Biomimetic Tissue Engineering

### **Chapter 3**

Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric Materials with Biomimetic 3D Microarchitecture for Tissue Engineering and Medical Applications

*Ching-Cheng Huang and Masashi Shiotsuki*

#### **Abstract**

Continuous work and developments in biomedical materials used in three-dimensional (3D) bioprinting have contributed to significant growth of 3D bioprinting applications in the production of personalized tissue-repairing membrane, skin graft, prostheses, medication delivery system, and 3D tissue engineering and regenerative medicine scaffolds. The design of clinic products and devices focus on new natural and synthetic biomedical materials employed for therapeutic applications in different 3D bioprinting technologies. Design and characterization of natural and synthetic soft polymeric materials with biomimetic 3D microarchitecture were considered. The natural soft polymeric materials would focus on new design bioinspired membranes containing supercritical fluids-decellularized dermal scaffolds for 3D bioprinting potential applications. Synthetic soft polymeric materials would focus on bioinspired polyvinyl alcohol (b-PVA) matrix with structural foam-wall microarchitectures. Characterization, thermal stability, and cell morphology of the b-PVA and the corresponding collagen-modified b-PVA were employed to evaluate their potential tissue engineering applications. Also, the b-PVA materials were conductive to HepG2 cells proliferation, migration, and expression, which might serve as a promising liver cell culture carrier to be used in the biological artificial liver reactor. TGA, DTG, DSC, SEM, and FTIR were employed to build up the effective system identification approach for biomimetic structure, stability, purity, and safety of target soft matrix.

**Keywords:** biomimetic, three-dimensional, microarchitecture, design-thinking, supercritical fluids

#### **1. Introduction**

Bioprinting, a type of three-dimensional (3D) bioprinting, uses cells and other biological materials as "inks" to fabricate 3D biological structures. Bioprinted materials have the potential to repair damaged organs, cells, and tissues in the human body. Skin, bone, and blood vessels may be bioprinted. Although a variety of tissue engineering strategies combining cells and biomaterial scaffolds have been investigated for human tissues regeneration (e.g., using natural biomaterials and/or synthetic polymers [1]), the success of this approach still strongly depends on the development of more suitable scaffolds. In this sense, one of the main challenges is to accurately reproduce the complex 3D anatomy with personalized shape and size [2–4]. Efforts to address these issues have led to an increased interest in the application of threedimensional (3D) bioprinting in the tissue engineering field. 3D bioprinting is a biofabrication method that uses computer-aided design (CAD) and additive layer manufacturing technique to precisely deposit bioinks (basically comprising a mixture of biomaterials with cells, with or without bioactive molecules) in a predesigned manner to create 3D bioengineered living structures and to generate artificial tissue and organs [5, 6]. Using this novel printing technology, biomaterials were possible to fabricate a patient-specific scaffold reproducing individual shape, size, and macrostructure of the native tissue [7, 8]. An implant's effectiveness depends upon the form of biomaterial used in its manufacture. A suitable material for implants should be biocompatible, sterile, mechanically stable, and simple to shape [9]. 3D printing technologies have been breaking new ground in the medical industries in order to build patient-specific devices embedded in bioactive drugs, cells, and proteins. 3D printing was a non-exclusive concept that defines various techniques for layer-bylayer construction. 3D bioprinting has progressed the method of printing standard biocompatible materials and even actual cells into difficult 3D dimensional tissue buildings [10], with the ability to create ideal tissues and organs appropriate for different biomedical uses, such as organ transplantation [11]. In recent years, many bioprinting techniques have been developed to store cells and hydrogels together, altered and used as cell printers to print polymers [12]. In such printers, cell suspensions or cell totals are mounted on printer extruder, and the printing mechanism is controlled. Continuous work and developments in biomaterials used in 3D printing have contributed to significant growth of 3D printing applications in the production of personalized tissue-repairing membrane, skin graft, prostheses, medication delivery system, and 3D tissue engineering and regenerative medicine scaffolds. The design of clinic products and devices focus on new natural and synthetic biomedical materials used for therapeutic applications in different 3D bioprinting technologies. Many specific forms of medical 3D bioprinting technology are explored in depth, including extrusion-based bioprinting and inkjet printing processes, the specific therapeutic uses, various types of biomaterial used today, and the major shortcomings are being studied in depth.

Bioinspired structure-guided tissue/cell in vivo growth or vitro culturing tissue/ cell in vivo growth and in vitro culturing are remarkably important for implantable medical devices and drug evaluation [13]. To achieve desired growth and reproduction, it is necessary to construct the culturing conditions and the growing structure as close to living being as possible. Bioinspired surface was applied to the implants to improve the implanting quality. Bioinspired porous structure was built on the scaffolds to guide and improve the tissue growth.

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

#### **2. Design and preparation of natural soft polymeric materials with biomimetic 3D microarchitecture containing decellularized extracellular matrix**

Tree-dimensional (3D) bioprinting showed potential in tissue engineering and regenerative applications due to its overwhelming advantages over other approaches. In order to promote the functions of bioprinted tissues, the development of novel and versatile bioinks will have crucial implications [14]. Natural materials were famous for the excellent biocompatibility and abundance, among which sodium alginate mixed with gelatin has been widely used as bioink for extrusion-based 3D bioprinting [15]. Despite advance in bioprinting and bio-fabrication during the past decade, fabricating complex and functional tissue constructs that mimic their natural counterparts still remains a challenge. Bioink optimization is considered as one of main challenges in cell-laden 3D bioprinting [16].

Numerous materials have been proposed, modified, and employed for medical bioprinting applications such as scaffolds for skin and bone tissue reconstruction such as synthetic materials and natural materials [17–29]. In bioprinting applications, biomedical materials could be extruded through a print head either by pneumatic pressure or mechanical force. Since the process did not involve any heating procedures, it was most commonly used for preparing tissue engineering constructs with cells and growth hormones laden. Bioinks were the biomedical materials laden with cells and other biological materials and used for 3D bioprinting. The 3D bioprinting process allowed for the deposition of small units of cells accurately, with minimal process-induced cell damage. Advantages such as precise deposition of cells control over the rate of cell distribution and process speed had greatly increased the applications of this technology in fabricating living scaffolds. A wide range of materials with varied viscosities and high cell density aggregates could be 3D printed using this technique. A large variety of polymers were under research for the use in bioprinting technology [18]. Natural polymers, including collagen [19], gelatin [20], decellularized extracellular matrix(dECM) [21], alginate [22], chitosan [23], cellulose [24], agarose [25], silk protein [26], and hyaluronic acid (HA) [27, 28], and photopolymerizable gelatin and hyaluronic acid [29] were commonly used in bioinks for 3D bioprinting (**Figure 1**). Often these bioinks are post-processed either by chemical or UV cross-linking to enhance the constructs' mechanical properties. Depending on the type of polymer used in the bioink, biological tissues and scaffolds of varied complexity can be fabricated.

Decellularized extracellular matrix (dECM) scaffolds had a lot of collagens, which constitute the main structural element of the dECM, provide tensile strength, regulate cell adhesion, support migration, and direct tissue development. Dense connective tissue is an abundant source of dECM scaffolds, which can be prepared and purified by a defatting and decellularizing procedure [21, 30]. The treatments combined with supercritical carbon dioxide and specific enzymes could be employed to prepare dECM scaffolds. Furthermore, a series of new biomedical composite materials containing dECM scaffolds and alginate could be designed and prepared for tissue engineering and bioprinting applications. The composite materials containing dECM scaffolds for biomimetic bioinks could be characterized by Fourier transform infrared spectroscopy (FTIR), thermo-gravimetric analysis (TGA), and scanning electron microscope (SEM) to get the results of identifications, thermal stabilities, and microstructures.

#### **Figure 1.**

*Chemical structures of natural and synthetic soft polymeric materials with biomimetic 3D microarchitecture for tissue nanogineering, bioprinting, and medical applications. A) Natural soft polymeric material. B) Synthetic soft polymeric material.*

#### **2.1 Natural soft polymeric materials with biomimetic 3D microarchitecture containing decellularized extracellular matrix via supercritical fluid treatments**

Decellularized extracellular matrix could be an important natural soft polymeric material with biomimetic 3D microarchitecture for bioprinting applications. In previous study, supercritical fluid of carbon dioxide (ScCO2) was employed for preparation of designed decellularized extracellular matrix scaffolds for bioprinting applications. The ScCO2 could be employed before or after decellularization treatments for effectively removing most fatty acids and tissues [21]. The ScCO2 extraction could be performed for complete decellularization. Also, the steady thickness of about 0.5 mm of thinly sliced tissue sample, which could be obtained by using a designed tissuecutting machine (Taiwan PARSD Pharm. Tech. Consulting Ltd Co.), was employed for complete decellularization.

Fourier transform infrared spectroscopy analysis was important for characterization of the natural soft polymeric material with biomimetic 3D microarchitecture containing decellularized extracellular matrix because of collagen segments, which could exhibit remarkable typical absorption bands of specific functional groups such as Amide I, Amide II, Amide III, Amide A, and Amide B could be an important for bioprinting applications. From the FTIR analysis of the original tissue such as porcine

#### *Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

skin, absorption bands at 1452, 1400, 1337, 1240, 1203, and 1080 cm−1 were attributed to the amides III containing δ(CH2), δ(CH3), ν(C–N), and δ(N–H) absorptions of collagens in the original porcine skin. Amides I and amides II absorptions were found at 1632 and 1551 cm−1, respectively. The absorption band at 3301 cm−1 δ(C–H) was attributed to the fatty acid of the original porcine skin. The absorption band at 1744 cm−1 δ(C=O) was attributed to the fatty acid. After supercritical carbon dioxide treatment and decellularization, the decellularized extracellular matrix could be formed and the absorption bands of fatty acids could not be observed in FTIR spectrum, demonstrating the effectiveness of the supercritical carbon dioxide treatment and the formation of natural soft polymeric materials. The resulting natural soft polymeric materials could be considered as a nano-bioscaffold. The microstructures of resulting membranes with dNBS-S were characterized by scanning electron microscope (SEM). Scanning electron micrographs of the decellularized samples such as decellularized porcine skin (dNBS-S), decellularized porcine liver(dNBS-L), decellularized porcine costal cartilages (dNBS-CC), and decellularized porcine elastic cartilage (dNBS-EC) after treatment with supercritical carbon dioxide are shown in **Figure 2(A)–(D)**, respectively. The pore space of the resulting natural soft polymeric materials from different tissues such as the skin, liver, costal cartilage, and elastic cartilage with diameters in a range of 10–250 μm was observed, which could be good nano-bioscaffolds for cell migration [21].

Further, L929 cells were cultured on the resulting natural soft polymeric material for a potential evaluation of good nano-bioscaffolds for tissue engineering. The morphology of L929 cells cultured on the resulting natural soft polymeric materials was also investigated by SEM (**Figure 3**). Significantly, most area was covered by L929

#### **Figure 2.**

*Scanning electron micrographs of the samples: (A) decellularized porcine skin (dNBS-S), (B) decellularized porcine liver(dNBS-L), (C) decellularized porcine Costal cartilages(dNBS-CC), and (D) decellularized porcine elastic cartilage (dNBS-EC).*

**Figure 3.** *SEM photograph of L929 cells grew on dNBS-S.*

cells on the resulting natural soft polymeric materials such as dNBS-S after 3 days of culture was observed [21]. The cells that grew upon the dNBS-S were significantly observed. Perez-Puyana et al. also reported that the pore space of nano-bioscaffold was suggested to be small enough to establish a high specific surface area and large enough to allow cells to migrate into the microstructure (20–120 μm) [31].

