**Meet the editor**

Dr. Ilser Turkyilmaz obtained his dental degree from Hacettepe University, Ankara, Turkey in 1998. He completed his PhD program at the same university in 2004. Then he worked in the Department of Biomaterials, Institute of Clinical Sciences, Sahlgrenska Academy, Goteborg University, Goteborg, Sweden in 2005. He worked as an Implant Prosthodontic Fellow in the Department

of Restorative and Prosthetic Dentistry, the Ohio State University, Columbus, Ohio from 2007 to 2008. He has been working as an Assistant Professor in the Department of Prosthodontics at the University of Texas Health Science Center in San Antonio, Texas since 2008. He has also been serving as the Director of Dental School Implant Clinic since 07/2011. Dr. Turkyilmaz earned Diplomate status within the International Congress of Oral Implantologists in 2011, which is the highest honor showing efforts in education, research and actual clinical experience with dental implants.

### Contents

#### **Preface XI**


Chapter 9 **Miniscrew Applications in Orthodontics 211** Fatma Deniz Uzuner and Belma Işık Aslan

#### Chapter 10 **Drug Delivery Systems in Bone Regeneration and Implant Dentistry 239** Sukumaran Anil, Asala F. Al-Sulaimani, Ansar E. Beeran, Elna P.

Chalisserry, Harikrishna P.R. Varma and Mohammad D. Al Amri

### Preface

Chapter 9 **Miniscrew Applications in Orthodontics 211** Fatma Deniz Uzuner and Belma Işık Aslan

**Dentistry 239**

**VI** Contents

Chapter 10 **Drug Delivery Systems in Bone Regeneration and Implant**

Sukumaran Anil, Asala F. Al-Sulaimani, Ansar E. Beeran, Elna P. Chalisserry, Harikrishna P.R. Varma and Mohammad D. Al Amri

> Implant dentistry has changed and enhanced significantly since the introduction of osseoin‐ tegration concept with dental implants. Because the benefits of therapy became apparent, implant treatment earned a widespread acceptance. Therefore, the need for dental implants has caused a rapid expansion of the market worldwide. Nowadays, general dentists and a variety of specialists provide implants to replace partial and complete edentulism.

> Dental implantology continues to excel with the developments of new surgical and prostho‐ dontic techniques, and armamentarium. The purpose of this book named *"Current Concepts in Dental Implantology"* is to present a novel resource for dentists who want to replace miss‐ ing teeth with dental implants. It is a carefully organized book, which blends basic science, clinical experience, and current and future concepts.

> This book includes ten chapters and our aim is to provide chapters that people from all over the world can easily understand, and advance the discipline of dental implantology. We contemplate that our book, *"Current Concepts in Dental Implantology"* , will be a valuable source for dental students, post-graduate residents and clinicians who want to know more about dental implants.

> In bringing this book to life, I sincerely owe my gratitude to many people. Firstly, I would like to thank all contributors in this book project, who worked hard in creating their chap‐ ters and getting them to me in the allocated time. Secondly, I would like to thank InTech Publisher for believing in the value of this book.

> I dedicate this book to my mother, *Servet* , father, *Ilhan* , and sister, *Ezgi* for their tremendous love and support in all my life.

> > **Ilser Turkyilmaz, DDS, PhD** Assistant Professor, Director, Dental School Implant Clinic, Department of Comprehensive Dentistry, The University of Texas Health Science Center at San Antonio, Texas, USA

### **Chapter 1**

### **Rationale for Dental Implants**

Ilser Turkyilmaz and Gokce Soganci

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/59815

#### **1. Introduction**

The loss of just one tooth will eventually have a global impact on the entire stomatognathic system. Bone loss, shifting of teeth, occlusal changes, decreased bite force and many more effects are felt throughout the entire system [1-3]. In attempt to prevent the progression of these effects, dentistry has continually searched for the ideal tooth replacement. With the advent of dental implants, clinicians can now restore patients higher levels of health and function than ever before [4-8].

The deleterious effects of tooth loss have been well know for centuries. As early as 600AD we have evidence of early Honduran civilizations attempting to implant seashells as replacements of a missing tooth and root complex [9]. As an alternative to replacing the entire tooth complex, the profession of dentistry has also created innovations targeted at replacing just the coronal aspect of the deficient site. An example of this would be the classic three unit fixed dental prosthesis to replace an extracted maxillary molar. This modality of treatment presents many attractive features. The time involved to restore only the coronal deficiency is minimal, often times being accomplished in as little as one hour. Commonly, this will involve alteration of existing, and sometimes virgin, teeth to support a tooth borne, fixed dental prostheses. The unfortunate side effect of this treatment lies in the eventual development of future complica‐ tions on those abutment teeth. [10]. Whether it be recurrent decay, material failure, or a different ailment, at some point the prosthesis will start to breakdown and the next restoration will be more invasive, costly and time consuming to both the patient and the practitioner [11]. More importantly, entire system will still experience negative effects because the root was never replaced. Both hard and soft tissues underneath the pontic site are still subjected to the cycle of breakdown as if the tooth was never replaced. Even with known future flaws in this design, the speed and affordability of these restorations have kept them as a popular method to replace missing teeth.

© 2015 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

The patient driven treatment plan classically places emphasis on speed of restoration and direct cost to the consumer. Until recently, implant dentistry has performed poorly in those two categories when compared to tooth borne restorations. Continued development in both macroscopic and microscopic elements in implant design have ushered in the era of speedier implant treatment. Traditional dental implant protocols were known to prescribe long time periods of healing. Patient and doctor demand have recently placed a high value to shortening the time period involved in implant dentistry. From dual stage, to single stage to immediate loading, the trend is consistent in shortening the treatment times to allow for immediate results [12-14]. Further, the increase in the number of companies in the industry, and improved methods of manufacturing have helped keep the cost of implant treatment attainable to the vast majority of patients. Contemporary implant dentistry has not only started to rival classic tooth borne care, but it is becoming the clear choice for tooth replacement. This has caused the number of implants being sold and surgically placed to grow exponentially [15]. With the advent of immediate placement and loading, this industry is poised to command the lion's share of the tooth replacement market as it will be satisfying all the demands of both the patients and practitioners with regard to speed, cost and healthy replacement of the all the missing components in the system.

Historically, there have been many valuable contributions from clinicians that helped implant dentistry evolve. Implant dentistry main consistent feature has been constant evolution in design, materials and protocols. The list of contributors is a different topic of discussion than what is targeted in this book. However Dr. Per-Ingvar Branemark is deserving of special attention.

In the 1950's Dr. Branemark was involved with in-vivo blood flow experiments on rabbits [16]. Initially, titanium chambers were being embedded in the ears of rabbits to record data for their investigations. When Dr. Branemark moved those chambers into the femurs of rabbits he later discovered he could not remove the chambers from the bone into which he had placed the chambers. He found the bone to have grown around the chambers and thus integrated to the titanium surface. Following this discovery, Dr. Branemark performed additional studies that verified the phenomenon of osseointegration [17]. His collaborative efforts verified pure titanium to be the material of choice. His efforts from that point on were largely targeted to the development of dental implants and improving the quality of life in the edentulous population or those suffering from maxillofacial defects [17,18].

#### **2. Tooth loss and edentulism**

Although the profession of dentistry is developing osteoinductive, osseoconductive and regenerative products. The native alveolar bone is still the ideal support apparatus for teeth and dental implants. The lack of osseous stimulation from the tooth complex results in bone loss. This loss is manifested in both density and volume. Once the tooth and periodontal ligament are no longer in place, the body initiates changes to remove the alveolar bony support it had once provided. Osteoclastic activity increases and the alveolar bone is eroded away. If the loss of a tooth is followed by placement of a dental implant, the loss of hard and soft tissue in the patient will be greatly reduced. If this process is not intercepted in a timely manner there will be a number of negative consequences dealt to the patient. Severe resorption of the bony processes harms both the quality of life and the quality of dental restorations that are able to be offered to the patients. The patient that waits to replace their teeth will often be informed that extensive grafting is needed to support dental implants. This results in increased cost and complexity. Whereas patients that are proactive in the transition from the dentate to edentate phases afford the clinician a better scenario to design for optimum results. Procedures such as "All on 4" have been designed to take an unhealthy, failing dentition to a healthy and fully restored state in as little as one day [19-21].

#### **3. Expansion of the market**

The patient driven treatment plan classically places emphasis on speed of restoration and direct cost to the consumer. Until recently, implant dentistry has performed poorly in those two categories when compared to tooth borne restorations. Continued development in both macroscopic and microscopic elements in implant design have ushered in the era of speedier implant treatment. Traditional dental implant protocols were known to prescribe long time periods of healing. Patient and doctor demand have recently placed a high value to shortening the time period involved in implant dentistry. From dual stage, to single stage to immediate loading, the trend is consistent in shortening the treatment times to allow for immediate results [12-14]. Further, the increase in the number of companies in the industry, and improved methods of manufacturing have helped keep the cost of implant treatment attainable to the vast majority of patients. Contemporary implant dentistry has not only started to rival classic tooth borne care, but it is becoming the clear choice for tooth replacement. This has caused the number of implants being sold and surgically placed to grow exponentially [15]. With the advent of immediate placement and loading, this industry is poised to command the lion's share of the tooth replacement market as it will be satisfying all the demands of both the patients and practitioners with regard to speed, cost and healthy replacement of the all the

Historically, there have been many valuable contributions from clinicians that helped implant dentistry evolve. Implant dentistry main consistent feature has been constant evolution in design, materials and protocols. The list of contributors is a different topic of discussion than what is targeted in this book. However Dr. Per-Ingvar Branemark is deserving of special

In the 1950's Dr. Branemark was involved with in-vivo blood flow experiments on rabbits [16]. Initially, titanium chambers were being embedded in the ears of rabbits to record data for their investigations. When Dr. Branemark moved those chambers into the femurs of rabbits he later discovered he could not remove the chambers from the bone into which he had placed the chambers. He found the bone to have grown around the chambers and thus integrated to the titanium surface. Following this discovery, Dr. Branemark performed additional studies that verified the phenomenon of osseointegration [17]. His collaborative efforts verified pure titanium to be the material of choice. His efforts from that point on were largely targeted to the development of dental implants and improving the quality of life in the edentulous

Although the profession of dentistry is developing osteoinductive, osseoconductive and regenerative products. The native alveolar bone is still the ideal support apparatus for teeth and dental implants. The lack of osseous stimulation from the tooth complex results in bone loss. This loss is manifested in both density and volume. Once the tooth and periodontal ligament are no longer in place, the body initiates changes to remove the alveolar bony support it had once provided. Osteoclastic activity increases and the alveolar bone is eroded away. If

population or those suffering from maxillofacial defects [17,18].

missing components in the system.

2 Current Concepts in Dental Implantology

**2. Tooth loss and edentulism**

attention.

Currently, global populations are living longer. At the time when Dr Branemark discovered osseointegration, the worlds life expectancy was 52. Currently, the life expectancy worldwide is 69.2 years. As our populations continue to live longer, there will be an increased demand on the dental profession's ability to both maintain oral health and effectively treat the eden‐ tulous population. Although there is speculation that the edentulous rate is dropping, the increased number of people entering the elderly population counters that number to yield an increase in the number of patients entering edentulism [15]. In fact, the total number of edentulous arches will climb to 37.9 million by the year 2020. This translates into a rise in the number of patients requiring at least one full arch of tooth replacement. Current evidence suggest that the restoration of the edentulous mandible with a conventional denture is no longer the most appropriate first choice of prosthodontic treatment [22,23].

While this demographic evolution may place strain on the worlds medical model, it serves as an ideal situation for the dental practitioner. Opportunistic clinicians are recognizing this trend and learning the skills to provide the great services that can be offered using dental implants.

Modern society has placed a high value on appearances. In the midst of an economic recession in 2009 the United States of America's population spent 10.5 billion dollars on cosmetic surgery. Patients exert a demand upon the dental practitioner to provide esthetics and function. The days of patients succumbing to edentulism and alteration of lifestyle are over. Through various forms of marketing, the modern population is aware of our ability to restore lost function and esthetics. The global market for dental implants is currently 3.4 billion dollars, with expected growth in the coming years.

Contemporary dental practices are in an ideal position to provide implant dentistry to patients. Through marketing and patient to patient interactions, the public is becoming aware of what implant dentistry can provide to the world. Improvements in surgical protocols and implant designs have enabled the clinician to immediately restore missing pieces of the stomatognathic system. However, it is up to the clinician to take the time and learn the techniques and protocols if they wish to capitalize on this market.

#### **4. Surface technologies**

Through the initial experiments of Dr. Branemark and coworkers in 1977 [17] and recent researchers [24-26], the dental profession adopted commercially pure grade 4 (high oxygen content) titanium as the material of choice for the implant body. Recently, the alloy form of Ti-6Al-4V has also been adopted into the dental implant industry to improve strength, corrosion resistance and density [27-29]. While the use of an alloy gives added strength to an implant, the lower grade titanium will give an increased osseointegration. Research by Johansson and coworkers showed only slight differences in removal torque values after periods of healing when placing implants of various grades in rabbits [30]. These authors concluded that the level of integration was sufficient in the alloy group and an argument can be made to use the alloys which give improved strength characteristics. Current dental research has allowed for further modifications to both microscopic and macroscopic aspects of dental implants that have improved success rates and healing times.

Surgical integration in combination with healing and loading dynamics are the main factors of whether or not an implant is integrated successfully. The general purpose of surface technologies is targeted to specific goals. Increasing bioacceptance, speeding up the healing of the surgical site and osseiointegration of the implant. Previous improvements on the micron level have been helpful, but the control of tissue response at the nano technological level is the current goal of researchers [31-33]. The implant itself will fall into one or a combination of the three possible categories. Metal, ceramic, or polymer are the three broad chemical classifica‐ tions of the materials.

Metals have enjoyed a long successful history in various areas of medical and dental implant practice. Biomechanical properties and suitability to sterilization are two advantages to this type of material. One must always remember that when the implant, abutment, or connecting screw are of dissimilar chemical composition, the risk of galvanic interactions exists [34-36]. Further, a galvanic reaction can yield corrosion, oxidation and even the production of pain in the host. This sort of complication is rarely reported, but the whenever we use dissimilar metals in our treatment plans we should be aware of this potential.

Ceramics can be seen as the entire implant or as a surface modification to the metal implant body. Common forms of coatings are hydroxyapatite, tricalcium phosphate or a form of bioglass [37-39]. The possibility of surface degradation, especially with hydroxyapatite, has been an area of contention with many pointing to this element when adverse implant to bone interactions occur.

Polymers were once thought to have advantageous qualities to be incorporated into implant design. Specifically, the shock absorbing capability was once thought to counteract the lack of periodontal ligaments with regards to occlusion. However, research and clinical reports have shown this material to be inferior to those previously discussed and is seldom incorporated today.

Surfaces are generally going to be further classified by the biodynamic response they illicit from the body [40]. No material is completely accepted by the body, but to optimize the implant's performance emphasis is placed on minimizing biologic response while allowing adequate function. Bioinert, bioactive or biotolerant are the current terms used in this area of investigation [41,42]. All three of these descriptive adjectives imply biocompatibility to the host.

A biotolerant material is one that is not rejected by the host but, rather is surrounded by a fibrous layer. Bioinert materials are described as allowing close apposition of bone to the surface, lending itself to contact osteogenesis. Bioactive refers to allowing formation of new bone onto the surface but ion exchange with host tissue leading to formation of chemical bonds along the interface.

When the implant is inserted into the osteotomy site it will have an effect on the bone and blood clot that it is in intimate contact with. Osseoconductive and osseoinductive are common terms to describe the body's response to dental materials. Bioinert and bioactive materials are grouped into the osseoconductive category [41,42]. This refers to the ability to act as a scaffold, or allowing bone formation on their surface. Osseoinductive refers to a materials ability to induce bone formation de novo. An example of this is seen in recombinant human bone morphogenetic protein 2 [43,44].

A number of microscopic surface coating changes have been shown to provide improved healing to the implant surface. Generally, surface coatings are sprayed onto the implant. One must realize that surface coatings rely on adhesive qualities to remain on the implant during insertion. Bond strengths are currently reported to be in the range of 15-30 MPa. This low strength brings into question how practical a surface coating may be in the clinical environ‐ ment. Speculation exists whether or not the coating is maintained during the placement of the implant into a osteotomy. However, many manufacturers are using this technology on their implants which suggests positive feedback from the clinical results.

Turned surfaces, sandblasted, plasma sprayed, acid etched, anodized, HA, zirconia, and more have been heavily advertised as additions to the to pure titanium body. This list will continue to grow as implant companies position themselves to achieve faster healing times and thus allow for immediate loading. The common theme advertised from all the manufacturers is increasing bone to implant contact in both volume and speed. Examples of popular surfaces will be discussed. Currently there is over 80 companies producing over 250 different types of dental implants. Caution is recommended to the dentist with regards to this aspect of implant dentistry. As this field is rapidly changing. It is up to the clinician to use professional judgement on whether or not to adopt a new surface into their implant practice. Food and Drug Admin‐ istration (FDA) clearance is often a good sign of whether or not a manufacturer's claim has undergone any actual scientific investigation.

#### **4.1. Microscopic topography**

**4. Surface technologies**

4 Current Concepts in Dental Implantology

tions of the materials.

interactions occur.

today.

Through the initial experiments of Dr. Branemark and coworkers in 1977 [17] and recent researchers [24-26], the dental profession adopted commercially pure grade 4 (high oxygen content) titanium as the material of choice for the implant body. Recently, the alloy form of Ti-6Al-4V has also been adopted into the dental implant industry to improve strength, corrosion resistance and density [27-29]. While the use of an alloy gives added strength to an implant, the lower grade titanium will give an increased osseointegration. Research by Johansson and coworkers showed only slight differences in removal torque values after periods of healing when placing implants of various grades in rabbits [30]. These authors concluded that the level of integration was sufficient in the alloy group and an argument can be made to use the alloys which give improved strength characteristics. Current dental research has allowed for further modifications to both microscopic and macroscopic aspects

Surgical integration in combination with healing and loading dynamics are the main factors of whether or not an implant is integrated successfully. The general purpose of surface technologies is targeted to specific goals. Increasing bioacceptance, speeding up the healing of the surgical site and osseiointegration of the implant. Previous improvements on the micron level have been helpful, but the control of tissue response at the nano technological level is the current goal of researchers [31-33]. The implant itself will fall into one or a combination of the three possible categories. Metal, ceramic, or polymer are the three broad chemical classifica‐

Metals have enjoyed a long successful history in various areas of medical and dental implant practice. Biomechanical properties and suitability to sterilization are two advantages to this type of material. One must always remember that when the implant, abutment, or connecting screw are of dissimilar chemical composition, the risk of galvanic interactions exists [34-36]. Further, a galvanic reaction can yield corrosion, oxidation and even the production of pain in the host. This sort of complication is rarely reported, but the whenever we use dissimilar metals

Ceramics can be seen as the entire implant or as a surface modification to the metal implant body. Common forms of coatings are hydroxyapatite, tricalcium phosphate or a form of bioglass [37-39]. The possibility of surface degradation, especially with hydroxyapatite, has been an area of contention with many pointing to this element when adverse implant to bone

Polymers were once thought to have advantageous qualities to be incorporated into implant design. Specifically, the shock absorbing capability was once thought to counteract the lack of periodontal ligaments with regards to occlusion. However, research and clinical reports have shown this material to be inferior to those previously discussed and is seldom incorporated

Surfaces are generally going to be further classified by the biodynamic response they illicit from the body [40]. No material is completely accepted by the body, but to optimize the

of dental implants that have improved success rates and healing times.

in our treatment plans we should be aware of this potential.

