In vitro and in vivo evaluation of customized polycaprolactone tricalcium phosphate scaffolds for bone tissue engineering

IN VITRO AND IN VIVO EVALUATION OF CUSTOMIZED POLYCAPROLACTONE TRICALCIUM PHOSPHATE SCAFFOLDS FOR BONE TISSUE ENGINEERING ERVI SJU (B.Eng.(Hons.), NUS) A THESIS SUBMITTED FOR THE DEGREE OF MASTER OF ENGINEERING DEPARTMENT OF MECHANICAL ENGINEERING NATIONAL UNIVERSITY OF SINGAPORE 2010 PREFACE The thesis is submitted for the degree of Master of Engineering in the Department of Mechanical Engineering at the National University of Singapore under the supervision of Professor Teoh Swee Hin and Dr Alvin Yeo. No part of this thesis has been submitted for other degree at other university or institution. Parts of this thesis have been published or presented in the following: INTERNATIONAL JOURNAL PUBLICATION A. Yeo, E. Sju, B. Rai, S.H. Teoh. Customizing the degradation and load-bearing profile of 3D polycaprolactone-tricalcium phosphate scaffolds under enzymatic and hydrolytic conditions. Journal of Biomedical Materials Research Part B: Applied Biomaterials. (Published online: 10 June 2008). CONFERENCE PAPERS E. Sju, A. Yeo, B. Rai, S.H. Teoh. In vitro and in vivo degradation profile of untreated, sodium hydroxide- and lipase-treated PCL-TCP scaffolds. International Conference on Advances in Bioresorbable Biomaterials for Tissue Engineering, Singapore, 2008. E. Sju, A. Yeo, B. Rai, S.H. Teoh. Enzymatic and hydrolytic degradation of poly(εcaprolactone) tricalcium phosphate composite scaffolds. 4th International Conference on Materials for Advanced Technologies (ICMAT), Singapore, 2007. i ACKNOWLEDGEMENTS The author wishes to express her sincere gratitude and heartfelt appreciation to the following people who have rendered generous support and technical assistance leading toward the accomplishment of this project:  Professor Teoh Swee Hin (Department of Mechanical Engineering), supervisor, for offering the privileged opportunity to work on this project and allowing the author to join his team, for his expertise, kindness, and most of all, his patience. His enthusiasm in research and continuous support have truly been a source of inspiration and motivation for this project throughout.  Dr. Alvin Yeo (Department of Mechanical Engineering and National Dental Centre), co-supervisor, for his patience and guidance on supervising the author throughout the whole process. He has been an immense driving force behind this project. One simply could not wish for a better or friendlier supervisor.  Dr. Bina Rai, mentor, for graciously sharing her knowledge and encouragement in this project. Her kind assistance and time spent are greatly appreciated.  Dr. Zhang Zhiyong, Ms. Erin Teo Yi Ling and Mr. Mark Chong Seow Khoon, PhD students, for their constructive feedbacks and for being excellent mentors. They have gone out of their way to render assistance on many occasions.  Mr. Cheong Jia Jian, NUS alumnus, whom was unreserved in sharing his knowledge and experience in this research field.  Mdm. Zhong Xiang Li (Materials Science Lab) for the use of the SEM (JEOL JSM 5600LV) and the gold-sputtering machine. ii  Dr. Zhang Yanzhong (Biomechanics Lab) and Ms. Eunice Tan Phay Sing (NanoBiomechanics Lab) for the use of the Instron 3345 compressive mechanical tester machine.  Dr. Jeremy Teo Choon Meng (DSO National Lab) and Ms. Lei Yang (Biosignal and Instrumentation Lab) for the use of the Skyscan 1076 Micro-CT.  Dr. Lin Jian Hua and Ms. Juline Sim Siew Hong (PSB corporation) for their assistance in Gel Permeation Chromatography.  Ms. Irene Kee (SingHealth Experimental Medicine Centre, Singapore General Hospital) for her assistance in the animal handling and maintenance.  Ms. Han Tok Lin (Faculty of Dentistry, NUS) for her assistance in Histology.  Mr. Jackson Ong Sing Kiat and Dr. Chui Chee Kong (BIOMAT), Mr. Zhang Jing (Biosignal and Instrumentation Lab), and fellow students at BIOMAT for their support and encouragement throughout the fulfilling years.  To all, who have given contribution in one way or another.  And to all close friends, for being understanding during this challenging period. Thank you for always being there during both good and bad times.  Last but not least, the author would like to thank her parents Mr. Sju Tjing Kwang and Mrs. Tea Giok Tjian, and younger sister Ms. Lydia Sju, for their constant love and support, without which this study would not have been possible. Their undaunting confidence gave the author the strength to overcome any difficulties. To them the author dedicates this thesis. The author acknowledged the financial support by the following grants:  No. 016/06 from National Dental Centre (SingHealth), Singapore.  TDF/CD003/2006 from SingHealth (Talent Development Fund), Singapore. iii TABLE OF CONTENTS PREFACE i ACKNOWLEDGEMENTS ii TABLE OF CONTENTS iv SUMMARY ix LIST OF TABLES xi LIST OF FIGURES xii LIST OF SYMBOLS xviii LIST OF ABBREVIATIONS xx CHAPTER 1: INTRODUCTION 1.1 BACKGROUND 1 1.1.1 Bone tissue engineering 1 1.1.2 Application in dentoalveolar defects 3 1.1.3 PCL-TCP scaffolds: Current drawbacks 4 1.2 RESEARCH OBJECTIVES 5 1.3 RESEARCH SCOPE 6 1.3.1 Part 1: Selective modification of PCL-TCP scaffolds targeted for dentoalveolar reconstruction application 6 1.3.2 Part 2: Optimization of native and customized scaffolds in vitro and their effects in initial bone healing 6 1.3.3 Part 3: Evaluation of PCL-TCP scaffolds in a clinically relevant defect model 7 iv CHAPTER 2: LITERATURE REVIEW 2.1 BONE PHYSIOLOGY 8 2.1.1 Composition 8 2.1.2 Morphology 11 2.2 POLY(ε-CAPROLACTONE) 13 2.2.1 Degradation of PCL polymer 14 2.2.1.1 Hydrolysis mechanism 16 2.2.1.2 Enzymatic degradation 17 2.3 TRICALCIUM PHOSPHATE (TCP) 2.3.1 Degradation mechanisms of calcium phosphate ceramics 18 19 2.3.1.1 Physicochemical degradation 19 2.3.1.2 Cellular degradation 21 2.3.1.3 Mechanical degradation 21 2.4 PCL-TCP SCAFFOLDS 22 2.4.1 Fabrication method of PCL-TCP scaffolds 24 CHAPTER 3: SELECTIVE MODIFICATION OF PCL-TCP SCAFFOLDS TARGETED FOR DENTOALVEOLAR RECONSTRUCTION APPLICATION 3.1 INTRODUCTION 29 3.2 MATERIALS AND METHODS 30 3.2.1 Scaffold design and fabrication 30 3.2.2 Sterilization of scaffolds 30 3.2.3 Scaffold characterizations 31 3.2.3.1 Micro-computed tomography analysis 31 3.2.3.2 Gravimetric analysis 31 3.2.3.3 Compressive mechanical testing 32 3.2.3.4 Electron microscopy preparation and analysis 32 v 3.2.3.5 Molecular weight testing 3.2.4 Statistical analysis 33 33 3.3 RESULTS 34 3.3.1 Porosity measurements and 3D model analysis 34 3.3.2 Weight loss analysis 36 3.3.3 Compressive mechanical properties 37 3.3.4 Surface morphology analysis 38 3.3.5 Molecular weight analysis 40 3.4 DISCUSSION 40 3.5 CONCLUSION 43 CHAPTER 4: OPTIMIZATION OF NATIVE AND CUSTOMIZED SCAFFOLDS IN VITRO AND THEIR EFFECTS IN INITIAL BONE HEALING 4.1 INTRODUCTION 44 4.1.1 In vitro degradation study 44 4.1.2 In vivo degradation study 45 4.2 MATERIALS AND METHODS 46 4.2.1 Scaffold design and fabrication 46 4.2.2 Sterilization of scaffolds 47 4.2.3 Animal husbandry 47 4.2.4 Scaffold implantation 48 4.2.5 Scaffold characterizations 49 4.2.5.1 Micro-computed tomography analysis 50 4.2.5.2 Gravimetric analysis 50 4.2.5.3 Compressive mechanical testing 50 4.2.5.4 Electron microscopy preparation and analysis 50 4.2.5.5 Molecular weight testing 50 vi 4.2.5.6 Histology preparation and analysis 50 4.2.6 Statistical analysis 51 4.3 RESULTS - In vitro degradation study 51 4.3.1 Porosity measurements and 3D model analysis 51 4.3.2 Weight loss analysis 57 4.3.3 Compressive mechanical properties 58 4.3.4 Surface morphology analysis 60 4.3.5 Molecular weight analysis 66 4.4 RESULTS - In vivo degradation study 67 4.4.1 Porosity measurements and 3D model analysis 67 4.4.2 Weight loss analysis 70 4.4.3 Compressive mechanical properties 71 4.4.4 Surface morphology analysis 72 4.4.5 Molecular weight analysis 75 4.4.6 Histology analysis 76 4.5 DISCUSSION 79 4.5.1 Comparison between in vitro and in vivo studies 4.6 CONCLUSION 82 83 CHAPTER 5: EVALUATION OF PCL-TCP SCAFFOLDS IN A CLINICALLY RELEVANT DEFECT MODEL 5.1 INTRODUCTION 85 5.2 MATERIALS AND METHODS 88 5.2.1 Implant design and fabrication 88 5.2.2 Animal husbandry 89 5.2.3 Pre- and postoperative medication 90 5.2.4 Surgery 1 (Extraction and defect creation) 91 vii 5.2.5 Surgery 2 (Ridge augmentation) 93 5.2.6 Sacrifice 95 5.2.7 Micro-computed tomography analysis 96 5.3 RESULTS 96 5.3.1 Gross examinations 96 5.3.2 New bone formation 98 5.3.3 Ratio of bone volume fraction for PCL-TCP scaffolds with respect to autografts 100 5.3.4 3D model analysis 101 5.3.5 Two-dimensional x-ray radiographs evaluation 102 5.4 DISCUSSION 104 5.5 CONCLUSION 109 CHAPTER 6: FINAL CONCLUSIONS AND RECOMMENDATIONS 6.1 FINAL CONCLUSIONS 110 6.2 RECOMMENDATIONS FOR FUTURE WORK 112 REFERENCES 114 APPENDICES 122 APPENDIX A – PART 1 STUDY 122 APPENDIX B – PART 2 STUDY 134 APPENDIX C – PART 3 STUDY 148 viii SUMMARY The research scope encompasses the degradation and load-bearing profile of 3D bioresorable polycaprolactone-20% tricalcium phosphate (PCL-TCP) scaffolds under enzymatic and hydrolytic conditions and subsequently to evaluate the efficacy of the scaffolds in both small and large animal models. The purpose was to develop scaffolds with desirable customized properties and increased degradation rates suitable for application in dentoalveolar defects treatment. The scope of this thesis ended with a large animal study, a stage just before preclinical trials. Initially, the PCL-TCP scaffolds were degraded in either sodium hydroxide or lipase solution for 0, 12, 24, 36, 48, 60, 72, 84, 96, and 108 hours. Samples were recovered at each time point and the following properties of the scaffolds were measured: porosity, 3D structure, weight loss, compressive strength and modulus, surface morphology, polymer molecular weight, and histology. In the second part of the study, in vitro and in vivo degradation behaviours of these treated scaffolds were investigated. PCL-TCP scaffolds were monitored after immersion in standard culture medium for 0, 6, 12, 18 and 24 weeks in vitro. In vivo degradation of the scaffolds was performed by implanting these scaffolds subcutaneously at the back of rats for 12 and 24 weeks. Upon retrieval, analyses similar to those described above were performed. Lastly, another in vivo study was conducted whereby PCL-TCP scaffolds and sheets were evaluated as defect fillers and barrier membranes respectively for novel guided bone regeneration technique in the reconstruction of localized ix dentoalveolar defects in a micropig model for up to 6 months. The possibility of the PCL-TCP scaffold for use as a bone substitute was compared to the current gold standard of using autogenous bone. The first objective of the study was achieved with scaffolds of approximately 85% porosity obtained after 96 hours of treatment in 3M NaOH and 12 hours in 0.1% lipase. These pre-treated scaffolds demonstrated favourable mechanical strength, structure, and surface morphology. Secondly, the in vivo degradation profile of porous PCL-TCP scaffolds are comparable with the obtained in vitro profile. Further, the degradation rate of the lipase-treated scaffolds was noted to be the highest. This is followed by NaOH-treated scaffolds and native untreated scaffolds. Overall, the data suggest that NaOH-treated scaffolds demonstrate the best degradation profile and physical properties for dentoalveolar reconstruction applications. They possess the potential to degrade in a controlled and predictable fashion and still display favourable mechanical strength within a desired time period for new bone formation to occur. Lastly, healing was uneventful in all micropigs showed that the PCL-TCP scaffolds exhibited good biocompatibility. Across the tested treatment options, defect sites augmented with autografts and collagen membranes showed the most promising results with greater bone formation detected as compared to PCL-TCP scaffolds and collagen membranes which were about 64% efficient. The collagen membranes were found to offer the advantage of a reduced frequency of soft tissue dehiscence in comparison to PCL-TCP sheets. More improvements are needed to increase the efficiency of the PCL-TCP scaffolds in bone healing as they could ruled out the need for harvesting grafts. x LIST OF TABLES Table 3.1 Mw, Mn, and PDI of NaOH-treated and lipase-treated PCL-TCP Scaffolds. 40 Table 4.1 Mw, Mn, and PDI of native, NaOH-treated, and lipase-treated PCL-TCP Scaffolds in vitro. 66 Table 4.2 Mw, Mn, and PDI of native, NaOH-treated, and lipase-treated PCL-TCP Scaffolds in vivo. 75 Table 5.1 Number of sites with soft tissue dehiscence for the implanted autograft, collagen membranes, PCL-TCP scaffolds, and PCLTCP sheets. 98 xi LIST OF FIGURES Figure 1.1 Schematic diagram of research scope. 6 Figure 1.2 Schematic diagram of part 1 and part 2 study. 7 Figure 2.1 Composition of bone. 9 Figure 2.2 The assembly of collagen fibrils and fibers and bone mineral crystals. 9 Figure 2.3 Microscopic structure of cortical and cancellous bone. 11 Figure 2.4 The hierarchical structure of bone from macrostructure to subnanostructure. 12 Figure 2.5 Chemical structure of PCL (as circled). 13 Figure 2.6 The degradation rate of PGA, PLA, and PCL. 16 Figure 2.7 The chemical structure of TCP. 19 Figure 2.8 Schematic diagram of the FDM process. 25 Figure 2.9 Sequence of the data preparation for FDM model fabrication. 26 Figure 3.1 Centrifuge tubes. 29 Figure 3.2 Illustration of scaffold with 0/60/120º lay-down pattern. 30 Figure 3.3 5x5x3mm PCL-TCP scaffold. 31 Figure 3.4 Porosity measurements of NaOH-treated and lipase-treated PCL-TCP Scaffolds over time. 34 Figure 3.5 3D model of original scaffold (of 75% porosity) at 0 hour: (L) top view, (R) tilted view. 35 Figure 3.6 3D model of scaffolds after 96 hours immersion in 3M NaOH: (L) top view, and (R) tilted view. 35 Figure 3.7 3D model of scaffolds after 12 hours immersion in 0.1% lipase: (L) top view, and (R) tilted view. 36 xii Figure 3.8 Comparison of weight loss between NaOH-treated and lipase-treated PCL-TCP scaffolds. 36 Figure 3.9 Compressive strength of PCL-TCP scaffolds when treated with NaOH and lipase at pre-determined time intervals. 37 Figure 3.10 Compressive modulus of PCL-TCP scaffolds when treated with NaOH and lipase at pre-determined time intervals. Electron micrographs of original scaffold (of porosity 75%) at 0 hour: (L) overall view, and (R) higher-magnification view. 38 Figure 3.12 Electron micrographs of scaffold after 96 hours immersion in 3M NaOH: (L) overall view, and (R) higher-magnification view. 39 Figure 3.13 Electron micrographs of scaffold after 12 hours immersion in 0.1% lipase: (L) overall view, and (R) higher-magnification view. 39 Figure 4.1 Native (left), NaOH-treated (middle), and lipase-treated (right) scaffolds. 45 Figure 4.2 Rat at the start of experiment (left) and at the end after 6 months (right). 46 Figure 4.3 50x50x3mm PCL-TCP scaffold. 46 Figure 4.4 5x5x3mm PCL-TCP scaffold. 47 Figure 4.5 Rat cages. 47 Figure 4.6 Rat shaved and scrubbed with iodine. 48 Figure 4.7 Scaffolds’ positioning. 48 Figure 4.8 Incision made (left), implanted scaffold (left, inset), scaffold positions (right). 49 Figure 4.9 Sacrifice of rats. 49 Figure 4.10 Removal of scaffolds. 49 Figure 4.11 Porosity measurements of native, NaOH-treated, and lipasetreated PCL-TCP scaffolds after immersion in DMEM for 6, 12, and 18 weeks. 52 Figure 4.12 3D model of native scaffold (of 85% porosity) at week 0: (L) top view, and (R) tilted view. 52 Figure 3.11 xiii 38 Figure 4.13 3D model of native scaffold after 6 weeks immersion in DMEM: (L) top view, and (R) tilted view. 53 Figure 4.14 3D model of NaOH-treated scaffold after 6 weeks immersion in DMEM: (L) top view, and (R) tilted view. 53 Figure 4.15 3D model of lipase-treated scaffold after 6 weeks immersion in DMEM: (L) top view, and (R) tilted view. 53 Figure 4.16 3D model of native scaffold after 12 weeks immersion in DMEM: (L) top view, and (R) tilted view. 54 Figure 4.17 3D model of NaOH-treated scaffold after 12 weeks immersion in DMEM: (L) top view, and (R) tilted view. 54 Figure 4.18 3D model of lipase-treated scaffold after 12 weeks immersion in DMEM: (L) top view, and (R) tilted view. 54 Figure 4.19 3D model of native scaffold after 18 weeks immersion in DMEM: (L) top view, and (R) tilted view. 55 Figure 4.20 3D model of NaOH-treated scaffold after 18 weeks immersion in DMEM: (L) top view, and (R) tilted view. 55 Figure 4.21 3D model of lipase-treated scaffold after 18 weeks immersion in DMEM: (L) top view, and (R) tilted view. 55 Figure 4.22 3D model of native scaffold after 24 weeks immersion in DMEM: (L) top view, and (R) tilted view. 56 Figure 4.23 3D model of NaOH-treated scaffold after 24 weeks immersion in DMEM: (L) top view, and (R) tilted view. 56 Figure 4.24 3D model of lipase-treated scaffold after 24 weeks immersion in DMEM: (L) top view, and (R) tilted view. 56 Figure 4.25 Weight loss of PCL-TCP Scaffolds In vitro. 58 Figure 4.26 Relative compressive strength of PCL-TCP Scaffolds In vitro. 59 Figure 4.27 Relative compressive modulus of PCL-TCP Scaffolds In vitro. 59 Figure 4.28 Electron micrographs taken after 6 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipasetreated scaffolds. (L) overall view, and (R) highermagnification view. 62 xiv Figure 4.29 Electron micrographs taken after 12 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipasetreated scaffolds. (L) overall view, and (R) highermagnification view. 63 Figure 4.30 Electron micrographs taken after 18 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipasetreated scaffolds. (L) overall view, and (R) highermagnification view. 64 Figure 4.31 Electron micrographs taken after 24 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipasetreated scaffolds. (L) overall view, and (R) highermagnification view. 65 Figure 4.32 Electron micrographs of native scaffold (of 85% porosity) at week 0: (L) overall view, and (R) higher-magnification view. 66 Figure 4.33 Porosity of PCL-TCP Scaffolds In vivo. 67 Figure 4.34 3D model of native scaffold after 3 months implantation: (L) top view, and (R) tilted view. 68 Figure 4.35 3D model of NaOH-treated scaffold after 3 months implantation: (L) top view, and (R) tilted view. 68 Figure 4.36 3D model of lipase-treated scaffold after 3 months implantation: (L) top view, and (R) tilted view. 68 Figure 4.37 3D model of native scaffold after 6 months implantation: (L) top view, and (R) tilted view. 69 Figure 4.38 3D model of NaOH-treated scaffold after 6 months implantation: (L) top view, and (R) tilted view. 69 Figure 4.39 3D model of lipase-treated scaffold after 6 months implantation: (L) top view, and (R) tilted view. 69 Figure 4.40 Weight loss of PCL-TCP Scaffolds In vivo. 70 Figure 4.41 Relative compressive strength of PCL-TCP Scaffolds In vivo. 71 Figure 4.42 Relative compressive modulus of PCL-TCP Scaffolds In vivo. 72 Figure 4.43 Electron micrographs taken after 3 months implantation: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 73 xv Figure 4.44 Electron micrographs taken after 6 months implantation: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 74 Figure 4.45 H&E stain of native scaffolds after 3 months implantation. 76 Figure 4.46 H&E stain of native scaffolds after 6 months implantation. 76 Figure 4.47 H&E stain of NaOH-treated scaffolds after 3 months implantation. 77 Figure 4.48 H&E stain of NaOH-treated scaffolds after 6 months implantation. 77 Figure 4.49 H&E stain of lipase-treated scaffolds after 3 months implantation. 78 Figure 4.50 H&E stain of lipase-treated scaffolds after 6 months implantation. 78 Figure 5.1 Timeline for the complete micropig study. 87 Figure 5.2 15x10x8mm PCL-TCP scaffold (left) and 25x25x1mm PCLTCP sheet (right). 88 Figure 5.3 Bioresorbable collagen membrane from BioGide (left) and temperature-controlled hot water bath (right). 89 Figure 5.4 Micropig housing facility at SEMC, SGH (left) and weighing of micropig prior to the experiment (right). 