In vitro and in vivo assessments of PCL TCP composites for bone tissue engineering

IN VITRO AND IN VIVO ASSESSMENTS OF PCL-TCP COMPOSITIES FOR BONE TISSUE ENGINEERING WONG WAH JIE (B. Eng. (Hons), NUS) A THESIS SUBMITTED FOR THE DEGREE OF MASTER OF ENGINEERING DEPARTMENT OF MECHANICAL ENGINEERING NATIONAL UNIVERSITY OF SINGAPORE 2010 INTERNATIONAL JOURNAL PUBLICATIONS A. Yeo, W. J. Wong, H. H. Khoo and S. H. Teoh “Surface modification of PCL-TCP scaffolds improve interfacial mechanical interlock and enhance early bone formation: An in vitro and in vivo characterization” Journal of Biomedical Materials Research: Part A v. 92A, p. 311-321 (2010) A. Yeo, W. J. Wong and S. H. Teoh “Surface modification of PCL-TCP scaffolds in rabbit calvaria defects: Evaluation of scaffold degradation profile, biomechanical properties and bone healing patterns” Journal of Biomedical Materials Research: Part A v. 93A, p. 1358-1367 (2010) i ACKNOWLEDGEMENTS The author would like to express his sincere gratitude and heartfelt thanks to the following individuals who have rendered assistance or gave valuable advice leading towards the successful accomplishment of this research project:  Professor Teoh Swee Hin (Dept of Mechanical Engineering), project supervisor, for giving valuable advice and support through the project. The author appreciates the trust and independence that has been given to him.  Dr. Alvin Yeo (National Dental Centre), project co-supervisor, for his constant supervision and guidance in this project. His active role in coordinating and participating guarantees the success of this study.  Dr. Bina Rai (IMB), project mentor, for giving valuable pointers on the project throughout the study despite her busy work schedule. The author would like to thank her for reviewing the drafts.  Dr. Simon Cool (IMB), project mentor, for his suggestions given on the progress of this project during presentations.  Richard Lin (3M), for his assistance in interferometery of composite thin films.  Dr Amber Sawyyer and Ivy See Hoo, for their assistance in histology and histomorphometry work. ii  Mr Low Chee Wah and Mr Abdul Malik Bin Baba (Impact Mechanics Lab), for rendering support and assistance on Instron micro tester for mechanical testing.  Ms Irene Kee (Dept of Experimental Surgery, SGH), for her valuable assistance and support during animal surgeries and taking care for them after implantations.  Everyone at VNSC and BIOMAT, for their encouragement and laughter throughout the whole period, which fills up the entire durations with wonderful memories.  60 rabbits (R2 to R24) that have been sacrificed in this project till now. They did not have a choice but it is them that made everything possible. And last but not least, to ALL those who has contributed in one way or another in this project. iii TABLE OF CONTENTS INTERNATIONAL JOURNAL PUBLICATIONS.................................................. i ACKNOWLEDGEMENTS ...................................................................................... ii TABLE OF CONTENTS ......................................................................................... iv SUMMARY............................................................................................................. vii LIST OF TABLES .................................................................................................... ix LIST OF FIGURES.................................................................................................... x CHAPTER 1: INTRODUCTION..............................................................................1 1.1 BACKGROUND .........................................................................................1 1.1.1 Current trends in BTE .........................................................................1 1.1.2 Limitations of current treatments for bone defects ..........................2 1.1.3 Strategies in BTE ..................................................................................3 1.2 RESEARCH OBJECTIVES..........................................................................6 1.3 RESEARCH SCOPE ....................................................................................7 CHAPTER 2: LITERATURE REVIEW ....................................................................8 2.1 BONE PHYSIOLOGY .................................................................................8 2.2 BIOMATERIALS.......................................................................................11 2.2.1 Polycaprolactone (PCL) ....................................................................11 2.2.2 Biodegradation ..................................................................................12 2.2.3 Tri-Calcium Phosphate (TCP) ..........................................................14 2.2.4 PCL-TCP scaffolds .............................................................................14 2.2.5 Bone Morphogenetic Proteins (BMP-2) ...........................................17 2.2.6 Heparin...............................................................................................20 CHAPTER 3: EFFECTS OF POROSITIES OF PCL-TCP SCAFFOLDS ON BONE REGENERATION, SCAFFOLD DEGRADATION AND MECHANICAL PROPERTIES. ..........................................................................................................23 3.1 INTRODUCTION .....................................................................................23 3.2 MATERIALS AND METHODS...............................................................26 3.2.1 Scaffold Fabrication ...........................................................................26 iv 3.2.2 Porosity Calculation ..........................................................................26 3.2.3 Experimental Design .........................................................................26 3.2.4 Animal husbandry and scaffold implantation ................................27 3.2.5 Micro-CT analysis ..............................................................................29 3.2.6 Mechanical strength testing ..............................................................30 3.2.7 Histological Analysis ........................................................................31 3.2.8 Histomorphometric Analysis ...........................................................31 3.2.9 Mineral Apposition Rate (MAR) ......................................................32 3.2.10 Statistical Analysis .............................................................................33 3.3 RESULTS ...................................................................................................33 3.3.1 Scaffolds characterisations ................................................................33 3.3.2 µ-CT analysis .....................................................................................34 3.3.3 Compressive strength .......................................................................37 3.3.4 Push Out test ......................................................................................38 3.3.5 Histology ............................................................................................39 3.3.6 Histomorphometric Analysis ...........................................................40 3.3.7 Mineral Apposition Rate (MAR) ......................................................41 3.4 DISCUSSIONS ..........................................................................................42 3.5 CONCLUSIONS .......................................................................................49 CHAPTER 4: PRELIMINARY EVALUATION OF PCL-TCP SCAFFOLDS AS CO-DELIVERY SYSTEMS FOR HEPARIN AND BMP-2 IN VITRO..................51 4.1 INTRODUCTION .....................................................................................51 4.2 MATERIALS AND METHODS...............................................................53 4.2.1 Porcine osteoblasts culture ...............................................................53 4.2.2 Cell culture and BMP-2 treatment ...................................................53 4.2.3 Heparin and BMP-2 treatment .........................................................54 4.2.4 Protein determination .......................................................................54 4.2.5 Alkaline phosphatase activity ..........................................................55 4.2.6 Western blot .......................................................................................55 4.2.7 Alizarin red staining..........................................................................55 v 4.2.8 Release profile studies .......................................................................56 4.2.9 BMP-2 Release....................................................................................56 4.2.10 Statistical Analysis .............................................................................56 4.3 RESULTS ...................................................................................................57 4.3.1 Optimal BMP-2 concentration ..........................................................57 4.3.2 Western blot analysis ........................................................................58 4.3.3 Alizarin red staining..........................................................................59 4.3.4 Optimal heparin concentration ........................................................59 4.3.5 Protein release profile .......................................................................60 4.3.6 BMP-2 release profile ........................................................................61 4.3.7 ALP bioactivity of eluted BMP-2 ......................................................62 4.4 DISCUSSIONS ..........................................................................................62 4.5 CONCLUSIONS .......................................................................................67 CHAPTER 5: FINAL RECOMMENDATIONS ....................................................68 5.1 Effects of porosities of PCL-TCP scaffolds on in vivo bone regeneration. ....................................................................................................68 5.2 Preliminary in vitro evaluation of PCL-TCP scaffolds as co-delivery systems for heparin and BMP-2 .........................................................................68 BIBLIOGRAPHY .....................................................................................................70 APPENDIX (PUBLICATIONS) .............................................................................82 vi SUMMARY This project consists of two chapters and revolves around PCL—TCP composite scaffolds. Pore size and porosity has always been an important property which affects cell and tissue infiltration. This has direct influence on bone regeneration and scaffold degradation upon implantation. In this study, acellular scaffolds of varying pore size and porosity were left in rabbit calvaria defects and explanted at 2, 4, 8, 12 and 24 weeks. Porosities (Group A: 74.9 ± 1.7%, Group B: 86.7 ± 0.2%) and pore size (Group A: 500 ± 73µm, Group B: 723 ± 92µm) were determined through Micro CT. %BV/TV from Micro CT demonstrated an increase up to 8 week and stabilizes thereafter. Power law relationship governed scaffold degradation rate regardless of porosity. Group B scaffolds (6 weeks) reached 50% scaffold loss faster than Group A (10 weeks). Mechanical properties between both groups were comparable throughout the study. Lastly, histology and histomorphometry detected bone formation and active vascularisation in all defects. In summary, porosities and pore size of PCL-TCP scaffolds has negligible effects on bone regeneration and scaffold degradation. Chapter 4 focused on the effects of BMP-2 and heparin on pig osteoblasts. Porcine osteoblasts obtained through explant culture displayed highest ALP activity in the presence of 100ng/ml of BMP-2 and 300ng/ml of vii heparin. Alizarin red staining and western blot confirmed the bioactivity of BMP-2 used. Protein release from Group D scaffolds has a biphasic release similar to Group B. This is in agreement with BMP-2 release profile. ALP activities of various groups were comparable mainly due to low concentration of eluted BMP-2. In conclusion, heparin and BMP-2 has demonstrated its potential as co delivery system in this preliminary study. More studies should be carried out to confirm the hypothesis. viii LIST OF TABLES Table 3.1: Physical parameters of PCL-TCP scaffolds ix LIST OF FIGURES Figure 2.1: Diagram showing skeletal long bone structure which comprises of cortical and trabecular bone (Biomedical Tissue Research Group, 1996) Figure 2.2 Schematic diagram showing stages of bone remodelling process (Biomedical Tissue Research Group, 2007) Figure 2.3: Chemical Structure of PCL polymer (Wikipedia, 2006) Figure 2.4 Chemical structure of tricalcium phosphate (CambridgeSoft Corporation, 2004) Figure 2.5: Activation of SMAD proteins after BMP mediation (Sakou, 1998) Figure 2.6: Chemical structure of Heparin (Gray et al., 2008) Figure 3.1: Micro CT images of scaffolds before implantation (A) Group A (B) Group B Figure 3.2: Implantation of scaffolds into rabbit calvarial defects Figure 3.3: Calculation of interlabel distances using Bioquant Image Analysis® software Figure 3.4: Representative µ-CT images of explants specimens with PCLTCP scaffold of (A) 75% porosity (B) 85% porosity (blue: scaffold, yellow: new bone growth and beige: calvaria bone) Figure 3.5: Percentage BV/TV in PCL-TCP scaffolds by µ-CT with varying porosities from 2 to 24 weeks of implantation. Figure 3.6: Scaffold volumes with varying porosities over a time period of 24 weeks. Figure 3.7: Percentage of PCL-TCP scaffolds volume loss with varying porosities from 2 to 24 weeks of implantation Figure 3.8: Compressive strength of PCL-TCP scaffolds with varying porosities from 2 to 24 weeks of implantation (* denotes p < 0.05) Figure 3.9: Shear strength of PCL-TCP scaffolds with varying porosities from 2 to 24 weeks of implantation (* denotes p < 0.05) x Figure 3.10: Representative images of ex vivo specimens at (A) 4 weeks (B) 8 weeks stained for Goldner’s Trichome (green sections: mineralised bone; areas labelled ‘S’ denote PCL-TCP scaffolds) Figure 3.11: Percentage BV/TV of PCL-TCP scaffolds by histomorphometry with varying porosities from 8 to 24 weeks of implantation Figure 3.12: Representative image showing diffused flurochrome labels between 2 and 8 weeks (red label – alizarin, green label – calcein) Figure 3.13: Mineral Apposition Rate (MAR) of PCL-TCP scaffolds with varying porosities from 8 to 24 weeks of implantation Figure 4.1: Alkaline phosphatase activity per mg protein of cells at different BMP-2 concentrations (* denotes p < 0.05) Figure 4.2: Western blot for pig osteoblasts treated with 100ng/ml BMP-2 at different treatment times. (top - phosphorylated SMAD 1/5/8 and bottom – Total SMAD 1/5/8) Figure 4.3: Alizarin red staining for pig osteoblasts with and without BMP-2 (control) treatment for 3 weeks. Figure 4.4: Alkaline phosphatase activity per unit protein of cells at different BMP-2 and/or heparin concentrations (* denotes p < 0.05) Figure 4.5: Amount of total protein release at various time points (Group A: PCL-TCP scaffolds loaded with PBS. Group B: PCL-TCP scaffolds loaded with 100ng/ml BMP-2. Group C: PCL-TCP scaffolds loaded with 300ng/ml of heparin. Group D: PCL-TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP2) Figure 4.6: Amount of BMP-2 release at various time points. (Group B: PCLTCP scaffolds loaded with 100ng/ml BMP-2. Group D: PCL-TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP2) Group A and C showed no release of BMP-2 at all time points. Figure 4.7: Bioactivity of eluted BMP-2 at different time points (Group A: PCL-TCP scaffolds loaded with PBS. Group B: PCL-TCP scaffolds loaded with 100ng/ml BMP-2. Group C: PCL-TCP scaffolds loaded with 300ng/ml of heparin. Group D: PCL-TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP2) xi CHAPTER 1: INTRODUCTION 1.1 BACKGROUND This chapter aims to give the reader an overview of the current trends in bone tissue engineering (BTE). Following that, limitations of available treatments for bone defects and various strategies of BTE will be discussed. 