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Controlling Whole Blood Activation and Resultant Clot
Properties on Various Material Surfaces:
A Possible Therapeutic Approach for Enhancing Bone
Healing
Hoi Ting Shiu, BBiomedSc (Hons)
Thesis submitted in fulfilment of the requirements for the degree of
Doctor of Philosophy
Institute of Health and Biomedical Innovation
Science and Engineering Faculty
Queensland University of Technology
2012
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Hoi Ting Shiu Page i
Keywords
Bone healing, bone graft substitutes, biomaterials, biomaterial-blood interactions,
biocompatibility, blood clot formation, clot lysis, coagulation, complement, fibrin
network, leukocytes, platelets, platelet-derived growth factor, surface chemistry, surface
functional groups.
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Hoi Ting Shiu Page ii
Abstract
Injured bone initiates the healing process by forming a blood clot at the
damaged site. However, in severe damage, synthetic bone implants are used to
provide structural integrity and restore the healing process. The implant unavoidably
comes into direct contact with whole blood, leading to a blood clot formation on its
surface. Despite this, most research in bone tissue engineering virtually ignores the
important role of a blood clot in supporting healing.
Surface chemistry of a biomaterial is a crucial property in mediating blood-
biomaterials interactions, and hence the formation of the resultant blood clot. Surfaces
presenting mixtures of functional groups carboxyl (–COOH) and methyl (–CH3) have
been shown to enhance platelet response and coagulation activation, leading to the
formation of fibrin fibres. In addition, it has been shown that varying the compositions of
these functional groups and the length of alkyl groups further modulate the immune
complement response.
In this study, we hypothesised that a biomaterial surface with mixture of –
COOH/–CH3(methyl), –CH2CH3 (ethyl) or –(CH2)3CH3 (butyl) groups at different
ratios would modulate blood coagulation and complement activation, and eventually
tailor the structural and functional properties of the blood clot formed on the surface,
which subsequently impacts new bone formation.
Firstly, we synthesised a series of materials composed of acrylic acid (AA),
and methyl (MMA), ethyl (EMA) or butyl methacrylates (BMA) at different ratios
and coated on the inner surfaces of incubation vials. Our surface analysis showed
that the amount of –COOH groups on the surface coatings was lower than the ratios
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of AA prepared in the materials even though the surface content of –COOH groups
increased with increasing in AA ratios. It was indicated that the surface
hydrophobicity increased with increasing alkyl chain length: –CH3 > –CH2CH3 > –
(CH2)3CH3, and decreased with increasing –COOH groups. No significant
differences in surface hydrophobicity was found on surfaces with –CH3 and –
CH2CH3 groups in the presence of –COOH groups. The material coating was as
smooth as uncoated glass and without any major flaws. The average roughness of
material-coated surface (3.99 ± 0.54 nm) was slightly higher than that of uncoated
glass surface (2.22 ± 0.29 nm). However, no significant differences in surface
average roughness was found among surfaces with the same functionalities at
different –COOH ratios nor among surfaces with different alkyl groups but the same
–COOH ratios. These suggested that the surface functional groups and their
compositions had a combined effect on modulating surface hydrophobicity but not
surface roughness.
The second part of our study was to investigate the effect of surface functional
groups and their compositions on blood cascade activation and structural properties of
the formed clots. It was found that surfaces with –COOH/–(CH2)3CH3 induced a faster
coagulation activation than those with –COOH/–CH3 and –CH2CH3, regardless of the –
COOH ratios. An increase in –COOH ratios on –COOH/–CH3 and –CH2CH3 surfaces
decreased the rate of activation. Moreover, all material-coated surfaces markedly
reduced the complement activation compared to uncoated glass surfaces, and the pattern
of complement activation was entirely similar to that of surface-induced coagulation,
suggesting there is an interaction between two cascades. The clots formed on material-
coated surfaces had thicker fibrin with a tighter network at the exterior when compared
to uncoated glass surfaces. Compared to the clot exteriors, thicker fibrins with a loose
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network were found in clot interiors. Coated surfaces resulted in more rigid clots with a
significantly slower fibrinolysis after 1 h of lysis when compared to uncoated glass
surfaces. Significant differences in fibrinolysis after 1 h of lysis among clots on material-
coated surfaces correlated well with the differences in fibrin thickness and density at clot
exterior. In addition, more growth factors were released during clot formation than
during clot lysis. From an intact clot, there was a correlation between the amount of
PDGF-AB release and fibrin density. Highest amount of PDGF-AB was released from
clots formed on surfaces with 40% –COOH/60% –CH3 (i.e. 65MMA). During clot lysis,
the release of PDGF-AB also correlated with the fibrinolytic rate while the release of
TGF-β1 was influenced by the fibrin thickness. This suggested that different clot
structures led to different release profiles of growth factors in clot intact and degrading
stages.
We further validated whether the clots formed on material-coatings provide
the microenvironment for improved bone healing by using a rabbit femoral defect
model. In this pilot study, the implantation of clots formed on 65MMA coatings
significantly increased new bone formation with enhanced chondrogenesis,
osteoblasts activity and vascularisation, but decreased inflammatory macrophage
number at the defects after 4 weeks when compared to commercial bone grafts
ChronOSTM β-TCP granules. Empty defects were observed when blood clot
formation was inhibited.
In summary, our study demonstrated that surface functional groups and their
relative ratios on material coatings synergistically modulate activation of blood
cascades, resultant fibrin architecture, rigidity, susceptibility to fibrinolysis as well as
growth factor release of the formed clots, which ultimately alter the healing
microenvironment of injured bones.
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Table of Contents
Keywords .......................................................................................................................................... i
Abstract ............................................................................................................................................ ii
Table of Contents.............................................................................................................................. v
List of Figures ................................................................................................................................ vii
List of Tables ................................................................................................................................... ix
List of Abbreviations ........................................................................................................................ x
Statement of Original Authorship .................................................................................................... xii
Acknowledgments ......................................................................................................................... xiii
CHAPTER 1: INTRODUCTION .................................................................................................. 1 1.1 INTRODUCTION ................................................................................................................. 1
1.2 HYPOTHESIS ....................................................................................................................... 5
1.3 SPECIFIC AIMS OF THIS THESIS ...................................................................................... 5
CHAPTER 2: LITERATURE REVIEW ...................................................................................... 7 2.1 BONE NON-UNION IN CRITICAL SIZED DEFECTS ......................................................... 7
2.2 OVERVIEW OF BONE HEALING ....................................................................................... 8 2.2.1 Haemostasis ................................................................................................................ 8 2.2.2 Inflammation............................................................................................................. 18 2.2.3 Proliferation .............................................................................................................. 20 2.2.4 Remodelling .............................................................................................................. 22
2.3 CURRENT THERAPEUTIC APPROACH .......................................................................... 24 2.3.3 Natural Bone Grafts .................................................................................................. 24 2.3.2 Bone Graft Substitutes ............................................................................................... 25 2.3.3 Bone Tissue Engineering ........................................................................................... 26
2.4 PLATELET-RICH PLASMA............................................................................................... 30 2.4.1 Beneficial Role of Platelets & Uses of PRP ................................................................ 30 2.4.2 Role of Thrombin Concentration in Clot Structure ..................................................... 34 2.4.3 Effect of Clot Structure on Fibrinolysis ...................................................................... 35 2.4.4 Effect of Clot Structure on Viscoelastic Properties ..................................................... 37 2.4.5 In Vivo Implications of Altered Clot Structure, Properties and Stability ...................... 37 2.4.6 Differences Between a PRP gel and a Haematoma ..................................................... 41
2.5 IN SITU BLOOD CLOT FORMATION & MODIFICATION - A POSSIBLE TREATMENT FOR BONE DEFECTS ................................................................................................................... 44
2.5.1 Blood and Host Response towards Biomaterial Implants ............................................ 44 2.5.2 Influence of Biomaterial Surface Chemistry on Blood Response ................................ 48
CHAPTER 3: SYNTHESIS AND CHARACTERISATION OF MATERIAL-COATED SURFACES ................................................................................................................................... 56 3.1 INTRODUCTION ............................................................................................................... 56
3.2 MATERIALS ...................................................................................................................... 59
3.3 METHODS .......................................................................................................................... 59 3.3.1 Synthesis of materials ................................................................................................ 59 3.3.2 Preparation of surface coatings .................................................................................. 61 3.3.3 Characterisation of surface ........................................................................................ 61 3.3.4 Statistical analysis ..................................................................................................... 65
3.4 RESULTS ........................................................................................................................... 66
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3.4.1 Surface coating on incubation vials ............................................................................ 66 3.4.2 XPS analysis of material-coated surfaces ................................................................... 67 3.4.3 Surface hydrophobicity.............................................................................................. 72 3.4.4 Surface morphology and roughness ........................................................................... 75
3.5 DISCUSSION ...................................................................................................................... 79
3.6 CONCLUSION.................................................................................................................... 85
CHAPTER 4: THE INFLUENCE OF CARBOXYL AND ALKYL FUNCTIONAL GROUPS AND THEIR RELATIVE COMPOSITIONS ON BLOOD-BIOMATERIAL INTERACTIONS AND CLOT PROPERTIES .......................................................................................................... 86 4.1 INTRODUCTION ............................................................................................................... 86
4.2 MATERIALS & METHODS ............................................................................................... 89 4.2.1 Blood sampling and in vitro incubation ...................................................................... 89 4.2.2 In vitro coagulation activation ................................................................................... 89 4.2.3 In vitro complement activation .................................................................................. 90 4.2.4 Characterisation of clots formed on material-coated surfaces ...................................... 91 4.2.5 Statistical analysis ..................................................................................................... 95
4.3 RESULTS & DISCUSSION ................................................................................................ 96 4.3.1 Surface-initiated coagulation response ....................................................................... 96 4.3.2 Surface-initiated complement response .....................................................................104 4.3.3 SEM analysis of clot morphology and structure ........................................................110 4.3.4 Assessment of clot rigidity by compaction ................................................................121 4.3.5 Clot lysis ..................................................................................................................124 4.3.6 Quantification of PDGF-AB in serum and during clot lysis .......................................134 4.3.7 Quantification of TGF-beta 1 in serum and during clot lysis ......................................138
4.4 CONCLUSION...................................................................................................................146
CHAPTER 5: A PILOT STUDY OF THE OSTEOGENIC PROPERTIES OF EX VIVO BLOOD CLOTS FORMED ON MATERIALS IN A RABBIT FEMORAL DEFECT ............ 148 5.1 INTRODUCTION ..............................................................................................................148
5.2 MATERIALS & METHODS ..............................................................................................150 5.2.1 Preparation of coatings on scaffold surfaces ..............................................................150 5.2.2 Animals ...................................................................................................................151 5.2.3 Ex vivo blood clot formation .....................................................................................151 5.2.4 Surgical procedures in rabbit femur ..........................................................................152 5.2.5 Examination of defects .............................................................................................154 5.2.6 Statistical analysis ....................................................................................................158
5.3 RESULTS ..........................................................................................................................159 5.3.1 Micro-CT analysis for calcification in the defects......................................................159 5.3.2 Histological examination of de novo bone formation with H & E staining .................161 5.3.3 Histological examination of chondrogenesis in the defects ........................................164 5.3.4 Histological examinations of ALP expression ...........................................................167 5.3.5 Vascularisation revealed by vWF ..............................................................................170 5.3.6 Inflammation response revealed by CD68 .................................................................173
5.4 DISCUSSION .....................................................................................................................176
5.5 CONCLUSION...................................................................................................................181
CHAPTER 6: GENERAL DISCUSSION AND FUTURE DIRECTIONS ............................... 182 6.1 GENERAL DISCUSSION ..................................................................................................182
6.2 FUTURE DIRECTIONS .....................................................................................................191
6.3 GENERAL CONCLUSION ................................................................................................192
BIBLIOGRAPHY ....................................................................................................................... 194
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List of Figures
Figure 2-1. Vascular spasm in response to blood vessel damage.. ..................................................... 10
Figure 2-2. Platelet adhesion and aggregation .................................................................................. 12
Figure 2-3. The intrinsic and extrinsic pathways of coagulation system ............................................ 14
Figure 2-4. Conversion of fibrinogen to fibrin is mediated by thrombin ........................................... 16
Figure 2-5. Components of a haemostatic blood clot ........................................................................ 17
Figure 2-6. Haemostasis results in the formation of a haematoma at the injured bone ....................... 17
Figure 2-7. Inflammatory phase in injured bone. .............................................................................. 19
Figure 2-8. Proliferative phase of bone healing ................................................................................ 21
Figure 2-9. Remodelling phase of bone healing. .............................................................................. 22
Figure 2-10. Time sequence of four main phases of bone healing ..................................................... 23 Figure 2-11. The thrombin concentration present at the time of gelation dictates the fibre thickness and
density .................................................................................................................................... 35
Figure 2-12. Plasmin-mediated fibrinolysis ..................................................................................... 36
Figure 2-13. Two distinct patterns of peri-implant endosseous healing ............................................. 39
Figure 2-14. Schematic pictures illustrating differences in cellular components and fibrin scaffold between a) platelet-rich plasma (PRP) gel, and b) a normal haematoma. .................................. 43
Figure 2-15. Activation of complement system ................................................................................ 45
Figure 2-16. A multinucleated foreign body giant cell ..................................................................... 47
Figure 2-17. Self-assembled monolayers (SAMs) ............................................................................ 49
Figure 2-18. Copolymerisation of two monomers M1 and M2 .......................................................... 53
Figure 3-1. Free-radical polymerisation is initiated by benzoyl peroxide .......................................... 60 Figure 3-2. Copolymer surfaces displaying various functional groups .............................................. 63
Figure 3-3. Water contact angle measurement ................................................................................. 64
Figure 3-4. a) Materials formed from free-radical polymerisation. b) An incubation vial treated with material solution resulted in a clear coating. ............................................................................ 66
Figure 3-5. XPS survey spectra of a) uncoated glass, b) PAA, c) PBMA, and d) 45BMA (45% AA/BMA) coated surfaces. ..................................................................................................... 69
Figure 3-6. XPS C1s spectra of a) PAA, b) PMMA, d) PEMA, f) PBMA, and c) 45MMA, e) 45EMA, and g) 45BMA coated surfaces. .............................................................................................. 70
Figure 3-7. Ratio of –COOH groups on the surface coating as a function of mole fraction of –COOH group-containing AA composed with a) MMA, b) EMA or c) BMA........................................ 72
Figure 3-8. Advancing contact angles of different surface coatings. ................................................. 74
Figure 3-9. Representative SEM images of a) uncoated glass, b) PEMA and c) 45 EMA coated surfaces, taken at magnification of 25000 x. ............................................................................ 76
Figure 3-10. AFM images (5 µm x 5 µm areas) of uncoated and coated surfaces. ............................. 77
Figure 4-1. The serum levels of prothrombin F1+2 after 30 min of whole blood incubation with material-coated surfaces relative to the uncoated glass surfaces (%) ........................................ 97
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Figure 4-2. The serum levels of C5a-desArg after 2 h of whole blood incubation with material-coated surfaces relative to the uncoated glass surfaces (%) ................................................................105
Figure 4-3. Cross-talk between complement and coagulation cascades ............................................107
Figure 4-4. Scanning electron microscopy analysis of whole blood clot structures formed on 45BMA, 55BMA, 65BMA and uncoated glass surfaces ........................................................................112
Figure 4-5. Scanning electron microscopy analysis of structure of clots formed on 55MMA, 55EMA, 55BMA and uncoated glass surfaces ......................................................................................114
Figure 4-6. Scanning electron microscopy analysis of structure of clots formed on 65MMA, 65EMA, 65BMA and uncoated glass surfaces ......................................................................................116
Figure 4-7. Compaction studies of clots formed on various material-coated surfaces compared to the uncoated glass surfaces ..........................................................................................................122
Figure 4-8. Release of D-dimer and weight loss over 24 h lysis of clots formed on BMA surfaces compared with uncoated glass surfaces ..................................................................................128
Figure 4-9. Release of D-dimer and weight loss over 24 h lysis of clots formed on 55MMA, 55EMA and 55BMA surfaces compared with 65MMA, 65EMA and 65BMA surfaces ........................132
Figure 4-10. The serum levels of PDGF-AB after 2 h of whole blood incubation with material-coated surfaces compared to the uncoated glass surfaces and the plasma baseline ..............................135
Figure 4-11. In vitro release of PDGF-AB during lysis of clots formed on material-coated surfaces and uncoated glass surfaces ..........................................................................................................137
Figure 4-12. The serum levels of TGF-β1 after 2 h of whole blood incubation with material-coated surfaces compared to the uncoated glass surfaces and plasma baseline....................................139
Figure 4-13. In vitro release of TGF-β1 during lysis of clots formed on material-coated surfaces and uncoated glass surfaces ..........................................................................................................141
Figure 5-1. Stainless steel scaffold coated with material solution ....................................................150
Figure 5-2. Implantation of ex vivo blood clots formed on material-coated scaffolds in rabbit femoral defects...................................................................................................................................153
Figure 5-3. Micro-CT scanning analysis on the femoral defects ......................................................160
Figure 5-4. De novo bone formation in the defects shown by H&E staining ....................................163
Figure 5-5. Chondrogenesis in the defects shown by Alcian Blue staining .......................................166
Figure 5-6. Osteoblasts in the defect shown by ALP staining ..........................................................169
Figure 5-7. Vascularisation of the defects shown by vWF staining ..................................................172
Figure 5-8. Inflammatory response evaluated by the number of macrophages at the defects using CD68 staining .......................................................................................................................175
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List of Tables
Table 2-1. Growth factors of platelets and their functions ................................................................ 31
Table 3-1. Ratio of –COOH groups measured on surface coatings (XCOOH coating) compared to mole fraction of –COOH group-containing AA (XCOOH material) composed with MMA, EMA or BMA ...................................................................................................................................... 71
Table 3-2. Advancing contact angles of surfaces coated with materials composed of varied mole fraction of acrylic acid and alkyl methacrylates ....................................................................... 74
Table 3-3. Average surface roughness measured by AFM ................................................................ 78
Table 5-1. Three treatment groups in the animal study. ...................................................................153 Table 5-2. Primary and secondary antibodies used in immunohistochemistry and immunofluorescence
studies. ..................................................................................................................................158
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List of Abbreviations
AA Acrylic acid ABC Avidin-biotin complex ADP Adenosine-5’ diphosphate AFM Atomic force microscopy ALP Alkaline phosphatase BPO Benzoyl peroxide BMA Butyl methacrylate BMP Bone morphogenic protein Β-TCP Beta tricalcium phosphate CSD Critical sized defect DAB Diaminobenzidine DAPI 4’, 6-diamindino-2-pheylindole EDTA Ethylenediamine tetraacetic acid EGF Epidermal growth factor EMA Ethyl methacrylate ELISA Enzyme-linked immunosorbent assay FBR Foreign body reaction FBGCs Foreign body giant cells FDPs Fibrin degradation products FGF Fibroblast growth factor FII Prothrombin (factor II) FIIa Thrombin (activated factor II) FIXa-FVIIIa Intrinsic tenase complex (Factor IXa–Factor VIIIa) FVII Factor VII FXa-FVa Extrinsic tenase complex (Factor Xa-Factor Va) FXII Factor XII FXIIa Activated factor XII F1+2 Prothrombin fragment 1+2 GP Ib/IX Glycoprotein Ib/IX receptor GP IIb/IIIa Glycoprotein IIb/IIIa receptor HA Hydroxyapaptite H&E Haematoxylin and eosin HMWK High molecular weight kininogen IGF-1 Insulin-like growth factor -1 IL-1 Interleukin-1 L-PRP Leukocyte- and platelet-rich plasma L-PRF Leukocyte- and platelet-rich fibrin MAC Membrane attack complex MMA Methyl methacrylate MP Microparticle PAI Plasminogen activator inhibitor PBS Phosphate buffered saline PDGF Platelet-derived growth factor
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PDEGF Platelet-derived epidermal growth factor PDAF Platelet-derived angiogenesis factor PEG Polyethylene glycol PF-4 Platelet factor 4 PIBMA Poly(isobutyl methacrylate) PL Phospholipid PLGA Poly(lactic-co-glycolic acid) PMMA Poly(methyl methacrylate) P-PPP Pure platelet-rich plasma P-PRF Pure platelet-rich fibrin PRP Platelet-rich plasma RGD Arginine - Glycine - Asparatic acid RSFs Relative sensitivity factors SAMs Self-assembled monolayers SEM Scanning electron microscopy tPA Tissue-type plasminogen activator TAFI Thrombin activatable fibrinolysis inhibitor TCC Terminal complement complex TF Tissue factor TF-FVIIa Extrinsic tenase complex (Tissue factor - Factor VIIa) TFPI Tissue factor pathway inhibitor TGF-β Transforming growth factor- β TNF-α Tumor necrosis factor-α VEGF Vascular endothelial growth factor vWF von Willebrand factor XPS X-ray photoelectron spectroscopy
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Statement of Original Authorship
The work contained in this thesis has not been previously submitted to meet
requirements for an award at this or any other higher education institution. To the best of
my knowledge and belief, the thesis contains no material previously published or written
by another person except where due reference is made.
Signature : _________________________
Hoi Ting Shiu
Date : _________________________ 7th June, 2012
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Acknowledgments
This thesis would not have been possible without the help and encouragement
of many people.
First and foremost, I am indebted to my principal supervisor, Dr Ben Goss, for
providing me with this great opportunity to work on a project in such a dynamic field.
His invaluable suggestions, patience and continuous support throughout my candidature
is greatly appreciated and undoubtedly led to my professional maturation.
My sincere thanks go to my associate supervisor, Dr Cameron Lutton, for his
guidance and enthusiasm which has given me confidence whenever experiments did not
go as planned. His willingness to share his innovative ideas has been truly inspiring and
of great model.
I would like to express my sincere gratitude to my associate supervisor,
Professor Yin Xiao who participated in our team during my PhD journey, for his
motivation, immense knowledge and constructive comments from the initial conception
to the end of this work. Without his technical assistance and scientific criticism, my
project would not have progressed to the stage that it did. Learning and working with
him has been beneficial and enjoyable by his training and his lively attitude.
I also wish to thank my associate supervisor, Professor Ross Crawford for his
stimulating clinical insights as well as persistence in research which have inspired me to
tackle the challenges faced during the period.
A warm thank you goes to all my colleagues, Will, Kunnika, Rinku, Shirly,
Indira, Willa and Navdeep for assisting me with laboratory techniques and for their
friendship which has enriched my PhD years with conversations ranging from food
science to jokes of having “ permanent head damaged (P-h-D)”.
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I wish to thank Dr Sadahiro Sugiyama for his veterinary expertise with animal
surgery. I also thank Dr Barry Wood of UQ for his support with X-ray photoelectron
microscopy and Associate Professor Nunzio Motta for assisting with atomic force
microscopy.
The financial support from IHBI and the Queensland University of Technology
is also gratefully acknowledged.
To my parents goes my deepest gratitude, for their endless love and support
throughout my life. I am grateful to my father for his care and trust, who initially held
opposed opinions on doing research but finally allowed me to explore myself more
through taking honours research project and subsequent this PhD. To my mother, no
words can sufficiently describe my deep gratitude to her, for putting an enormous effort
to provide the best possible environment for me to grow and accepting me as whom I
am. Her accompany during this journey for reminding me to balance my life has been
very important to me. I know no one would be more delighted than her for my
achievement. I wish to dedicate this thesis on her name.
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1Chapter 1: Introduction
1.1 INTRODUCTION
Healing of injured bone begins with haemostasis, leading to the formation of a
blood clot at the injured site (Marieb, 2001c). In severe damage, however, bone healing
is impaired due to a number of reasons, including deficient biological factors, instability
at the site, and uncontrollable inflammation (Doblare et al., 2004, Tsiridis et al., 2006).