90 Figure 5.5 Removal of all premolars and first molar (left), and the extraction sites (right). 92 Figure 5.6 The flaps were re-approximated with Vicryl sutures (left), and the defect sites were closed (right). 92 Figure 5.7 Schematic illustrations of the four tested grafting procedures. 93 Figure 5.8 Placement of PCL-TCP scaffolds and autografts (left), followed by PCL-TCP sheets and collagen membranes (right). 94 Figure 5.9 Micropig under euthanasia (left), and the mandible was block resected using an oscillating autopsy saw (right). 95 Figure 5.10 The recovered segment of mandible (left), the site after removal (right). 95 Figure 5.11 The recovered segment of the mandible of a micropig. 97 xvi Figure 5.12 Soft tissue dehiscence observed for the majority of grafts covered with PCL-TCP sheets. 97 Figure 5.13 Bone volume fraction detected after 6 months of implantation of autografts and PCL-TCP scaffolds for individual micropigs. 99 Figure 5.14 The average values of bone volume fraction detected after 6 months of implantation of autografts and PCL-TCP scaffolds. 100 Figure 5.15 The ratio of bone volume fraction for PCL-TCP scaffolds with respect to autografts for individual micropigs. 101 Figure 5.16 PCL-TCP scaffold treated site: overview (left) and crosssection (right). 102 Figure 5.17 Autograft-treated site: overview (left) and cross-section (right). 102 Figure 5.18 X-ray image of a micropig’s left mandible treated with autograft (posterior) and PCL-TCP scaffold (anterior), and covered with collagen membrane. 103 Figure 5.19 X-ray image of a micropig’s right mandible treated with PCLTCP scaffold (posterior) and autograft (anterior), and covered with collagen membrane. 103 Figure 5.20 X-ray image of a micropig’s left mandible treated with autograft (posterior) and PCL-TCP scaffold (anterior), and covered with collagen membrane. 104 xvii LIST OF SYMBOLS ºC Celcius CaCl2 Calcium Chloride CO2 Carbondioxide H&E Hematoxylin & Eosin H2O Water KCl Potassium Chloride KH2PO4 Potassium Dihydrogen Phosphate kN Kilonewton kV Kilovolt mm Milimeter Mn Number-average Molecular weight MPa Mega-pascal Mw Weight-average Molecular weight NaCl Sodium Chloride Na2HPO4 Sodium Hydrogen Phosphate NaOH Sodium Hydroxide O2 Oxygen P Probability rpm Revolution per minute Tg Glass transition temperature xviii Tm Melting point W0 Initial dry weight Wdry Dry weight at time t μA Microampere μm Micrometer xix LIST OF ABBREVIATIONS 3D Three Dimensional ABG Autogenous Bone Graft BMP Bone Morphogenetic Protein BV Bone Volume BVF Bone Volume Fraction CAD Computer-aided design CT Computed Tomography DMEM Dulbecco’s modified Eagle’s medium ECM Extracellular matrix FDA US Food and Drug Administration FDM Fused Deposition Modeling GA Gravimetric Analysis GBR Guided Bone Regeneration GPC Gel Permeation Chromatography IM Intramuscular ISO International Standards Organization IV Intravenous Lipase PS Pseudomonas Lipase Micro-CT Micro-computed Tomography PBS Phosphate Buffered Saline xx PCL Poly(ε-caprolactone) PDI Polydispersity Index PGA Poly(glycolic acid) PLA Poly(lactic acid) QS QuickSlice rhBMP-2 Recombinant human Bone Morphogenetic Protein-2 RP Rapid Prototyping SD Standard Deviation SEM Scanning Electron Microscope SEMC SingHealth Experimental Medicine Centre SFF Solid Free-form fabrication SGH Singapore General Hospital STL Stereolithography TCP Tricalcium Phosphate THF Tetrahydrofuran TV Tissue Volume xxi CHAPTER 1: INTRODUCTION 1.1 BACKGROUND This section aims to provide background information regarding bone tissue engineering strategy and the application in implant dentistry, as well as the current drawback of PCL-TCP scaffolds in dentoalveolar defects treatment that lead the author to pursue this research. Detailed research objectives and research scope are discussed in the next and last sections respectively. 1.1.1 Bone tissue engineering Loss of human tissues or organs is a devastating problem that can affect individuals of all ages. Bone, a complex living tissue that provides internal support for all higher vertebrates, is currently heralded as the most commonly replaced organ of the body. In fact, with over 1.3 million bone repair procedures performed per year in the United States alone [Chim, 2006], the ability to come up with an innovative and effective defects treatment to satisfy the major clinical need has indeed been a great challenge for many researchers. Historically, autogenous or allogenic bone grafts have been used for treatment in bone defects. Often, the bone repair mechanism fails as a result of magnitude, infection or other causes. Autogenous bone grafts are those made of tissue obtained from the patient who receives the graft, while allogenic bone grafts are those made of tissue from a human donor, usually post-mortem. However, these techniques 1 have some drawbacks. Harvesting of autogenous bone grafts induces additional trauma and morbidity, increase operation times, and are often limited in supply. At the site of bone transplantation, the risks of wound infection, necrosis, and resorption, representing up to 30% of transplanted material have also been experienced [Betz, 2002; Horch, 2006]. Allogenic bone grafts present risks of possible disease transmission and problems of religious implications [Hutmacher, 2005; Celil, 2006]. These limitations have then instigated new research aiming to provide a bone graft engineered in the laboratory and readily available. The ultimate goal of this approach was the regeneration rather than just the repair of skeletal tissue, and this treatment strategy was later coined as “bone tissue engineering”. A key component in tissue engineering for bone regeneration is the scaffold that serves as a 3D template for initial cell interactions and the formation of boneextracellular matrix to provide structural support to the newly formed tissue. The porous scaffold provides the necessary support for cells to attach, proliferate, and maintain their differentiated function. The ability of the scaffold to be metabolized by the body allows it to be gradually replaced by cells to form functional tissues [Pollok, 1996]. A well-designed scaffold for bone tissue engineering then plays an important role in facilitating bone healing. To do so effectively, several qualities of an effective scaffold material must be satisfied. Ideally, a scaffold should possess the following properties: (1) a 3D structure with an increased porosity and a highly interconnected pore network for cellular or vascular ingrowth and transport of nutrients and metabolic waste; (2) biocompatibility and bioresorbability with controlled degradation and resorption rates to match tissue replacement; (3) suitable surface properties for cell adhesion, proliferation, and differentiation; and (4) sufficient mechanical 2 properties to match those of the tissues at the site of implantation [Hutmacher, 2001]. The latter is extremely crucial in skeletal tissue such as bone and cartilage where certain mechanical properties are required. These scaffolds serve as temporary load-bearing devices that provide adequate strength and help maintain space for new bone formation to occur [Hutmacher, 2000; Rezwan, 2006; Zhou, 2007]. 1.1.2 Application in dentoalveolar defects In implant dentistry, clinical situations involving major defects or deformities as the result of trauma or diseases are often faced. The outcome is a compromised and deficient alveolar ridge, which is often extended and non-contained and frequently requiring extensive guided bone regeneration (GBR) procedures. In the dentoalveolar skeleton, an inadequate bone volume always creates problems in the prosthetic and esthetic reconstruction of partially and completely edentulous situations. In an era where implant borne tooth restorations have became the standard of care for the replacement of missing teeth, the quantity and quality of the available bony ridge is critical in determining whether ridge augmentation is required prior to dental implant placement [Adell, 1990; Jemt, 1993]. This will not only determine the outcome of a favorable ridge shape and the contour of the overlying soft tissue, but also the optimal three-dimensional placement of the dental implant. This is where the role of scaffolds come into the picture as they may eliminate the need for an extensive bone harvesting procedure from a donor site. However in facing a complex biological system as the human body, the requirements of scaffold materials for bone tissue engineering in dentoalveolar application can be extremely challenging. 3 1.1.3 PCL-TCP scaffolds: Current drawback The use of synthetic polymers in the field of tissue engineering has been widely investigated in recent years, with advances in the scaffold technology extending their usage to clinical applications such as bone regeneration. In particular of such interest is poly(ε-caprolactone)-tricalcium phosphate (PCL-TCP) composite scaffold, a synthetic biodegradable polymer frequently investigated for bone tissue engineering applications. Recent studies on PCL-TCP scaffolds have demonstrated favourable biocompatibility, bioactivity, and mechanical characteristics [Rai, 2004; Schantz, 2003; Rai, 2005]. However due to their high molecular weight and hydrophobicity, they degrade at a slow rate [Jeong, 2003; Ha, 1997]. This is a disadvantage for bone tissue engineering purposes in dentoalveolar application, as the new bone replacing the scaffold are inserted with dental implants for prosthetic rehabilitation [Wu, 2004; Lei, 2006]. Degradation behaviors of porous scaffolds play an important role in the engineering of new tissue, since the degradation rate is intrinsically linked to cell vitality, growth, as well as host response. In order for a biodegradable scaffold to be successful over the long term, the material must have a rate of degradation that acts in concert with the ingrowth of new bone. Ideally, the degradation and resorption kinetics of composite scaffolds should be designed such that the cells are allowed to adhere, proliferate, and secrete their own extracellular matrix (ECM) as the scaffolds gradually resorbs, creating space for new cell and tissue growth. The physical support provided by the three-dimensional (3D) scaffold should also be maintained until the regenerated tissue has sufficient mechanical integrity to support itself [Putnam, 1996]. Thus, it would be desirable to control the degradation of the PCL-TCP scaffolds to be in sync with the formation of new bone 4 targeted for dentoalveolar defects treatment (which takes approximately 5-6 months). 1.2 RESEARCH OBJECTIVES The interest of this study was to investigate the degradation and load-bearing profile of 3D bioresorable PCL-TCP scaffolds under enzymatic and hydrolytic conditions and subsequently to evaluate the efficacy of the scaffolds in both small and large animal models. The purpose was to develop scaffolds with desirable customized properties and increased degradation rates suitable for application in dentoalveolar defects treatment. In this research, specific aims have been identified: 1. To obtain PCL-TCP scaffolds with the desired higher porosity of 85% by treating them with 3M NaOH or 0.1% lipase-PBS medium under physiological conditions for up to 108 hours. 2. To compare the degradation profile of treated and untreated PCL-TCP scaffolds in vitro when immersed in standard culture medium for up to 24 weeks, and in vivo when implanted in the subcutaneous back of rats for 24 weeks (6 months). 3. To evaluate the rate and extent of bone formation of PCL-TCP scaffolds in vivo when implanted in a larger, clinically relevant defect model for up to 6 months. Micropigs were chosen as the animal models. 5 1.3 RESEARCH SCOPE In order to meet the objectives stated in the previous section, the research scope (Figure 1.1) has been divided into three parts as follows: In vitro Small animal model Large animal model Figure 1.1: Schematic diagram of research scope. 1.3.1 Part 1: Selective modification of PCL-TCP scaffolds targeted for dentoalveolar reconstruction application (in Chapter 3) PCL-TCP scaffolds (75% porosity) were treated using 3M NaOH or 0.1% lipase for 0, 12, 24, 36, 48, 60, 72, 84, 96, and 108 hours. Samples were recovered at each time intervals and properties such as porosity, mechanical strength, surface degradation and surface characteristics of the scaffolds were evaluated. This part serves as an initial stage of a larger project, in order to develop a scaffold of a higher porosity that allows for a more rapid degradation whilst maintaining favourable mechanical properties. A final porosity of about 85% was targeted for. 1.3.2 Part 2: Optimization of native and customized scaffolds in vitro and their effects in initial bone healing (in Chapter 4) In the second part of the study, PCL-TCP scaffolds of a higher porosity (85%) were tested. The scaffolds were divided into 3 experimental groups: NaOH-treated, lipasetreated and untreated. They were (a) implanted subcutaneously into the back of rats, 6 or (b) immersed in DMEM growth media, for various time periods of up to 6 months. Analysis similar to those described in part 1 were performed. Pre-degradation Study 0, 12, 24, 36, 48, 60, 72, 84, 96, 108 hours In Vitro Degradation Study In Vivo Degradation Study 0, 6, 12, 18, 24 weeks Micro-CT Gravimetric → Porosity → Structure → Weight Loss 0, 12, 24 weeks Mechanical Testing → Strength → Stiffness SEM GPC Histology → Surface → Molecular → Inflammation Morphology Weight → Vascularisation Figure 1.2: Schematic diagram of part 1 and part 2 study. 1.3.3 Part 3: Evaluation of PCL-TCP scaffolds in a clinically relevant defect model (in Chapter 5) In the third and last part of the study, PCL-TCP scaffolds and thin sheets were implanted in the posterior mandible of micropigs, after two lateral ridge defects were initially created in each side of the mandible. Following a healing period of 6 months, the micropigs were sacrificed and the harvested specimens were characterized. The scope of this thesis ended at the preclinical stage, which was this large animal study. 7 CHAPTER 2: LITERATURE REVIEW 2.1 BONE PHYSIOLOGY 2.1.1 Composition Bone, a subset of a large and diverse group of tissues collectively referred to as connective tissue, is the main building block of the human skeletal system. Bone is made up of organic and inorganic (mineral) matter, cells, and water (Figure 2.1). The organic matter is concentrated in the bone matrix, which amounts to about 35% of the dry weight of bone. It consists of 90% collagen, which is thus by far the most abundant bone protein. Collagen assembles in an organised pattern within the bone microstructure and modulates bone calcification sites (Figure 2.2). Its complex threedimensional structure, comparable to that of a rope, gives bone tissue its tensile strength. The remainder of the bone matrix is made up of various noncollagenous proteins such as cytokines, osteonectin, osteopontin, osteocalcin, growth factors, bone sialoprotein, hyaluronan, thrombospondin, proteoglycans, phospholipids, and phosphoproteins [Rho, 1998; Wang, 2001; Glimcher, 1989; Fleisch, 2000]. Together they play an important role in bone remodelling and in osteogenesis. The mineral matter of bone consists mainly of mineral salts known as hydroxyapatites, which are largely made up of calcium phosphates. Tiny crystals of these salts lie within and between the collagen fibers in the extracellular matrix, producing the compressive strength and stiffness that is so characteristic of bone [van Gaalen, 2008]. The 8 proper combination of the fibers and salts then allows bones to be both strong and durable without being brittle [Glimcher, 1998; Baron, 1996]. Figure 2.1 (above): Composition of bone [Fleisch, 2000]. Figure 2.2 (right): The assembly of collagen fibrils and fibers and bone mineral crystals [Rho, 1998]. Bone’s function is both biomechanical and metabolic. Biomechanically, bone acts to: (1) maintain the shape of the skeleton, (2) protect soft tissues in the cranial, thoracic and pelvic cavities, (3) transmit the forces of muscular contraction during movement, and (4) supply a framework for bone marrow. Metabolically, bone (1) serves as a reservoir for ions, especially calcium ions, and (2) contributes to the regulation of the extracellular matrix composition [Ferrer, 2007]. Bone is a self-repairing structural material; it is capable of adapting its mass, shape and properties to the changes in mechanical requirements and endures voluntary physical activity for life without breaking. This capacity stems from the fact that bone is in fact alive, and contains cells which work continuously to regenerate and repair it 9 [Bronner, 1999; Ferrer, 2007]. Bone tissue contains five basic types of bone cells: osteogenic cells, osteoblasts, osteocytes, osteoclasts, and bone-lining cells. Osteogenic cells respond to traumas, such as fractures, and begin the healing process immediately by giving rise to bone-forming cells (osteoblasts) and bonedestroying cells (osteoclasts). They can be found in the bone tissue which contacts the endosteum and the periosteum. Osteoblasts are cell which synthesize and secrete basic un-mineralized compound to help in the process of bone repair, bone growth, or bone regrowth. Osteoblast-secreted extracellular matrix may initially be amorphous and noncrystalline, but it gradually transforms into more crystalline forms [Boskey, 2003]. Mineralization is a process of bone formation promoted by osteoblasts and is thought to be initiated by the matrix vesicles that bud from the plasma membrane of osteoblasts to create an environment for the concentration of calcium and phosphate, allowing crystallization [Barckhaus, 1978; Celil, 2006]. Where the bone tissue has higher metabolism, the osteoblast cells are more plentiful, this includes the border of the medullary cavity and under the periosteum. A mature osteoblast is known as an osteocyte. While osteocytes are technically a different bone cell altogether, the osteoblast changes into an osteocyte over time. Osteoblasts have the unique ability to secrete bone tissue and form the tissue around itself like a protective wall of bone tissue. They are responsible for the maintenance of healthy bone by secreting enzymes and directing the bone mineral content. They also control the calcium release from the bone tissue to the blood. The cells which are responsible for the breakdown of bone tissue, which releases calcium, are known as osteoclasts. Osteoclasts are vital to the process of bone growth, remodeling, and healing. The last type of cells are bone-lining cells. They are made from osteoblasts along the surface of most bones in an adult, and are 10 thought to regulate the movement of calcium and phosphate into and out of the bone [Chenu, 1998]. 2.1.2 Morphology Macroscopically, bone can be divided into an outer part called cortical or compact bone, which makes about 80% of the total skeleton, and an inner part named cancellous, trabecular, or spongy bone. Cortical bone is very dense and contains only microscopic channels. Forming the outer wall of bones, it bears most of the supportive and protective function of the skeleton. Cancellous bone, on the other hand, makes up the remaining 20% of bone mass in the body. It consists of trabeculae which form an interconnected lattice. Cancellous bone can be found in vertebrae, fracture joints, ends of long bones and in foetuses. The whole structure, an outer cortical sheath and an inner three-dimensional trabecular network, allows optimal mechanical function under customary loads [van Gaalen, 2008; Brickley, 2008; Rho, 1998; Ferrer 2007]. Figure 2.3: Microscopic structure of cortical and cancellous bone [US National Cancer Institute’s SEER Program, 2009]. 