.** Character of distribution of synthesized microparticles of Fe3

is relatively small (from 10−2 to 10−3 А·м<sup>2</sup>

conditions:

realized through vibration magnetometers [26].

have no influence on a calibration constant;

be subtracted from it; and

• frequency and amplitude of vibrations remain constant;

• signal voltage—linear function of the magnetic moment of the sample.

As one can see, curves of Fe3 O4 ‐NP sample remagnetization have a nonhysteresis form with a zero coercive force Нс and residual magnetization Іr (**Figure 10**(а)). Probably particles of Fe3 O4 due to a high degree of dispersivity are in a superparamagnetic state—state that is typical of microscopic and nanoscopic particles of ferromagnetic materials. In this case, a magnetic moment of the like particle changes its orientation spontaneously and randomly or due to thermal fluctuations. When external magnetic field is absent, superparamagnetics

**Figure 10.** Curves of remagnetization of Fe3 O4 ‐NP (а) and Fe<sup>3</sup> O4 ‐MP (b) samples.

have on average a zero magnetic moment, i.e., they behave like paramagnetics with a high magnetic susceptibility. As known, superparamagnetic properties of Fe3 O4 particles at room temperature are exhibited, when reaching an average diameter D < 25 nm. In our case, the particles size is 5–10 nm.

AMR technology provides advantages over other magnetic sensor technologies. These aniso‐ tropic, directional sensors feature precision in‐axis sensitivity, linearity, and low cross‐axis sensitivity. For example, the HMC5883L is a small‐size surface‐mount, multi‐chip module designed for low‐field 3D magnetic sensing with a digital interface for applications such as low‐cost compassing and magnetometry. The HMC5883L includes the state‐of‐the‐art, high‐ resolution HMC118X series magnetoresistive sensors plus an application‐specific integrated circuit, containing amplification, automatic degaussing strap drivers, offset cancellation and

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393

The second type, Hall sensors, is based on Hall effect which is due to the nature of the current in a conductor. Current consists of the movement of many small charge carriers, typically electrons, holes or ions. When magnetic field is present, these charges experience a force, called the Lorentz force. When such a magnetic field is absent, the charges follow approximately straight "line of sight" paths between collisions with impurities, phonons and so on. However, when a magnetic field with a perpendicular component is applied, their paths between collisions are curved so that moving charges accumulate on one face of the material. This leaves equal and opposite charges exposed on the other face where there is a scarcity of mobile charges. The result is an asymmetric distribution of charge density across the Hall element, arising from the force that is perpendicular to both the "line of sight" path and the applied magnetic field. The separation of charge establishes an electric field that opposes the migration of further charge, so a steady elec‐

We have proposed new type of thin‐film magnetic field sensors for three‐dimensional sensing of orthogonal BX, BY and BZ projection of magnetic field vector [32–37], new method for mea‐

The general view of such thin‐film 3D sensors, its active area, structure and photographs are presented in **Figures 11**–**13**, correspondingly. The sensor (**Figure 12(a)**) active area is the thin sensitive semiconductor film (2) of corresponding configuration formed at the semi‐insu‐ lating GaAs substrate (1). Sensor contacts are formed by the metallization layer (3), which typically is the gold film together with other metals, for example, with titanium sublayer. For manufacturing of the structures, three photolithographies are used. The first one is intended

suring magnetic field [38] as well as their modeling and signal processing [39–44].

a 12‐bit analog‐to‐digit converter.

trical potential is established for as long as the charge is flowing.

**Figure 11.** General view of thin film 3D sensors (a) and its active area (b).

With increasing sizes of Fe3 O4 particles to microsizes, the magnetic properties do not change. These particles have already a hysteresis loop (**Figure 10(b)**). Samples' coercive force is Нс =4 кА/m, and ratio is І/І200=0.1.

#### **2.5. Use of the Internet of Things to measure of magnetic properties**

For a remote study of magnetic properties of nanoparticles, it is proposed in this work to use the approaches of the Internet of Things—a network that consists of the interrelated physi‐ cal objects or devices that contain mounted gauges. Their software allows one to pass and exchange data between physical world and computer systems using the conventional proce‐ dure of communication. In addition to gauges, the network contains executing devices, built‐ in physical objects interrelated by wire or wireless nets. These interrelated objects possess function of reading, putting in operation, programming and identification of objects and also allow one to exclude the necessity of person participation due to application of the intellectual interfaces.

Technology of magnetic particle detection and characterization is based on many approaches of magnetic sensing techniques [27, 28].

Here we present novel 3D magnetic sensors which offer three‐dimensional sensing of orthog‐ onal BX, BY, and BZ projections of magnetic field vector. Nowadays, many types of techniques and magnetic sensors provide such possibility [29, 30]. Among them, three types, namely magnetoresistors, Hall sensors and magnetotransistors, are preferable.

The first type, magnetoresistors, is based on the effect of magnetoresistance which is a ten‐ dency of material to change the value of its electrical resistance in an externally applied mag‐ netic field. There are a variety of effects that can be called magnetoresistance: some occur in bulk nonmagnetic metals and semiconductors, such as geometrical magnetoresistance, Shubnikov de Haas oscillations, or the common positive magnetoresistance in metals. Other effects occur in magnetic metals, such as negative magnetoresistance in ferromagnets or anisotropic magnetoresistance (AMR). Finally, in multicomponent or multilayer systems (e.g., magnetic tunnel junctions), giant magnetoresistance, tunnel magnetoresistance and extraordinary magnetoresistance can be observed [31].

As an example, Honeywell International Inc. produced a series of magnetic sensors based on anisotropic magnetoresistance effect. AMR is a property of material in which the dependence of electrical resistance on the angle between the direction of electric current and the direction of magnetization is observed. AMR arises from the simultaneous action of magnetization and spin‐orbit interaction, and its mechanism depends on the material. It can be for example due to a larger probability of s‐d scattering of electrons in the direction of magnetization, which is controlled by the applied magnetic field.

AMR technology provides advantages over other magnetic sensor technologies. These aniso‐ tropic, directional sensors feature precision in‐axis sensitivity, linearity, and low cross‐axis sensitivity. For example, the HMC5883L is a small‐size surface‐mount, multi‐chip module designed for low‐field 3D magnetic sensing with a digital interface for applications such as low‐cost compassing and magnetometry. The HMC5883L includes the state‐of‐the‐art, high‐ resolution HMC118X series magnetoresistive sensors plus an application‐specific integrated circuit, containing amplification, automatic degaussing strap drivers, offset cancellation and a 12‐bit analog‐to‐digit converter.

have on average a zero magnetic moment, i.e., they behave like paramagnetics with a high

temperature are exhibited, when reaching an average diameter D < 25 nm. In our case, the

change. These particles have already a hysteresis loop (**Figure 10(b)**). Samples' coercive

For a remote study of magnetic properties of nanoparticles, it is proposed in this work to use the approaches of the Internet of Things—a network that consists of the interrelated physi‐ cal objects or devices that contain mounted gauges. Their software allows one to pass and exchange data between physical world and computer systems using the conventional proce‐ dure of communication. In addition to gauges, the network contains executing devices, built‐ in physical objects interrelated by wire or wireless nets. These interrelated objects possess function of reading, putting in operation, programming and identification of objects and also allow one to exclude the necessity of person participation due to application of the intellectual

Technology of magnetic particle detection and characterization is based on many approaches

Here we present novel 3D magnetic sensors which offer three‐dimensional sensing of orthog‐ onal BX, BY, and BZ projections of magnetic field vector. Nowadays, many types of techniques and magnetic sensors provide such possibility [29, 30]. Among them, three types, namely

The first type, magnetoresistors, is based on the effect of magnetoresistance which is a ten‐ dency of material to change the value of its electrical resistance in an externally applied mag‐ netic field. There are a variety of effects that can be called magnetoresistance: some occur in bulk nonmagnetic metals and semiconductors, such as geometrical magnetoresistance, Shubnikov de Haas oscillations, or the common positive magnetoresistance in metals. Other effects occur in magnetic metals, such as negative magnetoresistance in ferromagnets or anisotropic magnetoresistance (AMR). Finally, in multicomponent or multilayer systems (e.g., magnetic tunnel junctions), giant magnetoresistance, tunnel magnetoresistance and

As an example, Honeywell International Inc. produced a series of magnetic sensors based on anisotropic magnetoresistance effect. AMR is a property of material in which the dependence of electrical resistance on the angle between the direction of electric current and the direction of magnetization is observed. AMR arises from the simultaneous action of magnetization and spin‐orbit interaction, and its mechanism depends on the material. It can be for example due to a larger probability of s‐d scattering of electrons in the direction of magnetization, which is

O4

particles to microsizes, the magnetic properties do not

particles at room

magnetic susceptibility. As known, superparamagnetic properties of Fe3

O4

**2.5. Use of the Internet of Things to measure of magnetic properties**

magnetoresistors, Hall sensors and magnetotransistors, are preferable.

extraordinary magnetoresistance can be observed [31].

controlled by the applied magnetic field.

=4 кА/m, and ratio is І/І200=0.1.

particles size is 5–10 nm.

392 Biomaterials in Regenerative Medicine

force is Нс

interfaces.

With increasing sizes of Fe3

of magnetic sensing techniques [27, 28].

The second type, Hall sensors, is based on Hall effect which is due to the nature of the current in a conductor. Current consists of the movement of many small charge carriers, typically electrons, holes or ions. When magnetic field is present, these charges experience a force, called the Lorentz force. When such a magnetic field is absent, the charges follow approximately straight "line of sight" paths between collisions with impurities, phonons and so on. However, when a magnetic field with a perpendicular component is applied, their paths between collisions are curved so that moving charges accumulate on one face of the material. This leaves equal and opposite charges exposed on the other face where there is a scarcity of mobile charges. The result is an asymmetric distribution of charge density across the Hall element, arising from the force that is perpendicular to both the "line of sight" path and the applied magnetic field. The separation of charge establishes an electric field that opposes the migration of further charge, so a steady elec‐ trical potential is established for as long as the charge is flowing.

We have proposed new type of thin‐film magnetic field sensors for three‐dimensional sensing of orthogonal BX, BY and BZ projection of magnetic field vector [32–37], new method for mea‐ suring magnetic field [38] as well as their modeling and signal processing [39–44].

The general view of such thin‐film 3D sensors, its active area, structure and photographs are presented in **Figures 11**–**13**, correspondingly. The sensor (**Figure 12(a)**) active area is the thin sensitive semiconductor film (2) of corresponding configuration formed at the semi‐insu‐ lating GaAs substrate (1). Sensor contacts are formed by the metallization layer (3), which typically is the gold film together with other metals, for example, with titanium sublayer. For manufacturing of the structures, three photolithographies are used. The first one is intended

**Figure 11.** General view of thin film 3D sensors (a) and its active area (b).

**Figure 12.** Structure and its formation stages (a), (b), (c), (d) of 3D thin film sensor: 1—substrate, 2—semiconductor active area; 3—contact metallization layer.

for creation of the contact system (**Figure 12(b)**); the second one is for etching a semiconductor film into the whole depth (**Figure 12(c)**); and the third one is for etching the semiconductor film to approximately 10% of their thickness (**Figure 12(d)**).

To show how a 3D thin‐film sensor works, let us consider its main component—transducer #1 (**Figure 14**). The magnetic field vector projections BX, BY are located in the transducer plane, and BZ projection is perpendicular to this plane.

The operation principle of transducer of type #1 is the following: transducer is connected to the power source, typically direct current source; for this, the central current contact 3 is connected to the first power supply output, and side current outputs 4, 5 are connected together to the second power supply output. Thus, the current in the active area of the transducer is divided equally and run in mutually antithetical directions in relation to the current output.

11. So the construction of the type #1 transducer provides the possibility of forming the output voltage at the 6 and 7 voltage contacts, which is equal to the voltage difference between the areas of lower surface of the active area in the neighborhood for the contacting to the inter‐ mediate areas 10 and 11. The voltage measurement at the areas of lower active area surface is obvious that the thickness of intermediate areas 10, 11 was considerably thinner (at least 10

**Figure 14.** Structure of type #1 transducer: 1—substrate, 2—semiconductor active area; 3—central current contact; 4, 5—

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395

side current contacts; 6, 7—voltage contacts; 8, 9—auxiliary areas; 10, 11—intermediate areas.

The principal difference of the type #1 transducer from analogs, for example Refs. [29, 30], is that in the first one the voltage difference is measured between the areas of lower active area surface, and in the analogs—between the areas of the upper active area surface. In its turn, the absence of contacts with the upper active area surface makes it possible not to perform the

In general case, the voltages at the voltage outputs 6, 7 are composed of three components. The first component VR is caused by the voltage drop at the semiconductor active area 2. Taking into consideration the symmetry of the active area in relation to the first current con‐ tact 3, the first voltage component at both voltage contacts 6, 7 is equal VR(6) = VR(7). The sec‐ ond component VZ is caused by the influence of BZ magnetic field vector projection, which is perpendicular to the transducer plane. Taking into consideration the transducer construction and current flow directions in it, the second voltage component at the voltage contacts 6 and 7 is also equal VZ(6) = VZ(7). The third component VX is caused by the influence of BX projection

times) than the thickness of active area 2.

active area surface insulation.

Neglecting the current in the 6 and 7 voltage contacts' circuit, one may consider that the volt‐ ages at those outputs are equal to the corresponding voltages of the intermediate areas 10 and

**Figure 13.** Photos of 3D thin‐film sensors: wafer (a), chip (b) and active area (c).

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods http://dx.doi.org/10.5772/intechopen.69717 395

for creation of the contact system (**Figure 12(b)**); the second one is for etching a semiconductor film into the whole depth (**Figure 12(c)**); and the third one is for etching the semiconductor

**Figure 12.** Structure and its formation stages (a), (b), (c), (d) of 3D thin film sensor: 1—substrate, 2—semiconductor active

To show how a 3D thin‐film sensor works, let us consider its main component—transducer #1 (**Figure 14**). The magnetic field vector projections BX, BY are located in the transducer plane,

The operation principle of transducer of type #1 is the following: transducer is connected to the power source, typically direct current source; for this, the central current contact 3 is connected to the first power supply output, and side current outputs 4, 5 are connected together to the second power supply output. Thus, the current in the active area of the transducer is divided equally and run in mutually antithetical directions in relation to the

Neglecting the current in the 6 and 7 voltage contacts' circuit, one may consider that the volt‐ ages at those outputs are equal to the corresponding voltages of the intermediate areas 10 and

film to approximately 10% of their thickness (**Figure 12(d)**).

**Figure 13.** Photos of 3D thin‐film sensors: wafer (a), chip (b) and active area (c).

and BZ projection is perpendicular to this plane.

area; 3—contact metallization layer.

394 Biomaterials in Regenerative Medicine

current output.

**Figure 14.** Structure of type #1 transducer: 1—substrate, 2—semiconductor active area; 3—central current contact; 4, 5 side current contacts; 6, 7—voltage contacts; 8, 9—auxiliary areas; 10, 11—intermediate areas.

11. So the construction of the type #1 transducer provides the possibility of forming the output voltage at the 6 and 7 voltage contacts, which is equal to the voltage difference between the areas of lower surface of the active area in the neighborhood for the contacting to the inter‐ mediate areas 10 and 11. The voltage measurement at the areas of lower active area surface is obvious that the thickness of intermediate areas 10, 11 was considerably thinner (at least 10 times) than the thickness of active area 2.

The principal difference of the type #1 transducer from analogs, for example Refs. [29, 30], is that in the first one the voltage difference is measured between the areas of lower active area surface, and in the analogs—between the areas of the upper active area surface. In its turn, the absence of contacts with the upper active area surface makes it possible not to perform the active area surface insulation.

In general case, the voltages at the voltage outputs 6, 7 are composed of three components. The first component VR is caused by the voltage drop at the semiconductor active area 2. Taking into consideration the symmetry of the active area in relation to the first current con‐ tact 3, the first voltage component at both voltage contacts 6, 7 is equal VR(6) = VR(7). The sec‐ ond component VZ is caused by the influence of BZ magnetic field vector projection, which is perpendicular to the transducer plane. Taking into consideration the transducer construction and current flow directions in it, the second voltage component at the voltage contacts 6 and 7 is also equal VZ(6) = VZ(7). The third component VX is caused by the influence of BX projection of induction vector. Contrary to the two stated above, this component at the voltage contacts 6 and 7 has the opposite signs VX(6) = −VX(7). In particular, in the upper part of active area, two charge carriers under the influence of electromotive force deviated to the upper semi‐ conductor layer surface, and then, in the lower part—to the lower surface (in the direction of substrate). The voltage difference caused by this carrier deviation is transferred to the voltage contacts 6 and 7 through the intermediate areas 10 and 11. It is to be noted that the BY magnetic field induction vector projection, parallel to the direction of current flow through the active area, does not cause the carrier deviation, and therefore, it may be neglected: VY(6) = VY(7) = 0.

*VX* = *V*<sup>8</sup> − *V*10; *VY* = *V*<sup>7</sup> − *V*<sup>9</sup>

The structure of type #3 transducer is shown in **Figure 16**.

area under the electromotive force influence.

contacts; 7, 8, 9, 10—voltage contacts; and 11—inner insulating area.

the informative signals.