Based on these limiting factors, intensive research aims to develop an ideal bone
substitute providing three key elements for bone healing: the scaffolding for
osteoconduction1, growth factors for osteoinduction2, and progenitor cells for
osteogenesis3 (Burg et al., 2000, Parikh, 2002, Crawford and Hatton, 2008, Schliephake,
2009). To date, no engineered material outperforms autograft in bone-forming ability
(Sammarco and Chang, 2002, Avramoglou et al., 2005, Harwood et al., 2010b).
Despite significant progress in biomaterial development and modification, the
use of non-toxic materials is often complicated with immune response and foreign body
reaction, leading to fibrotic encapsulation and implant dysfunction (Anderson et al.,
2004b, Tsai, 2004, Williams, 2008). This indicates that the reduction of immune
response towards the implant and initial formation of a microenvironment at the
1 Capable of supporting the biological process of bone healing such as ingrowth of capillaries and attachment of osteoprogenitor cells by providing a structural scaffold.
2 Capable of inducing bone formation by supplying biological factors.
3 The process of new bone formation by osteoblasts.
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implant/tissue interface are important for bone healing to proceed (Anderson et al.,
2008).
Platelet-rich plasma (PRP) which has been applied in clinical dentistry for over
a decade provides an insight into the microenvironment required. PRP is a fraction of
plasma with a high concentration of platelets and serves as an autologous source of
growth factors (Marx, 2001, Carlson and Roach, 2002, Griffin et al., 2009). Studies have
shown that formation of PRP gels to bone grafts or biomaterials increases the rate of
bone formation and bone density (Marx et al., 1998b, Choi et al., 2004, Wiltfang et al.,
2004, Roussy et al., 2007). Increased concentrations of growth factors derived from
platelets, and adhesive binding of graft particles by the fibrin network are believed to be
responsible for the beneficial effects of PRP gels (Lieberman et al., 2002a, Eppley et al.,
2004, Dolder et al., 2006). However, other studies have shown conflicting results when
PRP gels were prepared with different platelet numbers and thrombin concentration
(Lacoste et al., 2003, Weibrich et al., 2004, Gimeno et al., 2006). Collectively, these
findings imply that the effect of clots on bone healing depends on how the clots are
formed.
Thrombin concentration is known to affect fibrin thickness and density during
clot formation (Carr et al., 2002b, Wolberg, 2007b). Whilst abnormal changes in clot
structure have been shown to influence the viscoelastic properties (Collet et al., 2005,
Liu et al., 2006, Jámbor et al., 2009) and the lysis rate of clots (Gabriel et al., 1992,
Collet et al., 2000), leading to pathological conditions such as cardiovascular thrombosis
and bleeding disorders (Collet et al., 1993a, Mills et al., 2002). Indeed, dental
implantology has long demonstrated that a peri-implant clot plays an important role in
endosseous healing (Davies, 2003a, Carlsson et al., 2004). Chemokines released from
entrapped blood cells and the fibrin network of the peri-implant have been shown to
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support continuous recruitment of osteogenic cells, resulting in a direct bone formation
on the implant surfaces (Davies, 1998, Oprea et al., 2003). These reports indicate that the
properties of a blood clot formed on the implant would influence cellular response and
tissue ingrowth in bone healing.
However, most current approaches in bone tissue engineering focus solely on
the replacement of viable cells, growth factors and functional structure using engineered
scaffolds. It often overlooks the importance of blood clot formation around scaffolds in
controlling bone healing. We hypothesise that the bone-inducing ability of a biomaterial
may be improved by controlling the blood-biomaterial interactions and forming a
desirable peri-implant blood clot with appropriate properties.
Surface chemistry is one of the most crucial parameters in modulating
interactions between blood and biomaterial (Mrksich, 2000, Thevenot et al., 2008,
Tzoneva et al., 2008). Surfaces with mixtures of carboxyl (–COOH) and methyl (–CH3)
groups have been shown to increase platelet adhesion and activation, as well as thrombin
generation leading to the formation of sizeable fibrin fibres on the surfaces, compared to
the surfaces with either groups only. In addition, the extent of complement activation and
leukocyte response were also found to be modulated by the surfaces with –COOH/–CH3
functionalities at different ratios (Sperling et al., 2005a, 2009, Fischer et al., 2010b).
Moreover, the studies of Berglin et al (2004, 2009) indicated that the chain length of
alkyl groups has a regulatory role in surface-mediated coagulation and complement
activation. These findings lead us to consider binary mixtures of –COOH/–CH3(methyl),
–CH2CH3 (ethyl) or –(CH2)3CH3 (butyl) functionalities at different ratios to modulate
blood clot formation and immune response to biomaterials.
To elucidate closely the in vivo blood-biomaterial interactions, several issues
needed to be addressed. Firstly, whole blood must be used. General approaches in
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investigating blood-biomaterial interactions simplifies the whole blood system by
dividing it into plasma, isolated protein solution or single cell types (Eriksson and
Nygren, 2001, Gemmell, 2001, Sperling et al., 2005a, Thor et al., 2007, Fischer et al.,
2010b). Although the simplification permits close examination of protein and cell
reactions involved, which might be otherwise greatly affected by contributions of
interactive events. Secondly, anticoagulation of whole blood is performed in most
studies to achieve sensitivity of the test at manageable incubation periods (Streller et al.,
2003). Heparin is a commonly used anticoagulant (Berry and Chan, 2008). Besides
inactivating coagulation (Fushimi et al., 1998, Klement et al., 2002), heparin also
reduces complement response (Kopp et al., 2002, Lappegård et al., 2004, Sperling et al.,
2006). As such, using an incubation assay of whole blood without anticoagulant would
enable a more accurate evaluation of the coagulation/complement activation, and reflects
closely the structural features of peri-implant clots modified by surface chemistry.
Furthermore, an improvement in in vitro investigations of the surface chemistry would
require a feasible biomaterial, rather than a standardised flat surface model such as self-
assembled monolayer of alkanethiols (SAMs) on gold, which have been employed to
display various functional groups (Faucheux et al., 2004, Tsai et al., 2007, Arima and
Iwata, 2007).
In this study, novel materials composed of acrylic acid and alkyl (methyl, ethyl,
or butyl ) methacrylate at different ratios were formed to present –COOH/–CH3, –
CH2CH3 or –(CH2)3CH3 functionalities. Customised incubation vials coated with the
materials were used to investigate the impact of surface functional groups and their
relative ratios on blood coagulation and complement activation, the formation of a three-
dimensional blood clot, in term of the clot structure, rigidity, susceptibility to lysis and
release of growth factors. To validate the concept that a peri-implant blood clot
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modulated by material coatings could enhance the healing microenvironment, therefore
improving bone regeneration, a pilot study of animal defects treated with an ex vivo
blood clot formed on an optimal material coating was also performed.
1.2 HYPOTHESIS
We hypothesise that the surface functional groups at various compositions on a
biomaterial surface can modulate blood coagulation and complement activation, and
subsequently tailor the architecture and functional properties of the blood clot formed on
the surface, which eventually impacts new bone formation.
1.3 SPECIFIC AIMS OF THIS THESIS
Three specific aims are addressed in this study:
1. The first aim was to synthesise materials of acrylic acid and alkyl
methacrylates and characterise the physiochemical properties of the material
coatings. Materials were prepared by free radical polymerisation of acrylic
acid and alkyl methacrylates ranging from methyl, ethyl or butyl terminal
groups at various ratios. An incubation vial coated with material was
designed to investigate whole blood response to the surface in a three-
dimensional manner. Surface analysis was performed by X-ray photoelectron
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spectroscopy, advancing water contact angle measurement, atomic force
microscopy and scanning electron microscopy.
2. The second aim was to evaluate the influence of surface chemistries of
material coatings on blood coagulation and complement response, and clot
formation after incubation. The extent of coagulation/complement activation
was determined by quantifying cascades related end products using enzyme-
linked immunosorbent assay. The blood clots formed in material-coated vials
were characterised by measuring fibrin thickness and density, clot rigidity,
susceptibility to lysis and the release of growth factors: platelet-derived
growth factors-AB (PDGF-AB) and transforming growth factor-beta 1 (TGF-
β1).
3. The third aim was to perform a pilot animal study to validate the concept that
the blood clots formed on an implant provide the initial microenvironment for
bone healing and assess the in vivo osteogenic potential of a blood clot
formed on material-coated stainless steel scaffolds. The optimal material was
selected based on in vitro analysis from previous aims. A rabbit femoral
defect model was established and autologous rabbit blood was incubated with
the scaffolds to generate an ex vivo blood clot prior to implantation.
Calcification, osteogenesis and chondrogenesis as well as angiogenesis and
prolonged inflammation were analysed using Micro-CT scanning, histology
and immunohistochemistry after 4 weeks.
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2Chapter 2: Literature review
2.1 BONE NON-UNION IN CRITICAL SIZED DEFECTS
Bone healing is a response of injured bone to restore its loss of structure and
function by forming new bone (Marieb, 2001b, Hench, 2005b). However when
extensive bone is lost in large defects such as trauma and tumour resections, the healing
may proceed abnormally slowly or completely attenuate (Caputo, 1999). A critical sized
defect (CSD) is defined as the smallest size of a damaged bone that would not heal
spontaneously during the life time of the animal (e.g. approximately a loss of bone
segment with length exceeding 3 times the diameter of the affected long bone in sheep)
(Hertel et al., 2001, Lindsey et al., 2006, Liu et al., 2008, Reichert et al., 2009). This
leads to a condition termed non-union, as commonly seen in bone fractures (Phillips,
2005, Jahagirdar and Scammell, 2009, Harwood et al., 2010a). Patients with incomplete
healing suffer from prolonged morbidity and dysfunction, resulting in decreased quality
of life (Tsiridis et al., 2006).
More than 500,000 bone graft procedures are performed in the United States
and approximate 2.2 million worldwide every year (Giannoudis et al., 2005, Tosounidis
et al., 2009). Currently, the treatment using autograft is limited by graft availability from
the patients and donor site morbidity (Goulet et al., 1997). In addition, no other implants
are known to perform as well as or better than autograft (Sammarco and Chang, 2002,
Avramoglou et al., 2005). Therefore, a new strategy to improve the capability of
artificial bone grafts in enhancing bone healing is urgently needed. To develop effective
therapeutic interventions, it is important to understand the mechanism of bone healing.
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2.2 OVERVIEW OF BONE HEALING
Bone repairs itself through similar phases of normal wound healing and
some bone tissue-specific processes. Typically, there is an overlap of four major
phases (Caputo, 1999, Marieb, 2001b):
1. Haemostasis
2 Inflammation
3. Proliferation
4. Remodelling
These dynamic processes begin immediately upon injury and include a
series of interactions among various cell types, mediators and extracellular matrix
(Dimitriou et al., 2005, Nurden et al., 2008). Two distinct mechanisms of bone
formation occur in bone repair, namely intramembranous and endochodral
ossification (Marieb, 2001b, Doblare et al., 2004, Harwood et al., 2010a). A bone
fracture where there is a loss of continuity is used herein to illustrate the bone healing
(Tsiridis et al., 2006, Tosounidis et al., 2009).
2.2.1 Haemostasis
As a vascularised tissue, a fracture in bone is associated with the damage of
blood vessels in the periosteum, endosteum and surrounding soft tissue (Harwood et
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al., 2010a). Haemorrhage and oedema at the fracture site disrupt oxygen and nutrient
supply to osteocytes and eventually cause necrosis of broken bone ends (Caputo,
1999, Carano and Filvaroff, 2003, Tosounidis et al., 2009). Haemostasis is a tightly
regulated process that prevents excessive blood loss and maintains blood flow to the
rest of body. It involves vasoconstriction, formation of platelet plugs and coagulation
(Allford and Machin, 2007).
2.2.1.1 Vasoconstriction of damaged vessels
On damage of the endothelium which lines the lumen of blood vessels,
underlying extracellular matrix proteins are exposed to whole blood. Subendothelium
collagen (principally type I and III), von Willebrand factor (vWF) and fibronectin
interact with platelets through various glycoprotein receptors, thereby supporting
platelet adhesion to the damaged site (Hantgan et al., 1990, Sugimoto and Miyata,
2002, Tsai et al., 2002, Ruggeri, 2003, Schmugge et al., 2003, Nurden, 2007, Bennett
et al., 2009). The interactions also stimulate platelets to release contents from their
granules, including serotonin and thromboxane A2 (Watanabe and Kobayashi, 1988,
Gobbi et al., 2003). These vasoactive factors induce contraction of vascular smooth
muscle cells (Golino et al., 1989) and together with cytokine-stimulated endothelial
cells, damaged blood vessels are constricted to reduce extravasation of blood
constituents (Minors, 2004, Anitua et al., 2004) (Figure 2-1).
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Figure 2-1. Vascular spasm in response to blood vessel damage. Exposed subendothelium proteins bind and activate platelets to release vasoactive factors, together with contraction of smooth muscle cells leading to vessel constriction.
2.2.1.2 Platelets and Formation of a Platelet Plug
Platelets (also termed thrombocytes) play a central role in regulating
haemostasis. They are anuclear, disc-shaped cytoplasmic fragments (diameter 3 - 4
µm) derived from megakaryocytes in the bone marrow. They circulate at an average
concentration of 200 million per mL in the blood (Marieb, 2001d, Nurden et al.,
2008). Platelets become active when they are exposed to thrombogenic surfaces (e.g.
injured endothelium and subendothelium collagen) and soluble components such as
adenosine-5’diphosphate (ADP) and thrombin (Brummel et al., 2002, Andrews and
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Berndt, 2004, Allford and Machin, 2007). Activated platelets exhibit five physiologic
responses, including morphological change to spherical with psedudopodia, exposure
of phosphatidylserine and receptors on cell membrane, degranulation, activation of
cytoskeletal contractile apparatus, and formation of platelet microparticles (MP)
(Blockmans et al., 1995, Blair and Flaumenhaft, 2009, Rand et al., 2010, Nurden,
2011). All these responses have been shown to augment platelet adhesion and
activation at the damaged vessels (Andrews and Berndt, 2004, Gibbins, 2004). In
particular, binding of blood protein, fibrinogen, with platelet glycoprotein IIb/IIIa
receptor (GP IIb/IIIa) further facilitates aggregation of adjacent platelets, leading to
the formation of a platelet plug (Bennett, 2001, 2009, Dubois et al., 2004, Sivaraman
and Latour, 2011) (Figure 2-2). This platelet plug not only preserves the integrity of
the vessel, but also supports coagulation activation (Gorbet and Sefton, 2004,
Minors, 2004).
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Figure 2-2. Platelet adhesion and aggregation. a) Platelets adhesion on subendothelium is mediated by the interaction between adsorbed von Willebrand factor (vWF) and surface glycoprotein Ib/IX receptor (GP Ib/IX). Plasma fibrinogen aggregates more platelets onto the subendothelium by binding glycoprotein IIb/IIIa receptors (GP IIb/IIIa) on nearby platelets. b) Scanning electron micrograph of platelet aggregates (pale pink) and red blood cells (pink). Magnification 4000x. Adapted from http://www. inmagine.com/spl012/spl012882-photo.
2.2.1.3 Coagulation and Formation of a Haematoma
The main purpose of coagulation is to produce a stable haemostatic clot by
forming fibrin mesh on the platelet plugs. Coagulation system can be initiated by two
pathways: intrinsic and extrinsic. For both pathways, a series of proteolytic reactions
where an enzyme precursor becomes active will trigger the activation of another
precursor in the downstream cascade. Both pathways converge to a common final
pathway resulting in the formation of thrombin (Davie et al., 1991, Minors, 2004).
Thrombin is an enzyme which mediates the formation of fibrin.
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The intrinsic pathway begins when prekallikrein, high molecular weight
kininogen (HMWK) and factor XII (FXII) contact with thrombogenic surfaces
(Yarovaya et al., 2002, Zhuo et al., 2005, Renné and Gailani, 2007). Upon cleavage
on the surfaces, FXII becomes activated FXIIa and subsequently leads to the
formation of an intrinsic tenase complex (FIXa-FVIIIa) (Vogler and Siedlecki, 2009)
(Figure 2-3).
On the other hand, the initiation of extrinsic pathway is dependent on the
presence of tissue factor (TF). TF is highly expressed on or released from various
cells (e.g. endothelial cells, activated platelets and monocytes/macrophages)
following vascular damage and inflammatory stimuli such as endotoxin, tumor
necrosis factor α (TNFα) and interleukin-1α (IL-1α) (Altieri, 1995, Ernofsson et al.,
1997, Bouchard and Tracy, 2001, Esmon, 2004). TF binds and activates factor VII
(FVII) in the plasma, thereby becoming the extrinsic tenase complex (TF-FVIIa).
For each pathway, it has been revealed that activated platelets provide the
phospholipid surfaces for the assembly and function of the complexes, and hence
greatly propagate the coagulation activation (Esmon, 1995, Bouchard and Tracy,
2001, Smyth et al., 2009). Both tenase complexes convert factor X (FX) to FXa via
the common pathway. When FXa binds to an activated factor V (FVa), they form a
prothrombinase complex (FXa-FVa) locally on the platelet membrane. This complex
cleaves prothrombin (Factor II) and produces thrombin (Factor IIa) (Mann et al.,
2003, Minors, 2004, Vogler and Siedlecki, 2009). As a result, thrombin mediates the
conversion of fibrinogen to fibrin, including fibrinogen bound to the platelet plugs,
thereby restricting formation of fibrin mesh to the location of damaged vessel
(Jirousková et al., 1997, Clark, 2001, Di Cera, 2003) (Figure 2-3).
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Figure 2-3. The intrinsic and extrinsic pathways of coagulation system. The intrinsic pathway starting with surface contact activations of prekallikrein, high molecular weight kininogen (HMWK) and factor XII (FXII) is shown as a linear cascade of zymogen activation steps, leading to the formation of intrinsic tenase complex (FIX-FVIIIa). In parallel, the extrinsic pathway is initiated by tissue factor (TF) generated during trauma. TF activates factor VII (FVII) into FVIIa, and form extrinsic tenase complex (TF-FVIIa). Both tenase complexes from respective pathways merge at the common pathway in which factor X (FX) is converted factor Xa (FXa). FXa which in turn binds to activated factor V (FVa) forming the prothrombinase complex (FXa-FVa) that converts prothrombin to thrombin. Thrombin as the end product of coagulation activation subsequently catalyses the formation of fibrin from fibrinogen. Phospholipids (PL) membrane of platelets and calcium ion (Ca2+) serve as cofactors of the process (Minors, 2004).
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In the conversion of fibrinogen to fibrin, thrombin sequentially cleaves
fibrinopeptides A and B from fibrinogen, resulting in the generation of fibrin
monomers (Weisel et al., 1993, Blomback and Bark, 2004, Riedel et al., 2011). The
monomers interact in a half-staggered end-to-end fashion and become double-
stranded protofibrils. Following lateral aggregation of the protofibrils, fibrin fibres
are formed, branched out and eventuated in a three-dimensional network on the
platelet plugs (Brummel et al., 2002, Mosesson, 2005, Wolberg, 2007a). Thrombin-
activated factor XIII forms cross-links between neighbouring fibres in the network.
Hence, the resultant clot is strengthened against flow, mechanical and proteolytic
impacts (Laurens et al., 2006, Rojkjaer and Rojkjaer, 2007, Jámbor et al., 2009)
(Figure 2-4).
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Figure 2-4. Conversion of fibrinogen to fibrin is mediated by thrombin. Fibrinogen is a trinodular protein consisting two sets of three different polypeptide chains: Aα, Bβ, and γ, assembling with their N-termini in a central E domain. The C-termini of Bβ, and γ chains extend to form distal D domains while the C-termini of Aα chains follow to the D domains and extend back to interact intramolecularly with the central E domain (Mosesson, 2005, Weisel, 2005). Thrombin releases the fibrinopeptides A and B from E domain of fibrinogen sequentially, forming the fibrin monomers. Interacting in a half-staggering and end to end manner, fibrin monomers polymerise into protofibrils. By lateral aggregation and branching out of protofibrils, fibrin polymers are formed in a three-dimension. A stable fibrin clot is formed by factor XIIIa-mediated crosslinking between the γ chains in D-domian in the fibrin network.
An array of proteins are incorporated into the clot during fibrin formation.
For example, vWF, fibronectin, collagen, albumin, tissue-type plasminogen activator
(tPA), plasminogen activator inhibitor (PAI), α2-antiplamin and fibroblast growth
factor-2 (FGF-2) (Standeven et al., 2005, Weisel, 2005, Mosesson, 2005). Also,
platelets and other blood cells such as erythrocytes and leukocytes are entrapped in
the growing clot, resulting in a more definitive haemostatic fibrin clot (Laurens et al.,
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2006) (Figure 2-5). A clot, which is formed between the fractured bone ends is called
a haematoma (Figure 2-6) (Street et al., 2000, Marieb, 2001b).
Figure 2-5. Components of a haemostatic blood clot. a) It is composed of platelet aggregates, erythrocytes and b) leukocytes entrapped in the fibrin network. Scanning electron micrograph b) at 5000 x, scale bar 20 µm. a) Adapted from http://www.biocurious.com/new-perspective-on-blood-clot-mechanics.html.
Figure 2-6. Haemostasis results in the formation of a haematoma at the injured bone. Modified from www.uwo.capatholcasesSkeletalfracture.html.
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2.2.2 Inflammation
The second stage of healing is inflammation. In the affected tissue, histamine
and prostaglandins released by local mast cells trigger inflammation (Marieb, 2001a,
Esmon, 2003, Nurden, 2011). These mediators cause vasodilation and increase vascular
permeability, leading to the influx of inflammatory cells, plasma proteins and fluid into
the injured site. The influx produces the cardinal signs of redness, sweating, heat and
pain (Marieb, 2001a, 2001e, Hench, 2005a). In addition, activated platelets in the
haematoma degranulate a series of cytokine and growth factors including platelet factor
4 (PF-4), platelet-derived growth factor (PDGF) and transforming growth factor-beta
(TGF-β) (Hosgood, 1993, Anitua et al., 2004, Frechette et al., 2005, Blair and
Flaumenhaft, 2009, Müller et al., 2009). These factors are revealed to stimulate
chemotaxis of neutrophils, monocytes, fibroblasts and progenitor cells into the damaged
site (Deuel et al., 1981, Pierce et al., 1991, Cross and Mustoe, 2003, Andrae et al., 2008).
Neutrophils are the first type of inflammatory cells that migrate from blood into
the damaged tissue (Marieb, 2001a, Esmon, 2003). Their migration is also mediated by
complement protein C5a, an end product of complement cascade in innate immune
response (Guo and Ward, 2005, Ritis et al., 2006, Nilsson et al., 2007a). Neutrophils
remove necrotic cellular debris, pathogens and foreign materials by phagocytosis. After
24-48 hours, the population of neutrophils diminish due to their short lifespan and are
replaced by blood monocytes (Park and Barbul, 2004, Janeway et al., 2005a, Anderson
et al., 2008). The monocytes differentiate into long-life (up to months) macrophages in
the tissue in response to cytokines, extracellular metabolites and the hypoxic
environment (Hench, 2005a, Anderson et al., 2008). Osteoclasts are also activated to
reabsorb bone debris. In addition to performing phagocytosis, macrophages release
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reactive oxygen species and growth factors. They release fibroblast growth factor (FGF)
and epidermal growth factor (EGF) to mediate the growth of fibroblasts, new blood
vessels and epithelial cells (Harwood et al., 2010a). Fibroblasts and vascular endothelial
cells migrate into the haematoma and contribute to form granulation tissue
approximately 3 to 5 days following the fractures. Undifferentiated mesenchymal cells
which originate from the periosteum of broken bone also migrate into the site (Doblare
et al., 2004, Tsiridis et al., 2006, Tosounidis et al., 2009, Harwood et al., 2010a) (Figure
2-7).