11 Microscopically, woven and lamellar bone can be distinguished. Woven bone is the type formed initially in the embryo and during growth, and is characterized by an irregular array of loosely packed collagen fibrils. It is then replaced by lamellar bone, so that it is practically absent from the adult skeleton, except under pathological conditions of rapid bone formation, such as occur in Paget's disease, fluorosis, or fracture healing. In contrast, lamellar bone is the form present in the adult, both in cortical and in cancellous bone. It is made of well-ordered parallel collagen fibers, organized in a lamellar pattern called osteons or haversian systems. The osteon consists of a central canal called the osteonic (haversian) canal, which is surrounded by concentric rings (lamellae) of matrix. Between the rings of matrix, the osteocytes are located in the lacunae. The osteonic canals contain blood vessels that are parallel to the long axis of the bone. These blood vessels interconnect, by way of the canaliculi, with vessels on the surface of the bone [van Gaalen, 2008; Rho, 1998; Ferrer, 2007]. Figure 2.4: The hierarchical structure of bone from macrostructure to subnanostructure [Rho, 1998]. 12 2.2 POLY(ε-CAPROLACTONE) Poly(ε-caprolactone) (PCL) is a semi-crystalline, biodegradable, and bioresorbable polymer widely used in tissue engineering recently [Teoh, 2004]. It has a melting point (Tm) of 60ºC and a low glass transition temperature (Tg) of -60ºC that gives it rubbery characteristics and be relatively ductile at room temperature [Gan, 1999]. It is synthesized by ring-opening polymerization of ε-caprolactone monomers. As a homopolymer belonging to the aliphatic polyester family, the repeating molecular structure of PCL consists of a 5 non-polar methylene group and a single relatively polar ester group. The presence of this hydrolytically unstable aliphatic-ester linkage along the polymer backbone attributed to the biodegradability of the polymer [Perrin, 1997]. When the polymer is implanted in the body, hydrolysis of polymer backbone reduces the molecular weight of polymer and the degraded products can be metabolized in the body. The presence of methylene groups on PCL also renders it non-polar; hence, PCL is hydrophobic and its resistance to a number of medium such as water, oil and solvent gives it a slow degradation rate. Figure 2.5: Chemical structure of PCL (as circled) [Wikimedia, 2007]. The biocompatibility of PCL has been confirmed through extensive in vitro and in vivo studies and approved by US Food and Drug Administration (FDA) for its usage 13 in various medical applications namely sutures and drug delivery systems [Zein, 2002; Pitt, 1981]. The in vitro biocompatibility of PCL scaffolds was investigated by Hutmacher et al. It was found that both human fibroblasts and osteoblasts colonized the struts and bars and formed a cell-to-cell and cell-to-extracellular matrix interconnective network throughout the entire 3D honeycomb-like architecture [Hutmacher, 2000]. In an in vivo study, intramedullary pins made of PCL were implanted into a rat humerus osteotomy model. Gross post mortem examination revealed normal soft tissue and callus formation. Nonunion, lymhadenopathy, infection and sinus drainage were not seen in any of the PCL specimens. Histology verified the absence of osteolytic regression around the implant site and foreign body giant cell reactions. Decalcified humeri demonstrated osteoblastic and osteoclastic activity [Lowry, 1997]. Hence based on a large number of tests, the polymer PCL is currently regarded as non-toxic and tissue compatible materials. Besides being bioresorbable and biocompatible, the polymer can also be processed with ease into many shapes and forms [Rezwan, 2006]. All the abovementioned qualities make PCL an ideal candidate for biomedical applications including controlled drug releases and resorbable matrices as scaffolds for tissue engineering. 2.2.1 Degradation of PCL polymer Degradation behaviours of scaffolds play an essential role in the engineering of new tissue, as the rate of degradation is intrinsically linked to many cellular processes including cell viability, tissue growth, as well as the host response [Lei, 2006]. Once implanted in the body, a porous scaffold should maintain its mechanical properties and structural integrity until the ingrowth of new tissue could adapt to the environment and excrete sufficient amount of extracellular matrix. During this time, it 14 is expected of the scaffold to be largely degraded and absorbed by the body, enabling the space occupied by porous scaffolds to be replaced by the newly formed tissue [Alsberg, 2003]. Ideally, the degradation rate should match to or be slightly slower than the rate of tissue formation [Hedberg, 2005; Rai, 2006]. Different factors may affect the degradation kinetics of a scaffold. This include the chemical composition and configurational structure, processing history, porosity, polydispersity, environmental conditions, stress and strain, crystallinity, size, surface morphology, chain orientation, distribution of chemically reactive compounds within the matrix, additives, presence of original monomers and overall hydrophilicity. [Rezwan, 2006] In general, PCL, like other members of this family of aliphatic polyesters such as poly(glycolic acid) (PGA) and poly(lactic acid) (PLA), is degraded by non-enzymatic random hydrolytic scission of esters linkage [Coombes, 2004]. In the case of PCL, several reports have shown that enzymes might play a role to some extents [Gan, 1999; Jeong, 2004]. Based on the hydrophilicity of monomeric units, PGA degrades fast, PLA slow and PCL very slow. PLA is much more hydrophobic than PGA due to the additional methyl group in the structure of PLA. Hence PGA degrades much quicker in weeks time than PLA, which the latter can remain stable for over 1 year or more depending on its degree of crystallinity [Mano, 2004]. It has been found that the degradation of PCL with a high molecular weight (Mn of about 50,000) is remarkably slow, requiring 3 years for complete removal from the host body [Rezwan, 2006]. 15 Fast PGA Slow PLA PCL Figure 2.6: The degradation rate of PGA, PLA, and PCL. One of the main advantages of PCL is the non-toxic nature of the degradation products, reported mainly to be CO2 and H2O [Pitt, 1981; Woodard, 1985], making it safe for medical applications. 2.2.1.1 Hydrolysis mechanism The degradation of poly(α-hydroxy esters) in the aqueous media generally proceeds via a random, bulk hydrolysis of the ester bonds in the polymer chain. This process was mainly due to the ends of the carboxylic chains that are produced during the ester hydrolysis. During degradation, the soluble oligomers which are close to the surface leach out towards the aqueous medium faster than the chains located inside the matrix. This gradient of concentration in acidic groups then leads to the formation of a layer composed of less degraded polymer [Mano, 2004]. Woodard et al. have also extensively studied the intracellular degradation of PCL polymer [Woodard, 1985]. They reported that polymer encapsulation by collagen filaments containing only occasional giant cells was observed during the first stage (non-enzymatic bulk hydrolysis). Significant weight loss of the polymer was not observed during this stage that lasted about 9 months. After this time period, the molecular weight decreased to about 5000, followed by the onset of the second stage of degradation. The rate of chain scission slowed, the hydrolytic process began to produce short chain oligomers and weight loss was observed. Eventually the polymer was observed to 16 fragment into a powder that was observed inside the phagosomes of macrophages and giant cells [Lei, 2006]. Inside these cells, the degradation was rapid, requiring only 13 days for complete absorption in some cases. It was noted that PCL fibers were susceptible to enzymatic degradation as well. 2.2.1.2 Enzymatic degradation The studies of both in vivo and in vitro biodegradation of a given polymer are important for biomedical applications. Special research interests have also been paid to the enzymatic biodegradation [Gan, 1999]. One of the available model is the classical Michael-Menten enzymatic model. However, this model is usually valid for homogeneous systems in which both enzyme and substrate are water-soluble. Most polymers are water-insoluble, so the enzymatic degradation is more likely a heterogeneous kinetic process [Timmins, 1997]. It was proposed that those enzymes soluble in water will first bind to the polymer substrate and then slowly catalyze the hydrolytic scission of polymer chains [Mukai, 1993]. The surface area of polymer materials will then have a greater influence on the enzymatic degradation. In the case of an enzymatic biodegradation between PCL and lipase PS, the process mainly involved two essential steps: (1) the adsorption of Lipase PS onto the PCL and; (2) the interaction between Lipase PS and PCL. In principle, the second step is dependent on the characteristics of Lipase PS and PCL, while the first step is related to the total concentration of Lipase PS and PCL. It was reported that within the Lipase PS-PCL system, the degradation rate was mainly dependent on the first step [Gan, 1999]. In addition, the amount of degradation and the degradation rate of PCL depended only on the concentration of Lipase PS and independent of the PCL 17 concentration. Several results also showed that enzymatic degradation is a rapid method to study the degradation of PCL [Gan, 1999]. 2.3 TRICALCIUM PHOSPHATE (TCP) Calcium phosphates, or more accurately calcium orthophosphates, are salts of the orthophosphoric acid. They were one of the first bioceramics that were specifically developed for bone repair [Barrère, 2008]. The main driving force behind the development of calcium orthophosphates as bone substitute materials is their chemical similarity to the mineral component to mammalian bones and teeth. As a result, in addition to being non-toxic, they are biocompatible, not recognized as foreign materials in the body and, most importantly, both exhibit bioactive behavior and integrate into living tissue by the same processes active in remodeling healthy bone. They exhibit excellent bone-bonding properties that are related to the surface reactivity, via dissolution/precipitation mechanisms. This leads to an intimate physicochemical bond between the implants and bone, termed osteointegration. In addition, their degradation products are entirely metabolized in a natural way by our body [den Hollander, 1991; Lai, 2005]. These features are unique and contribute to their potential in bone tissue engineering. The first clinical attempt to use calcium phosphate compound was in the successful repair of bony defect reported by Albee and Morrison in 1920 [LeGeros, 2002]. Since then, several calcium phosphate biomaterials have been developed and used successfully in clinics. One of them is tricalcium phosphate (TCP), which belongs to 18 the categories of bioresorbable and bioactive compounds. Dental applications of tricalcium phosphate ceramics include the filling of defects due to periodontal loss, as well as repairing cleft palates. In orthopaedics, tricalcium phosphate remains an implant material for defect filling where a resorbable material is indicated [Barrère, 2008]. Tricalcium phosphate is a white crystalline powder (hexagonal crystals) Figure 2.7: The chemical structure of TCP [Chemical land, 2007]. with a melting point of 1670ºC. It is insoluble in cold water, but decomposes in hot water. 2.3.1 Degradation mechanisms of calcium phosphate ceramics In artificial or natural aqueous environments calcium phosphates can degrade via: 1. Solution-mediated mechanisms leading to physicochemical dissolution of the ceramic with possibly phase transformation 2. Cell-mediated mechanisms via macrophages and osteoclasts 3. Loss of mechanical integrity as a result of the aforementioned mechanisms. In biological systems, degradation of calcium phosphates is a combination of nonequilibrium processes that occur simultaneously or in competition with each other [Barrère, 2008]. 2.3.1.1 Physicochemical degradation The physicochemical degradation, or dissolution, of calcium phosphate ceramics can be described as a dissolution-reprecipitation cascade which is the result of exchanges at a solid-liquid interface. In artificial or natural aqueous environments these ceramics dissolve, this physicochemical process is typical of inorganic 19 substrates, i.e. having dominant ionic features. It is the result of a multi-component dynamic process that cannot be mimicked in vitro. However, in vitro experiments simplifying the biological environment have led to conclusions that, in general, fit with in vivo observations [Barrère, 2008]. From a thermodynamic point of view, most calcium phosphates are sparingly soluble in water, while some are very insoluble, but all dissolve in acids. Their solubility, defined as the amount of dissolved solute contained in a saturated solution when particles of solute are continually passing into solution (dissolving) while other particles are returning to the solid solute phase (growth) at exactly the same rate [Wu, 1998], decreases with the increase of pH [de Groot, 1983]. From a surface reactivity viewpoint, physicochemical dissolution can be seen as ionic transfer from the solid phase to the aqueous liquid via surface hydration of calcium, phosphate species and possible impurities present in the biomaterial. As a result of these ionic transfers, phase transformations occur at the ceramic surface. A phase transformation occurs for calcium phosphate phases which are unstable under physiological conditions, such as tricalcium phosphate [Barrère, 2008]. The physicochemical dissolution behavior in vitro and in vivo can be affected by the crystalline features, the thermodynamical solubility, the structure, and the presence of additives [Barrère, 2000; Elliot, 1994; LeGeros, 1991; Radin and Ducheyne, 1994]. 20 2.3.1.2 Cellular degradation The typical cellular degradation of calcium phosphates is mediated by osteoclasts. Osteoclasts are multinucleated bone cells derived from hematopoietic stem cells that differentiate along the monocyte/macrophage lineage. They are responsible for bone resorption by acidification of bone mineral leading to its dissolution and by enzymatic degradation of demineralized extracellular bone matrix. The mature osteoclast is a functionally polarized cell that attaches via its apical pole to the mineralized bone matrix by forming a tight ring-like zone of adhesion, the sealing zone. This attachment involves the specific interaction between the cell membrane and some bone matrix proteins via integrins (transmembrane adhesion proteins mediating cellsubstratum and cell-cell interactions). In the resorbing compartment, situated under the cell and delimited by the sealing zone, osteoclasts generate a milieu acidification resulting in the dissolution of bone mineral. This osteoclastic acidification is mediated by the action of carbonic anhydrase that produces carbon dioxide, water and protons that are extruded across the cell membrane into the resorbing compartment [Kartsogiannis, 2004]. The degree of osteoclastic activity and dissolution process of a calcium phosphate ceramic depends on the nature of the calcium phosphate. In the case of the degradation of highly soluble tricalcium phosphate ceramics in vivo, Zerbo et al. have shown that the degree of physicochemical dissolution was higher than osteoclastic resorption [Zerbo, 2005]. 2.3.1.3 Mechanical degradation The mechanical degradation of calcium phosphates is the result of both previous degradation mechanisms. From a mechanical point of view, the calcium phosphate ceramics are brittle polycrystalline materials for which mechanical properties are 21 governed by the grain size, grain boundaries and porosity [LeGeros, 2002]. Under humid conditions, e.g. in liquids or physiological fluids, and as a consequence of the physicochemical dissolution mechanisms, calcium phosphate ceramics undergo a decrease of mechanical strength [de Groot, 1983; Mirtchi, 1989; Pilliar, 2001, Raynaud, 1998] and of resistance to fatigue [de Groot, 1983; Raynaud, 1998]. The mechanical strength of a material can be seen as its resistance to fracture formation under specific and acute stress at a time point, while failure by fatigue includes an additional parameter which is the long-duration strength. Normally, decrease of strength of brittle ceramic materials is caused by slow or subcritical crack growth, occurring under stress, sometimes assisted by environment factors [de Groot, 1983]. Parameters influencing the mechanical strength degradation in vitro and in vivo are directly related to the parameters influencing the physicochemical dissolution [Barrère, 2008]. 2.4 PCL-TCP SCAFFOLDS Many methods have been developed to enhance the properties of biodegradable polymers in order to improve the rate of degradation and the mechanical properties. One way to do so is by physical blending [Gan, 1999]. The incorporation of a tricalcium phosphate (TCP) into a poly(ε-caprolactone) polymer matrix produces a hybrid or composite material. This bioceramic allows to tailor the desired degradation and resorption kinetics of the polymer matrix [Hutmacher, 2000]. Our team has hypothesized that the addition of TCP can 22 accelerate the degradation of the PCL polymer. It was shown that the TCP particles were only physically blended into the polymer and they occupied random spaces in the polymer. After the scaffold was immersed in solution, the TCP particles being hydrophilic tend to fall off and interact with the surrounding medium. The falling of TCP then created voids within the polymer, thus exposing their surfaces to hydrolytic attack and weakening the overall structure of the PCL [Lei, 2006]. The composite approach can circumvent the limitations of pure ceramics for example it offsets the problems of brittleness and the difficulty of shaping hard ceramic materials to fit bone defects [Hu, 2007]. A composite material would also improve biocompatibility and hard tissue integration in a way that ceramic particles, which are embedded into the polymer matrix, allow for increased initial quick spread of serum proteins compared to the more hydrophobic polymer surface [Hutmacher, 2000]. In addition, the basic resorption products of TCP would buffer the acidic resorption byproducts of the aliphatic polyester like PCL and may thereby help to avoid the formation of an unfavorable environment for the cells due to a decreased pH [Hutmacher, 2000]. Finally, the addition of ceramics dispersed throughout the polymeric matrix results in a superior compressive strength of the composite compared to non-reinforced materials. All these qualities thus, render a promising future for the application of PCL-TCP scaffolds in tissue engineering purposes. The proposed mechanism of degradation manifested by PCL-TCP composite scaffolds are expected to be a result of combining PCL polymer with TCP ceramic [Lei, 2006]. 