1.1.1 Current trends in BTE Healthcare spending in US can be represented by National Health Expenditure (NHE). It is defined as the “total amount spent to purchase healthcare goods and services as well as investment in the medical sector to produce healthcare services”(NHED, 2006). NHE has been rising rapidly throughout the years from $153 billion in 1976 to $1990 billion in 2006 (NHED, 2006). Healthcare comprises of 16% of GDP in 2006 and is expected to increase to 19% in a decade. Average annual growth of healthcare expenses involving musculoskeletal conditions is ranked second at 8.5% (HCUP, 2006). The huge growth can be attributed to the rapidly aging population. Also, more highimpact accidents have led to an increase in serious limb trauma. Bone and Joint decade (2000-2010) was set up by United Nations and World Health Organisation to raise awareness of the growing costs and also to deepen understanding of musculoskeletal diseases through research (BJD, 2000-2010). 1 In Singapore, Government Health Expenditure/Total Government Expenditure increased from 6.5 to 7.1% between 2006 to 2008 (MOH, 2008). 1.1.2 Limitations of current treatments for bone defects For surgical procedures involving bone grafts, patients usually suffer from trauma-related injuries or bone fractures. Currently, the gold standard for bone grafts is autologous bone commonly taken from iliac crest. Autologous bone are bones extracted from another part of patient’s own body (Casey K C, 2006). The drawbacks include an additional surgery site for harvesting and limited availability (Enneking et al., 1980). Other types of bone grafts include tissues taken from human donors (allografts) or xenografts obtained from animals. Allografts are not as popular due to the increased risk of disease transmission, high cost, graft rejection (Gitelis and Saiz, 2002), and the acceptance of xenografts in certain races or religion may be highly controversial due to its animal origin. Allografts which have undergone chemical treatment to remove minerals are known as demineralised bone matrix. Collagen and other growth factors are still being retained but they have low mechanical strength due to loss of minerals. Even though it is widely used as bone substitute, its effectiveness fluctuates in different patients (Drosos et al., 2007; Kay, 2007). Large clinical defects resulting from bone cancer or trauma are currently treated with titanium plates. This may lead to 2 complications like rejection and stress shielding. In serious cases, revision surgery has to be conducted to replace the implant. oLimitations to current bone grafts as discussed propel researchers to look into the field of bone tissue engineering for an ideal bone substitute. 1.1.3 Strategies in BTE Tissue engineering is the restoration, improvement, maintenance and substitution of damaged tissues and organs using principles of biology and engineering (Langer and Vacanti, 1993). BTE consists of an interplay of scaffold technology, growth factors and cells. In this thesis, the focus will be on using composite scaffold technology in enhancing bone regeneration through in vitro and in vivo studies. Scaffold intended for BTE should possess the following properties: 1. High porosity and pore interconnectivity to allow cell growth, migration and promote vascularisation (Sundelacruz and Kaplan, 2009). 2. Biocompatible, bioresorbable and controllable degradation rate to match surrounding tissue growth. 3. Suitable surface topography whereby cells are able to attach, proliferate and differentiate (Stevens et al., 2008). 3 4. Mechanical properties that closely resemble the defect site and has the strength to withstand load upon implantation (Hutmacher D.W, 2001). 5. The ability to impregnate cells, growth factors and drugs which can trigger surrounding cells for bone regeneration; controlled drug or growth factor delivery can also be effectively targeted at the defect site. 6. Ease of manufacturing is necessary for the scaffold to be mass produced. 7. Customizability of the shape and size of scaffold will be favourable for use in different clinical applications (Jones, 2005). It must be noted that interdependent relationships exist among the desired properties discussed above. One example is when the porosity of scaffold is increased, cells are able to infiltrate easily but mechanical strength will decrease. Scaffold degradation time will also shorten due to lower scaffold volume. This particular scaffold may be suitable for non load bearing anatomical sites but not for the reverse. In this study, PCL-TCP composites were selected as the delivery vehicle. Hydrophobic nature of PCL is improved by adding bioactive TCP particles. Mechanical strength of composite is increased as TCP scaffolds are originally brittle. Furthermore, PCL-TCP composites have shown to be biocompatible, controlled degradation rate and effective delivery systems for growth factors (Rai et al., 2005b; Yeo et al., 2008a). 4 Growth factors are “secreted by a wide range of cell types to transmit signals that activate specific developmental programs controlling cell migration, differentiation and proliferation” (Chen and Mooney, 2003). They transmit signals by attaching onto receptors on the cell surface. Signals will then be passed through the cell membrane and results in the expression of a target gene. This process is extremely complex and may involve multiple growth factors and receptors for one particular signal (Johnson et al., 1988; Pimentel, 1994). Some of the highly researched osteoinductive growth factors include Bone morphogenetic proteins (BMP-2), Transforming growth factor-β (TGF-β), Fibroblast growth factor (FGF), Insulin-like growth factor (IGF) and Platelet-derived growth factor (PDGF). BMP-2 is arguably the most potent osteoinductive growth factor and will be elaborated on in the next chapter (Bessa et al., 2008b). FGF stimulates neo-angiogenesis, indirectly augmenting bone regeneration by providing necessary nutrients to the core of defect site (Hurley et al., 1993; Rifkin and Moscatelli, 1989). IGF enhances the proliferation of osteoblasts ,osteoclasts and mineralisation (Khan et al., 2000). PDGF exhibits stimulatory effects on osteoblasts proliferation (Canalis et al., 1989) especially in bone fracture healing (Andrew et al., 1995). Drug delivery system (DDS) is a “technology that enables biological signalling molecules to enhance in vivo therapeutic efficacy by combination with biomaterials”(Tabata, 2005). Growth factors cannot be applied to the 5 defect site in solution because they will immediately diffuse away from the defect site. In addition, direct injection of growth factors at high doses has been shown to generate undesirable results (Yancopoulos et al., 2000). Negative feedback at high levels has been shown to induce heterotopic bone formation (Paramore et al., 1999) and even resulted in formation of antibodies in clinical trials (Walker and Wright, 2002). Here, a carrier is needed to ensure sustained delivery of a growth factor at the defect site. The release profile of the growth factor from the vehicle shall usually be gradual and controllable spatially and temporally to ensure maximum therapeutic effects (Chen and Mooney, 2003). 1.2 RESEARCH OBJECTIVES The general aim in this thesis was to evaluate and improve on the current properties of PCL-TCP composites from various aspects namely porosity and cell modulators in both in vitro and in vivo environments. The two specific aims of this research was 1. To investigate the effects of bone regeneration, scaffold degradation and mechanical properties of PCL-TCP scaffolds with different porosities in vivo. 2. To evaluate the effects of heparin on BMP-2 release and bioactivity from PCL-TCP scaffolds in vitro. 6 1.3 RESEARCH SCOPE In the first chapter, two different groups of scaffolds were randomly placed in rabbit calvaria defects and sacrificed after 2, 4, 8, 12 and 24 weeks. [Group A: ~75% porosity, Group B: ~82% porosity]. Upon sacrifice, at each interval, the specimens were subjected to µ-CT analysis, mechanical test and histomorphometric analysis. This will enable us to determine the effects of porosity differences on bone regeneration, scaffold degradation and mechanical integrity. The final chapter firstly examined the effectiveness of BMP-2 and heparin on pig osteoblasts in enhancing osteoblast differentiation. Pig osteoblasts was chosen to simulate implantation conditions which is in line with our future plan to assess the co-delivery system in a porcine model. Optimal concentration of both BMP-2 and heparin was then determined and adapted for subsequent analysis. Release profile of BMP-2 from PCL-TCP scaffolds was plotted for signs of sustained delivery when immersed in PBS solution. Heparin was chosen to improve binding and regulate release of BMP-2 here and concurrently act as a framework for binding of endogenous growth factors upon implantation. 7 CHAPTER 2: LITERATURE REVIEW 2.1 BONE PHYSIOLOGY In order to regenerate bone using BTE techniques, the understanding of the structure and cells which participate in bone repair is of utmost importance. Bone is composed of around 70-90% of minerals with the rest in the form of proteins. Within the proteins in bone, the ratio of collagenous to noncollagenous stands at 9:1. This is in stark contrast with other tissues consisting of only 10% collagenous proteins (Gokhale et al., 2001). High strength and rigidity of the bone stem is attributed to its mineral component, which is similar to hydroxyapatite (Ca 10 (PO 4 )(OH) 2 ). Bone has an elastic nature and it is also resistant to tension due to the high amount of collagen fibres. In Figure 2.1, there are two layers of bone namely: cortical (compact) and trabecular (spongy) bone. Cortical bone surrounds the outer layer of bone with thick and compact walls. It houses the medullary cavity where bone marrow resides during life. Trabecular bone, which has a spongy honeycomb structure, is only located at the epiphysis ends of long bone. Haematopoietic bone marrow also resides within the pores of trabecular network. With the exception of the articulating surfaces, the cortical bone is surrounded by the periosteum which is a thin layer of connective tissue made of a collagen rich layer and osteoprogenitor cells. 8 Figure 2.3: Diagram showing skeletal long bone structure which comprises of cortical and trabecular bone (Biomedical Tissue Research Group, 1996) There are a wide variety of cells that participate in the bone remodelling and regeneration process (Figure 2.2). The main duty of osteoclasts is to remove and resorb bone. When osteoclasts determine a bone site to be resorbed, it will create a barrier on its surface using its apical membrane. The pH level beneath osteoclast will be decreased, which triggers the formation of hydrogen ions and lysomal enzymes. After the resorption phase, Howship’s lacuna, which is a depression with ruffled border, is created (Raisz and Seeman, 2001). 9 Figure 2.4: Schematic diagram showing stages of bone remodelling process (Biomedical Tissue Research Group, 2007) In the renewal phase, macrophages are present at resorbed site. Osteoblasts, which had the ability to synthesis new bone, will continue to mineralise. During this process, it maintains and develops various channels with surrounding cells to facilitate various cellular actions through receptors and transmembrane proteins. When osteoblasts have completed the bone formation process, it will experience multiple transformations. It can convert itself into bone lining cells, which cover bone surfaces after quinesence phase. Some osteoblasts may be programmed to die after fulfilling its duties. The rest will become osteocytes and reside in bone matrix. Osteocytes are classified as mature osteoblasts and will not mineralise any further (Lian J 1999). 10 2.2 BIOMATERIALS 2.2.1 Polycaprolactone (PCL) Poly(ε-caprolactone) (PCL) is a semi crystalline resorbable polyester. PCL belong to aliphatic polyester family and thus share similar properties with other members such as polyglycolide (PGA) and polylactide (PLA). It has a low melting point of between 59 to 64°C, depending on its level of crystallinity. Low melting temperature enhances its processibility. It has a low glass transition temperature of around -60°C which explains its ductile and rubbery state at room temperature (Juan Pena, 2006). PCL has a higher decomposition temperature (350°C) relative to other aliphatic polyesters which will decompose between 235°C and 255°C. PCL also possess favorable mechanical properties: Elastic modulus between 300 to 400MPa which matches the stiffness of cancellous bone (100 – 300MPa) and a tensile strength which ranges from 15 to 60MPa (Zein I, 2002). Figure 2.3: Chemical Structure of PCL polymer (Wikipedia, 2006) Each monomer of PCL consists of five methylene groups and one ester group. PCL is hydrophobic due to the presence of non polar methylene groups 11 (Figure 2.3). Aliphatic ester linkage in PCL makes it susceptible to hydrolytic degradation. Low glass transition temperature contributes to the high permeability of PCL. It is this property that allows PCL to form copolymer blends with other polymers. PCL is widely used in its copolymer state in controlled release drug delivery applications (James M Pachence, 2000). PCL is an FDA approved material used widely in biomedical applications eg in sutures (Rezwan K, 2006). 2.2.2 Biodegradation The degradation mechanism of polymers used for bone tissue regeneration must be elucidated thoroughly before the product can be released into the market. The degradation profile of the scaffold will have a significant effect on the mechanical properties and various cellular activities that include host tissue response (Y. Lei, 2007). If the scaffold degrades well before sufficient bone regeneration take place, implant failure may result. Conversely, if the scaffold fails to degrade fast enough, it will act as a barrier and hinder new bone formation. PCL degrades completely in vitro and vivo to release harmless by-products. This is one advantage that PCL possess which make it highly suitable for use in medical devices. Unlike PCL, PLGA degrades upon implantation to form acidic byproducts which is toxic to the body and will affect cell growth and proliferation directly (Hak-Joon Sung, 2004). 