#### **2.2 Natural soft polymeric composite materials with biomimetic 3D microarchitecture containing alginate and decellularized extracellular matrix via supercritical fluid treatments**

Decellularized extracellular matrix(dECM)/alginate composite materials with biomimetic 3D microarchitecture and cross-linked decellularized extracellular matrix/alginate composite materials with biomimetic 3D Microarchitecture could also be designed for tissue engineering and medical applications. Briefly, the desired amount of dECM powder was first dispersed completely in water. Then, alginate aqueous solution was homogenized thoroughly with the dispersed dECM solution. The alginate/dECM solutions were employed to form 3D printed membranes by using extrusion-based bioprinting, molding, and freeze-drying procedures for evaluation of new 3D bioprinting materials. Further, cross-linked decellularized extracellular matrix/alginate composite materials with biomimetic 3D Microarchitecture were designed and prepared by using ionic crosslinking procedure with CaCl2 aqueous solution.

The incorporation of nano-bioscaffold(dNBS) in the alginate/decellularized extracellular matrix composite materials(AdNBS) with biomimetic 3D microarchitecture was done. The FTIR spectroscopy analysis was carried out to confirm the structures. In the spectrum of the AdNBS, the main absorption bands at around 1632 cm−1 (amide I, C-O, and C-N stretching), 1537 cm−1 (amide II) and 1242 cm−1 (amide III) were also observed. Because of the introduction of alginate segments, retaining the characteristic bands of pure sodium alginate was observed. The characteristic bands of AdNBS at around 1595 cm−1 (the carbonyl (C=O) bond) and 1408 cm−1 (asymmetric and symmetric stretching peaks of carboxylate salt groups) were visible [21].

#### *Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

It showed a stronger absorption band at 1595 cm−1 and two remarkably shoulders at 1632 and 1537 cm−1, which were characteristic absorption of carbonyl groups of amide (amide I and amide II) of dNBS segments, which would confirm the formation of AdNBS composite materials remarkably. A higher absorption from 3600 to 3200 cm−1 would be observed in the spectrum, which suggested an increase of hydrogen bonds resulting from the interaction between dNBS molecule and alginate(ALG) molecule. The results of FTIR indicated the presence of dNBS in the composite materials as wells as the interaction between them. After cross-linking reaction, the cross-linked decellularized extracellular matrix/alginate composite materials with biomimetic 3D microarchitecture were obtained and the calcium alginate segments were formed. The -C-O-O-Ca-O-CO- structure made the C-O stretching vibration absorption increase and had an obvious absorption band at 1024 cm−1, which indicated the formation of -CO-O-Ca-O-CO- structure in AdNBS materials [21].

The microstructure of resulting AdNBS composite materials was characterized by scanning electron microscopy (SEM). The micro-scaffold structure could be observed in the dNBS-S derived from porcine skin after treatment with supercritical carbon dioxide (**Figure 2(B**)) showed quite different morphology form ALG, which showed a micro-scaffold shape with narrow boundaries [21]. The averaged diameter of narrow boundaries was found in a range of 1–3 μm. The remarkable micro-scaffold structures were observed in AdNBS composite materials with various introduction ratios of dNBS-S and ALG. With the increasing introduction ratio of dNBS-S and ALG, the new combined micro-scaffold shapes of composite materials were observed with the spaces and smooth mixed boundaries. The averaged diameter of smooth mixed boundary was found in a range of 1–35 μm [21], which might provide relatively high structural and thermal stabilities.

Thermal stability of resulting natural soft polymeric materials with biomimetic 3D microarchitecture such as AdNBS-S could be evaluated by TGA (**Figure 4**). Thermogravimetric analysis of the ALG showed a maximum pyrolysis temperature (Tdmax) lower than 300 degrees. To enhance the thermal stability of ALG-based materials, the decellularized nano-bioscaffolds, dNBS-S, were introduced. The maximum pyrolysis temperature (Tdmax) of dNBS-S was higher than 300 degrees. The resulting AdNBS-S would be a new heat-resistant biomaterial for bioprinting applications. From TGA analysis of AdNBS-S, the main loss is presented in two different temperature ranges given by stage I(<150°C), stage II(150–250°C), and stage III (250–500°C). From TGA results of ALG, initial weight loss up to 150°C is found to be 15, 20, and 20 for non-cross-linked alginate(NCA), high cross-linked alginate(HCA) (5wt% CaCl2(aq)), and relative low cross-linked alginate(LCA)(1wt% CaCl2(aq)), respectively, due to the elimination of absorbed and bounded water molecules in the membrane. In case of cross-linked AdNBS-S (dNBS-S/ALG= 20/80), composite material with CaCl2(aq) such as relatively low-cross-linked AdNBS-S(dNBS-S/ALG=20/80) composite material(LCLAdNBS-S) (1wt% CaCl2(aq)) and high-cross-linked AdNBS-S(dNBS-S/ALG=20/80) composite material(HCLAdNBS-S)(5wt% CaCl2(aq)), weight loss would be increased comparing with the non-cross-linked AdNBS-S(dNBS-S/ ALG=20/80) composite material sample (**Figures 4(C)**, **(F)**, and **(I)**). This increase may be due to more adsorption of water molecules present along with Ca2+ molecules while cross-linking with CaCl2 aqueous solution. From TGA results, relatively high Tdmax(i) values of AdNBS-S composite materials were observed at 270–300°C in stage II comparing with the 250°C of non-cross-linked alginate(NCA) (**Figure 4(A)**). When the high concentration of CaCl2 (5wt%) added, the relative high Tdmax(i) value of high-cross-linked alginate(HCA) was observed at ca. 270°C comparing with

#### **Figure 4.**

*Thermogravimetric analysis of the composite membranes with/without cross-linking reaction, (A) non-crosslinked alginate material (NCA); (B) non-cross-linked AdNBS-S (5/95) (NCHAdNBS-S); (C) non-cross-linked AdNBS-S (20/80) (NCLAdNBS-S); (D) relative low-cross-linked alginate(LCA) (1wt% CaCl2(aq); (E) relative low-cross-linked AdNBS-S (5/95) (LCHAdNBS-S) (1wt% CaCl2(aq) ); (F) relative low-cross-linked AdNBS-S (20/80) (LCLAdNBS-S) (1wt% CaCl2(aq)); (G) high-cross-linked alginate(HCA) (5wt% CaCl2(aq) ); (H) high-cross-linked AdNBS-S (5/95)(HCHAdNBS-S) (5wt% CaCl2(aq)); (I) high cross-linked AdNBS-S (20/80) (HCLAdNBS-S) (5wt% CaCl2(aq)).*

the 250°C of non-cross-linked alginate(NCA) and 252°C of relative low-cross-linked alginate molecule(LCA)(1wt%). Furthermore, when the high concentration of CaCl2 (5wt%) was added, the relatively high Tdmax(i) values of high-cross-linked AdNBS-S such as HCHAdNBS-S and HCLAdNBS-S were observed at ca. 300°C comparing with the 270°C of high-cross-linked alginate (5wt%)(HCA) (**Figure 4(G)**). It would be due to the association between the ALG molecule and dNBS-S. Similarly, the relatively high Tdmax(ii) value of cross-linked AdNBS-S composite materials would be observed at ca.370°C. Particularly, much higher Tdmax(iii) values than Tdmax(ii) values of cross-linked AdNBS-S composite materials were observed at 400°C, which might be contributed to the formation of new mixed cross-linked network microstructures of ALG and dNBS-S molecules.

When a little amount of dNBS-S was introduced into the AdNBS-S composite materials without CaCl2, weak ionic association between –COOH group of alginate molecule and -NH2 group of dNBS-S molecule was formed and which is difficult to

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

build up the cross-linking structure. With an increasing introduction of dNBS-S to AdNBS-S composite materials, ordinary ionic association between –COOH group of alginate molecule and –NH2 group of dNBS-S was employed to build weak ionic crosslinking microstructure. When a large amount of dNBS-S was introduced into the AdNBS-S composite materials without CaCl2(NCLAdNBS-S), strong ionic association between -COOH group of ALG molecule and –NH2 group of dNBS-S was employed to build up strong ionic cross-linking microstructure.

When a little amount of dNBS-S was introduced into the dNBS-S/alginate composite membrane with 1wt% CaCl2(LCHAdNBS-S), weak ionic association between –COOH group of alginate molecule and -NH2 group of dNBS-S and weak ionic associations among –COOH group of alginate molecule, Ca2+, and –COOH group of alginate molecule could be found. The remarkable cross-linked microstructure was difficult to be observed. Tdmax(iii) value could not be found in DTG of LCHAdNBS-S results (Figure (E)). Weak ionic association between –COOH group of alginate molecule and –NH2 group of dNBS-S and weak ionic associations among –COOH group of alginate molecule, Ca2+, and –COOH group of alginate molecule were employed to build weak ionic cross-linking microstructure. The overlapped Tdmax(ii,iii) values of AdNBS-S (LCHAdNBS-S) could be observed at 390°C. With an increasing introduction of dNBS-S to AdNBS-S composite materials with 5wt% CaCl2, some strong ionic associations were employed to build strong mixed ionic cross-linked microstructure. The remarkable high Tdmax(iii) values of AdNBS-S could be observed at 400°C (HCLAdNBS-S). That is, when the high concentrations of CaCl2(5wt%) and dNBS-S were added, some different associations would be enhanced, such as association between –COOH group of alginate molecule and –NH2 group of dNBS-S, associations among –COOH group of alginate molecule, –COOH group of dNBS-S, and Ca2+ ion, associations among –COOH group of alginate molecule, –COOH group of alginate molecule, and Ca2+ ion, and associations among –COOH group of dNBS-S, –COOH group of dNBS-S, and Ca2+ ion.

With an increasing additions of dNBS-S to composite membranes with 1wt% CaCl2, ordinary ionic association between –COOH group of alginate and –NH2 group of dNBS-S was employed to build weak ionic cross-linked microstructure (**Figure 5**). When a large amount of dNBS-S was introduced into the dNBS-S/alginate composite membrane with 1wt% CaCl2, some ordinary ionic associations such as ionic association between –COOH group of alginate molecule and –NH2 group of dNBS-S, ionic between among –COOH group of alginate molecule, Ca2+, and –COOH group of dNBS-S, and ionic between among –COOH group of dNBS-S, Ca2+, and –COOH group of dNBS-S. The remarkable high Tdmax(iii) values of AdNBS-S could be observed at 400°C.