23 2.4.1 Fabrication method of PCL-TCP scaffolds Along with the material selection, fabrication methods are also critical for designing biological scaffolds. Numerous fabrication technologies have been applied to process biodegradable and bioresorbable materials into scaffolds of high porosity and surface area. One of them is solid free-form fabrication (SFF) or rapid prototyping (RP) technique known as fused deposition modeling (FDM), which has been applied to fabricate complex-shaped tissue engineering constructs. Unlike conventional machining which involves constant removal of materials, RP is able to build scaffolds by selectively adding materials, layer-by-layer, as specified by a computer program. Each layer represents the shape of the cross-section of the computer-aided design (CAD) model at a specific level. In addition, another potential benefit offered by RP technology is the ability to create parts with highly reproducible architecture and compositional variation across the entire matrix due to its computer controlled fabrication [Hutmacher, 2004]. A traditional FDM machine consists of a head-heated-liquefier attached to a carriage moving in the horizontal x-y plane. The function of the liquefier is to heat and pump the filament material through a noozle to fabricate the scaffold following a programmed path which is based on CAD model and the slice parameters. Figure 2.8 shows a schematic representation of the FDM process. The FDM method involves the melt extrusion of filament materials through a heated nozzle and deposition as thin solid layers on a platform. The nozzle is positioned on the surface of a build platform at the start of fabrication. It is part of the extruder head (FDM 24 head), which also encloses a liquefier to melt the filament material fed through two counter-rotating rollers. Figure 2.8: Schematic diagram of the FDM process [Hutmacher, 2001]. Figure 2.9 shows how each layer is made of raster roads deposited in the x and y directions. A fill gap can be programmed between the roads to provide horizontal channels. Subsequent layers are deposited with the x-y direction of deposition, the raster angle, programmed to provide different lay-down patterns. The sequence of data preparation from step 1 to 4 is: importing of computer-aided design (CAD) data in STL (stereolithography) format into QS (QuickSlice), slicing of the CAD model into horizontal layers and conversion into SLC format, creation of a deposition path for each layer and conversion into SML format for downloading to the FDM machine, and FDM-fabrication process with a filament modeling material to build the actual 25 physical part in an additive manner layer by layer, respectively [Hutmacher, 2001; Hutmacher, 2008]. Figure 2.9: Sequence of the data preparation for FDM model fabrication [Hutmacher, 2001]. 26 The method of synthesis for PCL-TCP scaffolds used in this study was described in detail in recent literatures [Hutmacher, 2001; Rai, 2006]. The first step is filament fabrication. Pellets of PCL (catalog no. 440744) from Aldrich Chemical Co., Inc. (Milwaukee, WI) are used. The polymer has an average number-average molecular weight of 80,000 with a melt index of 1.0 g/10 min. The polymer pellets are kept in a desiccator prior to usage. PCL pellets are physically blended with TCP granules prior to filament fabrication. Filament fabrication then is performed with a fiber-spinning machine (Alex James & Associates Inc., Greensville, SC). The pellets are melted at 190°C in a cylinder with an external heating jacket. After a hold time of 15 min, the temperature is lowered to 140°C and the polymer melt is extruded through spinnerets with a die exit diameter of 0.064 in. (1.63 mm). Each batch of PCL pellets weighs about 30 ± 1 g. The piston speed is set at 10 mm/min. The extrudate is quenched in chilled water placed 40 mm below the die exit. The combination of temperature, piston speed and height drop to water quenching settings produces a filament diameter of 1.70 ± 0.10 mm. The PCL filaments are fabricated to have a consistent diameter to fit the drive wheels of the FDM system. The filaments are vacuum-dried and kept in a desiccator prior to usage. The next step is scaffold design and fabrication [Zein, 2002; Hutmacher, 2000]. The PCL filaments are fed into a FDM 3D Modeler RP system from Stratasys Inc. (Eden Prairie, MN). Stratasys Quickslice software is manipulated to produce the desired dimensions. The head speed, fill gap, and raster angle for every layer are programmed through this software and saved in the Slice file format. Lay-down patterns of 0/60/120° are used to give a honeycomb, fully interconnected matrix architecture and mechanical properties suitable for rapid vascularization and 27 maintenance of the structural integrity of tissue engineered bone grafts in loadbearing applications [Schantz 2002; Schantz, 2003; Hutmacher, 2001]. The use of the highly reproducible and computer-controlled FDM technique allows the fabrication of bone grafts that can be designed based on computed tomography (CT) scans of individual defect sites [Endres, 2003; Hutmacher, 2000]. 28 CHAPTER 3: SELECTIVE MODIFICATION OF PCL-TCP SCAFFOLDS TARGETED FOR DENTOALVEOLAR RECONSTRUCTION APPLICATION 3.1 INTRODUCTION This pre-degradation study served as an initial stage to develop a scaffold of a higher porosity that allows for a more rapid degradation whilst maintaining favourable mechanical properties. A final porosity of about 85% was targeted for. The time frame for the action of lipase and NaOH required to achieve this desired porosity on the PCL-TCP scaffold degradation was also determined. 180 small blocks of PCL-TCP scaffolds with porosity 75% were equally divided into two groups. Each composite were placed individually into clean centrifuge tubes and completely immersed in either 3M NaOH or 0.1% lipasePBS solution (Amano Lipase PS, Sigma-Aldrich, USA). The tubes were then sealed and put into an incubator of constant temperature 37ºC. Ten scaffolds were removed Figure 3.1: Centrifuge tubes. at each predetermined time intervals of 12, 24, 36, 48, 60, 72, 84, 96, and 108 hours. These samples were gently rinsed with PBS before drying for 48 hours in an incubator that maintained a temperature of 37ºC and controlled relative humidity of 29 30%. The scaffolds were then characterized in terms of their porosity, mechanical strength, surface degradation and surface morphology. 3.2 MATERIALS AND METHODS 3.2.1 Scaffold design and fabrication Scaffold specimens were fabricated with PCL-TCP (80:20%) filaments by using a fused deposition modeling (FDM) 3D Modeler RP system from Stratasys Inc (Eden Prairie, MN). Blocks of 50 x 50 Figure 3.2: Illustration of scaffold with 0/60/120º laydown pattern [Zein, 2002]. x 3 mm were purchased directly from Osteopore International Pte Ltd, Singapore. Each composite manifested a lay-down pattern of 0/60/120º with a typical honeycomb array of interconnected equilateral triangle, and porosity of about 75%. TCP existed as nonuniformly distributed particles on the rods of PCL. The specimens were then cut into smaller blocks of 5 x 5 x 3mm dimension by using a cutter under aseptic conditions. 3.2.2 Sterilization of scaffolds The raw pieces of scaffold blocks were pre–treated before used in the degradation experiment. After being rinsed 3x with phosphate buffered saline (PBS, 137mM NaCl, 2.7mM KCL, 10mM Na2HPO4, 1.8mM KH2PO4, pH7.4), the PCL-TCP scaffolds were sterilized in 70% ethanol for 24h. This was followed by rinsing twice in PBS with centrifugation at 1000 rpm for 10 min. The scaffolds were dried in humidified 30 atmosphere at 37oC and 5% CO2 for 1h and soaked for 3h in PBS before loading. This latter process of pre–wetting the porous scaffolds was to ensure that PBS solution had permeated through all the pores of the scaffolds and to let the scaffolds became more hydrophillic. Figure 3.3: 5x5x3mm PCL-TCP scaffold. 3.2.3 Scaffold characterizations Upon retrieval from the respective mediums, the scaffold specimens were subjected to characterizations for analysis of their porosity, structure, percentage weight loss, compressive mechanical properties, surface morphology, and molecular weight. Several methods employed to characterize the PCL-TCP scaffold samples were: 3.2.3.1 Micro-computed tomography analysis (n = 3) Microcomputed tomography (SkyScan 1076 In Vivo X-Ray Microtomograph, Belgium) was performed to monitor the surface and porosity changes of the scaffolds before and after implantation. The parameters were set at 104 kV energy, 98 μA intensity, and 35 μm pixel size. The specimens were scanned through 180o, and the image data from the scanned planes were subsequently thresholded and reconstructed to create 3-D images for quantitative histomorphometric analyses. 3.2.3.2 Gravimetric analysis (n = 3) Scaffolds’ weight losses during degradation were measured by the changes in dry weight after incubation or implantation for specifed time periods. Weights were normalized with respect to zero-time measurements. For such tests, three scaffolds 31 were removed, gently rinsed in PBS and dried in vacuum oven for 48 hours. Values obtained for triplicate samples were averaged. Percent weight loss was computed according to the following equations (3.1): W0 − Wdry W0 Weight loss (%) = Where ×100 (3.1) W0 is the initial dry weight as measured at time 0, and Wdry is the dry weight at time t. 3.2.3.3 Compressive mechanical testing (n = 5) Evaluation of mechanical properties of the specimens were performed with an Instron 3345 uniaxial testing system (table-top tester 3345; Instron, Canton, MA), with a 1kN load cell. Five specimens from each group were to be tested for each concentration study at each time point. Each specimen was to be placed between 2 flat plates for compressional testing. The scaffolds were compressed at a crosshead speed of 1mm/min up to 80% of the scaffolds total thickness. A stress-strain curve was plotted using the experimental data obtained to determine the compressive modulus and compressive strength. 3.2.3.4 Electron microscopy preparation and analysis (n = 2) Surface morphological changes of the PCL-TCP scaffolds were characterized using the JEOL JSM – 5600LV SEM operating at an accelerating voltage of 15kV under high vacuum mode. Prior to the usage of the SEM, the PCL-TCP scaffolds have to be pre-treated first. The scaffolds samples were immersed in 2.5% gluteraldehyde (Sigma, Germany) at 4°C overnight. They were then rinsed in PBS for 10 minutes 32 and dehydrated in a graded ethanol series of 25% (5 min), 50% (10min), 75% (10 min), 95% (10 min) and 100% (10 min, 3 times). Following that, the scaffolds were placed in an oven dessicator overnight to dry. As the scaffolds were viewed under high vacuum condition in the SEM to attain high magnification, they have to be gold coated first which is a destructive method. The gold splattering of the fracture surfaces was conducted with JEOL JFC – 1200 fine coater in a high vacuum chamber for 40 seconds at a current of 10mA. 3.2.3.5 Molecular weight testing (n = 3) Each of the scaffold samples was selected and partially dissolved into Tetrahydrofuran (THF). The THF solute was then analyzed for molecular weight distribution by Gel Permeation Chromatography (GPC) using Waters 600_717_2414 System. The measurements were carried out at an elution rate of 1 ml/min using THF as the mobile phase solvent through Styragel column refractors. A total of four 7.8 x 300mm column were used: Styragel HR 0.5, Styragel HR 1, Styragel HR 3 and Styragel HR 4. Polystyrene Standards from Waters were used to obtain the calibration curve. 3.2.4 Statistical analysis All quantitative data (the mechanical strength and molecular weight loss) were expressed as mean values ± the standard deviation (SD) of the mean. Data analyses and comparisons were performed using Student’s paired t-test. A value of p< 0.05 was considered to be statistically significant. 33 3.3 RESULTS 3.3.1 Porosity measurements and 3D model analysis Results from the micro CT analysis demonstrated an increase in porosity values in both treatment groups. Lipase-treated scaffolds showed a faster rate of degradation than NaOH-treated ones. Pre-degraded scaffolds of the desired 85% porosity were obtained when they were immersed in 3M NaOH for 96 hours and in 0.1% lipase for 12 hours (Figure 3.4). Refer to Appendix A1 – Table A.1 for the complete data. 140 3M NaOH 0.1% lipase 120 Porosity (%) 100 81.78 84.44 80 75.61 75.61 80.92 81.08 80.85 81.14 98.23 98.12 97.69 96.47 95.45 93.4 80.01 82.01 98.51 98.32 83.41 87.44 60 40 20 0 0 12 24 36 48 60 72 84 96 108 Degradation Time (Hours) Figure 3.4: Porosity measurements of NaOH-treated and lipase-treated PCL-TCP scaffolds over time. The 3 dimensional images reconstructed from these two treatment groups similarly showed a more significant degradation action by the lipase than the NaOH treatment. It has been suggested that the chemical action of sodium hydroxide mainly causes surface erosion, whereas the enzymatic action involves scission of the polymer chain backbone and hence a more significant degradation activity [Wan, 2005]. 34 Whilst attaining the desired 85% porosity, both the treated PCL-TCP scaffolds maintained favourable 3D morphology with the usual interconnected pore network (Figure 3.6 and 3.7). An untreated PCL-TCP scaffold of around 75% porosity was shown in Figure 3.5. Figure 3.5: 3D model of original scaffold (of 75% porosity) at 0 hour: (L) top view, (R) tilted view. Figure 3.6: 3D model of scaffolds after 96 hours immersion in 3M NaOH: (L) top view, and (R) tilted view. 35 Figure 3.7: 3D model of scaffolds after 12 hours immersion in 0.1% lipase: (L) top view, and (R) tilted view. 3.3.2 Weight loss analysis There is an increasing trend for the weight-loss percentage as shown in Figure 3.8. 100 82.3 Weight loss (%) 80 92.77 91.28 90.36 90 92.8 90.61 85.66 73.71 70 60 52.8 50 39.48 45.13 40 30 15.76 20 10 22.33 20.32 21.94 24.59 26.18 17.06 3M NaOH 0.1% lipase 0 0 0 0 10 20 30 40 50 60 70 80 90 100 110 120 Degradation Time (Hours) Figure 3.8: Comparison of weight loss between NaOH-treated and lipase-treated PCL-TCP scaffolds. The values for NaOH-treated at 96 hours and lipase-treated at 12 hours scaffolds were found to increase logarithmically to 39.48±2.25 % and linearly to 45.13±7.41 % 36 respectively. These values demonstrated a significant loss of mass as the scaffolds were degraded over the course of time. Refer to Appendix A4 – Table A.4 for the complete data. 3.3.3 Compressive mechanical properties The mechanical properties of the treated PCL-TCP scaffolds over time are shown in Figures 3.9 and 3.10. For the NaOH-treated scaffolds at 96 hours, both the compressive strength and compressive modulus have decreased significantly by 41.6% (4.6±0.8 MPa) and 51.3% (10.8±2.75) MPa respectively. Likewise for the lipase-treated scaffolds, at 12 hours, both the compressive strength and compressive modulus have reduced significantly by 44.7% (4.36±1.64 MPa) and 46.4% (11.87±2.58 MPa) respectively. As expected, the reduction in mechanical properties was accompanied following in increases in the scaffold porosity. Refer to Appendix A2 – Table A.2 and Table A.3 for the complete data. Compressive Strength (MPa) 10 8 3M NaOH 0.1% lipase 7.88 7.01 6.85 6.94 7.25 6.90 6.66 7.21 6 4.60 4.36 4 3.19 2.35 2 2.22 1.57 0 0 0 0 10 20 30 40 50 60 70 0 80 0 90 100 0 110 -2 Degradation Time (Hours) Figure 3.9: Compressive strength of PCL-TCP scaffolds when treated with NaOH and lipase at pre-determined time intervals. 37 120 30 Compressive Modulus (MPa) 3M NaOH 25 0.1% lipase 22.15 20 19.47 19.00 20.17 18.07 17.32 17.48 15.38 15 11.87 10.78 10 7.98 7.15 7.10 5.65 5 0 0 0 10 20 30 40 50 60 0 70 0 80 0 90 100 0 110 120 -5 Degradation Time (Hours) Figure 3.10: Compressive modulus of PCL-TCP scaffolds when treated with NaOH and lipase at pre-determined time intervals. 3.3.4 Surface morphology analysis Electron microscopy analysis was conducted to investigate the surface morphology of the PCL-TCP scaffolds. Figure 3.11 shows the SEM micrograph of the original untreated scaffold (that is, scaffold at 0 hour). In general, the surface was smooth with numerous small TCP particles protruding out. Figure 3.11: Electron micrographs of original scaffold (of porosity 75%) at 0 hour: (L) overall view, and (R) higher-magnification view. 38 Figures 3.12 and 3.13 display SEM micrographs taken at 96 hours immersion in NaOH and at 12 hours in lipase respectively. Viewed under higher magnification, it could be observed that the smooth surface of the scaffolds have roughened and displayed small cracks in both cases. Numerous tiny and shallow pores have appeared as well. However, the degree of the degradation seemed to be slightly higher in lipase-treated scaffolds as seen from the more significant reduction in the diameter of the rods. Refer to Appendix A3 for the complete images. Figure 3.12: Electron micrographs of scaffold after 96 hours immersion in 3M NaOH: (L) overall view, and (R) higher-magnification view. Figure 3.13: Electron micrographs of scaffold after 12 hours immersion in 0.1% lipase: (L) overall view, and (R) higher-magnification view. 39 3.3.5 Molecular weight analysis The values of Mw, Mn, and PDI are shown in Table 3.1. Overall, the NaOH-treated PCL–TCP scaffolds exhibited no appreciable decrease in molecular weight throughout the degradation period of up to 108 hours. At 96 hours immersion, the value of Mw, Mn, and PDI is 58940 Daltons, 41279 Daltons, and 1.43 respectively. On the other hand, PCL-TCP scaffolds immersed in lipase maintained their molecular weights for up to 72 hours and showed a sudden drop when it reached 96 hours with nearly 50% reduction in the values. At 12 hours, the lipase-treated PCLTCP scaffolds declined in its Mw, Mn, and PDI to 54491 Daltons, 39763 Daltons, and 1.37 respectively. Table 3.1: Mw, Mn, and PDI of NaOH-treated and lipase-treated PCL-TCP scaffolds. Hour 0 Hour 12 NaOH-treated PCL-TCP Scaffolds 54548 54457 Mw (Dalton) Hour 24 Hour 48 Hour 72 Hour 96 56575 55330 59849 58940 Mn (Dalton) 40415 38010 42283 40090 43041 41279 PDI 1.35 1.43 1.34 1.38 1.39 1.43 54180 50364 54526 29337 Lipase-treated PCL-TCP Scaffolds 53548 54491 Mw (Dalton) Mn (Dalton) 40415 39763 39842 34859 38529 24667 PDI 1.35 1.37 1.36 1.44 1.42 1.19 3.4 DISCUSSION In this part of the study, sodium hydroxide solution and lipase enzyme were used to treat the original 75% porosity scaffolds with the objective of obtaining customized 40 scaffolds of higher porosity value. The desired 85% porosity was targeted for in this study. The concentration of sodium hydroxide and lipase was selected as 3M and 0.1% respectively, which was considered low-medium in terms of the strength. The scaffolds were allowed to degrade but in a slower and controlled manner. This was to ensure that the treated scaffolds still retained most of their essential properties such as strength, stiffness, pores interconnectivity, molecular weight, and others, prior immersion in the DMEM growth media or implantation in vivo in rats. Sodium hydroxide solution is a strong alkaline reagent. Previous studies have shown that it has the capability to enhance the hydrophilicity and accelerate the degradation of PCL polymer under laboratory condition [Park, 2005]. However not many have specifically focused on PCL-TCP scaffolds. The non-enzymatic breakdown of the scaffolds due to sodium hydroxide involves mainly surface erosion. Reports have also demonstrated that PCL, as opposed to its related aliphatic polymers, such as PGA and PLA, have the capability to undergo enzymatic degradation [Gan, 1999]. Among the various types of enzymes, lipase is found to be the most widely investigated, and that lipase PS is considered the best candidate [Gan, 1997; He, 2003]. And for this reason, Lipase PS was chosen in this experiment. In the case like this study, whereby the enzyme was water-soluble but the PCL-TCP composite as substrate was water-insoluble, the degradation mechanism will most likely to follow the heterogenous kinetics model. The model proposed that the enzyme being soluble in water will first bind to the polymer substrate, and then catalyze the hydrolytic scission of polymer chains [Mukai, 1993]. Due to this catalytic property, thus it is hypothesized that once occurred, enzymatic degradation is a rapid process. One report showed that the biodegradation of a macroscopic PCL film with a 41 dimension of 10 x 10 x 0.1 mm3 could be completed within 4 days in a buffer solution containing 5.0 x 10-4 g/ml Lipase PS [Gan, 1997]. As expected, it has been demonstrated in this experiment that lipase-treated scaffolds are the most rapidly degraded compared to NaOH-treated scaffolds within the same degradation time frame. After 7-8 days of immersion in lipase solution, a total breakdown of the PCLTCP scaffolds was actually observed. Findings from the pre-degradation section of the study also demonstrated that scaffolds with the desired 85% porosity were obtained after being submerged in 3M NaOH for 96 hours (83.41±3.04 %) and in 0.1% lipase for 12 hours (84.44±2.80 %). Porosity is one of the key parameters to be considered when designing a scaffold as it would determine degradation rate, cell seeding efficiency, diffusion and the mechanical strength of the scaffold [Rai, 2006]. A high porosity and high surface area to volume ratio are required for uniform cell delivery, cellular attachment and neo-tissue in growth. However, it was reported that scaffolds with a too high porosity would possess very low mechanical properties. Based on this, compressive mechanical testings were conducted and the results showed that the decrease in the strength and stiffness recorded for the NaOH-treated and lipase-treated scaffolds at 85% porosity were still within the acceptable range to support load bearing tissues such as bone and cartilage. For the NaOH-treated scaffolds at 96 hours, the compressive strength and stiffness decreased by 41.6% (4.6±0.8 MPa) and 51.3% (10.8±2.75) MPa respectively. In the lipase-treated group, both the compressive strength and compressive modulus of the scaffolds were reduced by 44.7% (4.36±1.64 MPa) and 46.4% (11.87±2.58 MPa) respectively after 12 hours. 