; *VZ* = {(*V*<sup>8</sup> − *V*<sup>9</sup> ) +(*V*<sup>10</sup> − *V*<sup>7</sup> )}/2 (2)

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397

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

The difference of type #2 transducer from its analog [29, 30] is the limitation of the area of cur‐ rent transition between the 2nd outer and 11 inner insulating areas. This provides the increase of the sensor sensitivity to the BX, BY magnetic fields and decreases the cross‐impact between

The sensitivity increase is explained by the reduction of the transducer active area size. Contrary to the traditional Hall transducers (sensitive to the field BZ perpendicular to the transducer plane), where the size that determines the sensitivity is the active area thickness, in case of the BX, BY field transducers, the determinative size is the current scattering region. The smaller area occupies the current scattering region, the higher is the voltage difference between the two corresponding voltage contacts, and therefore the higher sensitivity is.

Arms 2, 4 and contacts 6, 8, 10, 12 form the first vertical Hall transducer, and arms 3, 5 and contacts 7, 9, 11, 13—the second vertical Hall transducer. The first transducer is intended for measuring the BX magnetic field vector projection, and the second one—for measurement of BY projection. The measurement principle of Hall transducers consists in forming the voltage difference at the voltage outputs during the deviation of charge carriers in the semiconductor

**Figure 15.** Structure of type #2 transducer: 1—semiconductor active area; 2—outer insulating area; 3, 4, 5, 6—current

Thus, the voltage difference between the voltage outputs 6 and 7 (output voltage of type#1 transducer) is determined only by BX magnetic field induction vector projection, and in first approximation, it does not depend on other vector projections:

$$\mathbf{V}\_{\rm OUT} = \begin{bmatrix} \mathbf{V}\_{\rm x}(6) + \mathbf{V}\_{\chi}(6) + \mathbf{V}\_{\chi}(6) + \mathbf{V}\_{\chi}(6) \end{bmatrix} - \begin{bmatrix} \mathbf{V}\_{\chi}(7) + \mathbf{V}\_{\chi}(7) + \mathbf{V}\_{\chi}(7) + \mathbf{V}\_{\chi}(7) \end{bmatrix} = \mathbf{V}\_{\chi'} \tag{1}$$

where VX = VX(6)+VX(7).

The especially high efficiency of transducer of type#1 application is realized in case of their application in 3D sensor for simultaneous measurement of three projections BX, BY and BZ of the magnetic field induction vector. Such 3D sensor contains three coupled transducers at a single substrate. The construction of two transducers of #1 type is similar to **Figure 14**. They are mutually orthogonally rotated and provide the sensitivity to the BX, BY vector projections of the field induction vector. The third transducer, which provides the sensitivity to the BZ projection, is a traditional Hall transducer [29].

To provide the equal sensitivity levels for all three transducers of such 3D sensor, it is neces‐ sary to take into account the following: contrary to the traditional Hall transducers, which have the sensitivity as an inverse function of semiconductor active layer thickness d, the sen‐ sitivity of the transducer of type #1 is inversely proportional function of the semiconductor area width W. So it is recommended that the active area width W should be minimal and approximately equal to the semiconductor layer thickness d.

There are two other options to build up 3D magnetic thin film sensors. They are based on type #2 and type #3 transducers.

The structure of type #2 transducer is shown in **Figure 15**.

The type #2 transducer is fed by the direct current source, and one pair of mutually opposed current contacts (e.g., 3 and 5) is connected to one output of the power source (e.g., positive), and another pair (4 and 6 correspondently) to the other power source contact (negative, cor‐ respondently). In accordance with this connection scheme, the four current flow circuits are formed: I34, I36, I54 and I56, where the indices in the marked currents correspond to the numbers of current contacts. These current flow circuits geometrically form the square sides. The mat‐ ter of principal importance is that the currents that flow in opposed square sides are equal by value and opposed by sign: →*<sup>I</sup>* <sup>34</sup> = − →*I* 65; <sup>→</sup>*<sup>I</sup>* <sup>36</sup> = − →*I* 45. The output signal of transducer is the voltage difference between the voltage contacts. The informative signals about the magnetic field vector projections BX, BY, BZ are the voltages VX, VY, VZ, which in first approximation are determined as:

$$V\_{\chi} = V\_s - V\_{10'} \; \because \; V\_{\chi} = V\_{\tau} - V\_{\varphi} \; \; \; V\_{\mathcal{Z}} = \left\{ (V\_s - V\_{\varphi}) \star (V\_{10} - V\_{\tau}) \right\} / 2 \tag{2}$$

The difference of type #2 transducer from its analog [29, 30] is the limitation of the area of cur‐ rent transition between the 2nd outer and 11 inner insulating areas. This provides the increase of the sensor sensitivity to the BX, BY magnetic fields and decreases the cross‐impact between the informative signals.

The sensitivity increase is explained by the reduction of the transducer active area size. Contrary to the traditional Hall transducers (sensitive to the field BZ perpendicular to the transducer plane), where the size that determines the sensitivity is the active area thickness, in case of the BX, BY field transducers, the determinative size is the current scattering region. The smaller area occupies the current scattering region, the higher is the voltage difference between the two corresponding voltage contacts, and therefore the higher sensitivity is.

The structure of type #3 transducer is shown in **Figure 16**.

of induction vector. Contrary to the two stated above, this component at the voltage contacts 6 and 7 has the opposite signs VX(6) = −VX(7). In particular, in the upper part of active area, two charge carriers under the influence of electromotive force deviated to the upper semi‐ conductor layer surface, and then, in the lower part—to the lower surface (in the direction of substrate). The voltage difference caused by this carrier deviation is transferred to the voltage contacts 6 and 7 through the intermediate areas 10 and 11. It is to be noted that the BY magnetic field induction vector projection, parallel to the direction of current flow through the active area, does not cause the carrier deviation, and therefore, it may be neglected: VY(6) = VY(7) = 0. Thus, the voltage difference between the voltage outputs 6 and 7 (output voltage of type#1 transducer) is determined only by BX magnetic field induction vector projection, and in first

approximation, it does not depend on other vector projections:

(6) + VY

(6) + VZ

(6) ] –[VR

The especially high efficiency of transducer of type#1 application is realized in case of their application in 3D sensor for simultaneous measurement of three projections BX, BY and BZ of the magnetic field induction vector. Such 3D sensor contains three coupled transducers at a single substrate. The construction of two transducers of #1 type is similar to **Figure 14**. They are mutually orthogonally rotated and provide the sensitivity to the BX, BY vector projections of the field induction vector. The third transducer, which provides the sensitivity to the BZ

To provide the equal sensitivity levels for all three transducers of such 3D sensor, it is neces‐ sary to take into account the following: contrary to the traditional Hall transducers, which have the sensitivity as an inverse function of semiconductor active layer thickness d, the sen‐ sitivity of the transducer of type #1 is inversely proportional function of the semiconductor area width W. So it is recommended that the active area width W should be minimal and

There are two other options to build up 3D magnetic thin film sensors. They are based on type

The type #2 transducer is fed by the direct current source, and one pair of mutually opposed current contacts (e.g., 3 and 5) is connected to one output of the power source (e.g., positive), and another pair (4 and 6 correspondently) to the other power source contact (negative, cor‐ respondently). In accordance with this connection scheme, the four current flow circuits are formed: I34, I36, I54 and I56, where the indices in the marked currents correspond to the numbers of current contacts. These current flow circuits geometrically form the square sides. The mat‐ ter of principal importance is that the currents that flow in opposed square sides are equal

difference between the voltage contacts. The informative signals about the magnetic field vector projections BX, BY, BZ are the voltages VX, VY, VZ, which in first approximation are determined as:

(7) + VX

(7) + VY

(7) + VZ

45. The output signal of transducer is the voltage

(7)

] = VX, (1)

(6) + VX

projection, is a traditional Hall transducer [29].

approximately equal to the semiconductor layer thickness d.

The structure of type #2 transducer is shown in **Figure 15**.

<sup>34</sup> = − →*I* <sup>65</sup>; <sup>→</sup>*<sup>I</sup>* <sup>36</sup> = − →*I*

VOUT = [VR

where VX = VX(6)+VX(7).

396 Biomaterials in Regenerative Medicine

#2 and type #3 transducers.

by value and opposed by sign: →*<sup>I</sup>*

Arms 2, 4 and contacts 6, 8, 10, 12 form the first vertical Hall transducer, and arms 3, 5 and contacts 7, 9, 11, 13—the second vertical Hall transducer. The first transducer is intended for measuring the BX magnetic field vector projection, and the second one—for measurement of BY projection. The measurement principle of Hall transducers consists in forming the voltage difference at the voltage outputs during the deviation of charge carriers in the semiconductor area under the electromotive force influence.

**Figure 15.** Structure of type #2 transducer: 1—semiconductor active area; 2—outer insulating area; 3, 4, 5, 6—current contacts; 7, 8, 9, 10—voltage contacts; and 11—inner insulating area.

where V(11), V(12), V(13) and V(14) are the voltages at the voltage outputs 11, 12, 13 and 14, correspondingly; VX, VY and KX and KY—output signals and transducing coefficients of the first and second transducers, correspondingly; I—operational current, W—the width of

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

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From the physical point of view, the appearance of voltage difference at the voltage contacts of the vertical Hall transducers is explained by the fact that due to the opposed current flow directions in both transducer arms, the deviation of current carriers in those arms has also the opposed direction. In particular, in arm 2 of the first transducer, the carriers are deviated in the direction away from the substrate to the surface of the semiconductor area, and then, in

In a high‐gradient magnetic field, the voltage difference formed at the voltage outputs of the vertical Hall transducers is the informative value of the averaged field induction value. Taking into the consideration that all the voltage outputs of the vertical Hall transducers are equidis‐ tant from its center (crossing area), the measured averaged induction value corresponds to the

The second power‐supplying mode presumes the application of only one pair of current con‐ tacts, namely 6 and 8 of the first vertical transducer. Then, output 6 is connected to the first power supply output, and output 8 to the second one. This provides the linear trajectory of

The output signal of the horizontal Hall transducer, as a voltage difference, is proportional to the multiplication of the transducer power supply current by the BZ projection of the magnetic field induction vector, formed at the voltage outputs 11 and 13 of the second vertical Hall

From the physical point of view, the appearance of the voltage difference at the voltage out‐ puts 11 and 13 is explained by the fact that charge carrier deviation in semiconductor area 14 of the horizontal Hall transducer under the influence of BZ magnetic field induction vector

As in the vertical transducer, the horizontal Hall transducer measures the value of the mag‐ netic field induction value in the transducer center spatial point. Thus, type #3 transducer allows measuring of all three projections BX, BY and BZ of the magnetic field induction vector

ness of the horizontal Hall transducer semiconductor area. This allows the formation of the transducer with equal sensitivity value to all three magnetic field induction vector projec‐ tions: VX/BX = VY/BY = VZ/BZ. Taking into consideration that the sensitivity of vertical Hall transducers is inversely proportional to the semiconductor area width W, and for horizontal

The other distinctive feature of the type #3 transducer is the possibility to change the d<sup>2</sup>

thickness of the vertical Hall transducer areas, the d2

zontal transducer semiconductor area may change after the output formation. In particular,

thickness may be realized by partial etching of the semiconductor layer.

, where KZ is the coefficient of transduction and d<sup>2</sup>

is

399

thick‐

, the equality of the stated values

ratio. It is important that

thickness of the hori‐

arm 4 of this transducer—in the direction from the surface to the substrate.

the charge carriers in semiconductor area 14 of the horizontal Hall transducer.

the thickness of semiconductor area of the horizontal Hall transducer.

projection is in the direction from arm 5 to arm 3, or backwards.

transducer—to the thickness of its semiconductor area d2

of sensitivity is provided by the correspondent selection of W/d2

semiconductor areas.

spatial point of the transducer center.

transducer: VZ = V(13)–V(11) = KZ.I.BZ/d2

in a single spatial point.

contrary to the d1

the decrease of d2

**Figure 16.** Structure of type #3 transducer: 1—substrate; 2, 3, 4, 5—four arms of cross‐shaped figure, formed by the crossing of two semiconductor areas of vertical Hall transducers; 6, 7, 8, 9—current contacts; 10, 11, 12, 13—voltage contacts; and 14—semiconductor area of horizontal Hall transducer.

The operation of type #3 transducer presumes two power supplying modes. The first one provides the operation of the vertical Hall transducers, and the second one for operation of the horizontal one.

The first power‐supplying mode presumes the connecting of the current contacts of the verti‐ cal Hall transducers, when the currents in their semiconductor areas are flowing in mutu‐ ally opposed directions. For this, the current contacts of both vertical Hall transducers are connected into one circuit. Namely, contacts 6 and 8 of the first vertical Hall transducer are connected together and to the first, for example, positive output of the power source, and contacts 7 and 9 of the second vertical Hall transducer are also connected together to the sec‐ ond, therefore, negative output of the power source. So, in the first vertical Hall transducer, the currents are flowing from the top to the bottom (current I<sup>2</sup> in arm 2) and from the bottom to the top (current I4 in arm 4), and in the second—from left to right (current I3 in arm 3) and from right to left (current I5 in arm 5). In case of ideal symmetry of the transducer structure, the following equality takes place: <sup>→</sup>*I* = − →*I*; <sup>→</sup>*I* = − →*I*.

Output signals of vertical Hall transducers are formed at the voltage outputs as a voltage dif‐ ference, which is proportional to the multiplication of the power supply current value of the transducer by the corresponding magnetic field vector projection:

$$\mathbf{V}\_{\chi} = \text{ V(12)} - \text{V(10)} = \mathbf{K}\_{\chi} \times \mathbf{I} \times \mathbf{B}\_{\chi} / \text{ W-for the first transformer and} \tag{3}$$

and

$$\mathbf{V}\_{\chi} = \mathbf{V}(12) - \mathbf{V}(10) \ = \mathbf{K}\_{\chi} \times \mathbf{I} \times \mathbf{B}\_{\chi} / \ \mathbf{W} \text{-for the second transformer and} \tag{4}$$

where V(11), V(12), V(13) and V(14) are the voltages at the voltage outputs 11, 12, 13 and 14, correspondingly; VX, VY and KX and KY—output signals and transducing coefficients of the first and second transducers, correspondingly; I—operational current, W—the width of semiconductor areas.

From the physical point of view, the appearance of voltage difference at the voltage contacts of the vertical Hall transducers is explained by the fact that due to the opposed current flow directions in both transducer arms, the deviation of current carriers in those arms has also the opposed direction. In particular, in arm 2 of the first transducer, the carriers are deviated in the direction away from the substrate to the surface of the semiconductor area, and then, in arm 4 of this transducer—in the direction from the surface to the substrate.

In a high‐gradient magnetic field, the voltage difference formed at the voltage outputs of the vertical Hall transducers is the informative value of the averaged field induction value. Taking into the consideration that all the voltage outputs of the vertical Hall transducers are equidis‐ tant from its center (crossing area), the measured averaged induction value corresponds to the spatial point of the transducer center.

The second power‐supplying mode presumes the application of only one pair of current con‐ tacts, namely 6 and 8 of the first vertical transducer. Then, output 6 is connected to the first power supply output, and output 8 to the second one. This provides the linear trajectory of the charge carriers in semiconductor area 14 of the horizontal Hall transducer.

The output signal of the horizontal Hall transducer, as a voltage difference, is proportional to the multiplication of the transducer power supply current by the BZ projection of the magnetic field induction vector, formed at the voltage outputs 11 and 13 of the second vertical Hall transducer: VZ = V(13)–V(11) = KZ.I.BZ/d2 , where KZ is the coefficient of transduction and d<sup>2</sup> is the thickness of semiconductor area of the horizontal Hall transducer.

The operation of type #3 transducer presumes two power supplying modes. The first one provides the operation of the vertical Hall transducers, and the second one for operation of

**Figure 16.** Structure of type #3 transducer: 1—substrate; 2, 3, 4, 5—four arms of cross‐shaped figure, formed by the crossing of two semiconductor areas of vertical Hall transducers; 6, 7, 8, 9—current contacts; 10, 11, 12, 13—voltage

The first power‐supplying mode presumes the connecting of the current contacts of the verti‐ cal Hall transducers, when the currents in their semiconductor areas are flowing in mutu‐ ally opposed directions. For this, the current contacts of both vertical Hall transducers are connected into one circuit. Namely, contacts 6 and 8 of the first vertical Hall transducer are connected together and to the first, for example, positive output of the power source, and contacts 7 and 9 of the second vertical Hall transducer are also connected together to the sec‐ ond, therefore, negative output of the power source. So, in the first vertical Hall transducer,

in arm 4), and in the second—from left to right (current I3

Output signals of vertical Hall transducers are formed at the voltage outputs as a voltage dif‐ ference, which is proportional to the multiplication of the power supply current value of the

VX = V(12) – V(10) = K<sup>X</sup> × I × BX / W–for the first transducer and (3)

VX = V(12) – V(10) = K<sup>X</sup> × I × BX / W–for the second transducer and (4)

→*I*; <sup>→</sup>*I* = − →*I*.

in arm 5). In case of ideal symmetry of the transducer structure,

in arm 2) and from the bottom

in arm 3) and

the currents are flowing from the top to the bottom (current I<sup>2</sup>

contacts; and 14—semiconductor area of horizontal Hall transducer.

transducer by the corresponding magnetic field vector projection:

the horizontal one.

398 Biomaterials in Regenerative Medicine

to the top (current I4

and

from right to left (current I5

the following equality takes place: <sup>→</sup>*I* = −

From the physical point of view, the appearance of the voltage difference at the voltage out‐ puts 11 and 13 is explained by the fact that charge carrier deviation in semiconductor area 14 of the horizontal Hall transducer under the influence of BZ magnetic field induction vector projection is in the direction from arm 5 to arm 3, or backwards.