12 The process of polymer degradation at different pH has been scrutinized by Burkersrodaa et al. The study concluded that polymer can either break down by surface erosion or bulk degradation. The mechanism of degradation is dependent on three factors: 1. size of matrix, 2. water diffusivity into scaffold centre, 3. rate of degradation of polymer reactive groups (A.S.Htay, 2004; Friederike von Burkersrodaa, 2002). PCL follows a two step degradation process when it is placed in an in vivo environment. The first step is a non-enzymatic, random hydrolytic ester cleavage which is triggered automatically by carboxyl end groups of the polymer chain. Chemical structure and molecular weight of polymer will affect the duration of the first step of degradation. When the molecular weight of polymer decreases to about 5000, second step of degradation will commence. The rate of chain scission and weight of polymer decreases as a result of the formation and removal of short chains of oligomers from the scaffold matrix. Fragmentation of polymer precedes the absorption and digestion of polymer particles by phagocytes or enzymes (C.G Pitt, 1981; Vert, 2002). 13 2.2.3 Tri-Calcium Phosphate (TCP) Figure 2.4: Chemical structure of tricalcium phosphate (CambridgeSoft Corporation, 2004) TCP is a biocompatible and biodegradable material used widely and successfully for bone replacement for many years. Some forms of calcium phosphates include: mono-, di- and tetra calcium phosphate, hydroxyapatite and β-whitlockie. The ideal Ca/P ratio is 1.6, which is similar to hydroxyapatite. The higher the Ca/P ratio, the more stable the compound will be in solutions (Lakes, 2007). TCP is found naturally in the inorganic phase of bone in form of hydroxyapatite. TCP is also responsible for the hardness of bone, dentine and enamel. TCP exhibit excellent regenerative activity when placed in vivo (Beruto et al., 2000). However, it has poor mechanical properties such as low compressive strength. This contributed to its brittleness when fabricated in blocks and scaffolds (K A Hing, 1998). 2.2.4 PCL-TCP scaffolds The purpose of using composites for medical applications usually is to reduce drawbacks of individual materials and the benefits of both are 14 combined together. Here, PCL is highly hydrophobic which leads to a longer degradation period (>2 years) in vitro and in vivo. TCP alone, when fabricated into a scaffold is brittle and weak in strength. By using PCL-TCP scaffolds for guided bone regeneration, the above disadvantages will be minimized. Previous research has showed that adding TCP to PCL by physically blending to produce composite scaffold, the degradation rate of PCL can be accelerated. In particular, under accelerated hydrolytic conditions of 5M NaOH, PCL-TCP scaffolds completely degrades at 48 hrs where PCL scaffolds require 6 weeks for degradation to complete (Christopher XF Lam, 2007). Our research team found that PCL-TCP scaffolds degrade to 40% by weight when it was immersed in standard culture media after six months. Recalling that the optimal degradation rate of scaffolds intended for dentoalveolar defects is about five to six months, the need for accelerated degradation propels our team to look into the possibility of alkaline and enzymatic degradation (Yeo et al., 2008a). 3M NaOH produced a more favourable surface morphology for bone regeneration relative to lipase treated PCL-TCP scaffolds. Alkaline treated scaffolds have a slower and more predictable degradation profile; whereas lipase treated ones have lower mechanical properties at each treatment point (Yeo et al., 2008b). Selective surface modification can be used to improve the surface hydrophilicity and pore morphology of biodegradable polyester scaffolds 15 without affecting the core of the rods. It was reported that cellular adhesion and proliferation are closely dependent on the topographical nature of the biomaterial surface (Boyan BD, 1995). An increase in surface area or roughness of scaffold matrices enhanced osteoblast response, which lead to improved osteoconductivity (Brett PM, 2004; Price RL, 2004). We showed that after NaOH treatment, surface wettability of PCL-TCP scaffolds increased significantly but the overall pore dimensions and honeycomb structure remains unaffected. Scaffolds subjected to longer alkaline treatments exhibited larger and deeper micro pits sizes, thus increasing the surface area to volume ratio favourable for better cell adhesion and bone growth (Yeo et al., 2010). PCL-TCP scaffolds had also been investigated as a delivery vehicle for BMP-2. In the novel DDS, fibrin sealant and BMP-2 were loaded onto PCLTCP scaffolds and their elution profile and bioactivity in different stages were analysed. Even though loading efficiency of PCL-TCP scaffolds stood at 43%, they were more uniformly distributed as compared with PCL scaffolds. PCLTCP scaffolds when loaded with 20µg/ml exhibit a triphasic release profile which had a delayed release profile than PCL scaffolds. BMP-2 also retained its bioactivity upon release at all time points (Rai et al., 2007; Rai et al., 2005b). PCL-TCP scaffolds have also proved to be a suitable delivery system for platelet-rich plasma as demonstrated by its sustained release in PBS and simulated body fluid (Rai et al., 2007). 16 2.2.5 Bone Morphogenetic Proteins (BMP-2) It has been discovered a century ago that bone has excellent regenerative capabilities. Ectopic bone formation was induced using decalcified bone or injected bone extracts in one of the earliest study on bone regeneration (Senn, 1889). The breakthrough came about when Marshall Urist discovered that bone formed at ectopic sites in rodents upon addition of proteins extracted from demineralised bone matrix. He named the protein “Bone Morphogenetic Proteins (BMP)” as its regenerative capability closely matched inherent bone repair process (Urist, 1965). BMP consists of a long hydrophobic stretch between 50-100 amino acids in length. BMP-2, prior to cell secretion, is made up of signal peptide, prodomain and mature peptide. Upon secretion, the signal peptide is cleaved (Xiao et al., 2007). BMP-2 belongs to the superfamily of transforming growth factor (TGF)-β. Members in the family mainly have roles in bone and cartilage development. Some other functions of BMPs include heart development (Callis et al., 2005; Simic and Vukicevic, 2005) and kidney formation (Simic and Vukicevic, 2005). There exists a heparin binding site in N-terminal region of mature BMP-2 polypeptide. In pioneering work by Ruppert in 1996, it was shown that BMP-2 activity was increased upon interactions with heparins present in ECM. In the presence of N terminus of BMP-2 and heparin, there is a five-fold increased in bioactivity (Ruppert et al., 1996). This lead to various studies on heparin effects on BMP-2 in mind of effective bone regeneration. 17 BMP signalling is important for morphogenesis to occur at a cellular level initially. There are two types of receptors: BMPR-1 and BMPR-2. Both have to work together for the signalling process. After BMP has bound itself strongly to the heteromeric complex of receptors, Smads proteins are activated instantaneously. Smads are nuclear effector proteins that are part of signalling pathway in BMP signalling cascades. There are three different groups of Smads: Common mediated Smads (C-Smads) – Smad 4, Receptor regulated Smads (R-Smads) – Smad 1, 5, 8 and inhibitory Smads (I-Smads) – Smad 6 and 7. In Figure 2.5, following the adhesion of BMP-2 to BMPR-1 receptor, phosphylation of R-Smads will be follow. Phosphylated R-smads will form heteromeric complex with Smad 4 and be translocated into the nucleus. Transcription of target gene occurs in the presence of transcription factors and heteromeric complex. Signalling is regulated by inhibitory Smad 6/7 (Vukicevic and Sampath, 2008). 18 Figure 2.5: Activation of SMAD proteins after BMP mediation (Sakou, 1998) However, BMP-2 has several disadvantages. It degrades rapidly in vivo (Yamamoto et al., 2003). Excessive dosages of BMP-2 were shown to trigger bone formation away from defect site (Valentin-Opran et al., 2002). Furthermore, subsequent administrations of BMP-2 will be costly. Hence, for BMP-2 to function effectively in treatment of bone defects, more need to be done in the following areas (Bessa et al., 2008a): 1. Optimised variables for clinical translation. 2. Good carrier biocompatibility and biodegradability. 3. Efficient BMP-2 loading method. 4. Sustained released targeted at defect site. 5. Bioactivity of eluted BMP-2 maintained. 19 2.2.6 Heparin The discovery of heparin occurred in John Hopkins University in 1916. While researching on the cause of blood clotting, Jay McLean accidentally collected substances that inhibit clotting. It was named heparin thereafter. Following this discovery, more efforts were targeted at understanding the structure of heparin. The first commercialisation of bovine lung and porcine intestinal heparin was carried out in Toronto and Stockholm. Figure 2.6: Chemical structure of Heparin (Gray et al., 2008) Heparin is glycosaminoglycan a sulphated (GAG) family polysaccharide (Figure 2.6). which They belongs are to linear heteropolysaccharides that alternates between glucosamine and iduronic acid (Lever and Page, 2002). It is one of the most negatively charged molecules relative to its small size. Sulfates and carboxylates groups contributed to the high net negative charge (Caughey, 2003). Size of low molecular weight heparin used in this study varies from 2-10kD. Heparin is mainly found in mast or granulated cells in various organs. Heparin is also widely used as an anticoagulant in the treatment of stroke and coronary artery disease. 20 There are heparin binding sites in some growth factors which played an important role in their modulation of cell activities. It has been shown that heparin binds strongly to proteins with highly positive-charged binding sites. However, heparin sulphates favoured sites where basic residues are far apart from each other (Fromm et al., 1997). Aside from the usual ionic bonding, heparin also interacts with proteins through hydrogen bonding and hydrophobic forces (Bae et al., 1994). Heparin and heparan sulphates are structurally 70% similar to each other. Heparan sulphates are less sulphated and thus have a lower overall negative charge than heparin. Even though they are made up of same units, heparan sulphates have a higher glucosamine and lower iduronic acid component (Lever and Page, 2002). Despite the differences, heparin has been used widely as a model for the costly heparan sulphate. Heparin has been demonstrated to possess binding affinities for growth factors including VEGF and BMP-2 (Ruppert et al., 1996). As a result, heparin has been incorporated into biomaterials for the purpose of binding to endogenous growth factors (Nillesen et al., 2007; Steffens et al., 2004). Heparin has been shown previously to enhance osteoblast differentiation through BMP2 activity in vitro (Ruppert et al., 1996). In a recent study, it was reported that heparin prolonged BMP-2 degradation by 20-folds in culture medium. In 21 addition, higher bone mineral density was observed in subcutaneous implants when both heparin and BMP-2 were present (Zhao et al., 2006). 22 CHAPTER 3: EFFECTS OF POROSITIES OF PCL-TCP SCAFFOLDS ON BONE REGENERATION, SCAFFOLD DEGRADATION AND MECHANICAL PROPERTIES. 3.1 INTRODUCTION As highlighted in chapter 1, shortcomings of current bone grafts motivate researchers to look into polymeric scaffolds for better substitutes. Important characteristics for a scaffold to function effectively as bone void filler at defect sites include: 1) Highly porous and well connected pores to facilitate cellular and vascular infiltration. 2) Biodegradable and predictable degradation profile which coincides with surrounding tissue growth. 3) Surface characteristics that allows for greater osteoblast attachment and function. 4) Mechanical properties that is similar to defect site and is able to withstand load right after implantation. 5) Easy loading of cells and proteins that are able to induce bone regeneration. (Hutmacher et al., 2001; Jones, 2005; Temenoff and Mikos, 2000). In this chapter, we will focus on the porosity of scaffold. Pores are a necessary feature in scaffolds as they allow for cellular migration and proliferation. Porosity of a scaffold is defined as the ratio of voids to the overall volume occupied by the scaffold. Porosity and pore size of a scaffold are closely interrelated with each other. Scaffolds with higher porosity will have larger pore size provided that they remained constant throughout. These are structural properties of scaffolds which is independent of the material. In 23 addition, it determines the flow of nutrients and metabolic wastes through the scaffold (Kuboki et al., 1998). Optimal pore size has always been highly debated amongst researchers. A wide range of pore sizes from 10 – 600 µm have been tested in BTE with porosities from 43 to 87.5% (El-Ghannam, 2004; Lickorish et al., 2004; Roy et al., 2003; Zhang and Zhang, 2002). New bone growth was observed in all defects in above studies. Noteworthy, scaffolds with engineered channels exhibited larger new bone area compared to non porous ones and those that were left unfilled (Roy et al., 2003). It has been previously shown that direct osteogenesis was seen in pore sizes larger than ~300µm due to increased vascularisation. Conversely, osteochondral ossification was facilitated when pore size falls below 300µm (Gotz et al., 2004; Karageorgiou and Kaplan, 2005; Kuboki et al., 2001; Tsuruga et al., 1997). Thus, it can be seen that porosity and pore size has a great influence on the bone regeneration mechanism at bone defects. PCL-TCP scaffolds fabricated by Fused Deposition Modelling (FDM) have a unique and consistent architecture. PCL-TCP scaffolds used in this study have a 0/60/120° lay-down pattern. Angles here are with respect to the first layer and parallel to polymer rods spaced evenly apart from each other. At the forth layer, the pattern repeats itself to produce scaffold with triangular pores when viewed from above. A regular distribution of pores is visible from the 24 side. By changing FDM parameters, scaffolds of different porosity and pore size can be customised. Fully interconnected pores and porosity of 65% allowed canine osteoblasts to attach and proliferate on PCL-TCP scaffolds (Rai et al., 2004). Combination with platelet-rich plasma in dog mandible demonstrated higher bone regeneration and similar scaffold degradation relative to controls (Rai et al., 2007; Rai et al., 2005a; Rai et al., 2005b). Previous analysis by our group showed that surface modification with alkaline treatment created micropores on rods of scaffolds and increases its surface roughness concurrently. However, its mechanical properties were not compromised; instead increased bone growth was observed within surface modified scaffolds upon implantation (Yeo et al., 2008b; Yeo et al., 2009a, 2010). The purpose of this study was to evaluate the effects of varying porosities of PCL-TCP scaffolds on bone regeneration, scaffold degradation and mechanical properties in a rabbit calvarial model. Compressive and shear strength of scaffolds were obtained through mechanical testing. Micro CT and histomorphometry analyses provided us with information on scaffold degradation and bone ingrowth. Histology was used to examine scaffoldtissue interactions at a cellular level. 