42 Meanwhile, the treated scaffolds still retained their usual framework with proper interconnected pore networks as desired. This is advantageous as a high pore interconnectivity would encourage the development of a vascular network to allow for the infiltration of nutrients, cells and growth factors within the system in vivo. As for the molecular weights, the results for NaOH-treated scaffolds were relatively constant and exhibited no particular trend throughout the degradation period of up to 108 hours. This finding then further confirmed that the chemical action of sodium hydroxide mainly involves surface erosion. In contrast, lipase enzyme causes scission of the polymer chain backbone and as observed, the molecular weights of the lipase-treated scaffolds showed a decreasing trend and demonstrated a fall when it reached a certain time point. In addition, scanning electron microscopy were conducted and the results displayed that there is an increase in the surface roughness of the treated scaffolds as clearly shown from the increase in indentations on the surfaces of the scaffold rods. This is an encouraging morphology as bone forming cells, like osteoblasts, favour rough surfaces which is optimal for promoting better cell attachment, proliferation, and differentiation [Park, 2005]. 3.5 CONCLUSION PCL-TCP scaffolds have shown to undergo both chemical and enzymatic degradation in this part of the study. The objective of the experiment was achieved with scaffolds of approximately 85% porosity obtained after 96 hours of treatment in 3M NaOH and 12 hours in 0.1% lipase. These pre-treated scaffolds demonstrated acceptable mechanical strength, structure, and surface morphology. 43 CHAPTER 4: OPTIMIZATION OF NATIVE AND CUSTOMIZED SCAFFOLDS IN VITRO AND THEIR EFFECTS IN INITIAL BONE HEALING 4.1 INTRODUCTION In Chapter 3, the PCL-TCP scaffolds have been subjected to both chemical and enzymatic degradation. Scaffolds with desirable higher porosity with acceptable mechanical strength, structure, and surface morphology were successfully obtained. However the degradation behavior of these customized scaffolds must be further tested. It was then the interest of this present study to investigate the degradation kinetics of treated and untreated PCL-TCP scaffolds when immersed in standard culture medium and when implanted subcutaneously to the back of rats. Both these in vitro and in vivo studies were described under this chapter for better comparison purpose. 4.1.1 In vitro degradation study The in vitro degradation study of the PCL–TCP scaffolds was done closely in accordance to the ISO 10993 standards. The medium used was Dulbecco’s Modified Eagle Media (DMEM D1152, Sigma, USA. Refer to Appendix B.5 – Figure B.19) growth medium. In this study, three group of scaffolds namely native, sodium hydroxide-treated, and lipase-treated of 85% porosity were fully submerged separately in 1ml DMEM solution. The samples were placed individually in small 44 cryogenic tubes and they were stored in an oven desiccator at a temperature of 37ºC inside the cleanroom for the duration of 6, 12, 18, and 24 weeks. 10 samples were taken and characterized at each time points. In addition, the DMEM solution was replaced with a fresh batch twice a week (every Monday and Thursday) to avoid any built-up excess or bacterial contamination. Figure 4.1: Native (left), NaOH-treated (middle), and lipase-treated (right) scaffolds. 4.1.2 In vivo degradation study In order to acquire better undertanding of the degradation behaviour of the scaffolds in the living system and prior conducting larger animal studies, an in vivo study on smaller animal model is a necessity. In this project, rat models were chosen. A total of 24 untreated PCL-TCP (native) and 48 treated (24 lipase-treated and 24 NaOHtreated) scaffolds with porosity of 85% were implanted in the subcutaneous backs of rats for the duration of 3 and 6 months. 45 Figure 4.2: Rat at the start of experiment (left) and at the end after 6 months (right). 4.2 MATERIALS AND METHODS 4.2.1 Scaffold design and fabrication Scaffold specimens were fabricated with PCL-TCP (80:20%) filaments by using a fused deposition modeling (FDM) 3D Modeler RP system from Stratasys Inc (Eden Prairie, MN). Blocks of 50 x 50 x 3 mm were purchased directly from Osteopore International Pte Ltd, Singapore. Each composite Figure 4.3: 50x50x3mm PCL-TCP scaffold. manifested a lay-down pattern of 0/60/120º with a typical honeycomb array of interconnected equilateral triangle, and porosity of about 75%. TCP existed as non-uniformly distributed particles on the rods of PCL. The specimens were then cut into smaller blocks of 5 x 5 x 3mm dimension by using a cutter under aseptic conditions. 46 4.2.2 Sterilization of scaffolds The raw pieces of scaffold blocks were pre–treated before used in the degradation experiment. After being rinsed 3x with phosphate buffered saline (PBS, 137mM NaCl, 2.7mM KCL, 10mM Na2HPO4, 1.8mM KH2PO4, pH7.4), the PCL-TCP scaffolds were sterilized in 70% ethanol for 24h. This was followed by rinsing twice in PBS with centrifugation at 1000 rpm for 10 min. The scaffolds were dried in humidified atmosphere at 37oC and 5% CO2 for 1h and soaked for 3h in PBS before Figure 4.4: 5x5x3mm PCL-TCP scaffold. loading. This latter process of pre–wetting the porous scaffolds was to ensure that PBS solution had permeated through all the pores of the scaffolds and to let the scaffolds became more hydrophillic. 4.2.3 Animal husbandry Twelve, 7-8 week old male Wistar rats (260.83 ± 39.43 g) were employed for this in vivo study. The animals were housed in the animal holding facility at the SingHealth Experimental Medicine Centre (SEMC), Singapore General Hospital, for the entire duration of the experiment. Housing and feeding were according to standard animal-care protocols. The study has been approved by the Animal Welfare Figure 4.5: Rat cages. Committee of the Singapore General Hospital and 47 has been licensed by the National Institute of Health’s Guide for Care and Use of Laboratory Animals. 4.2.4 Scaffold implantation The rats were randomly divided into two groups of 6 rats each for the implantation of scaffolds. They were operated on under general anaesthesia which consists of an intraperitoneal injection of ketamine and xylazine mixture (75 mg/kg + 10 mg/kg). Under anesthesia, the neck and back region of the rat was shaved and Figure 4.6: Rat shaved and scrubbed with iodine. scrubbed with iodine, followed by disinfection with 70 % ethyl alcohol (shown in Figure 4.6). A midline incision was made in the skin of the back with scissors paralleling the line connecting the apical most point of the shoulder blades and to the point most coronal to the line. A cylinder shaped dissection was then made subdermal with a pair of needle holders and the cavities were exposed for the placement of the samples in the right and left flanks of the dorsal lumbar regions. For each rat, 6 samples of either treated or untreated PCL-TCP scaffolds (2 from each group) with dimensions of 5 x 5 x 3 mm were placed into the empty spaces in the cavity Figure 4.7: Scaffolds’ positioning. and the skin was closed with sutures. The rats were given buprenorphine (100-500 ug/kg) and cephalexin (15-20 mg/kg) subcutaneously for 3 and 5 days respectively. 48 Figure 4.8: Incision made (left), implanted scaffold (left, inset), and scaffold positions (right). A group of 6 rats were euthanised by an overdose of carbon dioxide inhalation at 3 and 6 months respectively. The tissues surrounding the implanted scaffolds were removed and fixed in either 10 % neutral buffered formalin or saline solution. After sufficient fixing, the samples were processed accordingly for their degradation and physical properties. Figure 4.9: Sacrifice of rats. Figure 4.10: Removal of scaffolds. 4.2.5 Scaffold characterizations Upon retrieval from the respective mediums, the scaffold specimens were subjected to characterizations for analysis of their porosity, structure, percentage weight loss, 49 compressive mechanical properties, surface morphology, and molecular weight. Several methods employed to characterize the PCL-TCP scaffold samples were: 4.2.5.1 Micro-computed tomography analysis (n = 3) Please refer to section 3.2.3.1 for the details. 4.2.5.2 Gravimetric analysis (n = 3) Please refer to section 3.2.3.2 for the details. 4.2.5.3 Compressive mechanical testing (n = 5) Please refer to section 3.2.3.3 for the details. 4.2.5.4 Electron microscopy preparation and analysis (n = 2) Please refer to section 3.2.3.4 for the details. 4.2.5.5 Molecular weight testing (n = 3) Please refer to section 3.2.3.5 for the details. 4.2.5.5 Histological preparation and analysis (n = 2) The two specimens removed after 3 months and stored in neutral buffered formalin were dehydrated in ascending series of alcohol rinses and embedded using a process that produced ground sections with the glycol metacrylate resin. Once polymerized, the block was trimmed to remove excess plastic with an industrial vertical band saw and cut along its long axis with a diamond band saw (EXAKT standard saw). Ground polished sections of 10 µm thickness were made using the 50 EXAKT micro grinder system from EXAKT Technologies, Inc., Oklahoma City, OK. Two slides were created for each scaffold. The slides were stained with Hematoxylin & Eosin (H & E). 4.2.6 Statistical analysis All quantitative data (the mechanical strength and molecular weight loss) were expressed as mean values ± the standard deviation (SD) of the mean. Data analyses and comparisons were performed using Student’s paired t-test. A value of p< 0.05 was considered to be statistically significant. 4.3 RESULTS - IN VITRO DEGRADATION STUDY 4.3.1 Porosity measurements and 3D model analysis Upon retrieval from the growth culture medium at each respective time points, the porosity of the PCL-TCP scaffolds were measured. The graph as shown in Figure 4.11 revealed an increasing trend in the porosity of all three scaffold groups. The increasing rate in the porosity of the lipase-treated scaffolds has been noted to be the highest, followed by NaOH-treated scaffolds and native scaffolds respectively. After 6, 12, and 18 weeks of immersion, the porosity change of the lipase-treated scaffolds is 9.82%, 13.49%, and 10.37% respectively. For NaOH-treated scaffolds, the porosity change after 6, 12, and 18 weeks of immersion is 4.02%, 2.77%, 4.62% respectively. The porosity change of the native scaffolds after 6, 12, and 18 weeks of 51 immersion was the least among the three, which is 1.03%, 4.18%, and 3.27% respectively. Refer to Appendix B1 – Table B.1 for the complete data. 140 Native NaOH-treated 120 Lipase-treated Porosity (%) 100 92.73 85.24 83.41 84.44 86.12 86.76 95.83 88.80 85.72 88.03 87.26 93.20 80 60 40 20 0 0 6 12 18 Degradation Time (Weeks) Figure 4.11: Porosity measurements of native, NaOH-treated, and lipase-treated PCL-TCP scaffolds after immersion in DMEM for 6, 12, and 18 weeks. Figure 4.12: 3D model of native scaffold (of 85% porosity) at week 0: (L) top view, and (R) tilted view. 52 Figure 4.13: 3D model of native scaffold after 6 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.14: 3D model of NaOH-treated scaffold after 6 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.15: 3D model of lipase-treated scaffold after 6 weeks immersion in DMEM: (L) top view, and (R) tilted view. 53 Figure 4.16: 3D model of native scaffold after 12 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.17: 3D model of NaOH-treated scaffold after 12 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.18: 3D model of lipase-treated scaffold after 12 weeks immersion in DMEM: (L) top view, and (R) tilted view. 54 Figure 4.19: 3D model of native scaffold after 18 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.20: 3D model of NaOH-treated scaffold after 18 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.21: 3D model of lipase-treated scaffold after 18 weeks immersion in DMEM: (L) top view, and (R) tilted view. 55 Figure 4.22: 3D model of native scaffold after 24 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.23: 3D model of NaOH-treated scaffold after 24 weeks immersion in DMEM: (L) top view, and (R) tilted view. Figure 4.24: 3D model of lipase-treated scaffold after 24 weeks immersion in DMEM: (L) top view, and (R) tilted view. 56 Figure 4.12 displayed the 3D construct of native PCL-TCP scaffold which was manufactured of 85% porosity at week 0 (before immersion in DMEM). For the 85% porosity NaOH-treated and lipase-treated scaffolds used at week 0 (prior immersion in DMEM), please refer to Figure 3.6 and 3.7 respectively. Figure 4.13 to 4.24 illustrate the 3D models of ±85% porosity native, NaOH-treated, and lipase-treated PCL-TCP scaffolds after being submerged for 6, 12, 18, and 24 weeks in DMEM. There were obvious changes to the 3D morphology of the scaffolds. The rods of the scaffolds of all three groups have became thinner and looked fragile over time. In the case of native scaffolds, up to 18 weeks of immersion, the finding shows that they still retained the original framework. The same goes to NaOHtreated scaffolds, although slight disconnectivity of the rods started to be seen only later at week 18. In contrast to these two groups, lipase-treated scaffolds seemed to have undergone significant structural distortion. Loss of original shape and honeycomb-like pattern, as well as disconnectivity of the rods, were observed as early as 6 weeks of immersion. By week 24, all scaffolds were significantly degraded. 4.3.2 Weight loss analysis Figure 4.25 show the percentage weight loss of scaffolds in DMEM growth medium over various time intervals. Overall there was an increase in the percentage of weight loss for the native, NaOH-treated, and lipase-treated scaffolds. Again, the highest being lipase-treated scaffolds, which have undergone significant degradation during the immersion. The huge loss of mass was mainly due to the reduction in volume caused by the disconnection of the rods, which correlate to the results 57 obtained from the Micro-CT analysis as discussed earlier. Apart from lipase-treated scaffolds, the other two scaffold groups have demonstrated an increase in the weight loss as well. Although the amount is much lower compared to lipase-treated scaffolds, the student’s t-test show that the rise in percentage weight loss for native and NaOH-treated scaffolds was statistically significant. Refer to Appendix B4 – Table B.7 for the complete data. 100 Native 90 NaOH-treated 76.77 80 Lipase-treated 70.77 66.11 Weight loss (%) 70 60 50 40 30 20 22.38 14.29 13.07 6.23 10 10.89 6.56 0 6 12 18 -10 Degradation Time (Weeks) Figure 4.25: Weight loss of PCL-TCP Scaffolds In vitro. 4.3.3 Compressive mechanical properties Compressive properties of degrading porous PCL-TCP scaffolds from three different groups as a function of degradation time are illustrated in Figure 4.26 and 4.27. 58 Figure 4.26: Relative compressive strength of PCL-TCP Scaffolds In vitro. Figure 4.27: Relative compressive modulus of PCL-TCP Scaffolds In vitro. Referring to the graphs, the mechanical properties were shown to decrease over the course of the study up to 18 weeks. But there are variations in the decreasing rate in the properties across the three groups. While compressive strength and modulus of 59 the native scaffold decline in a gradual manner (close to exponentially), those of the NaOH-treated and lipase-treated scaffolds behaved much different within the same observation period. The latter two demonstrated a drastic fall between week 0 and week 6, with reduction in compressive strength and modulus by 57.17% and 45.08% respectively for NaOH-treated scaffolds, and by 72.94% and 84.08% respectively for lipase-treated scaffolds. This sharp decline have then supported the finding in the previous section and could be explained that it is likely due to the extreme weight loss of the scaffolds during the degradation process. However interestingly between week 6 up to week 18, the compressive strength and modulus of both NaOH-treated and lipase-treated scaffolds seemed to stall around their value at week 6, with only slight fluctuations observed. By the end of the study at week 18, the compressive strength of the native, NaOH-treated, and lipase-treated scaffolds have diminished by 41.89%, 55%, and 70.64% respectively. As for the stiffness, the values for native, NaOH-treated, and lipase-treated scaffolds have reduced by 48.45%, 37.94%, and 88.04% respectively. Refer to Appendix B2 – Table B.3 and Table B.4 for the complete data. 4.3.4 Surface morphology analysis The PCL-TCP scaffolds immersed in culture medium demonstrated significant changes to their surface morphology over time as depicted in Figure 4.28 to 4.31. At week 0, all the three scaffold groups had a consistent interconnected architecture. For native scaffolds (manufactured 85% porosity), the TCP was evident as particles protruding out of the rods’ surfaces, contributing to the rough texture (Figure 4.32). Whereas for the other two groups, their surfaces have some cracks and microporous 60 as resulted from being treated with NaOH and lipase prior immersion in the growth media (refer to Figure 3.12 and 3.13 respectively). After soaking in growth medium for 6 weeks, thinning of the rods were observed across the three groups, and more significantly in NaOH-treated and lipase-treated scaffolds. In fact for the case of lipase-treated scaffolds, which has undergone the highest degradation rate, disconnectivity of the rods can actually be seen as early as 6 weeks (Figure 4.28 (e) and (f)). Viewed under higher magnification, it is observed that the surface of the rods of native scaffolds have roughened to a small extent, but no significant micropores can be seen yet. The surface of the rods in NaOH-treated and lipase-treated scaffolds still remained microporous however the size of the pores have slightly increased. This implies that to a certain extent, degradation of the scaffolds have occured. Refer to Appendix B3 for the complete images. At week 12, the surfaces of the scaffold from all groups seemed to be increasingly eroded, enhancing the surface contact area with the culture medium. The degradation of the scaffolds continued to accelerate and by the end of week 24, there was an obvious shrinkage in the size of the three scaffold groups indicating significant loss of material. Surface line cracks with considerable gaps, and slight distortion has started to appear on the rods of native scaffolds. In addition, the lipase-treated scaffolds have severely lost their architexture, although those of the other two groups still retained their framework. 