As in the vertical transducer, the horizontal Hall transducer measures the value of the mag‐ netic field induction value in the transducer center spatial point. Thus, type #3 transducer allows measuring of all three projections BX, BY and BZ of the magnetic field induction vector in a single spatial point.

The other distinctive feature of the type #3 transducer is the possibility to change the d<sup>2</sup> thick‐ ness of the horizontal Hall transducer semiconductor area. This allows the formation of the transducer with equal sensitivity value to all three magnetic field induction vector projec‐ tions: VX/BX = VY/BY = VZ/BZ. Taking into consideration that the sensitivity of vertical Hall transducers is inversely proportional to the semiconductor area width W, and for horizontal transducer—to the thickness of its semiconductor area d2 , the equality of the stated values of sensitivity is provided by the correspondent selection of W/d2 ratio. It is important that contrary to the d1 thickness of the vertical Hall transducer areas, the d2 thickness of the hori‐ zontal transducer semiconductor area may change after the output formation. In particular, the decrease of d2 thickness may be realized by partial etching of the semiconductor layer.

In relation to the analogs [29], the type #3 transducer provides the increase of the measure‐ ment accuracy and construction simplification.

The increasing accuracy is caused by the fact that horizontal Hall transducer is placed in the center (crossing area) of vertical Hall transducer. This provides the high spatial alignment of all transducers (two vertical and one horizontal), and therefore, all the three projections BX, BY and BZ of the magnetic field induction vector are measured at a single spatial point. During the measurement of high‐gradient fields, this gives the ability to decrease the magnetic field induction vector measurement error in several times.

The third type, magnetoresistors, is mostly bipolar or field‐effect transistors whose structures and operating conditions are optimized with respect to the magnetic sensitivity of its output currents. There are three major effects of magnetotransistors, namely the current deflection effect, the injection modulation and the magnetodiode effect.

We have proposed the structure of magnetotransistors adapted to three dimensional sensing of orthogonal BX, BY and BZ projections of the magnetic field vector (**Figures 17**–**19**). These are drift‐aided lateral double‐collector p‐n‐p transistors. The emitter area E and two collector p‐type areas C1, C2 are embedded into plate‐like n‐type base region. The two base contact n+ type areas B1, B2 are used to apply the bias voltage needed to establish the lateral accelerating electric field in the region. The emitter injects holes into the base region. Under the influence of the accelerating field in the base, these holes form a minority carrier beam. In the presence of normal BZ magnetic field (**Figure 17**), the electric field in the base region rotates for the Hall angle of majority carries, and the hole beam tilts with respect to the electric field for the Hall

angle. The deflection of a beam leads to an imbalance in the two C1, C2 collector currents. In the presence of BY magnetic field vector projection (**Figure 18**), the hole beam tilts up to collectors

**Figure 18.** Structure of magnetotransistor (a) and current deflection (b) I = f(BZ) under BZ magnetic field vector projection:

type)—base contacts; E (p type)—emitter area; C1, C2 (p type)—collector areas; G—a substrate connected to

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401

**Figure 19.** Photos of magnetotransistors: single (a), doubled (b) and quad (c) structures.

tures are used to provide a set of signals in correspondence with three projections of magnetic

Based on the sensors proposed and described above, a set of magnetometers for detection and characterization of magnetic particles has been developed. For illustration only, an example of software windows for control and measurement result visualization, as well as a measure‐ ment resolution archived (in this case about 10−7T), are shown in **Figures 20** and **21**. As it was noted, the main advantage of such sensors is a possibility of measuring three orthogonal projections of a magnetic field vector that is quite new in characterization of magnetic particle

type substrate connected to ground. Doubled and quad (**Figure 19**) struc‐

or down toward a p<sup>−</sup>

field vector.

B1, B2 (n+

ground (p<sup>−</sup>

type).

technique.

**Figure 17.** Structure of magnetotransistor (a) and current deflection (b) I = f(BZ) under BZ magnetic field vector projection: B1, B2—base contacts; E—emitter area; C1, C2—collector areas; and G–connected to ground a p‐n junction wall insulation.

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods http://dx.doi.org/10.5772/intechopen.69717 401

In relation to the analogs [29], the type #3 transducer provides the increase of the measure‐

The increasing accuracy is caused by the fact that horizontal Hall transducer is placed in the center (crossing area) of vertical Hall transducer. This provides the high spatial alignment of all transducers (two vertical and one horizontal), and therefore, all the three projections BX, BY and BZ of the magnetic field induction vector are measured at a single spatial point. During the measurement of high‐gradient fields, this gives the ability to decrease the magnetic field

The third type, magnetoresistors, is mostly bipolar or field‐effect transistors whose structures and operating conditions are optimized with respect to the magnetic sensitivity of its output currents. There are three major effects of magnetotransistors, namely the current deflection

We have proposed the structure of magnetotransistors adapted to three dimensional sensing of orthogonal BX, BY and BZ projections of the magnetic field vector (**Figures 17**–**19**). These are drift‐aided lateral double‐collector p‐n‐p transistors. The emitter area E and two collector p‐type areas C1, C2 are embedded into plate‐like n‐type base region. The two base contact n+ type areas B1, B2 are used to apply the bias voltage needed to establish the lateral accelerating electric field in the region. The emitter injects holes into the base region. Under the influence of the accelerating field in the base, these holes form a minority carrier beam. In the presence of normal BZ magnetic field (**Figure 17**), the electric field in the base region rotates for the Hall angle of majority carries, and the hole beam tilts with respect to the electric field for the Hall

**Figure 17.** Structure of magnetotransistor (a) and current deflection (b) I = f(BZ) under BZ magnetic field vector projection: B1, B2—base contacts; E—emitter area; C1, C2—collector areas; and G–connected to ground a p‐n junction wall insulation.

ment accuracy and construction simplification.

400 Biomaterials in Regenerative Medicine

induction vector measurement error in several times.

effect, the injection modulation and the magnetodiode effect.

**Figure 18.** Structure of magnetotransistor (a) and current deflection (b) I = f(BZ) under BZ magnetic field vector projection: B1, B2 (n+ type)—base contacts; E (p type)—emitter area; C1, C2 (p type)—collector areas; G—a substrate connected to ground (p<sup>−</sup> type).

**Figure 19.** Photos of magnetotransistors: single (a), doubled (b) and quad (c) structures.

angle. The deflection of a beam leads to an imbalance in the two C1, C2 collector currents. In the presence of BY magnetic field vector projection (**Figure 18**), the hole beam tilts up to collectors or down toward a p<sup>−</sup> type substrate connected to ground. Doubled and quad (**Figure 19**) struc‐ tures are used to provide a set of signals in correspondence with three projections of magnetic field vector.

Based on the sensors proposed and described above, a set of magnetometers for detection and characterization of magnetic particles has been developed. For illustration only, an example of software windows for control and measurement result visualization, as well as a measure‐ ment resolution archived (in this case about 10−7T), are shown in **Figures 20** and **21**. As it was noted, the main advantage of such sensors is a possibility of measuring three orthogonal projections of a magnetic field vector that is quite new in characterization of magnetic particle technique.

properties and methods of identification should be considered. Special attention should be paid to sizes, topography and biocompatibility of initial nanoparticles. Just these parameters

Identification of Fe3O4 Nanoparticles Biomedical Purpose by Magnetometric Methods

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403

It is shown in the paper that identification of synthesized nanoparticles can be realized most effectively by the change of their magnetic properties. This is possible both at the stage of their synthesis and after biofunctionalization. Success of the proposed approach formed the basis of the ideology of the Internet of Things use for monitoring the stability of nanoparticles

We are infinitely grateful to Mykhailo Gonchar and Tetiana Prokopiv from the Institute of Cell Biology of the National Academy of Sciences of Ukraine, Department of Analytical Biotechnology, for kindly provided samples of Fe3O4 powders for our research. Also we would like to acknowledge financial support of the Ministry of Education and Science of Ukriaine under Grant 0116U004142 and partially by funds of 543994‐TEMPUS‐1‐2013‐1‐BE‐ TEMPUS‐JPCR MMATENG (Modernization of two cycles (MA, BA) of competence‐based curricula in Material Engineering according to the best experience of Bologna Process – EU).

[1] De Crozals G, Bonnet R, Farre C, Chaix C. Nanoparticles with multiple properties for

[2] Nikiforov VN. Biomedical applications of magnetic nanoparticles. Nauka i tehnologii v

[3] Filippousi M, Angelakeris M, Katsikini M, Paloura EC, Efthimiopoulos I, Wang Y, Zamboulis D, Van Tendeloo G. Surfactant effects on the structural and on the magnetic properties of iron oxide nanoparticles. Journal of Physical Chemistry C. 2014;**118**(29):16209‐16217. DOI: 10.1021/jp5037266

biomedical applications: A strategic guide. Nano Today. 2016;**11**(4):435‐463

, Volodymyr Kulyk<sup>1</sup>

, Peter Arras2

and

microstructure and their functional properties in application in medicine.

, Tetiana Tepla1

finally determine the functional properties.

\*, Roman Holyaka1

2 KU Leuven, Sint‐Katelijne‐Waver, Belgium

promyishlennosti. 2011;**1**:90‐99

3 Technische Universität Berlin, Berlin, Germany

\*Address all correspondence to: zduriagina@ukr.net 1 Lviv Polytechnic National University, Lviv, Ukraine

**Acknowledgements**

**Author details**

Zoia Duriagina1

Elena Eyngorn3

**References**

**Figure 20.** Software of magnetometer based on 3D sensors.

**Figure 21.** Measurement resolution (noise) of magnetometer based on 3D sensors.

#### **3. Conclusion**

It is possible to say that the interest to nanomaterials with magnetic properties increases every year. These materials have found their applications in medicine for diagnostics and treat‐ ment of serious diseases, where treatment by other methods is more expensive and durable. However, synthesis of such nanomaterials with stable properties is not completely elabo‐ rated, since besides the size factor, the character of their building microgeometry, functional properties and methods of identification should be considered. Special attention should be paid to sizes, topography and biocompatibility of initial nanoparticles. Just these parameters finally determine the functional properties.

It is shown in the paper that identification of synthesized nanoparticles can be realized most effectively by the change of their magnetic properties. This is possible both at the stage of their synthesis and after biofunctionalization. Success of the proposed approach formed the basis of the ideology of the Internet of Things use for monitoring the stability of nanoparticles microstructure and their functional properties in application in medicine.

## **Acknowledgements**

We are infinitely grateful to Mykhailo Gonchar and Tetiana Prokopiv from the Institute of Cell Biology of the National Academy of Sciences of Ukraine, Department of Analytical Biotechnology, for kindly provided samples of Fe3O4 powders for our research. Also we would like to acknowledge financial support of the Ministry of Education and Science of Ukriaine under Grant 0116U004142 and partially by funds of 543994‐TEMPUS‐1‐2013‐1‐BE‐ TEMPUS‐JPCR MMATENG (Modernization of two cycles (MA, BA) of competence‐based curricula in Material Engineering according to the best experience of Bologna Process – EU).

## **Author details**

Zoia Duriagina1 \*, Roman Holyaka1 , Tetiana Tepla1 , Volodymyr Kulyk<sup>1</sup> , Peter Arras2 and Elena Eyngorn3


## **References**

**3. Conclusion**

It is possible to say that the interest to nanomaterials with magnetic properties increases every year. These materials have found their applications in medicine for diagnostics and treat‐ ment of serious diseases, where treatment by other methods is more expensive and durable. However, synthesis of such nanomaterials with stable properties is not completely elabo‐ rated, since besides the size factor, the character of their building microgeometry, functional

**Figure 21.** Measurement resolution (noise) of magnetometer based on 3D sensors.

**Figure 20.** Software of magnetometer based on 3D sensors.

402 Biomaterials in Regenerative Medicine


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**Chapter 18**

**Provisional chapter**

**Biomaterials for Tissue Engineering Applications in**

The search for ideal implants or alternative scaffolds is a challenge for biomedical science researchers, especially in diabetic patients. Many alternative bioactive materials have been used in the regenerative medicine, especially in patient with complex metabolic disorder as diabetes mellitus. Among them, we discussed the following alternative material scaffolds, including amniotic membrane (AM), homogenous demineralized dentin matrix (HDDM), platelet-rich plasma (PRP), and alloplastic materials as porous polyethylene and polyurethane. These biomaterials were applied in the craniomaxillofacial complex and liver injury, resulting in tissue regeneration and microstructural reconstruction due to their effective inductive and conductive properties. Additionally, diabetes disease and its general biophysical mechanism and systemic complications were described in order to improve the comprehension of the physiopathology of this comorbidity and its effects in the tissues. The AM, HDDM, and PRP in implantation sites initiated an inductive cascade as chemotaxis of progenitor cells, mitogenesis, angiogenesis, and differentiation into wide variety of cells. The cell recruitment, division rate, and differentiation of cell lines are under the direct control of several growth factors and stem cells which are present in these biomaterials. Further, some alloplastic materials have triggered satisfactory tissue responses when used in treatments of craniofacial deformities or in anatomi-

**Biomaterials for Tissue Engineering Applications in** 

DOI: 10.5772/intechopen.69719

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution,

© 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

and reproduction in any medium, provided the original work is properly cited.

**Keywords:** regenerative medicine, amniotic membrane, homogenous dentin demineralized

matrix, platelet-rich plasma, alloplastic material, diabetes mellitus

Mônica Fernandes Gomes, José Benedito Amorim,

Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

Mônica Fernandes Gomes, José Benedito Amorim, Lilian Chrystiane Giannasi and

**Diabetes Mellitus**

**Diabetes Mellitus**

Lilian Chrystiane Giannasi and Miguel Angel Castillo Salgado

Miguel Angel Castillo Salgado

http://dx.doi.org/10.5772/intechopen.69719

**Abstract**

cal reconstructions.

**Provisional chapter**

## **Biomaterials for Tissue Engineering Applications in Diabetes Mellitus Diabetes Mellitus**

**Biomaterials for Tissue Engineering Applications in** 

DOI: 10.5772/intechopen.69719

Mônica Fernandes Gomes, José Benedito Amorim, Lilian Chrystiane Giannasi and Miguel Angel Castillo Salgado Amorim, Lilian Chrystiane Giannasi and Miguel Angel Castillo Salgado Additional information is available at the end of the chapter

Additional information is available at the end of the chapter

Mônica Fernandes Gomes, José Benedito

http://dx.doi.org/10.5772/intechopen.69719

#### **Abstract**

The search for ideal implants or alternative scaffolds is a challenge for biomedical science researchers, especially in diabetic patients. Many alternative bioactive materials have been used in the regenerative medicine, especially in patient with complex metabolic disorder as diabetes mellitus. Among them, we discussed the following alternative material scaffolds, including amniotic membrane (AM), homogenous demineralized dentin matrix (HDDM), platelet-rich plasma (PRP), and alloplastic materials as porous polyethylene and polyurethane. These biomaterials were applied in the craniomaxillofacial complex and liver injury, resulting in tissue regeneration and microstructural reconstruction due to their effective inductive and conductive properties. Additionally, diabetes disease and its general biophysical mechanism and systemic complications were described in order to improve the comprehension of the physiopathology of this comorbidity and its effects in the tissues. The AM, HDDM, and PRP in implantation sites initiated an inductive cascade as chemotaxis of progenitor cells, mitogenesis, angiogenesis, and differentiation into wide variety of cells. The cell recruitment, division rate, and differentiation of cell lines are under the direct control of several growth factors and stem cells which are present in these biomaterials. Further, some alloplastic materials have triggered satisfactory tissue responses when used in treatments of craniofacial deformities or in anatomical reconstructions.

**Keywords:** regenerative medicine, amniotic membrane, homogenous dentin demineralized matrix, platelet-rich plasma, alloplastic material, diabetes mellitus

© 2016 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. © 2018 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

## **1. Introduction**

Diabetes mellitus (DM), a complex metabolic disorder, is a syndrome characterized by abnormalities in carbohydrate, lipid, and protein metabolism, which results either from a partial or an absolute insulin deficiency or from target tissue resistance to its cellular metabolic effects. This disease is characterized by the presence of few or by the absence of functional pancreatic β-cells in the islets of Langerhans and by a substantial reduction or inexistence of insulin secretion. This cellular dysfunction is an important defect in the pathogenesis of type 2 diabetes [1, 29].

and protein metabolism, may develop in, called diabetes mellitus (DM). According to the International Diabetes Federation, 8.8% of the adult population worldwide has diabetes. Of all individuals with diabetes, only 10–15% have type 1 diabetes mellitus (T1DM); type 2 diabetes mellitus (T2DM) is the most common form. However, T1DM is the most common form of diabetes in children (<15 years of age), and >500,000 children are currently living with this condition globally. The classification of diabetes mellitus disease defines both process and stage of the disease. The processes include type 1, autoimmune and nonautoimmune, with beta-cell destruction; type 2 with varying degrees of insulin resistance and insulin hyposecre-

Biomaterials for Tissue Engineering Applications in Diabetes Mellitus

http://dx.doi.org/10.5772/intechopen.69719

411

Alterations in the molecular mechanisms of via insulin signaling may provide a deficiency, resistance, and/or decrease of insulin release into bloodstream, leading to development of the diabetes mellitus. To improve the comprehension of the physiopathology of this comorbidity and its effects in the tissues, firstly, it is important to know the biophysical mechanism involved in the glucose transport and the mechanism of insulin release, as well as the relation-

Fasting blood glucose levels are between 70 and 99 mg/dL in healthy individuals. These values can vary depending on food intake, renal function, physical activity, and the method used for testing. The major dietary polysaccharides, also called carbohydrates, for many species are starch and glycogen which are a storage form of glucose arisen from plants and animals, respectively. After food intake, the digestion process of vegetable (starch) and animal (glycogen) polysaccharides initiates from mouth, which are broken down by the salivary amylase (ptyalin), leading to the formation of oligosaccharides. Afterward, pancreatic and digestive enzymes break down these macromolecules to monosaccharides (glucose, fructose, and galactose) from upper small intestine. Posteriorly, these monosaccharides enter into enterocytes by luminal membrane transporters. It is worth highlighting that the pancreatic amylase is produced by the pancreatic acinar tissue, while the maltase, isomaltase, saccharase, and lactase are generated by the enterocytes. Pancreas is known as a glandular organ in the upper abdomen, which serves as two glands in one: a digestive exocrine gland (e.g., production of pancreatic amylase) and a hormone-producing endocrine gland especially composed of islets Langerhans. These islets are constituted of four hormone-producing cell types: insulin-secreting beta cells, glucagon-secreting alpha cells, somatostatin-secreting delta cells, and pancreatic, polypeptide-

tion; gestational diabetes mellitus; and other types where the cause is known [1].