Currently, most all manufacturers have made the shift from smooth implant surfaces to a rough surface [5,24,45,46]. Recently, even the smooth collar model that was promoted for increased hygiene has seen reduced promotion and use. This signals that most contemporary research points to rough surfaces functioning better in the role of promoting the mechanical interlocking of the surrounding tissues. On the microscopic level of this element lies cell differentiation responses to different microscopic topographies. The appositional response of the extracellular matrices in the bone to implant environment have shown potential for providing improvement in implant performance [47,48]. Similar to the computer industry, the major advances in this area are found in nanotechnological engineering [32,33]. The word nano-lithography may be the next buzz word in advertisements from implant manufacturers. As a profession we will get there, but as technologically advanced as this sounds, the reality of current manufacturing is a surface is being textured by some sort of grit blasting process.

TiUnite is the current surface advertised by NobelBiocare. This adds an osseoconductive element to implants manufactured by NobelBiocare. It is a highly crystalline and phosphate enriched titanium oxide characterized by a micro structured surface with open pores in the low micrometer range. The surface is generated by spark anodization and consists of titanium oxide [49,50]. The following ptohos show an implant with TiUnite surface, and scanning electron microscopic (SEM) images of TiUnite surface during osseointegration (Figures 1-4).

**Figure 1.** NobelReplace Straight Groovy Implant with TiUnite surface.

Strauman currently promotes a surface by the name of SLActive [51-53]. This title denotes how the implant is conditioned for optimizations. **S**andblasting with **L**arge grit followed by **A**cid etching is how the manufacturer achieves the surface topography. To create the 'active' surface, the implant is conditioned with nitrogen and preserved in an isotonic saline solution.

Astratech dental implants are currently promoting a TiOblast and Osseospeed surface [54,55]. Essentially the surface of the implant is grit blasted with titanium dioxide particles to achieve

**Figure 2.** SEM view of TiUnite surface during osteoblasts are attaching on it.

interlocking of the surrounding tissues. On the microscopic level of this element lies cell differentiation responses to different microscopic topographies. The appositional response of the extracellular matrices in the bone to implant environment have shown potential for providing improvement in implant performance [47,48]. Similar to the computer industry, the major advances in this area are found in nanotechnological engineering [32,33]. The word nano-lithography may be the next buzz word in advertisements from implant manufacturers. As a profession we will get there, but as technologically advanced as this sounds, the reality of current manufacturing is a surface is being textured by some sort of grit blasting process.

6 Current Concepts in Dental Implantology

TiUnite is the current surface advertised by NobelBiocare. This adds an osseoconductive element to implants manufactured by NobelBiocare. It is a highly crystalline and phosphate enriched titanium oxide characterized by a micro structured surface with open pores in the low micrometer range. The surface is generated by spark anodization and consists of titanium oxide [49,50]. The following ptohos show an implant with TiUnite surface, and scanning electron microscopic (SEM) images of TiUnite surface during osseointegration (Figures 1-4).

Strauman currently promotes a surface by the name of SLActive [51-53]. This title denotes how the implant is conditioned for optimizations. **S**andblasting with **L**arge grit followed by **A**cid etching is how the manufacturer achieves the surface topography. To create the 'active' surface,

Astratech dental implants are currently promoting a TiOblast and Osseospeed surface [54,55]. Essentially the surface of the implant is grit blasted with titanium dioxide particles to achieve

the implant is conditioned with nitrogen and preserved in an isotonic saline solution.

**Figure 1.** NobelReplace Straight Groovy Implant with TiUnite surface.

**Figure 3.** SEM view of TiUnite surface when osteoblasts have filled pores on the implant surface.

an isotropic, moderately roughened surface. Later the implant is chemically conditioned with fluoride to gain slight topographical changes.

Zimmer contemporary surface is called MTX [56,57]. This acronym denotes 'micro-texturing' the implant surface. The implant is Grit blasted with hydroxyapatite particles and then conditioned a non-etching environment to remove residual blasting material.

3i, or implant innovations Inc, uses surface technology termed nano-tite [58,59]. After microtexturing like the companies previously listed, the implant is then conditioned into a calcium phosphate solution.

**Figure 4.** Another SEM view of TiUnite surface when osteoblasts have filled pores.

If all five of these surfaces from the five major manufacturers are compared, not much difference exists. Currently, a textured surface is created which is then followed by some element of conditioning thought to improve bioactivity.

An additional step to spraying on coatings or roughening the surface of implants is seen in the chemical treatment of the implant surface. The overriding goal in this treatment modality is to improve the wettability of the implant surface itself, or otherwise, to make the implant surface more hydrophilic [60]. Clinicians are advised that the contact angle of pre-existing surfaces was never poor and may be sufficient without additional modification. Early experi‐ ments have been promising in showing improvements in this area. However, it is not know to what extent this actually plays in implant success.

It is impossible to predict what the next big thing in implant dentistry will be. In fact, dentistry as a profession is changing so rapidly, it is a challenge for the practicing clinician to remain current with what the research world can produce. An over-riding principle must always be to be critical of what is advertised.

#### **5. Macroscopic design**

Dental implants have assumed a variety of shapes through the years. From frames to baskets and cylinders to tapered screw threaded forms, the macroscopic design has seen numerous functional advances. Currently the threaded implant body enjoys the majority share of the market. Experiments have shown that screw type implants maintain a higher bone to implant contact through years of function [9]. With this body shape dominating the market, a discus‐ sion in the elements of the screw shape is deserving.

There are four basic types of threads seen in a screw shape [9]. V-thread, buttress thread, reverse buttress thread and square threads. All of these designs will exert different forces on the surrounding tissues when subjected to various load patterns. The thread pitch denotes the distance between adjacent threads. Thread depth will refer to the distance between the major and minor diameter of the implant. In addition to load distribution, the geometry of the thread will impact the surgical behavior of the implant during placement.

Implant bodies are commonly available as tapered or parallel. With regard to immediate load, clinicians are often looking for immediate stability. The tapered design imparts the ability to place an implant into an underprepared osteotomy site resulting in higher insertional torque values [61]. Controversy exists over what is the maximum torque that results in negative effects on the supporting tissues. Modern implants are being designed to withstand the high torquing forces on the implant body itself. However, some argue these high forces placed on the surrounding bone have the ability to cause compression necrosis. This is a current point of contention amongst various researchers. With regards to implant length, some manufacturers/ researchers are promoting the use of shorter length implants [62,63]. However, caution is advised to the clinicians in this area, especially for implants shorter than 10mm.

When considering force distribution in the final prosthesis it is imperative to consider both biomechanics and limitations of biology [64]. By using longer and wider implants, the surface area of a load is increased. This in-turn lowers the force on the overall system (F=M/A). In contrast, if a wide platform implant is chosen for a given osteotomy site, one must be careful not to exceed the biologic parameters of the patient. For example, if a wide implant results in insufficient buccal bone, the gain in force distribution will be negated by the decrease of vascularity to the buccal bone in that site and potential implant complications.

If all five of these surfaces from the five major manufacturers are compared, not much difference exists. Currently, a textured surface is created which is then followed by some

An additional step to spraying on coatings or roughening the surface of implants is seen in the chemical treatment of the implant surface. The overriding goal in this treatment modality is to improve the wettability of the implant surface itself, or otherwise, to make the implant surface more hydrophilic [60]. Clinicians are advised that the contact angle of pre-existing surfaces was never poor and may be sufficient without additional modification. Early experi‐ ments have been promising in showing improvements in this area. However, it is not know

It is impossible to predict what the next big thing in implant dentistry will be. In fact, dentistry as a profession is changing so rapidly, it is a challenge for the practicing clinician to remain current with what the research world can produce. An over-riding principle must always be

Dental implants have assumed a variety of shapes through the years. From frames to baskets and cylinders to tapered screw threaded forms, the macroscopic design has seen numerous functional advances. Currently the threaded implant body enjoys the majority share of the market. Experiments have shown that screw type implants maintain a higher bone to implant contact through years of function [9]. With this body shape dominating the market, a discus‐

There are four basic types of threads seen in a screw shape [9]. V-thread, buttress thread, reverse buttress thread and square threads. All of these designs will exert different forces on

element of conditioning thought to improve bioactivity.

**Figure 4.** Another SEM view of TiUnite surface when osteoblasts have filled pores.

to what extent this actually plays in implant success.

sion in the elements of the screw shape is deserving.

to be critical of what is advertised.

8 Current Concepts in Dental Implantology

**5. Macroscopic design**

The design of the implant to abutment connection is another aspect of treatment that the clinician must decide upon prior to treatment [65,66]. Whether to use external or internal hex, trilobe, conical, morse taper, platform switching are all decisions that must be made by the dentist (Figures 5,6). As in other areas of dentistry, there is a blend of art and science. Some clinicians use what works best in their hands or make decisions based on feel. Hopefully, as evidence based dentistry matures and actually starts to produce tangible recommendations, the decision making tree will become more research based.

**Figure 5.** Immediate implant placement with NobelReplace implants with internal trilobe connection.

**Figure 6.** Cast with Branemark implant replicas with external hex connection.

In 1982, Dr. Gerald Niznick introduced the internal hex connection [67]. The purpose of this design was to create an implant to abutment connection that shifted the force from the implant screw to the platform connection. Prior to this innovation, screw fractures were a common complication [10]. Numerous studies have shown this to be a structural improvement with regards to reducing the stress placed upon the abutment screw [68,69]. Most clinicians prescribe implants with internal connections. Even with this said, the external hex remains a viable option and is still used by many dentists. After the decision of making your connection internal or external (Figures 7-9), the next choice is whether or not to use a platform shift.

**Figure 7.** Engaging and non-engaging UCLA abutments for NobelBiocare Replace implants with internal trilobe con‐ nection.

The term platform shift refers to a mismatched fit of the implant platform and that of the abutment. In the late 1980s the benefits of platform switching was unforeseen by the practi‐ tioners using this mismatched design. Wide diameter implants did not have a matching platform for the abutments so a regular platform was used. Upon follow up examination, the crestal bone levels were thought to be equal or better than platform matched connections [70, 71]. Current research suggests the medial movement of the implant / abutment junction is beneficial in reducing crestal bone loss. The marginal gap is thought to exert a sphere of influence on the biological reaction from the bone and soft tissues. A mismatched connection of.4 mm or greater appears to result in statistically significant less bone loss [72].

**Figure 8.** Engaging and non-engaging UCLA abutments for Zimmer implants with internal hex connection.

**Figure 6.** Cast with Branemark implant replicas with external hex connection.

10 Current Concepts in Dental Implantology

nection.

In 1982, Dr. Gerald Niznick introduced the internal hex connection [67]. The purpose of this design was to create an implant to abutment connection that shifted the force from the implant screw to the platform connection. Prior to this innovation, screw fractures were a common complication [10]. Numerous studies have shown this to be a structural improvement with regards to reducing the stress placed upon the abutment screw [68,69]. Most clinicians prescribe implants with internal connections. Even with this said, the external hex remains a viable option and is still used by many dentists. After the decision of making your connection internal or external (Figures 7-9), the next choice is whether or not to use a platform shift.

**Figure 7.** Engaging and non-engaging UCLA abutments for NobelBiocare Replace implants with internal trilobe con‐

The term platform shift refers to a mismatched fit of the implant platform and that of the abutment. In the late 1980s the benefits of platform switching was unforeseen by the practi‐ tioners using this mismatched design. Wide diameter implants did not have a matching platform for the abutments so a regular platform was used. Upon follow up examination, the crestal bone levels were thought to be equal or better than platform matched connections [70,

**Figure 9.** Engaging and non-engaging UCLA abutments for implants with external hex connection.

In implant dentistry, current paradigms for treatment success are based not only on true clinical outcomes such as implant survival, restoration survival, and patient satisfaction but also on surrogate clinical outcomes such as dentogingival esthetics and health of surrounding soft tissues [73]. This is especially important for implant therapy in maxillary and mandibular anterior regions, where esthetics play a predominant role in treatment success. A variety of abutments, and restorations differing in design and biomaterials have been introduced to achieve optimal mechanical, biological, and esthetic treatment outcomes. As an abutment material, traditionally titanium is selected due to its mechanical properties. However, the color of underlying titanium abutments negatively affected the appearance of peri-implant mucosa. To provide more predictable results regarding esthetic aspects, all-ceramic abutments made out of alumina and zirconia were introduced about 10 years ago. In vitro and in vivo studies [74,75] demonstrated superior fracture resistance of zirconia abutments with esthetic outcomes (Figures 10-13).

**Figure 10.** Implant is placed to restore maxillary left lateral tooth.

**Figure 11.** Zirconia abutment is screwed on the implant.

**Figure 12.** All-ceramic crown is cemented on zirconia abutment.

**Figure 13.** Translucency of natural teeth and all-ceramic restorations is similar, giving more esthetic outcomes.

#### **6. Risk factors for implant candidates**

**Figure 10.** Implant is placed to restore maxillary left lateral tooth.

12 Current Concepts in Dental Implantology

**Figure 11.** Zirconia abutment is screwed on the implant.

**Figure 12.** All-ceramic crown is cemented on zirconia abutment.

Many attempts to use the phrase contraindications to dental implants have been made [76]. However those lists are often subject to controversy as the severity of a disease or patient condition exists on a sliding scale. For example, one diabetic patient may be at a higher risk than another [77,78]. Or one could ask, at what point does tobacco smoking effect implant survival? Case reports may exist for complications related to various patient conditions, but the doctor is reminded that those reports fall very low on the scale of strength of evidence the clinician can use and apply to their patient pool. Recently, the focus has fallen away from indications and contraindications and more emphasis is placed on risk factors. Risk factors are characteristics statistically associated with, although not necessarily causally related to, an increased risk of morbidity or mortality.

Multiple consensus review groups have recommended that risk factors be divided into two groups [76,79]. Systemic factors and local factors are the groups usually recommended and the latter is further subdivided into very high risk and significant risk. A noteworthy statement that resulted from these reports was in regard to the many attempts from other authors to create a list of relative and absolute contraindications to dental implant placement. This idea is discredited by this group because for many topics weak evidence exists in placing different conditions into an absolute contraindication. A case report of limited sample size is simply not enough evidence to create an absolute contraindication.

The chapter classified very high risk patients as those who could be attributed to having serious systemic disease, immunocompromised health status, drug abusers, and non-compliant patients [76]. A systemic disease can interfere with dental implant therapy at the level of local healing by altering tissue responses to implant placement and surgical treatments. Further, the medications that a patient may be taking for the systemic disease can interfere with normal cellular functions and thereby affect healing and osseointegration. The American Society of Anesthesiology (ASA) has a well known publication to help classify a patient's risk to anesthesia leading into a surgical procedure [76]. Although many dental implants are not placed under general anesthesia, this classification system is an effective way to gauge the patients status for receiving any surgical treatment. For patients that fall into categories, dental treatment is not generally recommended until the patients health status improves and they are placed in a lower category. Significant risk patients were those who had prior irradiation, severe diabetes, bleeding disorders and/ or heavy smoking habits. Local factors are of partic‐ ular concern with regard to implant survival. Some often highlighted factors are interdental / interimplant space, infected implant sites, soft tissue thickness, width of keratinized soft tissue, bone density, bone volume and implant stability.

In the era of immediate loading of dental implants, initial stability is of primary concern [46,80,81]. Reports have concluded through clinical research that initial stability is related to success with implant survival. There are a number of ways to measure the initial stability of an implant. The most common method of that is to measure the insertion torque of the implant during the final stage of placement using a torque wrench. Resonance frequency has recently been examined and verified to provide useful intrapatient information [80] (Figures 14, 15). Specifically, values of multiple implants in the same patient are useful gauges on implant stability throughout the life of the implant. A correlation between preoperative CBCT scans and resonance frequency values at the time of placement has shown that primary implant stability may be able to be calculated preoperatively [81,82].

**Figure 14.** Osstell instrument used to determine implant stability.

**Figure 15.** Transducer of Osstell instrument attached to implant for measurement.

cellular functions and thereby affect healing and osseointegration. The American Society of Anesthesiology (ASA) has a well known publication to help classify a patient's risk to anesthesia leading into a surgical procedure [76]. Although many dental implants are not placed under general anesthesia, this classification system is an effective way to gauge the patients status for receiving any surgical treatment. For patients that fall into categories, dental treatment is not generally recommended until the patients health status improves and they are placed in a lower category. Significant risk patients were those who had prior irradiation, severe diabetes, bleeding disorders and/ or heavy smoking habits. Local factors are of partic‐ ular concern with regard to implant survival. Some often highlighted factors are interdental / interimplant space, infected implant sites, soft tissue thickness, width of keratinized soft tissue,

In the era of immediate loading of dental implants, initial stability is of primary concern [46,80,81]. Reports have concluded through clinical research that initial stability is related to success with implant survival. There are a number of ways to measure the initial stability of an implant. The most common method of that is to measure the insertion torque of the implant during the final stage of placement using a torque wrench. Resonance frequency has recently been examined and verified to provide useful intrapatient information [80] (Figures 14, 15). Specifically, values of multiple implants in the same patient are useful gauges on implant stability throughout the life of the implant. A correlation between preoperative CBCT scans and resonance frequency values at the time of placement has shown that primary implant

bone density, bone volume and implant stability.

14 Current Concepts in Dental Implantology

stability may be able to be calculated preoperatively [81,82].

**Figure 14.** Osstell instrument used to determine implant stability.

It is questioned if an adequate amount of interdental space needs to exist between an implant and an adjacent tooth [83,84]. Studies have shown that interdental spaces of less than 3mm were associated with increased bone loss around the implants. In this particular study, cases where this space was compromised seemed to especially result in bone loss around maxillary lateral incisors.

