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

49 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 4.13 3D model of native sca

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 caprolactone) tricalcium phosphate composite scaffolds 4th International Conference on Materials for Advanced Technologies (ICMAT), Singapore, 2007

poly(ε-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

 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

LIST OF SYMBOLS xviii

CHAPTER 1: INTRODUCTION

1.1 BACKGROUND 1

1.1.3 PCL-TCP scaffolds: Current drawbacks 4

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

CHAPTER 2: LITERATURE REVIEW

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

3.2.3.5 Molecular weight testing 33

3.2.4 Statistical analysis 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

4.2.5.6 Histology preparation and analysis 50

4.3.1 Porosity measurements and 3D model analysis 51

4.4.1 Porosity measurements and 3D model analysis 67

4.5.1 Comparison between in vitro and in vivo studies 82

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

5.2.4 Surgery 1 (Extraction and defect creation) 91

5.2.5 Surgery 2 (Ridge augmentation) 93

5.3.3 Ratio of bone volume fraction for PCL-TCP scaffolds

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

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

Table 5.1 Number of sites with soft tissue dehiscence for the implanted

autograft, collagen membranes, TCP scaffolds, and TCP sheets

PCL-98

LIST OF FIGURES

Figure 1.2 Schematic diagram of part 1 and part 2 study 7

Figure 2.2 The assembly of collagen fibrils and fibers and bone mineral

Figure 2.6 The degradation rate of PGA, PLA, and PCL 16

Figure 2.9 Sequence of the data preparation for FDM model fabrication 26

Figure 3.2 Illustration of scaffold with 0/60/120º lay-down pattern 30

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

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

38

Figure 3.11 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

Figure 4.8 Incision made (left), implanted scaffold (left, inset), scaffold

positions (right)

49

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

52

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

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) treated scaffolds (L) overall view, and (R) higher-

lipase-magnification view

62

Figure 4.29 Electron micrographs taken after 12 weeks immersion in

DMEM for: (a,b) native, (c,d) NaOH-treated, and (e,f) treated scaffolds (L) overall view, and (R) higher-

lipase-magnification 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) treated scaffolds (L) overall view, and (R) higher-

lipase-magnification 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) treated scaffolds (L) overall view, and (R) higher-

lipase-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

66

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

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

Figure 5.2 15x10x8mm TCP scaffold (left) and 25x25x1mm

PCL-TCP 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)

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

cross-section (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

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

104

T m Melting point

LIST OF ABBREVIATIONS

DMEM Dulbecco’s modified Eagle’s medium

Micro-CT Micro-computed Tomography

PCL Poly(ε-caprolactone)

rhBMP-2 Recombinant human Bone Morphogenetic Protein-2

SEMC SingHealth Experimental Medicine Centre

CHAPTER 1: INTRODUCTION

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

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 bone-extracellular 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

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 dento-alveolar 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

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

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

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:

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, lipase-treated and untreated They were (a) implanted subcutaneously into the back of rats,

In vitro

Small animal model

Large animal model

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

0, 12, 24, 36, 48, 60, 72, 84, 96, 108 hours

0, 6, 12, 18, 24 weeks 0, 12, 24 weeks

→ Porosity → Weight Loss → Strength → Surface → Molecular → Inflammation → Structure → Stiffness 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.

Mechanical Testing

Micro-CT

Pre-degradation Study

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 three-dimensional 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

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

[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 bone-destroying 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

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]

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

sub-nanostructure [Rho, 1998]

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

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

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]

Fast Slow

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

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

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