**3. Development of the diabetes disease**

ship of the glycation and diabetic complications.

**3.1. Biophysical mechanism involved in the glucose transport**

secreting F cells. Both the glands are vital to the body's survival (**Figure 1**).

physiological regulators of insulin release. The glucose molecule (C<sup>6</sup>

Several factors can influence in the insulin secretion, including amino acids, fatty acids, hormones, and drugs; however, the blood glucose concentration levels are considered the main

transporter protein to enter the cells, seeing that is impermeable to cell membrane. Thus, the glucose uptake from intestinal lumen occurs through ion-coupled membrane cotransporters

H12O<sup>6</sup>

) depends on a

Many strategies have been used as a biological dressing, an alternative bioactive material scaffold, and alloplastic material, including amniotic membrane (AM), demineralized dentin matrix (DDM), platelet-rich plasma (PRP), and alloplastic materials as porous polyethylene (porous PTFE) implant, in the degenerative complications in the diabetes disease. The AM is a high-throughput source for multipotent mesenchymal stem cells (MSCs) with the ability to differentiate into wide variety of cells, such as chondroblasts, osteoblasts, adipocytes and fibroblasts, myocytes, endothelial cells, neuronal cells, and hepatocytes, leading to formation of cartilage, bone, connective, muscle, blood vessel, nerve, and liver tissues, respectively. This membrane acts as a barrier prevents the entry of pathogens and toxins, preserves tissue structure, and consequently reduces the levels of local pro-inflammatory cytokines. The DDM is a bioactive tissue able to stimulate the bone repair process, to increase bone mass, and to improve bone microstructure without causing any rejection or infection. It is noteworthy that it acts as a reservoir for various proteins that are embedded within the matrix and are available to local cells, including bone morphogenetic proteins (BMPs), insulin-like growth factors I and II (IGF I and IGF II), platelet-derived growth factor (PDGF), fibroblast growth factor (FGF), transforming growth factor beta (TGF-β), and vascular endothelial growth factor (VEGF).

Based on the literature, we consider that the rich AM and DDM biological properties could be associated with other traditional graft materials, improving their biodynamic effects when implanted in host sites. Some alloplastic materials are well tolerated by the host tissues and present good stability in the implementation site, including porous polyethylene and polyurethane. These materials have been used in orthopedic, articular, orbital, cranial, and maxillofacial reconstruction. The presence of porous allows the ingrowth of the host tissue, which promotes better anchorage of the polymers in the host tissue.

Therefore, these alternative therapies may provide beneficial effects on tissue regeneration and reconstruction, in particular, in diabetes, since one of the main complications is delayed and abnormal repair process by accumulation of advanced glycation endproducts (AGEs) on tissues. These products cause severe widespread damage to tissue through upregulation of inflammation and cross linking of collagen and other proteins.

#### **2. Diabetes mellitus**

Diabetes mellitus is known as a metabolic disorder, characterized by deficiency of insulin secretion or action, leading to chronic hyperglycemia and disturbances of carbohydrate, fat, and protein metabolism, may develop in, called diabetes mellitus (DM). According to the International Diabetes Federation, 8.8% of the adult population worldwide has diabetes. Of all individuals with diabetes, only 10–15% have type 1 diabetes mellitus (T1DM); type 2 diabetes mellitus (T2DM) is the most common form. However, T1DM is the most common form of diabetes in children (<15 years of age), and >500,000 children are currently living with this condition globally. The classification of diabetes mellitus disease defines both process and stage of the disease. The processes include type 1, autoimmune and nonautoimmune, with beta-cell destruction; type 2 with varying degrees of insulin resistance and insulin hyposecretion; gestational diabetes mellitus; and other types where the cause is known [1].

## **3. Development of the diabetes disease**

**1. Introduction**

410 Biomaterials in Regenerative Medicine

Diabetes mellitus (DM), a complex metabolic disorder, is a syndrome characterized by abnormalities in carbohydrate, lipid, and protein metabolism, which results either from a partial or an absolute insulin deficiency or from target tissue resistance to its cellular metabolic effects. This disease is characterized by the presence of few or by the absence of functional pancreatic β-cells in the islets of Langerhans and by a substantial reduction or inexistence of insulin secretion. This cellular dysfunction is an important defect in the pathogenesis of type 2 diabetes [1, 29].

Many strategies have been used as a biological dressing, an alternative bioactive material scaffold, and alloplastic material, including amniotic membrane (AM), demineralized dentin matrix (DDM), platelet-rich plasma (PRP), and alloplastic materials as porous polyethylene (porous PTFE) implant, in the degenerative complications in the diabetes disease. The AM is a high-throughput source for multipotent mesenchymal stem cells (MSCs) with the ability to differentiate into wide variety of cells, such as chondroblasts, osteoblasts, adipocytes and fibroblasts, myocytes, endothelial cells, neuronal cells, and hepatocytes, leading to formation of cartilage, bone, connective, muscle, blood vessel, nerve, and liver tissues, respectively. This membrane acts as a barrier prevents the entry of pathogens and toxins, preserves tissue structure, and consequently reduces the levels of local pro-inflammatory cytokines. The DDM is a bioactive tissue able to stimulate the bone repair process, to increase bone mass, and to improve bone microstructure without causing any rejection or infection. It is noteworthy that it acts as a reservoir for various proteins that are embedded within the matrix and are available to local cells, including bone morphogenetic proteins (BMPs), insulin-like growth factors I and II (IGF I and IGF II), platelet-derived growth factor (PDGF), fibroblast growth factor (FGF), transforming growth factor beta (TGF-β), and vascular endothelial growth factor (VEGF).

Based on the literature, we consider that the rich AM and DDM biological properties could be associated with other traditional graft materials, improving their biodynamic effects when implanted in host sites. Some alloplastic materials are well tolerated by the host tissues and present good stability in the implementation site, including porous polyethylene and polyurethane. These materials have been used in orthopedic, articular, orbital, cranial, and maxillofacial reconstruction. The presence of porous allows the ingrowth of the host tissue, which

Therefore, these alternative therapies may provide beneficial effects on tissue regeneration and reconstruction, in particular, in diabetes, since one of the main complications is delayed and abnormal repair process by accumulation of advanced glycation endproducts (AGEs) on tissues. These products cause severe widespread damage to tissue through upregulation of

Diabetes mellitus is known as a metabolic disorder, characterized by deficiency of insulin secretion or action, leading to chronic hyperglycemia and disturbances of carbohydrate, fat,

promotes better anchorage of the polymers in the host tissue.

inflammation and cross linking of collagen and other proteins.

**2. Diabetes mellitus**

Alterations in the molecular mechanisms of via insulin signaling may provide a deficiency, resistance, and/or decrease of insulin release into bloodstream, leading to development of the diabetes mellitus. To improve the comprehension of the physiopathology of this comorbidity and its effects in the tissues, firstly, it is important to know the biophysical mechanism involved in the glucose transport and the mechanism of insulin release, as well as the relationship of the glycation and diabetic complications.

#### **3.1. Biophysical mechanism involved in the glucose transport**

Fasting blood glucose levels are between 70 and 99 mg/dL in healthy individuals. These values can vary depending on food intake, renal function, physical activity, and the method used for testing. The major dietary polysaccharides, also called carbohydrates, for many species are starch and glycogen which are a storage form of glucose arisen from plants and animals, respectively. After food intake, the digestion process of vegetable (starch) and animal (glycogen) polysaccharides initiates from mouth, which are broken down by the salivary amylase (ptyalin), leading to the formation of oligosaccharides. Afterward, pancreatic and digestive enzymes break down these macromolecules to monosaccharides (glucose, fructose, and galactose) from upper small intestine. Posteriorly, these monosaccharides enter into enterocytes by luminal membrane transporters. It is worth highlighting that the pancreatic amylase is produced by the pancreatic acinar tissue, while the maltase, isomaltase, saccharase, and lactase are generated by the enterocytes. Pancreas is known as a glandular organ in the upper abdomen, which serves as two glands in one: a digestive exocrine gland (e.g., production of pancreatic amylase) and a hormone-producing endocrine gland especially composed of islets Langerhans. These islets are constituted of four hormone-producing cell types: insulin-secreting beta cells, glucagon-secreting alpha cells, somatostatin-secreting delta cells, and pancreatic, polypeptidesecreting F cells. Both the glands are vital to the body's survival (**Figure 1**).

Several factors can influence in the insulin secretion, including amino acids, fatty acids, hormones, and drugs; however, the blood glucose concentration levels are considered the main physiological regulators of insulin release. The glucose molecule (C<sup>6</sup> H12O<sup>6</sup> ) depends on a transporter protein to enter the cells, seeing that is impermeable to cell membrane. Thus, the glucose uptake from intestinal lumen occurs through ion-coupled membrane cotransporters

literature [28]. None of these transport mechanisms requires cell energy expenditure, char-

Considering the intestinal epithelial cells, the glucose entry depends on the electrochemi-

(ECM) through specific channels, located at the basal membrane that maintains pump activity. The sodium-potassium pump is an active transport mechanism which requires energy expenditure using adenosine triphosphate (ATP). The internal cell membrane potential of the enterocytes is negative in relation to the external cell membrane, typically about −70 millivolts. This voltage arises from differences in concentration of the sodium and potassium electrolyte ions, leading to glucose entry into cells. The glucose molecule passes into the extracellular matrix (ECM) through the glucose transporter-2 protein (GLUT-2). Then, it enters in the bloodstream due to the differences of gradient concentration between the ECM and blood capillary (**Figure 2**). Considering the biological property of the blood vessels, intercellular clefts, fenestrations, and cell membrane discontinuity of the endothelial cells also facilitate the

As previous discussion, the elevated levels of blood glucose generate insulin release from pancreatic β-cells at the islets Langerhans, which plays a pivotal role in the glucose homeostasis and acts in a coordinated fashion on cellular events that regulate the metabolic and growth

brane. These differences of extracellular and intracellular concentrations of Na<sup>+</sup>

, as well as its translocation from luminal membrane to basolateral mem-

are kept by

413

) into cells pass to extracellular matrix

http://dx.doi.org/10.5772/intechopen.69719

Biomaterials for Tissue Engineering Applications in Diabetes Mellitus

acterizing a biological process of passive transport.

the sodium-potassium pump. The potassium ions (K<sup>+</sup>

gradient diffusion into bloodstream [24].

**Figure 2.** Diagram showing the glucose absorption by the enterocytes.

**3.2. Mechanism of insulin release**

processes in the human body.

cal gradient of Na<sup>+</sup>

**Figure 1.** Diagram showing carbohydrate degradation by oral, pancreatic, and intestinal enzymes.

which are molecular machines that use the electrochemical energy of a transmembrane ionic gradient to energize the transport of another solute [37]. Among them, it is known as the type 1 sodium-glucose cotransporter (SGLT-1) that is located at the cell membrane of the apical region of the enterocytes. This cotransporter is coded by the SLC5A1 gene, which was the first mammalian cotransporter carrier protein to be identified, cloned, and sequenced [4].

The present review discusses the members of the sodium-glucose cotransporter (SGLT) gene family (the SLC5A gene family) have been identified in a broad range of tissues, including intestine, kidney, muscle, neurons, and thyroid [21, 36]. The SGLT-1 protein has two allosteric sites that carry a galactose or glucose molecule and sodium ions (2Na<sup>+</sup> ) into the cells, while fructose molecule passes across the luminal membrane through the GLUT-5 (fructose transporter). Both the transport proteins are found at the apical surface of the enterocytes in the small intestine [9]. The glucose transporters (GLUTs) are sodium-independent membrane transport proteins which carry particularly monosaccharides, including glucose fructose and galactose. Various members of solute-carrier (SLC) gene family have been found in literature [28]. None of these transport mechanisms requires cell energy expenditure, characterizing a biological process of passive transport.

Considering the intestinal epithelial cells, the glucose entry depends on the electrochemical gradient of Na<sup>+</sup> , as well as its translocation from luminal membrane to basolateral membrane. These differences of extracellular and intracellular concentrations of Na<sup>+</sup> are kept by the sodium-potassium pump. The potassium ions (K<sup>+</sup> ) into cells pass to extracellular matrix (ECM) through specific channels, located at the basal membrane that maintains pump activity. The sodium-potassium pump is an active transport mechanism which requires energy expenditure using adenosine triphosphate (ATP). The internal cell membrane potential of the enterocytes is negative in relation to the external cell membrane, typically about −70 millivolts. This voltage arises from differences in concentration of the sodium and potassium electrolyte ions, leading to glucose entry into cells. The glucose molecule passes into the extracellular matrix (ECM) through the glucose transporter-2 protein (GLUT-2). Then, it enters in the bloodstream due to the differences of gradient concentration between the ECM and blood capillary (**Figure 2**). Considering the biological property of the blood vessels, intercellular clefts, fenestrations, and cell membrane discontinuity of the endothelial cells also facilitate the gradient diffusion into bloodstream [24].

#### **3.2. Mechanism of insulin release**

which are molecular machines that use the electrochemical energy of a transmembrane ionic gradient to energize the transport of another solute [37]. Among them, it is known as the type 1 sodium-glucose cotransporter (SGLT-1) that is located at the cell membrane of the apical region of the enterocytes. This cotransporter is coded by the SLC5A1 gene, which was the first

The present review discusses the members of the sodium-glucose cotransporter (SGLT) gene family (the SLC5A gene family) have been identified in a broad range of tissues, including intestine, kidney, muscle, neurons, and thyroid [21, 36]. The SGLT-1 protein has two allo-

while fructose molecule passes across the luminal membrane through the GLUT-5 (fructose transporter). Both the transport proteins are found at the apical surface of the enterocytes in the small intestine [9]. The glucose transporters (GLUTs) are sodium-independent membrane transport proteins which carry particularly monosaccharides, including glucose fructose and galactose. Various members of solute-carrier (SLC) gene family have been found in

) into the cells,

mammalian cotransporter carrier protein to be identified, cloned, and sequenced [4].

steric sites that carry a galactose or glucose molecule and sodium ions (2Na<sup>+</sup>

**Figure 1.** Diagram showing carbohydrate degradation by oral, pancreatic, and intestinal enzymes.

412 Biomaterials in Regenerative Medicine

As previous discussion, the elevated levels of blood glucose generate insulin release from pancreatic β-cells at the islets Langerhans, which plays a pivotal role in the glucose homeostasis and acts in a coordinated fashion on cellular events that regulate the metabolic and growth processes in the human body.

**Figure 2.** Diagram showing the glucose absorption by the enterocytes.

In humans, the threshold value for glucose-stimulated insulin secretion is about 100 mg/dL (5.6 mmol/L) of plasma glucose concentration. This hormone is secreted in a biphasic manner in response to a marked increase in blood glucose. An initial burst of insulin secretion may last 5 to 15 min, resulting from the secretion of preformed insulin secretory granules. This response is followed by more gradual and sustained insulin secretion that results largely from the synthesis of new insulin molecules.

The primary mechanism is through binding to sulfonylurea receptor (SUR-1) on functioning pancreatic β-cells. Afterward, binding closes the linked ATP-sensitive potassium channels (KATP channels), which leads to decreased potassium influx and subsequent depolarization of the β-cell membrane. This event provides voltage-dependent T-type calcium (Ca2+) and sodium (Na<sup>+</sup> ) channel opening, resulting in an influx of Ca<sup>+</sup> channel, as a consequence, causing translocation and exocytosis of insulin-secretory granules to the cell surface. Then, the insulin from the ECM enters into the bloodstream, following toward the target cells [5].