An infected tooth site is generally defined as one that exhibits signs or symptoms of pain, periapical radiolucency, fistula, suppuration, or a combination of these. The clinical scenario whereby an infected tooth is to be extracted and subsequently followed by implant placement in that site is commonplace in many practices. Whether or not placement of an implant in that site immediately is a key decision the dentist must face. Several clinical reports have been published on this topic, all with varying degrees of success [85,86]. However, studies like that of Villa have shown success in the placement of dental implants and immediate loading into previously infected sites. This idea is relatively new to dentistry, but the preliminary results do appear promising.

The subject of bone density and volume is of particular concern to the implant clinician. While bone density is often a topic of discussion, there exists little data on the relationship of bone density and implant success. With regard to bone volume, it is generally accepted that there are critical parameters in bone volume to support the success of a dental implant. An implant must be surrounded by bone that has adequate vascularity. If the surrounding bone does not have adequate thickness and therefore compromised vascularity, the implant has a higher chance of experiencing both soft and hard tissue attachment loss.

In addition to local and systemic biologic factors previously listed, a patient having a positive history to periodontitis and/or use of smoking tobacco should be noted and considered by the implant clinician. Drs. Heitz-Mayfield and Huynh-Ba performed a comprehensive review of the literature on this subject in 2009 [87]. They found numerous studies have targeted at success rates in patients that fit this demographic of past periodontal disease and tobacco use. With regards to patients that had a history of treated periodontal disease they were able to identify patterns and make following useful conclusions;


When discussing the issue of a positive history to tobacco smoking they found results that enabled them to make the following conclusions;


The take away message from these reviews for the clinician should be; patients with a history of treated periodontitis and or smoking have an increased risk of implant failure and periimplantitis. However, neither of these risk factors are absolute contraindications to implant therapy.

#### **7. Conclusion**

Implant dentistry has come a long way since the discovery of osseointegration of dental implants. In the last 40 years, the use of dental implants has dramatically increased. Initially, very few specialists were trained in surgical placement and subsequent restoration. As the treatment became more predictable, the benefits of therapy became evident. The tremendous demand for implants has fueled a rapid expansion of the market. Presently, general dentists and multiple specialists offer implant treatments. The field is evolving and expanding, from surgical techniques to types of restorations available.

In this chapter, general information regarding the need for dental implants, implant types and designs, and possible risks factors for patients who are looking for implant treatments have been provided. In the following chapters, more detailed information about several topics will be covered.

#### **Author details**

rates in patients that fit this demographic of past periodontal disease and tobacco use. With regards to patients that had a history of treated periodontal disease they were able to identify

**a.** implant survival in patients with a histoy of treated periodontitis ranged from 59% to

**b.** the majority of studies reported high implant survival rates >90% for implants with turned

When discussing the issue of a positive history to tobacco smoking they found results that

**a.** Implant outcomes in 45 patients who were rehabilitated following an immediate loading protocol in the mandible were evaluated following 1 year of loading. The results showed there was no statistically significant difference in the smokers and non-smokers with

**c.** Overall there is limited data on the survival and success rates of implants in former

**d.** There are studies that show an increased risk of peri-implantitis for patients that smoke. The take away message from these reviews for the clinician should be; patients with a history of treated periodontitis and or smoking have an increased risk of implant failure and periimplantitis. However, neither of these risk factors are absolute contraindications to implant

Implant dentistry has come a long way since the discovery of osseointegration of dental implants. In the last 40 years, the use of dental implants has dramatically increased. Initially, very few specialists were trained in surgical placement and subsequent restoration. As the treatment became more predictable, the benefits of therapy became evident. The tremendous demand for implants has fueled a rapid expansion of the market. Presently, general dentists and multiple specialists offer implant treatments. The field is evolving and expanding, from

In this chapter, general information regarding the need for dental implants, implant types and designs, and possible risks factors for patients who are looking for implant treatments have been provided. In the following chapters, more detailed information about several topics will

**b.** The majority of studies showed implant survival rates in smokers of 80% to 96%.

patterns and make following useful conclusions;

enabled them to make the following conclusions;

regards to immediate loading protocol.

surgical techniques to types of restorations available.

**c.** all studies reported regular supportive periodontal therapy.

or moderately rough surfaces,

100%,

16 Current Concepts in Dental Implantology

smokers.

**7. Conclusion**

be covered.

therapy.

Ilser Turkyilmaz1\* and Gokce Soganci2

\*Address all correspondence to: ilserturkyilmaz@yahoo.com

1 Department of Comprehensive Dentistry, University of Texas Health Science Center at San Antonio, Texas, USA

2 Department of Prosthodontics, Oral and Dental Health Center, Ankara, Turkey

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### **Bone Substitute Materials in Implant Dentistry**

Sybele Saska, Larissa Souza Mendes, Ana Maria Minarelli Gaspar and Ticiana Sidorenko de Oliveira Capote

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/59487

#### **1. Introduction**

Although bone autografts have been routinely used as "gold standard" for reconstruction/ replacement bone defects, because they have osteogenic, osteoinductive, osteoconductive properties, they have a high number of viable cells and are rich in growth factors. However, the use of autograft is limited by several factors, being one of them the insufficient amount of donor tissue. Therefore, bone substitute materials have been extensively studied in order to develop an ideal material for substitution of bone grafts, due to some disadvantages presented by autografts, allografts and xenografts, such as poor bone quality, an inadequate amount of bone and possible immunogenicity for allografts and xenografts, which limit the use of these grafts in specific surgical protocols. These disadvantages have led tissue engineering and biotechnology to develop new materials and promising methods for tissue repair, especially for bone tissue. Thus, bone substitutes, synthetic and/or biotechnologically processed have become potential materials for clinical applications in different areas of health.

An ideal bone substitute (BS) material should provide a variety of shapes and sizes with suitable mechanical properties to be used in sites where there are impact loading; moreover, these materials should be biocompatible, osteoconductive, preferably being resorbable and replaced by new bone formation. In general, resorbable BS materials are preferred, since these materials are expected to preserve the increased bone volume during the reconstruction and simultaneously are gradually replaced by newly formed bone.

Synthetic materials, denominated as alloplastics, may act as scaffolds for bone cells providing tissue growth inside the respective material.

© 2015 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

A scaffold must be highly porous with interconnected pores and have adequate mechanical properties. The surface of a scaffold should be similar to extracellular matrix (ECM). These properties enable the scaffold to act as a matrix for tissue regeneration to maintain and improve tissue/organs functions; therefore, it is considered the key element for the success in tissue engineering. Numerous physicochemical features of scaffolds, such as surface chemistry, surface roughness, topography, mechanical properties and interfacial free energy (hydropho‐ bic/hydrophilic balance) are important for cell attachment, proliferation and differentiation. These factors are also critically important to the overall biocompatibility and bioactivity of a particular material [1-3].

Resorption of a biomaterial is related to several factors, such as, particle size, porosity, chemical structure (composition and crystallinity), and pH of body fluids [4, 5]. Particles with nano‐ metric sizes are reabsorbed faster than micrometric particles, because osteoclasts or macro‐ phages act faster on a biomaterial surface. Biomaterial crystallinity also changes the resorption rate, since highly crystalline structures are more resistant to resorption than an amorphous or semi-crystalline structure. Moreover, the chemical composition is also important. Impurities such as calcium carbonate promote faster resorption [6]. The failure or the success of a material for bone fill or replacement may be related to the resorption rate of the material, as well as the regenerative capacity of bone tissue. This process can occur in three forms: 1. insufficient permanence of the material to promote bone apposition and to allow the osteoconductivity; 2. premature destabilization of newly formed bone due to the complete degradation of the material; 3. an exaggerated inflammatory response due to the degradation of the material [7]. Thus, bone substitute materials must have suitable resorption rate in accordance with the rate of tissue formation.

Despite recent advances in the development of new BS for bone tissue engineering, there is still a search for a material or a composite with mechanical properties and physicochemical characteristics similar to autograft and a structure closer to the natural ECM.

#### **2. Ceramic-based bone substitutes**

Ceramics are compounds between metallic and nonmetallic elements. Ceramic materials have a several of attractive advantages comparing to other materials. These include high melting points, great hardness, low densities and chemical and environmental stability. However, ceramics are severely affected by lack of toughness; they are extremely brittle, and are highly susceptible to fracture. They are most frequently oxides, nitrides, and carbides, for example, some of the common ceramic materials include aluminum oxide (or alumina, Al2O3) and silicon dioxide (or silica*,* SiO2), in addition, some traditional ceramics are referred as those composed of clay minerals (*i.e*., porcelain), as well as cement and glass [8].

The ability of ceramic materials to bond to the bone tissue is a unique property of bioactive ceramics. This property has been led their wide clinical application in both areas as orthopedics and dentistry. The use of ceramics for hard tissues reconstitution has been performed for centuries, but in clinical practice the use of these materials only began in the late eighteenth century with the use of porcelain for making dental prostheses. On the other hand, in ortho‐ pedics the use of ceramic materials happened in the late nineteenth century with the use of plaster of Paris (calcium sulfate hemihydrate, CaSO4 ½H2O) for bone defects filling [9].

A scaffold must be highly porous with interconnected pores and have adequate mechanical properties. The surface of a scaffold should be similar to extracellular matrix (ECM). These properties enable the scaffold to act as a matrix for tissue regeneration to maintain and improve tissue/organs functions; therefore, it is considered the key element for the success in tissue engineering. Numerous physicochemical features of scaffolds, such as surface chemistry, surface roughness, topography, mechanical properties and interfacial free energy (hydropho‐ bic/hydrophilic balance) are important for cell attachment, proliferation and differentiation. These factors are also critically important to the overall biocompatibility and bioactivity of a

Resorption of a biomaterial is related to several factors, such as, particle size, porosity, chemical structure (composition and crystallinity), and pH of body fluids [4, 5]. Particles with nano‐ metric sizes are reabsorbed faster than micrometric particles, because osteoclasts or macro‐ phages act faster on a biomaterial surface. Biomaterial crystallinity also changes the resorption rate, since highly crystalline structures are more resistant to resorption than an amorphous or semi-crystalline structure. Moreover, the chemical composition is also important. Impurities such as calcium carbonate promote faster resorption [6]. The failure or the success of a material for bone fill or replacement may be related to the resorption rate of the material, as well as the regenerative capacity of bone tissue. This process can occur in three forms: 1. insufficient permanence of the material to promote bone apposition and to allow the osteoconductivity; 2. premature destabilization of newly formed bone due to the complete degradation of the material; 3. an exaggerated inflammatory response due to the degradation of the material [7]. Thus, bone substitute materials must have suitable resorption rate in accordance with the rate

Despite recent advances in the development of new BS for bone tissue engineering, there is still a search for a material or a composite with mechanical properties and physicochemical

Ceramics are compounds between metallic and nonmetallic elements. Ceramic materials have a several of attractive advantages comparing to other materials. These include high melting points, great hardness, low densities and chemical and environmental stability. However, ceramics are severely affected by lack of toughness; they are extremely brittle, and are highly susceptible to fracture. They are most frequently oxides, nitrides, and carbides, for example, some of the common ceramic materials include aluminum oxide (or alumina, Al2O3) and silicon dioxide (or silica*,* SiO2), in addition, some traditional ceramics are referred as those composed

The ability of ceramic materials to bond to the bone tissue is a unique property of bioactive ceramics. This property has been led their wide clinical application in both areas as orthopedics and dentistry. The use of ceramics for hard tissues reconstitution has been performed for centuries, but in clinical practice the use of these materials only began in the late eighteenth

characteristics similar to autograft and a structure closer to the natural ECM.

of clay minerals (*i.e*., porcelain), as well as cement and glass [8].

particular material [1-3].

26 Current Concepts in Dental Implantology

of tissue formation.

**2. Ceramic-based bone substitutes**

The term "bioceramic" refers to biocompatible ceramic materials applied to biomedical and clinical use due to certain characteristics such as biocompatibility, excellent tribological properties and high chemical stability, which is superior to metals in different applications, moreover excellent osteoconductive [10].

Among bioceramics, calcium phosphates are ceramics with Ca/P molar ratio ranging from 0.5 to 2.0 and are found in different types [11], in which the best known form is hydroxyapatite (HA), a natural mineral component representing 30 to 70% of the mass of bones and teeth [12]. The chemical structure of biological HA is very complex, because it not presents a totally pure composition (non-stoichiometric), being frequently calcium-deficient hydroxyapatite enriched with carbonate ions forming the carbonate-apatite [13]. Some calcium phosphates of biological relevance are: amorphous calcium phosphate (ACP), dicalcium phosphate dihydrate (DCPD), dicalcium phosphate (DCP), octacalcium phosphate (OCP), tricalcium phosphate (TCP), calcium pyrophosphate (CPP) and hydroxyapatite (HA).

Pure HA, calcium hydroxyapatite specifically, is a stoichiometric composition of (Ca)10(PO4)6(OH)2 (Ca/P = 1.67). It is main inorganic component of bone tissue and teeth. For many years, different types of synthesis and applications of these calcium phosphates have been researched for regeneration/reconstruction of bone structures. Synthetic HA has been used for this purpose, because they are bioactive material and can have a Ca/P molar ratio less than 1.67; thus, they are more effective clinically due to its similarities with the composition of bone tissue and their osteoconductive properties [13, 14].

Bioceramics have different rates of *in vitro* solubility, which reflects in the *in vivo* degradation, *i.e*., as greater the Ca/P molar ratio lower is the solubility of bioceramics [15]. However, the rate of dissolution is not only influenced by Ca/P molar ratio, but also may be influenced by other factors such as local pH, chemical composition, crystallinity, particle size and porosity of material [5].

Bioceramics when in contact with body fluids and tissues, in this interface material-tissue, suffer reactions at the molecular scale of type dissolution preferably by the release of Ca2+ and PO4 3- ions; however, in this interface there is an increase of local pH promoted by Ca2+ ion release. This increase in pH stimulates alkaline phosphatase activity in pre-existing osteoblas‐ tic cells and in newly-differentiated active osteoblasts to synthesize more alkaline phosphatase, type I collagen, non-collagen proteins and others. Therefore, pH at the material-tissue interface is gradually reestablished, while occurs the nucleation of crystals of calcium phosphate to the collagen fibers until forming a chemically phase more stable. This event is related to PO4 3- ion release from ATP molecules, pyrophosphate and others, which contain PO4 3- ion from adjacent tissues. Moreover the action of biological buffers containing HCO3- ion, which favor the precipitation of carbonate-apatite as well as the decrease of chemical mediators locally, produced by leukocytes [13]. Table 1 shows the occurrence of calcium phosphates in biological systems.


**Table 1.** Main calcium phosphate phases. Apatite phase, chemical formula and Ca/P molar ratio.

Bioactive ceramics have been used as bone substitute materials for maxillary sinus lift, alveolar ridge augmentation, inlay bone grafting and as coatings for titanium and their respective alloys. However, bioceramics present a limitation in clinical application due to their low mechanical properties, for instance, low elastic modulus, when compared to other metallic and polymeric biomaterials. Therefore, these ceramic materials cannot be used in sites where there is a high mechanical loading, but can be used for bone fill materials and coatings of metallic surfaces or materials of high mechanical properties [16, 17]. These coatings may accelerate initial stabilization of implants and stimulating bone appositions on the implant surface, promoting a rapid fixation of these devices [18].

The bioceramics may be employed in dense and porous forms. Despite the increase in porosity decrease the mechanical strength of ceramics, the existence of isolated pores with suitable dimensions can favor the ingrowth of tissue through of these pores, promoting a strong entanglement between the material and newly formed tissue [19], moreover this porosity may promote circulation of biological fluids, increases the specific surface area, and thus acceler‐ ating the biodegradability.

Bioceramics can be single crystals (sapphire), polycrystalline [alumina, hydroxyapatite (HA), tricalcium phosphate (TCP)] or semi-crystalline structure as glass-ceramics (Ceravital® or A/W glass-ceramic) and composites, which have an amorphous phase and one or more crystalline phases. In addition, bioactive glasses (Bioglass®, PerioGlas®, BioGran®) which are a group of glass-ceramic consist in a structure of amorphous solids.

The initial medical applications of CS were documented in 1961 [20]. This material, plaster of Paris, was used in many bone defects of trauma. In the dental field, one of the first reports of the use of CS was in 1961 by Lebourg and Biou [21]. These authors implanted CS in alveoli after extraction of third molars, even in other bone defects in the mandible and maxilla. After three to four weeks it has been observed that the material had been completely resorbed, and bone healing was accelerated in the treated areas in comparison with the control. The authors concluded that CS was a favorable material for the treatment of bone defects and they justify it by the ability of the material to supply essential inorganic ions for the repair process.

Clinical studies showed positive results regarding the use of CS as material for bone fill and barrier to the preservation of alveolar ridge, post-dental extraction, providing a barrier which stabilizes the clot, assisting in the healing and bone regeneration of the local to receive the implant. The use of CS hemihydrates (CS) (powder, particulate or cement form) and CS associated with demineralized freeze-dried bone (DFDB) in bone defects, post-extraction dental and periodontal defects, promotes the increase of the quality and quantity of newly formed bone preserving the dimensions of alveolar ridge [22-24]. Moreover, CS or CS associ‐ ated with DFDB when used to maxillary sinus lift, this bone substitute, favors a good primary stability of dental implants and with relative bone density [25-27]. In addition to these advantages, CS is a BS rapidly resorbable and promotes angiogenesis [27-29]; however, in some clinical situations this rapid absorption *in vivo*, may be a disadvantage, due to its degradation which often occurs before the new bone formation.

**Apatite phase Formula Ca/P** Monocalcium phosphate monohydrate - MCPH Ca(H2PO4)2·H2O 0.5 Monocalcium phosphate anhydrous - MCP Ca(H2PO4)2 0.5 Dicalcium phosphate dihydrate (Brushite) - DCPD CaHPO4·2H2O 1.0 Dicalcium phosphate anhydrous (Monetite) - DCP CaHPO4 1.0 Octacalcium phosphate - OCP Ca8H2(PO4)6·5H2O 1.33 Amorphous calcium phosphate - ACP Ca*x*(PO4)*y*·*n*H2O 1.2 - 2.2 α or β-Tricalcium phosphate - TCP Ca3(PO4)2 1.48 - 1.50 Calcium-deficient hydroxyapatite - CDHA Ca9(HPO4)(PO4)5(OH) 1.5 Hydroxyapatite - HA Ca10(PO4)6(OH)2 1.67

**Table 1.** Main calcium phosphate phases. Apatite phase, chemical formula and Ca/P molar ratio.

promoting a rapid fixation of these devices [18].

ating the biodegradability.