#### **2.3 Natural soft polymeric composite materials with biomimetic 3D microarchitecture containing gelatin, alginate, and decellularized extracellular matrix via supercritical fluid treatments**

For specific clinic applications, the gelatin could also be employed to prepare crosslinked porous natural soft polymeric composite materials with biomimetic 3D microarchitecture containing gelatin, alginate, and decellularized nano-bioscaffolds. A series of composite materials with a biomimetic 3D microarchitecture were prepared based on a fixed weight ratio of gelatin, alginate, and decellularized nano-bioscaffolds. An aqueous gelatin(G) solution and an aqueous alginate(A) solution (weight ratio is 2:1) were homogenized thoroughly. A mixed aqueous solution of gelatin(G) and alginate(A) was obtained as an aqueous GA solution. Briefly, the desired amount of decellularized

#### **Figure 5.**

*Proposed model of ionic cross-linked structure from Ca2+ ions, alginate molecules, and dNBS-S. (A) non-crosslinked structure, (B) weak ionic cross-linked structure, (C) weak strong ionic cross-linked structure, (D) ionic cross-linked interpenetrating structure.*

nano-bioscaffolds powder was first dispersed completely in water. Then, the aqueous GA solution was homogenized thoroughly with the dispersed decellularized nanobioscaffolds solution. A mixed aqueous solution of gelatin/alginate(GA) and decellularized nano-bioscaffolds was then molded and frozen and then lyophilized. Polymeric porous natural soft polymeric composite materials(GA/dNBS) with biomimetic 3D microarchitecture containing gelatin, alginate, and decellularized nano-bioscaffolds were obtained. The resulting GA/dNBS was further soaked in CaCl2 or GTA aqueous solution with various soaking times for cross-linking reaction with magnet mixer. The cross-linked GA/dNBS solutions were then frozen and dried. A series of designed crosslinked porous composite materials with biomimetic 3D microarchitecture containing gelatin, alginate, and decellularized nano-bioscaffolds could be obtained [32].

#### **2.4 Natural soft polymeric composite materials with biomimetic 3D microarchitecture containing bovine Achilles tendon type I collagen and decellularized extracellular matrix via supercritical fluid treatments**

The bovine Achilles tendon type I collagen was dispersed in an acetic acid solution and stirred to get a full dispersion collagen gel. The decellularized extracellular matrix immersed in acetic acid solution at the same concentration with collagen gel resolution was respectively placed on the table concentrator of 25 and 37°C for 24 h, allowing decellularized extracellular matrix to fully swell in acetic acid solution to obtain a decellularized extracellular matrix gel. Afterward, the decellularized extracellular matrix gels were added into the collagen gel following continuous stirring to obtain the composite hydrogel with a final ratio of 9:1 (wt/wt) of collagen and decellularized extracellular matrix. The composite gel was lyophilized. The composite nanobioscaffolds were obtained. After, those scaffolds were fully cross-linked by immersion in a 0.1% (v/v) glutaraldehyde solution (75% ethanol aqueous solution) for 1 h, freeze-dried, and

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

sterilized at a dose of 25 kGy. Finally, the cross-linked collagen/decellularized fibrous extracellular matrix composite nano-bioscaffolds were obtained.

#### **3. The 3D bioprinting applications of natural soft polymeric materials with biomimetic 3D microarchitecture**

#### **3.1 Preparation of natural soft polymeric scaffolds with biomimetic 3D microarchitecture by extrusion-based bioprinting procedure with/without cells**

Preparation of natural soft polymeric scaffolds with biomimetic 3D microarchitecture could be carried out from dNBS-S and ALG by using an extrusion-based bioprinting procedure with/without cells (**Figure 6**). The uncross-linked natural soft polymeric scaffolds having biomimetic 3D microarchitectures (**Figure 6(A)**) were obtained. Subsequently, the cross-linked interpenetrating composite membranes with/without cells could be obtained after ionic cross-linking procedures (**Figure 6(B)**). On the other hand, the cross-linked interpenetrating composite membranes with/without cells (**Figure 6(D)**) could be directly obtained by using a dual extrusion-based bioprinting procedure (**Figure 6(C)**). Further, introduction of gelatin in designed natural soft polymeric materials with biomimetic 3D microarchitecture could enhance structural and thermal stabilities for tissue engineering and 3D bioprinting applications as a scaffold (**Figure 7**) [32].

#### **3.2 Design of natural soft polymeric materials with biomimetic 3D microarchitecture for dNBS-containing skin equivalents by using a layer-by-layer-dispensing method as active cell-containing wound dressings**

Recently, skin grafts could be fabricated using a commercially available bioprinter(CELLINK) [33]. PLG mech was used as a mech to build up a skin construct. For medical application of skin grafts, the designed alginate-based

#### **Figure 6.**

*Proposed model of ionic cross-linked structure from Ca2+ ions, alginate molecules, and dNBS-S/alginate. (A) Non-cross-linked structure, (B) weak ionic cross-linked structure, (C) weak strong ionic cross-linked structure, (D) ionic cross-linked interpenetrating structure.*

#### **Figure 7.**

*Proposed model of ionic cross-linked structure from Ca2+ ions, alginate molecules, and dNBS-S/alginate. (A) Non-cross-linked structure, (B) weak ionic cross-linked structure, (C) weak strong ionic cross-linked structure, (D) ionic cross-linked interpenetrating structure.*

biopolymeric mech containing dECM could be employed to displace PLG mech. The dECM-containing skin equivalents with a dermal layer could be designed using a layer-by-layer-dispensing method. In the concept, a designed dermal construct could be generated by dispensing dermal bioink containing human endothelial cells (ECs), fibroblasts (FBs), and pericytes (PCs) and transfer these cells such as FBs, PCs, and ECs dispensed from a syringe into a mold at an extrusion pressure. A sterile alginate-based biopolymeric mesh was added on top of the first printed dermal layer and incubated at 37°C. The alginate-based biopolymeric mesh was sterilized prior to printing by immersion in a 70% ethanol solution for 30min and allowed to dry for 1 h at room temperature. Upon gelation, additional dermal bioink containing FBs, PCs, and ECs was printed on top of the mesh, incubated at 37°C until complete gelation, and submerged in media. The resulting construct was further submerged in a larger volume of media to prevent rapid exhaustion of nutrients. Medium was changed daily. After 4days under medium submersion to allow self-assembly of endothelial networks, the dermal construct was carefully transferred for clinic applications such as an active cell-containing wound dressing (**Figure 8(A)** and **(B)**) or further was bioprinted an additional keratinocytes (KCs) layer to obtain an active KCs-containing wound dressings for wound management (**Figure 8(C)** and **(D)**).

#### **3.3 Design of natural soft polymeric materials with biomimetic 3D microarchitecture for dNBS-containing skin equivalents by using a wholedispensing method**

For specific medical application of skin grafts, the dECS-containing skin equivalents with a dermal layer could be designed using a *whole-*dispensing method. In the concept, *Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

#### **Figure 8.**

*Proposed model of active cell-containing and active keratinocytes-containing wound dressings for wound management. (A) preparation of an active cell-containing wound dressing, (B) clinical application of an active cell-containing wound dressing, (C) preparation of an active keratinocytes-containing wound dressing, and (D) clinical application of an active keratinocytes-containing wound dressing.*

a designed dermal construct could be generated by dispensing dermal bioink containing human endothelial cells (EC), fibroblasts (FBs), pericytes (PCs) , and dECS. The dermal bioink was transferred from a syringe into a mold by using extrusion-based bioprining. A bioprinted dermal layer was obtained and incubated at 37°C (**Figure 9**). The resulting dECS-containing dermal construct was further submerged in a larger volume of media to prevent rapid exhaustion of nutrients. Medium was changed daily. After 4days under medium submersion to allow self-assembly of endothelial networks, the dECS-containing dermal construct was carefully transferred for clinic applications such as an active dECS/cell-containing wound dressings (**Figure 9(A)**) or further

**Figure 9.**

*Proposed model of (A) an active dECS/cell-containing wound dressings and (B) an active dECS/KCs-containing wound dressings.*

was bioprinted an additional keratinocytes (KCs) layer to obtain an active dECS/KCscontaining wound dressings for wound management (**Figure 9(B)**).

#### **4. Design of synthetic soft polymeric materials with biomimetic 3D with 3D bioinspired structural foam-wall microarchitectures**

For regenerative applications, existing biomedical membranes such as collagen membranes have demonstrated the ability to promote or inhibit cell proliferation. Pure collagens suffer from uncontrollable rapid degradation and weak mechanical strength [34–37]. Generally, synthetic polymers have better mechanical properties and stability than collagen [38]. Polyvinyl alcohol(PVA), which is an FDA-approved material, was widely used in various biomedical applications, including surgical sponges, osteochondral grafts, contact lenses, artificial blood vessels, and implantable medical devices [39, 40] (e.g., bone regeneration [41], wound healing [42], and dental applications [43]) because of the desirable properties such as biocompatibility, nondegradability, low protein absorption, and easily tunable mechanical properties. PVA matrix was characterized by its super water absorbent property, great durability and cleaning ability, and its super soft texture when moist. Using sulfuric acid as catalyst and a suitable pore-forming agent, this porous PVA foam was prepared through PVA acetalization. In usual, some pore-forming agents, such as starch, surfactants, or the reagents [44, 45], which could be employed to produce gas during the cross-linking reaction to prepare PVA foam matrix. Also, water was used as a pore-forming agent to obtain porous structure without any other additional pore-forming agents [46]. Starch was a pore-forming agent that had been commonly used in the preparation of porous PVA matrix [47, 48]. The average pore diameter of PVA foam varied from 30 to 60 μm when wheat-starch was used and varied from 60 to 100 μm when potato starch was used [49, 50]. However, the volumes of PVA matrix gradually shrunk with the increasing acetalization degree and water was continuously excluded during the

#### *Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

acetalization process as observation. The PVA matrix with compacted or closed cell porous structure and a poor interconnectivity was obtained. The resulting PVA foam had residual pore-forming agent and undesirable degradable products, which might be harmful for biomedical applications such as tissue engineering or wound management [50]. The PVA/collagen composite materials were successfully developed in the forms of patches, nanofibers, hydrogels, polymer blend, mixed foam, and dual layers [17–55], which were mainly employed for osteochondral defects application [51], tissue-engineered corneas [52], cartilage tissue engineering [53], and wound healing [54, 55]. PVA was employed to modulate the mechanical properties, degradation properties, and cell regulation ability of PVA-based composite materials, which would show the ability to regulate the cell adhesion and proliferation behavior of cell and play a critical role in tissue regeneration [56, 57]. With a strong focus on therapeutic applications, the characteristics of composite materials for medical or regenerative applications examine the inspirations from nature [58, 59].

#### **4.1 Design of synthetic soft polyvinyl alcohol materials with 3D bioinspired structural foam-wall microarchitectures**

Biomimetic designs would bring effective materials that are sources of inspiration to biomedical engineers. For tissue engineering and medical applications, a novel bioinspired soft matrix with supporting interconnective foam-wall microarchitectures and/or struts made of cross-linked polyvinyl alcohol with air cavities inspired by avian skeleton and feather rachises was designed. The fully open-cell microstructures with air cavities, structural foam-walls, and structural pneumaticity bioinspired by avian feather rachises and pneumatic bone could be designed. The foam-wall showed a "foam-in-a-foam" microstructure and provided internal reinforcements and pneumaticity [60, 61].