Figure 2-7. Inflammatory phase in injured bone. It involves the influx of inflammatory cells, plasma proteins and fluid to the haematoma, followed by infiltration of fibroblasts which deposit collagen matrix, and endothelial cells which forms new capillaries. Modified from www.uwo.capatholcasesSkeletalfracture.html.
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2.2.3 Proliferation
Granulation tissue is a specialised type of tissue characterised by the
proliferation of new small blood vessels and fibroblasts (Marieb, 2001a). Fibroblasts
synthesise collagen to form a new matrix within the clot while mature endothelial cells
organise into new capillaries from pre-existing vessels in a process known as
angiogenesis (Glowacki, 1998, Carano and Filvaroff, 2003). Angiogenesis is important
in tissue regeneration; as newly established blood flow not only provides oxygen and
nutrient, but also mesenchymal stem cells and signalling molecules (Gerber and Ferrara,
2000, Carano and Filvaroff, 2003, Weiss et al., 2009). During this proliferative phase,
the clot is gradually degraded by plasmin-mediated fibrinolysis while neutrophils and
macrophages clean the remaining by-products of fibrinolysis (Erickson et al., 1985,
Simon et al., 1993, Lijnen and Collen, 1995). Eventually, the haematoma is replaced
with granulation tissue in a controlled pattern. Approximately 12 days after injury a
completely vascularised collagen matrix is formed.
For the majority of severe fractures, there will be a degree of motion during
healing. In this situation, injured bone heals through an indirect (secondary) pathway in
which both intramembranous and endochondral ossification occur (Hench, 2005b,
Phillips, 2005). Endochondral ossification takes place at the centre of the collagen matrix
where mesenchymal stem cells proliferate vigorously and differentiate into chondrocytes
(Figure 2-8). Chondrocytes actively lay down cartilaginous matrix, thus transforming the
matrix to a fibrocartiliaginous callus (Marieb, 2001b). Gradually, the callus is deposited
with osteoid secreted by osteoblasts and calcified with calcium hydroxyapatite. When
mineralization proceeds, the hypertrophic chondrocytes within the callus undergo
apoptosis and a bone callus is formed at the fracture site. Upon union, the callus is
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known as woven bone (Phillips, 2005, Harwood et al., 2010a). Simultaneously,
intramembranous ossification forms a bone callus without a cartilage scaffold at the
periphery of the collagen matrix (Marieb, 2001b, Tsiridis et al., 2006). Mesenchymal
stem cells differentiate to osteoprogenitor cells, which later form the periosteum together
with surrounding fibroblasts. Some osteoprogenitor cells further differentiate into
osteoblasts to synthesise type I collagen of osteoid, resulting in a direct generation of
calcified tissue at the outer surface of fracture (Tosounidis et al., 2009, Harwood et al.,
2010a).
Figure 2-8. Proliferative phase of bone healing. It involves endochondral (at central of matrix through cartilage formation) and intramembranous (at periphery of matrix) ossification leading to woven bone formation from the vascularised granulation tissue. Modified from www.uwo.capatholcasesSkeletalfracture.html.
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2.2.4 Remodelling
The final stage of bone healing is to remodel bone callus into normal lamellar
bone (Hench, 2005b). Collagen fibres are in random orientation in the newly formed
woven bone. This disorganises the calcified components on the fibres, resulting in
reduced mechanical properties of the woven bone (Jahagirdar and Scammell, 2009).
Over a period of months, the collagen fibres are aligned to conform as closely as
possible to the original tissue by dual actions of collagen degradation and synthesis
(Doblare et al., 2004, Phillips, 2005). Together with the coupled actions of osteoblasts
and osteoclasts in bone deposition and resorption respectively, the woven bone is
remodelled into lamellar bone with highly ordered micro-architecture, thereby restoring
the normal mechanical properties (Tsiridis et al., 2006, Jahagirdar and Scammell, 2009)
(Figure 2-9).
Figure 2-9. Remodelling phase of bone healing. It involves collagen alignment and transformation of woven bone into lamella bone as origin. Modified from www.uwo.capatholcasesSkeletalfracture.html.
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To conclude, the bone healing is a highly complex process influenced by
biological and mechanical factors. Despite this, the underlying mechanism is largely
driven by the early microenvironment of which a haematoma forms the critical
component. This suggests that modifying a haematoma might be a primary target for
enhancing bone healing in severe defects. Figure 2-10 summarises the sequence of bone
healing events.
Figure 2-10. Time sequence of four main phases of bone healing: haemostasis (bleeding and blood clotting), inflammation, proliferation and remodelling.
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2.3 CURRENT THERAPEUTIC APPROACH
With an improved understanding of the biology and molecular aspects of bone
healing after injuries, it is generally accepted that the interruption of healing sequences in
non-union is caused by deficiencies of osteogenic cells, growth factors, and a conductive
scaffold at the damaged site (Dimitriou et al., 2005, Tsiridis et al., 2006). Current
therapeutic strategy aims to expedite the natural healing mechanism by providing the
deficient factors to the bone defects. However, treatment options have not been
improved greatly for the past decades.
2.3.3 Natural Bone Grafts
Autograft remains the gold standard for treating CSD. Autograft is harvested
from local bone and implanted in the same patient (Ryzewicz et al., 2009, Cove and
Keenan, 2009). It has been proven clinically efficacious and widely used in
orthopaedics, dentistry, oral and maxillofacial surgeries (Sajjadian et al., 2010). One
important reason is that autograft contains living cells and tissue-inducing substances
with its scaffold, hence imparting the physical structure and supporting osteoconduction,
osteoinduction and osteogenesis (Bauer and Muschler, 2000, Carlsson et al., 2004,
Precheur, 2007). However, the use of autograft is complicated by graft availability,
donor site morbidity and post-operative pain (Goulet et al., 1997, Bimmel and Govaers,
2006, Kolomvos et al., 2010). Despite allograft and xenograft harvested from human
donors or animals respectively overcomes the drawbacks of autograft, their applications
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are also limited by immune rejection, risk of virus transmission and decreased
biomechanical integrity (Wheeler and Enneking, 2005, Garofalo, 2007).
2.3.2 Bone Graft Substitutes
Due to the problems associated with the uses of natural bone grafts, bone
graft substitutes have been developed. Bone graft substitutes are either proteins
derived from organic bone (e.g. collagen & demineralised bone matrix) or inorganic
(synthetic) materials (Frayssinet, 2004, Giannoudis et al., 2005). Metallic implants
were the first “bone substitutes” to be used in orthopaedic history because of their
high mechanical strength (Precheur, 2007, Kao and Scott, 2007). In particular,
titanium has shown some superior osteointegrating properties (Hong et al., 1999,
2005, Le Guéhennec et al., 2007). Other synthetic materials such as ceramics (e.g.
hydroxyapatite (HA), tricalcium phosphate (TCP)) (Kandaswamy et al., 2000, Hing
et al., 2004), polymers (e.g. poly (lactic-co-glycolic acid)(PLGA)) and composites
are also introduced (Giannoudis et al., 2005, Deb, 2008, Chen et al., 2008). They
could be relatively biodegradable, osteoconductive, and have mechanical strength
similar to bone (McAuliffe, 2003, Ilan and Ladd, 2003).
However, each material has specific disadvantages. For instance fatigue and
stress shielding for non-degradable metals, and deficiency of osteoinductive property
for synthetic biomaterials (Sammarco and Chang, 2002, Hallab et al., 2004, Kao and
Scott, 2007). None of these synthetic materials can perfectly substitute for autograft
in current clinical practice.
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2.3.3 Bone Tissue Engineering
Bone tissue engineering is the latest approach to treat CSD. It combines the
technologies of biomaterial science, cell and molecular biology (Meyer et al., 2004).
The strategy is to confer osteoinductive capabilities on a bone substitute by
incorporating growth factors and/or pre-seeding reparative cells on the biomaterials
(Habibovic and de Groot, 2007, Dawson and Oreffo, 2008, Schliephake, 2009). It is
proposed that a biodegradable material not only serves as an osteoconductive
scaffold for growth, but also as a delivery vehicle of cells and signalling molecules to
accelerate bone healing (Stevens, 2008, Hutmacher et al., 2012).
Growth factor therapy is based on the molecular processes of pro-osteogenic
(e.g. bone morphogenetic protein (BMP)-2, -7, TGF-β) and angiogenic factors (e.g.
PDGF, FGF) in mediating bone formation (Wozney and Rosen, 1998, Lieberman et
al., 2002c, Carano and Filvaroff, 2003, Cross and Mustoe, 2003, Schmidmaier et al.,
2007, Bosetti et al., 2007). However, the degree of bone healing achieved is largely
dependent on the amount of growth factor released by the used scaffold. The
structural properties of the carrier have a dramatic effect on the release profile of
growth factors (Yang et al., 2008). A wide range of materials such as collagen,
hyaluronic acid, polylactic acid, injectable cement and even a biomimetic coating
(e.g. bone-like apatite (carbonated hydroxyapatite) layer) have been investigated to
achieve a controllable and sustained release of growth factor along with their
degradation rate, chemical and physical properties (Seeherman et al., 2006,
Kamitakahara et al., 2007, Liu et al., 2010). So far, design of scaffold architecture
remains a challenging topic in bone tissue engineering (Burg et al., 2000).
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Cell therapy works by grafting normal cells to restore cellular function in
the defects. In bone defects, a number of cell types have been engrafted on
biomaterials in vitro prior to transplantation, including osteoblasts, chondroblasts,
periosteal cells, bone marrow cells, stem cells and vascular cells (Puelacher et al.,
1996, Breitbart et al., 1998, Meyer et al., 2004, Kraus and Kirker-Head, 2006, Nather
et al., 2010). To further enhance the cellular function in producing growth factors,
gene therapy has been combined with this approach. Engrafted cells are transfected
in vitro by viral vectors which contain a gene encoding the expression of desired
growth factor (Kang et al., 2000, Dai et al., 2004, Betz et al., 2010). In addition,
implants might also be coated directly with viral vectors to transduce a patient’s own
cells in vivo (Awad et al., 2007). Recent work by Lin et al. (2010) demonstrated that
the bone marrow stem cells modified to express BMP-2 and vascular endothelial
growth factor (VEGF) accelerated the repair of a CSD in a rabbit femur. Although
these cells execute distinct functions in the bone healing, the cell transplantation has
not been widely applied in clinical practice due to several issues.
Primarily, this approach requires substantial labour and time to isolate and
expand the cells in vitro prior to in vivo application (Avramoglou et al., 2005). While
stem cells derived from periosteum and bone marrow can be expanded readily in
vitro due to their unique self-renewal property, it remains a problem to amplify the
differentiated cells to sufficient populations to rebuild bone mass without losing
viability (Pountos et al., 2007, Shanti et al., 2007). It was also shown that
osteoprogenitor cells only represent approximately 0.001% of the nucleated cells in
healthy adult marrow, making this approach least practical to aged and injured
groups for whom it is most needed. Moreover, concerns of immune rejection, ethical
issues and risk of uncontrolled gene expression also impede the uses of allogeneic
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cells, embryonic stem cells or viral vector transduced cells (Baltzer and Lieberman,
2004).
Together, bone tissue engineering has shown some potential in regenerating
bone. However, there are challenges and limitations in this technique before it can be
routinely translated to clinical situations.
(1) The performance of bioengineered constructs are greatly dependent on
the architecture and property of the biomaterial used (Burg et al., 2000, Blom, 2007,
Stevens, 2008). Slow or incomplete vascularisation of the engineered constructs after
implantation often occurs, causing hypoxia and death of cells seeded deep within the
scaffold and dysfunction of the tissue construct (Avramoglou et al., 2005, Laschke et
al., 2006). This suggests that the scaffold design requires further modification to
achieve rapid and adequate blood vessel ingrowth to include the engraftment of
endothelial cells as well as the release of angiogenic factors. Also, it indicates that
the angiogenic potential of biomaterials with different chemical compounds remain
largely unclear (Burg et al., 2000, Eckhaus et al., 2008, Butler and Sefton, 2007).
Generally, it is suggested that an ideal bone graft substitute for tissue engineering
should have the following characteristics (Giannoudis et al., 2005, Hutmacher and
Williams, 2006, Anderson et al., 2008, Thevenot et al., 2008, Amini et al., 2011):
1. Biocompatible (i.e. defined as the ability of a biomaterial to perform with an
appropriate host response in a specific application without eliciting other
adverse effects) (Williams, 2008).
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2. Three-dimensional and highly interconnected porous network (i.e. for tissue
growth and efficient flow transport of nutrients and metabolic waste).
3. Biodegradable or bioadsorptable at an appropriate rate (i.e. to facilitate the
controlled release of growth factors and match the rate of tissue ingrowth).
4. Possess mechanical properties (i.e. to match those of the tissue at implant
site and support before the regenerated tissue bears an increasing load while
the scaffold is degrading).
5. Suitable surface chemistry for cell adhesion, proliferation and
differentiation.
(2) On the other hand, one limitation of bone tissue engineering is that the
strategy often focuses on replacing a single cell type (differentiated or pluripotent),
or one or more growth factors to enhance bone regeneration, which in reality has a
multi-factorial mechanism. One of the typical examples is that a chondrocyte-seeded
scaffold remains in a cartilaginous stage after implantation into the bone defect, with
no observation of endochondral ossification (Vacanti et al., 1995). This finding
suggests that the healing sequence is important to direct cellular response for bone
formation. It is conceivable that the formation of a haematoma-like clot structure at
the implant site would have a beneficial effect on eliciting bone healing. This idea is
confirmed by studies using platelet-rich plasma.
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2.4 PLATELET-RICH PLASMA
Platelet-rich plasma (PRP) is a fraction of plasma in which platelets are
concentrated in a small volume of plasma (Lozada et al., 2001, Dugrillon et al.,
2002). The rationale behind the use of PRP is to provide autologous platelets, which
secrete their storage pool of growth factors at high concentration to expedite bone
regeneration and soft tissue repair (Anitua et al., 2004, Cenni et al., 2010).
2.4.1 Beneficial Role of Platelets & Uses of PRP
Indeed, activated platelets release a range of osteogenic and angiogenic
growth factors from their alpha (α)-granules, including PDGF, TGF-β, platelet-
derived epidermal growth factor (PDEGF), platelet-derived angiogenesis factor
(PDAF), insulin-like growth factor-I (IGF-I) and platelet factor 4 (PF-4) (Sánchez et
al., 2003, Frechette et al., 2005, Blair and Flaumenhaft, 2009) (Table 2-1). These
growth factors are known to have positive effects on bone healing by stimulating the
proliferation and differentiation of undifferentiated mesenchymal cells and
osteoblasts (Lieberman et al., 2002c, Arpornmaeklong et al., 2003, Bosetti et al.,
2007), angiogenesis as well as chemotaxis for inflammatory cells (Oprea et al., 2003,
Frechette et al., 2005). Primarily, it is believed that the initiation of bone regeneration
begins with the releases of PDGF and TGF-β after platelet aggregation (Kells et al.,
1995, Lieberman et al., 2002b).
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Growth Factor Actions
Platelet-derived growth factor (PDGF)
Chemotaxis and activation: neutrophils, macrophages Chemotaxis, mitogenesis and activation: fibroblasts (collagen synthesis), endothelial cells & smooth muscle cells (angiogenesis) Mitogenesis: mesenchymal and bone marrow stem cells
Transforming growth factor-beta (TGF-β)
Chemotaxis and activation: monocytes, endothelial cells (angiogenesis), preosteoblasts Inhibition: osteoclasts (bone resorption) Stimulation: osteoblasts(osteoid formation, osteogenesis) Regulation of mitogenesis: endothelial cells & fibroblasts (collagen synthesis)
Platelet-derived epidermal growth factor (PDEGF)
Chemotaxis and angiogenesis Mitogenesis: epithelial cells & mesenchymal cells Regulation of collagen synthesis
Insulin-like growth factor-1 (IGF-1)
Stimulation: cartilage & bone matrix formation Mitogenesis: preosteoblasts & osteoblasts Enhancement with PDGF: increases rate & quality of wound healing
Platelet-derived angiogenesis factor (PDAF)
Mitogenesis: endothelial cells Angiogenesis and enhance vascular permeability Upregulated by: IGF-1, TGF-α, -β, PDEGF, FGF & IL-1β
Platelet factor 4 (PF-4)
Chemotaxis: neutrophils, fibroblasts Inhibition: heparin
Table 2-1. Growth factors of platelets and their functions (Sánchez et al., 2003, Frechette et al., 2005).
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Previous studies have shown that growth factor concentrations correlate
directly with platelet number in PRP (Dugrillon et al., 2002, Eppley et al., 2004).
After dual gradient density centrifugation of whole blood to obtain PRP, platelets are
activated by the addition of thrombin and calcium chloride, leading to platelet
degranulation and formation of a PRP gel (Lozada et al., 2001, Grageda, 2004). The
application of PRP offers theoretical advantages over the delivery of a single
recombinant growth factor since PRP releases high concentrations of multiple native
growth factors in their biological ratios. The complex and interdependent nature of
growth factors (i.e. TGF-β, PDAF and IGF-1) suggests that more than one signalling
pathway of bone regeneration could be targeted with the use of PRP (Kells et al.,
1995, Sánchez et al., 2003). Also, when PRP gel is combined with particulate grafts,
better handing characteristics and in vivo stability could be achieved (Roukis et al.,
2006, Mooren et al., 2007).
In vitro studies have shown that PRP supernatants support the viability and
proliferation of human fetal osteoblast-like cells (Slater et al., 1995), alveolar bone
cells (Choi et al., 2005), porcine articular chondrocytes (Akeda et al., 2006) and
human endothelial cells (Roussy et al., 2007, Cenni et al., 2009). Extensive animal
studies have also investigated the effect of PRP gel alone on bone regeneration
(Weibrich et al., 2004, Swennen et al., 2004, Gandhi et al., 2006, Roussy et al.,
2007), or in combination with bone grafts and graft substitutes (Marx et al., 1998a,
Wiltfang et al., 2004, Choi et al., 2004, Mooren et al., 2007, Kasten et al., 2008,
Messora et al., 2008).
However, there is some inconsistency in the literature regarding the benefits
of PRP. While some studies reported significant increases in bone formation and
maturation rates (Marx et al., 1998a, Wiltfang et al., 2004, Gandhi et al., 2006,
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Roussy et al., 2007), others did not observe any improvement or even inhibition of
new bone formation (Choi et al., 2004, Swennen et al., 2004, Weibrich et al., 2004,
Mooren et al., 2007).
One of the possible reasons is the differential use of platelet concentrations
among studies. A typical PRP is defined to have a 5-fold increase platelet
concentration (approximately 1,000,000/µl) over the physiological level (Marx,
2001). Platelet concentration varies greatly due to different baseline values of animal
species and the preparing procedures of PRP (Dolder et al., 2006, Roukis et al.,
2006). In fact, the platelet concentration required for a positive effect on bone
regeneration seems to span a very limited range. Weibrich et al. (2004) reported that
advantageous effects of PRP on peri-implant bone regeneration in rabbits only
occurred when a platelet concentration of approximately 1,000,000/µl was used. At
lower concentration (164,000-373,000/µl), the effect was suboptimal whereas higher
concentration (1,845,000-3,200,000/µl) led to a paradoxically inhibitory effect. This
finding was supported by in vitro work of Choi et al. (2005) and Tomoyasu et al.
(2007) who studied the effect of platelet concentration in PRP alone on human
alveolar cells, and in combination with BMPs on human osteoblasts respectively.
Different protocols for platelet activation may be another reason for the
discrepancy of results (Walkowiak et al., 2000, Lozada et al., 2001). Concentrations
of thrombin and calcium for platelet activation were shown to affect the release of
growth factors, endothelial cell division (Lacoste et al., 2003, Frechette et al., 2005,
Roussy et al., 2007), and the adhesive property of PRP clots on soft tissue (Gimeno
et al., 2006). However, the mechanism of how these activators vary the properties of
the PRP clot is not fully understood. This may be associated with the thrombin
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concentration, which alters platelet activation and fibrin polymerisation, leading to
different kinetics of growth factor release and clot structure (Martineau et al., 2004).
2.4.2 Role of Thrombin Concentration in Clot Structure
The thrombin concentration present at the time of blood clotting has been
shown to profoundly influence the fibrin architecture than fibrinogen concentration, pH
and ionic strength (Nair and Dhall, 1991, Carr et al., 2002b, Wolberg, 2007b). Clots
formed at low thrombin concentration (< 1 nM) are composed of thick fibrin fibres in a
loose configuration. While those formed at high thrombin concentration are composed of
thin fibres in a tight configuration (Figure 2-11) (Wolberg and Campbell, 2008). Using
turbidimetric analysis of plasma, it has been shown that the altered thrombin
concentration contributes to different clot structure through fibrin polymerisation
process. An increase in thrombin concentration leads to a shorter time required for
protofibrils to grow to a sufficient length before they aggregate. It also causes an increase
in maximum rate of turbidity development and a decrease in the maximum final
turbidity, indicating a faster fibrin formation and a decrease in fibrin size (Standeven et
al., 2005).
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Figure 2-11. The thrombin concentration present at the time of gelation dictates the fibre thickness and density. a) Scanning electron micrographs of fibrin clots formed by adding thrombin (0.5-20 nM) to solutions of purified fibrinogen (2 mg/mL). The scale bars indicate 20 μm (top row), and 1 μm (bottom row) Adapted from (Wolberg, 2007a). Laser confocal micrographs of fibrin clots formed by adding thrombin b) 2.5 nM, c) 10 nM to solutions of purified fibrinogen (1 mg/mL). Laser confocal micrographs at 63 x (scale bar 10 µm). Adapted from (Wolberg and Campbell, 2008).
2.4.3 Effect of Clot Structure on Fibrinolysis
Alterations in clot structure have been demonstrated to affect the clot
susceptibility to fibrinolysis. In fibrinolysis, fibrin fibres are digested by plasmin which
is produced from cleavage of inactive plasminogen by tissue type plasminogen activator
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(tPA) (Figure 2-12) (Medved and Nieuwenhuizen, 2003). Fibrin fibres are shown to be
transverse cut across, rather than by progressive cleavage uniformly around the fibres
(Veklich et al., 1998). A study by Collet et al. (2000) showed that while individual thick
fibres are lysed more slowly than thin fibres, clots with a loose conformation of thick
fibres were lysed more rapidly than those with tightly-packed thin fibres. These findings,
as confirmed by later study of Bhasin et al. (2008) indicate that the network
conformation is a more important determinant for fibrinolysis rate compared with fibre
thickness. The changes in network conformation are believed to regulate the lysis rate by
influencing the fibrin density, tPA bindings on fibrin and transport of fibrinolytic
components throughout the clot (van Gelder et al., 1995, Collet et al., 2000, Medved and
Nieuwenhuizen, 2003, He et al., 2005, Undas et al., 2006).
Figure 2-12. Plasmin-mediated fibrinolysis. Inactive plasminogen is converted to active plasmin by tissue type plasminogen activator (tPA). Plasmin digests fibrin and generates fibrin degradation products (FDPs): D-dimers and E fragments. Antifibrinolysis system which includes proteins such as plasminogen activator inhibitor (PAI) 1 & 2, α2-antiplasmin and thrombin-activatable fibrinolysis inhibitor (TAFI), inhibits the fibrinolysis at different steps.
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2.4.4 Effect of Clot Structure on Viscoelastic Properties
Individual fibrin fibre has been illustrated to possess extensibility and elasticity
(Liu et al., 2006). The effect of fibre thickness on the properties of individual fibres is
not known yet. Instead, it has been demonstrated that altered fibrin structure may
modulate the clot rigidity as a whole, depending on fibre thickness, length, density,
degree of branching and cross-linking (Standeven et al., 2005, Wolberg and Campbell,
2008). In particular, cross-links that reinforce fibrin contacts within the clot increase the
elasticity of individual fibres and the overall clot elasticity (Collet et al., 2005, Liu et al.,
2006, Jámbor et al., 2009).