28 Current Concepts in Dental Implantology

Bioactive ceramics have been used as bone substitute materials for maxillary sinus lift, alveolar ridge augmentation, inlay bone grafting and as coatings for titanium and their respective alloys. However, bioceramics present a limitation in clinical application due to their low mechanical properties, for instance, low elastic modulus, when compared to other metallic and polymeric biomaterials. Therefore, these ceramic materials cannot be used in sites where there is a high mechanical loading, but can be used for bone fill materials and coatings of metallic surfaces or materials of high mechanical properties [16, 17]. These coatings may accelerate initial stabilization of implants and stimulating bone appositions on the implant surface,

The bioceramics may be employed in dense and porous forms. Despite the increase in porosity decrease the mechanical strength of ceramics, the existence of isolated pores with suitable dimensions can favor the ingrowth of tissue through of these pores, promoting a strong entanglement between the material and newly formed tissue [19], moreover this porosity may promote circulation of biological fluids, increases the specific surface area, and thus acceler‐

Bioceramics can be single crystals (sapphire), polycrystalline [alumina, hydroxyapatite (HA), tricalcium phosphate (TCP)] or semi-crystalline structure as glass-ceramics (Ceravital® or A/W glass-ceramic) and composites, which have an amorphous phase and one or more crystalline phases. In addition, bioactive glasses (Bioglass®, PerioGlas®, BioGran®) which are

The initial medical applications of CS were documented in 1961 [20]. This material, plaster of Paris, was used in many bone defects of trauma. In the dental field, one of the first reports of the use of CS was in 1961 by Lebourg and Biou [21]. These authors implanted CS in alveoli after extraction of third molars, even in other bone defects in the mandible and maxilla. After three to four weeks it has been observed that the material had been completely resorbed, and bone healing was accelerated in the treated areas in comparison with the control. The authors concluded that CS was a favorable material for the treatment of bone defects and they justify it by the ability of the material to supply essential inorganic ions for the repair process.

Clinical studies showed positive results regarding the use of CS as material for bone fill and barrier to the preservation of alveolar ridge, post-dental extraction, providing a barrier which

a group of glass-ceramic consist in a structure of amorphous solids.

Other the bioactive ceramics most commonly investigated as bone substitute materials are HA, β-TCP and bioactive glasses. Synthetic HA, β-TCP and biphasic calcium phosphates (HA:β-TCP) are routinely employed as BS in block or granule forms. Furthermore, cements based on HA and/or β-TCP are excellent bone fill materials, due to their easy manipulation and favor the bone contour, moreover, are clinically used by their similarity to the bone inorganic composition and by osteoconductive property. On the other hand, bioactive glasses are most commonly used in granule forms.

For several years, synthetic HA was used as the main method in the reconstruction of bone defects involving the craniofacial region, oral surgery, orthopedic and implant dentistry [14, 30-32]. HA presents some disadvantages related to its resorption, because it is hardly absorbed, which hampers the remodeling and the new bone formation, and results in poor local stability or permanent stress concentration. Currently, biphasic calcium phosphates, mixtures contain‐ ing HA/TCP (α-TCP or β-TCP) are preferably used in clinical practice with varied proportions between HA and TCP [33-38] due to their considerably difference in the resorption rate, which HA reabsorbs very slowly compared with TCP. The difference in the resorption rate influences in the osteoconductive property of these materials, TCPs are more osteoconductive than HA, due to their greater biodegradability rate in relation to HA [13, 39, 40]. Clinical and experi‐ mental studies have shown that mixture HA/TCP promotes intense activity of bone formation with high osteoconductivity [34, 41-43], whose mixture has demonstrated to be an excellent material for sinus lift [34, 37, 38]. However, even the resorption rate of TCP being faster than HA, clinical studies that used just TCP reported presence of TCP particles after long-term postoperative of maxillary sinus lift and mandible defects [44-49]. Results show that β-TCP is a good material for grafting [44-51], on this account also promotes stability of implants increasing the survival rate [48].

Furthermore, these bioceramics when associated with biopolymers as hyaluronic acid [49] and collagen [52, 53] or other osteoinductive biomolecules (growth factors: bone morphogenetic protein-2 (BMP-2); fibroblast growth factor-2 (FGF-2)) [54-59] have displayed promising results for bone regeneration. These associations have promoted a better quality and quantity of newly formed bone [49, 52, 53, 56, 57, 59], consequently they can improve the primary stability of implants.

Other subgroups of bioceramics quite used as BS material are bioactive glasses and glassceramics. Silica glasses are generally classified as a subgroup of ceramics. The glass-ceramics are materials formed by a glass matrix reinforced by ceramic crystals obtained from controlled crystallization processes [60]. This crystallization process can take place by heat treatment, resulting in a material containing various crystal phases and controlled grain sizes [10]. The glass-ceramic materials have relatively high mechanical strengths, low coefficients of thermal expansion and good biological compatibility. Possibly the most attractive attribute of this class of materials is the ease with which they may be fabricated, in which conventional glass-forming techniques may be used [8]. On the other hand, bioactive glasses present limitations in certain mechanical properties such as low strength and toughness.

In the late 1960s and early 1970s, the several researches for developing implant materials with a better biocompatibility resulted in the new concept of bioceramic materials, which could mimic natural bone tissue. During this period, Hench and coworkers [61] developed a new biocompatible material, silica-based melt-derived glass, for bonding fractured bones, a bioactive glass denominate 45S5 Bioglass®. This denomination was given because the material mimicked normal bone and to stimulate the new bone formation between the fractures [62]. Bioglass® is a commercially available family of bioactive glasses, based on SiO2, Na2O, CaO and P2O5 in specific proportions, and was one of the first materials completely synthetic with excellent osteoconductive properties, which seamlessly binds to bone [61, 63]. The bioactive glasses, since their discovered, have been widely used in dentistry for bone defects repair/ reconstruction, because these glasses exhibit bone bonding, a phenomenon also observed with other bioactive ceramics [64]. Bioglass offers advantages such as control of resorption rate, excellent osteoconductivity, bioactivity, and capacity for delivering cells. This process is a result of the surface reactive silica, calcium, and phosphate groups that are characteristic of these materials. Silica is believed to play a critical role in bioactivity [64].

In the 1970s, Brömer and coworkers [65] developed a glass-ceramic, Ceravital® through reduction of alkali oxides in the composition and the phase precipitation of the glass matrix by heat treatment of Bioglass®. Ceravital® has been used for small bone defects/structure reconstruction as dentistry [66] as other medical applications, *e.g.* tympanoplasties [67].

The use of bioactive glasses as alloplastic bone graft materials for alveolar ridge augmentation [68-71] and maxillary sinus lift [72-74] procedures has received increasing attention in implant dentistry. Besides Bioglass® other commercial types of bioactive glass have been used for bone repair such as PerioGlas® [70, 75, 76] and BioGran® [71-73]. Studies have reported presence of bioactive glass long-term postoperative (1-2 years) [72, 73].

Moreover, several studies have shown that bioactive glasses and glass ceramics stimulates the secretion of angiogenic growth factors on fibroblasts and endothelial cell proliferation [77, 78].

Although they are quite biocompatible and exhibit bone bonding, bioactive glasses are not osteoinductive and are not capable of forming bone in ectopic sites (although they can be used to deliver osteopromotive growth factors) [64].

Another glass-ceramic with potential for application in implant dentistry is apatite/wollas‐ tonite (A/W), which was developed by Kokubo et al., in 1982 [79]. This material presents a great capacity for bone bonding and moderate mechanical strength [80], with excellent results in orthopedic applications [81-83]. The resorption rate of this glass-ceramic can be increased when associated with β-TCP [84]. According to Carrodeguas et al. (2008) [85] the report that the new ceramics containing wollastonite did not exhibit toxicity in cell culture with human fibroblasts. Moreover, they are biocompatible, resorbable and bioactive releasing ions of silica and calcium in the physiological environment, which are capable of stimulating cells to produce bone matrix [86, 87].

Biosilicate®, glass-ceramic developed by Zanotto and coworkers in 2004 [88], which is highly crystalline (~100%), has an elastic modulus value close to cortical bone, and displays high level of bioactivity [89-91]. It is biocompatible and provides efficient new bone formation in sockets preserving alveolar bone ridge height and allowing osseointegration of implants [92].


Table 2 summarizes some bioceramics used in clinical practice.

crystallization processes [60]. This crystallization process can take place by heat treatment, resulting in a material containing various crystal phases and controlled grain sizes [10]. The glass-ceramic materials have relatively high mechanical strengths, low coefficients of thermal expansion and good biological compatibility. Possibly the most attractive attribute of this class of materials is the ease with which they may be fabricated, in which conventional glass-forming techniques may be used [8]. On the other hand, bioactive glasses present limitations in certain

In the late 1960s and early 1970s, the several researches for developing implant materials with a better biocompatibility resulted in the new concept of bioceramic materials, which could mimic natural bone tissue. During this period, Hench and coworkers [61] developed a new biocompatible material, silica-based melt-derived glass, for bonding fractured bones, a bioactive glass denominate 45S5 Bioglass®. This denomination was given because the material mimicked normal bone and to stimulate the new bone formation between the fractures [62]. Bioglass® is a commercially available family of bioactive glasses, based on SiO2, Na2O, CaO and P2O5 in specific proportions, and was one of the first materials completely synthetic with excellent osteoconductive properties, which seamlessly binds to bone [61, 63]. The bioactive glasses, since their discovered, have been widely used in dentistry for bone defects repair/ reconstruction, because these glasses exhibit bone bonding, a phenomenon also observed with other bioactive ceramics [64]. Bioglass offers advantages such as control of resorption rate, excellent osteoconductivity, bioactivity, and capacity for delivering cells. This process is a result of the surface reactive silica, calcium, and phosphate groups that are characteristic of

In the 1970s, Brömer and coworkers [65] developed a glass-ceramic, Ceravital® through reduction of alkali oxides in the composition and the phase precipitation of the glass matrix by heat treatment of Bioglass®. Ceravital® has been used for small bone defects/structure reconstruction as dentistry [66] as other medical applications, *e.g.* tympanoplasties [67].

The use of bioactive glasses as alloplastic bone graft materials for alveolar ridge augmentation [68-71] and maxillary sinus lift [72-74] procedures has received increasing attention in implant dentistry. Besides Bioglass® other commercial types of bioactive glass have been used for bone repair such as PerioGlas® [70, 75, 76] and BioGran® [71-73]. Studies have reported presence

Moreover, several studies have shown that bioactive glasses and glass ceramics stimulates the secretion of angiogenic growth factors on fibroblasts and endothelial cell proliferation [77, 78]. Although they are quite biocompatible and exhibit bone bonding, bioactive glasses are not osteoinductive and are not capable of forming bone in ectopic sites (although they can be used

Another glass-ceramic with potential for application in implant dentistry is apatite/wollas‐ tonite (A/W), which was developed by Kokubo et al., in 1982 [79]. This material presents a great capacity for bone bonding and moderate mechanical strength [80], with excellent results in orthopedic applications [81-83]. The resorption rate of this glass-ceramic can be increased when associated with β-TCP [84]. According to Carrodeguas et al. (2008) [85] the report that the new ceramics containing wollastonite did not exhibit toxicity in cell culture with human

mechanical properties such as low strength and toughness.

30 Current Concepts in Dental Implantology

these materials. Silica is believed to play a critical role in bioactivity [64].

of bioactive glass long-term postoperative (1-2 years) [72, 73].

to deliver osteopromotive growth factors) [64].

**Table 2.** Some bioceramics used in clinical practice.

### **3. Composite and polymer-based bone substitutes**

Composite materials are described as those that have at least two components or two phases with distinct physical and chemical properties that are separated by an interface. The purpose of developing composites is to associate different materials to produce a single device with superior properties compared with the isolated components [10]. Separately the constituents of the composite maintains their features, however when mixed they constitute a compound with their own properties inherent to the new composition. Two examples of natural fiber composites are: 1. wood, which is basically formed of cellulose fibers and lignin (amorphous resin which binds the cellulose fibers); 2. bone tissue, which is formed by an inorganic phase, essentially carbonate-apatite, placed in an organic matrix, whose composition is about 95% type I collagen. Therefore, the composites are formed by the matrix, which is the continuous phase ("fiber network") and involves the other phase, the dispersed one. Among the several types of composites, polymer composites exhibit some advantages such as: low weight, corrosion resistance, high temperature resistance and good mechanical properties when compared to conventional engineering materials [98].

However, the current goal of tissue engineering is the development of polymer composites, metal-free, with mechanical properties similar to living tissue, especially bone tissue, for partial or total replacement or reconstruction of the organ or tissue being repaired.

Polymers can be classified as natural or synthetic and degradable or non-degradable. These compounds provide versatility in their structure and can modulate the mechanical properties of other compounds like ceramics. Degradable polymers may be advantageous in certain clinical situations.

Composites produced from a combination of natural polymers (collagen, cellulose, polyhy‐ droxybutyrate), or synthetic [poly (lactic-*co*-glycolic acid) (PLGA), poly(lactic acid) (PLA), poly(ε-caprolactone) (PCL)] associated with bioceramics have been highlighted in academic community [99-106] because they are biocompatible, excellent osteoconductors, bioactives, have satisfactory mechanical properties, and are absorbable, therefore they are potential materials for application in regenerative medicine therapies.

Natural polymers or biopolymers have attractive properties for the construction of 3D scaffolds, such as biocompatibility and biodegradability. Bioactivity of these polymers, if you need to improve, can also be controlled by the addition of chemicals, proteins, peptides, and cells. The most commonly studied natural polymers for the purpose of bone engineering are collagen/gelatin, chitosan, silk, alginate, hyaluronic acid, and peptides [107].

Currently, the most BS, available commercially, for clinical application in implant dentistry based on polymers are barrier membranes for guided bone regeneration (GBR) or collagen sponge/BMPs (INFUSE®) for bone reconstruction.

At the beginning of the use of the GBR the treatment was preferably with non-resorbable membranes based on expanded polytetrafluoroethylene (e-PTFE) [108, 109], because of its inert characteristic and their biological effective and predictable results as a mechanical barrier. However, resorbable membranes have been widely used to the development of new bioma‐ terials, due to the predictable and similar results compared to the non-resorbable membranes [110-112], moreover resorbable membranes can be used on peri-implant defects, *i. e.,* an advantage in relation to non-resorbable membranes. Among the membranes the most used as resorbable membranes are: collagen, PLA and PLGA [110-113].

**3. Composite and polymer-based bone substitutes**

32 Current Concepts in Dental Implantology

compared to conventional engineering materials [98].

materials for application in regenerative medicine therapies.

sponge/BMPs (INFUSE®) for bone reconstruction.

clinical situations.

Composite materials are described as those that have at least two components or two phases with distinct physical and chemical properties that are separated by an interface. The purpose of developing composites is to associate different materials to produce a single device with superior properties compared with the isolated components [10]. Separately the constituents of the composite maintains their features, however when mixed they constitute a compound with their own properties inherent to the new composition. Two examples of natural fiber composites are: 1. wood, which is basically formed of cellulose fibers and lignin (amorphous resin which binds the cellulose fibers); 2. bone tissue, which is formed by an inorganic phase, essentially carbonate-apatite, placed in an organic matrix, whose composition is about 95% type I collagen. Therefore, the composites are formed by the matrix, which is the continuous phase ("fiber network") and involves the other phase, the dispersed one. Among the several types of composites, polymer composites exhibit some advantages such as: low weight, corrosion resistance, high temperature resistance and good mechanical properties when

However, the current goal of tissue engineering is the development of polymer composites, metal-free, with mechanical properties similar to living tissue, especially bone tissue, for partial

Polymers can be classified as natural or synthetic and degradable or non-degradable. These compounds provide versatility in their structure and can modulate the mechanical properties of other compounds like ceramics. Degradable polymers may be advantageous in certain

Composites produced from a combination of natural polymers (collagen, cellulose, polyhy‐ droxybutyrate), or synthetic [poly (lactic-*co*-glycolic acid) (PLGA), poly(lactic acid) (PLA), poly(ε-caprolactone) (PCL)] associated with bioceramics have been highlighted in academic community [99-106] because they are biocompatible, excellent osteoconductors, bioactives, have satisfactory mechanical properties, and are absorbable, therefore they are potential

Natural polymers or biopolymers have attractive properties for the construction of 3D scaffolds, such as biocompatibility and biodegradability. Bioactivity of these polymers, if you need to improve, can also be controlled by the addition of chemicals, proteins, peptides, and cells. The most commonly studied natural polymers for the purpose of bone engineering are

Currently, the most BS, available commercially, for clinical application in implant dentistry based on polymers are barrier membranes for guided bone regeneration (GBR) or collagen

At the beginning of the use of the GBR the treatment was preferably with non-resorbable membranes based on expanded polytetrafluoroethylene (e-PTFE) [108, 109], because of its inert characteristic and their biological effective and predictable results as a mechanical barrier.

or total replacement or reconstruction of the organ or tissue being repaired.

collagen/gelatin, chitosan, silk, alginate, hyaluronic acid, and peptides [107].

The type I collagen is an example of biopolymers quite used to the development of BS. It is a matrix that provides a favorable environment for induction of osteoblasts differentiation *in vitro* and osteogenesis *in vivo* [114]. Type III collagen constitutes the reticular fibers of the tissues and is also widely used in the manufacture of membranes for GBR. The non-mineral‐ ized collagen membranes are usually weak (low tensile strength) making their clinical manipulation difficult. The great advantage of them is the excellent cell affinity stimulating the chemotaxis of fibroblasts and acting as support migration of these cells (osteoconduction). Other advantages are: good adaptation to bone surfaces, especially to dental roots and hemostatic effect [113]. When embedded in the bone matrix they are gradually metabolized by the action of collagenase, or can be partially embedded in the bone matrix.

The resorption of collagen occurs parallel to bone formation as well as by the formation of new periodontal tissue such as cementum and periodontal ligament. The resorption time ranges from 06 to 08 weeks depending on the strength of the material, however it can last from 04 to 06 months [115]. In this case, the new bone is protected against the growth of connective tissue within the defect area. Despite prevent cellular infiltration, this membrane is permeable to nutrients, and the degradation occurs through enzymatic reactions without irritating the surrounding tissues. These membranes have adequate mechanical resistance [116]; moreover, they can facilitate the maintenance of the space to be regenerated, similar to the non-resorbable membranes.