Also, the corresponding foaming process was established. The bioinspired soft matrix with high interconnectivity was fabricated by using a designed air stream pore-forming process [34], which could regulate different pore sizes and microstructure of resulting materials as shown in **Figure 10**. Furthermore, the bioinspired synthetic soft polyvinyl alcohol material was employed to prepare a series of biological modified synthetic soft polyvinyl alcohol materials to provide new functional characteristics for specific clinic and regenerative applications.

The traditional designs of polyvinyl alcohol matrix were prepared by traditional starch pore-foaming process or air-assisted starch pore-foaming process. The medical drainage materials with fully open-cell microstructure could not be obtained. It is difficult to form air cavities to provide interconnectivity and structural support. To build up the stable air cavities, foam walls, and structural pneumaticity bioinspired by avian feather rachises, the introduction of clean atmospheric flow in the foaming process of polyvinyl alcohol matrix would be considered (**Figure 11**). The clean air was incorporated during the pore-foaming process and cross-linking process for formation of air cavities, foam walls, and structural pneumaticity bioinspired by avian feather rachises and pneumatic bone, which could be considered as a biomimetic airstream pore-foaming process. The clean air current was employed to avoid impurity. Complete cross-linking reaction for preparation of polyvinyl alcohol foam materials was also important. During starch-containing pore-foaming process, the pore-foaming agent, starch, could not provide enough driving force to form atmospheric flow, which would promote formation of air cavities with structural support, foam walls, and structural pneumaticity. New biomimetic design of PVA matrix with

#### **Figure 10.**

*Macroscopic images of designed synthetic soft polyvinyl alcohol materials with different pore sizes and microstructure. (A) Diameters were ca. 300:(A1) dried type, (A2) subhumid type, (A3) moist type; (B) Diameters were ca. 500:(B1) dried type, (B2) subhumid type,(B3) moist type; and (C) Diameters were ca. 900:(C1) dried type, (C2) subhumid type, (C3) moist type.*

air cavities inspired by avian skeleton and feather rachises was prepared by using a designed air flow pore-foaming process without starch pore-foaming agent. A novel b-PVA with fully open cells and channels could be designed and prepared by using a clean airstream pore-foaming process. The reduced pressure could be employed to build up a clean airstream pore-foaming environment. Furthermore, a morphological evaluation of the resulting was carried by using SEM as shown in **Figure 6**. The resulting PVA matrix exhibited spongy structure with fully open-cell interconnecting porous network.

Porous matrices served as a guide for tissue regeneration as a three-dimensional substrate with temporary mechanical support for cell attachment, proliferation, infiltration [62]. The morphology was the key feature that affects both biological and mechanical efficiency of the medical matrices, which would be needed to provide a porous architecture with high interconnectivity to enable cell infiltration, nutrient flow, and integration of the material within the host tissue [63]. Various matrix manufacturing techniques such as emulsion templating [63], gas foaming of heterogeneous blends [64], electrospinning [65], and additive manufacturing [66] had been widely used to introduce porosity into tissue engineering scaffolds. However, PVA matrix always exhibited no interconnectivity or limited interconnectivity

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

#### **Figure 11.**

*Microstructures of new biomimetic design of bioinspired PVA matrix were inspired by avian skeleton and feather rachis containing foam-walls. (A) avian feather rachis, (B) microstructure with air cavities stents for structural support, (C) avian skeleton pneumatic bones, (D) microstructure with air-filled canals for pneumatized drainage, and (E) morphological evaluation of new biomimetic design of bioinspired PVA matrix by using scanning electron microscopy (SEM).*

#### **Figure 12.**

*The microstructural morphological evaluations of commercial soft materials with limited interconnectivity, (A) blind pore, (B) compacted pore, and (C) inaccessible pore.*

because of inaccessible pore, blind pore, and compacted pore within the microstructure (**Figure 12**), which might be due to poor foaming ability by using traditional starch pore-foaming process, air-assisted traditional starch pore-foaming process, or incomplete cross-linking reaction of PVA. Biomimetic airstream pore-foaming process was employed as a new matrix fabrication technique to prepare a bioinspired polyvinyl alcohol (b-PVA) with a foam-wall microstructure, which exhibited high porosity and high interconnectivity. The resulting high porosity and interconnectivity would enable cell migration, vascularization, and providing space for newly forming tissues to meet the requirements of specific tissue engineering applications and clinic treatments [66–68].

To understand advanced thermal stability of the synthetic soft polymeric materials with 3D bioinspired structural foam-wall microarchitectures matrix(b-PVA), the commercial medical PVA matrix was also studied by using TGA and DSC and compared with the results of b-PVA. The weight loss curves obtained from thermogravimetric analysis (TGA) could provide some information of thermal degradation such as small molecules, solvent, water, residual reagents, residual foaming agents, starch, weak structural molecules, and prepolymeric molecules, to identify the thermal structural stability. Most of commercial medical PVA matrix exhibited poor thermal stability, which might be due to the compacted and closed-cell microstructure [69]. With increasing temperature, several thermally degraded products and the weight loss of materials were observed. From TGA and DTG results of commercial medical PVA matrix, three stages of weight loss, such as stage I in 50–100°C, stage II in 100–250°C, and stage III in 250–500°C. First, weight loss of commercial medical PVA matrix in stage I at around 100°C was contributed to the loss of water molecules trapped in the hydrophilic PVA. The weight losses of stage I were observed in a range of 3–5 wt%. Second, TGA and DTG spectra of commercial medical PVA matrix in stage II exhibited slight thermal hydrolysis signals increasing with temperature, indicating that the materials might have a poor thermal stable structure and impurity. The weight losses of stage II were determined to be in a range of 2–9 wt%. The relative very low values of initial hydrolysis temperature in stage II were observed at 110–160°C, indicating the thermal degradation of the commercial medical PVA matrix would be happened easily and harm seriously for biomedical applications such as biocompatibility and tissue engineering. Third, TGA and DTG spectra of commercial medical PVA soft matrix in stage III exhibited quite broad peak of DTG curve and the relative low values of initial hydrolysis temperature in stage III were observed in a range of 220–250°C, indicating commercial medical PVA soft matrix might be incomplete cross-linked structures with poor thermal structural stability and incomplete entanglement microstructure with closed cells [34]. Differential scanning calorimeter analysis (DSC) of the commercial medical PVA matrix exhibited some peak below 100°C such as 56, 65, 68, 75, and 95°C, which might be due to the residual compounds such as unreacted PVA or starch pore-forming agents [70]. In general, the conventional method of manufacturing PVA matrix by using the pore-forming agents such as wheat starch (61.8°C [71], 56.2°C [72], and 68.2°C [73]), pea starch (59.8°C [72]), corn starch (71.2 and 80.0°C [71]), potato starch (66.6°C [71], 63.0°C [72], and 72.2°C [73]), cassava starch (65.1°C [73]), tapioca starch (69.9°C [73]), maize starch (76.7°C [73]), and rice starch (68.3°C [71] and 67.8°C/75.3°C [72]). The DSC values of the commercial medical PVA matrix observed at 84–110°C might be contributed to amylopectin part in starch. In the case of the amylopectin part in sago starch, the melting temperature peak was between 50 and 150 °C [74]. Most of the commercial products with compacted and closed cell microstructure were determined and characterized; the results might be contributed to the evident of residual starch after starch pore-foaming process, which could not provide a good support and drainage microstructure. The addition of starch and thermal depredated compounds would be harmful to the clinical applications and risks of pollution in storage system of the resulting soft medical drainage material.

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

**Figure 13.** *Thermal characteristics of new bioinspired PVA(b-PVA) matrix, (A) TGA and (B) DTG.*

Also, the residual starch would be degraded with an increasing temperature and enhance risks of treatments.

From **Figure 13**, weight loss curves obtained from thermogravimetric analysis (TGA) of biomimetic PVAF could exhibit excellent thermal structural stability and high purity. First, TGA and DTG curves of bioinspired PVA matrix in the region with the temperature lower than 100°C exhibited 6wt% of weight loss, which would be due to only water molecules escaping from the materials and a good water-absorption property. Also, DSC curve of bioinspired PVA matrix was employed to identify thermal stability and exhibited a smooth signal below 100°C, which was contributed to escaping water molecules. Second, TGA and DTG curves of bioinspired PVA matrix in the region with a temperature range between 100 and 300°C exhibited no change of thermal signals, indicating high purity of synthetic soft materials. Third, TGA and DTG curves of bioinspired PVA matrix in the region with the temperature higher than 300°C exhibited a high Tdmax value >425°C and a narrow peak of DTG curve, which indicated high thermal and structural stability and a fully cross-linked structure (**Figure 13**).

A novel bioinspired soft matrix derived from PVA was designed and prepared. Fourier-transform infrared spectra of the designed synthetic soft polymeric materials with 3D bioinspired structural foam-wall microarchitectures (b-PVA) were determined. The main bands of pure PVA at 3500–3400, 2917, 1425, 1324, and 839 cm−1 were assigned to the O-H stretching vibration of the hydroxy group, CH2 asymmetric stretching vibration, C-H bending vibration of CH2, C-H deformation vibration, and C-C stretching vibration by using FTIR. After the biomimetic airstream pore-foaming process, the molecular structure of PVA matrix was also characterized. The bands at 3500–3400cm−1 were weakened and shifted toward higher frequencies due to the consumption of -OH (due to condensation reaction be acetalization tween PVA and formaldehyde) and the corresponding cleavage of the intra- and intermolecular hydrogen bonding. The new bands at 2952, 2913, 2861, and 2675 cm−1 ascribed to symmetric stretching vibrations of the alkyl CH2 group, the new bands at 1239, 1172, 1129, and 1065 cm−1 ascribed to the stretching vibration of C-O in C-O-H groups, and a new peak at 1008 cm−1 ascribed to -C-O-C-O-C- stretching vibrations could confirm the formation of a formal structure of b-PVA [75, 76]. Starch of corn, cassava, and potato showed similar absorption bands below 1200 cm−1 [77, 78]. A absorbance band around 1150 cm−1,

bands at 1080 and 1020 cm−1, and bands around 700–900 cm−1 were contributed to vibrations of the glycosidic C–O–C bond [34, 76], the anhydroglucose ring O–C stretch [34, 77], and C-O-C ring vibration [34, 76], respectively [34, 76, 77]. The absorbance band at 924 cm−1 was assigned to vibrational modes of a skeletal glycoside bonds of starches [34, 77]. FTIR spectrum of commercial medical PVA matrix showed the similar absorbance band at 924 cm−1. However, the absorbance band at 924 cm−1 could not be observed in the FTIR spectra of new bioinspired PVA matrix. The absorbance band at 924 cm−1 might be an important evident for starch pore-foaming process and an evaluation for a structural support cell microstructure. The fingerprint region at wavenumbers 1200–600 cm−1 could be employed to identify the purity and molecular structure of suitable medical soft matrix. That is, the fingerprint region of new biomimetic design of medical soft matrix could identify the biomimetic super-clean airstream pore-foaming process different from those of traditional starch pore-foaming process. The residual starch pore-foaming agents were degraded easily with an increasing temperature, which might be harmful to the cleanliness of the medical soft matrix and enhance risks of pollution in biomedical application such as tissue engineering or wound management [79, 80].