2.4.5 In Vivo Implications of Altered Clot Structure, Properties and Stability
A number of studies have shown that an altered clot structure is a causative
mechanism of many thrombotic diseases and bleeding disorders (Collet et al., 1993b,
Mills et al., 2002, Terasawa et al., 2006, Bhasin et al., 2008, Undas et al., 2008, 2009).
Patients with acute ischemic stroke are found to produce in vitro plasma clots that are
denser with thicker fibres and more resistant to fibrinolysis compared to controls (Undas
et al., 2010). In contrast, haemophilia A patients who are deficient of factor VIII, are
shown to produce clots that are much more porous with thicker fibres, and overly more
susceptible to fibrinolysis (Wolberg, 2007a). These findings suggest that the abnormal
clot structure and susceptibility to fibrinolysis are principally related to pathological
conditions.
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In addition to the genetic factors, environmental determinants of fibrin structure
are also of clinical importance (Scott et al., 2004). Changes in thrombin concentration
may indirectly modulate clot architecture by activating factor XIII to cross-link adhesive
proteins to fibrin. Binding of fibronectin to fibrin is essential for cell adhesion and
migration into the clots (Okada et al., 1985, Carr et al., 1987, Weisel, 2005). While
cross-linked collagen also supports the formation of extracellular matrix at the injury
sites, aiding the healing process. Moreover, fibrin-bound actin, myosin and vinculin
together with platelet cytoskeleton have been shown to mediate clot retraction and
subsequent wound narrowing (Asijee et al., 1988, Mosesson, 2005, Weisel, 2005,
Wolberg, 2010). Hence, it possibly explains the changes in growth factor release, cell
division and adhesive properties of PRP clots as previously reported by studies using
thrombin at different concentrations. Moreover, these findings offer an insight into the
pivotal role of a blood clot and its structural properties, which will have a direct effect on
the bone healing by influencing macromolecule transport, cell behaviour and new tissue
ingrowth.
A dental implant inserted in the jaw bone is a typical example of endosseous
(in-bone) implants in which its clinical success is greatly influenced by a blood clot
formed around the implant. During implantation, tissue damage and bleeding are
inevitable. The first tissue that comes into contact with the implant is blood (Anderson,
2001). Subsequently, a blood clot is formed at the gap between host bone and implant, as
a result of coagulation activation. The blood clot not only detains blood flow and anchors
the implant to the endosseous wound site, but most importantly supports two types of
peri-implant endosseous healing: distance and contact osteogenesis (Davies, 2003b).
Distance osteogenesis occurs when new bone is initially formed on the surfaces of
surrounding old bone at a distance from the implant. Conversely, contact osteogenesis
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takes place when new bone is first formed on the implant surface. Hence, it is clinically
considered as a superior mode of healing in case there is insufficient cortex to provide
early stability (Figure 2-13).
Figure 2-13. Two distinct patterns of peri-implant endosseous healing: a) distance osteogenesis and b) contact osteogenesis where osteoblasts line initially on the host bone or implant surface respectively and lay down matrix. In distance osteogenesis, the bone surfaces provides a population of osteogenic cells which differentiate into osteoblasts and lay down a new matrix that grows slowly towards the implant and results in the approximation of the implant surface shape by the newly formed bone. An intervening layer of non-bone cells is often formed between the bone and implant. Contact osteogenesis occurs with osteoconduction and de novo bone formation, which predominantly require recruitment and migration of osteogenic cells towards the implant surfaces.
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Clearly, the prerequisite of contact osteogenesis is the continuous
recruitment and migration of osteogenic cells to the implant surface through the
three-dimensional blood clot. The osteoconduction is largely stimulated by activated
platelets and leukocytes entrapped in the compartment. They release a range of
cytokines and growth factors, creating chemoattractant gradients to recruit
undifferentiated or osteogenic cells to the implant site (Oprea et al., 2003, Park and
Barbul, 2004, Dohan Ehrenfest et al., 2006). On the other hand, the three-
dimensional network of fibrin as well as structural proteins of the clot serve as
physical scaffolds to support cell adhesion and migration (Choukroun et al., 2006).
To migrate to the implant, the osteogenic cells impose contractile forces on the fibrin
fibres that attached to the implant surface, where they differentiate into osteoblasts
and directly lay down bone matrix. Following mineralisation, a collagen-free cement
line appears and results in de novo bone formation (Davies, 2003b). Thus, the fibrin
architecture of the clot is important for effective fibrin retention of implant surface
and critically determines the process of contact osteogenesis (Choukroun et al., 2006,
Liu et al., 2006).
Overall, the presence of a blood clot with appropriate clot structural
properties ensures the bone-implant interface environment to support the bone
healing.
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2.4.6 Differences Between a PRP gel and a Haematoma
While PRP gels are mostly formed by adding a fixed and high amount of
thrombin to platelet-concentrated plasma, it is conceivable that a PRP gel would be a
platelet clump tightly networked by thin fibres. These abnormal clot structures and
properties, as well as their impacts on growth factor release, cell proliferation and
physical stability are likely attributed to the negative effect on bone healing as previously
reported. The confusion between the differences in cellular components and structures
between a PRP gel and a haematoma, and how these differences relate to their potentials
in enhancing bone healing may be the source of inconsistent results.
A haematoma contains mostly erythrocytes, approximately 5% of platelets and
less than 1% of leukocytes (Marieb, 2001b, Carlson and Roach, 2002). However, a PRP
gel which contains theoretically no other blood cells but platelets, possessing nearly a
reverse ratio of erythrocytes and platelets compared to a haematoma. In fact, platelets
and circulating blood cells, both their number and interactions play an important role in
clot features. It has long been proposed by Ulevitch and Johnston (1980) that
erythrocytes participate in the intrinsic pathway of coagulation, which is believed to be
associated with the negatively charged phospholipids of the cells (Peyrou et al., 1999,
Iwata et al., 2004). Also, it has been demonstrated that erythrocytes interact with
platelets by promoting platelet aggregation, and inversely activated platelets also
enhance erythrocyte agglomeration (Alkhamis et al., 1988, Valles et al., 1991). In
addition, leukocytes have been suggested to influence coagulation by expressing TF,
activating platelets and factor X (Plescia and Altieri, 1996, Bouchard and Tracy, 2001,
Gorbet and Sefton, 2004, Elalamy et al., 2007), as well as cleavage of tissue factor
pathway inhibitor (TFPI) (Petersen et al., 1992, Sundaram et al., 1996). All these studies
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are consistent with the findings of Thor et al. (2007), where whole blood induced higher
levels of thrombin generation and platelet activation than PRP on clinically used
titanium, and that the response of PRP could be partially restored in the presence of
erythrocytes (Hong et al., 2001, Horne et al., 2006).
Additionally, it has been shown that cells influence local fibrin structure by
direct interaction between cell surface integrins and fibrin. Platelets have been shown to
organise fibrin into tighter bundles near its cell surface than seen more distally within the
clot (Collet et al., 2002, Campbell et al., 2009). In contrast, the presence of erythrocytes
has been shown to form a more porous fibrin network, facilitating migration of cells into
the area and thus support wound healing (Carr and Hardin, 1987).
Furthermore, cell-associated fibrin is revealed to be more resistant to
fibrinolysis than distally located fibrin (He et al., 2005, Jerome et al., 2005, Campbell et
al., 2008). This might be due to soluble proteins released from the cells which can
regulate the equilibrium between clot formation and dissolution. FXIII and PAI-1,
released from platelets, are known to increase the resistance of the clot to fibrinolysis
(Korbut and Gryglewski, 1995, Handt et al., 1996, Carrieri et al., 2011). On the contrary,
haemoglobin as well as neutrophil elastase and cathepsin G, are shown to enhance
fibrinolysis (Bach-Gansmo et al., 1998, Yoshida et al., 2001). These studies agree that
during in vivo myocardic infarction, normal erythrocytes-rich clots are readily dissolved
by enzymatic lysis whereas platelet-rich clots are more resistant to be degraded (Jang et
al., 1989, Parise and Agnelli, 1991, Gold et al., 1991).
So far, the innovative uses of PRP in bone tissue engineering focus solely on
the biological value of platelet growth factors. There are conflicting results of PRP with
bone healing and its utility remains unsolved. Recently, a second-generation of platelet
concentrates has also been introduced based on leukocyte content and fibrin architecture
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including: pure platelet-rich plasma (P-PPP), leukocyte- and platelet-rich plasma (L-
PRP), pure platelet-rich fibrin (P-PRF), leukocyte- and platelet-rich fibrin (L-PRF)
(Anitua et al., 2006, Choukroun et al., 2006, Dohan Ehrenfest et al., 2006, Baeyens et al.,
2010, Simonpieri et al., 2012). Little is known about how these two parameters
influences the intrinsic biology of these products (Dohan Ehrenfest et al., 2010, 2012,
Simonpieri et al., 2011), and not to mention other cellular components in the clots are too
often neglected.
Taken together, a PRP gel is different from a natural haematoma in both
cellular components and structure (Figure 2-14). It is likely that these differences
contribute to different molecular and cellular activities as well as mechanical stability at
the injured bone, ultimately dictating the outcome of peri-implant bone healing.
Figure 2-14. Schematic pictures illustrating differences in cellular components and fibrin scaffold between a) platelet-rich plasma (PRP) gel, and b) a normal haematoma.
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2.5 IN SITU BLOOD CLOT FORMATION & MODIFICATION - A
POSSIBLE TREATMENT FOR BONE DEFECTS
Based on in vivo potentials in influencing bone healing, it seems that blood
clots could be classified as “ good” and “ bad”. The fact that the composition of a
blood clot is specific to the circumstances under which it formed opens up a possible
approach to make a “ good ” clot on artificial bone implant as a therapeutic agent for
treating bone defects. A central question is how blood interacts with biomaterials
after implantation.
2.5.1 Blood and Host Response towards Biomaterial Implants
As noted, surgical procedures of implantation induce bleeding which in turn
triggers haemostasis and acute inflammation immediately (Anderson, 2001, Gorbet
and Sefton, 2004). Blood contact with implants leads to rapid adsorption of proteins
on the implant surface. Within seconds, a layer of proteins is formed as a result of
dynamic collision, competitive adsorption and displacement in a process known as
the Vroman effect (Horbett, 1993). This protein deposition on biomaterial surface is
termed as provisional matrix formation (Anderson, 2001). Cells that migrate later
into the site interact with the adsorbed proteins, rather than directly with the material
surface itself. Hence, it is widely believed that the nature of the absorbed proteins
determines subsequent responses of coagulation, platelets, leukocytes, as well as the
immune complement reactions to the implant (Figure 2-15) (Eskin et al., 2004).
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Figure 2-15. Activation of complement system. As part of the innate immune system, complement is a blood cascade for the recognition and clearance of foreign materials. Similar to coagulation system, complement system consists of over 20 proteins and are activated upon cleavage by an upstream enzyme. It can be activated through classical, alternative, and lectin pathways. While classical pathway is initiated by C1q binding on foreign surface, the alternative and lectin pathways are triggered by spontaneous hydrolysis of C3 on foreign surface and mannose-binding lectin, respectively. All pathways converge at the formation of C3 convertase, which cleaves C3 to fragments C3a and C3b. C3a acts as an anaphylatoxin, producing local inflammatory response such as induction of smooth muscle contraction, increase of vascular permeability and chemotaxis. While C3b complexes with C3 convertase and become C5 convertase, which cleaves complement molecule C5 to C5a and C5b. C5a acts as the most potent anaphylatoxin whereas C5b participates in the formation of complement C5b-9 complex (also known as membrane attack complex (MAC) or the terminal complement complex (TCC)). Ultimately, incorporation of C5b-9 into cell membrane of foreign cells achieve defending effects of complement system of destruction of foreign particles by alternating membrane polarization to cause cell rupture, and facilitating recognition and opsonisation by phagocytes.
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Earlier in vitro studies noted that the adsorption of FXII, HMWK and
prekallikrein on biomaterial surfaces were accompanied with thrombin formation,
suggesting biomaterial-induced coagulation may be initiated by intrinsic pathway
(Mulzer and Brash, 1989, Ziats et al., 1990). Also, the adsorption of fibrinogen,
fibronectin and vWF are known to promote platelet adhesion and activation on
biomaterials (Keuren et al., 2002, Tsai et al., 2002, Kwak et al., 2005). On the other
hand, it is believed that immunglobulin G and complement fragment C3b deposited
on surfaces not only mediate surface-activated complement by classical and
alternative pathways, but also act as opsonins to modulate adhesion and phagocytosis
of neutrophils and macrophages (Gemmell, 1997, Wilson et al., 2005, Sellborn et al.,
2005, Nilsson et al., 2007a).
In acute inflammation, macrophages are beneficial as they replenish
cytokine and growth factors at the implant site, leading to the formation of a
granulation tissue on the implant (Anderson, 2001). However, in some situations,
persistent inflammatory stimuli, chemical and physical properties of the biomaterial
may lead to chronic inflammation that lasts the entire lifetime of the implant (Tsai,
2004, Gorbet and Sefton, 2004, Jones, 2008). Adherent macrophages on implant
surfaces cannot engulf the implant due to size disparity, and therefore experience
“frustrated phagocytosis” (Nilsson et al., 2007b, Anderson et al., 2008). Unlike
normal phagocytosis, they fuse to form multinucleated foreign body giant cells
(FBGCs) surrounding the implant (Figure 2-16). Moreover, the macrophages may
also respond by respiratory burst and protease release, leading to the deterioration of
the implant and injury to peripheral tissue (Janeway et al., 2005b). Thus, the FBGCs
together with granulation tissue in the presence of implant is referred to as foreign
body reaction (FBR) (Tang and Eaton, 1995, Hu et al., 2001, Anderson et al., 2008).
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Eventually, the host response to implants leads to fibrosis or fibrous
encapsulation. Fibrous capsule and interfacial FBR become a physiochemical barrier
which isolates the implant from the surrounding tissues, severely impairing implant
function and in vivo integration (Anderson et al., 2008, McNally and Anderson,
2011).
In summary, the mechanisms governing the biomaterial-blood interactions,
and the possible role of biomaterial surfaces affecting such responses are crucial for
the design of biocompatible endosseous implants.
Figure 2-16. A multinucleated foreign body giant cell. Adapted from http://www.manana tomy.com/basic-anatomy/reticuloendothelial-system-macrophage-system
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2.5.2 Influence of Biomaterial Surface Chemistry on Blood Response
To minimise adverse responses and ensure long term function of implants,
many research efforts have been placed on modifying the surface chemistry of
biomaterials.
The surface chemistry of titanium which offers thrombogenicity to the
endosseous implant has long been suggested to attribute to its superior
osteointegration (Hong et al., 1999). Upon air exposure, a titanium oxide layer is
formed naturally on the surfaces. It displays negative charges that closely resemble
extracellular matrix protein of damaged tissue, therefore resulting in enhanced
attachment and activation of FXII and platelets on the surface. Thrombin generation
was also greatly elevated by titanium when compared to stainless steel, which is
known to have a poorer osteointegration property (Nygren et al., 1997, Broberg et
al., 2002, Hong et al., 2005). Furthermore, using a rat subcutaneous pocket model, it
was demonstrated that the extent of neovascularisation in healing process was
modulated by pre-coating titanium with a thin plasma clot (100 nm) (Jansson et al.,
2001). Hence, these findings suggested that an endosseous implant might generate a
positive osteogenic response through its surface chemistry that modulates blood
cascade activation and the formation of a blood clot.
To determine more specifically what species and density of surface
chemical functionalities affect blood protein adsorption and subsequent cellular
interactions, alkythiols or alkysilanes self-assembled monolayers (SAMs) are widely
used as a flat surface model to presents a single or multiple functional groups (Figure
2-17).
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Figure 2-17. Self-assembled monolayers (SAMs). a) A simple unfunctionalized (methyl-terminated) alkanethiols monolayer on gold substrate. b) ω-functionalized alkanethiols on gold. c) Alkyl silane monolayer on oxidized silicon.
2.5.2.1 Carboxyl and Methyl Functional Groups
Surface chemical functionality has been demonstrated to influence both the
amount and structural changes of protein upon adsorption (Martins et al., 2003,
Ishizaki et al., 2007). SAMs bearing methyl (–CH3) groups were shown to bind more
fibrinogen and cause a greater conformational changes of adsorbed proteins than
carboxyl (–COOH), amino (–NH2) and hydroxyl (–OH) groups (Sit and Marchant,
1999a, Evans-Nguyen and Schoenfisch, 2005b, Roach et al., 2005, Sivaraman et al.,
2009, Sperling et al., 2009, Xu and Siedlecki, 2009, Fischer et al., 2010b).
Such conformational changes of adsorbed fibrinogen on –CH3 SAMs was
also evidenced to associate with stronger platelet adhesion and activation than –
COOH or –OH SAMs (Lin and Chuang, 2000, Sperling et al., 2005a, Rodrigues et
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al., 2006, Sivaraman and Latour, 2010). In addition, the work of Evans-Nguyen et
al.(2005b) indicated that surface functionalities affect the ability of adsorbed
fibrinogen to interact with thrombin in forming fibrin fibres. Hence, surface
functionality alters the adsorption characteristics of fibrinogen, its ability to support
platelet interaction and fibrin polymerisation on biomaterials.
With regards to surface-associated coagulation, –CH3 SAMs was shown to
induce a stronger overall activation than –COOH SAMs, despite the initiation of
intrinsic pathway being higher on –COOH SAMs. Interestingly, a study which
investigated –COOH and –CH3 mixed SAMs found an addition of –COOH groups
(50-83%) to –CH3 SAMs further elevated coagulation activation compared to –CH3
SAMs alone. Varying compositions of –COOH/–CH3 SAMs was also shown to
modulate platelet adhesion and activation (Sperling et al., 2005a, 2009, Fischer et al.,
2010b). This is consistent with the other studies suggesting that combining –CH3 and
other functional groups may further alter surface-mediated fibrinogen adsorption and
platelet responses (Rodrigues et al., 2006, Tsai et al., 2007).
An opposite trend in immune response was observed on the –COOH and –
CH3 mixed SAMs. Surfaces containing 47% –COOH/53% –CH3 resulted in a lower
level of complement activation followed by –CH3, –COOH and –OH SAMs
(Sperling et al., 2007). The stronger complement response to –OH or –NH2 SAMs is
believed to be mediated through an alternative pathway in which these functional
groups form a direct covalent thioester linkage of complement fragment C3b, leading
to the formations of C3 and C5 convertases. Increasing –OH content on mixed SAMs
was also found to increase complement activation (Tang et al., 1998, Hirata et al.,
2003, Sperling et al., 2005a). Salvador-Morales et al. (2009) who investigated lipid-
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polymer nanoparticles presenting –COOH/–CH3 groups also demonstrated an
increase in complement activation with increasing –COOH content.
Furthermore, it has been indicated that the complement C3 and
immunoglobulin G deposited on –OH SAMs are associated with the highest number
of leukocyte found on pure –OH3 SAMs than pure –CH3 SAMs which is accounted
for the lowest number of adherent cells in vitro and in vivo, suggesting a relationship
between complement and inflammation (Kalltorp et al., 1999, Barbosa et al., 2003,
2010). Unexpectedly, a noticeably enhanced leukocyte adhesion was seen on 83% –
COOH/17% –CH3 SAMs than either –COOH or –CH3 SAMs alone but no difference
in leukocyte activation was found in vitro (Sperling et al., 2009, Fischer et al., 2010b,
2010a). These studies indicate the surface functionalities modify specific leukocyte
response around implants.
2.5.2.2 Functional Groups from Copolymer Surfaces
Besides using SAMs, an alternative way to generate specific functionalities
on surfaces is by copolymerisation. It is a process of polymerising two monomers,
which contain functional groups. Polymeric materials have been widely generated for
tissue engineering applications.
In synthesis of poly (alkyl methacrylates), Berglin et al. (2004, 2009)
suggested that the chain length of alkyl group (i.e. number of carbons in the alkyl
side chain: 4, 6, 12, 18) had a major influence on the rate of thrombin generation
from coagulation, fibrin deposition and complement activation. A decrease in alkyl
chain length was shown to reduce the rates of thrombin generation and fibrin
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deposition, but increase complement activity. Thus, monomer alkyl methacrylate
with different alkyl chain length: methyl (–CH3), ethyl (–CH2CH3) or butyl (–
(CH2)3CH3)) may modulate activation of coagulation and complement system.
On the other hand, acrylic acid (AA) which contains –COOH group could
be an another desired monomer. A number of studies on graft polymerisation on
various polymers employed acrylic acid to develop functional interfaces for
immobilising biomolecules (Gupta et al., 2001, Alexander et al., 2004, Monien et al.,
2005, Huang and Jang, 2009). Most importantly, AA has been shown to interact with
methyl methacrylate and produce non-toxic copolymers (Yan and Gemeinhart,
2005). Hence, copolymerisation of acrylic acid and alkyl methacrylates may provide
a solid polymeric surface with tailored chemical functionality for a systematic study
in contrast to the model SAMs on gold substrates. However, very little has been
reported on modifying copoly (AA-co-alkyl methacrylate) compositions.
For copolymerisation of two monomers M1 and M2, the reactivity of a
monomer with one another (defined as relative reactivity ratio) and the monomer
concentrations are important. Four propagation reactions are possible with the
assumption that the propagation is dependent on the nature of the monomer at the
growing chain end (Odian, 2004) (Figure 2-18).
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Figure 2-18. Copolymerisation of two monomers M1 and M2. One propagating chain with M1 at the end and the other with M2. The asterisk represents the propagating species such as a radical. k11 and k22 are rate constants of self-propagation; k12 and k21 are rate constants of cross-propagation.
Monomer relative reactivity ratios r1 and r2 are determined as ratios of rate
constants (Ekpenyong, 1985):
The tendency of two monomers to copolymerise is determined by r values
while the type of copolymerisation are classified according to the products r1r2.
When both values of r1 and r2 are zero, it indicates that neither monomer is capable
of adding its own monomer, and propagation results in an alternating copolymer. If
both r1 and r2 values equal to one, it indicates that each monomer shows the same
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preference for one of the monomers, and the sequence of propagating chain is
completely random, resulting in ideal copolymerisation (r1 r2 = 1) and copolymer
composition is the same as the monomer feed (Odian, 2004). When either r values
greatly more than one, homopolymers or block copolymers are obtained.
During copolymerisation, monomers disappear by incorporating into
growing chains. Thus, the rates of disappearance of the two monomers are
synonymous with their rates of entry into the copolymer. The copolymer
composition is given by copolymerisation equation (Ekpenyong, 1985):
where [m1] and [m2] are molar concentrations of monomer M1 and M2 in
copolymer, respectively. [M1] and [M2] are their corresponding concentrations in the
monomer mixture prior to polymerisation.
Hence, different monomer relative ratios and monomer feed concentrations
influence the kinetic of copolymerisation and resultant copolymer composition. This
might result in a complex mixture of random or homopolymers in copolymers,
modulating the bulk and surface properties of the copolymers (Hermitte et al., 2004)
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Taken together, most studies on SAMs target to improve blood
compatibility (i.e. non-thrombogenic and non-immunogenic properties) of blood
contacting devices such as cardiovascular implants, catheters and extracorporeal
circulation. We believe that a blood clot formed on bone implants might constitute a
beneficial microenvironment in bone healing applications, which is currently
overlooked. An understanding of the bioactivities of surface functionalities informs a
rational strategy for developing “prothrombogenic” and non-immunogenic surfaces
on artificial bone implants. The combined use of the copolymer functionalised
surfaces (i.e. –COOH/–CH3, –CH2CH3 or –(CH2)3CH3) at different compositions
may provide a mean of synergistic modulation of the degree of coagulation and
related structure and properties of fibrin clot. In addition, diminishing the immune
response via reduced complement activation may also improve the bone healing
capability of artificial implants.