Another mechanism that promotes ATP-sensitive potassium channel closure is increased ATP levels, which are produced into two distinct metabolic pathways: glycolysis (anaerobic process) and mitochondrial respiratory chain (aerobic process). Initially, the glucose molecule passes across the β-cell membrane through the glucose transporter 2 (GLUT-2), and then, it is phosphorylated to form glucose-6-phospate (G-6-P) by two enzymes: hexokinase IV (HK-IV) or glucokinase (GK) of lower affinity for glucose and hexokinase I (HK-I) of higher affinity for glucose. However, the higher affinity enzyme is strongly inhibited by the G-6-P, transferring to the GK a key role in the glucose phosphorylation at the pancreatic β-cells. This process constitutes the first flux determining step for glycolysis. The G6P is broken down into 2-phosphoenolpyruvate by the pyruvate kinase which is transformed in two pyruvates molecules (glycolysis). After that, these molecules enter into mitochondria where two different biochemical processes occur, including the cycle of Krebs and the oxidative phosphorylation cycle. Finally, several ATP molecules are generated. The glycolysis and mitochondrial respiratory chain pathways result in the synthesis of 2 ATPs and 34–36 ATP molecules, respectively, totalizing about 36–38 ATP molecules/mol of glucose [5, 34]. It is noteworthy to highlight the promotor mechanisms to KATP channels closure lead to the insulin release into the blood stream (**Figure 3**).

Based on these basic concepts, we can consider that methods for SGLT and GLUT inhibition can be considered an suitable alternative means of managing glucose control and, as a conse-

Biomaterials for Tissue Engineering Applications in Diabetes Mellitus

http://dx.doi.org/10.5772/intechopen.69719

415

Persistently elevated glucose levels during long-standing diabetes induce structural and functional changes in various proteins in the body, including plasma protein (albumin, globulins, and fibrinogen) and collagens, through a glycation or nonenzymatic glycosylation process [3, 23, 39, 44]. This process involves an excessive chemical attachment of sugar molecule (e.g., glucose or fructose) to proteins, lipids, and nucleic acids, without the involvement of enzymes. Various deleterious effects from glycation of plasma protein may be produced, including alteration in drug binding in the plasma, platelet activation, generation of oxygen free radicals, impaired fibrinolysis, and impairment in immune system regulation. On the other hand, the structural impairment in collagen alters the osteoblast differentiation, leading

quence, a further treatment option to control the diabetes disease.

**Figure 3.** Mechanism of insulin release from pancreatic β-cell to bloodstream.

**3.3. Relationship of the glycation process and diabetic complications**

After the insulin arrival onto target cells, it links to a specific cell membrane receptor, called insulin receptor substrate (IRS-1) protein, initiating a cascade of phosphorylation events. As a consequence, vesicles containing glucose transporter proteins, especially the GLUT-4, move to the cell surface. The high GLUT-4 level onto cell membrane leads to increase of glucose uptake. Depending on the cell activity, two processes can occur due to high concentration of intracellular glucose: glycolysis or glycogen synthesis. To elucidate better this mechanism, GLUT-4 is an insulin-stimulated glucose transporter that has the primary form of the transporter present in skeletal muscle tissue and adipose tissue. It is present in cells and in intracellular vesicles of the smooth ER. In target cells, the effect of insulin is to promote the translocation of GLUT-4 transporter from intracellular pool to cell membranes. As a result, more transporters are available in the plasma membrane, and glucose uptake by target cells is, thereby, increased.

**Figure 3.** Mechanism of insulin release from pancreatic β-cell to bloodstream.

In humans, the threshold value for glucose-stimulated insulin secretion is about 100 mg/dL (5.6 mmol/L) of plasma glucose concentration. This hormone is secreted in a biphasic manner in response to a marked increase in blood glucose. An initial burst of insulin secretion may last 5 to 15 min, resulting from the secretion of preformed insulin secretory granules. This response is followed by more gradual and sustained insulin secretion that results largely from

The primary mechanism is through binding to sulfonylurea receptor (SUR-1) on functioning pancreatic β-cells. Afterward, binding closes the linked ATP-sensitive potassium channels (KATP channels), which leads to decreased potassium influx and subsequent depolarization of the β-cell membrane. This event provides voltage-dependent T-type calcium (Ca2+) and

ing translocation and exocytosis of insulin-secretory granules to the cell surface. Then, the insulin from the ECM enters into the bloodstream, following toward the target cells [5].

Another mechanism that promotes ATP-sensitive potassium channel closure is increased ATP levels, which are produced into two distinct metabolic pathways: glycolysis (anaerobic process) and mitochondrial respiratory chain (aerobic process). Initially, the glucose molecule passes across the β-cell membrane through the glucose transporter 2 (GLUT-2), and then, it is phosphorylated to form glucose-6-phospate (G-6-P) by two enzymes: hexokinase IV (HK-IV) or glucokinase (GK) of lower affinity for glucose and hexokinase I (HK-I) of higher affinity for glucose. However, the higher affinity enzyme is strongly inhibited by the G-6-P, transferring to the GK a key role in the glucose phosphorylation at the pancreatic β-cells. This process constitutes the first flux determining step for glycolysis. The G6P is broken down into 2-phosphoenolpyruvate by the pyruvate kinase which is transformed in two pyruvates molecules (glycolysis). After that, these molecules enter into mitochondria where two different biochemical processes occur, including the cycle of Krebs and the oxidative phosphorylation cycle. Finally, several ATP molecules are generated. The glycolysis and mitochondrial respiratory chain pathways result in the synthesis of 2 ATPs and 34–36 ATP molecules, respectively, totalizing about 36–38 ATP molecules/mol of glucose [5, 34]. It is noteworthy to highlight the promotor mechanisms to KATP channels closure lead to the insulin release into the blood

After the insulin arrival onto target cells, it links to a specific cell membrane receptor, called insulin receptor substrate (IRS-1) protein, initiating a cascade of phosphorylation events. As a consequence, vesicles containing glucose transporter proteins, especially the GLUT-4, move to the cell surface. The high GLUT-4 level onto cell membrane leads to increase of glucose uptake. Depending on the cell activity, two processes can occur due to high concentration of intracellular glucose: glycolysis or glycogen synthesis. To elucidate better this mechanism, GLUT-4 is an insulin-stimulated glucose transporter that has the primary form of the transporter present in skeletal muscle tissue and adipose tissue. It is present in cells and in intracellular vesicles of the smooth ER. In target cells, the effect of insulin is to promote the translocation of GLUT-4 transporter from intracellular pool to cell membranes. As a result, more transporters are available in the plasma membrane, and glucose uptake by target cells

channel, as a consequence, caus-

) channel opening, resulting in an influx of Ca<sup>+</sup>

the synthesis of new insulin molecules.

414 Biomaterials in Regenerative Medicine

sodium (Na<sup>+</sup>

stream (**Figure 3**).

is, thereby, increased.

Based on these basic concepts, we can consider that methods for SGLT and GLUT inhibition can be considered an suitable alternative means of managing glucose control and, as a consequence, a further treatment option to control the diabetes disease.

#### **3.3. Relationship of the glycation process and diabetic complications**

Persistently elevated glucose levels during long-standing diabetes induce structural and functional changes in various proteins in the body, including plasma protein (albumin, globulins, and fibrinogen) and collagens, through a glycation or nonenzymatic glycosylation process [3, 23, 39, 44]. This process involves an excessive chemical attachment of sugar molecule (e.g., glucose or fructose) to proteins, lipids, and nucleic acids, without the involvement of enzymes. Various deleterious effects from glycation of plasma protein may be produced, including alteration in drug binding in the plasma, platelet activation, generation of oxygen free radicals, impaired fibrinolysis, and impairment in immune system regulation. On the other hand, the structural impairment in collagen alters the osteoblast differentiation, leading to disorders of bone remodeling process and, as a consequence, skeletal fragility [20, 39]. The glycation and oxidation of lipids and nucleotides can also form a single product from various heterogeneous molecules (heterogeneous adducts) through amino-carbonyl reactions, known as advanced glycation endproducts (AGEs) [23, 27]. As the direct products of diabetic metabolic remodeling, AGEs are an important environmental medium for changes of cells, cytokines, and extracellular matrix components including collagens, proteoglycans, laminin, and vitronectin [11, 20]. These changes in the extracellular matrix cause specific alterations in bone formation and bone remodeling. Bone tissue turnover has been shown to be suppressed; studies have shown decreases in the percentage of osteoclasts and osteoblasts as well as a decrease in osteocalcin synthesis [8, 11, 16]. Moreover, AGEs exhibit the ability to inhibit cellular differentiation, proliferation, and migration, which could delay the repair process [20].

**4. Biological and alloplastic materials for tissue engineering**

and structural integrity with the absence of any kind of local infection [17].

polyADP-ribosylation that leads to depletion of cellular NAD<sup>+</sup>

centration cause a rapid destruction of the pancreatic β-cells [41].

mental model were described as follows.

**4.1. Diabetic experimental models**

(**Figure 4**) [11, 13, 42].

The goal of tissue engineering is to create grafts that enhance tissue repair following trauma, infection, or neoplasm, and for developmental abnormalities; these procedures represent a challenge in surgery. To obtain effective and suitable tissue regeneration or osseointegration, the following are necessary: osteoprogenitor and osteoconductive cells, which offer potential to differentiate and facilitate the various stages of tissue regeneration, growth factors (GFs),

Biomaterials for Tissue Engineering Applications in Diabetes Mellitus

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417

As a compensatory factor, some biomaterials for tissue engineering with promotive properties, including amniotic membrane (AM), demineralized dentin matrix (DDM), platelet-rich plasma (PRP), and porous polyethylene (porous PTFE) in diabetic and nondiabetics experi-

The main drugs widely used to induce diabetic experimental models are streptozotocin and alloxan. The cytotoxic action of both the diabetogenic agents is mediated by reactive oxygen species (ROS). The streptozotocin drug enters into pancreatic β-cells via a glucose transporter (GLUT-2) and causes alkylation of DNA. DNA damage induces activation of

dephosphorylation supplies a substrate for xanthine oxidase, resulting in the formation of superoxide radicals as hydrogen peroxide and hydroxyl radicals. Furthermore, toxic amounts of nitric oxide are release, causing inhibition of mitochondrial aconitase activity and DNA damage and triggering the destruction and necrosis of these cells. The alloxan drug promotes the superoxide radical formation, which undergoes the dismutation to hydrogen peroxide. Thereafter, highly reactive hydroxyl radicals are formed by the Fenton reaction. The ROS action and the simultaneous massive increase of cytosolic calcium con-

Experimental models for monohydrate alloxan-induced diabetes in rabbits and rats were developed. The diabetes was diagnosed due to the blood glucose levels above 200 mg/dL (**Figure 4**). This drug, in particular, promotes destruction of the pancreatic β-cells and parenchyma disorganization with presence of atypical acinar cells, scarce secretion vesicles, atrophy, and a decreased number of Langerhans islets in some regions, leading to severe pancreas dysfunction and, as a consequence, the development of diabetes type 2. Three months after the diagnosis of diabetes in experimental models, general clinical complications can be found, including weight loss, polyuria, polyphagia, ketoacidosis due to accentuated hypoglycemia, convulsions, diabetic foot, acute dermatitis, erythematous, and ulcerated lesion associated with fungal and bacterial infections located in the submandibular region extending to trunk and neck regions, and intracranial abscess in regions of surgical defect due to impaired bone repair and susceptibility to infections. Severe chronic gingivitis can be also evidenced

It is important to consider that diabetic wounds exhibit impaired angiogenesis, reduced growth factor levels, and reduced chemotactic ability to recruit inflammatory cells to the

and ATP. Enhanced ATP

Therefore, the protein glycation, elevated level of AGEs accumulation in tissue collagen, and high plasma AGE concentration lead to major diabetic complications like retinopathy, nephropathy, neuropathy, atherosclerosis, degenerative and cardiovascular diseases, inflammatory arthritis, osteoporosis, susceptibility to infections, delayed tissue repair, and severe periodontal disease [3, 17, 39]. Other generic complications are impaired neovascularization and microvascular complications, resulting in ulcers and chronic nonhealing wounds with the characteristics of high amputation rate [20, 46].

Currently, the major concerns with the use of AGEs inhibitors as therapeutic agents are low effectiveness, poor pharmacokinetics, and undesirable side effects. Several AGEs inhibitors possess potent antiglycation activity and are devoid of undesirable side effects. These small molecules inhibitors can, therefore, serve as scaffolds for the development and designing of new AGEs inhibitors as clinical agents [23].

Studies have described that high levels of AGEs induce increased index values of oxidative stress in bone cells due to the absence of the counterbalancing effects of endogenous antioxidants [17]. Thus, it is considered that bone loss can occur as a result of activated osteoclastogenesis, attenuated osteoblastogenesis, and/or stimulated osteoblast apoptosis, which can be induced by the elevated intracellular levels of reactive oxygen species (ROSs) [13, 17]. ROSs produced either endogenously or exogenously are often associated with the principle of oxidative stress, which can attack lipids, proteins, and nucleic acids simultaneously in living cells [45]. This free radical or byproducts of an aerobic metabolism, including superoxide anion (O2 − ), hydrogen peroxide (H2 O2 ), and hydroxyl radicals (OH), can confer reactivity to different biological targets. Increasing evidences have suggested that oxidative stress plays a major role in the pathogenesis of diabetes mellitus due to structural and/or functional damage of the pancreatic β cells [38, 45]. Some studies have described that the persistent hyperglycemia has reduced recruitment of mesenchymal stem cells (MSCs) toward the wounds, compromised their cellular differentiation, and affected the quantity of endothelial progenitor cells, preventing vasculogenesis and delaying the healing process [46]. Accumulating evidence also suggests that the impaired MSCs can compromise mobilization and function of bone cells, resulting in low bone turnover and formation [17].

Recently, alternative material scaffolds have been used to accelerate tissue repair process or to promote implant-based tissue reconstruction, besides the advent of stem cell therapy as a pivotal basis of tissue regeneration or replacing, bringing great hope to diabetic patients.

## **4. Biological and alloplastic materials for tissue engineering**

The goal of tissue engineering is to create grafts that enhance tissue repair following trauma, infection, or neoplasm, and for developmental abnormalities; these procedures represent a challenge in surgery. To obtain effective and suitable tissue regeneration or osseointegration, the following are necessary: osteoprogenitor and osteoconductive cells, which offer potential to differentiate and facilitate the various stages of tissue regeneration, growth factors (GFs), and structural integrity with the absence of any kind of local infection [17].

As a compensatory factor, some biomaterials for tissue engineering with promotive properties, including amniotic membrane (AM), demineralized dentin matrix (DDM), platelet-rich plasma (PRP), and porous polyethylene (porous PTFE) in diabetic and nondiabetics experimental model were described as follows.

#### **4.1. Diabetic experimental models**

to disorders of bone remodeling process and, as a consequence, skeletal fragility [20, 39]. The glycation and oxidation of lipids and nucleotides can also form a single product from various heterogeneous molecules (heterogeneous adducts) through amino-carbonyl reactions, known as advanced glycation endproducts (AGEs) [23, 27]. As the direct products of diabetic metabolic remodeling, AGEs are an important environmental medium for changes of cells, cytokines, and extracellular matrix components including collagens, proteoglycans, laminin, and vitronectin [11, 20]. These changes in the extracellular matrix cause specific alterations in bone formation and bone remodeling. Bone tissue turnover has been shown to be suppressed; studies have shown decreases in the percentage of osteoclasts and osteoblasts as well as a decrease in osteocalcin synthesis [8, 11, 16]. Moreover, AGEs exhibit the ability to inhibit cellular differentiation, proliferation, and migration, which could delay the repair process [20]. Therefore, the protein glycation, elevated level of AGEs accumulation in tissue collagen, and high plasma AGE concentration lead to major diabetic complications like retinopathy, nephropathy, neuropathy, atherosclerosis, degenerative and cardiovascular diseases, inflammatory arthritis, osteoporosis, susceptibility to infections, delayed tissue repair, and severe periodontal disease [3, 17, 39]. Other generic complications are impaired neovascularization and microvascular complications, resulting in ulcers and chronic nonhealing wounds with

Currently, the major concerns with the use of AGEs inhibitors as therapeutic agents are low effectiveness, poor pharmacokinetics, and undesirable side effects. Several AGEs inhibitors possess potent antiglycation activity and are devoid of undesirable side effects. These small molecules inhibitors can, therefore, serve as scaffolds for the development and designing of

Studies have described that high levels of AGEs induce increased index values of oxidative stress in bone cells due to the absence of the counterbalancing effects of endogenous antioxidants [17]. Thus, it is considered that bone loss can occur as a result of activated osteoclastogenesis, attenuated osteoblastogenesis, and/or stimulated osteoblast apoptosis, which can be induced by the elevated intracellular levels of reactive oxygen species (ROSs) [13, 17]. ROSs produced either endogenously or exogenously are often associated with the principle of oxidative stress, which can attack lipids, proteins, and nucleic acids simultaneously in living cells [45]. This free radical or byproducts of an aerobic metabolism, including superoxide anion

ent biological targets. Increasing evidences have suggested that oxidative stress plays a major role in the pathogenesis of diabetes mellitus due to structural and/or functional damage of the pancreatic β cells [38, 45]. Some studies have described that the persistent hyperglycemia has reduced recruitment of mesenchymal stem cells (MSCs) toward the wounds, compromised their cellular differentiation, and affected the quantity of endothelial progenitor cells, preventing vasculogenesis and delaying the healing process [46]. Accumulating evidence also suggests that the impaired MSCs can compromise mobilization and function of bone cells,

Recently, alternative material scaffolds have been used to accelerate tissue repair process or to promote implant-based tissue reconstruction, besides the advent of stem cell therapy as a pivotal basis of tissue regeneration or replacing, bringing great hope to diabetic patients.

), and hydroxyl radicals (OH), can confer reactivity to differ-

the characteristics of high amputation rate [20, 46].

new AGEs inhibitors as clinical agents [23].