For the therapeutic application of liver, bioartificial liver support system has been proposed to support the regeneration of the patient's liver [81, 82]. The PVA matrix might be considered as a suitable material. The b-PVA was a nontoxic material with favorable mechanical properties, chemical stability, and good cell adhesion, which could be easily processed and has a special three-dimensional porous structure [83, 84]. Fabrication of porous matrix as potential hepatocyte carriers for bioartificial liver support had been widely explored [85, 86]. Previous studies have tried to immobilize protein with biomaterials by using covalent binding for HepG2 cell culture [85, 86]. The designed b-PVA matrix showed a microstructure containing foam-walls with good structural support and interconnectivity, which might act as a feasible material for artificial liver devices because of biocompatibility, pore-foaming, and mechanical property [85, 86].

The human hepatoblastoma HepG2 cells could be employed for evaluation of liver therapeutic application by using an extrusion-based bioprinting procedure as shown in **Figure 14A**). The ability of cells to adhere and proliferate on the bioinspired polyvinyl alcohol (b-PVA) matrix was an important indicator for evaluating the in vivo application potential. The SEM images were used to evaluate the cell morphology on the bioinspired polyvinyl alcohol matrix after various times of human hepatoblastoma HepG2 cell culture such as 72 h as shown in **Figure 15**. After 72 h of culturing, HepG2 cells would be adhered to the scaffold and grew cover on the interface and inside of the open-cell pores. The foam-walls might provide a rough surface for HepG2 cell proliferation. The interconnected microporous microstructure of the matrix might well supports the penetration of HepG2 cells.

#### **4.2 Design of biological modified b-PVA with 3D bioinspired structural foam-wall microarchitectures**

The designed bioinspired PVA(b-PVA) was employed to form a relative biocompatible surface of a collagen-modified b-PVA, which might provide more suitable characteristics for cell adhesion and growth. Some domains in collagen molecules, such as Phe-Hyp-Gly segment, could act as ligands for the integrin family *Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

#### **Figure 14.**

*Bioprinting applications of novel active liver-repairing membranes derived from synthetic soft polymeric materials with biomimetic 3D microarchitecture. (A) Active liver tissue-repairing membranes and (B) active dNBS-containing liver tissue-repairing membranes.*

**Figure 15.** *SEM photographs of HepG2 cells grew on the b-PVA after 72 h (scale bar x10 μm).*

receptors of cell surface and potentially activate specific biological signals to promote cell attachment and proliferation. The fully open cell and open foam-wall microstructure of b-PVA was maintained after preparation of collagen-modified b-PVA (**Figure 16**). The b-PVA and bovine Achilles tendon type I collagen were employed to prepare a bioinspired bovine Achilles tendon type I collagen-modified b-PVA composite matrix with open-cell foam-wall microarchitectures. The hydrophilic hydroxyl group within b-PVA molecules might provide a good binding with the coated bovine Achilles tendon type I collagen molecules by using hydrogen bonding and Van der Waals forces [87, 88]. Similar behavior was observed by Zhou et al. [89]. The interaction between b-PVA and bovine Achilles tendon type I collagen molecules within collagen-modified b-PVA was based on non-covalent interactions such as hydrogen

#### **Figure 16.**

*Interconnected microporous microstructure of medical drainage materials, (A) the original bioinspired PVA matrix and (B) the bioinspired collagen-modified PVA composite matrix.*

bonding and Van der Waals forces. The hydrogen bonding would be the main force between PVA and collagen interactions, in which collagen molecule might act as a hydrogen donor and forms hydrogen bonds with the hydroxyl group of b-PVA [90]. The bioinspired soft matrix combined the structural properties of the synthetic PVA matrix with the high cell affinity of the type I collagen, thus enhancing the matrix cytocompatibility. Collagen was the dominant component of the extracellular matrix and had been widely used in tissue engineering due to its low antigenicity, good biocompatibility, biodegradability, and nontoxicity [91]. The FTIR analysis was employed to confirm changes of molecular structure after the surface modification on the PVA matrix with collagen molecules. The absorbances at ~3300, ~2930, ~1636, ~1571, and ~1150 cm−1 were observed and characterized for amide A (N-H stretching), amide B(the asymmetrical stretching of CH2 vibration), amide I(hydrogen bonding between N-H stretching and C=O), amide II(N-H bending and C-N stretching), and amide III(C-N bending, and N-H stretching), respectively. Remarkably, the collagen-modified bioinspired PVA composite matrix showed quite different absorbance from original bioinspired PVA matrix in the region of 1200–1800 cm−1. After the biological surface modification on the b-PVA with collagen molecules, the resulting biological collagen-modified b-PVA still sustained a favorable biomimetic interconnected porous structure**.**

The morphology of the bioinspired collagen-modified b-PVA with high interconnectivity was investigated and confirmed. After 72 h of culturing, HepG2 cells were remarkably adhered to the bioinspired collagen-modified b-PVA and grew cover on the interface and inside of the open-cell porous microstructure. The SEM images intuitively displayed the growth and distribution of cells on the resulting bioinspired and collagen-modified b-PVA. HepG2 cells adhering to the type I collagen-modified b-PVA showed filopodia, which indicated that human hepatoblastoma HepG2 cells on type I collagen-modified b-PVA in a relative better proliferative condition than pure b-PVA. The stable adhesion and complete spreading could be observed in **Figure 17**. Majhy et al. reported that facile surface modification of substrate could enhance the cell-substrate interaction from a non-adherent state for promoting cell culture [62, 92]. The HepG2 cells attached and grew in cluster on the bioinspired and collagen-modified b-PVA, which was quite different from original b-PVA. The bioinspired and collagen-modified b-PVA was conductive to HepG2 proliferation, migration, expression, and functionality maintenance. *Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

#### **Figure 17.**

*Design of (A) bioinspired PVA composite matrix and the corresponding (B) dNBS-modified biological bioinspired PVA composite matrix.*

#### **Figure 18.**

*SEM photographs of HepG2 cells grew on the collagen modified b-PVA (scale bar x10 μm).*

HepG2 cells proliferated actively and formed cell clusters more efficiently in the collagen-modified b-PVA than original b-PVA (**Figure 18**). The collagen of collagen-modified b-PVA appeared to promote the growth and differentiation of the HepG2 cells. As shown in **Figure 18**, after 72 h, cells attached and grew well on the collagen-modified PVA composite matrix and a large number of pseudopods could be seen. The HepG2 cells would be evenly distributed on the surface of the pore of the b-PVA. Similarly, Moscato et al. reported that poly(vinyl alcohol)/gelatin hydrogels were good for growth of HepG2 cells [63, 93]. In the future, the bioinspired and collagen-modified b-PVA might serve as a promising liver cell culture carrier to be used in the biological artificial liver reactor. Both the type I collagen-modified b-PVA and the corresponding b-PVA could be considered as good materials for different tissue engineering applications. Furthermore, the type I collagen could be combined with dNBS to prepare new designed collagen/dNBS-modified b-PVA, which might provide additional 3D microenvironments within porous structure for cell adhesion and growth as shown in **Figure 17**. Also, the human hepatoblastoma HepG2 cells could be employed for evaluation of liver therapeutic application such as active liver-tissue-repairing membranes by using an extrusion-based bioprinting procedure as shown in **Figure 14(B)**.

#### **4.3 Design of synthetic chitosan oligosaccharides-modified PVA composite matrix(COS/b-PVA) with 3D bioinspired structural foam-wall microarchitectures**

Chitosan oligosaccharide (COS) is a biomaterial obtained by chemical or enzymatic degradation of chitosan derived from shrimp and crab shells [94]. The resulting chitosan oligosaccharide(COS) is composed of 2–10 glucosamines linked by β-1, 4 glycosidic bonds and the molecular weight is up to 3900 Da. The b-PVA was soaked in chitosan oligosaccharide (COS) solution and then frozen-dried. Synthetic chitosan oligosaccharides-modified b-PVA composite matrix(COS/b-PVA) with 3D bioinspired structural foam-wall microarchitectures was obtained. The open-cell microstructures were retained in chitosan oligosaccharides-modified polyvinyl alcohol(COS/b-PVA) (**Figure 19**). The connecting holes and the corresponding open channels were fully covered and filled with chitosan oligosaccharides, which provide a controlled release system of active chitosan oligosaccharide molecules for wound managements. The water permeability of the resulting chitosan oligosaccharide-modified COS/b-PVA dressings was determined by ASTM D4491 standard test methods. The good water permeability was observed in a relative higher level than 80%, which would be contributed to specific clinic applications [94–96].

#### **4.4 Design of soft polymeric materials with 3D bioinspired structural foam-wall microarchitectures using a mixed foaming procedure combined synthetic polyvinyl alcohol with natural chitosan**

A new form of polymer blend, macroporous chitosan/poly(vinyl alcohol) (PVA) foams made by a starch expansion process, exhibits the functionalities of chitosan while avoiding its poor mechanical properties and chemical instabilities. The appropriate conditions for foaming are discussed using both insoluble and water-soluble chitosan. Both insoluble/PVA and water-soluble chitosan/PVA foams demonstrated interconnected and open-cell structures with large pore size from tens to hundreds of micrometers and high porosities from 73.6 to 84.3%. The chitosan/PVA with high interconnected and open-cell structures might have potential in the tissue

#### **Figure 19.**

*(A) Photo of the designed permeable chitosan oligosaccharide modified polyvinyl alcohol complex foam (COS/b-PVA) and (B) morphology of designed chitosan oligosaccharide modified polyvinyl alcohol complex foam (COS/b-PVA) dressing with fully open-cell and open-channel microstructures.*

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

#### **Figure 20.**

*Active tissue-repairing sponges made from chitosan/poly(vinyl alcohol) (Supported by Cenefom and PARSD).*

engineering and bioprinting applications [97]. The clean airstream pore-foaming process for b-PVA could be employed to build up a blend-foaming process for preparation of chitosan/poly(vinyl alcohol) (b-PVA). The chitosan/poly(vinyl alcohol) (PVA) was obtained and applied for clinic treatments such as active tissue-repairing sponges (**Figure 20**).

#### **5. Clinic application of active tissue-repairing membranes by using 3D bioprinting and cells culture procedure**

New biomimetic biomedical inventive design-thinking methods could be established for design of new clinic application such as active tissue-repairing membranes by using 3D bioprinting and cells culture procedure [98].

#### **5.1 Clinic application of synthetic soft active tissue-repairing membranes by using 3D bioprinting and cells culture procedure**

The collagen-coating PVA could be employed as a 3D-bioprinting supporting scaffold for 3D bioprinting applications of a tissue-repairing membrane such as active skin-tissue-repairing membranes and liver-repairing membranes.