25 3.2 MATERIALS AND METHODS 3.2.1 Scaffold Fabrication Scaffold specimens (Osteopore International Pte Ltd, Singapore) were fabricated with PCL- 20% TCP filaments using a fused deposition modeling (FDM) 3D Modeler RP system from Stratasys Inc (Eden Prairie, MN). Blocks of 50 x 50 x 2mm were created directly in Stratasys Quickslice (QS) software. A lay-down pattern of 0/60/120o was used to give a honey-combed like pattern of triangular pores. The specimens were cut into smaller discs of 6mm in diameter and 2mm in thickness subsequently. PCL-TCP scaffolds were immersed in ethanol for sterilisation. This was followed by rinsing 3x in phosphate buffer saline (PBS, 137 mM NaCl, 2.7 mM KCL, 10 mM Na 2 HPO 4 , 1.8 mM KH 2 PO 4 , pH 7.4). Scaffolds were dried overnight prior to implantation. 3.2.2 Porosity Calculation Porosities were calculated by first measuring the weight and volume of each sample. Apparent density of the scaffolds was calculated using the following formula: ρ * = m (g) / V (cm3). Finally, scaffold porosities = ε = 1- ρ * / ρ x 100 % (ρ = 1.17 g/cm3) were obtained. 3.2.3 Experimental Design The scaffolds were randomly assigned to the defects made in the calvaria of rabbits and followed up for 2, 4, 8, 12 and 24 weeks. 26 Two groups of PCL-TCP scaffolds (Figure 3.1) were analysed: Group A: ~75% porosity Group B: ~85% porosity B A 1 mm 1 mm Figure 3.1: Micro CT images of scaffolds before implantation (A) Group A (B) Group B Micro-CT, mechanical strength testing (compressive strength and push out test), histology and histomorphometric analyses were performed. A minimum of 12 samples were required for each experimental group at 2, 4, 8, 12 and 24 weeks to have sufficient data for analysis as recommended by ISO standard 10993-6. Since we had a total of 24 samples at each time point and 2 samples were implanted in each rabbit, 60 rabbits were needed for the entire study. 3.2.4 Animal husbandry and scaffold implantation Sixty, 6-8 month old New Zealand White male rabbits were used. The study was approved by the SingHealth Institutional Animal Care and Use Committee (IACUC) and conformed to the respective guidelines. The rabbits 27 were operated on under general anesthesia, which consisted of an intraperitoneal injection of ketamine and xylazine mixture (75 mg/ kg + 10 mg/ kg). Under anesthesia, the skull region of the rabbit was shaved and scrubbed with iodine, followed by disinfection with 70 % ethyl alcohol. A midline incision was made in the skin of the calvaria along the sagittal suture line. The soft tissue and periosteum are elevated and reflected. Under constant saline irrigation, 6 mm diameter circular and 2 mm deep defects were made using the appropriate trephine drills. A total of 2 circular defects were made on the calvarium of each rabbit. Care is taken to preserve the dura. Defects were randomly assigned to receive 1 of the 2 test scaffolds (Figure 3.2). Prior closure, a non-resorbable membrane was positioned over the defects to prevent soft tissue ingrowth. This was followed by repositioning of the periosteum to cover the scaffolds followed by closure of the skin with sutures. The rabbits were then given carpofen (1-2 mg/ kg) and cephalexin (1520 mg/ kg) subcutaneously for 3 and 5 days respectively. Figure 3.2: Implantation of scaffolds into rabbit calvarial defects 28 Twelve rabbits were euthanized at 2, 4, 8, 12 and 24 weeks respectively. All samples were processed and analyzed accordingly. The tissues surrounding the selected implanted scaffolds were carefully removed and stored in 10 % neutral buffered formalin (NBF) for histology (n = 3). The remaining samples were wrapped in PBS-soaked gauze and frozen at -20 ºC and subsequently subjected for micro-CT (n = 4; non-destructive), push out (n = 4), compressive strength testing (n = 4) analyses. The rabbits were closely monitored everyday for the first week for presence of swelling, pain and infection. They were then observed weekly for severe weight loss (20-25 %) and any other complications. 3.2.5 Micro-CT analysis PCL-TCP scaffolds were scanned isotropically at 14 μm resolution with an SMX-100CT micro-CT scanner (Shimadzu, Japan) using a cone CT scanning technique. To ensure a consistent CT image resolution among all the datasets, the scanner turntable location was fixed at a specific source-to-object distance (SOD; 48.03 mm) and source-to-image distance (SID; 361.30 mm) respectively. X-ray parameters were set at 33 kV and 156 μA and the CT images were processed at a scaling coefficient of 100 and averaged 3 times. Region of interest (6 mm diameter) was drawn at the site of implantation. Resultant micro-CT datasets for each bone cube were evaluated for microarchitectural parameters using VG Studio Max software (Heidelberg, Germany). Overall bone formation and scaffold volume loss were obtained and analyzed. 29 3.2.6 Mechanical strength testing Compressive test Two kinds of mechanical testing (compressive test and push out test) were performed using the Instron 5500 micro tester (Instron, Canton, MA) with a 1kN load cell. Each specimen was placed between 2 flat plates for compression testing. The scaffolds were compressed at a speed of 1mm/ min up to 80% of scaffold original thickness at room temperature. The mechanical results of load and extension were used to calculate the σ, Compressive Stress (MPa) = [Load (N) / Area (m2)] x (1 x 10-6); ε, Strain = Extension (mm) / Original Length (mm); and E, Elastic Modulus (MPa) = Compressive Stress (MPa) / Strain. A stress-strain curve was then plotted using the experimental data (load versus deformation) and the compressive modulus and strength was recorded for each specimen, with the stiffness being measured as the slope of the linear portion of the curve. Push out test Utilizing a similar Instron 5500 micro tester, push out test was done using the same configuration except that a support jig with a hole of 7.2 mm was used. Schematic diagram was shown previously (Yeo et al., 2009a). Interfacial shear strength between old bone and the newly regenerated bonescaffold composite was calculated by dividing peak force with cross sectional area of specimen. 30 3.2.7 Histological Analysis The specimens were removed and stored in neutral buffered formalin (NBF) 4% and 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 EXAKT micro grinder system (EXAKT Technologies, Inc., Oklahoma City, OK) and were subsequently stained with Goldner’s trichrome to identify new bone formation. Three defects per group were used at each time point per analysis. 3.2.8 Histomorphometric Analysis Images of stained sections were taken with an Olympus SZX12 microscope connected to a CCD camera and measured using Bioquant Image Analysis® software (Nashville, TN, USA). The region of interest (ROI) was defined as area containing tissue within the defect site reaching from the periphery of the host bone. New bone was quantified by selecting a fixed threshold for positive stain pixels (black for Von Kossa). Percentage of bone volume per tissue volume (%BV/TV) was calculated by the following formula. 31 3.2.9 Mineral Apposition Rate (MAR) Prior to the animal sacrifice, fluorochrome markers alizarin red and calcein green were administered into the rabbits at a time interval of 10 days. Using unstained slides under fluorescence, high magnification (20x) images were obtained for regenerated bone in the defect site using an epifluorescence microscope (model BX51; Olympus). Up to 10 fields of view were taken for each specimen. Two images were taken at each site using a green and a red filter. Images were then merged with Adobe Photoshop and the inter-label distances were measured using Bioquant Image Analysis® software (Nashville, TN, USA) (Figure 3.3). The distances were taken from the center of the first label to the center of the second label perpendicular to the labelled surfaces. Mineral apposition rate was calculated using the formula below: Figure 3.3: Calculation of interlabel distances using Bioquant Image Analysis® software 32 3.2.10 Statistical Analysis All values in this study were presented as mean values ± standard deviation (SD). Data analyses and comparisons were performed using Student’s paired t-test (Microsoft Excel). P-values of < 0.05 were considered as statistical significance. 3.3 RESULTS Throughout the surgery and post implantation period, rabbits showed no signs of infections and were sacrificed at respective time points. Implanted scaffolds remained in position throughout the entire study. Upon explantation after 12 weeks, all defects created were completely filled. Samples felt hard and vessels could be observed in all specimens. 3.3.1 Scaffolds characterisations Group + Measured Porosity+ (%) Pore size (µm) Rod size A 74.9 ± 1.7 500 ± 73 0.163 ± 0.02 B 86.7 ± 0.2 723 ± 92 0.143 ± 0.003 (mm) Values obtained from manufacturer. P < 0.05 was observed for all parameters Table 3.2: Physical parameters of PCL-TCP scaffolds TCP particles appeared as bright white particles relative to grey PCL as shown in µ-CT images (Figure 3.1). They were shown to be evenly distributed throughout the scaffold. The fused deposition modelled scaffolds were fully 33 interconnected and highly porous. (Group A: 74.9 ± 1.7%, Group B: 86.7 ± 0.2%) The pore size distribution is relatively consistent from the small standard deviation. The pore size of Group B was 36.9 – 42.3% higher than Group A. 3.3.2 A µ-CT analysis B Figure 3.4: Representative µ-CT images of explants specimens with PCL-TCP scaffold of (A) 75% porosity (B) 85% porosity (blue: scaffold, yellow: new bone growth and beige: calvaria bone) Figure 3.5: Percentage BV/TV in PCL-TCP scaffolds by µ-CT with varying porosities from 2 to 24 weeks of implantation. 34 Before analysis of specimens, calibration was done on rabbit calvarial bone and scaffolds which served as a benchmark for detecting and to distinguish various components present in specimens (Figure 3.4). Bone volume (BV) here refers to the amount of new bone growth detected at defect site. Total volume (TV) is total volume available for growth at the defect site. %BV/TV was observed to increase from the start to 8 weeks and stabilised thereafter. There were no significant differences between Group A and B throughout the experiment (Figure 3.5). Figure 3.6: Scaffold volumes with varying porosities over a time period of 24 weeks. Scaffold degradation in rabbit model seemed to follow a power law relationship, regardless of porosity (Figure 3.6). However, there is an additional constant which indicates that scaffold volume will fall to a specific volume after prolonged periods of scaffold implantation. Scaffolds were observed to degrade gradually throughout the study. At 24 weeks, scaffolds 35 with both porosities had volumes that closely matched each other (Group A: 8.34 ± 0.7mm3; Group B: 7.5 ± 0.25mm3). Figure 3.7: Percentage of PCL-TCP scaffolds volume loss with varying porosities from 2 to 24 weeks of implantation Figure 3.7 indicates that 50% scaffold loss was achieved at different time points for both groups. Group B scaffolds reported 50% degradation at around 6 weeks. This was faster than Group A scaffolds that demonstrated 50% scaffold loss at around 10 weeks. Between 12 to 24 weeks, degradation of scaffolds seemed to stabilize at 50-60% and there was no difference between groups. 36 3.3.3 Compressive strength Figure 3.8: Compressive strength of PCL-TCP scaffolds with varying porosities from 2 to 24 weeks of implantation (* denotes p < 0.05) Compressive strength in this study was taken at 50% strain to compensate for the unevenness of the specimens. Compressive strength for Group B was significantly lower at 2 weeks than Group A. Group A scaffolds experienced a increase of 103% from 2 to 12 weeks and Group B scaffolds peaked at 8 weeks and decreased after that (Figure 3.8). Porosity variations did not have a significant effect on the amount of compressive strength. 37 3.3.4 Push Out test Figure 3.9: Shear strength of PCL-TCP scaffolds with varying porosities from 2 to 24 weeks of implantation (* denotes p < 0.05) Groups A and B showed a similar trend for shear strength at different time points even though the decrease in Group B compared to Group A was significant at 12 weeks (Figure 3.9). This suggests that the differences in porosity had no immediate effect on shear strength. Shear strength of both groups seemed to increase from 2 to 4 weeks and stabilises thereafter. 38 3.3.5 Histology Figure 3.10: Representative images of ex vivo specimens at (A) 4 weeks (B) 8 weeks stained for Goldner’s Trichome (green sections: mineralised bone; areas labelled ‘S’ denote PCL-TCP scaffolds) Harvested samples were stained with Golder’s Trichome to identify new bone formation in the defect site. New mineralised bone (green stains) was detected between scaffolds rods and more evidently at later part of the study. No significant differences were observed between groups. Overall, the absence of fibrous encapsulation was an indication that PCL-TCP scaffolds were compatible with its surroundings. The non-resorbable membrane (grey lining) 39 served as an orientation due to its position as top layer in the specimens. Multi-nucleated giant cells were observed on the surface of scaffolds but not on mineralised bone. Neovascularisation, specifically in the presence of vessels was also observed at the defect site. 3.3.6 Histomorphometric Analysis Figure 3.11: Percentage BV/TV of PCL-TCP scaffolds by histomorphometry with varying porosities from 8 to 24 weeks of implantation New bone formation within defect site was divided by total defect area to obtain %BV/TV by histomorphometry (Figure 3.11). Similar trend was observed using µ-CT in Figure 3.5. Bone mineralisation rate was measured by fluorescence quantification from 8 to 24 weeks. The images demonstrated distinct clear labels in the defect sites. This was an indication that new bone formation was actively taking place. Flurochrome labels for 2 and 4 weeks were spread out and diffused which did not allow for accurate measurements (Figure 3.12). 40 Figure 3.12: Representative image showing diffused flurochrome labels between 2 and 8 weeks (red label – alizarin, green label – calcein) 3.3.7 Mineral Apposition Rate (MAR) Figure 3.13: Mineral Apposition Rate (MAR) of PCL-TCP scaffolds with varying porosities from 8 to 24 weeks of implantation There was no significant differences between MAR of Group A and Group B. MAR for Group A peaked at 12 weeks (1.38µm/day) and Group B at 8 weeks (1.52µm/day). Both groups exhibited a reduction in MAR to 0.92 - 0.97 µm/day at 24 weeks. 