416 Biomaterials in Regenerative Medicine

), hydrogen peroxide (H2

O2

resulting in low bone turnover and formation [17].

(O2 − The main drugs widely used to induce diabetic experimental models are streptozotocin and alloxan. The cytotoxic action of both the diabetogenic agents is mediated by reactive oxygen species (ROS). The streptozotocin drug enters into pancreatic β-cells via a glucose transporter (GLUT-2) and causes alkylation of DNA. DNA damage induces activation of polyADP-ribosylation that leads to depletion of cellular NAD<sup>+</sup> and ATP. Enhanced ATP dephosphorylation supplies a substrate for xanthine oxidase, resulting in the formation of superoxide radicals as hydrogen peroxide and hydroxyl radicals. Furthermore, toxic amounts of nitric oxide are release, causing inhibition of mitochondrial aconitase activity and DNA damage and triggering the destruction and necrosis of these cells. The alloxan drug promotes the superoxide radical formation, which undergoes the dismutation to hydrogen peroxide. Thereafter, highly reactive hydroxyl radicals are formed by the Fenton reaction. The ROS action and the simultaneous massive increase of cytosolic calcium concentration cause a rapid destruction of the pancreatic β-cells [41].

Experimental models for monohydrate alloxan-induced diabetes in rabbits and rats were developed. The diabetes was diagnosed due to the blood glucose levels above 200 mg/dL (**Figure 4**). This drug, in particular, promotes destruction of the pancreatic β-cells and parenchyma disorganization with presence of atypical acinar cells, scarce secretion vesicles, atrophy, and a decreased number of Langerhans islets in some regions, leading to severe pancreas dysfunction and, as a consequence, the development of diabetes type 2. Three months after the diagnosis of diabetes in experimental models, general clinical complications can be found, including weight loss, polyuria, polyphagia, ketoacidosis due to accentuated hypoglycemia, convulsions, diabetic foot, acute dermatitis, erythematous, and ulcerated lesion associated with fungal and bacterial infections located in the submandibular region extending to trunk and neck regions, and intracranial abscess in regions of surgical defect due to impaired bone repair and susceptibility to infections. Severe chronic gingivitis can be also evidenced (**Figure 4**) [11, 13, 42].

It is important to consider that diabetic wounds exhibit impaired angiogenesis, reduced growth factor levels, and reduced chemotactic ability to recruit inflammatory cells to the

to clinical applications in regenerative medicine, which leads to processes of tissue replacing,

Biomaterials for Tissue Engineering Applications in Diabetes Mellitus

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419

Several researches developed in the Center of Biosciences Applied to Persons with Special Care Needs (CEBAPE) in the São Paulo State University (UNESP) has applied the AM as a biologic dressing and a bioactive scaffold material to treat oral mucositis, surgical hepatic resection areas in experiments *in vivo*, and wound areas from incisional and excisional biopsies of oral lesions originated of inflammatory and development processes. This membrane was also applied to stimulate the guided bone regeneration (GBR) in experimental bone

The methods of preparation and storage of the human (h) and homogenous (H) AM were developed by the CEBAPE which are described in the literature [15, 18, 26, 43]. Additionally, the clinical procedures were performed in diabetic and nondiabetic patients or in different animal models, leading to highly favorable results that are reported as follows. We can prove its effects on the tissues through the clinical, biochemical, histological and immunohistochemical outcomes, in particular, in experimental models. As a xenogenic material, a human amniotic membrane (hAM) was used for guided bone regenerative on surgical defects in healthy rabbits. The biocompatibility and low immunogenicity of the hAM were shown since we evidenced newly formed bone tissue in intimate contact with hAM in several surgical defect areas and discrete infiltration of mononuclear inflammatory cells surrounding the hAM in some regions; moreover, its osteogenic and mitogenic effects were very evident (**Figure 5**).

Other application could be in oral biopsy areas to lesion diagnosis, especially in areas of difficult suture as alveolar bone ridge in maxilla and mandible. Studies also showed the use of hAM on wounds produced from incisional and excisional biopsies of alveolar gingiva of the upper premolar region and cheek mucosa, respectively. The diagnosis was plan lichen (inflammatory process) in the alveolar gingiva and intramucosal pigmented nevus (development process). The outcomes were highly satisfactory in both processes, inducing the rapid closure of

**Figure 5.** Human amniotic membrane (arrow-head) used as a bioactive xenogenic material which was applied in the guided bone regeneration on the surgical defect in rabbits: newly formed bone formation from the membrane and few infiltrations of mononuclear inflammatory cells, especially lymphocytes, were evidenced (Source: Gomes et al., [15]).

engineering or regenerating, and preservation of organ functions [12, 26, 43].

defects and to accelerate the tissue repair.

**Figure 4.** Classical clinical complications in diabetes showing diabetic foot and chronic gingivitis (arrows) in rabbit after 3 months persistent hyperglycemia.

wound, besides that the poor vascularization and maintenance of a chronic inflammatory state limit the healing capacity. Therefore, we discuss in this chapter some conductive and inductive materials for tissue engineering, including AM, DDM, PRP, and porous PTFE, which were applied in diabetic and nondiabetic experimental models or individuals.

#### **4.2. Biological materials**

#### *4.2.1. Amniotic membrane (AM)*

The AM is a tissue of fetal origin and is composed of three major layers: a single epithelial layer, a thick basement membrane, and an avascular mesenchyme that contain growth factors, cytokines, and other active substances [26]. AM has two types of cells of different embryological origins: amnion epithelial cells derived from embryonic ectoderm and amnion mesenchymal cells from embryonic mesoderm. It is adjacent to the trophoblast cells and lines the amniotic cavity. It can be easily separated from the underlying chorion, with which it never truly fuses at the cellular level. Its nutrition and oxygen are obtained from surrounding chorionic fluid, the amniotic fluid, and the fetal surface vessels. This structure protects the developing embryo against mechanical aggression, provides an environment where the embryo can grow without distortion by pressure from surrounding structures, and also plays an important role during parturition [15].

As a highlight of great potential for clinical application, the amniotic membrane has been widely used in regenerative medicine due to its effective biological properties (biocompatibility, low immunogenicity, anti-inflammatory, and antimicrobial activities), adequate physicomechanical action (permeability, stability, elasticity, flexibility, and resorbability), excellent tissue adhesion, easy delivery of biomodulatory agents including growth factors and genetic materials, and rich MSC source [22, 31, 33]. Current investigations described that pluripotent stem cell lines derived from AM have the ability to differentiate into endothelial and neuronal cells, osteogenic, chondrogenic, adipogenic, skeletal myogenic, hepatic lineages, and able suppress T-cell proliferation [19, 31, 35]. These properties show the greatest prospects to clinical applications in regenerative medicine, which leads to processes of tissue replacing, engineering or regenerating, and preservation of organ functions [12, 26, 43].

Several researches developed in the Center of Biosciences Applied to Persons with Special Care Needs (CEBAPE) in the São Paulo State University (UNESP) has applied the AM as a biologic dressing and a bioactive scaffold material to treat oral mucositis, surgical hepatic resection areas in experiments *in vivo*, and wound areas from incisional and excisional biopsies of oral lesions originated of inflammatory and development processes. This membrane was also applied to stimulate the guided bone regeneration (GBR) in experimental bone defects and to accelerate the tissue repair.

The methods of preparation and storage of the human (h) and homogenous (H) AM were developed by the CEBAPE which are described in the literature [15, 18, 26, 43]. Additionally, the clinical procedures were performed in diabetic and nondiabetic patients or in different animal models, leading to highly favorable results that are reported as follows. We can prove its effects on the tissues through the clinical, biochemical, histological and immunohistochemical outcomes, in particular, in experimental models. As a xenogenic material, a human amniotic membrane (hAM) was used for guided bone regenerative on surgical defects in healthy rabbits. The biocompatibility and low immunogenicity of the hAM were shown since we evidenced newly formed bone tissue in intimate contact with hAM in several surgical defect areas and discrete infiltration of mononuclear inflammatory cells surrounding the hAM in some regions; moreover, its osteogenic and mitogenic effects were very evident (**Figure 5**).

wound, besides that the poor vascularization and maintenance of a chronic inflammatory state limit the healing capacity. Therefore, we discuss in this chapter some conductive and inductive materials for tissue engineering, including AM, DDM, PRP, and porous PTFE,

**Figure 4.** Classical clinical complications in diabetes showing diabetic foot and chronic gingivitis (arrows) in rabbit after

The AM is a tissue of fetal origin and is composed of three major layers: a single epithelial layer, a thick basement membrane, and an avascular mesenchyme that contain growth factors, cytokines, and other active substances [26]. AM has two types of cells of different embryological origins: amnion epithelial cells derived from embryonic ectoderm and amnion mesenchymal cells from embryonic mesoderm. It is adjacent to the trophoblast cells and lines the amniotic cavity. It can be easily separated from the underlying chorion, with which it never truly fuses at the cellular level. Its nutrition and oxygen are obtained from surrounding chorionic fluid, the amniotic fluid, and the fetal surface vessels. This structure protects the developing embryo against mechanical aggression, provides an environment where the embryo can grow without distortion by pressure from surrounding structures, and also plays

As a highlight of great potential for clinical application, the amniotic membrane has been widely used in regenerative medicine due to its effective biological properties (biocompatibility, low immunogenicity, anti-inflammatory, and antimicrobial activities), adequate physicomechanical action (permeability, stability, elasticity, flexibility, and resorbability), excellent tissue adhesion, easy delivery of biomodulatory agents including growth factors and genetic materials, and rich MSC source [22, 31, 33]. Current investigations described that pluripotent stem cell lines derived from AM have the ability to differentiate into endothelial and neuronal cells, osteogenic, chondrogenic, adipogenic, skeletal myogenic, hepatic lineages, and able suppress T-cell proliferation [19, 31, 35]. These properties show the greatest prospects

which were applied in diabetic and nondiabetic experimental models or individuals.

**4.2. Biological materials**

3 months persistent hyperglycemia.

418 Biomaterials in Regenerative Medicine

*4.2.1. Amniotic membrane (AM)*

an important role during parturition [15].

Other application could be in oral biopsy areas to lesion diagnosis, especially in areas of difficult suture as alveolar bone ridge in maxilla and mandible. Studies also showed the use of hAM on wounds produced from incisional and excisional biopsies of alveolar gingiva of the upper premolar region and cheek mucosa, respectively. The diagnosis was plan lichen (inflammatory process) in the alveolar gingiva and intramucosal pigmented nevus (development process). The outcomes were highly satisfactory in both processes, inducing the rapid closure of

**Figure 5.** Human amniotic membrane (arrow-head) used as a bioactive xenogenic material which was applied in the guided bone regeneration on the surgical defect in rabbits: newly formed bone formation from the membrane and few infiltrations of mononuclear inflammatory cells, especially lymphocytes, were evidenced (Source: Gomes et al., [15]).

wounds due to stimulation of the tissue repair, relief of pain, and no sign of local infection [18]. Probably, relief of pain occurs due to its strong action of mechanism barrier and good stability on the wound surface, covering nerve endings.

In previous research using homogenous amniotic membrane (HAM) on the surface of surgical liver, resection in rats was performed. The HAM was obtained female rats by cesarean at the 20th day of pregnancy. Fetal amniotic membrane was separated aseptically from the chorionic membrane and rinsed several times in a sterile physiological saline solution and a phosphate buffer at pH 7.4 until the removal of all debris (**Figure 6**). After its laboratory preparation, the HAM was placed on the surface of surgical liver resection. Clinically, this membrane showed an immediate stability and a strong hemostatic action after its implantation on the wound. Histologically, the HAM behavior was observed in 10, 20, 30, and 40 days, showing the following results. Initially, the HAM was biocompatible to hepatic tissue once it intimately adhered to hepatic tissue in regenerative process; subsequently, it underwent metaplasia turning into a highly well-cellularized and vascularized tissue called aminohepatic tissue. Discrete infiltration of mononuclear inflammatory cells was seen surrounding this biological membrane. This histological characteristic was one of the prominent aspects that can suggest its low immunogenicity in surgically impaired liver tissue. HAM remnants can be found on the injured liver tissue surface in some specimens, after 40-day postsurgery (**Figure 7**). In biochemical analysis of plasma and tissue samples, the aspartate aminotransferase (AST), alanine aminotransferase (ALT), and gamma-glutamyltransferase (GGT) enzymatic levels were significantly low; in contrast, the alkaline phosphatase (ALP) enzymatic levels and the total protein concentrations were strongly high in the rats treated with HAM. Thus, angiogenic, mitogenic, and hepatic-protector effects of the HAM were proved, leading to a better microstructural reconstruction of the injured liver [12].

In our study, other beneficial application of the HAM was as a potent biological dressing in 5-fluorouracil-induced oral mucosites (OM) in rats (**Figure 8**). The pathophysiology of the analyzed OM and its histomorphological features were divided into four phases: inflammation (3 days), cell proliferation (7 days), tissue organization (14 days), and tissue repair (21 days). Among the periods of inflammation and cell proliferation phases, the HAM was incorporated with the wound surface, acting as a protective barrier against microbial infection and preventing any infection signs. Other aspects were discrete to moderate infiltration of inflammatory cells in the connective tissue and early total re-epithelialization in most specimens. The incorporated HAM into connective tissue was degraded and, then resorbed. Regarding the tissue organization and repair phases, the histological features were described the follow. At 14 days, discrete mononuclear inflammatory cells surrounding the connective tissue, muscle layer with normal aspect, a tissue maturation with regularly disposed, remodeled collagen fibers, and rare blood vessels were found. Finally, at 21 days, the oral mucosa showed normal aspects (**Figure 9**). Thus, we can consider that the HAM provides an excellent environment for cell proliferation and neovascularization, stimulating healing process and functioning as a promising biological dressing. Previous studies confirm the growth factors are found into amniotic membrane, including transforming growth factor-β1 and β2 (TGF-β1 and β2), basic fibroblast growth factor (bFGF), epidermal growth factor (EGF), transforming growth factor-α (TGF-α), keratinocyte growth factor (KGF), and hepatocyte growth factor (HGF) [33], reinforcing our findings. Regarding the biodegradable property of the HAM, we hypothesize that the avascular stromal matrix of the HAM is similar

extracellular matrix of the connective tissue in the submucosa region, favoring its degradation during the healing dynamic processes. Investigations report the AM presence of fibronectin, elas-

**Figure 6.** (a) Fetus, amniotic, and chorionic membrane (arrow) and (b) fresh HAM separated from the chorionic

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Homogenous demineralized dentin matrix (HDDM) arises from tooth structure which has been widely used to stimulate bone regeneration, inducing osteoblastic differentiation and proliferation and chemotaxis. Hence, HDDM accelerates the bone repair process, increases bone mass, and improves bone quality without causing any rejection or infection. HDDM

tin, nidogen, collagen types I, III, IV, V, and VI, elastin, and hyaluronic acid [31, 33].

membrane and placed in sterile physiological saline solution and phosphate buffer at pH 7.4.

*4.2.2. Homogenous demineralized dentin matrix (HDDM)*

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**Figure 6.** (a) Fetus, amniotic, and chorionic membrane (arrow) and (b) fresh HAM separated from the chorionic membrane and placed in sterile physiological saline solution and phosphate buffer at pH 7.4.

extracellular matrix of the connective tissue in the submucosa region, favoring its degradation during the healing dynamic processes. Investigations report the AM presence of fibronectin, elastin, nidogen, collagen types I, III, IV, V, and VI, elastin, and hyaluronic acid [31, 33].

#### *4.2.2. Homogenous demineralized dentin matrix (HDDM)*

wounds due to stimulation of the tissue repair, relief of pain, and no sign of local infection [18]. Probably, relief of pain occurs due to its strong action of mechanism barrier and good stability

In previous research using homogenous amniotic membrane (HAM) on the surface of surgical liver, resection in rats was performed. The HAM was obtained female rats by cesarean at the 20th day of pregnancy. Fetal amniotic membrane was separated aseptically from the chorionic membrane and rinsed several times in a sterile physiological saline solution and a phosphate buffer at pH 7.4 until the removal of all debris (**Figure 6**). After its laboratory preparation, the HAM was placed on the surface of surgical liver resection. Clinically, this membrane showed an immediate stability and a strong hemostatic action after its implantation on the wound. Histologically, the HAM behavior was observed in 10, 20, 30, and 40 days, showing the following results. Initially, the HAM was biocompatible to hepatic tissue once it intimately adhered to hepatic tissue in regenerative process; subsequently, it underwent metaplasia turning into a highly well-cellularized and vascularized tissue called aminohepatic tissue. Discrete infiltration of mononuclear inflammatory cells was seen surrounding this biological membrane. This histological characteristic was one of the prominent aspects that can suggest its low immunogenicity in surgically impaired liver tissue. HAM remnants can be found on the injured liver tissue surface in some specimens, after 40-day postsurgery (**Figure 7**). In biochemical analysis of plasma and tissue samples, the aspartate aminotransferase (AST), alanine aminotransferase (ALT), and gamma-glutamyltransferase (GGT) enzymatic levels were significantly low; in contrast, the alkaline phosphatase (ALP) enzymatic levels and the total protein concentrations were strongly high in the rats treated with HAM. Thus, angiogenic, mitogenic, and hepatic-protector effects of the HAM were proved, leading

In our study, other beneficial application of the HAM was as a potent biological dressing in 5-fluorouracil-induced oral mucosites (OM) in rats (**Figure 8**). The pathophysiology of the analyzed OM and its histomorphological features were divided into four phases: inflammation (3 days), cell proliferation (7 days), tissue organization (14 days), and tissue repair (21 days). Among the periods of inflammation and cell proliferation phases, the HAM was incorporated with the wound surface, acting as a protective barrier against microbial infection and preventing any infection signs. Other aspects were discrete to moderate infiltration of inflammatory cells in the connective tissue and early total re-epithelialization in most specimens. The incorporated HAM into connective tissue was degraded and, then resorbed. Regarding the tissue organization and repair phases, the histological features were described the follow. At 14 days, discrete mononuclear inflammatory cells surrounding the connective tissue, muscle layer with normal aspect, a tissue maturation with regularly disposed, remodeled collagen fibers, and rare blood vessels were found. Finally, at 21 days, the oral mucosa showed normal aspects (**Figure 9**). Thus, we can consider that the HAM provides an excellent environment for cell proliferation and neovascularization, stimulating healing process and functioning as a promising biological dressing. Previous studies confirm the growth factors are found into amniotic membrane, including transforming growth factor-β1 and β2 (TGF-β1 and β2), basic fibroblast growth factor (bFGF), epidermal growth factor (EGF), transforming growth factor-α (TGF-α), keratinocyte growth factor (KGF), and hepatocyte growth factor (HGF) [33], reinforcing our findings. Regarding the biodegradable property of the HAM, we hypothesize that the avascular stromal matrix of the HAM is similar

on the wound surface, covering nerve endings.