The collagen-coating b-PVA could be cultured HepG2 cell to design a liverrepairing membrane as shown in **Figure 15**. Also, the bioprinting of a human skin substitute containing xeno-free cultured human EC, FBs, and PCs in a xeno-free bioink (**Figure 21(A)(A1)**) containing human collagen type I and fibronectin layered in a biocompatible collagen-coating b-PVA supporting scaffold (**Figure 21(A)(A2)**). After cells of EC, FBs, and PCs being cultured and transferred into the biocompatible collagen-coating b-PVA supporting scaffold (**Figure 21(A)(A3)**), a designed active skin-tissue-repairing membranes containing human EC, FBs, and PCs could be obtained for clinical treatments (**Figure 21(A)(A4)**). Furthermore, dNBS could be introduced into the bioink (**Figure 21(B)(B1)**) containing EC, FBs, PCs, and dNBS, human collagen type I, and fibronectin and layered in a biocompatible

#### **Figure 21.**

*Bioprinting applications of novel active skin-repairing membranes derived from natural and synthetic soft polymeric materials with biomimetic 3D microarchitecture. (A) active skin dermal tissue-repairing membranes, (B) active dNBS-containing skin dermal tissue-repairing membranes, and (C) active skin epidermal tissuerepairing membranes.*

collagen-coating b-PVA supporting scaffold (**Figure 20(B**)(**B2**)). The collagencoating b-PVA could be cultured EC, FBs, and PCs with dNBS (**Figure 20(B)(B3)**) to obtain a novel active dNBS-containing skin-tissue-repairing membranes containing human EC, FBs, PCs, and dNBS (**Figure 21(B)(B4)**). Independently, human KCs were introduced into a xeno-free bioink (**Figure 21(C)(C1)**). Furthermore, the collagen-coated PVA could be bioprinted with the resulting KCs bioinks (**Figure 20(C)**(**C2**)) and cultured with xeno-free human KCs (**Figure 21(C)(**C3**)**) to form active epidermal tissue-repairing membranes for clinic application in wound management (**Figure 21(C)(C4)**).

For the clinic applications, the novel active skin dermal tissue-repairing membranes, dNBS-containing active skin dermal tissue-repairing membranes, and active skin epidermal tissue-repairing membranes could be applied directly on the target wounds for different repairing steps such as dermal repairing (**Figures 22(A)** and **23(A)**) and subsequently epidermal repairing (**Figures 22(C)** and **23(C)**). Also, the additional PBS or saline could be employed to transfer the cells within the wound dressings onto the surface of wound depending on the clinic needs for different repairing steps such as dermal repairing (**Figures 22(B)** and **23(B)** and subsequently epidermal repairing (**Figures 22(D)** and **23D)**). The dNBS-containing active skin repairing wound dressings could provide a natural biomimetic 3D microarchitecture

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

#### **Figure 22.**

*Novel active skin-repairing membranes derived from synthetics oft polymeric materials with biomimetic 3D microarchitecture for wound managements. (A) Active skin dermal tissue-repairing membranes without transferring assistances, (B) active skin dermal tissue-repairing membranes with transferring assistances, (C) active skin epidermal tissue-repairing membranes without transferring assistances, and (D) active skin epidermal tissue-repairing membranes with transferring assistances.*

for promotion of cell growth. The collagen-coating PVA matrix was considered as a good cell carrier for cell culturing and transferring.

#### **6. Conclusions**

Design and characterization of natural and synthetic soft polymeric materials with biomimetic 3D microarchitecture were considered. The natural soft polymeric

#### **Figure 23.**

*Novel active dNBS-containing skin-repairing membranes derived from synthetic soft polymeric materials with biomimetic 3D microarchitecture for wound managements. (A) active dNBS-containing skin dermal tissue-repairing membranes without transferring assistances, (B) active dNBS-containing skin dermal tissuerepairing membranes with transferring assistances, (C) active dNBS-containing skin epidermal tissue-repairing membranes without transferring assistances, and (D) active dNBS-containing skin epidermal tissue-repairing membranes with transferring assistances.*

materials would focus on new design bioinspired membranes containing supercritical fluids-decellularized dermal scaffolds for 3D bioprinting potential applications. Synthetic soft polymeric materials would focus on b-PVA with structural foam-wall microarchitectures. The b-PVA was designed, prepared, and showed relative higher interconnectivity than commercial medical soft matrix, which could be considered as a good potential scaffold for tissue engineering. Furthermore, the b-PVA was employed to prepare a highly biocompatible collagen-modified b-PVA composite

*Perspective Chapter: Design and Characterization of Natural and Synthetic Soft Polymeric… DOI: http://dx.doi.org/10.5772/intechopen.106471*

matrix after surface modification with type I collagen. In the future, the collagenmodified b-PVA composite matrix might serve as a promising liver cell culture carrier to be employed in the biological artificial liver reactor.

#### **Acknowledgements**

Author would like to acknowledge the Taiwan PARSD Pharmaceutical Technology Consultants Ltd. Company for financial and technical support. Additionally, the author would like to thank Gui-Feng Zhang and Yuguang Du for their technical assistance from the State Key Laboratory of Biochemical Engineering, Institute of Process Engineering, and Chinese Academy of Sciences, Beijing.

#### **Conflicts of interest**

The authors declare no conflict of interest.

#### **Author details**

Ching-Cheng Huang1,2\* and Masashi Shiotsuki3

1 Department of Biomedical Engineering, Ming-Chuan University, Taiwan


\*Address all correspondence to: jcchuang.mcu@gmail.com

© 2022 The Author(s). Licensee IntechOpen. 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.

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#### **Chapter 4**

## Novel Composites for Bone Tissue Engineering

*Pugalanthipandian Sankaralingam, Poornimadevi Sakthivel and Vijayakumar Chinnaswamy Thangavel*

#### **Abstract**

Novel metal oxide-doped fluorophosphates nano-glass powders were synthesized by melt quenching method, and their non-toxicity is proved by MTT. Their efficacy in bone formation is confirmed by osteocalcin and ALP secretion. Composites were made using PLA, PDLLA, PPF, or 1,2-diol with fluorophosphates nano-glass powders (AgFp/MgFp/ZnFp). Their non-toxicity was assessed by cell adhesion and MTT. The ability of the composite for bioconversion was assessed by RT-PCR estimation for osteocalcin, Collagen II, RUNX2, Chondroitin sulfate, and ALP secretion accessed by ELISA method. The animal study in rabbit showed good callus formation by bioconduction and bioinduction. The bioconversion of the composite itself was proved by modified Tetrachrome staining. From the 12 different composites with different composition, the composite PPF+PDLLA+PPF+ZnFp showed the best results. These obtained results of the composites made from common biological molecules are better than the standards and so they do biomimic as bone substitutes. The composites can be made as strips or granules or cylinders and will be a boon to the operating surgeon. The composite meets nearly all the requirements for bone tissue engineering and nullifies the defect in the existing ceramic composites.

**Keywords:** fluorophosphate, nano bioactive glass, bioconversion, bone substitute, synthetic bone graft

#### **1. Introduction**

Scientists have tried natural polymer, synthetic polymer, various ceramics, and composites, etc., as bone graft substitutes. An ideal bone substitute should be nontoxic, osteoconductive, osteoinductive, bioconvertable, physical properties should be near to that of normal bone, should be cost-effective, easily sterilizable, and be feasible for bulk production. The commonly available bone graft substitutes do not meet all the above requirements [1, 2].

The materials that were put to use initially were ceramics (hydroxyapatite and tricalcium phosphate) as bone graft substitutes. Hydroxyapatite was only osteoconductive and rarely was converted to bone even after years [3, 4]. It was not useful in replacing weight-bearing function. Tricalcium phosphate had minimal osteoinductive capacity along with osteoconduction but had no bioconversion capability [5, 6].

To have the advantage of bioconversion, certain specific bone hormones such as "Bone Morphogenic Protein" (BMP) came into use. "Demineralized Bone Matrix" (DBM) was also marketed as bone graft substitute, which was discovered by Urist [7]. The essential problem in their use is the phenomenal cost involved, and they had good osteoinduction but were not good osteoconductors. Hench [8] came out with the 45S5 glass, which was a breakthrough as it was made from cheap chemicals, was osteoconductive as well as osteoinductive, was able to merge with the natural bone, and is commercially available. The drawback with 45S5 glass is their very slow resorption, the longer time taken for bioconversion, and their inability to be used as a weightbearing implant. The time taken for bioconversion was a drawback in reducing the duration of morbidity of the patients. To circumvent these problems, silica-free phosphate bioglasses and metal-oxide-doped bioglasses entered into these fields.

Over decades, orthopedic surgeons know chronic ingestion of fluoride in abundance leads to fluorosis, an ectopic bone-forming condition [9]. This was used to our advantage, and the ideal concentration of fluoride in the ceramics network was standardized. The evaporation of fluorine from the fluoride compound was circumvented by the new methodology of preparing fluoride bioglasses by melt quenching.

Standardization of the ideal mole percentage of fluoride resulted in the invention of fluorophosphates glasses, which are much more bio active than other types of phosphate and silica glasses and had a higher rate of bioconversion and faster resorption [10, 11]. Doping them with metal oxides improved their physical properties and brought the elastic moduli close to that of the human bone. Scaffolding the fluorophosphate glasses was essential to bring the molecule for clinical use that will bridge the gap between the need of the surgeon and the capability of the scientist.

This novel synthetic composite is made from bioinert polymers comprising poly lactic acid, poly D, L -Lactic acid, and bioactive polymers consisting of polypropylene fumarate, diester of fumaric acid, and 1,2 propylene diol and a bioactive inorganic component consisting of a metal-doped fluorophosphates nano glass powder. It can be fabricated as granules, scaffolds such as strip and cylinder. The plurality of the shapes that can be made gives additional advantage to the surgeon.

The standard biomaterial should mimic the biological process in the microenvironment simply denoted as "Biomimetics." The selected three different metals have beneficial properties as follows: (a) Silver: The best antimicrobial agent is especially in wound healing and skin care. (b) Magnesium: It plays a vital role in bone structure development and also acts as dietary supplement in medicine. (c) Zinc: It involves in DNA synthesis, gene expression, and wound healing. Also, the degradation products of the biocompatible polymers (PLA, PDLLA, PPF, and 1,2-diol) all enter the Kreb's cycle and excreted. So, the metal oxides and the polymers induce bone formation by stimulating the genetic pathway of bone formation and so are biomimetic.

#### **2. Materials**

Poly lactic acid (PLA) and poly DL-lactic acid (PDLLA) were procured from BioDegmer® Japan. Polypropylene fumarate (PPF) was prepared by

*Novel Composites for Bone Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.106255*

transesterification method [12]. The di-1,2-propanediol ester of fumaric acid (1,2 diol) was prepared by azeotropic distillation, and its detailed explanation is available in previous publication [13].

The fluorophosphate glasses were prepared by melt quenching method and converted to nano particles by mechanical milling. The measured quantities of the required chemicals (Na2CO3, CaCO3, CaF2, P2O5, and ZnO/Ag2O/MgO) were taken in a ball mill and homogenized. The mixture was heated in a 10% alumina crucible for 1 h up to 120°C and cooled to room temperature. The material was again ball milled for 1 hr. The components were taken in a platinum crucible and kept in a furnace preheated to 1100°C for 90 min. Then the crucible with the material was quenched by plunging in liquid nitrogen. The formed glass was broken to pieces and milled for 48 h to obtain metal-oxide-doped nano powder of the specific fluorophosphates glass [14].