The collagen membranes developed in recent years have shown optimal physicochemical characteristics for clinical application [117-121]. According to the literature, the determination of the density of crosslinking reaction (cross-linking) directly influences the physical properties of collagen matrices, *i. e.*, the increased crosslinking of collagen fibrils provides increased tensile strength and enzymatic degradation, and higher thermal stability [118, 119, 121, 122].

The membrane Bio-Gide® has been the most membrane widely used for GBR in the last years [113, 123-126], which is composed of type I and III collagen from porcine. This membrane has a bilayer structure with a compact layer and other porous. The porous layer (inner face) promotes a three-dimensional matrix for bone integration. The natural collagen structure of the Bio-Gide ® is ideal for tissue adhesion, while the newly formed bone is protected against the growth of connective tissue into the defect region; while preventing cellular infiltration this membrane is permeable to nutrients, and the degradation occurs through enzymatic reactions without irritating the surrounding tissues.

Studies with collagen membrane for GBR have reported satisfactory results *in vivo*, for example, the rate of bone regeneration has a similar efficacy to the e-PTFE membranes. This occurred due to the advent of collagen membranes with good mechanical strength. In the past, it was difficult to obtain such satisfactory and predictable results due to the difficulty of producing collagen membranes with these characteristics [123, 127].

#### **4. New perspectives for bone substitutes**

#### **4.1. Bone tissue engineering**

In recent years, a new generation of bone substitutes have been developed in an attempt to obtain materials closer to the autograft standard by using biomaterials capable of inducing specific cellular responses at the molecular level, by integrating the bioactivity and biode‐ gradability of these materials [107]. These BS are being based on the concept of bone tissue engineering. Tissue engineering/regenerative medicine has emerged as an interdisciplinary field that includes cell-based therapies and use of porous-bioactive materials for development of functional substitutes for the repair or replacement of damaged tissues or organs [128]. Tissue engineering has achieved great progress in the development of three-dimensional materials (scaffolds) for repair or replacement of damaged tissues or organs, including alloplastic materials such as bioceramics, bioactive glasses and polymers [60, 61, 63, 129, 130] in association to the signaling pathway, molecular and/or biophysical stimulation. Thus, tissue engineering is based on three elements that must be in synergism: matrix (scaffolds), cells and signals (mechanical and/or molecules: proteins, peptides and cytokines) [131, 132]; the absence or dysfunction of one element will halt or delay tissue regeneration. Furthermore, the tissue formation inside the scaffolds is directly influenced by porosity rate and pore size. In the case of bone formation, scaffolds should preferably have pores greater than 300 µM for promoting a good vascularization and a new bone formation, preventing hypoxia and induction of endochondral formation before the osteogenesis [133].

Porous scaffolds have been developed by variety of conventional methods from alloplastic materials, such as particle/salt leaching, chemical/gas foaming, fiber bonding, solvent casting, melting molding, phase separation and freeze drying [134, 135]. However, these methods present some limitations due to their lack of the controlled formation of pores and do not produce interconnected structures to favoring cell growth inside the structure. For overcome these disadvantages, additive manufacturing (AM), also otherwise known as three-dimen‐ sional (3D) printing, is a promising option for the production of scaffolds particularly for bone substitutes.

This technique consists in constructing 3D scaffolds by a tool for direct digital fabrication that selectively prints a respective material (layer-by-layer) into/onto a bed, whose shape is given by CAD specifications [136]. A distinctive feature of this layer-by-layer printing process, is the printing of structures with high geometric complexity and well-defined architecture as well as patient-specific implant designs, which are not possible to be constructed by any other manufacturing method (Figure 1).

Some of the commercially available AM techniques are 3DP (ExOne, PA), fused deposition modeling (FDM, Stratasys, MN), selective laser sintering (SLS, 3D Systems, CA), stereolithog‐

it was difficult to obtain such satisfactory and predictable results due to the difficulty of

In recent years, a new generation of bone substitutes have been developed in an attempt to obtain materials closer to the autograft standard by using biomaterials capable of inducing specific cellular responses at the molecular level, by integrating the bioactivity and biode‐ gradability of these materials [107]. These BS are being based on the concept of bone tissue engineering. Tissue engineering/regenerative medicine has emerged as an interdisciplinary field that includes cell-based therapies and use of porous-bioactive materials for development of functional substitutes for the repair or replacement of damaged tissues or organs [128]. Tissue engineering has achieved great progress in the development of three-dimensional materials (scaffolds) for repair or replacement of damaged tissues or organs, including alloplastic materials such as bioceramics, bioactive glasses and polymers [60, 61, 63, 129, 130] in association to the signaling pathway, molecular and/or biophysical stimulation. Thus, tissue engineering is based on three elements that must be in synergism: matrix (scaffolds), cells and signals (mechanical and/or molecules: proteins, peptides and cytokines) [131, 132]; the absence or dysfunction of one element will halt or delay tissue regeneration. Furthermore, the tissue formation inside the scaffolds is directly influenced by porosity rate and pore size. In the case of bone formation, scaffolds should preferably have pores greater than 300 µM for promoting a good vascularization and a new bone formation, preventing hypoxia and induction of

Porous scaffolds have been developed by variety of conventional methods from alloplastic materials, such as particle/salt leaching, chemical/gas foaming, fiber bonding, solvent casting, melting molding, phase separation and freeze drying [134, 135]. However, these methods present some limitations due to their lack of the controlled formation of pores and do not produce interconnected structures to favoring cell growth inside the structure. For overcome these disadvantages, additive manufacturing (AM), also otherwise known as three-dimen‐ sional (3D) printing, is a promising option for the production of scaffolds particularly for bone

This technique consists in constructing 3D scaffolds by a tool for direct digital fabrication that selectively prints a respective material (layer-by-layer) into/onto a bed, whose shape is given by CAD specifications [136]. A distinctive feature of this layer-by-layer printing process, is the printing of structures with high geometric complexity and well-defined architecture as well as patient-specific implant designs, which are not possible to be constructed by any other

Some of the commercially available AM techniques are 3DP (ExOne, PA), fused deposition modeling (FDM, Stratasys, MN), selective laser sintering (SLS, 3D Systems, CA), stereolithog‐

producing collagen membranes with these characteristics [123, 127].

**4. New perspectives for bone substitutes**

endochondral formation before the osteogenesis [133].

**4.1. Bone tissue engineering**

34 Current Concepts in Dental Implantology

substitutes.

manufacturing method (Figure 1).

**Figure 1.** Printed mandibular condyle by Fused Deposition Modeling (FDM) process. *Image provided by Centro de Tecno‐ logia de Informação (CTI) – Renato Archer (Campinas, Brazil).*

raphy (3D Systems, CA), 3D plotting (Fraunhofer Institute for Materials Research and BeamTechnology, Germany), as well as various methods [135]. These AM techniques can be classified as – (a) extrusion (deformation + solidification), (b) polymerization, (c) laser-assisted sintering, and (d) direct writing-based processes [135, 136].

Recently, some researches have performed using 3D printed scaffolds for bone regeneration. Among them we can highlight the use of this 3D printed BS based on bioceramics for vertical bone augmentation as onlay graft. The results shown by these studies are promising and efficient as bone graft compared to autografts [137-139]. Li et al (2011) [140] reported a case report of a 3D printed mandibular condyle implant made of nano-hydroxyapatite/polyamide. The clinical results suggest that this type of 3D printed implant can be a viable alternative to the autografts for maxillofacial defects. 3D printed scaffolds base on PLGA have also demon‐ strated good results for bone regeneration [141]. So far, as signaling pathway, growth factor and drug delivery, have been reported the use recombinant human BMP-2 (rhBMP-2) [142] and alendronate [143]. PCL/PLGA/gelatin scaffolds containing rhBMP-2 did not induce the osteogenic differentiation of mesenchymal stem cells *in vitro*, however, in the preclinical experiements, PCL/PLGA/collagen/rhBMP-2 showed the best bone healing quality at both weeks (4 and 8 after implantation) without inflammatory response. On the other hand, a large number of macrophages indicated severe inflammation caused by burst release of rhBMP-2 [142]. In addition, a study about 3D-printed bioceramic scaffolds containing alendronate shows that *in vivo* local alendronate delivery from PCL-coated 3DP TCP scaffolds could further induce increased early bone formation [143].

#### **4.2. Growth factors**

Osteoprogenitor cells, osteoblasts and osteoclasts are under growth factors activity. The role of growth factors is not only to stimulate cell proliferation through cell cycle regulation by initiating mitosis but also to maintain cell survival and to stimulate the migration, differen‐ tiation and apoptosis as well. Osteoblasts proliferate mediated by growth factors released by themselves and by the bone during the resorption process. Among the most important are the TGF-β and the factors released by the bone matrix, such as growth factor similar to insulin (IGF-1 and 2), the fibroblast growth factor (FGF-2) and growth factor derived from platelets (PDGF) [144, 145] which are potent mitogens [146, 147].

Moreover, other factors are secreted during the repair process, such as BMPs and angiogen‐ ic factors (vascular endothelial growth factor - VEGF) [147]. TGF-β presents activity in embryonic development, cell differentiation, hormone secretion and immune function, and acts synergistically with TGF-α in the induction of phenotypic transformation [146]. The TGF-β superfamily includes TGF-β1, TGF-β2, TGF-β3 and other important factors, such as BMPs 1-8, which promote several stages of intramembranous and endochondral ossifica‐ tion during bone repair [148].

Among these several growth factors, BMP-2 has received attention from the scientific com‐ munity due its use in combination with different scaffolds to promote bone repair, especially in tissue engineering. The literature, in the constant search for developing a biomaterial with excellent osteoinductive properties, such as autogenous bone for reconstructive surgery, has recently shown that some polymers and bioceramics can be great carriers for BMPs, especially the collagen [54, 57, 59, 143, 149-152].

The concept of osteoinduction was first described by Urist in 1965 [153] when he observed new bone formation inside the demineralized bone matrices. Since then, these proteins, BMPs, have been reported as factors responsible for bone neoformation [149, 154]. BMPs attract mesen‐ chymal cells to the site of bone formation by chemotaxis, and induces the conversion of these cells to a pre-osteoblastic lineage. BMP-2, 6 and 9 are described as important for the initiation of the differentiation of mesenchymal cells into pre-osteoblasts, while BMP-4 and 7 promote the differentiation of pre-osteoblasts into osteoblasts [155].

Clinical studies with rhBMP-2 using collagen as a carrier for surgical protocol of vertebral column showed similar or better results compared to autografts [150, 156-158]. However, the cost-effectiveness ratio of BMPs is questionable because of the large required amount (12 mg, 1.5 mg.mL-1 therapeutic dose of INFUSE®) to obtain an effective bone repair in comparison to conventional surgical techniques [159].

In recent years, the use of synthetic peptides has been highlighted due to the ease of recognition and binding to specific sites of the extracellular matrix proteins increasing the material-cell interaction and for do not promote an immunogenic reaction. In this context, the specific amino acid sequence Arg-Gly-Asp (RGD) of the extracellular matrix proteins, such as fibronectin and osteopontin is recognized by the transmembrane receptors (integrins) [160, 161], and promotes better adhesion, and consequently a greater proliferation of osteoblastic cells. This RGD sequence has been widely used for functionalization of biomaterials in order to stimulate the initial process of cell adhesion [162-164].

#### **5. Cytototoxic, genotoxic and mutagenic tests of biomaterials**

**4.2. Growth factors**

36 Current Concepts in Dental Implantology

tion during bone repair [148].

the collagen [54, 57, 59, 143, 149-152].

conventional surgical techniques [159].

Osteoprogenitor cells, osteoblasts and osteoclasts are under growth factors activity. The role of growth factors is not only to stimulate cell proliferation through cell cycle regulation by initiating mitosis but also to maintain cell survival and to stimulate the migration, differen‐ tiation and apoptosis as well. Osteoblasts proliferate mediated by growth factors released by themselves and by the bone during the resorption process. Among the most important are the TGF-β and the factors released by the bone matrix, such as growth factor similar to insulin (IGF-1 and 2), the fibroblast growth factor (FGF-2) and growth factor derived from platelets

Moreover, other factors are secreted during the repair process, such as BMPs and angiogen‐ ic factors (vascular endothelial growth factor - VEGF) [147]. TGF-β presents activity in embryonic development, cell differentiation, hormone secretion and immune function, and acts synergistically with TGF-α in the induction of phenotypic transformation [146]. The TGF-β superfamily includes TGF-β1, TGF-β2, TGF-β3 and other important factors, such as BMPs 1-8, which promote several stages of intramembranous and endochondral ossifica‐

Among these several growth factors, BMP-2 has received attention from the scientific com‐ munity due its use in combination with different scaffolds to promote bone repair, especially in tissue engineering. The literature, in the constant search for developing a biomaterial with excellent osteoinductive properties, such as autogenous bone for reconstructive surgery, has recently shown that some polymers and bioceramics can be great carriers for BMPs, especially

The concept of osteoinduction was first described by Urist in 1965 [153] when he observed new bone formation inside the demineralized bone matrices. Since then, these proteins, BMPs, have been reported as factors responsible for bone neoformation [149, 154]. BMPs attract mesen‐ chymal cells to the site of bone formation by chemotaxis, and induces the conversion of these cells to a pre-osteoblastic lineage. BMP-2, 6 and 9 are described as important for the initiation of the differentiation of mesenchymal cells into pre-osteoblasts, while BMP-4 and 7 promote

Clinical studies with rhBMP-2 using collagen as a carrier for surgical protocol of vertebral column showed similar or better results compared to autografts [150, 156-158]. However, the cost-effectiveness ratio of BMPs is questionable because of the large required amount (12 mg, 1.5 mg.mL-1 therapeutic dose of INFUSE®) to obtain an effective bone repair in comparison to

In recent years, the use of synthetic peptides has been highlighted due to the ease of recognition and binding to specific sites of the extracellular matrix proteins increasing the material-cell interaction and for do not promote an immunogenic reaction. In this context, the specific amino acid sequence Arg-Gly-Asp (RGD) of the extracellular matrix proteins, such as fibronectin and osteopontin is recognized by the transmembrane receptors (integrins) [160, 161], and promotes better adhesion, and consequently a greater proliferation of osteoblastic cells. This RGD

(PDGF) [144, 145] which are potent mitogens [146, 147].

the differentiation of pre-osteoblasts into osteoblasts [155].

Biomaterials may have low, medium or high potential risk to human safety, depending on the type and extent of the patient contact. Safety assessments of medical biomaterials are guided by the toxicological guidelines recommended by the International Organization of Standard‐ ization (ISO 10993-1/EN 30993-1). One of the recommended and appropriate steps for the biological assessment of potential medical biomaterials consists of an *in vitro* evaluation of cytotoxicity and genotoxicity [165].

It is important to consider the possible impact of the composition on processes linked to cell proliferation and survival. It is essential to ensure that the proportional amounts of each component do not impoverish the cytocompatibility of the final composite, due to the release of toxic or irritating components. Therefore, *in vitro* cytotoxicity tests represent critical requirements previous to the clinical application of such materials (ISO 10993-12; [166])The choice of one or more cytotoxic tests depends on the nature of the sample to be evaluated, the potential site of use and the nature of the use (ISO 10993-5).

Cytotoxicity can be evaluated regarding the cell viability. XTT is a soluble variation of the widely employed MTT test, which accounts for mitochondrial activity in the tested material [166, 167]. Dimethyl sulfoxide solubilization of cellular-generated 3-(4,5-dimethylthiazol-2 yl)-2,5-diphenyltetrazolium bromide (MTT) - formazan presents several inherent disadvan‐ tages of this assay, including the safety hazard of personnel exposure to large quantities of dimethyl sulfoxide, the deleterious effects of this solvent on laboratory equipment, and the inefficient metabolism of MTT by some human cell lines [167, 168]. Recognition of these limitations prompted development of possible alternative microculture tetrazolium assays utilizing a different tetrazolium reagent, 2,3-bis(2-methoxy-4-nitro-5-sulfophenyl)-5-[(phenyl‐ amino)carbonyl]-2H-tetrazolium hydroxide (XTT), which is metabolically reduced in viable cells to a water-soluble formazan product. This reagent allows direct absorbance readings, therefore eliminating a solubilization step and shortening the microculture growth assay procedure [167]. Therefore, in XTT test mitochondrial dehydrogenase activity is measured by the ability of such enzymes to reduce the reagent XTT to soluble formazan salts, with differing color.

To evaluate cell survival, Neutral Red uptake cytotoxicity test detects membrane intact viable cells by incorporation of the dye in their lysosomes [166, 169]. It is one of the most used cytotoxicity tests with many biomedical and environmental applications and most primary cells and cell lines from diverse origin may be used [169]. The procedure is cheaper and more sensitive than other cytotoxicity tests (tetrazolium salts, enzyme leakage or protein content) [169].

Bone substitute and implant materials have been evaluated regarding cytotoxicity by different assays [166, 170-173].

It is inherent in the provision of safe medical devices that the risk of serious and irreversible effects, such as cancer or second-generation abnormalities, can be minimized to the greatest extent feasible. The assessment of mutagenic, carcinogenic and reproductive hazards is an essential component of the control of these risks (ISO 10993-3). An international standard (ISO 10993) lays down specific requirements for biocompatibility, including the tests based on the nature of the contact and the duration of implantation of the biomaterial. The standard stipulates that all materials that will be in contact with mucous, bone, or dentinal tissue if the contact exceeds 30 days, as well as all implantable devices if the contact exceeds 24h, must undergo genotoxicity testing [174].

A useful approach for assessing genotoxic activity is the single cell gel electrophoresis (SCGE) or Comet assay. Singh et al. (1988) [175] introduced a microgel technique involving electro‐ phoresis under alkaline conditions for detecting DNA damage in single cells which led to a sensitive version of the assay that could assess both double- and single-strand DNA breaks as well as the alkali labile sites expressed as frank strand breaks in the DNA. In this technique, cells are embedded in agarose gel on microscope slides, lysed by detergents and high salt, and then electrophoresed for a short period under alkaline conditions [175]. The assay is called a comet assay because the damaged cells look like a comet under a microscope. Cells with increased DNA damage display increased migration of DNA from the nucleus toward anode [176], so it appears like a comet tail that moves away from the unbroken DNA ("comet head") (Figure 2). Cells with increased DNA damage display increased migration of DNA from the nucleus toward anode [175]. Staining with different fluorescent dyes like ethidium bromide, propidium iodide, SYBR green quantifies the migrating DNA [176]. The most flexible approach for collecting comet data involves the application of image analysis techniques to individual cells, and several software programs are commercially available [176].

Some advantages of the SCGE assay is its sensitivity for detecting low levels of DNA damage, the requirement for small numbers of cells per sample, its flexibility and the short time needed to complete a study [176].