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to a better microstructural reconstruction of the injured liver [12].

Homogenous demineralized dentin matrix (HDDM) arises from tooth structure which has been widely used to stimulate bone regeneration, inducing osteoblastic differentiation and proliferation and chemotaxis. Hence, HDDM accelerates the bone repair process, increases bone mass, and improves bone quality without causing any rejection or infection. HDDM

control of HDDM growth factors, including BMPs [11, 14, 15]. Others growth factors include

**Figure 8.** Oral mucositis in labial fornix region of inferior incisors (a), HAM placed on the wound (b), and normal oral

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HDDM slices strongly favor the chemotaxis of appropriate cells into the surgical defect, the transformation of undifferentiated mesenchymal cells into osteoprogenitor cells, osteoblast proliferation and differentiation, and the synthesis of the extracellular matrix with mineralization of osteoid tissue that leads to accelerated bone maturation and remodeling. It is worth highlighting that the HDDM slices initially yielded hemostatic effects by mechanical action after their implantation in the periphery of the defect. Furthermore, it was demonstrated that the HDDM was incorporated into the newly formed bone matrix and resorbed during the bone remodeling process. Moreover, numerous osteoblasts and osteoprogenitor cells were also located adjacent to the HDDM, confirming its strong chemotactic activity. Studies suggested that HDDM provides bioactive molecules for the host of proteins and growth factors can be released through dentin matrix degradation and exposed dentinal tubules when HDDM is used in the form of slices [6, 11, 14, 15, 17]. Furthermore, when this biomaterial is used in the periphery of surgical defect, it is able to stimulate the reconstruction of the bone architectural microstructure,

**Figure 9.** 5-Fluorouracil-induced oral mucosites. Histomorphological features divided into four phases: (a) inflammation (3 days), (b) cell proliferation (7 days), (c) tissue organization (14 days), and (d) tissue repair (21 days). Asterisk: HAM remnants incorporated into connective tissue in the submucosa region (proliferation phase) (hematoxylin-eosin,

TGF-β, IGF I and IGF II, PDGF, FGF, TGF-β, and VEGF [17].

mucosa surface after 7 days of treatment with HAM (c).

magnification ×25).

**Figure 7.** HAM placed on the surgically induced liver wound, HAM intimately adhered to the surface of the injured liver tissue and underwent metaplasia turning into amnion-hepatic tissue in 10 days post-surgery, and remnants of HAM on liver surface in the final step of the liver tissue regeneration in 40 days post-surgery. (hematoxylin-eosin, magnification ×25).

slices comprise an extracellular matrix rich in collagen, without vessels, and act as a reservoir for various proteins that are embedded within the matrix and are available to local cells. Regarding its main property, it is proved that the osteoinductive cascade begins with chemotaxis of bone progenitor cells, mitogenesis, angiogenesis, and bone cell differentiation. The cell recruitment, division rate, and differentiation of these cell lines are under the direct

**Figure 8.** Oral mucositis in labial fornix region of inferior incisors (a), HAM placed on the wound (b), and normal oral mucosa surface after 7 days of treatment with HAM (c).

control of HDDM growth factors, including BMPs [11, 14, 15]. Others growth factors include TGF-β, IGF I and IGF II, PDGF, FGF, TGF-β, and VEGF [17].

HDDM slices strongly favor the chemotaxis of appropriate cells into the surgical defect, the transformation of undifferentiated mesenchymal cells into osteoprogenitor cells, osteoblast proliferation and differentiation, and the synthesis of the extracellular matrix with mineralization of osteoid tissue that leads to accelerated bone maturation and remodeling. It is worth highlighting that the HDDM slices initially yielded hemostatic effects by mechanical action after their implantation in the periphery of the defect. Furthermore, it was demonstrated that the HDDM was incorporated into the newly formed bone matrix and resorbed during the bone remodeling process. Moreover, numerous osteoblasts and osteoprogenitor cells were also located adjacent to the HDDM, confirming its strong chemotactic activity. Studies suggested that HDDM provides bioactive molecules for the host of proteins and growth factors can be released through dentin matrix degradation and exposed dentinal tubules when HDDM is used in the form of slices [6, 11, 14, 15, 17]. Furthermore, when this biomaterial is used in the periphery of surgical defect, it is able to stimulate the reconstruction of the bone architectural microstructure,

**Figure 9.** 5-Fluorouracil-induced oral mucosites. Histomorphological features divided into four phases: (a) inflammation (3 days), (b) cell proliferation (7 days), (c) tissue organization (14 days), and (d) tissue repair (21 days). Asterisk: HAM remnants incorporated into connective tissue in the submucosa region (proliferation phase) (hematoxylin-eosin, magnification ×25).

slices comprise an extracellular matrix rich in collagen, without vessels, and act as a reservoir for various proteins that are embedded within the matrix and are available to local cells. Regarding its main property, it is proved that the osteoinductive cascade begins with chemotaxis of bone progenitor cells, mitogenesis, angiogenesis, and bone cell differentiation. The cell recruitment, division rate, and differentiation of these cell lines are under the direct

**Figure 7.** HAM placed on the surgically induced liver wound, HAM intimately adhered to the surface of the injured liver tissue and underwent metaplasia turning into amnion-hepatic tissue in 10 days post-surgery, and remnants of HAM on liver surface in the final step of the liver tissue regeneration in 40 days post-surgery. (hematoxylin-eosin, magnification ×25).

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resulting in formation of mature bone trabeculae and bone marrow. Initially, bone trabeculae were mature at the periphery and immature at the centre of the bone defect; thus, centripetal bone growth was seen in the bone repair process (**Figures 10**, **11**). Finally, an exophytic bone growth can be evidenced due to the highest chemotactic, mitogenic, and osteogenic potential of the HDDM. Dependent of the animal's age, a physiological conversion from red to yellow bone marrow can be observed, leading to large and regular medullary spaces (**Figure 12**).

Concerning the evidence of HDDM biocompatibility, two hypotheses are suggested by the present study. First, cell surface histocompatibility antigens and sequestered antigens are absent; however, they may be present throughout the extension of the remnant odontoblastic processes within the HDDM. Second, the decalcification process for teeth could cause the denaturation of surface proteins on the plasma membrane of the odontoblastic processes. It was considered that the plasma membrane of the odontoblastic processes could present not only cell surface histocompatibility antigens, but also sequestered antigens that, when not exposed to the cells of the immune system, do not trigger immunological reactions.

The resorption process of the HDDM slices presented the following phases: (1) the dentin matrix was first incorporated into the newly formed bone tissue; (2) next, the dentin was degraded during bone remodeling; and (3) the area in which the dentin matrix was located was replaced by new bone tissue (**Figure 13**).

Therefore, the HDDM could be considered as an ideal bone graft material since it stimulates osteoblast differentiation and proliferation and potentially leads to efficient bone regeneration. We suggested that it could be used as a scaffold for stem cells and bone growth factors, since its biological properties have shown an excellent bioactive material for bone tissue engineering in diabetes.

Autogenous demineralized dentin matrix (ADDM) has been used to accelerate the dental socket wound healing process in humans. The ADDM is considerate biocompatible with surgical human dental sockets. Initially, radiopaque image suggestive of remaining ADDM slices can be found. Moreover, the radiographic bone density of the dental sockets treated with ADDM was similar to that of the surrounding normal bone from 90th day. It is worth highlighting that alveolar bone architecture is improved when the ADDM is used as an engineering tissue (**Figure 14**).

**Figure 11.** Performance of the HDDM inside the surgical bone defect after 15 days in diabetic rabbit: (a) intense osteogenic connective tissue showing numerous osteoblasts (asterisk) showing chemotactic activity; (b) HDDM slices in resorption process by osteoclasts (OC), resorbed HDDM slices (blue arrow), and HDDM slices incorporated into the newly formed

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**Figure 12.** Radiographic image of surgical defect in the parietal bone of diabetic rabbits taken at 90 days and defect region completely filled with mature bone trabeculae a bone marrow (hematoxylin-eosin, magnification ×25) (Source:

bone tissue (black arrows) (hematoxylin-eosin, magnification ×200).

Gomes et al. [17]).

**Figure 10.** Clinical and radiographic images of surgical defect in the parietal bone of diabetic rabbits taken at 15 days. HDDM slices promoting hemostatic effects by mechanical action after their implantation and their osteogenic activity after 15 days in diabetic rabbit (H.E., original magnification ×25) (Source: Gomes et al. [17]).

resulting in formation of mature bone trabeculae and bone marrow. Initially, bone trabeculae were mature at the periphery and immature at the centre of the bone defect; thus, centripetal bone growth was seen in the bone repair process (**Figures 10**, **11**). Finally, an exophytic bone growth can be evidenced due to the highest chemotactic, mitogenic, and osteogenic potential of the HDDM. Dependent of the animal's age, a physiological conversion from red to yellow bone

Concerning the evidence of HDDM biocompatibility, two hypotheses are suggested by the present study. First, cell surface histocompatibility antigens and sequestered antigens are absent; however, they may be present throughout the extension of the remnant odontoblastic processes within the HDDM. Second, the decalcification process for teeth could cause the denaturation of surface proteins on the plasma membrane of the odontoblastic processes. It was considered that the plasma membrane of the odontoblastic processes could present not only cell surface histocompatibility antigens, but also sequestered antigens that, when not

The resorption process of the HDDM slices presented the following phases: (1) the dentin matrix was first incorporated into the newly formed bone tissue; (2) next, the dentin was degraded during bone remodeling; and (3) the area in which the dentin matrix was located

Therefore, the HDDM could be considered as an ideal bone graft material since it stimulates osteoblast differentiation and proliferation and potentially leads to efficient bone regeneration. We suggested that it could be used as a scaffold for stem cells and bone growth factors, since its biological properties have shown an excellent bioactive material for bone tissue engineering in diabetes.

Autogenous demineralized dentin matrix (ADDM) has been used to accelerate the dental socket wound healing process in humans. The ADDM is considerate biocompatible with surgical human dental sockets. Initially, radiopaque image suggestive of remaining ADDM slices can be found. Moreover, the radiographic bone density of the dental sockets treated with ADDM was similar to that of the surrounding normal bone from 90th day. It is worth highlighting that alveolar bone architecture is improved when the ADDM is used as an engineering tissue (**Figure 14**).

**Figure 10.** Clinical and radiographic images of surgical defect in the parietal bone of diabetic rabbits taken at 15 days. HDDM slices promoting hemostatic effects by mechanical action after their implantation and their osteogenic activity

after 15 days in diabetic rabbit (H.E., original magnification ×25) (Source: Gomes et al. [17]).

marrow can be observed, leading to large and regular medullary spaces (**Figure 12**).

exposed to the cells of the immune system, do not trigger immunological reactions.

was replaced by new bone tissue (**Figure 13**).

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**Figure 11.** Performance of the HDDM inside the surgical bone defect after 15 days in diabetic rabbit: (a) intense osteogenic connective tissue showing numerous osteoblasts (asterisk) showing chemotactic activity; (b) HDDM slices in resorption process by osteoclasts (OC), resorbed HDDM slices (blue arrow), and HDDM slices incorporated into the newly formed bone tissue (black arrows) (hematoxylin-eosin, magnification ×200).

**Figure 12.** Radiographic image of surgical defect in the parietal bone of diabetic rabbits taken at 90 days and defect region completely filled with mature bone trabeculae a bone marrow (hematoxylin-eosin, magnification ×25) (Source: Gomes et al. [17]).

**Figure 13.** Surgical bone defect region after 90 days in diabetic rabbit. Phases of the HDDM degradation during the bone remodeling process: HDDM in degradation (white arrows), exchange of HDDM by well-cellularized and immature bone tissue with remnants of dentin matrix (asterisk), and replacement of the preexisting HDDM by mature bone tissue (yellow arrows) (Schmorl's stain, magnification ×200).

recruit resident stem cells to the site of injury, where they are stimulated to secrete additional growth factors and anti-inflammatory cytokines, causing more increase in collagen and matrix synthesis. Furthermore, the recruited stem cells will react with the environment to differentiate

**Figure 14.** ADDM cut into slices with frozen microtomy (a) and placed into human dental socket covering its surface (b). Periapical radiographs after 15 (c) and 90 days (d) post-surgery. Arrow: radiopaque image suggests remaining of

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Some investigators have reported that platelet-rich plasma (PRP) becomes more efficient in tissue repair when associated with bone grafts, bioactive material scaffold, or mesenchymal stem cells (MSCs) which are recruited at the wound site, accelerating tissue repair. This complex could exhibit a synergistic effect. PRP contains numerous proteins including PDGF (αα, ββ, and αβ isomers), TGF-b (β1 and β2 isomers), VEGF, angiopoietins, FGF, platelet-derived epidermal growth factor (PDEGF), epidermal growth factor (EGF), IGF, and fibronectin. When it is placed into bone surgical defect, it is quickly filled by a loose connective tissue and newly formed bone trabeculae in a linear arrangement. This connective tissue exhibits myxoid features and numerous newly formed blood vessels, showing intense angiogenic activity and stimulating the formation of bone marrow tissue. The myxoid tissue consisted of very delicate collagen fibers, numerous spindle cells, and a large amount of extracellular amorphous matrix (**Figure 15**). It is possible that the myxoid tissue and intense vascularization have occurred due to the preservation of bioactive angiogenic factors in the PRP, e.g., VEGF

in parenchymal cells or replace the injured tissue [32].

ADDM slices (Source: Gomes et al. [10]).

and angiopoietins, after its confection in laboratory [17].

#### *4.2.3. Platelet-rich plasma (PRP)*

Platelet-rich plasma (PRP) has been described as a source of growth factors and a potentially ideal scaffold for tissue engineering in regenerative medicine. PRP is an autologous product that contains highly concentrated number of platelet in a small volume of plasma from whole blood by gradient density centrifugation. It is a proven source of growth factors like platelet-derived growth factor (PGDF) and transforming growth factor β1 and β2 (TGF-β1 and β2), which is obtained by sequestering and concentrating platelets by gradient density centrifugation [25, 47]. This substrate is considered to ameliorate tissue regeneration due to presence of essential various cytokines and growth factors (GFs). Among them, platelet-derived growth factor (PDGF), transforming growth factor β1 (TGF-β1), insulin-like growth factors (IGF), platelet factor 4 (PF-4), fibroblast growth factor 2 (FGF-2), vascular endothelial growth factor (VEGF), and epidermal growth factor (EGF) are considered to be the most important. Subsequently, through stimulation of vascular ingrowth, macrophages arrive and start producing their own cytokines and GFs, some similar to those produced by platelets. This results in a new and continued local tissue repair and regrowth [40]. The cytokines and growth factors also

**Figure 14.** ADDM cut into slices with frozen microtomy (a) and placed into human dental socket covering its surface (b). Periapical radiographs after 15 (c) and 90 days (d) post-surgery. Arrow: radiopaque image suggests remaining of ADDM slices (Source: Gomes et al. [10]).

recruit resident stem cells to the site of injury, where they are stimulated to secrete additional growth factors and anti-inflammatory cytokines, causing more increase in collagen and matrix synthesis. Furthermore, the recruited stem cells will react with the environment to differentiate in parenchymal cells or replace the injured tissue [32].

*4.2.3. Platelet-rich plasma (PRP)*

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(yellow arrows) (Schmorl's stain, magnification ×200).