#### **2.1 Gel foam casting under rapid heating**

The required amounts of the PLA [15], PDLLA, PPF/1,2-diol, and AgFP/ ZnFP/MgFP were taken and dispersed in dichloromethane using a magnetic stirrer (300 rpm). Once the mixture was homogenized, the material was slowly poured over a hot glass plate (70°C). The rapid evaporation of dichloromethane leads to the generation of random pores in the composite scaffold formed. After complete evaporation of the solvent, highly interconnected porous scaffold with homogeneous distribution of the components was obtained. The scaffold thus made can be ground to granules, cut into strip, or rolled into cylinder [16].

To know the significance of PDLLA, scaffolds were fabricated with or without PDLLA [17]. The fabricated composites are:1. PLA+1,2-diol+AgFp, 2. PLA+PDLLA+ 1,2-diol+AgFp, 3. PLA+PPF+AgFp, 4. PLA+PDLLA+PPF+AgFp, 5. PLA+1,2-diol+ ZnFp, 6. PLA+PDLLA+1,2-diol+ZnFp, 7. PLA+PPF+ZnFp, 8. PLA+PDLLA+PPF+ZnFp, 9. PLA+1,2-diol+MgFp, 10.PLA+PDLLA+1,2-diol+MgFp, 11. PLA+PPF+MgFp, 12. PLA+PDLLA+PPF+MgFp.

#### **3. Methodology**

#### **3.1 In vitro evaluation of the scaffolds**

#### *3.1.1 Metal-doped fluorophosphates nano bioglasses.*

The biological activity (cell viability, attachment, and proliferation) of the metal-doped fluorophophate nano bioglass was ascertained by cell adhesion and MTT assay; Potential osteogenic differentiation was ascertained by (Alkaline Phosphatase Activity) and non-collagenous protein (intra and extracellular osteocalcin) of the nano glass, in relation with MG63 cell lines.

#### *3.1.2 Scaffolds (PLA+PDLLA+PPF/1,2-diol+AgFp/ZnFp/MgFp)*

The significance of the pores in the scaffold was assessed by calcein AM study and MTT evaluation. By following standard Kokubo protocol, simulated body fluid (SBF) was prepared. All the fabricated scaffolds were cut into 2X2cm2 size. The scaffolds were placed in 20 mL SBF filled glass container, for a period of 21 days at 5% CO2 incubator (Heraus – Germany). The pH variation was noted everyday using pH meter E1 model. After 21 days, the scaffolds were carefully removed; dried in laminar air flow for 48 h. The variation in the pH over 21 days was noted. The deposition of hydroxyapatite and fluorapatite was investigated by FTIR and SEM\_EDAX. The toxicity of the different composites with different composition of the components was ascertained by the MTT in relation to the SaOS2 and human endothelial cell lines. Their efficiency in enhancing secretion of alkaline phosphatase and chondroitin sulfate, the ground substance of the bone was measured by ELISA method. The ability of the composites in the secretion of osteocalcin, Collagen II, RUNX2 was assessed by RT-PCR method. The porosity and micro-architecture in the multilayered scaffold were assessed by Micro-CT evaluation with GE SRμCT analyzer.

The primers used for PCR were as follows:


The analysis of the results was performed using ABI PRISM® 7000 Sequence Detection System software that enables more sensitive and accurate estimation of the relative gene expression.

### **3.2 In vivo studies**

#### *3.2.1 For granules*

The *in vivo* studies were conducted with the Ethics committee approval (Ethical committee approval no. ABS/IAEC/18-10-2019/003-). A single species of *Orictologuscuniculus* was purchased from King Institute, Chennai, India, *Novel Composites for Bone Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.106255*

and domesticated over a period of 2 weeks. The day-night rhythm was maintained and was fed on good nourishing food. The adaptation was confirmed by the gain in weight of 150 g in 2-week time. (1800–1950 g).The composite (PLA+PDLLA+PPF+AgFp) granules were prepared by grounding the scaffold and sterilized by ethylene oxide gas.

The animal was given a premedication of pedichloryl syrup (2.5 mL) 30 minutes before surgery. Intramuscular ketamine anesthesia was given in the dose of 45 mg per kilogram body weight and waited for 10 minutes to get the full dissociated anesthetic effect. The anesthetic effect was maintained by oxygen and sevoprim inhalation through mask (**Figure 1**).

The left thigh was repeatedly painted with 10% povidone iodine and ethylene alcohol. Xylocaine 2% with adrenaline was injected in the line of incision as an additional analgesia and also a hemostatic agent. The skin incision made on the anterolateral aspect was rolled down to expose the posterior boarder of the quadriceps muscle. Using sharp dissection, the muscle was slit open and enlarged by thin bone spikes to expose the anterolateral aspect of the thigh bone. Using an electric dental burr of 1 mm, a trough was made for a length of 2 cm.This exposed the medullary cavity. It was packed with the sterile composite powder. Liberal saline wash was given to wash off the spilled over composite materials. The spikes once removed the muscle fell back into position completely covering the bony trough. The two 3–0 vicryl stitches were used to close the muscle. The skin incision, which was far away from the bone work, was closed with 3–0 ethilon. A single dose of ceftrioxazone 250 mg was given intramuscularly.

#### *3.2.2 For strip*

The *in vivo* studies were conducted with the Ethics committee approval (Ethical committee approval no. ABS/IAEC/18-10-2019/003-). Three male rabbits were

#### **Figure 1.**

*In vivo study in rabbit (granules-PLA+PDLLA+PPF+AgFp) [A-thigh preparation, B - femur exposed, C - gutter in the femur, D - gutter filled with granules E - Wound closed by sutures.*

**Figure 2.**

*In-vivo study in rabbits (strip -PLA+PDLLA+PPF+AgFp). [A - Thigh preparation, B - Femur exposed, C - Cut in the femur, D - Two strips one above another over the cut, E - Maintain position by vicryl knot].*

procured and domesticated in the same way as explained before. The AgFp/ZnFp/ MgFp composites were made with PLA+PDLLA+PPF by gel foam casting under rapid heating. They were of 1 mm thickness and cut into size of 2X20mm. The cut specimens were sterilized by ethylene oxide gas sterilization.

The animals were anesthetized, limb prepared, and femur exposed as described in the previous section. Narrow cuts were made with no.701 conical dental burr at an angle of 45° to the femur to make it extremely thin cut. The 3–0 vicryl was threaded around the femur and both the ends were kept free. Two layers of the 2X20mm sterilized composite strips were kept over the cut made allowing the marrow blood to soak the specimen. The vicryl was tied around the specimen so that the specimen does not slip or move away and the wound was closed in layers. The procedure was done for all the three specimens, one on each animal (**Figure 2**).

The animals were cared for the postoperative period with nourishing food. The day 1 X-ray did not show the specimen in either view as the specimens were translucent to the X-ray. The X-ray evaluation was done under sedation on the first, ninth, and 16th day. Clinical union occurred as early as the 15th day. The CT evaluation was done on the 19th day. The animals were euthanized as per the protocol and the limb harvested, denuded of skin and muscles, and bone preserved in 10% formalin. The X-rays of the specimens taken and then sent for histopathological evaluation in both EH stain and modified tetrachrome stain. The procedure adapted is shown in the serial photographs in **Figure 2**, where two layers of 1 mm thick strips have been placed over a very narrow corticotomy wound in the shaft of femur and have been retained in position by a single 3–0 vicryl encircling knot.

#### **4. Results and discussions**

#### **4.1 Metal-doped fluorophosphates nanobioglasses**

Alkaline phosphatase (ALP) is an osteogenic differentiation marker at all the stages from the differentiation of the mesenchymal cells to the mineralization front. Hence its enhanced secretion is considered as a vital factor to choose the ingredient for the composite for Bone Tissue Engineering (BTE). The obtained results (**Table 1**) indicated that AgFp and MgFp showed consistently raised levels at all concentration from 0.1 to 100 μg/mL, whereas ZnFp showed increased secretion only at lower concentrations of 0.01 and 1 μg/mL [18, 19].

Bone is a composite of the ground substance reinforced by multiple collagens and mineralized by hydroxyapatite [20]. Though various collagens are present in various parts of the body, osteocalcin is found exclusively in bone. It is also an excellent gene marker of bone induction. The ability of the ionic dissolution products of various FP glasses in various concentrations was evaluated for their efficiency to promote osteocalcin secretion. While the extracellular expression of osteocalcin [21] showed increase than the control only with ZnFp and MgFp, (**Table 2**) intracellular osteocalcin was raised in most of the glasses, but significant raise was present in ZnFp, MgFp, and AgFp glasses and was more when the concentration of the products of dissolution was 10 μg/mL.

The calcein AM study was used to assess the cell wall integrity and the double staining to assess the cytotoxicity showed specific features. The control group of cells was not only brilliantly green but also showed homogeneous spindle shape, indicating the integrity of cell wall and the metabolic potential. The addition of PPF to the basic


**Table 1.** *Alkaline phosphatase activity of metal-doped fluorophosphates nano bioglass powders.*


#### **Table 2.**

*Osteocalcin (intracellular and extracellular) activity of metal-doped fluorophosphates nano bioglass powders.*

**Figure 3.** *Calcein AM study of composites.*

components PLA+PDLLA increased the cell wall integrity and the addition of pores to the same increased the number of spindle shaped cells [22].

The addition of FP glass to the basic components PLA+PDLLA [23] with pores or without pores increased the number of cells phenomenally, but the quality of them was poor exhibited by their round shape rather than the spindle shape of the healthy cell. When all the components PLA+PDLLA+PPF and ZnFp glasses were added, both the intensity of fluorescence and the quality of the cells also increased, and it was more so when pores were present in the composite (**Figure 3**).The XPPF (cross-linked PPF) when replaced the PPF in the composite, there was only deleterious effect both in the fluorescence and the quality of the cells (**Table 3**).

From the above study, it can be inferred that the least toxic composite was that of PLA+PDLLA+PPF+ZnFp glass.

The three essential gene markers in the synthesis of bone from the stage of mesenchymal stem cells to that of the osteocyte maturation are Osteocalcin [24], Collagen II [25], and RUNX2 [26](**Table 4**). When the results were charted to scrutinize the fold change than the control, the fold increase in collagen II was highest with PLA+PDLLA+PPF+AgFp, and the highest fold increase in osteocalcin was also with AgFp but when constituted with 1,2-diol than with PPF. The highest fold change in

*Novel Composites for Bone Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.106255*


#### **Table 3.**

*Cytotoxicity of the prepared composites.*


#### **Table 4.**

*MT, ALP, collagen II, osteocalcin, RUNX2, and chondroitin sulfate levels of composites.*

RUNX2 than the control was with ZnFp when combined with PLA+PDLLA+PPF. All the Mg-based composites fared poorly with all the three types of gene markers (**Figure 4** and **Table 4**).

#### **Figure 4.**

The results showed that all the 12 composites showed many fold increase in the secretion of chondroitin sulfate [27] than the control, immaterial of the component having 1,2-diol or PPF and the FPglass being either Ag, Zn, or Mg.

The pH variation of all the compression-molded specimens showed uniformly a reduction in the first 2 days, which is because of phosphoric acid formation. And all the specimens bounced back to 7 on the third day due to the alkaline earth metal (Na+ and/or Ca2+) release. The dissolution of the ions thus replaces H+ ions by cations (Na+ and/or Ca2+) leading to an increase in hydroxyl ion concentration. None of them went below 6.5 even in the first 2 days. From then on it showed a steady variation between 7 and 6.7.