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3Chapter 3: Synthesis and Characterisation
of Material-Coated Surfaces
3.1 INTRODUCTION
Modification of surface properties is one of the approaches to enhance the
bone healing and in vivo integration of artificial bone substitutes (Ma et al., 2007).
Oral implant with surface roughness of about 1.5 µm Sa value 4 were shown to result
in a stronger bone ingrowth and osteointegration than smoother (< 0.5 µm) and
rougher ( > 2 µm) surfaces (Cooper, 2000, Shalabi et al., 2006, Le Guehennec et al.,
2007, Wennerberg and Albrektsson, 2009). An increase in surface hydrophobicity
was also found to decrease density and spreading of osteoblasts (Lin and Lin-Gibson,
2009, Wei et al., 2009).
Although these studies reported that surface properties affect osteoblast
behaviour and osteointegration, few studies have explored early blood/endosseous
implant interactions and resulting blood clot formation, even if these events occur
prior to bone formation (Hong et al., 2001, Thor et al., 2007).
The effect of surface chemistry on blood-biomaterial interactions has been
intensively studied in developing cardiovascular devices. SAMs displaying 47% –
COOH/53%–CH3 chemical functionalities were shown to increase platelet
interactions and coagulation activation leading to a strong fibrin fibre deposition,
4 Sa value is the arithmetic mean deviation of a surface
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with decreased leukocyte accumulation and complement initiation when compared
with pure –COOH and –CH3 SAMs (Sperling et al., 2005b). In addition, polymer
particles presenting –COOH/–CH3 groups at different ratios, and poly (alkyl
methacrylates) with various alkyl chain lengths were also found to modulate
complement activation, the rates of thrombin generation and fibrin deposition
(Berglin et al., 2004, 2009, Salvador-Morales et al., 2009). These findings highlight
the possible roles of –COOH/–CH3 (methyl), –CH2CH3 (ethyl) or –(CH2)3CH3
(butyl) in the design of a prothrombogenic and immunocompatible polymer surface
for synthetic bone implants.
Methacrylic and acrylic polymers have been long used in medical
applications. Of poly (alkyl methacrylates), poly (methyl methacrylate) (PMMA) is
widely used as a bone cement, dental filling and in intraocular lenses (Lee et al.,
2007). Polymers containing ethyl (EMA) or butyl methacrylate (BMA) have also
been shown to be capable of modulating chondrocyte and osteoblast attachment
(Hutcheon et al., 2001), and induce angiogenesis respectively (Butler and Sefton,
2007). Furthermore, Yan and Gemeinhart (2005) used copolymer microparticles
composed of methyl methacrylate (MMA) and acrylic acid (AA) as a drug delivery
system and confirmed that the copolymer was non-toxic both in vitro and in vivo.
Hence, AA which possess hydrophilic negatively charged –COOH groups provides
an opportunity to present the surface functional groups of our interest together with
hydrophobic non-ionic alkyl methacrylates at different ratios.
In this part of our study, we aim: (1) To synthesise a series of materials by
varying the monomer ratios of AA/MMA, EMA or BMA; (2) To coat materials on
the inner surface of a customised incubation chamber; and (3) To analyse the surface
compositions of functionalities –COOH/–CH3, –CH2CH3 or –(CH2)3CH3,
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hydrophobicity and roughness on coatings. Such knowledge would allow a priori
prediction of physiochemical or even possibly prothrombogenic and non-
immunogenic properties directly from the material formulation.
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3.2 MATERIALS
Acrylic acid (AA), methyl, ethyl and butyl methacrylates were all purchased
from Sigma-Aldrich (New South Wales, Australia) and used as received. Benzoyl
peroxide (BPO; Luperox®, Sigma-Aldrich, Australia), which contained 25% water,
was purified by dissolving in chloroform and recovered by precipitation in excess
methanol. The precipitated BPO was filtered and dried under vacuum before use.
3.3 METHODS
3.3.1 Synthesis of materials
Acrylic acid and alkyl methacrylate were reacted via free-radical
polymerisation using BPO as an initiator. Three types of alkyl methacrylates were
employed: methyl methacrylates (MMA), ethyl methacrylates (EMA) and butyl
methacrylates (BMA). The monomer solutions were added at molar ratios (AA: alkyl
methacrylate; 45, 55 or 65 %) in a glass vial containing 0.5% of BPO. The solutions
were deoxygenated by bubbling argon gas with a syringe through the septum caps on
the vials. The vials were then incubated in an oil bath with increasing temperatures
(45˚C, 55˚C, 65˚C, 75˚C, 85˚C) for 1 h intervals, and at 90 ˚C and 100˚C for 20 min
intervals. Figure 3-1 illustrates the free-radical polymerisation initiated by BPO and
the chemical structures of the comonomers: AA, MMA, EMA and BMA.
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Figure 3-1. Free-radical polymerisation is initiated by benzoyl peroxide, which generates free radicals by heat. Chemical structures of comonomers: acrylic acid, methyl, ethyl and butyl methacrylates were shown with the changes in alkyl chain length highlighted in red boxes.
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3.3.2 Preparation of surface coatings
The impact of the surface functionalities and their relative ratios on blood
responses was studied by incubation vials with surface coatings. The materials
prepared as described above were dissolved in acetone to give a 5 % (w/v) solution.
The solutions were coated on the internal surface of glass vials (4 mL, 15x45 mm;
Waters, Australia) by a solvent-evaporation technique (Fujishita et al., 2009).
Solutions of 1.5 mL were added to the vials and dried in an oven at 50 ˚C until
complete evaporation of acetone. The procedure was repeated three times to ensure
full coverage on the glass surfaces. After drying, the coated vials were capped at
room temperature until used.
For water contact angle measurement, glass coverslips (No. 1, diameter 13
mm, ProSciTech, Australia) were used to provide a flat surface. Coverslips were
placed at the bottom of glass vials before the addition of solution and drying step.
Uncoated glass vials or coverslips were used as controls.
3.3.3 Characterisation of surface
To analyse the surface properties of the coating, the coated vials or
coverslips were cut manually. Macroscopic contamination or dust on the coated
surfaces were removed by purging with argon. Samples were stored in clean sealed
containers until analysis.
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3.3.3.1 X ray photoelectron spectroscopy
The functionalization of coatings was analysed by X-ray photoelectron
spectroscopy (XPS) (Centre for Microscopy and Microanalysis, University of
Queensland). XPS is a surface analysis technique to determine the composition and
chemical state of elements that exist in a material. A spectrometer (Axis ULTRA Kratos
Analytical, Shimadzu, UK) equipped with a monochromatic Al Kα X-rays (1486.6eV)
was operated at 150 W and incident at 45 ⁰ to the sample surface. Photoelectrons emitted
from the surface were collected at a take-off angle of theta 90 ⁰ with a 165 mm
hemispherical electron energy analyser. Elements present in the surface were identified
by survey scans taken at pass energy of 160 eV and at resolution of 1.0 eV. These scans
recorded binding energies of the photoelectrons ranging from 0 - 1200 eV. Multiplex
scans at pass energy of 20 eV and at a higher resolution of 0.05 eV were also performed
to determine the chemical states of carbon atoms. The base pressure in the chamber was
1.0 x 10-8 torr during analysis.
The carbon 1s (C1s) high resolution spectra were processed to determine the
relative oxidation states of carbon atoms. Curve fitting of the spectra was performed
using the Casa XPS software (version 2.3.14) and a linear baseline with Kratos library
Relative Sensitivity Factors (RSFs). The binding energies that are indicated by peaks in
spectra were referenced to the C1s aliphatic carbon peak at 285.0 eV (Takemoto et al.,
2004). This corrected the effect of surface charging during analysis on shifting the peak
positions. Peak areas were normalised and adjusted to obtain a full width at half
maximum between 0.9 and 1.1 eV. By determining the area ratios of carbonyl (C-O)
component derived from methacrylate monomer, and carboxyl (O-C=O) component
derived from both types of comonomers, the compositions of corresponding surface
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functionalities were obtained. Copoly (AA-co-MMA) is shown as an example (Figure 3-
2 a) and the surface ratio of –COOH groups was calculated by using the equation shown
in Figure 3-2 b. Average data were collected from three measurements of each surface.
Figure 3-2. Copolymer surfaces displaying various functional groups. a) Schematic diagram of structure of copoly (AA-co-MMA) illustrates the relative ratios of carbon components C-O (in red box) and O-C=O (in blue circle) which were quantified to determine the compositions of comonomers and corresponding surface functional groups. b) Equation used to calculate the surface ratio of –COOH groups.
3.3.3.2 Water contact angle measurement
Surface hydrophobicity of coatings were assessed by measuring advancing
water contact angle using a FTÅ 200 system (First Ten Ångstroms, Poly-Instruments
Pty. Ltd., Australia) (Figure 3-3 a). Glass coverslips with surface coatings were
employed. After placing the samples on the stage of the goniometer in an environmental
chamber, a volume of degassed deionised water (MiliQ quality) was suspended from the
tip of a microliter syringe. Samples were gently lifted up until the surfaces made contact
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with the droplet. The needle was kept in contact with the droplet and continued
dispensing. Advancing contact angles were achieved when the volume of the droplet
increased to its maximum before spreading across the surfaces. Images were captured
with a horizontal CCD video camera and contact angles were measured from the images
using drop shape analysis software (Fta32 version 2.0) (Figure 3-3 b). Data were the
average of at least six regions of each surface.
Figure 3-3. Water contact angle measurement. a) FTÅ 200 water contact angle goniometer with a horizontal camera. b) The contact angle is the angle at which a liquid/vapour interface meets the solid surface. b) Adapted from http://superhydrophobiccoating.com
3.3.3.3 Scanning electron microscopy
Scanning electron microscopy (SEM) was used to examine the surface
morphology of coatings on glass vials. Specimens were mounted on aluminium stubs,
gold-coated and examined with a Quanta 200 scanning electron microscope (FEI, USA).
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Several fields on each surface were imaged and a representative field was chosen (at a
magnification of 25000 x).
3.3.3.4 Atomic force microscopy
The roughness of surface coating was measured by atomic force microscopy
(AFM) using a Solver P47 Pro scanning probe microscope (NT-MDT Co., Russia).
Sample surfaces were scanned in contact mode (constant force) with a golden silicon
probe (CSG 11 No. 2 rectangular, NT-MDT Co.). Scans size of 5 x 5 μm2 were
recorded under ambient laboratory conditions. Average surface roughness was then
measured from the AFM images using NT-MDT Image Analysis software (version
2.2.0). Data was presented as mean of at least six regions of each surface.
3.3.4 Statistical analysis
Data from the experiments were expressed as the mean values ± standard
derivation. Analysis was performed using SigmaPlot (version 11.0; Systat software
Inc). For multiple comparisons, one-way analysis of variance (ANOVA) was used
with the Holm-Sidak’s test. The significance level was set at p ≤ 0.05.
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3.4 RESULTS
3.4.1 Surface coating on incubation vials
As shown in Figure 3-4, material blocks synthesised from polymerisation were
used to form a translucent coating on the vials. By changing the comonomers and their
molar fractions in the polymerisation, the surface coatings on incubation vials can
display different surface functional groups: –COOH and –CH3, –CH2CH3 or –
(CH2)3CH3. In the present study, the coatings containing homopolymers of acrylic acid,
methyl, ethyl, butyl methacrylates are referred to as PAA, PMMA, PEMA, and PBMA,
respectively, while those containing both acrylic acid and methyl/ethyl/butyl
methacrylates are named according to the mole fractions of acrylic acid (45, 55, 65%),
for example 45% AA/ 55% MMA as 45MMA etc.
Figure 3-4. a) Materials formed from free-radical polymerisation. b) An incubation vial treated with material solution resulted in a clear coating.
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3.4.2 XPS analysis of material-coated surfaces
To confirm the presence of coating on the inner surfaces of glass vials and
to determine the surface compositions of functional groups on the coating, XPS
analysis was performed on uncoated and coated surfaces. A broader range of AA
mole fractions (0.0 to 0.65) were utilised to validate the gradual change in surface
content of –COOH groups.
Figure 3-5 compares XPS survey spectra of the uncoated glass surfaces and
material-coated surfaces, using PAA, PBMA and 45BMA as examples. As expected,
all coated surfaces showed two peaks of element C1s (285.0 eV) and O1s (532.0 eV)
(Figure 3-5 b-d). The oxygen peak was derived from the ester linkage. For PBMA
and 45BMA, the carbon peak increased due to the presence of methyl groups bonded
to the carbon backbone and the terminating –COO((CH2)3CH3) group compared to
that from –COOH group of PAA. This indicated successful surface modification on
the uncoated substrate by these coatings. Detected Si on the spectrum of PAA
reflected trace silicone contaminants.
XPS C1s spectra, which demonstrate different oxidation states of carbon,
were used to investigate the changes in chemical functionalities at surfaces after
coating. Alkyl methacrylates differed from acrylic acid by a prominent peak
corresponding to carbon component C-O (286.8 eV), whereas both comonomers
showed three common components: C-C (285.0 eV), C-COO (285.7 eV), and O-
C=O (289.1 eV) (Figure 3-6 a, b, d, f) (Chuang and Lin, 2007, Minelli et al., 2008).
This sole peak indicated that carbonyl-containing alkyl functional groups of alkyl
methacrylates displayed on the coated surfaces. A minor peak at around that of C-O
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was observed at PAA coated surface which may be due to the contamination. Coated
surfaces showed all four components, as illustrated by spectra of 45MMA, 45EMA
and 45BMA (Figure 3-6 c, e, g). The relative compositions of C-O and O-C=O on
various surface coatings were measured from the peak areas. The surface ratios of –
COOH/–CH3, –CH2CH3 or –(CH2)3CH3 are summarised in Table 3-1.
Overall, the proportion of –COOH group presented on surface coating
(XCOOH coating) was lower than the proportion of AA monomer fed in the material
(XCOOH material). The relationships between XCOOH coating and XCOOH material are
presented in Figure 3-7. As clearly demonstrated by BMA surfaces, the surface
coatings became increasingly richer in –COOH group (i.e. increased XCOOH coating)
with the increase in the AA monomer feed (i.e. increased XCOOH material).
Interestingly, XCOOH material of 0.45 and 0.55 at MMA and EMA surfaces showed
similar XCOOH coating concentrations (i.e. 0.32 – 0.34). Also, XCOOH coating at
MMA, EMA and BMA surfaces were found to be comparable at XCOOH material of
0.55 (0.33, 0.32 &0.34 respectively), and of 0.65 (0.41, 0. 39 & 0.41 respectively)
(Table 3-1).
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Figure 3-5. XPS survey spectra of a) uncoated glass, b) PAA, c) PBMA, and d) 45BMA (45% AA/BMA) coated surfaces.
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Figure 3-6. XPS C1s spectra of a) PAA, b) PMMA, d) PEMA, f) PBMA, and c) 45MMA, e) 45EMA, and g) 45BMA coated surfaces.
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Table 3-1. Ratio of –COOH groups measured on surface coatings (XCOOH coating) compared to mole fraction of –COOH group-containing AA (XCOOH material) composed
with MMA, EMA or BMA. Data were presented as the average value of three
measurements.
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Figure 3-7. Ratio of –COOH groups on the surface coating as a function of mole fraction of –COOH group-containing AA composed with a) MMA, b) EMA or c) BMA.
3.4.3 Surface hydrophobicity
The advancing contact angle of an uncoated glass substrate was 54.1 ± 4.2 º.
With addition of a coating, the values increased to 77.6 ± 3.1 º, 82.9 ± 1.0 º and 91.8
± 4.4 º for PMMA, PEMA and PBMA, respectively (Table 3-2). These results were
in agreement with published data (Takemoto et al., 2004). The contact angle values
of homopolymer-derived coatings increased significantly in the order: –CH3 < –
CH2CH3 < –(CH2)3CH3, suggesting the surface hydrophobicity increased with the
alkyl chain length (p ≤ 0.001).
Similarly, among coated surfaces displaying –COOH/–CH3, –CH2CH3 or –
(CH2)3CH3 functionalities at the same –COOH ratios, a significantly higher contact
angle was observed on surfaces with –(CH2)3CH3 groups compared to those with –
CH3 and –CH2CH3 groups (i.e. At an average 33% –COOH from XCOOH material of
0.55, p=0.005; At an average 40% –COOH from XCOOH material of 0.65, p ≤ 0.001).
This further indicated that surface hydrophobicity was strongly influenced by the
more hydrophobic –(CH2)3CH3 groups. No significant differences in the contact
angles were noted between surfaces with –CH3 and –CH2CH3 groups at both values
of XCOOH material. This suggested that the differences in chain length and
hydrophobic property of –CH3 and –CH2CH3 groups were less likely to affect
surface hydrophobicity in the presence of –COOH groups, unlike on homopolymer-
derived surfaces.
On the other hand, an increase in –COOH ratio (0% to 41%) on –COOH/–
(CH2)3CH3 surfaces led to a decrease in contact angle, suggesting the contact angle
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correlated well with the –COOH ratio (Figure 3-8). This correlation was also
observed on MMA and EMA surfaces in which the relatively similar –COOH ratios
at XCOOH material of 0.45 and 0.55 resulted in similar contact angles. Overall, these
results indicated that the surface functional groups and their compositions had a
combined effect on surface hydrophobicity.
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Table 3-2. Advancing contact angles of surfaces coated with materials composed of varied
mole fraction of acrylic acid and alkyl methacrylates. Measurements were reported as the
average value of contact angles of at least six data points.
Figure 3-8. Advancing contact angles of different surface coatings.
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3.4.4 Surface morphology and roughness
SEM showed that all coated surfaces exhibited as smooth as the uncoated
surfaces (Figure 3-9 a-c).
The smoothness of coatings was further analysed by AFM. The topography
images demonstrated that all coated surfaces had microscopic concave–convex
structures when compared to irregular protrusions on uncoated glass surfaces (Figure
3-10). These surface features were quantitatively verified by analysing the average
roughness (Table 3-3). The average roughness of coated surfaces was 3.99 ± 0.54
nm. No significant differences were found in the average roughness among surfaces
with similar –COOH ratios (i.e. at XCOOH material of 0.55; p = 0.334 or 0.65; p =
0.775) nor among surfaces with the same functionalities but different –COOH ratios
(i.e. –COOH/–(CH2)3CH3 surfaces; p = 0.066). This indicated that the type of
surface functional groups and their relative ratios did not affect the surface
roughness. Indeed, the roughness values of the coated surfaces were higher than that
of uncoated glass surfaces (2.22 ± 0.29 nm), suggesting the application of coatings
did create a small increase in surface roughness.
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Figure 3-9. Representative SEM images of a) uncoated glass, b) PEMA and c) 45 EMA coated surfaces, taken at magnification of 25000 x.
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Figure 3-10. AFM images (5 µm x 5 µm areas) of uncoated and coated surfaces.
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Table 3-3. Average surface roughness measured by AFM. Values are the means of 6 measurements ± SD.
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3.5 DISCUSSION
In this chapter, we developed a series of materials (AA-co-
MMA/EMA/BMA) at varied ratios and coated them on the inner surfaces of
incubation vials. The fabrication of a stable coating onto incubation vials was
achieved by consecutive additions of material solutions (Fujishita et al., 2009). Such
a surface coating technique is a common and efficient way to modify surfaces of
implants having complex geometries without altering the bulk properties for a
specific application (Ma et al., 2007, Werner et al., 2007). By using this model, the
surfaces should exhibit distinct surface chemistry with compositional variation of –
COOH/–CH3, –CH2CH3 or –(CH2)3CH3 functionalities on the coated surfaces.
Chemical composition of surface coatings
Using XPS analysis, it was confirmed that glass substrates were successfully
coated with materials of differing compositions. This was done by identification of the
elements and an additional peak representing C–O binding, appearing on the surfaces
after an alkyl methacrylate was composed with AA. Using this C–O peak contributed
solely from alkyl methacrylates, we quantified the relative compositions of the
comonomers and corresponding functional groups exposed on the surfaces. XPS is a
more sensitive approach in determining surface density of –COOH groups (McArthur,
2006) compared to colorimetric methods using dyes such as Rhodamine 6G (Kang et al.,
1993), toludine blue or thionin acetate (Tzoneva et al., 2008), which are based on an ion
exchange mechanism and the assumption that –COOH groups on surface are able to
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bind equal molar amounts of dye molecules. Since the binding energy differences
between C-O and O-C=O can be resolved clearly by XPS for quantifying the
concentrations of carboxyl and alkyl functionalities, chemical derivatisation in which a
functional group is tagged with an unique element prior to XPS analysis, in this case
trifluoroethanol for –COOH groups, is not necessary (Alexander et al., 2004).
From our XPS results, the surface coating generally had an increase in –COOH
group with increasing AA monomer fed in materials. However, it was found that the
surface –COOH ratio was lower than the expected AA mole fraction. This lower yield of
–COOH groups on the surface is probably due to three factors: the degree of
copolymerisation, polymer chain mobility, and functional group reorientation.
During free-radical copolymerisation, some comonomers may not incorporate
successfully into the propagating polymer chains due to hindered accessibility to the
primary radicals, or undergo homopolymerisation due to the differences in monomer
reactivity (Gupta et al., 2001, Li et al., 2005). It has been shown that in the bulk
copolymerisation of AA and MMA the monomer reactivity ratios are r1= 1.51 and r2 =
0.48, respectively. This indicates that AA tends to form homopolymers subunits in the
resultant copolymers. This homopolymerisation occurs at low conversion at low
temperature (50 ˚C) (Ekpenyong, 1985, Odian, 2004, El-Newehy et al., 2010). As such,
we performed two additional heating steps at 90 ˚C and 100 ˚C to ensure complete
curing of polymerisation to reduce comonomer residues. Hence, this factor is less likely
the primary contributor to a reduced amount of surface –COOH groups.
Similar to our findings, a lower yield of PAA on the surface after grafting AA
onto poly(ethylene terephthalate) films has also been reported by Gupta et al.(2002). It
has been suggested that the PAA chain underwent rearrangement, resulting in chain
distribution not only on the surface but also in the subsurface layer, which was beyond
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the normal sampling depth (<10 nm) of XPS. Such polymer chain flexibility has been
observed on surfaces prepared from PBMA solutions (Berglin et al., 2009, Xu et al.,
2009). In line with these findings, Berglin et al. (2004) identified that the mobility of
poly(alkyl methacrylates) chains was increased systematically with an increase in alkyl
chain length. Moreover, it has been shown that hydrophobic functional groups on
surfaces tend to expose to air whereas hydrophilic ones behave vice versa to minimize
surface energy (Hermitte et al., 2004, Ozcan and Hasirci, 2007, Michiardi et al., 2007).
Therefore, the general lower surface –COOH ratio in regard of the AA mole fraction in
our study is possibly due to the reorientation of hydrophobic alkyl groups to dominate
outmost surface while the hydrophilic carboxyl groups are buried.
Materials composed of BMA contain longer and more hydrophobic side chains
leading to an enhanced polymer chain mobility. This may explain a linear increase in
surface –COOH ratio with increasing AA mole fraction found on these surfaces. In
contrast, the hydrophilic –COOH groups on MMA and EMA surfaces will be less likely
able to orient away from air due to reduction in polymer chain mobility and accessibility
of surface groups reorientation mediated by the shorter and less hydrophobic –CH3 and
–CH2CH3, respectively. This may partly elucidate the higher concentration of –COOH
groups found on MMA and EMA surfaces at a lower AA mole fractions when compared
with BMA surfaces.