The SCGE assay has the capability to assess an increasing genotoxicity of a biomaterial model, whatever the cause and mechanism of the genotoxicity [174].

The *in vitro* micronucleus assay is well established in the field of toxicology for screening the effects of physical and chemical agents that may damage the DNA of eukaryotic cells [177]. The micronucleus assays have emerged as one of the preferred methods for assessing chro‐ mosome damage because they enable both chromosome loss and chromosome breakage to be measured reliably [178]. Because of the uncertainty of the fate of micronuclei following more than one nuclear division it is important to identify cells that have completed one nuclear division only [178]. In the cytokinesis-block micronucleus (CBMN) assay the cytokinesis is blocked using cytochalasin-B (Cyt-B). Cyt-B is an inhibitor of actin polymerization required for the formation of the microfilament ring that constricts the cytoplasm between the daughter nuclei during cytokinesis [178].

Micronuclei (MNi) are acentric chromosome fragments or whole chromosomes that are left behind during mitotic cellular division and appear in the cytoplasm of interphase cells as small additional nuclei [179]. MNi are morphologically identical to nuclei but smaller (Figure 3). The diameter of MNi usually varies between 1/16th and 1/3rd of the mean diameter of the main nuclei [180]. The number of micronuclei in 1000 binucleated cells should be scored and the frequency of MN per 1000 binucleated cells calculated [178].

It is inherent in the provision of safe medical devices that the risk of serious and irreversible effects, such as cancer or second-generation abnormalities, can be minimized to the greatest extent feasible. The assessment of mutagenic, carcinogenic and reproductive hazards is an essential component of the control of these risks (ISO 10993-3). An international standard (ISO 10993) lays down specific requirements for biocompatibility, including the tests based on the nature of the contact and the duration of implantation of the biomaterial. The standard stipulates that all materials that will be in contact with mucous, bone, or dentinal tissue if the contact exceeds 30 days, as well as all implantable devices if the contact exceeds 24h, must

A useful approach for assessing genotoxic activity is the single cell gel electrophoresis (SCGE) or Comet assay. Singh et al. (1988) [175] introduced a microgel technique involving electro‐ phoresis under alkaline conditions for detecting DNA damage in single cells which led to a sensitive version of the assay that could assess both double- and single-strand DNA breaks as well as the alkali labile sites expressed as frank strand breaks in the DNA. In this technique, cells are embedded in agarose gel on microscope slides, lysed by detergents and high salt, and then electrophoresed for a short period under alkaline conditions [175]. The assay is called a comet assay because the damaged cells look like a comet under a microscope. Cells with increased DNA damage display increased migration of DNA from the nucleus toward anode [176], so it appears like a comet tail that moves away from the unbroken DNA ("comet head") (Figure 2). Cells with increased DNA damage display increased migration of DNA from the nucleus toward anode [175]. Staining with different fluorescent dyes like ethidium bromide, propidium iodide, SYBR green quantifies the migrating DNA [176]. The most flexible approach for collecting comet data involves the application of image analysis techniques to individual

Some advantages of the SCGE assay is its sensitivity for detecting low levels of DNA damage, the requirement for small numbers of cells per sample, its flexibility and the short time needed

The SCGE assay has the capability to assess an increasing genotoxicity of a biomaterial model,

The *in vitro* micronucleus assay is well established in the field of toxicology for screening the effects of physical and chemical agents that may damage the DNA of eukaryotic cells [177]. The micronucleus assays have emerged as one of the preferred methods for assessing chro‐ mosome damage because they enable both chromosome loss and chromosome breakage to be measured reliably [178]. Because of the uncertainty of the fate of micronuclei following more than one nuclear division it is important to identify cells that have completed one nuclear division only [178]. In the cytokinesis-block micronucleus (CBMN) assay the cytokinesis is blocked using cytochalasin-B (Cyt-B). Cyt-B is an inhibitor of actin polymerization required for the formation of the microfilament ring that constricts the cytoplasm between the daughter

Micronuclei (MNi) are acentric chromosome fragments or whole chromosomes that are left behind during mitotic cellular division and appear in the cytoplasm of interphase cells as small

cells, and several software programs are commercially available [176].

whatever the cause and mechanism of the genotoxicity [174].

undergo genotoxicity testing [174].

38 Current Concepts in Dental Implantology

to complete a study [176].

nuclei during cytokinesis [178].

**Figure 2.** CHO-K1 cells exposed to different treatments. We can observe cells with different quantity of DNA damage obtained from Comet Assay. CHO-K1 cells stained by SYBR green. The cell located more superiorly presents minimal damage (about 5%) and the other cells show higher DNA damage. The longer is the tail of the "comet", the greater is the migration of damaged DNA.

Due to CBMN assay reliability and good reproducibility, it has become one of the standard cytogenetic tests for genetic toxicology tests in human and mammalian cells [180].

The measurement of nucleoplasmic bridges (NPBs), nuclear buds (NBUDs) and MNi of binucleated cells led the development of the concept of the cytokinesis-block micronucleus cytome (CBMN Cyt) assay [180]. The frequency of binucleated cells with MNi, NPBs or NBUDs provides a measure of genome damage and/or chromosomal instability. An NPB is a contin‐ uous DNA-containing structure linking the nuclei in a binucleated cell which originates from dicentric chromosomes (resulting from misrepaired DNA breaks or telomere end fusions) in which the centromeres are pulled to opposite poles during anaphase [180]. NBUDs represent the mechanism by which a nucleus eliminates amplified DNA and DNA repair complexes. They are similar to MNi in appearance with the exception that they are connected with the nucleus by a bridge [180]. Figure 3 shows NPB and NBUD in binucleated cells.

Since no single test has proved to be capable of detecting mammalian mutagens and carcino‐ gens with an acceptable level of precision and reproducibility, a battery of tests is needed (ISO 10993-3).

**Figure 3.** CHO-K1 cells after CBMN assay. We can observe binucleated cells (A, B); a binucleated cell with one micro‐ nucleus (C); a binucleated cell with two micronuclei (D); a binucleated cell with NBUDs (E); a binucleated cell with micronuclei and a NPB between the main nuclei.

#### **5.1. Some biomaterial studies — Cytotoxic, genotoxic, mutagenic assays**

Because of the low biodegradation rates of hydroxyatatite (HA), beta-tricalcium phosphate was added to HA, generating a biphasic calcium phosphate (BCP) composite, which may play an important role during assisted bone regeneration [166]. The authors [166] evaluated the cytocompatibility of dense HA, porous HA, dense BCP and porous BCP by three different cell viability parameters (XTT, Crystal Violet Dye Elution, Neutral Red assay) on human mesen‐ chymal cells. No significant differences on mitochondrial activity (XTT) or cell density (Crystal Violet Dye Elution) were observed among groups. Dense materials induced lower levels of total viable cells by Neutral Red assay. It was concluded that porous BCP has shown better results than dense materials and these ceramics are suited for further studies [166].

Authors [165] evaluated cytotoxic, genotoxic and mutagenic effects of fluor- hydroxyapatite (FHA) and fluorapatite (FA) eluates on Chinese hamster V79 cells and compared them with the effects of hydroxyapatite (HA) eluate. The results showed that the highest test concentra‐ tions of the biomaterials (100% and 75% eluates) induced very weak inhibition of colony growth (about 10%). On the other hand, the reduction of cell number per colony induced by these concentrations was in the range from 43% to 31%. The comet assay showed that bioma‐ terials induced DNA breaks, which increased with increasing test concentrations in the order HA < FHA < FA. None of the biomaterials induced mutagenic effects compared with the positive control; and DNA breakage was probably the reason for the inhibition of cell division in V79 cell colonies.

Calcium phosphate cements are an important class of bone repair materials. Dicalcium phosphate dihydrate (DCPD) cements were prepared using monocalcium phosphate mono‐ hydrate (MCPM) and hydroxyapatite (HA) [170]. Degradation properties and cytocompati‐ bility of this cement were analyzed and compared with β-tricalcium phosphate (β-TCP). The percent of viable cells as well as the percent of necrotic and apoptotic ones were evaluated by flow cytometry-based cell viability/apoptosis assay. According to the results, although conversion to HA has been noted in DCPD cements prepared with β -TCP, the conversion occurred rapidly when HA was used as the base component. HA during cement preparation seemed to accelerate the process and led to a rapid pH drop, extensive mass loss, a complete loss of mechanical integrity, and reduced cytocompatibility [170].

Authors [173] evaluated poloxamines, i.e., X-shaped poly(ethylene oxide)-poly(propylene oxide) block copolymers with an ethylenediamine core (Tetronic®), as an active osteogenic component and as a vehicle for rhBMP-2 injectable implants [173]. After cytotoxicity screening of various poloxamine varieties, Tetronic® 304, 901, 904, 908, 1107, 1301, 1307 and 150R1 and poloxamer Pluronic® F127 were analyzed. Tetronic® 908, 1107, 1301 and 1307 solutions were the most cytocompatible and it was concluded that the intrinsic osteogenic activity of polox‐ amines offers novel perspectives for bone regeneration using minimally invasive procedures (i.e., injectable scaffolds) and overcoming the safety and the cost/effectiveness concerns associated with large scale clinical use of recombinant growth factors [173].

Recombinant human bone morphogenetic protein 2 (rhBMP-2) has been widely employed for the induction of bone growth in animal models and in clinical trials [177]. Authors [177] prepared their own rhBMP-2 and the micronucleus assay was used to evaluate the genotoxic effect of it. It was concluded that author's preparations of recombinant human BMP-2 prepared in E. coli do not promote DNA damage in the concentration range tested.

**5.1. Some biomaterial studies — Cytotoxic, genotoxic, mutagenic assays**

micronuclei and a NPB between the main nuclei.

40 Current Concepts in Dental Implantology

Because of the low biodegradation rates of hydroxyatatite (HA), beta-tricalcium phosphate was added to HA, generating a biphasic calcium phosphate (BCP) composite, which may play an important role during assisted bone regeneration [166]. The authors [166] evaluated the cytocompatibility of dense HA, porous HA, dense BCP and porous BCP by three different cell viability parameters (XTT, Crystal Violet Dye Elution, Neutral Red assay) on human mesen‐ chymal cells. No significant differences on mitochondrial activity (XTT) or cell density (Crystal Violet Dye Elution) were observed among groups. Dense materials induced lower levels of total viable cells by Neutral Red assay. It was concluded that porous BCP has shown better

**Figure 3.** CHO-K1 cells after CBMN assay. We can observe binucleated cells (A, B); a binucleated cell with one micro‐ nucleus (C); a binucleated cell with two micronuclei (D); a binucleated cell with NBUDs (E); a binucleated cell with

Authors [165] evaluated cytotoxic, genotoxic and mutagenic effects of fluor- hydroxyapatite (FHA) and fluorapatite (FA) eluates on Chinese hamster V79 cells and compared them with the effects of hydroxyapatite (HA) eluate. The results showed that the highest test concentra‐ tions of the biomaterials (100% and 75% eluates) induced very weak inhibition of colony growth (about 10%). On the other hand, the reduction of cell number per colony induced by these concentrations was in the range from 43% to 31%. The comet assay showed that bioma‐ terials induced DNA breaks, which increased with increasing test concentrations in the order

results than dense materials and these ceramics are suited for further studies [166].

A fully crystallized bioactive glass–ceramic material (Biosilicate®) for bone repair was developed and the biocompatibility was evaluated by means of histopathological (after subcutaneous test), cytotoxic (MTT) and genotoxic analysis (Comet assay). Neonatal murine calvarial osteoblastic (OSTEO-1) and murine fibroblasts (L929) were employed in this study. The results indicated that Biosilicate® scaffolds was biocompatible and noncytotoxic and did not induce DNA strand breaks at any evaluated period [172].

Polymethyl methacrylate (PMMA) is an acrylic resin which is widely used as a biomaterial due to its excellent biocompatibility and haemocompatibility [181]. *In vitro* micronucleus (MN) induction by PMMA bone cement was analyzed in cultured human lymphocyte [181]. The results showed a highly significant increase in MN frequency in human lymphocytes treated with PMMA and consequently a genotoxic effect of this substance or of the aphorised residual ingredients, which continue to be released in small amounts from the polymer. According to the authors, after the polymerization process, small quantities of ingredients usually present in self-curing methacrylate bone cements are released and their rate of diffusion depends on storage conditions.

Titanium has been one of the most clinically applicable metals in bone tissue to serve as fracture fixation devices and also as endosseous implants for the rehabilitation of various parts of human body, especially in the oral maxillofacial region [182]. Piozzi et al. (2009) [182] evaluated whether liver, kidney, and lung of rats were particularly sensitive organs for DNA damaging (Comet assay) and cytotoxicity (histopathological changes) following implantation of internal fixture materials composed by titanium alloy in rats. No histopathological changes in cells of lung, kidney or liver were observed in the negative control group and in the experimental groups. The liver, lung and kidney cells did not show any genotoxic effects along the time course experiment. In the same way, no cytotoxic effects were present since neither tissue alterations nor signals of metals deposition were evidenced in these organs, even after 180 days of titanium exposure [182].

Metallic implants can release not only biocompatible ions but also some particles from mechanical wear or degradation. After corrosion or mechanical wear, these metal biomaterials release toxic elements such as ions or particles to the environment. Biodegradable metals seem to be the suitable material for orthopedic applications. Screws and plates made of magnesium alloys may work as stable biodegradable implants, which avoids the instance of a second operation. However, despite their use in novel technology, there is no available information about the possible toxic effects of magnesium particles (MP) from wear debris on human health [171]. Authors [171] used Mg powder to simulate the presence of MP wear debris within a cell culture and cytotoxic and genotoxic effects (comet assay and micronucleus induction) were analyzed. Neutral red (NR) incorporation and acridine orange/ethidium bromide (AO/EB) staining techniques were used to analyze the cytotoxic effects at 25–1000 µg/mL concentration range. Changes in lysosome activity were observed after 24 h only at 1000 µg/mL. Accordingly, AO/EB staining showed a significant decrease in the number of living cells at 500 µg/mL. A significant dose-dependent increase in MN frequencies was observed at 25–100 µg/mL range (nontoxic range). DNA damage induction was observed by comet assay only at 500 µg/mL. Therefore, authors verified a dose-dependent cytotoxic and genotoxic effects of MP on UMR106 cells with different threshold values of MP concentration.

#### **6. Summary**

This chapter approaches the most current bone substitute materials used in implant dentistry, as in research as in clinical application, for alveolar ridge augmentation, maxillary sinus lift and guided bone regeneration, such as: alloplastic materials (bioceramics, bioactive glasses, glass-ceramics, polymers and composites) and bioactive molecules (peptides and growth factors). In addition, concepts of tissue engineering used for the development of the new materials and techniques for implant dentistry were approached. Moreover, this chapter approached some cytotoxic, genotoxic and mutagenic assays used to evaluate the safety of biomaterials. Some studies that evaluated cytotoxicity, genotoxicity and/or mutagenicity of biomaterials were presented.

Thus, the use of bone substitutes continues to increase along with the availability of new technologies. Many alternatives for the replacement of autografts, allografts and xenografts are emerging. Rigorous preclinical and clinical studies are necessary to confirm the costeffectiveness of these approaches over traditional bone grafts methods with benefits of technological advancement exceeding risks to the patient and costs of implantation.

#### **Author details**

in self-curing methacrylate bone cements are released and their rate of diffusion depends on

Titanium has been one of the most clinically applicable metals in bone tissue to serve as fracture fixation devices and also as endosseous implants for the rehabilitation of various parts of human body, especially in the oral maxillofacial region [182]. Piozzi et al. (2009) [182] evaluated whether liver, kidney, and lung of rats were particularly sensitive organs for DNA damaging (Comet assay) and cytotoxicity (histopathological changes) following implantation of internal fixture materials composed by titanium alloy in rats. No histopathological changes in cells of lung, kidney or liver were observed in the negative control group and in the experimental groups. The liver, lung and kidney cells did not show any genotoxic effects along the time course experiment. In the same way, no cytotoxic effects were present since neither tissue alterations nor signals of metals deposition were evidenced in these organs, even after 180 days

Metallic implants can release not only biocompatible ions but also some particles from mechanical wear or degradation. After corrosion or mechanical wear, these metal biomaterials release toxic elements such as ions or particles to the environment. Biodegradable metals seem to be the suitable material for orthopedic applications. Screws and plates made of magnesium alloys may work as stable biodegradable implants, which avoids the instance of a second operation. However, despite their use in novel technology, there is no available information about the possible toxic effects of magnesium particles (MP) from wear debris on human health [171]. Authors [171] used Mg powder to simulate the presence of MP wear debris within a cell culture and cytotoxic and genotoxic effects (comet assay and micronucleus induction) were analyzed. Neutral red (NR) incorporation and acridine orange/ethidium bromide (AO/EB) staining techniques were used to analyze the cytotoxic effects at 25–1000 µg/mL concentration range. Changes in lysosome activity were observed after 24 h only at 1000 µg/mL. Accordingly, AO/EB staining showed a significant decrease in the number of living cells at 500 µg/mL. A significant dose-dependent increase in MN frequencies was observed at 25–100 µg/mL range (nontoxic range). DNA damage induction was observed by comet assay only at 500 µg/mL. Therefore, authors verified a dose-dependent cytotoxic and genotoxic effects of MP on

This chapter approaches the most current bone substitute materials used in implant dentistry, as in research as in clinical application, for alveolar ridge augmentation, maxillary sinus lift and guided bone regeneration, such as: alloplastic materials (bioceramics, bioactive glasses, glass-ceramics, polymers and composites) and bioactive molecules (peptides and growth factors). In addition, concepts of tissue engineering used for the development of the new materials and techniques for implant dentistry were approached. Moreover, this chapter approached some cytotoxic, genotoxic and mutagenic assays used to evaluate the safety of biomaterials. Some studies that evaluated cytotoxicity, genotoxicity and/or mutagenicity of

UMR106 cells with different threshold values of MP concentration.

storage conditions.

42 Current Concepts in Dental Implantology

of titanium exposure [182].

**6. Summary**

biomaterials were presented.