The strip and cylinders made by gel foam casting under rapid heating showed a better pH even in the first 2 days and never went below 6.8, and the end stage also showed higher pH than the scaffolds. The highest pH reached was with the strip of scaffold made by rapid heating method, and it was 7.15. This variation shows the better homogenezity and the porosity achieved by the rapid heating method, which avoids high acidic environment that can lead onto graft rejection [28]. The crystal formation over the strip and cylinder after *in vitro* evaluation is shown in **Figure 5**.

The predominant functional groups present in the composites were studied using their respective FTIR spectra:alcohol (3200–3500 cm−1), alkanes (2850–3000 cm−1), saturated ketone (1735–1750 cm−1), alkenes (1630–1680 cm−1), asymmetric methyl bend (1450–1470 cm−1), and methyl bending (1350–1395 cm−1). The presence of P-O bend (560–500 cm−1) bands indicates the formation of calcium phosphate(CaO-P2O5) layer. The carbonate group (CO3) 2− (1400–1550 cm−1) bands showed the crystalline

#### **Figure 5.**

*Pre- and post-immersion photographs of the strip and cylinder.*

nature of the HAp layer. The bands are observed at above 3500 cm−1, which corresponds to the OH group. After 21 days of soaking in SBF, the strong intensity and frequency shift of the (CO3) 2 -,P-O-P stretch and P-O bend groups reveal the interaction of the composite and HAp precipitation [29].

The shoulder peak at 1450–1410 cm−1 coupled with the weaker peak at 870– 875 cm−1 corresponds to type B carbonate vibrations, whereas the vibration regions of type A carbonate are 1450–1410 cm−1 coupled with a band at 880 cm−1. The type A and B carbonates [30, 31] are indistinguishable in these scaffolds because the ester peaks also lie on the same region. Both type A and B carbonates are present in these scaffolds and their intensities are maximum at three selected scaffold composites (PLA+PDLLA+PPF+ZnFp, PLA+PDLLA+PPF+AgFp, PLA+PDLLA+PPF+MgFp), For the same composites, the corresponding peaks for HAp in rapid heating combined gel foam casting are higher than the compression-molded scaffolds.

Although the HAp precipitation was noted in all the fabricated scaffolds, the intensity of the carbonated group (CO3) 2−and phosphatebased group (P-O-P asymmetric and symmetric stretch, P-O bend) was observed as high in gel foam casting under rapid heating.

The SEM evaluation of the scaffolds and strips was done after gold sputtering. (Model Ultra 55; Zeiss, Oberkochen, Germany). After evaluating the surface apatite formation, the specimens were cut into two halves and turned by 90°, and the depth of the apatite formation was measured. There was no significant change in the percentage, and it was infered that all the composites have near equal conversion once the pores allow penetration of the SBF inside except the absence of PDLLA had negative significance in the extent of crystalline conversion (**Table 5**).

The similar specimens were subjected to *in vitro* evaluations, which were analyzed by the same way in the same Scanning Electron Microscope to assess the degree of surface pores and the change in crystallinity after *in vitro* study. The photograph of a stirp of composite and a cylindrical composite, both made by gel foam casting under rapid heating, shows the retention of the shape after SBF immersion for 21 days, but there was complete change in the color and the texture indicating the crystalline conversion. The SEM of scaffold in two different magnifications both before and after *in vitro* evaluation is shown in **Figure 6**, which shows very scarce


#### **Table 5.**

*Apatite conversion(%).*

amount of crystallization in the pre *in vitro* evaluation and the homogeneous pores being well exhibited. After 21 days of immersion in SBF, the crystalline conversion is observed, and all the pores have been near completely clogged with the crystals formed [32, 33].

The SEM micrographs of the scaffolds (pre- and post-immersion). In pre-immersion status shows specks of crystallization indicating the high hydrophilicity of the scaffold, and the post-immersion evaluation of the same shows complete formation to crystals, which proves the high bioresorbability of the scaffold. The SEM evaluation of multilayered scaffold made by rapid heating under low magnification shows the adequacy of pores. The pre-immersion and the post-immersion SEM micrographs clearly show the formation of sufficient crystals. The EDAX evaluation of the pre and post in *vitro* SEM confirms the high level of carbonated hydroxyapatite and fluroapatite formation in the scaffold.

The Micro-CT (**Figure 7**) evaluation of the cylindrical scaffold made by gel foam casting under rapid heating proved the following factors: a) The scaffold had no layering and was continuous. b) There was adequate porosity and the pore sizes were varying. c) The pores were all well connected by interpores. The same specimen after *in vitro* evaluations had sufficient fomation of crystals with the preservation of the deeper pores [34].

#### **4.2** *In vivo* **studies**

From the *in vitro* evaluation, PPF-based composites were shortlisted for *in vivo* studies.

*Novel Composites for Bone Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.106255*

#### **Figure 6.**

*Pre- and post-immersion of the scaffolds.*

#### **Figure 7.**

*Micro-CT studies of the cylindrical composite. (A - Pre-immersion and B – Post-immersion).*

#### *4.2.1 Granules*

It was found that the femur had fractured and the ends were apart (**Figure 8**). Neither immobilization of the femur or any form of fixation was done. The rabbit was not limping and was feeding well. There was only a flare of the ends of the fracture and there was no evidence of any callus on the ninth postoperative day. After another week (Day 16), the limb when examined clinically had sound union. The X-ray taken showed abundant callus not only in the fracture end but all along the femur where the trough had been made and even below.

The animal was euthanized, the limb harvested, skin and muscles were peeled off, and an abundant amount of callus was found to have united the fracture very strongly. The dissected specimen were studied by X-ray and the specimen were preserved in 10% formalin.

The specimen was prepared and the decalcified specimen was sectioned axially to exhibit the two segments of the femur with the intervening tissue formed. The

**Figure 8.** *X-ray photographs of the opearted site in rabbit.*

**Figure 9.** *Histopathological study of the dissected specimen.* *Novel Composites for Bone Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.106255*

specimen was stained using regular eosin-hematoxylin stain and also von kossa stain. (**Figure 9**).

The procedure done using granules of the composite has been serially shown in the photographs (**Figure 1**). Though the limb got fractured, it did not receive any specific treatment for it. Until ninth day, there was a scarce response to heal but by the 16th day it had soundly united. (**Figure 8**). The X-ray of the specimen after dissection showed the extension of the callus almost over the entire femur. The histopathological evaluation was specifically focused toward the tissue between the fractured ends where the granules had been packed. The significant observations were: a) Nearly the whole of the granules had resorbed except occasional trace of it. b) Abundant cartilage had formed between the ends indicating the enchondral ossification. c) Woven bone formed in between the ends of the fracture was a proof of the rapidity of the fusion occurring. d) The absence of multinucleated giant cells indicates the biocompatibility of the composite. e) Similar features were observed in both the staining (**Figure 9**). The modified tetrachrome staining throws much more information than the above two. a) The new lamellar bone formed in continuity with the resorbing composite granule. b) The sound union by the woven bone formed from chondral ossification. c) The abundance of osteoblasts and the osteoid. d) An exuberant neo vascularization among the fibroblasts is well seen (**Figure 10a**–**d**) [35, 36].

#### *4.2.2 Strip*

The X-ray photographs show no evidence of the placed composite strip or the corticotomy made as the composite is not radio opaque and the furrow is very narrow. But the X-rays taken on the ninth day showed all three animals had fractured their femur. No specific treatment such as immobilization or interference was done for the fracture. Clinical union occurred as early as 15th day and was confirmed by X-ray on 16th day (**Figure 11**) and CT scan on 19th day (**Figure 12**). The harvested limb after euthanizing

**Figure 10.**

*a-d represent the tetrachrome stain results of the granules (Composition: PLA+PDLLA+PPF+AgFp).*

**Figure 11.** *Studies using composite strips: X-ray photographs of the operated site in rabbit.*

**Figure 12.** *Studies using composite strips: CT scan of the fractured leg in rabbit.*

the animal showed the composite strip was adherent to the bone underneath. The X-ray of the specimens showed abundant callus along the fracture (**Figure 13**) and the composite strip was not seen in the X-ray.

The histopathological evaluation showed the following features: a) Both the layers of the scaffold had merged into one layer. b) The composite had attached to the bone beneath. c) There was abundant woven bone formed beneath the composite strip at the level of the corticotomy. d) The second layer of the composite strip kept away from the corticotomy had profuse infiltration of fibrocytes. e) The fibrous changeover in the superficial layer of the composite had abundant neovascularization These changes confirm the osteo induction potential of the composite, the ability of the composite to go for bioconversion, and high bioactivity of the composite [37–39].

The modified tetrachrome staining of the specimens with the cross section at the level of the composite confirmed the findings by EH stain and showed the additional features. **Figure 14a** shows conversion of the fragmented composite forming woven bone to heal the corticotomy made and the binding of the two layers of the composite

#### **Figure 13.**

*Studies using composite strips: Postoperative X-ray of the dissected specimen from rabbits. (A-PLA+PDLLA+PPF+AgFp; B-PLA+PDLLA+PPF+ZnFp; C-PLA+PDLLA+PPF+MgFp).*

**Figure 14.**

*a-d represent the tetrachrome stain results of the strip (Composition: PLA+DPLLA+PPF+ZnFp).*

strip and random infiltration of the layer close to the bone with fibroblasts and specks of osteiod. On higher magnification (**Figure 14b**), the fusion of the composite strip to the underlying bone by osteoid is well seen. On further magnification (**Figure 14c**), the infiltration of the composite by newly formed layers of osteoid is well made out replacing the dissolved area of the composite. **Figure 14d** shows the adhesion of the composite strip, the composite strip dissolving and disintegrating to form new woven bone healing the corticotomy, the abundant laying of new osteoid in the dissolved portion of the composite.

### **5. Conclusions**

The novel composite of PLA+PDLLA+PPF+AgFp/ZnFp/MgFp meets most of the requirements of an ideal bone substitute, and it bridges the gap between the need of the clinician and the biomaterial scientist, more than the available present-day commercial ceramic composites. It is not only conductive like HAp but also inductive. It is more inductive than Tri calcium phosphate. It takes less time for resorption than Bioglass and the presence of fluoride makes the rate of bioconversion high. As the end products of all the four components of the composite are natural ingredients of the body and induce bone formation by enhancing the genetic pathway, the composite is a real biomimetics. Though small animal studies have proved the usefulness of the composite, their efficacy has to be confirmed in clinical situations in large animals before human trial.

### **Conflict of interest**

The authors declare no conflict of interest.

### **Author details**

Pugalanthipandian Sankaralingam\*, Poornimadevi Sakthivel and Vijayakumar Chinnaswamy Thangavel Bone Substitutes, Madurai, India

\*Address all correspondence to: drpandian@bonesubstitutes.in

© 2022 The Author(s). Licensee IntechOpen. 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.

*Novel Composites for Bone Tissue Engineering DOI: http://dx.doi.org/10.5772/intechopen.106255*

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Section 3