Surface hydrophobicity of coatings
Surface hydrophobicity of implant plays an important role in mediating protein
adsorption and cell responses (Arima and Iwata, 2007, Wei et al., 2009, Menzies and
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Jones, 2010). By measuring advancing contact angle, we demonstrated that surface
hydrophobicity of coatings correlates well with the chemical compositions determined
by XPS, a dependence as other papers reported (Tsyganov et al., 2005, Ukiwe et al.,
2005, Lai et al., 2006). Using homopolymeric surfaces with a single functionality, we
found that the contact angle increased with an increase in carbon number in the side
chain: –CH3 < –CH2CH3 < –(CH2)3CH3, indicating a direct correlation of alkyl chain
length and hydrophobicity properties as has been indicated by other studies (Van
Damme et al., 1986, Wulf et al., 2000, Berglin et al., 2004, Hermitte et al., 2004).
Similarly, advancing contact angles on surfaces with two functionalities clearly
demonstrated that surface hydrophobicity was dependent on both the type of functional
groups and their compositions (Tsai et al., 2007). As expected, water contact angle was
higher on surfaces with –(CH2)3CH3 groups than those with –CH3 and –CH2CH3
groups at relatively the same –COOH ratios (at AA mole fraction of 0.55 and 0.65). The
contact angle of the surfaces with –CH3 and –CH2CH3 groups did not differ
significantly. This suggests that the difference in one-carbon length between –CH3 and –
CH2CH3 groups has limited impact on modulating surface hydrophobicity in the
presence of –COOH groups. In contrast, an increase in –COOH ratios on –COOH/–
(CH2)3CH3 surfaces decreased water contact angle. As such, it is reasonable to deduce
that the nature of functionalities and their relative compositions are the key factors in
controlling surface hydrophobicity.
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Surface morphology and roughness of coatings
Coated surfaces were examined by SEM and AFM in an attempt to characterise
surface morphology of coating and determine any relationship between surface chemical
composition or surface hydrophobicity and surface roughness. Our SEM observation on
coated surfaces showed no difference in morphology compared to uncoated surfaces.
However, the surface topography of coated surfaces as examined by AFM underwent
prominent changes as a result of the coating process. Characteristic concave–convex
structures were observed. This may be related to the relative of alkyl methacrylates in
materials with AA, in acetone which determined the evaporation rate of solvent during
coating process. We observed a decrease in material solubility with increasing AA mole
fraction and alkyl chain length. This might be explained by different monomer reactivity
ratios which leads to different copolymer composition and subsequent solubility. Such a
difference in solubility and resultant surface structure has also been reported in mixed
SAM prepared with alkanethiols dissolved in ethanol solution (Chuang and Lin, 2007,
Tsai et al., 2007). It was suggested that the solubility was governed by specific
interactions, such as hydrogen bonds, between terminal functional groups of solute and
solvent, or among the solutes (Xu et al., 2011). In this case, the longer chain length of –
(CH2)3CH3 groups may reduce the evaporation rate of the solvent and produce less
pores on surfaces by strongly interacting with –COOH groups from AA or acetone.
Whereas the shorter length of –CH3 or –CH2CH3 group may give rise to more tiny pores
due to faster evaporation of solvent due to weaker bonds with surrounding components.
To verify these surface structures quantitatively and correlate the surface
topography in relation to surface chemical composition and hydrophobicity, we
measured the average roughness of these coated surfaces. The roughness data showed
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that there was no nano-scale difference in roughness for surfaces with different alkyl
functional groups at the same –COOH ratios (at AA mole fraction of 0.55 and 0.65).
Also, the surface roughness did not differ significantly with an increase in surface –
COOH ratio, as illustrated by BMA surfaces. These results indicate that there is no
evident relationship between surface roughness and surface functional groups and their
compositions. Furthermore, no cracks or other surface imperfections were found.
Consequently, differences in blood response should not be a result of different surface
morphology and roughness between coatings.
The knowledge generated herein would allow a prior prediction of the surface
properties directly from the material formulation. Further studies of whole blood
response on these materials could yield to phenomenological links between the
biocompatibility and formulation.
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3.6 CONCLUSION
In this chapter, we established an incubation model with inner surface
coated with a series of materials consisting of AA and MMA, EMA or BMA at
different ratios. It was demonstrated that the coated surfaces displayed different
contents of carboxyl and alkyl groups and that –COOH ratios were lower than the
AA mole fraction fed in materials. The properties of surface functional groups and
their relative compositions correlated well with the surface hydrophobicity. It
increased with increasing alkyl chain length: –CH3 < –CH2CH3 < –(CH2)3CH3, and
decreased gradually with increasing –COOH groups, suggesting a combined effect of
surface functional groups and their compositions on surface hydrophobicity. No
significant differences in surface hydrophobicity were found on surfaces with –CH3
and –CH2CH3 groups in the presence of –COOH groups. The coating appeared
relative smooth and the surface average roughness was 3.99 ± 0.54 nm, which is
slightly higher than that of uncoated glass surfaces (2.22 ± 0.29 nm). However, we
did not detect significant difference among surfaces with same functionalities at
different –COOH ratios nor among surfaces with different alkyl groups but the same
–COOH ratios, suggesting the surface chemistry did not influence the surface
roughness. Overall, surface functional groups and their relative compositions have a
combined effect on modulating surface chemistry and hydrophobicity on coatings.
The similarity concerning surface roughness may enable a correlation of blood
response and clot formation with the parameters related to the surface chemistry and
hydrophobicity.
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4Chapter 4: The Influence of Carboxyl and
Alkyl Functional Groups and Their Relative
Compositions on Blood-Biomaterial
Interactions and Clot Properties
4.1 INTRODUCTION
Activation of blood coagulation occurs rapidly when whole blood makes
contact with the surface of a synthetic implant. The adsorption of plasma proteins is
believed to initiate platelet reactions and coagulation activation, leading to the
generation of thrombin and fibrin, and ultimately blood clot formation on the implant
surface (Horbett, 1993, Eskin et al., 2004, Gorbet and Sefton, 2004). Although a
peri-implant clot is regarded to be detrimental to the function of cardiovascular
devices (Knetsch, 2008), it may be beneficial for bone repair when the peri-implant
clot has biological and structural properties similar to those of the haematoma
formed on injured bone during normal healing (Tsiridis et al., 2006, Tosounidis et
al., 2009). Moreover, an important concern of the blood-biomaterial interactions is
the activation of the complement pathway and its potential to initiate chronic
inflammation or foreign body reaction (FBR). These adverse reactions may cause the
deterioration of the implanted biomaterial and secondary injury to surrounding tissue
(Tsai, 2004, Anderson et al., 2008).
It is well known that surface chemical functionalities and their relative ratios
on biomaterials affect protein adsorption which in turn affects the activation of
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coagulation and complement cascades (Sperling et al., 2005a, Rodrigues et al., 2006,
Tsai et al., 2007, Barbosa et al., 2010, Fischer et al., 2010b). Using fibrinogen
solution, it was shown that surface functional groups presented on self-assembled
monolayers (SAMs) markedly influenced the extent of thrombin-catalysed fibrin
polymerisation of adsorbed fibrinogen (Evans-Nguyen and Schoenfisch, 2005b,
Evans-Nguyen et al., 2005b). Moreover, other studies investigated the impact of
thrombin concentration on fibrin structure and showed that a lower thrombin
concentration increased the fibre thickness; but decreased the fibre density (Wolberg,
2007a, Wolberg and Campbell, 2008). However, these studies have mainly used
SAMs on flat surfaces and isolated fractions of blood e.g. protein solution, plasma or
single cell culture, which limit the understanding of the complex response of whole
blood in three dimensions. In addition, the influence that the surface functionalities
and their relative ratios have on the ultimate structural properties of a whole blood
clot has not previously been taken into consideration.
In fact, alterations in fibrin structure of blood clots were shown to correlate
with differing clot elasticity and susceptibility to lysis (Collet et al., 2000, 2005, Liu
et al., 2006), and are believed to cause increased risks of bleeding or thrombosis
(Mills et al., 2002, Bhasin et al., 2008, Undas et al., 2009). Theoretically, these
factors could also affect the release of growth factors from the peri-implant clots,
alter the early stage of healing and consequently the extent of new bone formation at
the implant site (Laurens et al., 2006).
To systematically investigate both the surface chemistry-activated cascade
events and the ultimate clot structural properties upon whole blood contact in three-
dimensions, we developed an in vitro incubation vial where the inner surface was
coated with materials of AA/MMA, EMA or BMA at varied ratios. In this chapter,
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we investigated the effect of binary mixture of –COOH/–CH3, –CH2CH3 or –
(CH2)3CH3 surface functionalities and their relative compositions on coagulation
and complement activation as well as alterations in resultant clot structure, elastic
properties, susceptibility to fibrinolysis and the release of PDGF-AB and TGF- β.
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4.2 MATERIALS & METHODS
4.2.1 Blood sampling and in vitro incubation
Whole blood was collected from a healthy volunteer who was not on any
medication for at least 10 days and with no history of coagulation disorders.
Venipuncture was performed by a phlebotomist. Venous blood was drawn into
syringes with a 19-gauge needle and immediately transferred (1.5 mL) to the vials
and incubated at 37 ˚C. After the desired incubation time, all blood contents in the
vials were collected and processed for following experiments. This procedure was
approved by the Human Ethics Committee of the Queensland University of
Technology. Informed consent was also obtained from donor prior to blood
collection.
4.2.2 In vitro coagulation activation
Activation of the coagulation cascade leads to the conversion of
prothrombin into active thrombin, a process accompanied with the production of
prothrombin fragment 1+2 (F1+2). To assess in vitro coagulation activation on the
coated surfaces, prothrombin F1+2 was analysed using an enzyme-linked
immunosorbent assay (ELISA, Enzygnost F1+2; Dade Behring Marburg GmbH,
Germany). Based on our preliminary studies, the coagulation was analysed after 30
min of whole blood incubation as the formation of F1+2 demonstrated well
established differences among surfaces.
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Serum was isolated by transferring samples from incubation vials to
microcentrifuge tubes containing tri-sodium citrate (0.11 mol/L) (one part of tri-
sodium citrate with 9 parts of blood sample) and centrifuged according to the
manufacturer’s protocols. Briefly, the enzyme immunoassay was based on the
sandwich principle in which the monoclonal antibodies specific to F1+2 antigens
were coated on a 96-well plate. Standards and samples were added and any F1+2
antigens present were bound to the immobilised antibodies. After unbound
substances were rinsed away, peroxidise-conjugated antibodies to human
prothrombin were added to bind to F1+2 determinants, producing an antibody-
antigen-antibody “sandwich”. After a second wash, a substrate solution for
peroxidase was added. The enzymatic reaction between hydrogen peroxidase and its
substrate produced a blue colour in direct proportion to the amount of F1+2 present
in the sample. The reaction was terminated by a stop solution as indicated by a colour
change of blue to yellow. The colour intensity was determined by a
spectrophotometer (at 450nm) and quantified from a standard curve. Average data
from 6 replicates of each surface were presented. Plasma levels of F1+2 serve as a
baseline and were obtained by centrifuging blood in Vacuette test tubes containing
sodium citrate (3.5 mL blue capped, Greiner Labortechnik, Austria). All data were
obtained from 6 replicates of each surface.
4.2.3 In vitro complement activation
Initiation of the complement system leads to the formation of a common end
product C5a convertase, which cleaves complement protein C5 to C5a. C5a is
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rapidly transformed to C5a-des Arginine (C5a-desArg) by endogenous
carboxypeptidase N enzyme in plasma or serum (Janeway et al., 2005a). To
determine the extent of complement activation on material-coated surfaces, serum
C5a-desArg was quantified with Human C5a ELISA Kit II (OptEIATM; BD
Bioscience, USA). Serum collected from blood samples after 2 h incubation was
analysed as the formation of C5a-desArg were detectable and more pronounced
based on our preliminary study, which also agrees with earlier findings (Sperling et
al., 2005a). The sandwich ELISA assay was performed as previously described. A
background level of complement activation monitored as plasma level of C5a-
desArg was obtained by centrifuging blood in Vacuette test tubes containing
EDTA (4.0 mL purple capped, Greiner Labortechnik, Austria). All data were
obtained from 6 replicates of each surface.
4.2.4 Characterisation of clots formed on material-coated surfaces
4.2.4.1 Examination of clot structure
To study whether clot architecture was altered by surfaces presenting various
compositions of functional groups, the clots formed in incubation vials were examined
by SEM. After 2 h incubation, the clots were fixed with 4% paraformaldehyde (pH 7.4)
at 4 ˚C overnight. The clots were washed twice with phosphate buffered saline (PBS, pH
7.4) for 30 min, and dehydrated in grades of ethanol (50 %, 70 % and 100 %) for 1 h per
grade. A longitudinal cut was performed on the clots to allow examination of the
ultrastructures at clot/material interface and the centre of the clot. Following dehydration,
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the clots were processed through 100 % amyl acetate twice of 15 min intervals and dried
in a CO2 critical point dryer. The clots were then mounted, gold-coated and examined
with the SEM microscope. Representative images were captured.
4.2.4.2 Fibrin thickness measurement
The effect of material surfaces on the formation of fibrin in the clots was
assessed by measuring the diameter of fibrin strand from SEM images at 5000 x
magnification. The measurement was performed using Image J software (version 1.43)
according to the method of Lai et al. (2010). A transverse line was drawn perpendicular
to long axis of the fibre with clearly defined margins. The pixel value was related to that
obtained for the scale bar on the image. At least forty different fibrin strands were
measured at random fields approximate 50 µm away from the edge of clot and in the
centre of the clot. A minimum of two images were analysed for each sample. The
diameter of individual fibrin of each sample was reported as an average for all fibres
measured.
4.2.4.3 Fibrin density measurement
Quantitative fibrin network analysis was performed using Image J software
(version 1.43) using a modified method of Undas et al. (2008). A 64-field grid was
generated to cover each SEM image and at least twenty fields were selected
randomly at the edge and in the centre of the clots. The density of fibrin fibre was
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determined by counting the number of fibres per field (40 µm2) and the mean value
was presented.
4.2.4.4 Compaction study
To study the viscoelastic properties of the clots formed on material surfaces,
the clot rigidity was assessed by measuring compaction according to Undas et al.
(2009). After 2 h incubation, the clots formed in the incubation vials were transferred
to Eppendorff centrifuge tube (2.0 mL; Hamburg, Germany) and centrifuged at 6000
g for 60 seconds. The volume of fluid expelled from the network by centrifugation
was measured and expressed as a percentage of the initial volume of the clot and was
termed the compaction coefficient. All data were obtained from 6 replicates of each
surface.
4.2.4.5 Clot lysis assay
Clot lysis strongly correlated with fibrin thickness and density. The effect of
material surfaces on overall clot degradability and its relationship to the fibrin
architecture as determined above was studied. Clot lysis was evaluated by the generation
of fibrin degradation product (D-dimer) as the cross-linked clots were digested by
fibrinolytic enzymes in vitro.
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A suspended clot system modified from the protocols of Collet et al. (1993a)
was used in this study. Whole blood was incubated (1.5 mL) in the vials for 2 h at 37 ˚C
to allow complete clot formation and retraction. The clots were removed carefully from
the vials and suspended in 3 mL of PBS containing human plasminogen (Glu-
plasminogen, 5.4 μg/mL final concentration; American Diagnostica Inc., USA)
(Onundarson et al., 1992). Lysis of the clot was induced by adding recombinant tissue-
type plasminogen activator (tPA, 0.25 μg/mL final concentration; American Diagnostica
Inc., USA) at 37 ˚C with gentle agitation. This concentration of tPA was determined as
the lowest one able to induce clot lysis in preliminary study (data not shown). Aliquots
of 300 μL were removed at timed intervals and centrifuged at 1000 g for 3 min. The
supernatants were stored at -70 ˚C before analysis. The same volume of PBS was
supplemented after samplings. The extent of clot lysis was monitored by measuring the
amounts of D-dimer released from the clots using IMUCLONE D-Dimer ELISA
(American Diagnostica Inc., USA). Clots that were suspended in PBS only were used as
control of spontaneous fibrinolysis. Weight loss of clots during lysis was also traced
during the experiments (Prasad et al., 2006, Holland et al., 2008). All data was obtained
from triplicate of clots formed on each surface.
4.2.4.6 Quantification of growth factors
The release of PDGF-AB and TGF-β1 during both clot formation and clot
lysis was assayed by ELISA. After 2 h incubation, supernatant serum above the clots
was collected and the clots were directly subject to clot lysis as described previously.
Supernatant serum and buffer aliquots collected at varied intervals were centrifuged
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for 15 min at 1000 g and assayed according to the manufacturer’s instructions. All
Quantikine ELISA kits were purchased from R&D Systems (Minnesota, USA) and
the sandwich ELISA assay was performed as previously described. To detect the
circulating levels of growth factors, platelet-poor plasma was prepared by
centrifuging blood in Vacuette test tubes containing EDTA for 15 min at 1000 g
and subsequently for 10 min at 10,000g (4.0 mL purple capped, Greiner
Labortechnik, Austria). All data were obtained from triplicates of clots formed on
each surface.
4.2.5 Statistical analysis
Data from the experiments were expressed as the mean values ± standard
deviation. Analysis was performed using SigmaPlot (version 11.0; Systat software
Inc). The data from control and test surface were compared using the Student’s t-test.
For multiple comparisons, one-way analysis of variance (ANOVA) was used with
the Holm-Sidak’s test. The significance level was set at p ≤ 0.05.
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4.3 RESULTS & DISCUSSION
4.3.1 Surface-initiated coagulation response
To determine the effect of material surfaces on initiating coagulation cascade,
the serum level of prothrombin F1+2 after 30 min incubation was measured by ELISA.
Figure 4-1 shows F1+2 level relative to the uncoated glass surfaces. All material-coated
surfaces had higher levels than the plasma (0.014 ± 0.0007 %, i.e. 0.06 ± 0.003 nmol/L)
but lower levels compared to the uncoated glass surfaces (p ≤ 0.001). All surfaces
resulted in clot formation.
To evaluate the contribution of surface functional groups and their
compositions on activating the coagulation cascade, surfaces with the same
functionalities but different –COOH ratios , and those with different alkyl groups but
relatively similar –COOH ratios were grouped to compare (Figure 4-1). In general, all
BMA surfaces induced significantly higher F1+2 level than MMA and EMA surfaces,
regardless of the –COOH ratios (p ≤ 0.001). Amongst BMA surfaces with various –
COOH ratios, 55BMA had a slightly higher F1+2 level compared to 45BMA and
65BMA but no significant difference was detected (p = 0.305).
Among 55MMA, 55EMA & 55BMA surfaces with approximately 33% –
COOH, 55BMA showed the highest F1+2 level whereas 55EMA showed the lowest
level (p ≤ 0.001). A similar effect of alkyl groups on coagulation activation was also
found on surfaces with a higher content of –COOH (i.e. 40% on 65MMA, 65EMA &
65BMA) as 65BMA and 65EMA showed the higher and lower F1+2 level, respectively
(p ≤ 0.001). In contrast to BMA surfaces, an increase in –COOH ratio on MMA and
EMA surfaces reduced F1+2 level significantly (p ≤ 0.001).
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Figure 4-1. The serum levels of prothrombin F1+2 after 30 min of whole blood incubation with material-coated surfaces relative to the uncoated glass surfaces (%). Plasma levels served as baseline. Data was presented as mean of six replicates of each surface with SD. * p ≤ 0.001
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In all cases, these results demonstrated that the material-coated surfaces, and
not the glass substrate dictated the procoagulant property of the blood-contacting surface.
The assay of prothrombin F1+2 after 30 min incubation indicated that the instant rate of
coagulation initiated by material surfaces within the period. Since the half-life of F1+2 is
approximately 90 min, the level of F1+2 not only represents earlier and more discrete
instances of the prethrombotic state than other markers with shorter half-life (e.g. 3-5
min of fibrinopeptides A), but also is less susceptible to interference from other in vitro
activation, such as venepuncture (Mannucci et al., 1992, Greenberg et al., 1994).
Significant differences found in prothrombin F1+2 level among surfaces after
30 min of incubation indicated that the rates of thrombin generation and coagulation
initiated by the surfaces were different. All material-coated surfaces showed a reduced
rate of coagulation initiation compared to the uncoated glass surfaces. More importantly,
we observed surface functional groups and their compositions strongly influenced the
rate of coagulation activation.
Influence of varying ratios of surface functionalities on coagulation
The rate of coagulation activation was found to be a function of the –COOH/–
CH3 or –CH2CH3 ratio, decreasing as the ratio increases. This finding is surprising since
the initiation of intrinsic pathway on –COOH/–CH3 SAM surfaces has been shown to
increase with increasing –COOH ratio, as measured by bradykinin generation (Sperling
et al., 2005a) or kallikrein/FXIIa activity (Sperling et al., 2009, Fischer et al., 2010b).
Previous studies demonstrated that FXII activation is directly dependent on the amount
of negatively charged functional groups (Sanchez et al., 2002, Zhuo et al., 2006, Chen et
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al., 2007, Tzoneva et al., 2008). Indeed, Sperling et al. (2009) found that –COOH/–CH3
SAM surfaces with less than 50% –COOH did not show FXIIa activity in plasma phase.
Rather, these surfaces had a noticeable effect on activating FXII adsorbed on the
surfaces. Hence, we propose that increasing –COOH from 33% to 40% on our –
COOH/–CH3 surfaces would also increase the activation of absorbed FXII and hence the
rate of coagulation activation within 30 min. However, our results demonstrate a relative
alteration in the rate of coagulation activation on these surfaces. This may be related to
platelet-mediated FXII activation.
As pointed out in a review by van der Meijden and Heemskerk (2010),
activated platelets could trigger FXIIa-mediated intrinsic activation by secreting
polyphosphate, and lead to a high rate of thrombin generation in the presence of
negatively charged surfaces. Bäck et al. (2009, 2010) also provided evidence
suggesting that FXII is activated on the surfaces of platelets in a more physiological
environment, clotting whole blood. While it was showed that almost no platelets
adhered on 100% –COOH surfaces in vitro, however the adhesion increased with
increasing –CH3 content on –COOH/–CH3 surfaces and peaked at 100% –CH3
(Sperling et al., 2009, Fischer et al., 2010b). The platelets exposed to –CH3 surfaces
were also highly active compared to –COOH surfaces, and remained to be
moderately active on surfaces with a range of 50-100% –CH3 (Lin and Chuang,
2000, Sperling et al., 2005a, 2009, Fischer et al., 2010b). These findings support that
neither 100% –COOH surfaces with elevated intrinsic activation nor 100% –CH3
surfaces with strong platelet adhesion alone, was sufficient to boost a strong
coagulation activation (Sperling et al., 2009, Fischer et al., 2010b). It also indicates
that an interplay between FXIIa initiation and activated platelets propagation is
crucial for a substantial coagulation response. This probably explains our finding of
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an alteration of the rate of coagulation activation at 30 min. Increasing –COOH
content from 33% to 40% on –COOH/–CH3 or –CH2CH3 surfaces (which in turn
decreases the alkyl group concentration from 67% to 60%), may lead to a decline of
platelet-dependent amplification of coagulation and therefore decreased rate of
coagulation activation.
Specific activities of surface carboxyl and alkyl groups on coagulation
The surface carboxyl and alkyl groups also showed specific effects on the
rate of coagulation initiation. Regardless of varied –COOH ratios, we found that the
–(CH2)3CH3 groups induced a faster rate of coagulation activation than –CH3 and –
CH2CH3 groups and that varying –COOH ratios in the presence of –(CH2)3CH3
groups were less likely to affect the rate of coagulation. The faster rate of coagulation
response in the presence of more hydrophobic –(CH2)3CH3 groups may be attributed
to the amount and conformation of surface adsorbed fibrinogen, an important ligand
for platelets (Fuss et al., 2001, Tsai et al., 2002, Mosesson, 2005).