41 3.4 DISCUSSIONS Pore size and porosities serve as important material properties which enable comparison between different scaffolds in bone regenerative and scaffold degradation abilities. Pore size is important in bone regeneration as they act as channels for cells and vessels to enter matrix for effective bone formation. It also directly linked to mechanical strength of the scaffold. This is especially vital for implantation to load bearing sites. Here, scaffolds of 75% and 85% were chosen as previous studies by our group demonstrated promising results with porosities between 70 – 90% and also to determine whether there are any differences in bone regeneration in vivo with the porosity difference. Current literature reported conflicting results ranging from increased bone regeneration for scaffolds with higher porosity (Ikeda et al., 2009; Kuhne et al., 1994; Schliephake et al., 1991); lower bone growth with increased porosity (Eggli et al., 1988; Flautre et al., 2001) and limited or no effects at all. (Kasten et al., 2008; Roosa et al., 2009; Takahashi and Tabata, 2004). This propels our team to investigate the effects of pore size and porosity on previously researched PCL-TCP scaffolds that has proved its effectiveness in bone regeneration through numerous studies (Yeo et al., 2008a; Yeo et al., 2009b). The objective of this study was to compare the effects of two different 42 porosities of scaffolds on bone growth and scaffold degradation in a rabbit calvarial defect. µ-CT characterisation of native scaffolds indicated that Group A (75% porosity) had smaller pore sizes and larger rod diameters than Group B (85% porosity). All the parameters indicated in Table 1 are closely related to porosity. It has been shown previously that when the diameter of rods/fibers is comparable or smaller than cells, quality and quantity of cellular attachment will be reduced (Karageorgiou and Kaplan, 2005; Takahashi and Tabata, 2004). The rods diameters’ of PCL-TCP scaffolds used in this study are about 100x larger than size of cells which are about 10µm. µ-CT analysis has been used widely to determine the extent of bone regeneration in BTE. This is mainly because it is a non-destructive and fast technique. It also allowed us to distinguish between scaffold, new bone formation and surrounding calvaria bone within the specimen (Figure 3.6). The % BV/TV was found to increase by 68.6 - 100.6% from 2 to 8 weeks for both groups. Immediately after implantation, increased blood flow at defect site, as indicated by the presence of blood vessels in histology, initiated the cascade of wound healing and tissue ingrowth. Bone formation at defect site was clearly detected on µ-CT images (Figure 3.4). After 8 weeks of implantation, bone growth for all groups generally seems to have stabilized. The reduction of growth rate across all 43 groups suggests that bulk of bone remodelling is taking place (Yeo et al., 2009b). Following µ-CT analysis, the mechanical properties of scaffolds were analysed. An implant can collapse if bone fails to infiltrate enough before scaffold is resorbed considerably. However, scaffold degradation profile is often neglected in previous studies as the emphasis is usually on bone regeneration (Aronin et al., 2009; Heo et al., 2009). In this study, overall degradation profile of PCL-TCP scaffolds in a rabbit model was mapped out and analysed. We showed that PCL-TCP scaffolds degraded according to the power law relationship described below: (3.1) Where y refers to current scaffold volume x denotes time after implantation in weeks C and n refer to constants depending on scaffold properties and environment. d denotes scaffold volume that will remain after prolonged periods of implantation This equation allowed us to predict remaining scaffold volume (mm3) at a specific time, thus enabling unobstructed bone growth by using an appropriately selected scaffold. Group B (85% porosity) scaffolds broke down more rapidly than Group A (75% porosity) before 12 weeks. After which, the volume of scaffold remained constant. Here, degradation of PCL-TCP scaffolds followed the degradation profile of PCL which degraded in two 44 stages namely: 1) surface erosion 2) bulk degradation. The first step is the non enzymatic, random hydrolytic ester cleavage which is triggered automatically by carboxyl end groups of the polymer chain. Chemical structure and molecular weight of polymer will affect the duration of the first step of degradation. When the molecular weight of polymer decreased to about 5000, second step of degradation will commence. The rate of chain scission and weight of polymer will decrease as a result of the formation and removal of short chains of oligomers from the scaffold matrix. Fragmentation of polymer precedes the absorption and digestion of polymer particles by phagocytes or enzymes (C.G Pitt, 1981; Vert, 2002). TCP particles in PCL scaffolds are more hydrophilic thus preferentially eroded when implanted. This causes TCP particles to dislodge and increases surface area of PCL thereby accelerating surface erosion (Lam et al., 2009; Yeo et al., 2010). Further analysis of previous degradation profile of alkaline treated PCL-TCP scaffolds also pointed to power law relationships, but to a lesser extent (lower R2 value). This can be explained by formation of pits on scaffold surface after alkaline treatment leading to a faster surface erosion process (Yeo et al., 2009b). Here, first stage of degradation can be inferred to occur throughout the entire experiment. Thus, degradation profile here (Equation 3.1) is only valid during the surface erosion of PCL-TCP scaffolds. Bulk degradation, which is a slower process will follow thereafter, may have a degradation profile which is completely different from Equation 3.1. 45 Compressive strength of specimens can be regarded as a composite which is guided by the rules of mixture or Voigt model (Lakes, 2007). (3.2) Where σ refer to compressive strength, f refers to volume fraction, denotes total strength, subscript S refers to scaffold component and subscript B refers to bone component. Compressive strength is heavily influenced by the amount of bone in the defect site. This is evident in the close resemblance in the compressive strength and bone volume charts. Structural integrity of specimens was evaluated by measuring compressive and interfacial shear strength. There was no significant differences between Groups A and B, except at 2 weeks. The initial lower strength for Group B is probably due to low compressive strength of scaffold as bone in-growth is still insufficient to influence overall strength. Strength of bone scaffold is much larger than . Thus, the dominant factor in total strength changed from scaffold strength to new bone strength during the experiment. Low volume fraction of scaffolds ) (high porosities for Group A and B) also contribute to the above. Push out tests, which have been used widely to determine the mechanical strength of implants at the interfaces in animal studies (Müller et al., 2006; Sawyer et al., 2009). In this study, emphasis is placed on shear strength between scaffolds in defect site and surrounding calvaria walls. 46 Increase from 2 to 4 weeks has been discussed previously. There were no significant differences between Group A and B except at 12 weeks. The initial increase before 4 weeks can be attributed to the bone growth starting from the calvaria wall surrounding defect. After that, shear strength stabilised probably due to sufficient bone has already infiltrated defect site at 4 weeks and new bone continues its growing path into the centre of scaffolds. The above is supported by %BV/TV values and histology images. Histological analyses at various time points affirmed the findings from µ-CT in that bone has penetrated scaffold voids more evidently at a later stage. Adipose tissue was also residing alongside with bone. There was no evidence of fibrous encapsulation of scaffolds in all specimens observed. This confirmed PCL-TCP biocompatibility in an in vivo which attests to previous research findings (Lam et al., 2009; Rai et al., 2007; Yeo et al., 2009b). Multi-nucleated giant cells were observed to be only on the surface of scaffolds rods. Thus, they were probably involved in the breaking down of PCL-TCP scaffolds and not the newly formed bone. Vascularisation observed within regenerated defect sites for both scaffolds indicate that vessels were able to infiltrate scaffolds with pore sizes above 500µm, which is agreement with studies discussed previously (El-Ghannam, 2004). Histomorphometric analyses largely corroborated with µ-CT findings in terms of bone volume. However, two and four week’s data were omitted 47 due to failure in bone detection. Bone was actively forming and new woven bone does not have an organised structure that may have contributed to the diffused flurochrome labelling, which negates any measurements. Notably, MAR was able to provide us with bone growth rate in a lamellar fashion using distances between inter-label distances. Generally, MAR seems to be declining from 8 to 24 weeks and this might be due to remodelling that reduced osteoblastic activity (Yeo et al., 2009b). This may be attributed to the regeneration of cancellous calavria bone within defects at 12 weeks. Empty pores within new bone growth evident in histology images can thus be considered as trabecular structure of rabbit calavria. Previous studies in rabbits had resulted in similar findings compared to our current study. In an eight week study by Fisher et al, implantation of poly(propylene fumarate) scaffolds subcutaneously in rabbit calvaria did not indicate any significant differences when pore sizes (300 – 500µm and 600 800µm) and porosities (57 – 75%) were varied (Fisher et al., 2002). Identical findings were obtained when nitinol implants of varying pore sizes (179, 218 and 353 µm) and porosities (54, 51, 43%) were implanted into rabbits for 6 weeks (Ayers et al., 1999). Minor differences between scaffold groups may have diminished its effects on bone regeneration and scaffold degradation in this study. Only two groups of untreated scaffolds were presented here as a 48 concurrent study with alkaline treated PCL-TCP scaffolds which has been published elsewhere (Yeo et al., 2010; Yeo et al., 2009b). In the above studies, including our current work, the thickness of the specimen was in the same order of magnitude as the implant diameter. One limitation in adopting rabbit calvarial defect model lies in the thickness of implants must conform to that of calvaria. Here, PCL-TCP scaffolds has a thickness of 2mm. The impact of porosity variations on bone regeneration and scaffold degradation was reduced as cellular and vascular infiltration is better in thin models (Ayers et al., 1999). To address this problem, larger animal models which enable thicker scaffold implants should be used for studies with porosity and pore size as variables. This is in agreement with the hypothesis that there should be a critical thickness to pore size ratio that will have a significant impact on bone regeneration (Ayers et al., 1999). 3.5 CONCLUSIONS This study aimed to investigate the effects of porosity and pore sizes of PCL-TCP scaffolds on bone regeneration, scaffold degradation and mechanical properties in rabbit calvarial defects over 24 weeks. Results indicated that there were considerable bone growth and scaffold degradation especially at the later stages of the study. Bone remodelling has already taken place at 24 weeks evidently from histology images but shear strength of implants were 49 not compromised. Absence of inflammatory cells affirmed previous findings that PCL-TCP scaffolds are biocompatible. Histomorphometric analysis affirmed µ-CT and mechanical test results. In all, pore size and porosities have negligible effects on bone regeneration, scaffold degradation and structural integrity in a rabbit calvarial model. 50 CHAPTER 4: PRELIMINARY EVALUATION OF PCL-TCP SCAFFOLDS AS CO-DELIVERY SYSTEMS FOR HEPARIN AND BMP-2 IN VITRO 4.1 INTRODUCTION In Chapter 3, ability of PCL-TCP scaffolds as a biomaterial upon implantation into rabbit calvarial has been proven. However, acellular scaffolds may not be the best option when treating large defects. Initial rigidity of scaffold may not be sufficient to protect underlying brain unless new bone formation is accelerated (Salyer et al., 1995). Thus, the use of growth factors eg bone morphogenetic protein-2 (BMP-2) in this chapter come into play. Documentations of the effectiveness of BMPs can be traced back to the 1960s. Urist found that BMPs are the proteins involved in bone healing. Subsequently, BMP-2 has demonstrated its powerful osteoinductive abilities in various studies. BMP-2 belongs to the superfamily of Transforming growth factor (TGF)-β. BMPs ha s been shown to pla y an active role in cellular regulation and function in bone function and repair. Furthermore, BMP-2 is approved by FDA for use in fracture and spinal fusion applications (Jones et al., 2006; Mont et al., 2004). However, BMP-2 has short half life in vivo (Yamamoto et al., 2003). Sustained release of BMP-2 has not been achieved at defect sites, thus researchers turned to the usage of high and multiple doses (Jones et al., 2006). 51 This resulted in adverse effects including bone growth away from the defect site (Lieberman et al., 2002; Valentin-Opran et al., 2002) and high cost of BMP-2 incurred for repeat administrations. Therefore, it is essential to develop a system where the delivery of bioactive BMP-2 to the defect site is both localised and sustained. PCL-TCP have been shown to demonstrate excellent properties as bone void fillers in chapter 3 and in previous studies (Lam et al., 2007; Rai et al., 2007; Yeo et al., 2009a). Notably, PCL-TCP scaffolds had also been investigated for its delivery properties with BMP-2 in combination with fibrin sealant. PCLTCP scaffolds when loaded with 20µg/ml of BMP-2 exhibited a triphasic release profile, that was sustained. BMP-2 retained its bioactivity even after 21 days. (Rai et al., 2005b). Heparin is a sulphated polysaccharide which belongs to glycosaminoglycan (GAG) family. It has been demonstrated to have an affinity for growth factors namely BMP-2, FGF and VEGF (Steffens et al., 2004; Wissink et al., 2000). Heparin binds to BMP-2 directly through the N-terminal region of the mature polypeptide (Chung et al., 2007). BMP-2 bioactivity is prolonged as a result. The bioactivity and half-life of BMP-2 in culture medium were increased by 20-fold when heparin was present (Zhao et al., 2006). In view of the above, a co-delivery system of heparin and BMP-2 with PCL-TCP scaffolds was proposed in this study. The effects of heparin and 52 BMP-2 on pig osteoblasts were first investigated for increased osteogenic differentiation. The concentrations of heparin and BMP-2 with heightened ALP activities were chosen for subsequent analysis. Following that, BMP-2 delivery profile in PBS buffer was mapped out and evaluated. 4.2 4.2.1 MATERIALS AND METHODS Porcine osteoblasts culture Porcine bone chips were obtained from pigs following their sacrifice after surgical procedures as per ethical code. Bone chips (~2-3mm thickness) were harvested from the mandible and incubated in a solution of PBS (containing antibiotics/antimycotic) for transport back to the laboratory. They were then vortexed in PBS (containing antibiotics/antimycotic) until the hematopoeitic tissue was removed from the bone chips. The bone chips were dissected into pieces of 2mm x 2mm. Bone chips were then immersed in Alpha MEM media supplemented with 10% FBS (Hyclone, USA) and 1% PS at 37°C and 5% CO 2 , untouched for 1 week, before changing media. Medium was replaced every 3 days and cells were passaged when they are 70% confluent. 4.2.2 Cell culture and BMP-2 treatment Cells at passages 3-5 were used throughout the experiment. Pig osteoblasts (20000 cells/cm2) were plated in 24-well plates. Cells were inoculated in 3 ml of osteogenic medium (Alpha MEM, 10% FBS (Hyclone, USA), 1% PS, 1µM Dex, 53 0.05mM absorbic acid and 10mM glycerol-2-phosphate) for 24 hrs. Following that, medium was changed and replenished with treatment medium containing varying concentrations of BMP-2 (R&D Systems, USA). 30µl of BMP-2 (10, 100, 300 and 1000ng/ml) was added to the treatment medium and cells were cultured for 3 days. Cells were lysed by rinsing with 1ml PBS/1mM PMSF twice on a rotary shaker first. After addition of RIPA buffer (1% Triton X-100, 150mM NaCl, 10mM Tris pH7.4, 2mM EDTA, 0.5% Igepal, 0.1% SDS) and 1% of protease inhibitor(Calibiochem, Germany), cells were scraped using pipette tips and placed on rotary shaker (100rpm) at 4°C for 15 minutes. Suspensions were then centrifuged down at 11,000rpm at 4°C for 10 minutes and supernatants were extracted for further analysis. All experiments were carried out in triplicates. 