Sybele Saska1\*, Larissa Souza Mendes1 , Ana Maria Minarelli Gaspar2 and Ticiana Sidorenko de Oliveira Capote2

\*Address all correspondence to: sybele\_saska@yahoo.com.br

1 Institute of Chemistry, São Paulo State University-UNESP, Araraquara – SP, Brazil

2 Dental School at Araraquara, São Paulo State University-UNESP, Department of Morphology, Araraquara – SP, Brazil

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### **Biotechnology of Tissues and Materials in Dentistry — Future Prospects**

Andréa Cristina Barbosa da Silva, Diego Romário da Silva, Rafael Grotta Grempel, Manuel Antonio Gordón-Núñez and Gustavo Gomes Agripino

Additional information is available at the end of the chapter

http://dx.doi.org/10.5772/59644

**1. Introduction**

Long ago, humanity has sought alternatives to replacing living tissue, mainly due to birth defects, disease and accidents, using synthetic or natural substances as substitutes, best known as biomaterials. Thus, tissue engineering has emerged, a new and challenging field of modern medicine, which aims at recreating tissues and/or healthy organs to replace missing or diseased body parts [1].

Regenerative medicine which used medical devices and grafts underwent some changes in recent years, changing to a more biological approach, with use of specific biodegradable bioactive and supports (scaffolds) with cells and / or biological molecules to create a functional tissue repair in a diseased or damaged site. Thus, some newer and inter-related strategies are being used for the regeneration of tissues such as cell injection, cell induction and cells seeded in scaffolds (cell seeded scaffold) (detailed later in this chapter) [2]. These approaches depend on the use of one or more key elements, such as cells, growth factors and matrix for guiding tissue regeneration [3].

The technique used to obtain tissues (tissue engineering) is the regeneration of organs and living tissues, through recruitment of the patient's own tissue, which are dissociated into cells and cultured on synthetic or biological carriers, known as scaffolds (scaffolds, three-dimen‐ sional matrices, structures, etc.) and then being reinserted into the patient. As a multidiscipli‐ nary science, the work involves knowledge of the areas of biology, health sciences and engineering and materials science [4, 5].

© 2015 The Author(s). Licensee InTech. This chapter is distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/3.0), which permits unrestricted use, distribution, and eproduction in any medium, provided the original work is properly cited.

Thus, one important step for reconstruction of an organ or tissue is the scaffold selection to the cells, which must take into consideration the type, location and extent of injury. The scaffold structure provides mechanical support to the cell growth and allows transport of nutrients, metabolites, growth factors, and other regulatory molecules, both towards the extracellular environment to the cells, as in the opposite direction [6]. When prepared with bioresorbable polymeris, scaffolds, the scaffolds have specific implementation strategies [7].

After a degradable polymer is identified as a possible candidate for applications in tissue engineering, it must be used for manufacturing a porous scaffold [8, 9, 10, 11]. In this case, two methods are required for proper material manufacture: 1) a method that forms the polymer into a bulk material; 2) a method to make porous such material [12]. The optimal method of manufacturing depends in part on the chemical nature of the polymer. Long, saturated and linear polymers such as PLG are typically formed into bulk materials by entangling the individual polymer chains to form a loosely bound polymer network. Polymer chain entan‐ glement is often achieved by casting the polymer within a mold. The advantage to these methods is that they are relatively simple. However, since the material is elastic solid only because of entangled polymer chains, the material is generally lacking significant mechanical strength. This disadvantage is difficult to overcome without altering the chemical structure of the polymer [12].

Another method to form a bulk material from a linear polymer involves forming chemical bonds between polymer chains, known as polymer cross linking [13, 14]. Cross linking is most often performed between unsaturated carbon-carbon double bonds, and thus this moiety, or a similarly reactive one, is required to exist on somewhere along the polymer chain. An initiation system, typically either radical or ionic, is also needed to promote cross-linking. The initiator system is combined with the polymer and, in response to a signal such as heat, light, a chemical accelerant, or simply time, the initiator forms species that propagate cross-linking. As these polymers are formed into bulk materials by covalent cross-linking, they typically posses significant mechanical strength. Furthermore, their ability to cure in response to an applied signal allows these materials to be injected into the defect site and cure in situ. The major disadvantage of crosslinked materials is that the growing complexity of the material, in terms of the number of components and presence of a chemical reaction, often leads to problems with cytotoxicity and biocompatibility [12].

In this context, biomaterials are extremely important for tissue regeneration process, and can be defined as any substance constructed in such a way that, alone or as part of a complex system, is used for driving, through the control of interactions with components a living system, the course of a diagnostic or therapeutic procedure, whether in humans or animals [15].

In recent decades, biomaterials have been used to repair tissue function, such as metal implants, without concern for its effect on local tissues or on the cells. Thus, polymers and other synthetic materials with biological properties were then developed. More recently, degradable and natural scaffolds, considered a breakthrough for regenerative medicine have been used. Thus, there was an evolution of the use of biomaterials that simply replaced the damaged tissue, to others more specific, allowing the development in three dimensions of a tissue regenerated in full operation and structurally acceptable [2].

To use a material with the purpose of replacing a part of the body or induce the formation of a given tissue, a range of tests and assessments are necessary to establish the potential benefits and possible adverse effects that the material may have. Thus, biomaterials should have the following characteristics: not inducing thrombus formation as a result of contact between the blood and the biomaterial, not inducing adverse immune response, not being toxic or carci‐ nogenic, not disturbing the blood flow, and not producing chronic or acute inflammatory response that prevents the proper differentiation of adjacent tissues [16].

In other words, the biomaterial must be fully biocompatible, that is, must have the ability to perform its desired function with respect to a medical therapy without inducing any undesir‐ able local or systemic effect to the body; but generating cellular and tissue responses beneficial in that specific situation, and optimizing the clinically relevant responses of that therapy [15]. However, it is worth noting that despite the material having been considered inert for a considerable time, it was suggested that they may induce physical and chemical changes after deployment. Thus, before a biological perspective, no material can be considered in fact inert.

#### **2. Strategies for formation and development of tissues**

Thus, one important step for reconstruction of an organ or tissue is the scaffold selection to the cells, which must take into consideration the type, location and extent of injury. The scaffold structure provides mechanical support to the cell growth and allows transport of nutrients, metabolites, growth factors, and other regulatory molecules, both towards the extracellular environment to the cells, as in the opposite direction [6]. When prepared with bioresorbable

After a degradable polymer is identified as a possible candidate for applications in tissue engineering, it must be used for manufacturing a porous scaffold [8, 9, 10, 11]. In this case, two methods are required for proper material manufacture: 1) a method that forms the polymer into a bulk material; 2) a method to make porous such material [12]. The optimal method of manufacturing depends in part on the chemical nature of the polymer. Long, saturated and linear polymers such as PLG are typically formed into bulk materials by entangling the individual polymer chains to form a loosely bound polymer network. Polymer chain entan‐ glement is often achieved by casting the polymer within a mold. The advantage to these methods is that they are relatively simple. However, since the material is elastic solid only because of entangled polymer chains, the material is generally lacking significant mechanical strength. This disadvantage is difficult to overcome without altering the chemical structure of

Another method to form a bulk material from a linear polymer involves forming chemical bonds between polymer chains, known as polymer cross linking [13, 14]. Cross linking is most often performed between unsaturated carbon-carbon double bonds, and thus this moiety, or a similarly reactive one, is required to exist on somewhere along the polymer chain. An initiation system, typically either radical or ionic, is also needed to promote cross-linking. The initiator system is combined with the polymer and, in response to a signal such as heat, light, a chemical accelerant, or simply time, the initiator forms species that propagate cross-linking. As these polymers are formed into bulk materials by covalent cross-linking, they typically posses significant mechanical strength. Furthermore, their ability to cure in response to an applied signal allows these materials to be injected into the defect site and cure in situ. The major disadvantage of crosslinked materials is that the growing complexity of the material, in terms of the number of components and presence of a chemical reaction, often leads to

In this context, biomaterials are extremely important for tissue regeneration process, and can be defined as any substance constructed in such a way that, alone or as part of a complex system, is used for driving, through the control of interactions with components a living system, the course of a diagnostic or therapeutic procedure, whether in humans or animals [15]. In recent decades, biomaterials have been used to repair tissue function, such as metal implants, without concern for its effect on local tissues or on the cells. Thus, polymers and other synthetic materials with biological properties were then developed. More recently, degradable and natural scaffolds, considered a breakthrough for regenerative medicine have been used. Thus, there was an evolution of the use of biomaterials that simply replaced the damaged tissue, to others more specific, allowing the development in three dimensions of a

polymeris, scaffolds, the scaffolds have specific implementation strategies [7].

the polymer [12].

60 Current Concepts in Dental Implantology

problems with cytotoxicity and biocompatibility [12].

tissue regenerated in full operation and structurally acceptable [2].

The strategies employed for tissue engineering can be classified into three main classes: conductive or inductive approaches and cell transplantation.

The conductor/conductive approaches using biomaterials in a passive manner to facilitate the growth or regeneration capacity of existing tissue such as, for example, use of membranes or barriers for applied regeneration, adhesion molecules, growth factors, etc. in cases of perio‐ dontal diseases [1, 17, 18] or dental implant itself, which is a relatively simple implementation because the apparatus used does not include the use of living cells or other diffusible biological signals [19]. In the conductive techniques is usually accomplished the neoformation of periodontal complex structures, including cementum and periodontal ligament fibers [1]. The periodontium regeneration is the first engineering technology for dental tissue [17].

In 1965, Urist [20] demonstrated for the first time that the new bone formation could occur in a non-mineralized site after implantation of powder bone. This discovery led to the isolation of the active ingredients (specific growth factors - proteins) from bone powder, and the cloning of the genes encoding these proteins. These concepts have been used by many companies for production and expansion of these factors on a large scale [21]. Another method employed is the induction type or inductive approach, which involves the activation of cells near the defect site with specific biological signals that stimulate proliferation and assist in regeneration and repair of tissues by use of materials such bone morphogenetic proteins (BMPs) [20, 22] with promising results for supplementation therapies and the regeneration and bone repair in cases of fractures and periodontal disease [1].

In other words, an alternative approach is the use of diffusible growth factors, and consists of placing specific extracellular matrix molecules on a scaffold to allow the tissue growth. These molecules have the ability to direct or induce the function of cells already present in this location, and in consequence, promoting the formation of a tissue type or a particular desired structure at the location [23].

For the tissue induction can be clinically successful, it is necessary that the biologically active factors are delivered properly to the desired location and in the correct dose for the time period necessary. Typically, many such proteins have a short half-life in the body, but must be present for a long time to be effective. Doctors and researchers have shown these concerns so far by offering large doses of protein at the sites of interest [19]. The most recent research involves the development of a controlled release system of these proteins (inducing factors) [24] and, with the advent of genetic engineering in current biotechnology, a somewhat similar approach involves transfection of a gene encoding the inducing factor, instead of delivering the protein itself [19].

Cell transplantation is the third method, which consists of the direct transplantation of cells grown in the laboratory [25]. This approach is a strategy whose importance is based on the need for a multidisciplinary team for performing tissue engineering, since it requires the physician or surgeon in charge of obtaining tissue samples by biopsy, the bioengineer, who usually participates in manipulating the tissues in bioreactors and prepares the means necessary for placing the cells obtained from biopsy samples, besides cell biologist, who will apply the principles of cell biology required for multiplication and maintenance of cells in the laboratory [1, 18, 26, 27].

Despite having different mechanisms, the three strategies for tissue formation have one characteristic in common: the use of polymeric materials. In conducting approaches, polymer is mainly used as a membrane barrier for exclusion of particular cells that can disturb the regenerative process. In the inductive approaches, these materials act as a carrier for delivery of proteins (e.g., BMP) or the DNA encoding the protein [24, 28]. With regard to approaches used to achieve control of the dose and bioavailability of biodegradable polymer carriers enable localized and sustained release of inductive molecules. The dose rate and the molecule to be delivered are controlled generally by gradual breakdown of the vehicle [24].

These delivery vehicles are often used in cell transplantation approaches. However, in this approach the vehicle serves as a carrier of intact cells and even partial tissues [1].

Besides acting as vehicles for the simple delivery of cells, the vehicles also serve as scaffolds to guide new tissue to grow in a predictable way from the interaction between cells or transplanted tissue and host cells. The collagen derived from animal sources, and synthetic polymers of lactic acid and glycolic acid are the main absorbable materials used for tissue repair in three types of approaches. The collagen is degraded by cells in the tissue during its development, whereas the synthetic polymers are degraded into natural metabolites of lactic acid and glycolic acid by the water action at the implant site. From the development and innovation of biotechnology in tissue engineering various new materials are also being developed for these applications, such as injectable materials that enable a minimally invasive delivery of inductive molecules or transplanted cells [1].

Below (Figure 1), a schematic view of the three types of approaches in tissue engineering:

location, and in consequence, promoting the formation of a tissue type or a particular desired

For the tissue induction can be clinically successful, it is necessary that the biologically active factors are delivered properly to the desired location and in the correct dose for the time period necessary. Typically, many such proteins have a short half-life in the body, but must be present for a long time to be effective. Doctors and researchers have shown these concerns so far by offering large doses of protein at the sites of interest [19]. The most recent research involves the development of a controlled release system of these proteins (inducing factors) [24] and, with the advent of genetic engineering in current biotechnology, a somewhat similar approach involves transfection of a gene encoding the inducing factor, instead of delivering the protein

Cell transplantation is the third method, which consists of the direct transplantation of cells grown in the laboratory [25]. This approach is a strategy whose importance is based on the need for a multidisciplinary team for performing tissue engineering, since it requires the physician or surgeon in charge of obtaining tissue samples by biopsy, the bioengineer, who usually participates in manipulating the tissues in bioreactors and prepares the means necessary for placing the cells obtained from biopsy samples, besides cell biologist, who will apply the principles of cell biology required for multiplication and maintenance of cells in the

Despite having different mechanisms, the three strategies for tissue formation have one characteristic in common: the use of polymeric materials. In conducting approaches, polymer is mainly used as a membrane barrier for exclusion of particular cells that can disturb the regenerative process. In the inductive approaches, these materials act as a carrier for delivery of proteins (e.g., BMP) or the DNA encoding the protein [24, 28]. With regard to approaches used to achieve control of the dose and bioavailability of biodegradable polymer carriers enable localized and sustained release of inductive molecules. The dose rate and the molecule

These delivery vehicles are often used in cell transplantation approaches. However, in this

Besides acting as vehicles for the simple delivery of cells, the vehicles also serve as scaffolds to guide new tissue to grow in a predictable way from the interaction between cells or transplanted tissue and host cells. The collagen derived from animal sources, and synthetic polymers of lactic acid and glycolic acid are the main absorbable materials used for tissue repair in three types of approaches. The collagen is degraded by cells in the tissue during its development, whereas the synthetic polymers are degraded into natural metabolites of lactic acid and glycolic acid by the water action at the implant site. From the development and innovation of biotechnology in tissue engineering various new materials are also being developed for these applications, such as injectable materials that enable a minimally invasive

Below (Figure 1), a schematic view of the three types of approaches in tissue engineering:

to be delivered are controlled generally by gradual breakdown of the vehicle [24].

approach the vehicle serves as a carrier of intact cells and even partial tissues [1].

delivery of inductive molecules or transplanted cells [1].

structure at the location [23].

62 Current Concepts in Dental Implantology

itself [19].

laboratory [1, 18, 26, 27].

 **Figure 1** Schematic representation of the three main approaches for tissue rebuilding in tissue engineering in jaw: I) by the conductive method where use is made of a barrier that is able to exclude connective tissue cells that may interfere with the regeneration process and at the same time enables the desired host cells to populate the site to be regenerated. **II)** by the inductive method, in which a scaffold of the biodegradable polymer is used as a delivery vehicle for growth factors and / or genes encoding this factor in the desired location. As the polymer is being degraded, the growth factor is being released gradually. **III)** by the strategy of cell transplantation, which uses a delivery vehicle, similar to that used in an inductive approach, with the goal of transplanting cells and partial tissues to the place where we want to regenerate tissue. In this approach can be transplanted only tissues or cells previously formed in the laboratory from scaffolds. **Figure 1.** Schematic representation of the three main approaches for tissue rebuilding in tissue engineering in jaw: I) by the conductive method where use is made of a barrier that is able to exclude connective tissue cells that may interfere with the regeneration process and at the same time enables the desired host cells to populate the site to be regenerated. **II)** by the inductive method, in which a scaffold of the biodegradable polymer is used as a delivery vehicle for growth factors and / or genes encoding this factor in the desired location. As the polymer is being degraded, the growth factor is being released gradually. **III)** by the strategy of cell transplantation, which uses a delivery vehicle, similar to that used in an inductive approach, with the goal of transplanting cells and partial tissues to the place where we want to regenerate tissue. In this approach can be transplanted only tissues or cells previously formed in the laboratory from scaffolds.

important approach in the engineering scope for bone tissue formation [1] (Figure 2).

Tissue engineering seeks solutions for the regeneration of various tissues associated with the oral cavity, such as, bones, cartilage, skin and oral mucosa, dentin and dental pulp, and salivary glands. But in fact, this science will probably have its most significant impact in dentistry through bone reconstruction and regeneration. The fact that cell transplantation approaches may offer the possibility of pre-formation of bone structures of large dimensions (for example, full jaw), which may not be possible to use the other two strategies, makes it the most Tissue engineering seeks solutions for the regeneration of various tissues associated with the oral cavity, such as, bones, cartilage, skin and oral mucosa, dentin and dental pulp, and salivary glands. But in fact, this science will probably have its most significant impact in dentistry through bone reconstruction and regeneration. The fact that cell transplantation approaches may offer the possibility of pre-formation of bone structures of large dimensions (for example, full jaw), which may not be possible to use the other two strategies, makes it the most important approach in the engineering scope for bone tissue formation [1] (Figure 2).

**Figure 2.** Schematic representation of the advances in tissue engineering to regenerate part of the jaw by means of cell transplantation. A scaffold consisting of biodegradable polymer in the shape of half of the jaw is built (I). Thereafter, bone precursor cells are seeded on the polymer (beige dots) and stimulated to grow in a bioreactor (II). The scaffold will then be gradually degraded, while facilitating growth of jaw-shaped bone (III) (Scheme adapted from [1].

Thus, the tissue repair from the in vitro tissue engineering requires the use of cells to comple‐ tion and production of similar matrix to the native tissue. The main successful developments in this field have been using the transplant of primary cells taken from patient and used in combination with scaffolds to produce the required tissue to re-implant. However, this strategy has limitations due to the invasive nature of how the cells are removed. Thus, attention has turned to the use of stem cells, including embryonic stem cells and mesenchymal cells derived from bone marrow. In addition to being able to turn into all body tissues, these cells have the capability and advantage of being maintained in culture for long periods, thus having the potential to obtaining large amounts of cells to tissue. The extraordinary ability of these pluripotent cells is linked to their ability to form teratoma [29]. Besides the potential to differentiate into osteoblasts, the possibility of rejection of these cells is greatly reduced.