It was shown that a higher amount of fibrinogen adsorbed on hydrophobic
materials was associated with stronger platelet adhesion and activation when
compared to hydrophilic materials (Nygren, 1996, Keuren et al., 2002, Tsyganov et
al., 2005, Zha et al., 2009, Li et al., 2009, Faxälv et al., 2010). Similar findings were
also reported when comparing hydrophobic –CH3 to hydrophilic –OH or –COOH
groups on SAMs (Evans-Nguyen and Schoenfisch, 2005a, Rodrigues et al., 2006).
The higher affinity of fibrinogen toward hydrophobic surfaces is proposed to be
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driven by a stronger hydrophobic interaction between D-domain of adsorbed
fibrinogen and the substrate, at the expense of less favourable water-protein
interaction. Conversely, the electrostatic interaction between positively charged αC-
domain of fibrinogen and negatively charged hydrophilic surfaces is weak in nature
(Ta et al., 1998, Krishnan et al., 2006). Moreover, the strong thermodynamic driving
force is believed to cause adsorbed fibrinogen on hydrophobic –CH3 SAMs to
undergo a greater conformational change when compared to hydrophilic –COOH or
–OH SAMs (Sit and Marchant, 1999b, Evans-Nguyen et al., 2005a, Roach et al.,
2005, Xu and Siedlecki, 2009, Koo et al., 2010). A greater extent of unfolding of
adsorbed fibrinogen as indicated by a lower ratio of α-helix to β sheet was shown to
correlate with increased thrombogenicity as two distinctly binding sites for platelet
adhesion and activation were exposed (Koh et al., 2010a, Sivaraman and Latour,
2010).
Hence, these findings explained a polymer surface developed by Sperling et
al. (2007), where –(CH2)3CH3 groups derived from serine-(tert-butyl)-methylester
yielded the highest fibrinogen adsorption, platelet activation and the fastest rate of
coagulation activation when compared to –CH3 and –OH groups after 2 h incubation.
It is also in accordance with our results concerning fastest rate of coagulation
activation on –(CH2)3CH3 bearing surfaces than those with –CH3 or –CH2CH3.
Accordingly, the weaker electrostatic force between negatively charged –COOH
groups and fibrinogen compared to the strong hydrophobic force between
hydrophobic –(CH2)3CH3 and fibrinogen implies that the possible effect of varying
the –COOH ratios in the presence –(CH2)3CH3 groups would be too minor or
limited to influence the overall coagulation activation, unlike that seen on –CH3 and
–CH2CH3 bearing surfaces.
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Interestingly, between two less hydrophobic –CH3 and –CH2CH3 groups,
the –CH3 group displayed a specific activity in enhancing the rate of coagulation
initiation. As the surfaces with these two alkyl groups did not show a significant
difference in water contact angles at the same –COOH ratios, this finding suggests
that the surface hydrophobicity (surface free energy) is not the sole factor that
governs blood coagulation response.
Indeed, Sivaraman et al. (2009) convincingly demonstrated that –COOH
and –OH SAMs with similar levels of surface hydrophilicity induced a significant
difference in the degree of structural change of adsorbed fibrinogen and albumin. A
complex and non-linear relationship between platelet adhesion and water contact
angles has also been reported on –COOH/–NH2 SAM surfaces with graded mole
fractions (Chuang and Lin, 2007) or on a common clinical material oxidized
titanium, on which there is almost no adherent platelets in the range of 20-30 º water
contact angle, but a dramatic increase in platelet number up to 140 % at 80 º, and
then decreased above 110 º (Takemoto et al., 2004). Since surface hydrophobicity is
actually determined by the surface chemical species, the characteristics of the surface
functionalities would play a more significant role in modulating coagulation
activation.
Based on the study of Arima and Iwata (2007), preabsorbed albumin, the
most abundant blood protein, is effectively displaced by cell adhesive proteins on –
COOH/–CH3 surfaces with water contact angle less than 90 º. Surfaces with –
COOH/–CH2CH3 functionalities were not studied. Our data showed that at the same
–COOH ratios all –COOH/–CH3 and –COOH/–CH2CH3 surfaces presented water
contact angle less than 90 º. We postulate that absorbed albumin on –COOH/–CH3
surfaces is more efficiently replaced by fibrinogen than on –COOH/–CH2CH3
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surfaces, leading to increased rate of coagulation activation. However, further studies
such as a detailed time course are required to confirm this.
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4.3.2 Surface-initiated complement response
To assess complement cascade initiated by material surfaces, the serum level of
C5a-desArg after 2 h incubation was measured by ELISA. Figure 4-2 shows C5a-
desArg level relative to the uncoated glass surfaces. All material-coated surfaces had
dramatically reduced level compared with the uncoated glass surfaces, though
expectedly higher than the plasma level (6 ± 1%, i.e. 7 ± 1 ng/mL) (p ≤ 0.001).
Similar to the trends observed for coagulation activation, BMA surfaces
generally had higher C5a-desArg level than MMA and EMA surfaces (p ≤ 0.001).
Among BMA surfaces, 55BMA had a significantly lower C5a-desArg level than
45BMA and 65BMA (p ≤ 0.001). No difference was found between the latter two
surfaces.
For surfaces with approximately 33% –COOH (55MMA, 55EMA & 55BMA),
and 40% –COOH (65MMA, 65EMA & 65BMA), a lower C5a-desArg level was found
on EMA surface for both percentages (p ≤ 0.001). An increase in –COOH ratio on
MMA and EMA surfaces also further reduced C5a-desArg level (p = 0.036; p = 0.04,
respectively). The lowest C5a-desArg level among all coated surfaces was on the
65EMA surface.
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Figure 4-2. The serum levels of C5a-desArg after 2 h of whole blood incubation with material-coated surfaces relative to the uncoated glass surfaces (%). Plasma levels served as baseline. Data was presented as mean of six replicates of each surface with SD.* p ≤ 0.001
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Overall, these results demonstrated that all material-coated surfaces remarkably
reduced complement response compared with uncoated glass surfaces. Moreover, we
observed the effect of surface functional groups and their compositions on complement
activation followed an entirely similar pattern of surface-induced coagulation. The extent
of complement activation was significantly elevated in the presence of –(CH2)3CH3
groups, but reduced in the presence of –CH2CH3 groups. A further reduction in
complement response was observed with increasing –COOH ratio on the surfaces with –
CH3 and –CH2CH3. These data indicate that there is an association between these two
blood cascades.
Cross-talk between complement and coagulation system
In fact, the haematology community has long realised that the complement and
coagulation cascades appear to interact significantly in vivo. Interplay between various
proteins and cells is involved in both cascades (Figure 4-3).
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Figure 4-3. Cross-talk between complement and coagulation cascades. Red arrows illustrate
the amplification of complement activation by coagulation system. Black dash lines show the
procoagulant activities of complement system. i) Binding of C1q, an initiator of classical
pathway of complement is known to activate platelets to express P-selectins (Peerschke et
al., 1993, Peerschke and Ghebrehiwet, 1998). ii) Anaphylatoxin C3a activates platelets,
enhancing their aggregation. iii) Anaphylatoxin C5a initiates the extrinsic pathway of
coagulation by inducing TF expression from endothelial cells (Ikeda et al., 1997, Tedesco et
al., 1997), neutrophils and monocytes (Guo and Ward, 2005, Ritis et al., 2006, Kourtzelis et
al., 2010). iv) It also enhances blood thrombogenicity by upregulating PAI-1 on mast cells
and basophils. v) Furthermore, incorporation of the terminal complement complex (also
known as C5b-9 complex) into platelet membrane activates platelets to expose procoagulant
lipids, TF and release of TF-bearing microparticles (MPs) (Sims and Wiedmer, 1991, 1995).
On the other hand, coagulation contributes significantly to complement response. vi)
Intrinsic coagulation factors FXIIa (Ghebrehiwet et al., 1981, 1983) and kallikrein (Wiggins
et al., 1981, Thoman et al., 1984) are shown to trigger the classical pathway of complement.
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vii) FXIa, FXa and thrombin are also reported to have enzymatic activities for C3 and C5
cleavage in vitro and ex vivo (Amara et al., 2008, 2010, Huber-Lang et al., 2006). Activated
platelets not only viii) trigger classical pathway of complement through reciprocal
interaction of C1q (Peerschke et al., 2006, 2010), xi) stimulate alternative pathway by
expression of MP (Yin et al., 2008), P-selectins (del Conde et al., 2005, Peerschke et al.,
2008, Hamad et al., 2010), release of chondroitin sulphate from its granules (Hamad et al.,
2008), xii) but also amplify complement through phosphorylation of C3b.
The high interdependency between complement and coagulation activation
on material surfaces confirms our whole blood incubation system is able to maintain
free cross-talk between both systems. These data show unambiguously that surface
functional groups and their relative ratios have a synergistic effect on modulating the
activation of both cascades.
We found that complement response on material surfaces was significantly
reduced with the alkyl length in the order: –(CH2)3CH3 > –CH3 > –CH2CH3.
Berglin et al. (2004) suggested that complement activation is reduced with increased
alkyl chain length of poly(alkyl methacrylates) ranged from 4 to 18 carbons.
However, similar to our findings, they also found that PMMA with one carbon in its
alkyl chain (i.e. –CH3) induces slightly less activation than PIBMA (poly (isobutyl
methacrylate) with four carbons in its alkyl chain (i.e. –(CH2)3CH3), of which the
complement activity also did not differ from that of PBMA. On the other hand, we
found that an increase in –COOH contents (from 33% to 40%) on –COOH/–CH3 or
–CH2CH3 surface reduced complement activation upon clot formation, a similar
trend observed in coagulation response. Inconsistently, Salvador-Morales et al.
(2009) demonstrated an increase in –COOH content (0, 25, 50, 75, 100%) on lipid-
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polymer nanoparticles presenting –COOH/–CH3 increases complement activation,
and virtually have no effect on coagulation system after serum incubation. Moreover,
other studies using anticoagulated whole blood also did not find any significant
difference in complement-generated C5a and C3b adsorption among –COOH/–CH3
SAM surfaces with a range of –COOH contents (0, 50, 83, 100%), and reported that
the surface chemistry-initiated coagulation is independent of the complement
response (Fischer et al., 2010a, Sperling et al., 2009). We believe that the difference
in experimental conditions such as uses of heparin or isolated components may
inhibit both systems for cross-talk and thus be attributable to this discrepancy of
results.
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4.3.3 SEM analysis of clot morphology and structure
Despite various rates of coagulation were initiated by the material coatings, all
these surfaces effectively supported clot formation within 30 min. To investigate the
effect of material coating-initiated coagulation on ultimate clot structure, the fibrin
thickness and density at the edge and in the centre of clots were examined by SEM after
2 h.
4.3.3.1 Effect of various ratios of surface carboxyl groups on fibrin structures
As evident from the micrographs shown in Figure 4-4 a-d, the clots formed on
BMA surfaces showed a network of many thicker and highly branched fibrin fibers,
displaying small interstitial pores at the edge. In contrast, the clots formed on uncoated
glass surfaces showed a network of thinner, less branched and small numbers of fibres
displaying large interstitial pores. The fibrin fibres at the edge of clots on glass surfaces
were significantly smaller in diameter (Figure 4-4 i) and lower in density (Figure 4-4 k)
than those of BMA surfaces (p ≤ 0.001).
Among BMA surfaces, the fibrin diameter at the edge of clots was significantly
smaller on 55BMA, while the fibrin density was significantly lower on 45BMA (p ≤
0.001). No difference in fibrin density at the edge was found between 55BMA and
65BMA (p = 0.237).
Fibrin architecture changed dramatically from the edge to the centre of the clot.
The fibrin of all surfaces except 45BMA increased in diameter (Figure 4-4 j), while the
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densities of all surfaces decreased approximately 3 to 5 times (p ≤ 0.001) (Figure 4-4 l).
In the centre of the clots, 45BMA produced significantly thinner fibres (p ≤ 0.05) at
higher density (p ≤ 0.001) than all other surfaces. These results suggested that variation
of the –COOH ratio on –(CH2)3CH3 bearing surfaces led to significant changes in the
fibrin thickness and network density.
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Figure 4-4. Scanning electron microscopy analysis of whole blood clot structures formed on 45BMA, 55BMA, 65BMA and uncoated glass surfaces. Micrographs of the edge of clots (top panel; a-d) and the centre of clots (bottom panel; e–h), scale bar represents 20 µm. Comparison of fibrin thickness (diameter; nm) i) at the edge; j) at the centre of clots. Comparison of fibrin density (fibre number per 40 µm2) k) at the edge; l) at the centre of clots. Data of fibrin thickness was presented as mean of at least 40 fibrin fibres measured at random field while data of fibrin density was presented as mean of fibre numbers quantified in at least 20 random areas of 40 µm2 at the edge and at the centre of the clots of each surface with SD.* p ≤ 0.001
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4.3.3.2 Effect of various surface alkyl groups on fibrin structures
To evaluate the effect of surface alkyl groups –CH3, –CH2CH3 and –
(CH2)3CH3 on clot ultrastructures, the clots formed on surfaces with approximately
33% –COOH (55MMA, 55EMA & 55BMA) and uncoated glass surfaces were
compared.
As shown in Figure 4-5 a-d, the edge of clots formed on 55MMA were a
dense and highly homogenous network with small pores. For 55EMA, the clots were
a dense network similar to that of 55MMA but also had multiple fibres bundled
together resulting in a highly heterogeneous architecture. While the clots of 55BMA
displayed a more porous network relative to the others, a region with tightly packed
fibres was also observed leading to a slightly heterogeneous network. However,
55MMA, 55EMA and 55BMA surfaces did not differ significantly in the fibrin fibre
diameters (p = 0.878) (Figure 4-5 i) and fibrin densities (p = 0.404) (Figure 4-5 k).
Instead, the fibrin fibres at the edge of clots formed on these surfaces were
significantly larger in diameter and higher in density than those of glass surfaces (p ≤
0.001).
Compared to the fibres at the edge, the fibres at the centre of the clots formed
on all surfaces except 55EMA increased in diameter (Figure 4-5 j), and the fibrin
densities of all surfaces decreased approximately 4 to 5 times (p ≤ 0.001) (Figure 4-5 l).
Among 55MMA, 55EMA and 55BMA surfaces, the fibres at the centre of the clots of
55BMA were significantly larger in diameter while those of 55EMA were significantly
higher in density (p ≤ 0.001). Overall, 55MMA, 55EMA and 55BMA surfaces showed a
similar trend of fibrin density at the edge and at the centre of clots, though the difference
at the edge of clots was not significant.
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Figure 4-5. Scanning electron microscopy analysis of structure of clots formed on 55MMA, 55EMA, 55BMA and uncoated glass surfaces. Micrographs of the edge of clot (top panel; a-d) and the centre of clot (bottom panel; e–h), scale bar represents 20 µm. Comparison of fibrin diameter (nm) i) at the edge; j) at the centre of clots. Comparison of fibrin density (fibre number per 40 µm2) k) at the edge; l) at the centre of clots. Data of fibrin thickness was presented as mean of at least 40 fibrin fibres measured at random field while data of fibrin density was presented as mean of fibre numbers quantified in at least 20 random areas of 40 µm2 at the edge and at the centre of the clots of each surface with SD.* p ≤ 0.001
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The clots formed on surfaces with approximately 40% –COOH (65MMA,
65EMA & 65BMA) and uncoated glass surfaces were observed in Figure 4-6 a-d.
Compared to the uncoated glass surfaces, the edge of clots formed on
65MMA, 65EMA and 65BMA showed a highly branched network composed of
numerous, thicker and longer fibres with smaller pores. The fibrin fibres at the edge
of clots formed on 65MMA, 65EMA and 65BMA were significantly larger in
diameter and higher in density than those on glass surfaces (p ≤ 0.001) (Figure 4-6 i,
k). Among 65MMA, 65EMA and 65BMA surfaces, the fibrin diameter at the edge of
the clots of 65BMA was significantly larger than the others (p ≤ 0.001). No
difference was found between those of 65MMA and 65EMA (p = 0.287). Moreover,
the fibrin density was significantly higher on 65MMA while lower on 65EMA (p ≤
0.001).
Compared to the fibres at the edge, the fibres at the centre of the clots on all
surfaces except 65EMA increased in diameter (Figure 4-6 j), and the fibrin densities of
all surfaces decreased approximately 5 times (p ≤ 0.001) (Figure 4-6 l). While the fibrin
diameter of 65EMA was significantly smaller than those of 65MMA and 65BMA (p ≤
0.001), no significant differences were found between the latter (p = 0.105). For the
fibrin density at the centre of the clots, it was significantly higher on 65MMA than all
the others (p ≤ 0.001). Although the mean fibrin densities of 65EMA and 65BMA were
not significant different, a similar trend in fibrin density was observed at the edge and at
the centre of clots formed on 65MMA, 65EMA and 65BMA.
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Figure 4-6. Scanning electron microscopy analysis of structure of clots formed on 65MMA, 65EMA, 65BMA and uncoated glass surfaces. SEM micrographs of the edge of clot (top panel; a-d) and the centre of clot (bottom panel; e–h), scale bar represents 20 µm. Comparison of fibrin diameter (nm) i) at the edge; j) at the centre of clots. Comparison of fibrin density (fibre number per 40 µm2) k) at the edge; l) at the centre of clots. Data of fibrin thickness was presented as mean of at least 40 fibrin fibres measured at random field while data of fibrin density was presented as mean of fibre numbers quantified in at least 20 random areas of 40 µm2 at the edge and at the centre of the clots of each surface with SD.* p ≤ 0.001
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In summary, 65BMA (40% –COOH/60% –(CH2)3CH3) surfaces resulted in
the fibrin fibres at the edge of clots that were significantly larger in diameter than all
other surfaces (p ≤ 0.001). The mean fibrin diameter at the centre of clots on 65BMA
was also larger than all the others, suggesting this composition of –COOH/–(CH2)3CH3
functionalities have an influence throughout the clots. In contrast, the uncoated glass
surfaces produced fibrin with significantly smaller diameter and in lower density at the
edge of clots compared to all other surfaces (p ≤ 0.001). On the other hand, at the centre
of the clots, 65EMA (40% –COOH/60% –CH2CH3) seemed to be more effective in
decreasing both the fibrin diameter and density as the mean values of these two
parameters of this surface were lower than all the others. Furthermore, 65MMA (40% –
COOH/60% –CH3) and 45BMA (25% –COOH/75% –(CH2)3CH3) surfaces led to a
higher fibrin density compared to all other surfaces, at the edge and the centre of the
clots, respectively (p ≤ 0.001).
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Together, our findings indicate that material-coated surfaces modulate the
whole blood clot structure. The changes in alkyl length of –CH3, –CH2CH3 and –
(CH2)3CH3) groups as well as the concentration of –COOH groups on the coated
surfaces showed a combined and direct effect against the fibrin formation, in term of
fibre thickness and density, from the exterior to the interior part of the clots.
The effect of material surfaces on clot fibrin architectures
In direct contact with material-coated surfaces, the clots display a tight
network with thicker fibrin compared to those in contact with uncoated glass
surfaces. Generally, the interior part of clots on all surfaces become an extremely
loose network with thicker fibres. The dramatic changes in fibrin structure at the clot
exterior among various surfaces, and from clot exterior to interior may be due to a
combination of two mechanisms.
(1) The pattern of in situ thrombin generation follows the initiation,
amplification and propagation phases of coagulation. These phases are in turn
profoundly affected by environmental factors (Sauls et al., 2003, Allen et al., 2004,
Scott et al., 2004, Wolberg et al., 2005, Machlus et al., 2009). As such, a dynamic
change in the thrombin concentration (1nM to greater than 500 nM) (Mann et al.,
2003) may lead to significant differences in kinetics of fibrinopeptide release,
protofibril and fibre formation (Blomback, 2000, Carr et al., 2002a, Wolberg and
Campbell, 2008). This is also tied into the fact that normal plasma clots usually
display a bimodal distribution of the fibre diameters (Shah et al., 1982, Collet et al.,
1993b).
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Using turbidity assays of fibrinogen solutions, it has been illustrated that
clots produced in the presence of polymers contains heterogeneous fibrin structure
due to changes in protofibril aggregation rate, the number and size of fibres formed
compared to controls with an addition of a single thrombin concentration (Lai et al.,
2010).
In line with this, it has been demonstrated that the surface functional groups
significantly affect the efficacy of adsorbed fibrinogen to convert to fibrin (Sit and
Marchant, 2001, Wang et al., 2007, Rodriguez Hernandez et al., 2009). With similar
amounts of adsorbed fibrinogen, a denser fibrin network with more branches was
found on –CH3 surface associating with a larger amount of fibrinopeptides released
at a faster rate compared to sparse fibrin observed on –COOH surfaces (Evans-
Nguyen and Schoenfisch, 2005b, Evans-Nguyen et al., 2005b, 2006). The extent of
fibrinopeptides release and fibrin proliferation have been shown to be related to
surface-dependent fibrinopeptide availability. Approximately 2.7 fold more
accessible fibrinopeptide A was found on fibrinogen adsorbed on –CH3 surfaces for
thrombin cleavage than those on –COOH surfaces. Hence, this may explain the
higher efficacy of fibrin proliferation observed on –CH3 surfaces (Geer et al., 2007).
Furthermore, different alkyl chain length of poly (alkyl methacrylates) have also
displayed a major effect on regulating the rate of thrombin generation and fibrin
deposition (Berglin et al., 2004, 2009). This implies that surfaces functionalities –
COOH/–CH3, –CH2CH3 or –(CH2)3CH3) at varied ratios very likely influence the
fibrinopeptides availability, extent and kinetic of fibrinopeptides release leading to
different fibrin architecture as observed in our study.
Indeed, we found that elevated levels of prothrombin F1+2 on uncoated
glass and 55BMA surfaces produced clots with much thinner fibres, compared to
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material-coated surfaces and BMA surfaces, respectively. These findings are
consistent with the work of Wolberg et al. (2003) in which elevated prothrombin
level triggers the formation of thinner fibrin due to increased initial rate, peak and
total amount of thrombin generation.
(2) Highly procoagulant cells (e.g. activated platelets) have been shown to
support the formation of denser fibrin networks that are more resistant to fibrinolysis
and that the density and stability decrease with increasing distance from the cell
surface (Campbell et al., 2009, Wolberg, 2010, Aleman et al., 2011). For instance,
procoagulant human fibroblasts have been shown to produce denser networks in 10-
µm region proximal than distal to (40-50 µm) its surface (Campbell et al., 2008,
2009). Given the differences in procoagulant properties of cells and plasma factors in
surrounding milieu, a thrombin gradient will be formed in space and therefore may
cause the formation of a range of fibre thicknesses and densities across a region of
growing clots (Ovanesov et al., 2005, Panteleev et al., 2006). This likely explains our
observation of spatially heterogeneous clot morphology with fibrin propagation away
from the site of initiation to the interior part of clots.
In addition, we observed a consistent trend on fibrin density at the clot
exterior and interior on surfaces containing same –COOH ratio but different alkyl
groups. This indicates that the surface functionalities and relative ratios have a
specific influence on fibrin density throughout the clots.
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4.3.4 Assessment of clot rigidity by compaction
Our previous results showed that clots formed on material-coated surfaces
had differing fibrin architectures. To determine whether such changes in fibrin
structure influences the mechanical properties of whole blood clot, a compaction
study was performed to investigate the clot rigidity.
The compaction coefficient for clots formed on all material-coated surfaces
was lower than that on uncoated glass surfaces (p ≤ 0.001), suggesting the clot
rigidity was greatly increased by the coated surfaces (Figure 4-7).
For surfaces presenting 33% –COOH (55MMA, 55EMA & 55BMA), and
40% –COOH (65MMA, 65EMA & 65BMA), the clot rigidity was significantly
elevated on MMA surfaces but reduced on BMA surfaces at both percentages (p ≤
0.001) (Figure 4-7 a-b). As evidenced from BMA surfaces, an increase in –COOH
groups resulted in a significant increase in clot rigidity (p ≤ 0.05) (Figure 4-7 c).