4.2.3 Heparin and BMP-2 treatment After the optimal concentration of BMP-2 with heightened ALP activity was determined, concentrations of heparin (10, 100, 1000ng/ml) was in turn varied. 30µl of heparin was incubated with 10µl of BMP-2 (10µg/ml) for 20 minutes. 1ml of PBS was resuspended with the mixture and added to each well containing the pig osteoblasts. 4.2.4 Protein determination The amount of proteins in lysed cells was determined using BCA protein assay kit (Thermo Scientific, USA). 90µl of mixed reagent (A:B = 50:1) was 54 added to 10µl of samples in 96-well plates. This was followed by incubation at 37°C for 30 minutes. Protein concentration was determined spectrophotometrically at 595nm using multilabel counter. 4.2.5 Alkaline phosphatase activity 20µl of the lysed samples with protein concentrations of 20µg/ml were placed into each well. The positive well consists of 1µl of Calf Intestinal Phosphatase (Research Biolabs, England) and 19µl of RIPA buffer. 40µl of assay buffer (1x pNPP buffer: Invitrogen, USA) was added to each well. After incubation at 37°C for 60 minutes, samples were read at 405nm for measurement of ALP activity. 4.2.6 Western blot Lysates (12 µg of proteins) were separated on 10% (w/v) SDS– polyacrylamide gels and electroblotted onto nitrocellulose membranes. Overnight incubation with primary antibodies (p-SMAD 1/5/8, Cell Signalling, USA; Smad1/5/8, Santa Cruz, USA) was carried out at 4°C, followed by secondary antibodies for 3 h at room temperature. Membranes were then exposed to X-ray films for band detection. 4.2.7 Alizarin red staining Confluent cells on well plates were rinsed 3x with PBS. They were then fixed in 4% formaldehyde for 10 mins and rinsed 3x with DI water. Alizarin 55 red solution (13.7g/L) was added to cells and incubated for 30mins on shaker. Cells were left at room temperature to dry for 2 days. 4.2.8 Release profile studies Pig osteoblasts, with optimal concentrations of 30µl of BMP-2 and equal volumes of heparin, mixed with fibrin glue (Tisseel kit, Baxter, USA) were seeded onto scaffolds directly in 24 wells plates. After 1 hr of incubation at 37°C, scaffolds were flipped over and resuspensed to ensure complete solidifying of mixture. After the second incubation period, 2ml of PBS was added to individual wells. PBS was replaced at selected intervals: 2h, Day 1, 4, 7, 14, 21, 28 and analysed for BMP-2 and total protein release. 4.2.9 BMP-2 Release BMP-2 release was assessed using an ELISA kit (Quantikine, R&D Systems, USA). Supernatants were collected and frozen down to -20°C immediately. Prior to analysis, they are thawed and 50µl of samples were assessed for BMP2 concentration according to manufacturer instructions. 4.2.10 Statistical Analysis All values in this study were presented as mean values ± standard deviation (SD) of the mean. Data analyses and comparisons were performed using Student’s paired t-test (Microsoft Excel). P-values of < 0.05 were considered as statistical significance. 56 4.3 4.3.1 RESULTS Optimal BMP-2 concentration Figure 4.1: Alkaline phosphatase activity per mg protein of cells at different BMP-2 concentrations (* denotes p < 0.05) Alkaline phosphatase (ALP) activity was monitored as it is a standard early marker for osteogenesis. The results showed increased ALP activity in pig osteoblasts when treated with 100, 300 and 1000ng/ml of BMP-2(Figure 4.1). ALP activity was highest at 1000 ng/ml of BMP-2. However, optimal concentration was chosen to be 100 ng/ml due to high cost of BMP-2 and similar ALP activity relative to 300 ng/ml. 57 4.3.2 Western blot analysis Figure 4.2: Western blot for pig osteoblasts treated with 100ng/ml BMP-2 at different treatment times. (top - phosphorylated SMAD 1/5/8 and bottom – Total SMAD 1/5/8) In the presence of BMP-2, total SMAD 1/5/8 was converted to p-SMAD 1/5/8, according to the TGF-β signaling pathway as described in Chapter 2. Fig 4.2 indicates a higher expression for both SMAD and p-SMAD after 24 hours of BMP-2 treatment. The total SMAD expression was decreased significantly by 48 hours. 58 4.3.3 Alizarin red staining Figure 4.3: Alizarin red staining for pig osteoblasts with and without BMP-2 (control) treatment for 3 weeks. Alizarin red staining was used to detect calcium deposition in cells. BMP-2 was shown to induce mineralization when added to C2C12 cells after 7 days (Bessa et al., 2009). Slightly darker staining was observed for BMP-2 treated pig osteoblasts relative to controls. This indicates increased mineralization of osteoblasts induced by BMP-2. 4.3.4 Optimal heparin concentration Figure 4.4: Alkaline phosphatase activity per unit protein of cells at different BMP-2 and/or heparin concentrations (* denotes p < 0.05) 59 Heparin had been shown to have a potentiating and stabilizing effect on BMP-2. ALP/protein activity of pig osteoblasts treated with 100 ng/ml of BMP2 was significantly higher than untreated ones. At 300ng/ml of heparin, ALP activity was higher than at other concentrations (13.7-33.5%). Thus, it was chosen as optimal concentration for further analyses. 4.3.5 Protein release profile Figure 4.5: Amount of total protein release at various time points (Group A: PCL-TCP scaffolds loaded with PBS. Group B: PCL-TCP scaffolds loaded with 100ng/ml BMP-2. Group C: PCL-TCP scaffolds loaded with 300ng/ml of heparin. Group D: PCL-TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP-2) The general trend of protein release was a biphasic release for all groups except heparin. High level of proteins was released at 2hr (0.55 – 0.74mg/ml) and day 14 (0.52 – 0.66mg/ml). Protein release from Group C exhibited a 60 different release profile relative to other groups where the concentration was 0.54mg/ml at 2hr and decreased throughout the study. 4.3.6 BMP-2 release profile Figure 4.6: Amount of BMP-2 release at various time points. (Group A: PCLTCP scaffolds loaded with PBS. Group B: PCL-TCP scaffolds loaded with 100ng/ml BMP-2. Group C: PCL-TCP scaffolds loaded with 300ng/ml of heparin. Group D: PCL-TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP-2) Group A and C showed no release of BMP-2 at all time points. The concentration of BMP-2 release in PBS was determined using ELISA kit. Group B demonstrated a biphasic release with heightened release at 2hr and day 14. However, Group D did not exhibit a significant BMP-2 release (0 - 25.6pg/ml) throughout the study. 61 4.3.7 ALP bioactivity of eluted BMP-2 Figure 4.7: Bioactivity of eluted BMP-2 at different time points (Group A: PCLTCP scaffolds loaded with PBS. Group B: PCL-TCP scaffolds loaded with 100ng/ml BMP-2. Group C: PCL-TCP scaffolds loaded with 300ng/ml of heparin. Group D: PCL-TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP-2) Generally, there were no significant differences between all groups at various time points. Group B and D exhibited higher ALP activity at Day 1 and 4 compared to other groups (25.5 – 66.5%). ALP activity range of Group A (7.8 – 25.4/mg protein) was similar to that of other groups (4.9 - 26.3/mg protein). 4.4 DISCUSSIONS As discussed previously, there are numerous studies on the delivery of BMP-2 using a 3D matrix (Hosseinkhani et al., 2007; Kempen et al., 2008; Ruhé 62 et al., 2005). Problems like initial bursts upon implantation and bone growth away from defect sites indicate that BMP-2 delivery is far from ideal. Our group proposes a new co-delivery method of using heparin to effectively deliver BMP-2 to the defect sites. Heparin has shown to potentiate, sustain and stimulate BMP-2 through binding to its N-terminal (Zhao et al., 2006). Here, BMP-2 and heparin were seeded onto biodegradable PCL-TCP scaffolds using Tisseel fibrin sealant. This study focused on the effects of heparin and BMP-2 on pig osteoblasts and their release profile over 4 weeks when seeded in PCL-TCP scaffolds using fibrin sealant. Pig osteoblasts obtained from mandible explants were used in this experiment. This is in line with our next step to implant PCL-TCP scaffolds loaded with heparin and BMP-2 into pig mandible defects. Thus, this will paint a more accurate picture on the effects of heparin and BMP-2 on pig osteoblasts upon implantation. The bioactivity of BMP-2 can be verified by the responsiveness of pig osteoblasts through ALP activity. It has been used widely as a biochemical marker for osteoblast differentiation (Cremers et al., 2008; Katagiri et al., 1994). Pig osteoblasts displayed heightened ALP activity when subjected to 100, 300 and 1000ng/ml of BMP-2 (Figure 4.1). Increased ALP activity on C2C12 cells in demineralised bone matrix had been proven by (Lin et al.). Similar effects have also been reported in another study (Zhao et al., 2006). Highest ALP activity was obtained at 1000ng/ml. However, 63 100ng/ml was chosen primarily because of high cost of BMP-2 involved. Another reason is that ALP activity of 300ng/ml seemed to have no significant difference compared to 100ng/ml. Western blot and alizarin red staining were then performed to verify translated protein expression and mineralisation of pig osteoblasts by BMP-2 treatment. According to cell signalling pathway, total SMADs are present in all cells and they will be converted to p-SMAD in the presence of BMP-2. A higher expression of p-SMAD is detected at 24h compared to 2 and 48h. This is in agreement with the cell signalling pathway. This was affirmed by previous findings by our group when BMP-2 treatment times on C2C12 cells were varied (data not shown). However in Figure 4.2, total SMAD were not fully expressed at 2 and 48h. This deviation from signalling pathway might be due to technical error during blotting or reduced bioactivity of primary antibodies. Repeated freeze thaw activity and prolonged period of storage may result in instability of its properties. Mineralisation of treated cells at 3 weeks was detected by alizarin red staining. BMP-2 treated cells displayed a darker shade of red which is in agreement with western blot findings. BMP-2 used in this study resulted in increased osteoblast differentiation, induced p-SMAD expression and mineralisation in pig osteoblasts. In Figure 4.4, optimal concentration ratio of BMP-2 to heparin was found to be 1:3. Similar study conducted on C2C12 cells with ratio of 1:5 64 demonstrated that heparin enhance osteoblast differentiation. Concentration obtained here indicates that heparin has a positive effect on BMP-2 activity. More testings on animal studies is needed to confirm the ratio. Method of introducing heparin to BMP-2 may influence their interactions, resulting in varied effects in bone regeneration. Direct addition of heparin to BMP-2 (Zhao et al., 2006), crosslinking of heparin to demineralised bone matrix (Lin et al., 2008) and conjugation of heparin to PLGA scaffolds (Jeon et al., 2007) were some of the methods that was previously used and with good results. Here, BMP-2 was allowed to react with heparin first during incubation before mixing with medium and applying to cells. This allows sufficient time for N terminal bonds to occur and prevent other reactions that may take place preferentially. It can also simulate conditions whereby heparin is incorporated in PCL-TCP scaffolds. Fibrin sealant is derived from blood clots and has been used effectively for healing of critical sized bone defects and non-unions in cats, dogs and rats (Schmoekel et al., 2005). It has also been established as an effective method to sustain BMP-2 diffusion. This is verified when fibrin is used in collagen sponge for BMP-2 delivery in a rat spinal model. It has shown successfully to control bone growth in undesirable areas such as nerve tissues (Patel et al., 2006). Therefore, fibrin sealant was selected as delivery system for heparin and BMP-2. 65 Protein release in PBS can be attributed to fibrin and BMP-2 eluted from scaffolds. Fibrin release is dominant here as concentration of BMP-2 is relatively lower at 100ng/ml (Figure 4.6). Biphasic protein release was evident in Group A, B and D. Profile was similar to BMP-2 released for Group B and D. Initial burst at 2h and Day 1 was mainly due to release of fibrin sealant and uncrosslinked fibrin on the scaffold surface. Between day 4 to 7, BMP-2 was locked in PCL-TCP scaffolds by fibrin sealant. At this time, scaffolds were bound to the surface of well plate. Subsequent release of BMP-2 at Day 14 can be attributed to the disintegration of fibrin within scaffolds whereby BMP-2 was released concurrently (Figure 4.6). Fibrin loaded scaffolds were also observed to dislodge from the well plates. Release of BMP-2 after Day 14 can be attributed to the release of TCP which has the ability to form intermolecular bonds with BMP-2. Previous studies confirmed the findings (Rai et al., 2005b; Wei et al., 2004). There was a huge difference between Group C release profile relative to the rest. Heparin seemed to react with fibrin sealant at the start of experiment and thus resulting in a constant decrease in protein throughout the experiment. Performance of fibrin in the presence of heparin has been elucidated by Marx et al. Heparin did not have any effect on clotting time but bonding strength was decreased by 20% (Marx and Mou, 2002). Here, Group C’s protein release profile exhibited an initial burst and declined thereafter. It 66 is well known that heparin prevents the formation of fibrin clot by a series of coagulation cascade activities. Even though different reaction has occurred when heparin reacts with fibrin sealant causing degradation profile to differ, clot formation and mechanical properties still remain intact. More in depth characterisations are needed to determine the cause of this phenomenon. ALP activity of the eluted BMP-2 did not exhibit any significant differences from controls. This is to be expected from the low concentration of BMP-2 eluted (Figure 4.6). Any concentration that is lower than 10ng/ml as depicted in Figure 4.1 will have negligible effects relative to controls. 4.5 CONCLUSIONS Heparin and BMP-2 has been investigated previously as supplementary factor in the realm of bone regeneration. Here, heparin incubated with BMP-2 demonstrated higher ALP activity in pig osteoblasts than BMP-2 alone. The bioactivity of BMP-2 was also verified at different time points. Interestingly, BMP-2 in the presence of heparin when loaded on PCL-TCP scaffolds did not showed a sustained release as hypothesized. More analysis is needed to confirm the hypothesis. Further studies should focus on the controlled release of BMP-2 and heparin loaded PCL-TCP scaffolds in an in vivo model. 67 CHAPTER 5: FINAL RECOMMENDATIONS 5.1 Effects of porosities of PCL-TCP scaffolds on in vivo bone regeneration. This particular chapter has its emphasis on two issues namely surface modification and porosity of acellular PCL-TCP scaffolds where the latter is the focus here. Thickness to pore size ratio of scaffolds was not studied in this study and can be an interesting research area to look at. A micro pig model will be appropriate for this study whereby thickness of the scaffolds can be varied. Critical sized circular defect can be used again which allow for more accurate determination of region of interest. Micro CT, histology and mechanical testing should be conducted at 3 and 6 months to determine extent of bone formation, cellular interactions and mechanical properties can be evaluated. Mechanical properties can provide us with information on structural integrity of specimens at that particular point. Similarly, degradation profile of scaffolds can be mapped using data from Micro CT. This will be more representative as it is conducted on a larger animal model. 