In cell transplantation, these units can be directly transplanted to the desired location or they may be cultured in the laboratory on scaffolding. In this case, those cells are stimulated to lay the groundwork matrix to produce a tissue for transplantation [29].

Currently, several products can be used to achieve tissue regeneration or reconstruction. These options are divided according to the approach to be used (Inducing, conductive or cell transplantation) as shown in the scheme below (Figure 3) adapted from [19].

Tissue engineering seeks solutions for the regeneration of various tissues associated with the oral cavity, such as, bones, cartilage, skin and oral mucosa, dentin and dental pulp, and salivary glands. But in fact, this science will probably have its most significant impact in dentistry through bone reconstruction and regeneration. The fact that cell transplantation approaches may offer the possibility of pre-formation of bone structures of large dimensions (for example, full jaw), which may not be possible to use the other two strategies, makes it the most important

**Figure 2.** Schematic representation of the advances in tissue engineering to regenerate part of the jaw by means of cell transplantation. A scaffold consisting of biodegradable polymer in the shape of half of the jaw is built (I). Thereafter, bone precursor cells are seeded on the polymer (beige dots) and stimulated to grow in a bioreactor (II). The scaffold will then be gradually degraded, while facilitating growth of jaw-shaped bone (III) (Scheme adapted from [1].

Thus, the tissue repair from the in vitro tissue engineering requires the use of cells to comple‐ tion and production of similar matrix to the native tissue. The main successful developments in this field have been using the transplant of primary cells taken from patient and used in combination with scaffolds to produce the required tissue to re-implant. However, this strategy has limitations due to the invasive nature of how the cells are removed. Thus, attention has turned to the use of stem cells, including embryonic stem cells and mesenchymal cells derived from bone marrow. In addition to being able to turn into all body tissues, these cells have the capability and advantage of being maintained in culture for long periods, thus having the potential to obtaining large amounts of cells to tissue. The extraordinary ability of these pluripotent cells is linked to their ability to form teratoma [29]. Besides the potential to differentiate into osteoblasts, the possibility of rejection of these cells is greatly reduced.

In cell transplantation, these units can be directly transplanted to the desired location or they may be cultured in the laboratory on scaffolding. In this case, those cells are stimulated to lay

Currently, several products can be used to achieve tissue regeneration or reconstruction. These options are divided according to the approach to be used (Inducing, conductive or cell

the groundwork matrix to produce a tissue for transplantation [29].

transplantation) as shown in the scheme below (Figure 3) adapted from [19].

approach in the engineering scope for bone tissue formation [1] (Figure 2).

64 Current Concepts in Dental Implantology

**Figure 3.** Products used for bone tissue repair in different types of approaches (Inducing, or conductive cell transplan‐ tation) (Adapted from de Kumar, Mukhtar-Un-Nisar and Zia, 2011) [19].

#### **3. Importance of tissues for maxillofacial complex**

The maxillofacial complex can be subjected to processes of physical, chemical and biological nature, which usually determine from minor tissue losses to the involvement of large areas of structures of this complex. In this context, dentistry has been explored new technologies in order to change this reality, adapting to new concepts, scientific innovations that include research on stem cells, tissue engineering, and molecular biology techniques, as tools to stimulate regeneration or replacement of damaged tissue by tissue engineering.

Considering the scenario of new technologies, however, still in 2001 it was asked: "What impact could have this engineering in dentistry?" And "What maxillofacial tissues have potential or are important for that engineering?" According to Kaigler and Mooney (2001) [1], at that time the answer to the first question was still being formulated, since the engineering probably would have a revolutionary effect on the field of Dentistry, once almost all types of tissues in the maxillofacial complex could have potential for engineering. Currently, reality has changed significantly due to which the tissue engineering has wide application to many different tissue types associated with the oral cavity, including bone, cartilage, skin, oral mucosa, dentin and dental pulp, and salivary glands.

As previously mentioned, inductive, conductive and cell transplantation strategies, which represent the most used techniques in tissue engineering, are of importance to typically use different material components in order to achieve the goal of regeneration and / or replacement of damaged tissues.

Absolutely, all tissues of the maxillofacial complex are important for its proper functioning, playing a crucial role also in facial aesthetics. Thus, some comments are required about the major oral tissues and their importance for tissue engineering.

With respect to bone, it can be said that tissue engineering has had a greater impact in dentistry, particularly with regard to bone regeneration. Bone loss associated with trauma, diseases or disorders can currently be handled through the use of biomaterials for auto-grafts, allografts or synthetic, morphogenetic proteins (BMPs) and growth factors. It is reported that even though these biomaterials stimulate, replace and / or restore the stability and function of tissues in a reasonably sufficient manner, there are still limitations in their use, which is of importance for research is increasingly carried out using the three main strategies of tissue engineering in order to optimize the mechanisms of regeneration in bone areas compromised by various damaging agents [1, 28, 30].

The importance of cartilage tissue to tissue engineering of structures of the maxillofacial complex lies in the possibility of reconstruction of craniofacial chondromatosous structures, the design of polymeric structures with defined mechanical and degradative properties that can serve as a support structures for cartilage cell proliferation of temporomandibular or intranasal joints if compromised by trauma or degenerative diseases. One of the limitations of the use of cartilage tissue in tissue engineering is due to its limited capacity for regeneration and lack of inductive molecules to the proliferation of their cells; thus it is one of the tissues of great interest among researchers to develop envisaging bioengineering techniques for transplanting of cartilage cells [1, 31, 32].

Researches have been and continue to be focused on the production of dentin and dental pulp by the use of tissue engineering strategies. The importance of these tissues for this engineering is associated with the possibility to replace material lost by carious processes. There is evidence that odontoblasts, even lost due to caries, it would be possible to induce the formation of new pulp tissue cells by tissue engineering based on the use of certain biomolecules stimulating or inducing odontoblast proliferation and / or nerve cells, and these new odontoblasts, in turn, could synthesize new dentin material. Furthermore, it is suggested that the tissue engineering of the dental pulp itself may be possible by using techniques of cultured fibroblasts in synthetic polymer matrices [33, 34, 35, 36, 37].

One of the most exploited tissues in research of tissue engineering in dentistry is the epithelial lining of the oral mucosa with significant advances in the use of these tissues in regeneration and / or replacement of structures of the oral mucosa damaged by various aggressors. Recently, the introduction of 3D reconstruction of the oral mucosa has significantly impacted the approaches to biocompatibility evaluation of tissues and materials to replace and / or regen‐ erate oral soft tissues [2, 38, 39, 40].

types associated with the oral cavity, including bone, cartilage, skin, oral mucosa, dentin and

As previously mentioned, inductive, conductive and cell transplantation strategies, which represent the most used techniques in tissue engineering, are of importance to typically use different material components in order to achieve the goal of regeneration and / or replacement

Absolutely, all tissues of the maxillofacial complex are important for its proper functioning, playing a crucial role also in facial aesthetics. Thus, some comments are required about the

With respect to bone, it can be said that tissue engineering has had a greater impact in dentistry, particularly with regard to bone regeneration. Bone loss associated with trauma, diseases or disorders can currently be handled through the use of biomaterials for auto-grafts, allografts or synthetic, morphogenetic proteins (BMPs) and growth factors. It is reported that even though these biomaterials stimulate, replace and / or restore the stability and function of tissues in a reasonably sufficient manner, there are still limitations in their use, which is of importance for research is increasingly carried out using the three main strategies of tissue engineering in order to optimize the mechanisms of regeneration in bone areas compromised by various

The importance of cartilage tissue to tissue engineering of structures of the maxillofacial complex lies in the possibility of reconstruction of craniofacial chondromatosous structures, the design of polymeric structures with defined mechanical and degradative properties that can serve as a support structures for cartilage cell proliferation of temporomandibular or intranasal joints if compromised by trauma or degenerative diseases. One of the limitations of the use of cartilage tissue in tissue engineering is due to its limited capacity for regeneration and lack of inductive molecules to the proliferation of their cells; thus it is one of the tissues of great interest among researchers to develop envisaging bioengineering techniques for

Researches have been and continue to be focused on the production of dentin and dental pulp by the use of tissue engineering strategies. The importance of these tissues for this engineering is associated with the possibility to replace material lost by carious processes. There is evidence that odontoblasts, even lost due to caries, it would be possible to induce the formation of new pulp tissue cells by tissue engineering based on the use of certain biomolecules stimulating or inducing odontoblast proliferation and / or nerve cells, and these new odontoblasts, in turn, could synthesize new dentin material. Furthermore, it is suggested that the tissue engineering of the dental pulp itself may be possible by using techniques of cultured fibroblasts in synthetic

One of the most exploited tissues in research of tissue engineering in dentistry is the epithelial lining of the oral mucosa with significant advances in the use of these tissues in regeneration and / or replacement of structures of the oral mucosa damaged by various aggressors. Recently, the introduction of 3D reconstruction of the oral mucosa has significantly impacted the

major oral tissues and their importance for tissue engineering.

dental pulp, and salivary glands.

66 Current Concepts in Dental Implantology

of damaged tissues.

damaging agents [1, 28, 30].

transplanting of cartilage cells [1, 31, 32].

polymer matrices [33, 34, 35, 36, 37].

One of the most challenging areas of genetic engineering applied to the structures of the maxillofacial complex is the replace of function of salivary glands, since these tissues play important roles in mastication, phonation and protection of hard and soft tissues of the mouth by saliva production. In this context, we study the possibility of salivary gland cells trans‐ plantation or creating a replacement for compromised glandular structures through the use of artificial salivary glands consisting of a polymer tube coated with salivary epithelial cells [41]. The success importance of future tissue engineering for these tissues might represent the possibility of new and more effective approaches to the treatment of conditions associated with loss of function of the salivary glands, including dysphagia, dysgeusia, rampant caries and mucosal infections [1].

Regarding the possibility of reproducing teeth, there are numerous growth factors involved in the development of dental organs and biological processes involved in odontogenesis are quite complex, reason why we still cannot form a complete tooth; however, some studies have shown the enamel and dentin formation from stem cells isolated from dental pulp [42, 43]. The replacement of missing teeth by tissue engineering in humans is still being researched, but with a real possibility of application in the future.

#### **4. Biomaterials used in craniofacial tissue regeneration**

Biomaterials play a crucial role in tissue engineering. They are used for the manufacture of supports or matrices which allow a suitable microenvironment for optimal cell regeneration.

Biomaterials for constructing scaffolds can be natural/synthetic and rigid/non rigid. Natural biomaterials offer good cellular compatibility i.e. ability to support cell survival and function thereby enhancing the cells' performance, and biocompatibility. Their disadvantages include source variability, immunogenicity, if not pure, limited range of mechanical properties and lack of control over pore size. Unlike natural biomaterials, synthetic biomaterials can be manufactured in unlimited supply under controlled conditions, are cheaper and can be tailored to obtain desired shape, cell differentiation properties and mechanical and chemical properties especially the strength, pore characteristics and degradation rate suited for intended applications. However, synthetic biomaterials lack cell adhesion sites and require chemical modifications to improve cell adhesion

During the last century, various natural or synthetic biomaterials have been used for the manufacture of supports for tissue engineering (fabrication of tissue engineering scaffolds) such as metals, ceramics and polymers. However, metals and ceramics are not biodegradable and its processing is limited, which prevents their application as effective supports (scaffolds) for tissue regeneration. Thus, the polymers has been the most commonly used because they have some important characteristics for tissue regeneration such as biodegradability, porosity, large surface area and ease of processing, among others [44, 45].

There are two types of polymers: natural and synthetic [46, 47]. The main biodegradable synthetic polymers include polyesters, polyanhydride, polyfumarate, polycaprolactone, polycarbonate and polyorthoester [7, 48]. The polyesters such as poly (glycolic acid) (PGA), poly (lactic acid) (PLA), and their copolymer of poly [lactic-co-(glycolic acid)] (PLGA) are most commonly used for tissue engineering. The natural polymers include proteins of natural extracellular matrices such as glycosaminoglycan, collagen, alginic acid and chitosan etc [49, 50]. These polymers of natural origin are biodegradable and possess known cell-binding sites. However, they have some disadvantages such as the level of immunogenicity and speed of degradation.

The tissue regeneration from cells transplanted into a polymer scaffold is summarized in Figure 4.

**Figure 4.** Schematic figure illustrating the steps performed in the laboratory for tissue regeneration from the use of trans‐ planted cells stimulated to grow on biomaterials. It is necessary to understand the importance of biomaterial to perform this technique. It can be natural or synthetic and should meet the requirements of biocompatibility and other features al‐ ready mentioned in this chapter. It is also important to realize the multidisciplinarity involved in this process. The physi‐ cian is needed in order to perform the tissue biopsy to remove the cells **(I).** This tissue/cell is then taken to the laboratory to be multiplied several times. Thereafter, the use of principles of cell biology, such as growth factors **(II)** to stimulate the cells to grow and maintain their functions will be necessary. It is also required the involvement of engineers for manufac‐ turing matrices of biodegradable polymers **(III)** and the bioreactor **(IV).** When cells grow in appropriate number, they are seeded on the polymer scaffold. The tissue is then allowed to grow in the bioreactor until the time of transplantation by clinical surgeon. Biomaterials can be used to stimulate the growth of several types of tissues, e.g. bone, cartilage or skin. After the appropriate development, the tissue is transplanted and the area is regenerated.

Other extracellular matrices used as scaffolds include fibrin and fibrinogen. [51, 52, 53]. According to some studies, both can induce angiogenesis during tissue regeneration [54, 55, 56]. Chitosan is a derivative of chitin, a natural biopolymer which is biocompatible, biode‐ gradable, antimicrobial and possesses tissue healing and osteoinductive effects. It has the ability to bind to growth factors, glycosaminoglycans and DNA and can be easily processed into membranes, gels, nanofibres, beads, scaffolds and sponges. Because of these properties, chitosan gel alone or in combination with demineralized bone matrix/collagenous membrane is quite promising in periodontal regeneration [57].

Considering the bone tissue engineering, porous scaffolds are designed to support the migration, proliferation, and differentiation of osteo-progenitor cells and aid in the organiza‐ tion of these cells in three dimensions. These scaffolds may be made from a wide variety of both natural and synthetic materials. The naturally derived materials include cornstarch-based polymers, [58] chitosan [59, 60] collagen, [61] and coral [62, 63]. Among these materials, the coral has been shown to be an effective clinical alternative to autogenous and allogenous bone grafts [64, 65].

Examples of synthetic materials include calcium phosphates [66, 67] and organic materi‐ als such as poly (phosphazenes), [68] poly (tyrosine carbonates), [69] poly (caprolactones) [70], poly (propylene fumarates) [71], and poly (α-hydroxy acids) [72, 73]. Composites of inorganic and organic materials have also been successfully used to create scaffolds for bone grafts [74, 75]. Poly (α-hydroxy acids) are the most commonly used polymeric materials for the creation of tissue-engineering scaffolds for bone. The most common of the poly (α-hydroxy acids) are poly (glycolic acid), poly (lactic acid) (PLA), and copolymers of poly (lactic-co-glycolic acid) (PLGA). These materials are readily metabolized and excret‐ ed when degraded by the body [44].

#### **5. Challenges and future prospects**

Tissue engineering is an emerging technology with potential application in various medical fields. The main focus of recent research is the development of techniques for manipulating stem cells, aiming at the achievement of restorative treatments of injured and/or lost tissues and organs. Apart from stem cells, bioengineering requires the presence of factors that allow their proliferation in a microenvironment closer to tissue reality, including the extracellular matrix and growth factors. The biomaterials, in turn, are necessary for serving as porous scaffold upon which tissue regeneration is set. As knowledge is acquired with respect to stem cells and biomaterials, the potential for treating diseases may extend beyond the craniofacial region of the body. However, the mechanisms of action of these biotechnologies are not yet fully understood and offer a promising future, so that research is needed to apply them clinically.

#### **Author details**

There are two types of polymers: natural and synthetic [46, 47]. The main biodegradable synthetic polymers include polyesters, polyanhydride, polyfumarate, polycaprolactone, polycarbonate and polyorthoester [7, 48]. The polyesters such as poly (glycolic acid) (PGA), poly (lactic acid) (PLA), and their copolymer of poly [lactic-co-(glycolic acid)] (PLGA) are most commonly used for tissue engineering. The natural polymers include proteins of natural extracellular matrices such as glycosaminoglycan, collagen, alginic acid and chitosan etc [49, 50]. These polymers of natural origin are biodegradable and possess known cell-binding sites. However, they have some disadvantages such as the level of immunogenicity and speed of

The tissue regeneration from cells transplanted into a polymer scaffold is summarized in

**Figure 4.** Schematic figure illustrating the steps performed in the laboratory for tissue regeneration from the use of trans‐ planted cells stimulated to grow on biomaterials. It is necessary to understand the importance of biomaterial to perform this technique. It can be natural or synthetic and should meet the requirements of biocompatibility and other features al‐ ready mentioned in this chapter. It is also important to realize the multidisciplinarity involved in this process. The physi‐ cian is needed in order to perform the tissue biopsy to remove the cells **(I).** This tissue/cell is then taken to the laboratory to be multiplied several times. Thereafter, the use of principles of cell biology, such as growth factors **(II)** to stimulate the cells to grow and maintain their functions will be necessary. It is also required the involvement of engineers for manufac‐ turing matrices of biodegradable polymers **(III)** and the bioreactor **(IV).** When cells grow in appropriate number, they are seeded on the polymer scaffold. The tissue is then allowed to grow in the bioreactor until the time of transplantation by clinical surgeon. Biomaterials can be used to stimulate the growth of several types of tissues, e.g. bone, cartilage or skin.

Other extracellular matrices used as scaffolds include fibrin and fibrinogen. [51, 52, 53]. According to some studies, both can induce angiogenesis during tissue regeneration [54, 55, 56]. Chitosan is a derivative of chitin, a natural biopolymer which is biocompatible, biode‐ gradable, antimicrobial and possesses tissue healing and osteoinductive effects. It has the ability to bind to growth factors, glycosaminoglycans and DNA and can be easily processed into membranes, gels, nanofibres, beads, scaffolds and sponges. Because of these properties, chitosan gel alone or in combination with demineralized bone matrix/collagenous membrane

After the appropriate development, the tissue is transplanted and the area is regenerated.

is quite promising in periodontal regeneration [57].

degradation.

68 Current Concepts in Dental Implantology

Figure 4.

Andréa Cristina Barbosa da Silva\* , Diego Romário da Silva, Rafael Grotta Grempel, Manuel Antonio Gordón-Núñez and Gustavo Gomes Agripino

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

Graduate Program in Dentistry, Center of Sciences, Technology and Health, State University of Paraiba, UEPB, Araruna, Paraíba, Brazil

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