61 (a) (b) (c) (d) (e) (f) Figure 4.28: Electron micrographs taken after 6 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 62 (a) (b) (c) (d) (e) (f) Figure 4.29: Electron micrographs taken after 12 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 63 (a) (b) (c) (d) (e) (f) Figure 4.30: Electron micrographs taken after 18 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 64 (a) (b) (c) (d) (e) (f) Figure 4.31: Electron micrographs taken after 24 weeks immersion in DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 65 Figure 4.32: Electron micrographs of native scaffold (of 85% porosity) at week 0: (L) overall view, and (R) higher-magnification view. 4.3.5 Molecular weight analysis Table 4.1: Mw, Mn, and PDI of native, NaOH-treated, and lipase-treated PCL-TCP Scaffolds in vitro. Week 0 6 12 55489 40023 1.39 55905 40880 1.37 NaOH-treated PCL-TCP scaffolds in DMEM 58940 50213 Mw (Dalton) 41279 36788 Mn (Dalton) 1.43 1.36 PDI 53315 37225 1.43 Lipase-treated PCL-TCP scaffolds in DMEM 54491 51717 Mw (Dalton) 39763 35215 Mn (Dalton) 1.37 1.47 PDI 52769 36598 1.44 Native PCL-TCP scaffolds in DMEM 56437 Mw (Dalton) 42101 Mn (Dalton) 1.34 PDI The molecular weights of the PCL-TCP scaffolds immersed in culture medium were measured with gel permeation chromatography for the degradation of up to 12 weeks and the results were tabulated as given in Table 4.1. The molecular weights remained relatively constant for all the three scaffold groups, with only slight decreasing trend observed for NaOH-treated scaffolds and lipase-treated scaffolds. 66 4.4 RESULTS - IN VIVO DEGRADATION STUDY 4.4.1 Porosity measurements and 3D models analysis After 12 and 24 weeks of implantation in rats, the PCL-TCP scaffolds were retrieved and their porosity was measured. The graph as shown in Figure 4.33 revealed that the porosity values demonstrated an increasing trend for all three scaffold groups. The increasing rate in the porosity of the lipase-treated scaffolds has been found to be the highest, followed by NaOH-treated scaffolds and native scaffolds respectively. At 12 and 24 weeks, the change in the porosity percentage recorded were 0.11% and 2.28% for native scaffolds, 2.42% and 5.56% for NaOH-treated scaffolds, and 5.97% and 6.87% for lipase-treated scaffolds respectively. Refer to Appendix B1 – Table B.2 for the complete data. 150 Native NaOH-treated 120 Porosity (%) Lipase-treated 90 85.24 83.41 84.44 85.33 85.43 89.48 87.18 88.05 60 30 0 0 12 Degradation Time (Weeks) Figure 4.33: Porosity of PCL-TCP Scaffolds In vivo. 67 24 90.24 Figure 4.34: 3D model of native scaffold after 3 months implantation: (L) top view, and (R) tilted view. Figure 4.35: 3D model of NaOH-treated scaffold after 3 months implantation: (L) top view, and (R) tilted view. Figure 4.36: 3D model of lipase-treated scaffold after 3 months implantation: (L) top view, and (R) tilted view. 68 Figure 4.37: 3D model of native scaffold after 6 months implantation: (L) top view, and (R) tilted view. Figure 4.38: 3D model of NaOH-treated scaffold after 6 months implantation: (L) top view, and (R) tilted view. Figure 4.39: 3D model of lipase-treated scaffold after 6 months implantation: (L) top view, and (R) tilted view. 69 4.4.2 Weight loss analysis The percentage weight loss of the PCL-TCP scaffolds harvested from rats after implantation for 12 and 24 weeks were shown in Figure 4.40. As expected, the graph revealed an increasing trend for all the three scaffold groups. After 12 weeks, the percentage loss in weight for native, NaOH-treated, and lipase-treated scaffolds were 2.54%, 7.59%, and 6.63% respectively. As the degradation period increased, the weight loss percentage also increased. A sudden leap in the weight loss indicating a significant loss of mass were observed at 24 weeks for lipase-treated scaffolds. By the end of the in vivo study, the weight loss recorded were 4.12%, 6.14%, and 17.98% for native, NaOH-treated, and lipase-treated scaffolds respectively. Refer to Appendix B4 – Table B.8 for the complete data. 35 30 Weight loss (%) 25 17.98 20 15 7.59 10 5 6.63 4.12 2.54 6.14 0 Native -5 NaOH-treated -10 Lipase-treated -15 12 24 Degradation Time (Weeks) Figure 4.40: Weight loss of PCL-TCP Scaffolds In vivo. 70 4.4.3 Compressive mechanical properties The compressive properties of the scaffolds as a function of degradation time are depicted in Figure 4.41 and 4.42. Both the compressive strength and modulus have decreased significantly during the first 12 weeks of implantation and then continued declining at a much slower rate as the degradation period increased. After 3 months of implantation, the compressive strength of native, NaOH-treated, and lipasetreated scaffolds have reduced by 13.41%, 49.78%, and 57.34% respectively (Figure 4.41). Likewise, the compressive modulus have decreased by 25.15%, 37.01%, and 62.93% for native, NaOH-treated, and lipase-treated scaffolds respectively (Figure 4.42). By the end of the study at 24 weeks, the decrease in the compressive strength and modulus recorded were 60.89% and 82.61% for native scaffolds, 52.61% and 46.75% for NaOH-treated scaffolds, and 58.03% and 69.25% for lipase-treated scaffolds respectively. Refer to Appendix B2 – Table B.5 and Table B.6 for the complete data. Figure 4.41: Compressive strength of PCL-TCP Scaffolds In vivo. 71 Figure 4.42: Compressive modulus of PCL-TCP Scaffolds In vivo. 4.4.4 Surface morphology analysis Figures 4.43 and 4.44 displayed SEM micrographs of the three scaffold groups after implantation for 12 and 24 weeks in vivo, under low and high magnifications respectively. The honeycomb-like pattern of triangular pores present in scaffolds at week 0 was lost at week 12 and subsequently at week 24, with the diagonal rod-like structures appeared to be melted and fused together. It is observed as well that the rods’ diameter have decreased in size as the degradation occured. This is particularly obvious for the lipase-treated scaffolds whose surface has became rough and highly distorted (Figure 4.43 and 4.44 (e) and (f)). This correlates with the findings from micro-CT analysis. Refer to Appendix B3 for the complete images. 72 (a) (b) (c) (d) (e) (f) Figure 4.43: Electron micrographs taken after 3 months implantation: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 73 (a) (b) (c) (d) (e) (f) Figure 4.44: Electron micrographs taken after 6 months implantation: (a,b) native, (c,d) NaOH-treated, and (e,f) lipase-treated scaffolds. (L) overall view, and (R) higher-magnification view. 74 4.4.5 Molecular weight analysis The tabulated molecular weight values for the PCL-TCP scaffolds tested in vivo (up to 12 weeks) were shown in Table 4.2. In general, the values obtained did not reveal any trend. The molecular weights remained relatively constant for all the three scaffold groups. Table 4.2: Mw, Mn, and PDI of native, NaOH-treated, and lipase-treated PCL-TCP Scaffolds in vivo. Month 0 3 Native PCL-TCP scaffolds in rat 56437 Mw (Dalton) 61054 Mn (Dalton) 42101 48094 PDI 1.34 1.27 NaOH-treated PCL-TCP scaffolds in rat 58940 Mw (Dalton) 55875 Mn (Dalton) 41279 41002 PDI 1.43 1.36 Lipase-treated PCL-TCP scaffolds in rat 54491 Mw (Dalton) 56240 Mn (Dalton) 39763 46196 PDI 1.37 1.22 75 4.4.6 Histology analysis Slides of horizontal cross-sections of the scaffold blocks were prepared and stained with Hematoxylin and Eosin (H & E) upon retrieval of the three scaffold groups after 3 months implantation in rat models. Figure 4.45 up to 4.50 displays (a) 5x representative slides of the native, N a O H - t r e a t e d , a n d l i p a s e - t re a t e d scaffolds after H & E staining. The pink and blue colour seen is a positive stain of cytoplasm due to Eosin and of nuclei due to Hematoxylin respectively. As observed, soft tissues, specifically (b) 20x adipose tissues as identified by their Figure 4.45: H&E stain of native scaffolds after 3 months implantation. distinct anatomical features, were (a) 10x (b) 20x Figure 4.46: H&E stain of native scaffolds after 6 months implantation. 76 evident throughout the scaffold. They were found covering the surface of the rods as well as infiltrating the pores of the scaffolds. Moreover, several vascular vessels were also detected. Some of them were highlighted in orange circle in Figure 4.45 (b), 4.47(b), and 4.49(b). No (a) 5x overt microscopic signs of inflammatory cells or fibrous encapsulation were recorded. Multinucleated giant cells or macro-phages were also absent. The lower magnification images captured were also able to provide a rough demonstration of the density of the rods (b) 20x of the three scaffold groups. The blacker Figure 4.47: H&E stain of NaOHtreated scaffolds after 3 months implantation. the region is observed (as pointed by the (a) 10x (b) 20x Figure 4.48: H&E stain of NaOH-treated scaffolds after 6 months implantation. 77 green arrows), the higher the density of the rod is, as white region constitute empty spaces. Out of the three groups, NaOH-treated scaffolds appear to have the highest density (Figure 4.47(a)), followed by native scaffolds (Figure 4.45(a)) and lipasetreated scaffolds (Figure 4.49(a)). (a) 5x (b) 40x Figure 4.49: H&E stain of lipase-treated scaffolds after 3 months implantation. (a) 10x (b) 20x Figure 4.50: H&E stain of lipase-treated scaffolds after 6 months implantation. 78 4.5 DISCUSSION The first objective for this chapter of study was to monitor the in vitro degradation profile of PCL-TCP scaffolds when immersed in standard culture medium that was used as it contains vitamins, amino acid and glucose and is closer to the human blood composition [Rai, 2006; Cheong, 2005]. The findings then would be more clinically relevant. Results demonstrated that when immersed in DMEM culture medium, PCL-TCP scaffolds in all three experimental groups displayed distinct degradation behaviours. Micro-CT analysis revealed that although there is an increasing trend in the porosity for all three scaffold groups, the rate has been noted to be the highest for lipase-treated scaffolds, and then followed by NaOH-treated scaffolds and native scaffolds respectively. By the end of week 18, the porosity has reached to considerably high values of 93.20%, 87.26%, and 88.03% respectively. This is in accordance to the gravimetric analysis, which showed a relatively huge jump in the weight loss percentage of the lipase-treated as compared to the other two groups. Further increase in surface roughness and thinning of the rods were observed from the scanning electron micrographs for all the three scaffold groups as the degradation time increased. In general, no significant nanopores can be detected yet on the surface of the native scaffolds at week 6 and 12. Surface line cracks with considerable gaps, and slight distortion has started to appear only after approximately 18 weeks. In contrast, for lipase-treated scaffolds which has undergone the highest degradation rate, disconnection of the rods can actually be seen as early as 6 weeks and by 24 weeks, severe degradation and reduction in the dimensions of the scaffolds were reported. In addition, compressive tests have also corresponded well with the degradation results. There was a drastic fall in the 79 strength and stiffness of the lipase-treated scaffolds during the first 6 weeks of immersion to an unfavourably low value (< 2MPa). For the native scaffolds, both their strength and stiffness have been low, with the values decreasing thereafter. As for the NaOH-treated scaffolds, a sudden decrease was noted between week 0 and week 6, after which the values of the strength and stiffness fluctuated around 2-3 MPa and 5-6 MPa. The data suggest that the NaOH-treated scaffolds were able to withstand considerable stress. Results here suggest that NaOH-treated scaffolds demonstrate better physical properties compared to the lipase-treated and native untreated ones. Lipase-treated scaffolds were shown to degrade in a much faster but uncontrollable manner, whereas the untreated native scaffolds displayed the lack of favorable surface properties. The next objective of the study was to investigate the in vivo degradation profile of untreated and treated PCL-TCP scaffolds implanted in the subcutaneous back of rats for 12 and 24 weeks (3 and 6 months respectively). During the period of implantation of up to 24 weeks, the PCL-TCP scaffolds exhibited major degradation and changes in the surface morphology. Scanning electron micrographs demonstrated obvious distortion and increased surface roughness of the scaffolds, especially for the lipase-treated scaffolds. Also, the distinct honeycomb-like architecture of the pores present in original PCL –TCP scaffolds was lost. Micro-CT scans revealed similar changes to the overall 3D framework and considerable reduction in size. The percentage of porosity reported an increasing trend for all three scaffold groups, with the highest rate seen for the lipase-treated scaffolds, followed by NaOH-treated scaffolds and native scaffolds respectively. By the end of week 24, the porosity has reached to considerably high values of 90.24%, 88.05%, 80 and 87.18% respectively. The weight loss recorded were 17.98%, 6.14%, and 4.12% for lipase-treated, NaOH-treated, and native scaffolds respectively. This increase helps to explain the porosity changes mentioned previously as weight-loss changes are related to porosity i.e the higher the porosity, the greater the weight loss [Wan, 2005]. In line with the increase in porosity and weight loss, the mechanical properties decreased significantly after 12 weeks of implantation. At the end of the in vivo study of 24 weeks, the strength of the lipase-treated scaffolds dropped to < 2MPa. As for the NaOH-treated scaffolds, the strength and stiffness were found to be above 2 MPa and 5 MPa respectively, indicating a more favourable tolerance in load bearing sites. Interestingly, the molecular weight of the three groups remained relatively constant during the degradation up to 12 weeks. Histological sections of the scaffolds after implantation in vivo for 3 and 6 months were stained with Hematoxylin and Eosin and analyzed. Throughout all scaffolds, the adipose tissues were found covering the surface of the rods as well as infiltrating the pores of the scaffolds. These cells were observed in direct apposition to the rods of the scaffolds and are generally presented with a healthy appearance, suggesting their biocompatibility nature. In addition, several vascular vessels were also detected; thus indicating that vascularization has occured. At the implantation sites, vasculature provides the main mode of transport. Molecular transport would include the exchange of oxygen, nutrient, metabolic wastes and molecular signaling. And these biochemical exchanges are essential for cell migration and proliferation. Moreover, the presence of overt inflammatory cells or fibrous encapsulation was not detected. One of the main characteristics of the pathological response (foreign body reaction) to a biomaterial is the presence of multinucleated giant cells. Macrophages 81 or multinucleated giant cells were not detected suggesting the absence of an inflammatory response by the scaffolds in vivo in a rat model. Based on this histology findings as well as the physical properties results for the PCL-TCP scaffolds during the in vivo study, NaOH-treated scaffolds appeared to provide a better performance compared to the other two groups (native and lipase-treated scaffolds) as they demonstrated a more favorable surface morphology and maintained sufficient mechanical properties while degraded predictably to a higher porosity value. 4.5.1 Comparison between in vitro and in vivo studies The studies in this chapter showed that in vitro and in vivo degradation of PCL-TCP scaffolds behaved quite similarly over the course of the degradation period. The extent of degradation of PCL-TCP scaffolds in vitro was found to be comparable to that in vivo. Data at a similar timepoint of 3 and 6 months were used in making the comparison to provide a fair analysis. When immersed in the DMEM growth medium, the scaffolds were found to degrade in a similar rate as when they were implanted subcutaneously into the back of rats. This was reflected through the relatively similar porosity increase, mechanical properties, and weight loss seen between in vitro and in vivo experiments. Results from SEM also showed that both groups similarly demonstrated significant changes to the surface morphology of the PCL-TCP scaffolds. The in vivo environment is much more complex and difficult to predict than the in vitro setting, particularly compared with acellular in vitro experiments [Hedberg, 82 2005]. Implant size and location, health of the animal, and enzymatic and local cellular activity are all factors that can influence the rate of degradation of a given polymer scaffold [Perrin, 1997]. Accordingly, in general, correlations between in vitro and in vivo results can be difficult to make. Nevertheless, in this study the in vivo degradation of the PCL-TCP composite scaffolds appears similar to that seen in the in vitro setting in the absence of any cells. Hence, the findings may suggest the possibility of conducting in vitro experiment by using DMEM solution to observe the degradation behaviour of PCL-TCP scaffolds in lieu of in vivo experiment in the future. In a study conducted previously, PCL-TCP scaffolds were implanted in the abdomen of rat model and results showed that the degradation rate was higher in vivo than in vitro [Yeo, 2007]. However, this is most likely due to the fact that rat’s abdomen was an aggressive environment and hence caused the scaffolds to degrade within a shorter time frame compared when they were immersed in the culture medium or when implanted subcutaneously to the back of rats. 4.6 CONCLUSION Some exciting observations were realized from this chapter of the study. The outcome demonstrated that the untreated and treated PCL-TCP scaffolds, which were separately immersed in culture medium for up to 24 weeks or implanted for 24 weeks inside rat model, had undergone a significant degradation. From the 12 and 24 weeks data, the extent of degradation in vivo was comparable to that observed in 83 vitro. However, regardless of in vitro or in vivo experiments, the degradation rate of the lipase-treated scaffolds was noted to be the highest, followed by NaOH-treated scaffolds and native scaffolds. Results from histology also supported that all the three group of scaffolds promote healthy cellular attachment as well as vascularizations. The absence of overt inflammatory response by the scaffolds in vivo in the rat model was also reported. In conclusion, our findings demonstrated that the NaOH-treated scaffolds performed most favourably as compared to the rest. In contrast, lipase-treated scaffolds degraded in a much faster and uncontrollable manner, and native scaffolds displayed the lack of favourable surface properties. For oral and maxillofacial applications, a scaffold that degrades around 6 months with controlled degradation rate and favorable mechanical properties is ideal for bone regeneration. 