Platelet-rich plasma (PRP) has been described as a source of growth factors and a potentially ideal scaffold for tissue engineering in regenerative medicine. PRP is an autologous product that contains highly concentrated number of platelet in a small volume of plasma from whole blood by gradient density centrifugation. It is a proven source of growth factors like platelet-derived growth factor (PGDF) and transforming growth factor β1 and β2 (TGF-β1 and β2), which is obtained by sequestering and concentrating platelets by gradient density centrifugation [25, 47]. This substrate is considered to ameliorate tissue regeneration due to presence of essential various cytokines and growth factors (GFs). Among them, platelet-derived growth factor (PDGF), transforming growth factor β1 (TGF-β1), insulin-like growth factors (IGF), platelet factor 4 (PF-4), fibroblast growth factor 2 (FGF-2), vascular endothelial growth factor (VEGF), and epidermal growth factor (EGF) are considered to be the most important. Subsequently, through stimulation of vascular ingrowth, macrophages arrive and start producing their own cytokines and GFs, some similar to those produced by platelets. This results in a new and continued local tissue repair and regrowth [40]. The cytokines and growth factors also

**Figure 13.** Surgical bone defect region after 90 days in diabetic rabbit. Phases of the HDDM degradation during the bone remodeling process: HDDM in degradation (white arrows), exchange of HDDM by well-cellularized and immature bone tissue with remnants of dentin matrix (asterisk), and replacement of the preexisting HDDM by mature bone tissue

> Some investigators have reported that platelet-rich plasma (PRP) becomes more efficient in tissue repair when associated with bone grafts, bioactive material scaffold, or mesenchymal stem cells (MSCs) which are recruited at the wound site, accelerating tissue repair. This complex could exhibit a synergistic effect. PRP contains numerous proteins including PDGF (αα, ββ, and αβ isomers), TGF-b (β1 and β2 isomers), VEGF, angiopoietins, FGF, platelet-derived epidermal growth factor (PDEGF), epidermal growth factor (EGF), IGF, and fibronectin. When it is placed into bone surgical defect, it is quickly filled by a loose connective tissue and newly formed bone trabeculae in a linear arrangement. This connective tissue exhibits myxoid features and numerous newly formed blood vessels, showing intense angiogenic activity and stimulating the formation of bone marrow tissue. The myxoid tissue consisted of very delicate collagen fibers, numerous spindle cells, and a large amount of extracellular amorphous matrix (**Figure 15**). It is possible that the myxoid tissue and intense vascularization have occurred due to the preservation of bioactive angiogenic factors in the PRP, e.g., VEGF and angiopoietins, after its confection in laboratory [17].

**Figure 15.** PRP into surgical bone defect stimulating angiogenic activity and myxoit tissue with very delicate collagen fibers, numerous spindle cells, and a large amount of extracellular amorphous matrix, in a period of 30 days (Source: Gomes et al. [17]).

#### **4.3. Alloplastic materials**

The ideal implant should be composed of material that is biocompatible, affordable, and nondegradable. Although many materials have proven suitable, alloplastic materials are an attractive implant type, in particular, for the treatment of maxillofacial deformities and for reconstruction [7, 30]. To choose an ideal alloplastic material, first, we must identify the usage needs and their biological and physical properties as biocompatibility, easy of obtaining, noncarcinogenic and nonallergenic material, few production of inflammatory response, essential predeterminant physical-chemical properties to tissue reconstruction and regeneration, no susceptible to infections, sequelae or tissue sequestrum, easy anchorage, handling and sterilization without losing your characteristics and physical structure, high porosity to promote surrounding tissue ingrowth, similar consistency to replaced tissue, stability of contact surface, and good permanence in the implanted region to allow the suitable tissue reconstruction, especially in large defects. Among the several types of available alloplastic materials, porous polyethylene (porous PTFE), hydroxyapatite, and polyurethane resin derived from castor oil implants have been widely indicated in literature. It is worth mentioning the efficient effects of the porous PTFE implant when used in hard and soft tissues. Scientific investigations have shown its application in partial or total ossicular replacement, restorative and esthetic rhinoplasty, orbital implants, ear reconstructions, and augmentation or correction of craniofacial deformities such as temporal and mandibular regions, zygomatic arch, nose, and paranasal areas [7, 8]. Clinical and experimental studies have proven that the PTFE porosity aids surrounding tissue invasion, triggering good adhesion, and stability of the implant in the host site (**Figure 16**). However, this material can be considered poorly biocompatible by tissues, especially in diabetes, due to moderate-intense infiltration of mononuclear inflammation cells as lymphocytes, macrophages, and foreign body-type multinucleated giant cells, especially in uncontrolled diabetes.

**Figure 16.** Porous PTFE implanted into surgical bone defect in diabetic rats after 15 days. The defect was filled by fibrous connective tissue with discrete infiltration of mononuclear inflammatory cells and few newly formed bone trabeculae (arrow). The cavities show negative images (asterisk) of the polyethylene, and the soft tissue represents the pore areas

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(hematoxylin-eosin, magnification ×200).

**4.3. Alloplastic materials**

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Gomes et al. [17]).

uncontrolled diabetes.

The ideal implant should be composed of material that is biocompatible, affordable, and nondegradable. Although many materials have proven suitable, alloplastic materials are an attractive implant type, in particular, for the treatment of maxillofacial deformities and for reconstruction [7, 30]. To choose an ideal alloplastic material, first, we must identify the usage needs and their biological and physical properties as biocompatibility, easy of obtaining, noncarcinogenic and nonallergenic material, few production of inflammatory response, essential predeterminant physical-chemical properties to tissue reconstruction and regeneration, no susceptible to infections, sequelae or tissue sequestrum, easy anchorage, handling and sterilization without losing your characteristics and physical structure, high porosity to promote surrounding tissue ingrowth, similar consistency to replaced tissue, stability of contact surface, and good permanence in the implanted region to allow the suitable tissue reconstruction, especially in large defects. Among the several types of available alloplastic materials, porous polyethylene (porous PTFE), hydroxyapatite, and polyurethane resin derived from castor oil implants have been widely indicated in literature. It is worth mentioning the efficient effects of the porous PTFE implant when used in hard and soft tissues. Scientific investigations have shown its application in partial or total ossicular replacement, restorative and esthetic rhinoplasty, orbital implants, ear reconstructions, and augmentation or correction of craniofacial deformities such as temporal and mandibular regions, zygomatic arch, nose, and paranasal areas [7, 8]. Clinical and experimental studies have proven that the PTFE porosity aids surrounding tissue invasion, triggering good adhesion, and stability of the implant in the host site (**Figure 16**). However, this material can be considered poorly biocompatible by tissues, especially in diabetes, due to moderate-intense infiltration of mononuclear inflammation cells as lymphocytes, macrophages, and foreign body-type multinucleated giant cells, especially in

**Figure 15.** PRP into surgical bone defect stimulating angiogenic activity and myxoit tissue with very delicate collagen fibers, numerous spindle cells, and a large amount of extracellular amorphous matrix, in a period of 30 days (Source:

> **Figure 16.** Porous PTFE implanted into surgical bone defect in diabetic rats after 15 days. The defect was filled by fibrous connective tissue with discrete infiltration of mononuclear inflammatory cells and few newly formed bone trabeculae (arrow). The cavities show negative images (asterisk) of the polyethylene, and the soft tissue represents the pore areas (hematoxylin-eosin, magnification ×200).

As a compensatory factor, some therapeutic drugs could be used in association with alloplastic material, such as salmon calcitonin and verapamil. Systemic administration of salmon calcitonin has been widely used in patients with pathologies that affect the bone metabolism, including osteoporosis, osteogenesis imperfect, and Paget's disease, because of its recognized ability to inhibit bone resorption [8]. For support therapy, the salmon calcitonin can be indicated to treat aggressive cherubism, promoting the inhibition of multinucleated giant cell formation and osteoclastic activity, leading to a regressive process of the disease [16].

to the donor or to the individual that will receive the cells, and to accomplish the isolation of

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In this regard, we discussed some property and effects of biomaterials and alloplastic materials that could surely be indicated in cell therapy and tissue engineering, especially in diabetic individuals. As a compensative factor in diabetes, we suggested the clinical application of the AM and DDM in isolation or in association with other materials commonly found in the literature. The highly effective biological properties of these scaffolds favored the development of impaired tissues through the restoring, maintaining, or improving tissue function, corroborating with the basic principles of tissue engineering. Special attention has been paid to AM due to rich amount of growth factor and MSCs in its mesenchymal stroma, playing a key role in the creation of implantable tissue. Moreover, the AM is considered a promising source since it is usually discarded after childbirth (in humans and animals); therefore, it could be used without ethical or religious troubles when collected. The use of this membrane in some centers can be submitted the same rules for the use of organ and tissue transplants; however, it is worth noting the importance to conduct strict control procedures as serological

The funding source of the CEBAPE researches was supported by São Paulo Research Foundation, known as FAPESP (grant numbers: 1996/3203-2, 2002/03952-8, 2003/06017-0,

Mônica Fernandes Gomes\*, José Benedito Amorim, Lilian Chrystiane Giannasi and

Center of Biosciences Applied to Persons with Special Care Needs (CEBAPE), Institute of Science and Technology, São Paulo State University (UNESP), São José dos Campos,

[1] American Diabetes Association (ADA). Diagnosis and classification of diabetes mellitus.

[2] Amorim JBO, Teixeira SS, Negrato GV, Vilela-Goulart MG, Gomes MF. Histomorphometric analysis of verapamil effects on alveolar socket repair on diabetic alloxan-induced rats.

cells with high potential of expansion and proliferation.

and microbiological screening of this membrane.

2004/04447-0, 2004/08656-0, and 2007/54996-9).

\*Address all correspondence to: mfgomes@ict.unesp.br

Diabetes Care. 2014;**37**(Suppl 1):S81–S90

Ciência Odontológica Brasileira. 2009;**12**(1):37-48

**Acknowledgements**

**Author details**

São Paulo, Brazil

**References**

Miguel Angel Castillo Salgado

Regarding to the verapamil, it is used for treatment of high blood pressure, cardiac arrhythmias, and pectoris angina, as well as it is often indicated for diabetic individuals. Diabetic individuals are great users of this drug. The verapamil, a potent calcium channel blocker, promotes the inhibition of voltage-dependent calcium channel, L-type, reducing the Ca2+ entry into the cells and, consequently, altering their functions. Osteoclast lineages are affected, resulting in the inhibition of bone resorption and, thus, favoring the new bone tissue formation.

Some researches demonstrated that the verapamil has influenced the calcium balance in the mineralized tissues. It changes the calcium trophic hormone action, altering the endocrine regulator effect in the bone turnover process, since it acts directly in the inhibition of the osteoclasts activity. On the other hand, the increase of intracellular calcium levels of osteoblast stimulates the IGF-I, IGF-II, and TGF-β production, favoring the bone tissue formation [2, 8].

In addition, implant neovascularization is essential to recruit inflammatory and progenitor cells, initiating the wound healing and/or leading to an exponential formation of new tissues [8, 17]. Thus, researches have suggested an additional bioactive material to the porous polyethylene implant, as mesenchymal stem cells (MSCs) or angiogenic and mitogenic factors, so that they can present synergistic effects and improve the environment and, consequently, the development of desirable tissues in a variety of reconstructive and restorative surgeries.

## **5. Considerations**

Considering the exponential growth trend of chronic diseases, in particular, diabetes mellitus, ideal bioactive biomaterials for hard and soft tissue reconstruction and regeneration have been a great challenge for many researchers in worldwide. To choose an ideal biomaterial to be implanted in diabetes, we must first know the cellular and molecular pathophysiological mechanisms following immeasurable damages in the tissue response against any extrinsic factors, including impacts or grafts. In regard to the biodynamic constraints of the repair process, the recruitment of cells, as inflammatory and progenitors cells, and growth factors and cytokine production or action are largely impaired in diabetics individuals, in particular, with uncontrolled hyperglycemia. As undesirable results, these materials could increase the risk for and severity of the diabetic complications instead of favoring the tissue regeneration process.

Based on basic concepts of biomaterial, it is important to highlight that the biocompatibility, material storage without loss of viability, ease of obtaining the material, favorable cost and benefit relationship, and mainly preservation of its inductive properties must be considered as pivotal parameters for tissue-engineered implants which could be used in diabetes. Other concerns of researchers are to identify the reliable material sources that do not pose any risk to the donor or to the individual that will receive the cells, and to accomplish the isolation of cells with high potential of expansion and proliferation.

In this regard, we discussed some property and effects of biomaterials and alloplastic materials that could surely be indicated in cell therapy and tissue engineering, especially in diabetic individuals. As a compensative factor in diabetes, we suggested the clinical application of the AM and DDM in isolation or in association with other materials commonly found in the literature. The highly effective biological properties of these scaffolds favored the development of impaired tissues through the restoring, maintaining, or improving tissue function, corroborating with the basic principles of tissue engineering. Special attention has been paid to AM due to rich amount of growth factor and MSCs in its mesenchymal stroma, playing a key role in the creation of implantable tissue. Moreover, the AM is considered a promising source since it is usually discarded after childbirth (in humans and animals); therefore, it could be used without ethical or religious troubles when collected. The use of this membrane in some centers can be submitted the same rules for the use of organ and tissue transplants; however, it is worth noting the importance to conduct strict control procedures as serological and microbiological screening of this membrane.

## **Acknowledgements**

As a compensatory factor, some therapeutic drugs could be used in association with alloplastic material, such as salmon calcitonin and verapamil. Systemic administration of salmon calcitonin has been widely used in patients with pathologies that affect the bone metabolism, including osteoporosis, osteogenesis imperfect, and Paget's disease, because of its recognized ability to inhibit bone resorption [8]. For support therapy, the salmon calcitonin can be indicated to treat aggressive cherubism, promoting the inhibition of multinucleated giant cell

Regarding to the verapamil, it is used for treatment of high blood pressure, cardiac arrhythmias, and pectoris angina, as well as it is often indicated for diabetic individuals. Diabetic individuals are great users of this drug. The verapamil, a potent calcium channel blocker, promotes the inhibition of voltage-dependent calcium channel, L-type, reducing the Ca2+ entry into the cells and, consequently, altering their functions. Osteoclast lineages are affected, resulting in

Some researches demonstrated that the verapamil has influenced the calcium balance in the mineralized tissues. It changes the calcium trophic hormone action, altering the endocrine regulator effect in the bone turnover process, since it acts directly in the inhibition of the osteoclasts activity. On the other hand, the increase of intracellular calcium levels of osteoblast stimulates the IGF-I, IGF-II, and TGF-β production, favoring the bone tissue formation [2, 8]. In addition, implant neovascularization is essential to recruit inflammatory and progenitor cells, initiating the wound healing and/or leading to an exponential formation of new tissues [8, 17]. Thus, researches have suggested an additional bioactive material to the porous polyethylene implant, as mesenchymal stem cells (MSCs) or angiogenic and mitogenic factors, so that they can present synergistic effects and improve the environment and, consequently, the development of desirable tissues in a variety of reconstructive and restorative surgeries.

Considering the exponential growth trend of chronic diseases, in particular, diabetes mellitus, ideal bioactive biomaterials for hard and soft tissue reconstruction and regeneration have been a great challenge for many researchers in worldwide. To choose an ideal biomaterial to be implanted in diabetes, we must first know the cellular and molecular pathophysiological mechanisms following immeasurable damages in the tissue response against any extrinsic factors, including impacts or grafts. In regard to the biodynamic constraints of the repair process, the recruitment of cells, as inflammatory and progenitors cells, and growth factors and cytokine production or action are largely impaired in diabetics individuals, in particular, with uncontrolled hyperglycemia. As undesirable results, these materials could increase the risk for and severity of the diabetic complications instead of favoring the tissue regeneration process. Based on basic concepts of biomaterial, it is important to highlight that the biocompatibility, material storage without loss of viability, ease of obtaining the material, favorable cost and benefit relationship, and mainly preservation of its inductive properties must be considered as pivotal parameters for tissue-engineered implants which could be used in diabetes. Other concerns of researchers are to identify the reliable material sources that do not pose any risk

formation and osteoclastic activity, leading to a regressive process of the disease [16].

the inhibition of bone resorption and, thus, favoring the new bone tissue formation.

**5. Considerations**

430 Biomaterials in Regenerative Medicine

The funding source of the CEBAPE researches was supported by São Paulo Research Foundation, known as FAPESP (grant numbers: 1996/3203-2, 2002/03952-8, 2003/06017-0, 2004/04447-0, 2004/08656-0, and 2007/54996-9).

## **Author details**

Mônica Fernandes Gomes\*, José Benedito Amorim, Lilian Chrystiane Giannasi and Miguel Angel Castillo Salgado

\*Address all correspondence to: mfgomes@ict.unesp.br

Center of Biosciences Applied to Persons with Special Care Needs (CEBAPE), Institute of Science and Technology, São Paulo State University (UNESP), São José dos Campos, São Paulo, Brazil

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Biomaterials for Tissue Engineering Applications in Diabetes Mellitus

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## *Edited by Leszek A. Dobrzański*

The book *Biomaterials in Regenerative Medicine* is addressed to the engineers and mainly medical practitioners as well as scientists and PhD degree students. The book indicates the progress in research and in the implementation of the ever-new biomaterials for the application of the advanced types of prosthesis, implants, scaffolds and implantscaffolds including personalised ones. The book presents a theoretical approach to the synergy of technical, biological and medical sciences concerning materials and technologies used for medical and dental implantable devices and on metallic biomaterials. The essential contents of the book are 16 case studies provided in each of the chapters, comprehensively describing the authors' accomplishments of numerous teams from different countries across the world in advanced research areas relating to the biomaterials applied in regenerative medicine and dentistry. The detailed information collected in the book, mainly deriving from own and original research and R&D works pursued by the authors, will be beneficial for the readers to develop their knowledge and harmonise specific information concerning these topics.

Biomaterials in Regenerative Medicine

Biomaterials

in Regenerative Medicine

*Edited by Leszek A. Dobrzański*

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