5.2 Preliminary in vitro evaluation of PCL-TCP scaffolds as co-delivery systems for heparin and BMP-2 Optimal concentrations and ratio of BMP-2 and heparin have been determined in Chapter 4. Release of BMP-2 is inconclusive which is mainly due to low concentration of BMP-2 used. Separate release study with a higher 68 concentration of BMP-2 and heparin whist maintaining the optimal ratio (1:3) can be conducted. In fact, heparin release using alexa fluro has been conducted concurrently in this study. Even though lowest detectable concentration stood at 24ng/ml, no heparin was detected at 485nm. There is a possible reaction that occurred between heparin and fibrin sealant as described in Chapter 4. Heparin can also be radiolabelled and release can be tracked using suitable counter. Next step can be to incorporate pig osteoblast into PCL-TCP scaffolds with optimal ratio of heparin and BMP-2. At various time points, characterisation like ALP, BCA and picogreen can be conducted to determine their extent of differentiation, protein concentration and DNA content. Alizarin red staining will also be able to determine the extent of characterisation. Viability of cells on surface and within scaffolds can also be determined using Live/dead stain. 69 BIBLIOGRAPHY A.S.Htay, S.H.T.a.D.W.H., Develpoment of perforated microthin poly(caprolactone) films as matrices for membrane tissue engineering.: J. Biamater. Sci. Polymer Edn, v. 15, p. 683-700. 2004 Andrew, J.G., Hoyland, J.A., Freemont, A.J., and Marsh, D.R., Platelet-derived growth factor expression in normally healing human fractures: Bone, v. 16, p. 455-60. 1995 Aronin, C.E.P., Sadik, K.W., Lay, A.L., Rion, D.B., Tholpady, S.S., Ogle, R.C., and Botchwey, E.A., Comparative effects of scaffold pore size, pore volume, and total void volume on cranial bone healing patterns using microspherebased scaffolds: Journal of Biomedical Materials Research Part A, v. 89A, p. 632-641. 2009 Ayers, R.A., Simske, S.J., Bateman, T.A., Petkus, A., Sachdeva, R.L., and Gyunter, V.E., Effect of nitinol implant porosity on cranial bone ingrowth and apposition after 6 weeks: J Biomed Mater Res, v. 45, p. 42-7. 1999 Bae, J., Desai, U.R., Pervin, A., Caldwell, E.E., Weiler, J.M., and Linhardt, R.J., Interaction of heparin with synthetic antithrombin III peptide analogues: Biochem J, v. 301 ( Pt 1), p. 121-9. 1994 Beruto, D.T., Mezzasalma, S.A., Capurro, M., Botter, R., and Cirillo, P., Use of alpha-tricalcium phosphate (TCP) as powders and as an aqueous dispersion to modify processing, microstructure, and mechanical properties of polymethylmethacrylate (PMMA) bone cements and to produce bonesubstitute compounds: Journal of Biomedical Materials Research, v. 49, p. 498505. 2000 Bessa, P.C., Casal, M., and Reis, R.L., Bone morphogenetic proteins in tissue engineering: the road from laboratory to clinic, part II (BMP delivery): Journal of Tissue Engineering and Regenerative Medicine, v. 2, p. 81-96. 2008a —, Bone morphogenetic proteins in tissue engineering: the road from the laboratory to the clinic, part I (basic concepts): Journal of Tissue Engineering and Regenerative Medicine, v. 2, p. 1-13. 2008b Bessa, P.C.s., Balmayor, E.R., Hartinger, J., Zanoni, G., Dopler, D., Meinl, A., Banerjee, A., Casal, M., Redl, H., Reis, R.L., and van Griensven, M., Silk fibroin microparticles as carriers for delivery of human recombinant BMP-2: Tissue Engineering Part C: Methods, v. 0. 2009 70 Biomedical Tissue Research Group, Bone Structure, Volume 2009, University of York. 1996 Biomedical Tissue Research Group, The bone remodelling process, Volume 2009, Department of Biology, University of York. 2007 BJD, The Initiative, Bone and Joint Decade Musculoskeletal portal, Volume 2009. 2000-2010 Boyan BD, H.T., KieswetterK, Schraub D, dean DD, Schwartz Z., Effect of titanium surface characteristics on chondrocytes and osteoblasts in vitro: Cells Mater., v. 5, p. 323-335. 1995 Brett PM, H.J., Salih V, Mihoc R, Olsen I, Jones FH, Tonetti M',, Roughness response genes in osteoblasts.: Bone, v. 35, p. 124-33. 2004 C.G Pitt, M.M.G., G.L. Kimmel, J. Surles and A. Schindler, Aliphatic polyesters 2. The degradation of poly(DL-lactide), poly(caprolactone), and their copolymers in vivo: Biomaterials, v. 2, p. 215-220. 1981 Callis, T.E., Cao, D., and Wang, D.Z., Bone morphogenetic protein signaling modulates myocardin transactivation of cardiac genes: Circ Res, v. 97, p. 9921000. 2005 CambridgeSoft Corporation, Calcium Phosphate Tribasic ChemFinder.com Database and internet searching. 2004 [7758-87-4] Canalis, E., McCarthy, T.L., and Centrella, M., Effects of platelet-derived growth factor on bone formation in vitro: J Cell Physiol, v. 140, p. 530-7. 1989 Casey K C, T.S.K., S Liao, R Murugan, M Nigiam, S Ramakrishnan, Biomimetic nanocomposities for bone graft applications: Nanomedicine, v. 1, p. 177-188. 2006 Caughey, G.H., Building a better heparin: Am J Respir Cell Mol Biol, v. 28, p. 129-32. 2003 Chen, R.R., and Mooney, D.J., Polymeric growth factor delivery strategies for tissue engineering: Pharm Res, v. 20, p. 1103-12. 2003 71 Christopher XF Lam, S.H.T., Dietmar W Hutmacher, Comparison of the degradation of polycaprolactone and polycaprolactone-(tricalcium phosphate) scaffolds in alkaline solution: Polym Int, v. 56, p. 718-728. 2007 Chung, Y.I., Ahn, K.M., Jeon, S.H., Lee, S.Y., Lee, J.H., and Tae, G., Enhanced bone regeneration with BMP-2 loaded functional nanoparticle-hydrogel complex: J Control Release, v. 121, p. 91-9. 2007 Cremers, S., Garnero, P., and Seibel, M.J., Biochemical Markers of Bone Metabolism, in John, P.B., Lawrence, G.R., and Martin, T.J., eds., Principles of Bone Biology (Third Edition): San Diego, Academic Press, p. 1857-1881. 2008 Drosos, G.I., Kazakos, K.I., Kouzoumpasis, P., and Verettas, D.-A., Safety and efficacy of commercially available demineralised bone matrix preparations: A critical review of clinical studies: Injury, v. 38, p. S13-S21. 2007 Eggli, P.S., Muller, W., and Schenk, R.K., Porous hydroxyapatite and tricalcium phosphate cylinders with two different pore size ranges implanted in the cancellous bone of rabbits. A comparative histomorphometric and histologic study of bony ingrowth and implant substitution: Clin Orthop Relat Res, p. 127-38. 1988 El-Ghannam, A.R., Advanced bioceramic composite for bone tissue engineering: design principles and structure-bioactivity relationship: J Biomed Mater Res A, v. 69, p. 490-501. 2004 Enneking, W.F., Eady, J.L., and Burchardt, H., Autogenous cortical bone grafts in the reconstruction of segmental skeletal defects: J Bone Joint Surg Am, v. 62, p. 1039-58. 1980 Fisher, J.P., Vehof, J.W.M., Dean, D., Waerden, J.P.C.M.v.d., Holland, T.A., Mikos, A.G., and Jansen, J.A., Soft and hard tissue response to photocrosslinked poly(propylene fumarate) scaffolds in a rabbit model: Journal of Biomedical Materials Research, v. 59, p. 547-556. 2002 Flautre, B., Descamps, M., Delecourt, C., Blary, M.C., and Hardouin, P., Porous HA ceramic for bone replacement: role of the pores and interconnections experimental study in the rabbit: J Mater Sci Mater Med, v. 12, p. 679-82. 2001 Friederike von Burkersrodaa, L.S., Achim G.opferich, Why degradable polymers undergo surface erosion or bulk erosion: Biomaterials, v. 23, p. 42214231. 2002 72 Fromm, J.R., Hileman, R.E., Caldwell, E.E., Weiler, J.M., and Linhardt, R.J., Pattern and spacing of basic amino acids in heparin binding sites: Arch Biochem Biophys, v. 343, p. 92-100. 1997 Gitelis, S., and Saiz, P., What's new in orthopaedic surgery: J Am Coll Surg, v. 194, p. 788-91. 2002 Gokhale, J.A., Boskey, A.L., and Robey, P.G., The Biochemistry of Bone, in Robert, M., David, F., and Jennifer, K., eds., Osteoporosis (Second Edition): San Diego, Academic Press, p. 107-188. 2001 Gotz, H.E., Muller, M., Emmel, A., Holzwarth, U., Erben, R.G., and Stangl, R., Effect of surface finish on the osseointegration of laser-treated titanium alloy implants: Biomaterials, v. 25, p. 4057-64. 2004 Gray, E., Mulloy, B., and Barrowcliffe, T.W., Heparin and low-molecularweight heparin: Thromb Haemost, v. 99, p. 807-18. 2008 Hak-Joon Sung, C.M., Chad Johnson and Zorina S. Galis, The effect of scaffold degradation rate on three-dimensional cell growth and angiogenesis: Biomaterials, v. 25, p. 5735-5742. 2004 HCUP, Healthcare Cost and Utilization Project: Agency for Healthcare Research and Quality. 2006 Heo, S.-J., Kim, S.-E., Wei, J., Kim, D.H., Hyun, Y.-T., Yun, H.-S., Kim, H.K., Yoon, T.R., Kim, S.-H., Park, S.-A., Shin, J.W., and Shin, J.-W., In Vitro and Animal Study of Novel Nano-Hydroxyapatite/Poly(ɛ-Caprolactone) Composite Scaffolds Fabricated by Layer Manufacturing Process: Tissue Engineering Part A, v. 15, p. 977-989. 2009 Hosseinkhani, H., Hosseinkhani, M., Khademhosseini, A., and Kobayashi, H., Bone regeneration through controlled release of bone morphogenetic protein-2 from 3-D tissue engineered nano-scaffold: Journal of Controlled Release, v. 117, p. 380-386. 2007 Hurley, M.M., Abreu, C., Harrison, J.R., Lichtler, A.C., Raisz, L.G., and Kream, B.E., Basic fibroblast growth factor inhibits type I collagen gene expression in osteoblastic MC3T3-E1 cells: J Biol Chem, v. 268, p. 5588-93. 1993 Hutmacher D.W, S.T., Zein I, Ng K.W, Teoh S.H, Tan K.C., Mechanical properties and cell culture response of polycaprolactone scaffolds designed 73 and fabricated via fused deposition modeling: J. Biomed Mater Res, v. 55, p. 203-16. 2001 Hutmacher, D.W., Schantz, T., Zein, I., Ng, K.W., Teoh, S.H., and Tan, K.C., Mechanical properties and cell cultural response of polycaprolactone scaffolds designed and fabricated via fused deposition modeling: Journal of Biomedical Materials Research, v. 55, p. 203-216. 2001 Ikeda, R., Fujioka, H., Nagura, I., Kokubu, T., Toyokawa, N., Inui, A., Makino, T., Kaneko, H., Doita, M., and Kurosaka, M., The effect of porosity and mechanical property of a synthetic polymer scaffold on repair of osteochondral defects: International Orthopaedics, v. 33, p. 821-828. 2009 James M Pachence, J.K., Biodegradable polymers, Principles of tissue engineering, Academic press. 2000 Jeon, O., Song, S.J., Kang, S.W., Putnam, A.J., and Kim, B.S., Enhancement of ectopic bone formation by bone morphogenetic protein-2 released from a heparin-conjugated poly(L-lactic-co-glycolic acid) scaffold: Biomaterials, v. 28, p. 2763-71. 2007 Johnson, E.E., Urist, M.R., and Finerman, G.A., Repair of segmental defects of the tibia with cancellous bone grafts augmented with human bone morphogenetic protein. A preliminary report: Clin Orthop Relat Res, p. 249-57. 1988 Jones, A.L., Bucholz, R.W., Bosse, M.J., Mirza, S.K., Lyon, T.R., Webb, L.X., Pollak, A.N., Golden, J.D., and Valentin-Opran, A., Recombinant human BMP2 and allograft compared with autogenous bone graft for reconstruction of diaphyseal tibial fractures with cortical defects. A randomized, controlled trial: J Bone Joint Surg Am, v. 88, p. 1431-41. 2006 Jones, J.R., Scaffolds for tissue engineering, Biomaterials, artificial organs and tissue engineering: Imperial College London, UK, Woodward Publishing Limited, p. 201-214. 2005 Juan Pena, T.C., Isabel Izquierdo-Barba, Antonio L. Doadrio and Maria ValletRegi, Long term degradation of poly([var epsilon]-caprolactone) films in biologically related fluids: Polymer Degradation and Stability, v. 91, p. 14241432. 2006 74 K A Hing, S.M.B., K E Tanner, P A Revell W Bonfield, Histomorphological and biomechanical characterization of calcium phosphates in the osseous environment: Proc Istn Mech Engrs, v. 212H, p. 437-451. 1998 Karageorgiou, V., and Kaplan, D., Porosity of 3D biomaterial scaffolds and osteogenesis: Biomaterials, v. 26, p. 5474-91. 2005 Kasten, P., Beyen, I., Niemeyer, P., Luginbühl, R., Bohner, M., and Richter, W., Porosity and pore size of [beta]-tricalcium phosphate scaffold can influence protein production and osteogenic differentiation of human mesenchymal stem cells: An in vitro and in vivo study: Acta Biomaterialia, v. 4, p. 1904-1915. 2008 Katagiri, T., Yamaguchi, A., Komaki, M., Abe, E., Takahashi, N., Ikeda, T., Rosen, V., Wozney, J., Fujisawa-Sehara, A., and Suda, T., Bone morphogenetic protein-2 converts the differentiation pathway of C2C12 myoblasts into the osteoblast lineage [published erratum appears in J Cell Biol 1995 Feb;128(4):following 713]: J. Cell Biol., v. 127, p. 1755-1766. 1994 Kay, J.F., Tissue-engineered bone products, in Robert, L., Robert, L., and Joseph, V., eds., Principles of Tissue Engineering (Third Edition): Burlington, Academic Press, p. 1225-1236. 2007 Kempen, D.H., Lu, L., Hefferan, T.E., Creemers, L.B., Maran, A., Classic, K.L., Dhert, W.J., and Yaszemski, M.J., Retention of in vitro and in vivo BMP-2 bioactivities in sustained delivery vehicles for bone tissue engineering: Biomaterials, v. 29, p. 3245-52. 2008 Khan, S.N., Bostrom, M.P., and Lane, J.M., Bone growth factors: Orthop Clin North Am, v. 31, p. 375-88. 2000 Kuboki, Y., Jin, Q., and Takita, H., Geometry of carriers controlling phenotypic expression in BMP-induced osteogenesis and chondrogenesis: J Bone Joint Surg Am, v. 83-A Suppl 1, p. S105-15. 2001 Kuboki, Y., Takita, H., Kobayashi, D., Tsuruga, E., Inoue, M., Murata, M., Nagai, N., Dohi, Y., and Ohgushi, H., BMP-induced osteogenesis on the surface of hydroxyapatite with geometrically feasible and nonfeasible structures: topology of osteogenesis: J Biomed Mater Res, v. 39, p. 190-9. 1998 Kuhne, J.H., Bartl, R., Frisch, B., Hammer, C., Jansson, V., and Zimmer, M., Bone formation in coralline hydroxyapatite. Effects of pore size studied in rabbits: Acta Orthop Scand, v. 65, p. 246-52. 1994 75 Lakes, J.P.a.R.S., Biomaterials An Introduction, Springer Science+Business Media LLC. 2007 Lam, C.X., Hutmacher, D.W., Schantz, J.T., Woodruff, M.A., and Teoh, S.H., Evaluation of polycaprolactone scaffold degradation for 6 months in vitro and in vivo: J Biomed Mater Res A, v. 90, p. 906-19. 2009 Lam, C.X., Teoh, S.H., and Hutmacher, D.W., Comparison of the degradation of polycaprolactone and polycaprolactone-(beta-tricalcium phosphate) scaffolds in alkaline medium: Polymer International, v. 56, p. 718-728. 2007 Langer, R., and Vacanti, J.P., Tissue engineering: Science, v. 260, p. 920-6. 1993 Lever, R., and Page, C.P., Novel drug development opportunities for heparin: Nat Rev Drug Discov, v. 1, p. 140-8. 2002 Lian J , S.G., Canalis E , Gehron Robey P , and Boskey A, Bone formation: Osteoblast lineage cells, growth factors, matrix proteins and the mineralisation process . in Favus M ed., Primer on the metabolic bone diseases and disorders of mineral metabolism, Volume pp. 14 – 29: Philadelphia Lippincott William & Wilkins. 1999 Lickorish, D., Ramshaw, J.A., Werkmeister, J.A., Glattauer, V., and Howlett, C.R., Collagen-hydroxyapatite composite prepared by biomimetic process: J Biomed Mater Res A, v. 68, p. 19-27. 2004 Lieberman, J.R., Daluiski, A., and Einhorn, T.A., The role of growth factors in the repair of bone. Biology and clinical applications: J Bone Joint Surg Am, v. 84-A, p. 1032-44. 