84 CHAPTER 5: EVALUATION OF PCL-TCP SCAFFOLDS IN A CLINICALLY RELEVANT DEFECT MODEL 5.1 INTRODUCTION In the previous chapter, the degradation behavior of PCL-TCP scaffolds in smaller animal model has been evaluated. We discovered that the native and customized scaffolds were biocompatible and that scaffolds with faster degradation rate were needed when used for oral and maxillofacial applications. However the rat study alone was still considered insufficient if we planned to present the technique for clinical applications; that is to treat dentoalveolar defects in human. The main concern was that in the previous animal study, the implantation of the scaffolds (subcutaneous back of rats) was not exactly at clinically relevant sites. Ideally prior to human clinical study, a more demanding and clinically representative larger animal model was required. Hence with this in mind, another in vivo study was conducted whereby PCL-TCP scaffolds were implanted in the mandibles of micropigs. The possibility of the PCL-TCP scaffold for use as a bone substitute in bone regeneration would then be compared to the current gold standard of using autogenous bone [Betz, 2002; Horch, 2006; Schuckert, 2009]. In addition to scaffolds, the latest guided bone regeneration (GBR) technique for localized ridge augmentation also involves the usage of barrier membranes. Extensive experimental studies have been conducted and results from these studies 85 have demonstrated that the placement of a membrane promotes the osseous healing of bone defects, since competing non-osteogenic soft tissue cells are excluded from defect healing by the presence of the physical barrier. Simultaneously, membranes allow the ingrowth of angiogenic and osteogenic cells to populate and regenerate these defects with bone [Schenk, 1994; Buser, 1996; Dahlin, 1988; Dahlin, 1990; Seibert, 1990]. On the other hand, noncovered control sites demonstrated incomplete bone regeneration and the presence of scar tissue within the defects [Buser, 2002]. These observations were confirmed in a recently published study by Schenk et al. that gave detailed information about the sequence and pattern of bone regeneration underneath barrier membranes. This study in the canine mandible demonstrated that bone regeneration in membrane-protected defects closely followed the pattern of normal bone growth and development, and that tissue which had formed beneath membranes was normal bone [Schenk, 1994; von Arx, 2001]. Recent experimental results also indicate that the use of barrier membranes can direct bone fill not only to contour defects of the alveolar ridge, but also to grow beyond the level of the surrounding bone, thus forming excess bone to a considerable extent [Schliephake, 1994; Kostopoulos, 1994; Schliephake, 1998]. For this study, we have decided to test for the first time, the feasibility of PCL-TCP sheets when used as barrier membranes. They were designed to be 3 layers thick with varying porosity, and would be compared to the widely used resorbable collagen membranes, which are currently the gold standard. The specific aim of this study was then to evaluate PCL-TCP scaffolds and sheets as defect fillers and barrier membranes respectively for novel guided bone regeneration technique in the reconstruction of localized dentoalveolar defects in a 86 micropig model. A total of 10 micropigs were employed for the study. All premolars (P1-P4) and the 1st molars (M1) were initially extracted from the posterior mandible and removed bilaterally to create partially edentulous alveolar ridges. Simultaneously, 2 bilateral bone defects were created on each side (15 x 10 x 8 mm) of the mandible by removing the buccal cortex. The defects were left to heal for a period of 2 months. Upon re-entry, the defects were planned for lateral ridge augmentation using either autogenous graft or PCL-TCP scaffolds. A collagen membrane or a prototype PCL-TCP sheet was also used. The membranes or sheets were trimmed, immersed in a hot water bath to mold it into shape and placed over the augmented sites and the defect margins by about 2mm. Each micropig then had the 4 defects randomly treated with one of the following GBR techniques: Site 1: Collagen membrane + PCL-TCP scaffold Site 2: Collagen membrane + autograft Site 3: PCL-TCP sheet + autograft Site 4: PCL-TCP sheet + PCL-TCP scaffold After a healing period of 6 months following GBR (Figure 5.1), all micropigs were sacrificed and analyses of the bone regenerated were performed. Surgery 1 (Extraction and defect creation) 2 months Surgery 2 (Ridge augmentation) 6 months Sacrifice + Analysis Figure 5.1: Timeline for the complete micropig study. 87 5.2 MATERIALS AND METHODS 5.2.1 Implant design and fabrication Bioresorbable scaffold and sheet specimens were fabricated with PCL-TCP (80:20%) filaments by using a fused deposition modeling (FDM) 3D Modeler RP system from Stratasys Inc (Eden Prairie, MN). They were purchased directly from Osteopore International Pte Ltd, Singapore. Each composite manifested a lay-down pattern of 0/60/120º with a typical honeycomb array of interconnected equilateral triangle. TCP existed as non-uniformly distributed particles on the rods of PCL. The scaffolds came in a block of 15 x 10 x 8 mm with 70% porosity for the outer 2 mm and 85% porosity for the inner 3 mm. On the other hand, the sheets were 25 x 25 x 1 mm consisting of 3 layers: 1 outer layer of 30% porosity and 2 inner layers of 70% porosity. The sheets were light, resilient and malleable. All specimens (Figure 5.2) were produced in a class 10K clean room environment and sterility was achieved via ethidium oxide treatment. Figure 5.2: 15x10x8mm PCL-TCP scaffold (left) and 25x25x1mm PCL-TCP sheet (right). 88 The bioresorbable collagen membranes were purchased from BioGide, Geistlich Pharma AG, Wolhusen, Switzerland (Figure 5.3, left). Bio-Gide’s collagen membrane was developed particularly for periodontal, peri-implant applications or to improve the ossification of bone defects of any origin. It is a bilayer membrane; one compact and smooth layer is covered by a particularly dense film, designed to prevent the invasion of soft tissue in a membrane-protected bone defect. The other rough side is placed towards the bone defect in order to make bone ingrowth possible [Zhao, 2000]. Figure 5.3: Bioresorbable collagen membrane from BioGide (left) and temperaturecontrolled hot water bath (right). 5.2.2 Animal husbandry A total of 10 male micropigs were employed for the study. At the beginning, these animals were about 1-2 years old and weighed approximately 40-50 kg. The animals were housed in the animal holding facility at the SingHealth Experimental Medicine Centre (SEMC), Singapore General Hospital (SGH), for the entire duration of the experiment. Housing and feeding were according to standard animal-care protocols. The study was conducted according to the guidelines of the SingHealth Experimental Medicine Centre, Singapore General Hospital. 89 Figure 5.4: Micropig housing facility at SEMC, SGH (left) and weighing of micropig prior to the experiment (right). 5.2.3 Pre- and postoperative medication All surgical procedures were performed under general anesthesia in an operating room. For premedication, the following agents were used: Ketamine (11-15 mg/kg, Ketapex®, Apex Laboratories, Australia) and Atropine (0.4-0.5 mg/kg, Pharmacia Pte Ltd, Australia) IM. Subsequently, the micropigs were intubated and were administered an inhalation of 5% Isoflurane initially and thereafter maintained with 13% Isoflurane (Abbott Laboratories Ltd, England) in 02-maintenance. After disinfection of the surgical site with 10% povidone-iodine solution (Clinidine®, Clinipad Co., Guilford, CT, USA), local anesthetic (Lidocaine HCL 2% with epinephrine 1:100,000, Henry Schein Inc., Port Washington, NY, USA) was administered by infiltration at the respective buccal and lingual sites. Postoperatively, 90 the micropigs received Caprofen 2-4 mg/kg every 12-24 hours for 3 days (Vericore Ltd, Scotland, UK) as an analgesic IM. For antibiotic cover, 2.5 mg/ 50kg of Benzathine-Penicillin 150,000 + Procaine-Penicillin G 150,000 was delivered every 48 hours for 7-10 days IM (Pen-B®, Pfizer Inc., Lee’s summit, MO, USA). In addition, 1-2 mg/kg of the antibiotic Gentamicin was administered IM every 12 hours for 7-10 days (Il Dong, Pharmaceutical Co Ltd, Korea). For suture removal, Ketamine (11-15 mg/kg Ketapex®, Apex Laboratories, Australia) and Atropine (0.4-0.5 mg/kg, Pharmacia Pte Ltd, Australia) were administered IM and thereafter maintained with 1-3% Isoflurane (Abbott Laboratories Ltd, England) in 02-maintenance. Oral hygiene procedures were carried out two times a week using 0.2% chlorhexidine gel (PlakOut® Gel, Hawe Neos Dental, Biaggio, Switzerland). A soft diet was maintained throughout the study. 5.2.4 Surgery 1 (Extraction and defect creation) Sulcular incisions were made with subsequent reflection of full mucoperiosteal flaps. In the mandible, all premolars (P1-P4) and the first molars (M1) were removed, whereas in the maxilla the 2nd and 3rd premolars (P2 and P3) were also extracted. Prior to removal, all two-rooted teeth were sectioned and separated individually, employing a separating disk, prior to root extraction. Subsequently, a large “chronictype” bone defect (Length 45 mm, Height 12 mm, Depth 5 mm) were created in the mandible by removing the buccal bone plate. The large defect encompassed approximately the extraction sites of P2, P3, and P4. A small round bur was used to outline the defect margins on the buccal bone plate. Subsequently, the bur holes were connected employing a fissure bur and the buccal bone plate was removed 91 with a chisel placed in the cut groove. In order to accentuate the defect, a pearshaped bur was utilized. Caution was exercised to retain the lingual cortex and the height of the crest. All drilling was done with sterile saline irrigation. Finally, the flaps were re-approximated with single interrupted resorbable 4.0 Vicryl sutures (Ethicon, Norderstedt, Germany). These were removed two weeks postoperatively. Oral hygiene procedures were carried out two times a week using 0.2% chlorhexidine gel (Plak-Out® Gel, Hawe Neos Dental, Biaggio, Switzerland). A soft diet was maintained throughout the study. During the first week of post-operative healing, the animals were checked daily for signs of infection. Figure 5.5: Removal of all premolars and first molar (left), and the extraction sites (right). Figure 5.6: The flaps were re-approximated with Vicryl sutures (left), and the defect sites were closed (right). 92 5.2.5 Surgery 2 (Ridge augmentation) After a healing period of 2 months, the defect sites in the mandible were reopened using a mid-crestal incision from P1 to M1. Vertical releasing incisions enabled full access to the area. All granulation tissue was carefully removed from the formerly created ridge defects. To open up the bone marrow space around the chronic-type bone defect, small holes were drilled into the surrounding cancellous compartment. The bone defects were then augmented in four different ways with random assignment of each grafting treatment. Site 1: Collagen membrane + PCL-TCP scaffold Site 2: Collagen membrane + autograft Site 3: PCL-TCP sheet + autograft Site 4: PCL-TCP sheet + PCL-TCP scaffold Site 1 Site 2 PCL-TCP scaffold + collagen membrane Autograft + collagen membrane Site 3 Site 4 Autograft + PCL-TCP sheet PCL-TCP scaffold + PCL-TCP sheet Figure 5.7: Schematic illustrations of the four tested grafting procedures. 93 The autografts were harvested from the site of the formerly extracted M1 (12 x 8 x 5 mm), using the same method as before when the mandibular defects were created. Cortico-cancellous block grafts were procured from the buccal aspect using a small round bur to outline the grafts with a series of perforations. These were then connected with a side-cutting fissure bur and the fragments relieved with a chisel placed in the cut groove. Before the placement of the grafts and scaffolds, multiple small perforations were made into the recipient wall of the defect to encourage bleeding and the release of growth factors and cells into the defect sites. Figure 5.8: Placement of PCL-TCP scaffolds and autografts (left), followed by PCL-TCP sheets and collagen membranes (right). The block grafts were then immediately transplanted to their assigned defect sites. A porous PCL-TCP scaffold (12 x 8 x 5 mm) was used as an alternative bone substitute and placed in the assigned defects. The grafts and scaffolds were secured using a centrally located miniscrew. A prototype PCL-TCP sheet and a bioresorbable collagen membrane (BioGide®, Geistlich Pharma AG, Wolhusen, Switzerland) were again randomly selected and individually trimmed to overlap the defect margins by about 2-3 mm. To facilitate a fluid-tight and tension-free wound closure, the periosteum was released at its base. Wound margins were then re-approximated 94 and closed with horizontal mattress and interrupted resorbable 4.0 Vicryl sutures (Ethicon, Norderstedt, Germany). Sutures were removed two weeks postoperatively. Oral hygiene procedures were carried out two times a week using 0.2% chlorhexidine gel (Plak-Out® Gel, Hawe Neos Dental, Biaggio, Switzerland). A soft diet was maintained throughout the study. During the first week of post-operative healing, the animals were checked daily for signs of infection. 5.2.6 Sacrifice Figure 5.9: Micropig under euthanasia (left), and the mandible was block resected using an oscillating autopsy saw (right). Figure 5.10: The recovered segment of mandible (left), the site after removal (right). 95 All 10 micropigs were sacrificed 6 months after lateral ridge augmentation. Euthanasia was performed with an overdose of pentobarbital sodium 0.2 ml i.v. (=65 mg/kg, Euthanasia-5®, Henry Schein Inc). Subsequently, the mandibles were block resected using an oscillating autopsy saw and the recovered segments were immediately immersed in a solution of formaldehyde 4% combined with CaCl2 1%. 5.2.7 Micro-computed tomography analysis Upon retrieval from the animals, the PCL-TCP scaffold specimens were characterized using Micro-computed tomography (Micro-CT Skyscan 1076, Belgium). The specimens were placed in a sample holder and scanned through 180° at a spatial resolution of 30µm. The image data from the scanned planes were subsequently thresholded and reconstructed to create 3-D images for quantitative analysis. The 3D volumes were evaluated by direct transformation methods and subsequently the total scaffold volume was calculated within a volume defined by the boundaries of the constructs. All parameters were measured on the buccal and lingual aspects of the specimens. The bone volume fraction (BVF) obtained from BV/TV, represents the percentage of a volume of interest that was mineralized. 5.3 RESULTS 5.3.1 Gross examinations All 10 animals used in the study remained healthy throughout the experiment and survived the surgical procedure with minimal adverse effects from the implantation of 96 the autografts and the PCL-TCP constructs. Opened sutures, or soft tissue dehiscences, were noted for the majority of grafts covered with PCL-TCP sheets (70% occurrence for autograft + PCL-TCP sheet combination, and 90% occurrence for PCL-TCP scaffold + PCL-TCP sheet combination). However, only some were observed for those covered with collagen membranes (10% occurrence for autograft + collagen membrane combination, and 20% occurrence for PCL-TCP scaffold + collagen membrane combination). Figure 5.11: The recovered segment of the mandible of a micropig. Figure 5.12: Soft tissue dehiscence observed for the majority of grafts covered with PCL-TCP sheets. 97 Table 5.1: Number of sites with soft tissue dehiscence for the implanted autograft, collagen membranes, PCL-TCP scaffolds, and PCL-TCP sheets. Group Combinations Number of sites with soft tissue dehiscence (out of 10) 1 Autograft + Collagen membrane 1 2 Autograft + PCL-TCP sheet 7 3 PCL-TCP scaffold + Collagen membrane 2 4 PCL-TCP scaffold + PCL-TCP sheet 9 Due to the high occurrence of the soft tissue dehiscence for PCL-TCP sheets groups, it was decided that samples from this group would be omitted for analysis. Only those covered with collagen membranes that would be discussed from this point onwards (i.e. autograft means autograft covered with collagen membrane). 5.3.2 New bone formation Micro-CT was utilized to determine the value of early bone volume ingrowth detected at the 6 months after implantation of the autografts and the scaffolds. Results of new bone formation at 6 months showed a higher trend in the volume of bone ingrowth by the autografts than the scaffolds (as illustrated in Figure 5.13). Due to the missing screw in the autograft site covered with collagen membrane in P7, values from P7 were omitted from the computation for better comparison purpose. 98 Figure 5.13: Bone volume fraction detected after 6 months of implantation of autografts and PCL-TCP scaffolds for individual micropigs. The bone volume fraction represents the amount of mineral present at the defect site. The volume fraction of newly formed bone within the defect was calculated to be 22.67% for PCL-TCP scaffolds under collagen membrane, and therefore the pores of the PCL-TCP scaffolds were not completely filled with mineralized tissue after 6 months. Bone formation in the autograft site covered with collagen membrane was higher at 36.22% volume fraction. These values (shown in Figure 5.14) were based on the average of 9 micropigs, as data from P7 was incomplete and hence omitted. The considerably high standard deviation was likely due to the different healing pattern and capacity of each micropig tested. 99 Figure 5.14: The average values of bone volume fraction detected after 6 months of implantation of autografts and PCL-TCP scaffolds. 5.3.3 Ratio of bone volume fraction for PCL-TCP scaffolds with respect to autografts In order to quantify the efficiency of PCL-TCP scaffolds as compared to autograft in the bone reconstruction application, the ratio of bone volume fraction for PCL-TCP scaffolds with respect to autografts were computed. This was obtained by dividing the BV/TV value of a PCL-TCP scaffold with the BV/TV value of an autograft. Figure 5.14 displayed the calculated ratio for each individual micropigs tested. As shown from the graph, the efficiency for all the micropigs was found to be comparable. The average, based on 9 micropigs, was 0.64 ± 0.16. This means that when compared to the use of autografts in the reconstruction of bone, the PCL-TCP scaffolds were about 64% efficient. 100 Figure 5.15: The ratio of bone volume fraction for PCL-TCP scaffolds with respect to autografts for individual micropigs. 5.3.4 3D model analysis Representative 3D micro-CT images of autograft-treated and PCL-TCP scaffoldtreated defects under collagen membrane were constructed and examined to observe the formation of new bone at the defect area. It must be noted that the images were captured at a density that only picked up bone and miniscrews, and not scaffolds. The entire height of the defect was not fully occupied with bone. Regularsized gaps were evident in the new bone that corresponded to the scaffold porous architecture. 101 Figure 5.16: PCL-TCP scaffold treated site: overview (left) and cross-section (right). Figure 5.17: Autograft-treated site: overview (left) and cross-section (right). 5.3.5 Two-dimensional x-ray radiographs evaluation Radiographs of the defect sites after 6 months implantation of both autograft and PCL-TCP scaffold were also taken and findings similar to those from the Micro-CT images were revealed. The volume of mineralized tissue at the defect sites treated with autografts were observed to be higher than those treated with PCL-TCP scaffolds as shown from Figure 5.18 to 5.20 below. The images below show the side 102 of the mandible that has been treated with both autograft and PCL-TCP scaffold, and covered with collagen membrane. “P” and “A” refers to the posterior and anterior section of the mandible respectively. Figure 5.18: X-ray image of a micropig’s left mandible treated with autograft (posterior) and PCL-TCP scaffold (anterior), and covered with collagen membrane. Figure 5.19: X-ray image of a micropig’s right mandible treated with PCL-TCP scaffold (posterior) and autograft (anterior), and covered with collagen membrane. 103 Figure 5.20: X-ray image of a micropig’s left mandible treated with autograft (posterior) and PCL-TCP scaffold (anterior), and covered with collagen membrane. 