2002 Lin, H., Zhao, Y., Sun, W., Chen, B., Zhang, J., Zhao, W., Xiao, Z., and Dai, J., The effect of crosslinking heparin to demineralized bone matrix on mechanical strength and specific binding to human bone morphogenetic protein-2: Biomaterials, v. 29, p. 1189-97. 2008 Marx, G., and Mou, X., Characterizing fibrin glue performance as modulated by heparin, aprotinin, and factor XIII: Journal of Laboratory and Clinical Medicine, v. 140, p. 152-160. 2002 MOH, Ministry of Health: Government Health Expenditure. 2008 76 Mont, M.A., Ragland, P.S., Biggins, B., Friedlaender, G., Patel, T., Cook, S., Etienne, G., Shimmin, A., Kildey, R., Rueger, D.C., and Einhorn, T.A., Use of Bone Morphogenetic Proteins for Musculoskeletal Applications: An Overview: J Bone Joint Surg Am, v. 86, p. 41-55. 2004 Müller, M., Hennig, F.F., Hothorn, T., and Stangl, R., Bone-implant interface shear modulus and ultimate stress in a transcortical rabbit model of open-pore Ti6Al4V implants: Journal of Biomechanics, v. 39, p. 2123-2132. 2006 NHED, National Health Expenditure Data: Centers for Medicare and Medicaid Services U.S. Department of Health & Human Services. 2006 Nillesen, S.T., Geutjes, P.J., Wismans, R., Schalkwijk, J., Daamen, W.F., and van Kuppevelt, T.H., Increased angiogenesis and blood vessel maturation in acellular collagen-heparin scaffolds containing both FGF2 and VEGF: Biomaterials, v. 28, p. 1123-31. 2007 Paramore, C.G., Lauryssen, C., Rauzzino, M.J., Wadlington, V.R., Palmer, C.A., Brix, A., Cartner, S.C., and Hadley, M.N., The safety of OP-1 for lumbar fusion with decompression-- a canine study: Neurosurgery, v. 44, p. 1151-5; discussion 1155-6. 1999 Patel, V.V., Zhao, L., Wong, P., Kanim, L., Bae, H.W., Pradhan, B.B., and Delamarter, R.B., Controlling bone morphogenetic protein diffusion and bone morphogenetic protein-stimulated bone growth using fibrin glue: Spine (Phila Pa 1976), v. 31, p. 1201-6. 2006 Pimentel, E., Handbook of Growth Factors I: General Basic Aspects: Boca Raton, Florida, CRC Press. 1994 Price RL, E.K., Haberstroh KM, Webster TJ, Nanometer surface roughness increases select osteoblast adhesion on carbon nanofiber compacts.: J Biomed Mater Res A, v. 70, p. 129-38. 2004 Rai, B., Ho, K.H., Lei, Y., Si-Hoe, K.-M., Jeremy Teo, C.-M., Yacob, K.b., Chen, F., Ng, F.-C., and Teoh, S.H., Polycaprolactone-20% Tricalcium Phosphate Scaffolds in Combination With Platelet-Rich Plasma for the Treatment of Critical-Sized Defects of the Mandible: A Pilot Study: Journal of Oral and Maxillofacial Surgery, v. 65, p. 2195-2205. 2007 Rai, B., Teoh, S.H., and Ho, K.H., An in vitro evaluation of PCL-TCP composites as delivery systems for platelet-rich plasma: Journal of Controlled Release, v. 107, p. 330-342. 2005a 77 Rai, B., Teoh, S.H., Ho, K.H., Hutmacher, D.W., Cao, T., Chen, F., and Yacob, K., The effect of rhBMP-2 on canine osteoblasts seeded onto 3D bioactive polycaprolactone scaffolds: Biomaterials, v. 25, p. 5499-5506. 2004 Rai, B., Teoh, S.H., Hutmacher, D.W., Cao, T., and Ho, K.H., Novel PCL-based honeycomb scaffolds as drug delivery systems for rhBMP-2: Biomaterials, v. 26, p. 3739-3748. 2005b Raisz, L.G., and Seeman, E., Causes of age-related bone loss and bone fragility: an alternative view: J Bone Miner Res, v. 16, p. 1948-52. 2001 Rezwan K, C.Q.Z., Blaker J.J, Boccaccini A.R, Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering: Biomaterials, v. 27, p. 3413-31. 2006 Rifkin, D.B., and Moscatelli, D., Recent developments in the cell biology of basic fibroblast growth factor: J Cell Biol, v. 109, p. 1-6. 1989 Roosa, S.M., Kemppainen, J.M., Moffitt, E.N., Krebsbach, P.H., and Hollister, S.J., The pore size of polycaprolactone scaffolds has limited influence on bone regeneration in an in vivo model: J Biomed Mater Res A. 2009 Roy, T.D., Simon, J.L., Ricci, J.L., Rekow, E.D., Thompson, V.P., and Parsons, J.R., Performance of degradable composite bone repair products made via three-dimensional fabrication techniques: J Biomed Mater Res A, v. 66, p. 28391. 2003 Ruhé, P.Q., Boerman, O.C., Russel, F.G.M., Spauwen, P.H.M., Mikos, A.G., and Jansen, J.A., Controlled release of rhBMP-2 loaded poly(dl-lactic-co-glycolic acid)/calcium phosphate cement composites in vivo: Journal of Controlled Release, v. 106, p. 162-171. 2005 Ruppert, R., Hoffmann, E., and Sebald, W., Human bone morphogenetic protein 2 contains a heparin-binding site which modifies its biological activity: Eur J Biochem, v. 237, p. 295-302. 1996 Sakou, T., Bone morphogenetic proteins: from basic studies to clinical approaches: Bone, v. 22, p. 591-603. 1998 Salyer, K.E., Bardach, J., Squier, C.A., Gendler, E., and Kelly, K.M., Cranioplasty in the growing canine skull using demineralized perforated bone: Plast Reconstr Surg, v. 96, p. 770-9. 1995 78 Sawyer, A.A., Song, S.J., Susanto, E., Chuan, P., Lam, C.X.F., Woodruff, M.A., Hutmacher, D.W., and Cool, S.M., The stimulation of healing within a rat calvarial defect by mPCL-TCP/collagen scaffolds loaded with rhBMP-2: Biomaterials, v. 30, p. 2479-2488. 2009 Schliephake, H., Neukam, F.W., and Klosa, D., Influence of pore dimensions on bone ingrowth into porous hydroxylapatite blocks used as bone graft substitutes. A histometric study: Int J Oral Maxillofac Surg, v. 20, p. 53-8. 1991 Schmoekel, H.G., Weber, F.E., Schense, J.C., Gratz, K.W., Schawalder, P., and Hubbell, J.A., Bone repair with a form of BMP-2 engineered for incorporation into fibrin cell ingrowth matrices: Biotechnol Bioeng, v. 89, p. 253-62. 2005 Senn, On the Healing of Aseptic Bone Cavities by Implantation of Antiseptic Decalcified Bone: Ann Surg, v. 10, p. 352-68. 1889 Simic, P., and Vukicevic, S., Bone morphogenetic proteins in development and homeostasis of kidney: Cytokine Growth Factor Rev, v. 16, p. 299-308. 2005 Steffens, G.C., Yao, C., Prevel, P., Markowicz, M., Schenck, P., Noah, E.M., and Pallua, N., Modulation of angiogenic potential of collagen matrices by covalent incorporation of heparin and loading with vascular endothelial growth factor: Tissue Eng, v. 10, p. 1502-9. 2004 Stevens, B., Yang, Y., Mohandas, A., Stucker, B., and Nguyen, K.T., A review of materials, fabrication methods, and strategies used to enhance bone regeneration in engineered bone tissues: Journal of Biomedical Materials Research Part B: Applied Biomaterials, v. 85B, p. 573-582. 2008 Sundelacruz, S., and Kaplan, D.L., Stem cell- and scaffold-based tissue engineering approaches to osteochondral regenerative medicine: Seminars in Cell & Developmental Biology, v. 20, p. 646-655. 2009 Tabata, Y., Significance of release technology in tissue engineering: Drug Discov Today, v. 10, p. 1639-46. 2005 Takahashi, Y., and Tabata, Y., Effect of the fiber diameter and porosity of nonwoven PET fabrics on the osteogenic differentiation of mesenchymal stem cells: Journal of Biomaterials Science, Polymer Edition, v. 15, p. 41-57. 2004 Temenoff, J.S., and Mikos, A.G., Injectable biodegradable materials for orthopedic tissue engineering: Biomaterials, v. 21, p. 2405-12. 2000 79 Tsuruga, E., Takita, H., Itoh, H., Wakisaka, Y., and Kuboki, Y., Pore size of porous hydroxyapatite as the cell-substratum controls BMP-induced osteogenesis: J Biochem, v. 121, p. 317-24. 1997 Urist, M.R., Bone:formation by autoinduction: Science, v. 150. 1965 Valentin-Opran, A., Wozney, J., Csimma, C., Lilly, L., and Riedel, G.E., Clinical evaluation of recombinant human bone morphogenetic protein-2: Clin Orthop Relat Res, p. 110-20. 2002 Vert, S.L.a.M., Biodegradation of aliphatic polyesters, Degradable Polymers Principles and Applications, Kluwer Academic Publishers. 2002 Vukicevic, S., and Sampath, K., Introduction, p. 1-5. 2008 Walker, D.H., and Wright, N.M., Bone morphogenetic proteins and spinal fusion: Neurosurg Focus, v. 13, p. e3. 2002 Wei, G., Pettway, G.J., McCauley, L.K., and Ma, P.X., The release profiles and bioactivity of parathyroid hormone from poly(lactic-co-glycolic acid) microspheres: Biomaterials, v. 25, p. 345-52. 2004 Wikipedia, t.f.e., Polycaprolactone, http://en.wikipedia.org/wiki/Polycaprolactone, Wikimedia Foundation Incorporated. 2006 Wissink, M.J., Beernink, R., Scharenborg, N.M., Poot, A.A., Engbers, G.H., Beugeling, T., van Aken, W.G., and Feijen, J., Endothelial cell seeding of (heparinized) collagen matrices: effects of bFGF pre-loading on proliferation (after low density seeding) and pro-coagulant factors: J Control Release, v. 67, p. 141-55. 2000 Xiao, Y.T., Xiang, L.X., and Shao, J.Z., Bone morphogenetic protein: Biochem Biophys Res Commun, v. 362, p. 550-3. 2007 Y. Lei, B.R., K.H. Ho and S.H. Teoh, In vitro degradation of novel bioactive polycaprolactone--20% tricalcium phosphate composite scaffolds for bone engineering: Materials Science and Engineering, v. 27, p. 293-298. 2007 Yamamoto, M., Takahashi, Y., and Tabata, Y., Controlled release by biodegradable hydrogels enhances the ectopic bone formation of bone morphogenetic protein: Biomaterials, v. 24, p. 4375-83. 2003 80 Yancopoulos, G.D., Davis, S., Gale, N.W., Rudge, J.S., Wiegand, S.J., and Holash, J., Vascular-specific growth factors and blood vessel formation: Nature, v. 407, p. 242-8. 2000 Yeo, A., Rai, B., Sju, E., Cheong, J.J., and Teoh, S.H., The degradation profile of novel, bioresorbable PCL-TCP scaffolds: An in vitro and in vivo study: Journal of Biomedical Materials Research Part A, v. 84A, p. 208-218. 2008a Yeo, A., Sju, E., Rai, B., and Teoh, S.H., Customizing the degradation and loadbearing profile of 3D polycaprolactone-tricalcium phosphate scaffolds under enzymatic and hydrolytic conditions: Journal of Biomedical Materials Research Part B: Applied Biomaterials, v. 87B, p. 562-569. 2008b Yeo, A., Wong, W.J., Khoo, H.H., and Teoh, S.H., Surface modification of PCLTCP scaffolds improve interfacial mechanical interlock and enhance early bone formation: An in vitro and in vivo characterization: Journal of Biomedical Materials Research Part A, v. 92A, p. 311-321. 2010 Yeo, A., Wong, W.J., and Teoh, S.H., Surface modification of PCL-TCP scaffolds in rabbit calvaria defects: Evaluation of scaffold degradation profile, biomechanical properties and bone healing patterns Journal of Biomedical Materials Research Part A, v. Accepted July 2009. 2009b Zein I, H.D., Tan KC and Teoh SH, Fused deposition modeling of novel scaffold architeures for tissue engineering applications: Biomaterials, v. 23, p. 1169-1185. 2002 Zhang, Y., and Zhang, M., Three-dimensional macroporous calcium phosphate bioceramics with nested chitosan sponges for load-bearing bone implants: J Biomed Mater Res, v. 61, p. 1-8. 2002 Zhao, B., Katagiri, T., Toyoda, H., Takada, T., Yanai, T., Fukuda, T., Chung, U.I., Koike, T., Takaoka, K., and Kamijo, R., Heparin potentiates the in vivo ectopic bone formation induced by bone morphogenetic protein-2: J Biol Chem, v. 281, p. 23246-53. 2006 81 APPENDIX (PUBLICATIONS) 82 [...]... current bone grafts as discussed propel researchers to look into the field of bone tissue engineering for an ideal bone substitute 1.1.3 Strategies in BTE Tissue engineering is the restoration, improvement, maintenance and substitution of damaged tissues and organs using principles of biology and engineering (Langer and Vacanti, 1993) BTE consists of an interplay of scaffold technology, growth factors and. .. 2.1 BONE PHYSIOLOGY In order to regenerate bone using BTE techniques, the understanding of the structure and cells which participate in bone repair is of utmost importance Bone is composed of around 70-90% of minerals with the rest in the form of proteins Within the proteins in bone, the ratio of collagenous to noncollagenous stands at 9:1 This is in stark contrast with other tissues consisting of only... co-delivery system in a porcine model Optimal concentration of both BMP-2 and heparin was then determined and adapted for subsequent analysis Release profile of BMP-2 from PCL- TCP scaffolds was plotted for signs of sustained delivery when immersed in PBS solution Heparin was chosen to improve binding and regulate release of BMP-2 here and concurrently act as a framework for binding of endogenous growth... PCL- TCP scaffolds loaded with 100ng/ml BMP-2 Group C: PCL- TCP scaffolds loaded with 300ng/ml of heparin Group D: PCL- TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP2) xi CHAPTER 1: INTRODUCTION 1.1 BACKGROUND This chapter aims to give the reader an overview of the current trends in bone tissue engineering (BTE) Following that, limitations of available treatments for bone defects and. .. functions of BMPs include heart development (Callis et al., 2005; Simic and Vukicevic, 2005) and kidney formation (Simic and Vukicevic, 2005) There exists a heparin binding site in N-terminal region of mature BMP-2 polypeptide In pioneering work by Ruppert in 1996, it was shown that BMP-2 activity was increased upon interactions with heparins present in ECM In the presence of N terminus of BMP-2 and heparin,... The purpose of using composites for medical applications usually is to reduce drawbacks of individual materials and the benefits of both are 14 combined together Here, PCL is highly hydrophobic which leads to a longer degradation period (>2 years) in vitro and in vivo TCP alone, when fabricated into a scaffold is brittle and weak in strength By using PCL- TCP scaffolds for guided bone regeneration, the... (Chen and Mooney, 2003) 1.2 RESEARCH OBJECTIVES The general aim in this thesis was to evaluate and improve on the current properties of PCL- TCP composites from various aspects namely porosity and cell modulators in both in vitro and in vivo environments The two specific aims of this research was 1 To investigate the effects of bone regeneration, scaffold degradation and mechanical properties of PCL- TCP. .. loaded with 300ng/ml of heparin and 100ng/ml of BMP2) Figure 4.6: Amount of BMP-2 release at various time points (Group B: PCLTCP scaffolds loaded with 100ng/ml BMP-2 Group D: PCL- TCP scaffolds loaded with 300ng/ml of heparin and 100ng/ml of BMP2) Group A and C showed no release of BMP-2 at all time points Figure 4.7: Bioactivity of eluted BMP-2 at different time points (Group A: PCL- TCP scaffolds loaded... phase of bone in form of hydroxyapatite TCP is also responsible for the hardness of bone, dentine and enamel TCP exhibit excellent regenerative activity when placed in vivo (Beruto et al., 2000) However, it has poor mechanical properties such as low compressive strength This contributed to its brittleness when fabricated in blocks and scaffolds (K A Hing, 1998) 2.2.4 PCL- TCP scaffolds The purpose of using... regenerative capabilities Ectopic bone formation was induced using decalcified bone or injected bone extracts in one of the earliest study on bone regeneration (Senn, 1889) The breakthrough came about when Marshall Urist discovered that bone formed at ectopic sites in rodents upon addition of proteins extracted from demineralised bone matrix He named the protein Bone Morphogenetic Proteins (BMP)” as its regenerative ... bone tissue engineering for an ideal bone substitute 1.1.3 Strategies in BTE Tissue engineering is the restoration, improvement, maintenance and substitution of damaged tissues and organs using... 70-90% of minerals with the rest in the form of proteins Within the proteins in bone, the ratio of collagenous to noncollagenous stands at 9:1 This is in stark contrast with other tissues consisting... using principles of biology and engineering (Langer and Vacanti, 1993) BTE consists of an interplay of scaffold technology, growth factors and cells In this thesis, the focus will be on using composite