5.4 DISCUSSION The present study has evaluated the use of PCL-TCP scaffolds and sheets for ridge augmentation in an experimental micropig model. A pig model has been chosen, as it possesses similar healing properties to that of the human. The anatomy of the pig’s jaws and their dentition, bone metabolism, clotting parameters closely resembles to that of a human. Being omnivorous animals, the dietary pattern would also be similar. Initially the domestic pig model was chosen due to their availability and low cost. However, potential logistic and surgical challenges at sacrifice when the final weights of these pigs measure approximately 130-150 kg must be taken into consideration. A smaller micropig model was then finally decided. A clinically frequent situation with mandibular bone atrophy was simulated in this study by creating chronic bone defects on the buccal aspect of the alveolar ridge. The defects were treated using either autogenous grafts or PCL-TCP scaffolds in 104 combination with collagen membranes or prototype PCL-TCP sheets. Attention was then given to achieving a close adaptation of the membrane to the surrounding bone and good membrane stabilization. Throughout the experiment, all the autograft bone blocks and PCL-TCP scaffolds that were implanted within the assigned defects remained in situ with no signs of migration. The survival rate of the micropigs was excellent and they healed and recovered well after the surgeries. There were no reports of any complications detected, such as overt edema and infection during the healing period, suggesting the biocompatibility of PCL-TCP scaffolds in promoting a conducive in vivo environment for healing to take place. However, mid-way through the 6 months healing period, almost all the 10 micropigs presented with soft tissue dehiscence and biomaterial exposure in majority of sites that utilized PCL-TCP sheets as barrier membranes. This occurred despite the stringent post-operative care that was prescribed to the animals following GBR surgeries. It appears that the nature of the PCL-TCP sheets when moulded around the augmented sites had some form of ‘elastic memory’ which resulted in a partial relapsed of the PCL-TCP sheets and caused unwanted tension and perforation of the overlying soft tissue whilst healing. When a soft tissue dehiscence occurs, the exposure of the membrane could lead to its contamination with bacteria from the oral cavity and frequently to an infection in the membrane site [Buser, 1999; Buser, 2002]. For this reason, samples from the PCL-TCP sheets groups were omitted. Interestingly, GBR procedures utilizing collagen membranes with autografts or PCL blocks showed few soft tissue complications. The reason could be because collagen membrane is pliable when moist and conforms well to the surgical area. It provides a thrombogenic surface that is sealed coronally to the root surface by a fibrin clot, and does not elicit any allergic responses [Blumenthal, 1993]. Collagen itself has several advantages because it is 105 absorbable, does not require a second surgical procedure for removal, and has some unique biologic properties. It is the major extracellular macromolecule of the periodontal connective tissue and bone and is physiologically metabolized by these tissues; it is chemotactic for fibroblasts; it has been reported to act as a barrier for migrating epithelial cells in vitro; and it has been used experimentally in animals and humans [Pitaru, 1988; Owens, 2001]. The results from this study demonstrated that the combined placement of collagen membranes with autografts at the defect sites promotes the best osseous healing of bone defects as shown from the higher volume of mineralized tissue detected. This observation is consistent with the previous study conducted by Buser et al, comparing autografts with four alternative bone fillers in such defects in the mandible of micropigs. Utilizing autografts, bone regeneration is optimized during the initial bone healing period due to the presence of particulate grafts with excellent osteoconductivity [Buser, 1999; Lyford, 2003]. In addition, autografts also have osteoinductive properties in the sense of osteogenic transfer. With graft application, osteoblasts and osteoblast precursor cells as well as growth factors (transforming growth factor β) and bone-inducing factors (bone morphogenetic protein) entrapped in the grafted bone matrix are transferred to the augmentation site resulting in an activation of bone formation [Burchardt, 1983; Buser 1999, Buser, 1996]. This is crucial as bone formation is mainly activated by the release of growth factors and bone-inducing substances. This activation is manifested as a stimulation of neoangiogenesis, the recruitment of osteoblasts, and in the onset of bone matrix deposition, provided that the activators act upon committed responding cells. An influencing mechanism of collagen membrane placement may also be that 106 stimulating growth factors are locally concentrated in the osseous wound at inductive doses, leading to osseous repair of such defects that normally would not heal spontaneously [Schenk, 1994; Dahlin, 1993]. It is also essential to understand the biological behavior of autografts with respect to graft incorporation and repair and the differences between cortical and cancellous autografts. These details have been intensively studied in numerous experimental studies in orthopedic surgery [Burchardt, 1983; Burchardt, 1987]. Cancellous autografts are rapidly revascularized, and they are completely repaired by creeping substitution. In contrast, revascularization of cortical autografts is slow and occurs through existing haversian canals. Remodeling of cortical autografts is also slow and results in a mixture of necrotic and new viable bone [Lyford, 2003; Buser, 1999]. Based on this biological knowledge of graft incorporation and graft repair, corticocancellous block grafts placed in the center of the augmentation area were subsequently used in this study. They were appropriately applied to the recipient site with rigid fixation of the graft. A bone-graft fixation screw should be used because it allows precise positioning of the graft and prevents micromovements of the graft underneath the membrane during healing. In addition, the block graft must be placed with its cortical layer facing buccally and the cancellous portion of the graft in direct contact of the host bone. This surgical approach is based on two assumptions. First, the cortical portion of the graft facing to the buccal aspect of the crest is used to reestablish the missing buccal cortex. Although this new cortex will be a mixture of necrotic and new viable bone, it offers good mechanical stability and is less susceptible to resorption than cancellous bone. Second, the cancellous portion of the graft is placed in direct contact to the host bone in the area where the implant will 107 be placed during next surgery. The host bone surface is perforated during the surgical procedure to activate bone formation and to open the marrow space, allowing fast ingrowth of blood vessels. It can be expected that this portion of the graft will undergo rapid revascularization and graft remodeling. These assumptions, however, are based on orthopedic literature, and histologic details of graft incorporation and repair underneath barrier membranes are not yet documented [Buser, 2002; Buser, 1999]. Both autografts and PCL-TCP scaffolds are found to be able to support an applied membrane, thus preventing a membrane collapse and maintaining the created space. There is an agreement in the literature that the maintenance of the membrane-protected space is one of the essential prerequisites for a successful treatment outcome with guided bone regeneration procedures [Buser, 1990; Rominger, 1994; Buser, 1999]. When compared to the use of autografts in the reconstruction of bone, the PCL-TCP scaffolds were only about 64% efficient. However the application of PCL-TCP scaffolds, instead of autografts, would be highly preferable as this option would avoid the harvesting of autogenous bone from the patient. Hence, more studies and modifications were needed to further enhance the efficiency of PCL-TCP scaffolds as a bone substitute in bone regeneration for localized ridge augmentation. The development of a longer-lasting bioabsorbable membrane with the same qualities as the current collagen membrane concerning tissue compatibility, mechanical attributes, and intrasurgical handling would be highly desirable too. 108 5.5 CONCLUSION This part of the study was to investigate the in vivo behavior of early matrix deposition and early bone formation of PCL-TCP scaffolds as compared to the use of autografts, in conjunction with membrane coverage in the mandible of micropigs after 6 months of implantation. A clinically frequent situation with mandibular bone atrophy was simulated by creating chronic bone defects on the buccal aspect of the alveolar ridge. Healing was uneventful in all micropigs showed that the PCL-TCP scaffolds exhibited good biocompatibility. Across the tested treatment options in this study of 10 animals, defect sites augmented with autografts and collagen membranes showed the most promising results. The collagen membranes were found to offer the advantage of a reduced frequency of soft tissue dehiscence. More improvements are needed to increase the efficiency of the PCL-TCP scaffolds in bone healing as they could ruled out the need for harvesting grafts. 109 CHAPTER 6: FINAL CONCLUSIONS AND RECOMMENDATIONS 6.1 FINAL CONCLUSIONS The general aims of the present study were to investigate the degradation and loadbearing profile of 3D bioresorable polycaprolactone-20% tricalcium phosphate (PCLTCP) scaffolds under enzymatic and hydrolytic conditions and subsequently to evaluate the efficacy of the scaffolds in both small and large animal models. The purpose was to develop scaffolds with desirable customized properties and increased degradation rates suitable for application in dentoalveolar defects treatment. PCL-TCP scaffolds have shown to undergo both chemical and enzymatic degradation in the first part of the study when degraded in sodium hydroxide and lipase solution respectively for up to 108 hours. The objective of the experiment was achieved with scaffolds of approximately 85% porosity obtained after 96 hours of treatment in 3M NaOH and 12 hours in 0.1% lipase. These pre-treated scaffolds demonstrated acceptable mechanical strength, structure, and surface morphology. Some exciting observations were also realized from the second part of the study. The outcome demonstrated that the untreated and treated PCL-TCP scaffolds, which were separately immersed in culture medium for up to 24 weeks or implanted 110 for 24 weeks inside rat model, had undergone a significant degradation. From the 12 weeks data, the extent of degradation in vivo was comparable to that observed in vitro. However, regardless of in vitro or in vivo experiments, the degradation rate of the lipase-treated scaffolds was noted to be the highest, followed by NaOH-treated scaffolds and native scaffolds. Results from histology also supported that all the three group of scaffolds promote healthy cellular attachment as well as vascularizations. The absence of overt inflammatory response by the scaffolds in vivo in the rat model was also reported. Our findings demonstrated that the NaOHtreated scaffolds performed most favourably as compared to the rest. In contrast, lipase-treated scaffolds degraded in a much faster and uncontrollable manner, and native scaffolds displayed the lack of favourable surface properties. Lastly, another in vivo study was conducted whereby PCL-TCP scaffolds and sheets were evaluated as defect fillers and barrier membranes respectively for novel guided bone regeneration technique in the reconstruction of localized dentoalveolar defects in a micropig model for up to 6 months. This was essential as prior to human clinical study, a more demanding and clinically representative larger animal model was required. The possibility of the PCL-TCP scaffold for use as a bone substitute was compared to the current gold standard of using autogenous bone. Healing was found to be uneventful in all micropigs and this showed that the PCL-TCP scaffolds exhibited good biocompatibility. Across the tested treatment options, defect sites augmented with autografts and collagen membranes showed the most promising results with greater bone formation detected as compared to PCL-TCP scaffolds and collagen membranes which were about 64% efficient. The collagen membranes were found to offer the advantage of a reduced frequency of soft tissue dehiscence 111 in comparison to PCL-TCP sheets. For oral and maxillofacial applications, a scaffold that degrades around 6 months with controlled degradation rate and favorable mechanical properties is ideal for bone regeneration. More improvements are needed to increase the efficiency of the PCL-TCP scaffolds in bone healing as they could ruled out the need for harvesting grafts. 6.2 RECOMMENDATIONS FOR FUTURE WORK Future work may also entail the investigation of in vivo degradation behaviour of the bioactive factors-loaded PCL-TCP scaffolds in bone mandibular defects of large animal model such as micropig. The growth factor of choice could be bone morphogenetic proteins (BMPs), one of the most potent factors that regulate osteoblasts to form mineralized tissue [Yamaguchi, 2000]. The key research objective here is to investigate the capacity of BMP-loaded PCL-TCP composite scaffolds in improving the stimulation of bone repair and regeneration upon implantation into large segmental bone defects for a long-term period and ultimately to rule out the need for harvesting autografts. A large animal model is vital for advancement into clinical trials and a long time period is required to acquire useful insights of the in vivo degradation properties of the scaffolds and the time required for mature bone deposition and remodeling to occur. BMPs have the unique functions of inducing the differentiation of cells of the osteoblastic lineage, thus increasing the pool of mature cells, and of enhancing the differentiated function of the osteoblasts [Canalis, 2003; Yamaguchi, 2000]. They were originally identified as growth factors that regulate growth and differentiation of chondroblast and osteoblast 112 lineage cells in vitro. In accordance with their in vitro effects, BMPs induce bone and cartilage formation when implanted at ectopic sites in rats [Reddi, 1997; Urist, 1965]. PCL-TCP scaffolds seeded with recombinant human bone morphogenetic protein-2 (rhBMP-2) have previously been investigated both in vitro and in small animal models, and results have shown that the incorporation of rhBMP-2 seems to accelerate mineralization. When implanted for guided bone regeneration technique in the reconstruction of localized dentoalveolar defects in a micropig model, specific objectives could include investigating the amount and quality of bone formation in terms of the bone volume fraction, bone union formation and mechanical properties; the amount and type of cellular and vascular infiltration into the defect site; the longterm effects of the material on host tissue; and the long-term effects of the host environment on the material biodegradation. It is also interesting to compare the concentration of growth factor released at the local site of implantation against the amount that is transported in the bloodstream and excreted in urine. It is hoped that the local, controlled release of rhBMP-2 from PCL-TCP scaffolds in vivo will enhance the osteoinductivity of the scaffolds and that this strategy will provide a different and more sophisticated way of enhancing the efficacy of the PCL-TCP scaffolds for bone tissue engineering in dentoalveolar applications. In addition, more characterization methods and extensive analyses of the PCL-TCP scaffolds can be performed to provide better understanding of the changes in the scaffolds’ properties. 113 REFERENCES Adell R, Eriksson B, Lekholm U, Branemark P I, Jemt T. Long term follow up study of osseointegrated implants in the treatment of totally edentulous jaws. Int J Oral Maxillofac Implants 1990; 5: 347-359. Alsberg E, Kong H.J, Hirano Y, Smith M.K, Albeiruti A, Mooney D.J. 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Biomaterials 2007; 28: 814-824. 121 [...]... an initial stage of a larger project, in order to develop a scaffold of a higher porosity that allows for a more rapid degradation whilst maintaining favourable mechanical properties A final porosity of about 85% was targeted for 1.3.2 Part 2: Optimization of native and customized scaffolds in vitro and their effects in initial bone healing (in Chapter 4) In the second part of the study, PCL-TCP scaffolds. .. channels Forming the outer wall of bones, it bears most of the supportive and protective function of the skeleton Cancellous bone, on the other hand, makes up the remaining 20% of bone mass in the body It consists of trabeculae which form an interconnected lattice Cancellous bone can be found in vertebrae, fracture joints, ends of long bones and in foetuses The whole structure, an outer cortical sheath and. .. of 85% by treating them with 3M NaOH or 0.1% lipase-PBS medium under physiological conditions for up to 108 hours 2 To compare the degradation profile of treated and untreated PCL-TCP scaffolds in vitro when immersed in standard culture medium for up to 24 weeks, and in vivo when implanted in the subcutaneous back of rats for 24 weeks (6 months) 3 To evaluate the rate and extent of bone formation of. .. in bone healing as they could ruled out the need for harvesting grafts x LIST OF TABLES Table 3.1 Mw, Mn, and PDI of NaOH-treated and lipase-treated PCL-TCP Scaffolds 40 Table 4.1 Mw, Mn, and PDI of native, NaOH-treated, and lipase-treated PCL-TCP Scaffolds in vitro 66 Table 4.2 Mw, Mn, and PDI of native, NaOH-treated, and lipase-treated PCL-TCP Scaffolds in vivo 75 Table 5.1 Number of sites with soft... alive, and contains cells which work continuously to regenerate and repair it 9 [Bronner, 1999; Ferrer, 2007] Bone tissue contains five basic types of bone cells: osteogenic cells, osteoblasts, osteocytes, osteoclasts, and bone- lining cells Osteogenic cells respond to traumas, such as fractures, and begin the healing process immediately by giving rise to bone- forming cells (osteoblasts) and bonedestroying... polymers in the field of tissue engineering has been widely investigated in recent years, with advances in the scaffold technology extending their usage to clinical applications such as bone regeneration In particular of such interest is poly(ε-caprolactone) -tricalcium phosphate (PCL-TCP) composite scaffold, a synthetic biodegradable polymer frequently investigated for bone tissue engineering applications... engineering 2.2.1 Degradation of PCL polymer Degradation behaviours of scaffolds play an essential role in the engineering of new tissue, as the rate of degradation is intrinsically linked to many cellular processes including cell viability, tissue growth, as well as the host response [Lei, 2006] Once implanted in the body, a porous scaffold should maintain its mechanical properties and structural integrity... assembly of collagen fibrils and fibers and bone mineral crystals [Rho, 1998] Bone s function is both biomechanical and metabolic Biomechanically, bone acts to: (1) maintain the shape of the skeleton, (2) protect soft tissues in the cranial, thoracic and pelvic cavities, (3) transmit the forces of muscular contraction during movement, and (4) supply a framework for bone marrow Metabolically, bone (1)... regarding bone tissue engineering strategy and the application in implant dentistry, as well as the current drawback of PCL-TCP scaffolds in dentoalveolar defects treatment that lead the author to pursue this research Detailed research objectives and research scope are discussed in the next and last sections respectively 1.1.1 Bone tissue engineering Loss of human tissues or organs is a devastating problem... role of scaffolds come into the picture as they may eliminate the need for an extensive bone harvesting procedure from a donor site However in facing a complex biological system as the human body, the requirements of scaffold materials for bone tissue engineering in dentoalveolar application can be extremely challenging 3 1.1.3 PCL-TCP scaffolds: Current drawback The use of synthetic polymers in the ... submitted for the degree of Master of Engineering in the Department of Mechanical Engineering at the National University of Singapore under the supervision of Professor Teoh Swee Hin and Dr Alvin Yeo... secrete bone tissue and form the tissue around itself like a protective wall of bone tissue They are responsible for the maintenance of healthy bone by secreting enzymes and directing the bone mineral... CHAPTER 4: OPTIMIZATION OF NATIVE AND CUSTOMIZED SCAFFOLDS IN VITRO AND THEIR EFFECTS IN INITIAL BONE HEALING 4.1 INTRODUCTION 44 4.1.1 In vitro degradation study 44 4.1.2 In vivo degradation study