Similarly, an increase in –COOH groups on EMA surfaces also resulted in increased
clot rigidity as shown by comparing 55EMA and 65EMA (p ≤ 0.001). However, the
difference between clot rigidity on MMA surfaces due to an increase in –COOH ratio
was not significant (55MMA vs 65MMA; p = 0.921).
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Figure 4-7. Compaction studies of clots formed on various material-coated surfaces compared to the uncoated glass surfaces. a) Among 55MMA, 55EMA and 55BMA surfaces presenting 33% –COOH. b) Among 65MMA, 65EMA and 65BMA surfaces presenting 40% –COOH. c) Among 45BMA, 55BMA and 65BMA surfaces which contains different concentration of –COOH groups. Data was presented as mean of six replicates of each surface with SD.* p ≤ 0.001
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Overall, these findings demonstrated that the stiffness of clots formed on
material-coated surfaces was significantly enhanced by –CH3 groups compared to –
CH2CH3 and –(CH2)3CH3 groups and that the increase in –COOH ratio on surfaces
with –CH2CH3 and –(CH2)3CH3 groups also improved the clot rigidity.
Relationship between fibrin architectures and clot rigidity
With similar clot mass and approximately 3-5 fold less dense fibrin
network in the centre of clots, the densely packed and cross-linked thick fibres at the
clot exterior may contribute mostly to the network strength to resist collapse under
centrifugal force (Nair and Shats, 1997, Collet et al., 2005). Therefore the notable
difference in fibrin structure at the clot exterior between material-coated surfaces and
uncoated glass surfaces may explain the dramatic increase in clot rigidity of coated
surfaces assessed by compaction studies.
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4.3.5 Clot lysis
Having shown the differences in fibrin structural network and overall
rigidity of clots formed on various material-coated surfaces, we further determined
whether such changes in clot structure and elastic properties altered the clot
susceptibly to lysis. This was carried out by degrading the clot with tPA and
plasminogen, and assessing the release of D-dimer during fibrinolysis as a function
of time after an addition of tPA.
4.3.5.1 Effect of various ratios of surface carboxyl groups on fibrinolysis
The levels of D-dimer released from clots formed on BMA and uncoated
glass surfaces over 24 h of lysis were shown in Figure 4-8 b. After 1 h of lysis, a
significant increase in D-dimer concentration was detected from uncoated glass
surfaces compared to BMA surfaces (p ≤ 0.001) (Figure 4-8 a). As the level of D-
dimer is indicative of the rate of fibrinolysis, this finding suggested that the clots of
glass surfaces initially underwent a faster rate of fibrinolysis.
A significant difference was also detected among the BMA surfaces after 1
h of lysis, in which 55BMA led to a faster rate of fibrinolysis when compared to
45BMA and 65BMA (p ≤ 0.001) (Figure 4-8 a). Over the rest of the lysis period, the
mean D-dimer concentration of 55BMA was also higher than the other BMA and
glass surfaces, with a significant difference detected after 8 h of lysis (p ≤ 0.05)
(Figure 4-8 b). No significant differences in D-dimer concentration were found
among 45BMA, 65BMA and glass surfaces except at the early stage of lysis (1 h).
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However, it is worth noting that 45BMA led to a slower rate of fibrinolysis as the
mean D-dimer concentration was lower than all the other surfaces during the
intermediate lysis stage (4 h and 8 h). Whereas during the late stage (18 h and 24 h),
65BMA led to a slower rate of fibrinolysis as the mean value of 65BMA was lower
than other BMA surfaces. Thus, a similar release profile of BMA surfaces was
observed at the early stage (1 h) and late stage (18 h and 24 h) of lysis.
We also measured the weight loss of clots during lysis (Figure 4-8 d-e).
Significant differences in weight loss among BMA surfaces were detected at the late
stage of lysis (18 h and 24 h) with less weight loss from 45BMA than 65BMA (p ≤
0.05). When compared to the clots exposed to tPA and plasminogen in PBS buffer
(Figure 4-8 d-e), the control clots subjected to PBS buffer only showed a little weight
loss over time (Figure 4-8 f-g). A negligible amount of D-dimer was also detected
from control clots of uncoated glass surfaces (0.06 ± 0.005 µg/mL to 2.8 ± 0.287
µg/mL from 1 h to 24 h after lysis) when compared to the plasma level of D-dimer
(0.09 ± 0.008 µg/mL) (Figure 4-8 c). This suggested that the spontaneous fibrinolysis
was not profound under these experimental conditions and that the elevated level of
D-dimer was largely due to the clot lysis by tPA. Moreover, it indicated that our
suspended clot system supplemented with fibrinolytic enzymes was feasible for
assaying clot lysis.
Overall, the initial decrease in the D-dimer concentrations of BMA surfaces
compared to the uncoated glass surfaces indicated a delayed onset of fibrinolysis
during the first hour of lysis. Further, over the rest of lysis period, clots of 55BMA
showed a faster rate of fibrinolysis than other surfaces. Thus, the specific
composition (33% –COOH/ 67% –(CH2)3CH3 ) on 55BMA seemed to contribute to
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a higher clot susceptibility to lysis than lower or higher –COOH ratios presented on
45BMA (25% –COOH) and 65BMA (40% –COOH), respectively.
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Figure 4-8. Release of D-dimer and weight loss over 24 h lysis of clots formed on BMA surfaces compared with uncoated glass surfaces. a) The D-dimer levels of all surfaces after 1 h of lysis. The D-dimer levels over 24 h-lysis period of b) BMA surfaces and c) uncoated glass surfaces and relative control clots. Negligible amount of D-dimer was detected from control clots. The percentage of weight loss over 24 h lysis of clots formed on d) BMA surfaces and e) uncoated glass surfaces, and the control clots relative to f) BMA surfaces and g) uncoated glass surfaces, respectively. Data was presented as mean of three replicates of each surface with SD. * p ≤ 0.001 ** p ≤ 0.05
4.3.5.2 Effect of various surface alkyl groups on fibrinolysis
To evaluate the effect of surface alkyl groups on fibrinolytic potential of
clots, the lysis of clots formed on surfaces exhibiting 33% –COOH (55MMA,
55EMA & 55BMA), and 40% –COOH (65MMA, 65EMA & 65BMA) were
compared respectively. Figure 4-9 shows the D-dimer concentrations and weight loss
of the clots formed on these surfaces over 24 h of lysis. Similarly, after 1 h of lysis, a
faster rate of fibrinolysis was observed from uncoated glass surfaces as indicated by
a significantly higher D-dimer concentration than all other surfaces (p ≤ 0.001)
(Figure 4-9 a-b). Furthermore, 55BMA led to a significantly faster fibrinolysis
compared to 55MMA and 55EMA (Figure 4-9 a), whereas 65EMA also showed a
significantly faster fibrinolysis than 65MMA and 65BMA (p ≤ 0.001) (Figure 4-9 b).
Although no significant differences were detected among 55MMA, 55EMA
and 55BMA for the rest of the lysis period (Figure 4-9 c), a similar pattern was found
between 1 h and 4 h after lysis, in which 55BMA led to a faster rate of lysis while
55EMA led to a lower rate. In addition, the pattern between the late time points 18 h
and 24 h after lysis was also similar, with a faster rate of lysis occurring on 55MMA
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clots compared with a lower rate from 55BMA clots. A transitional pattern of D-
dimer release may be attributed to the difference between early and late time points
at the intermediate stage of lysis (8 h), when 55MMA clots began to display
enhanced lysis. The D-dimer release profile of these surfaces correlated well with the
weight loss measured from 8 h after lysis at which 55MMA showed slightly more
weight loss when compared to 55EMA and 55BMA. At 24 h of lysis, 55MMA
displayed significantly more weight loss compared with 55BMA (p ≤ 0.05) (Figure
4-9 e). The control clots of 55MMA showed more weight loss than that of 55BMA at
24 h of lysis (p ≤ 0.05) (Figure 4-9 g).
Thus, the initial decrease in the D-dimer concentrations compared to the
uncoated glass surfaces indicated that the clots formed on 55MMA, 55EMA and
55BMA surfaces were more resistant to lysis for up to 4 h. The significant
differences in the D-dimer concentration found among 55MMA, 55EMA and
55BMA surfaces after 1 h lysis, indicated that the surface alkyl groups with different
length altered the initial rate of fibrinolysis. It was showed that the clots formed on
33% –COOH/ 67% –(CH2)3CH3 surface were more prone to lysis initially whereas
that formed on 33% –COOH/ 67% –CH2CH3 surface were more resistant to lysis.
For clots formed on 65MMA, 65EMA and 65BMA, no significant
differences were found for the rest of the lysis period except at 1 h of lysis (Figure 4-
9 d). Instead, a similar pattern was found between 4 h and 8 h after lysis with a faster
rate of lysis occurring on 65MMA but a slower rate from 65BMA. In addition, the
pattern between the late time point 18 h and 24 h was also similar with a faster rate
of lysis from 65EMA while a slower rate from 65BMA. No significant differences
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were found in the weight loss of clots formed on these surfaces over time (Figure 4-9
f). In contrast, the weight loss of relative control clots showed a significant increase
from clots on 65MMA than that on 65BMA after 8 h of lysis (p ≤ 0.05) (Figure 4-9
h).
Thus, the initial decrease in the D-dimer concentrations compared to the
glass surfaces indicated that the clots formed on 65MMA, 65EMA and 65BMA
surfaces were also lysed more slowly for up to 4 h. The significant differences in the
D-dimer concentration found among 65MMA, 65EMA and 65BMA after 1 h lysis
not only reconfirmed that the initial rate of fibrinolysis was modulated by surface
alkyl groups but also by the concentration of –COOH groups. The increase in –
COOH ratio on –CH2CH3 bearing surfaces resulted in an increase in lysis
susceptibility (55EMA vs 65EMA; p ≤ 0.001). Whereas intermediate –COOH ratio
(i.e. 33%) on –(CH2)3CH3 bearing surfaces led to a higher susceptibility than low
(25%) or high (40%) ratio (55BMA vs 45BMA, 65BMA respectively; p ≤ 0.001). No
obvious influence of various –COOH ratios was found on –CH3 presenting surfaces
(55MMA vs 65MMA; p = 0.122). This discrepancy in the dependency of –COOH
ratios on surfaces with various alkyl groups suggested that specific functional groups
and their relative compositions had a major effect on the rate of fibrinolysis and
further supported the changes in the clot structure.
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Figure 4-9. Release of D-dimer and weight loss over 24 h lysis of clots formed on 55MMA, 55EMA and 55BMA surfaces compared with 65MMA, 65EMA and 65BMA surfaces. Compared to uncoated glass surfaces, D-dimer levels after 1 h of lysis of a) 55MMA, 55EMA and 55BMA surfaces, and b) 65MMA, 65EMA and 65BMA surfaces. The D-dimer levels over 24 h-lysis period of c) 55MMA, 55EMA and 55BMA surfaces, and d) 65MMA, 65EMA and 65BMA surfaces. The percentage of weight loss over 24 h lysis of clots formed on d) 55MMA, 55EMA and 55BMA surfaces, and e) 65MMA, 65EMA and 65BMA surfaces, and the control clots relative to f) 55MMA, 55EMA and 55BMA surfaces and g) 65MMA, 65EMA and 65BMA surfaces, respectively. Data was presented as mean of three replicates of each surface with SD. * p ≤ 0.001 ** p ≤ 0.05
Relationship between fibrin structures and clot susceptibility to fibrinolysis
Since the fibrin architectures at the clot exterior and interior are different,
we investigated the impact of fibrin structure modification on fibrinolysis using a
suspended clot system. In this case, the exogenous fibrinolytic enzymes would
initially interact with fibrin at the clot exterior and lysis would proceed from the clot
exterior to interior. Our results demonstrate that all material-coated surfaces lead to a
significantly slower fibrinolysis in the first hour of lysis compared to the uncoated
glass surfaces. This slower onset of fibrinolysis is in good agreement with the tight
network and thicker fibrin observed on the clot exterior on coated surfaces, in
accordance with previous studies which indicated that fibrinolysis occurs
predominantly faster on loose network and thinner fibrin (Gabriel et al., 1992, Collet
et al., 2000, Mullin et al., 2001).
Importantly, significant differences in the fibrinolytic rate after 1 h of lysis
in different material surface groups: BMA surfaces with different –COOH ratios,
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surfaces with different alkyl groups combined with 33% –COOH, or with 40% –
COOH, correlated strongly with surface-dependent differences in fibrin thickness
and density at the clot exterior. For instance, a significantly faster fibrinolysis rate on
65EMA clots compared with clots formed on 65BMA and 65MMA surfaces, is
associated with the ascending fibrin densities at the clot exterior. While a faster
fibrinolysis on 55BMA compared to 45BMA and 65BMA is associated with the
significantly thinner fibre observed at the clot exterior of 55BMA. Moreover,
increasing –COOH ratios on –CH2CH3 surfaces with a decrease in fibrin density at
the clot exterior may also account for increasing susceptibility to fibrinolysis. In
contract, an increase in –COOH ratios on –CH3 surfaces with increasing fibrin
density did not result in a significant difference in the initial rate of fibrinolysis.
These results suggest that the composition of the surface, –COOH groups and their
combination with –CH3, –CH2CH3 or –(CH2)3CH3 groups have a specific activity
on regulating initial rate of fibrinolysis through influence on the fibrin thickness and
density at the clot exterior.
After 4-8 h of lysis, we believe that the enzymatic lysis proceeds to the inner
part of the clots. This is illustrated by a slightly more D-dimer released from 45BMA
than other BMA surface during this period, correlating with the significantly thinner
fibres in the clot interior of 45BMA. Also, a shift of lysis pattern observed after 8 h
of lysis in which a slightly more D-dimer released from 55MMA than 55EMA
correlated well with a significantly lower density in the clot interior of 55MMA than
55EMA. Hence, the changes in D-dimer level from these clots during 4-8 h after
lysis may reflect the gross alteration of fibrin architectures from the clot exterior to
the interior.
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4.3.6 Quantification of PDGF-AB in serum and during clot lysis
Given that the clots formed on material-coated surfaces showed differences
in structural network, stability and lysis susceptibility, we next investigated whether
material-coated surfaces influences the biological function of clots for bone repair.
This was assessed by detecting the release of reparative PDGF-AB and TGF-β1 in
supernatant serum after clot formation and in the buffer during clot lysis.
The serum level of PDGF-AB after 2 h whole blood incubation with
surfaces was shown in Figure 4-10. Significantly elevated PDGF-AB level was
found on all surfaces compared to plasma baseline (308 ± 49 pg/mL), confirming the
growth factor is released upon clot formation. The mean values of 65MMA (8017 ±
330 pg/mL) and uncoated glass surfaces (8029 ± 689 pg/mL) was higher than the
other surfaces.
Among 65MMA, 65EMA & 65BMA surfaces, a significantly lower level of
PDGF-AB was released from 65EMA compared to 65MMA (p ≤ 0.05). For surfaces
with 33% –COOH (55MMA, 55EMA & 55BMA), 55MMA resulted in a
significantly lower level of PDGF-AB than 55EMA and 55BMA (p ≤ 0.001), but
there was no differences between the latter (p = 0.88). In addition, no significant
differences were found among the BMA surfaces.
Overall, 65MMA resulted in a significantly higher amount of PDGF-AB
released from intact clots than all other material surfaces.
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Figure 4-10. The serum levels of PDGF-AB after 2 h of whole blood incubation with material-coated surfaces compared to the uncoated glass surfaces and the plasma baseline. Data was presented as mean of triplicates of each surface with SD.* p ≤ 0.001 ** p ≤ 0.05
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The in vitro release of PDGF-AB during clot lysis was illustrated in Figure
4-11. Among the degrading clots of 55MMA, 55EMA and 55BMA surfaces, a
significantly higher amount of PDGF-AB was released from 55MMA throughout the
lysis period compared to the others (p ≤ 0.001) (Figure 4-11 a). A burst release was
observed after 1 h of lysis and peaked at 8 h with approximately 7-fold more than the
others. The level of growth factor remained high up to 24 h. In contrast, both 55EMA
and 55BMA showed similar release curves with no significance difference found.
In contrast, for 65MMA, 65EMA and 65BMA, similar release patterns were
observed with a peak at 8 h of lysis but overall 65MMA led to a higher amount of
PDGF-AB with a significant difference found at 1 h (p ≤ 0.001) and 4h after lysis (p
≤ 0.05) (Figure 4-11 b).
Degrading clots of BMA surfaces exhibited no difference in PDGF-AB
release profile except at the end of lysis period (24 h) 45BMA resulted in a
significantly higher level than the others (p ≤ 0.001) (Figure 4-11 c). The control
clots of glass surfaces showed higher level of PDGF-AB than that underwent
enzymatic lysis after 1 h (p ≤ 0.001) and 24 h of lysis (p ≤ 0.05) (Figure 4-11 d),
implying the presence of tPA and plasminogen may reduce the release of PDGF-AB.
Overall, the release of PDGF-AB during lysis was significantly elevated
from the clots formed on 55MMA than all other surfaces. However, the majority of
PDGF-AB was released during formation of intact clots instead of clot lysis.
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Figure 4-11. In vitro release of PDGF-AB during lysis of clots formed on material-coated surfaces and uncoated glass surfaces. Among clots formed on a) 55MMA, 55EMA and 55BMA surfaces, b) 65MMA, 65EMA and 65BMA surfaces, c) BMA surfaces and d) uncoated glass surfaces and the relative control clots. Data was presented as mean of triplicates of each surface with SD.* p ≤ 0.001 ** p ≤ 0.05
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4.3.7 Quantification of TGF-beta 1 in serum and during clot lysis
The serum level of TGF-β1 after 2 h whole blood incubation with surfaces
was shown in Figure 4-12. The clots formed on all surfaces released a significantly
elevated level of TGF-β1 compared to plasma level (1021 ± 31 pg/mL).
However, no significant differences were found among BMA surfaces
though the mean levels of TGF-β1 increased with increasing –COOH ratios.
Similarly, no difference was found among 65MMA, 65EMA and 65BMA surfaces,
nor among 55MMA, 55EMA and 55BMA surfaces. Also, uncoated glass surfaces
showed no difference to all the other surfaces.
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Figure 4-12. The serum levels of TGF-β1 after 2 h of whole blood incubation with material-coated surfaces compared to the uncoated glass surfaces and plasma baseline. Data was presented as mean of triplicates of each surface with SD.* p ≤ 0.001 ** p ≤ 0.05
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The in vitro release of TGF-β1 from the degrading clots was shown in
Figure 4-13. In general, the level of TGF-β1 increased over time for all surfaces.
Among the degrading clots formed on 55MMA, 55EMA and 55BMA surfaces, a
similar release pattern was observed initially up to 4 h of lysis. At 8 h of lysis,
55MMA showed an approximate two fold increase of TGF-β1 with up to 24 h of
lysis when compared to the others (p ≤ 0.001) (Figure 4-13 a).
The degrading clots of 65MMA, 65EMA and 65BMA also showed a similar
release pattern over the entire lysis period. At 1 h and 4 h of lysis, 65BMA showed a
significantly higher level of TGF-β1 than 65EMA (p ≤ 0.05) (Figure 4-13 b).
Among the BMA surfaces, the degrading clots of 55BMA and 65BMA
showed a very similar pattern of releasing a significantly higher amount of TGF-β1
than 45BMA at 1 h and 4 h of lysis. By 8 h of lysis, the level of TGF-β1 released
from the clot of 65BMA was also significantly higher than that of 55BMA (p ≤
0.001), and remained higher than the others by 24 h (p ≤ 0.05), and no difference was
found then between 55BMA and 45BMA (Figure 4-13 c). Unlike PDGF-AB, the
clots of uncoated glass surfaces subject to enzymatic lysis released a significantly
higher level of TGF-β1 than the control clots over the lysis period (p ≤ 0.001)
(Figure 4-13 d).
In summary, TGF-β1 was released in significantly higher amounts from the
degrading clots of 55MMA than all other surfaces. However, the majority of TGF-β1
was also released during formation of intact clots than during clot lysis.
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Figure 4-13. In vitro release of TGF-β1 during lysis of clots formed on material-coated surfaces and uncoated glass surfaces. Among clots formed on a) 55MMA, 55EMA and 55BMA surfaces, b) 65MMA, 65EMA and 65BMA surfaces, c) BMA surfaces and d) uncoated glass surfaces and the relative control clots. Data was presented as mean of triplicates of each surface with SD.* p ≤ 0.001 ** p ≤ 0.05
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Relationship between growth factor release, clot structure and early healing events
It has been well documented that various growth factors are expressed in
different phases of bone healing (Cross and Mustoe, 2003). In particular, the
initiation of bone regeneration is suggested to begin with the release of PDGF-AB
and TGF-β1 after a clot is formed (Hosgood, 1993, Kells et al., 1995, Lieberman et
al., 2002b). PDGF-AB is most abundant in platelet α-granules and is known to
support chemotaxis and proliferation of fibroblasts, smooth muscle cells as well as
endothelial cells, resulting in collagen synthesis and angiogenesis (Oprea et al., 2003,
Andrae et al., 2008). On the other hand, TGF-β1 is predominant in platelets, bone
and cartilage, and is shown to serve as a mitogen for osteoblasts, fibroblasts and
endothelial cells, as well as an inhibitor of osteoclasts (Zhang et al., 2005, Bosetti et
al., 2007). In addition, both PDGF-AB and TGF-β1 are chemotactic for
inflammatory cells such as neutrophils, monocytes or macrophages, which
establishes a positive feedback loop of growth factors within the injured bone
(Ashcroft, 1999). In view of their function in supporting bone healing, we evaluated
the potentials of modifications in fibrin structure and fibrin structure-dependent
fibrinolysis on affecting the release of these growth factors from the intact and
degrading clots.
Upon clot formation, increased amounts of PDGF-AB and TGF-β1 were
detected in serum compared to platelet-poor plasma. Interestingly, we found that
there was a correlation between the amount of PDGF-AB released in serum and
fibrin density at the clot exterior and interior. This may be related to fibrin-platelet
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interactions and the subsequent platelet-mediated clot retraction that normally occurs
in wound healing (Marieb, 2001b).
It has been shown that a direct interaction between fibrin fibres and platelet
GPIIb/IIIa receptors allows activated platelets to develop contractile forces on fibrin
with the microfilaments in their pseudopodia. These forces condense the fibrin
network, leading to clot retraction to approximately 1/10 of its original volume and
thereby expelling the entrapped the fluid content to the surrounding milieu
(Morgenstern et al., 1984, Katori et al., 2005). Hence, our observations of different
fibrin structures in clots formed on various surfaces may affect clot retraction,
resulting in a different extent of clot content expulsion. This phenomenon is also
supported by Carr and Zekert (1994), which reported an altered fibrin structure
affecting clot retraction.
The higher density of fibrin found throughout the clots of 65MMA when
compared to 65EMA and 65BMA may lead to a stronger retractile force and a higher
extent of clot content expulsion, correlating with the higher level of PDGF-AB
released from the clots of 65MMA. Furthermore, the clots of 65MMA which
released the highest amount of PDGF-AB in serum among all material-coated
surfaces might also be associated with its highest density of fibrin at the clot exterior
which retracts the clot very strongly. Similarly, a significantly lower level of PDGF-
AB released by 55MMA than 55EMA and 55BMA might be due to a lower fibrin
density found in the clots of 55MMA which retracts weakly. These findings suggest
that the surface functionalities and relative ratios influence the release of PDGF-AB
from intact clots through modification of fibrin network density.
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During clot lysis, the amount of PDGF-AB release also appeared to be
dependent on clot retraction. We found that the intact clots of 55MMA which
released the least amount of growth factor showed a burst release of PDGF-AB
during clot lysis, reaffirming a lower fibrin density associated with a weak clot
retraction retains PDGF-AB entrapped. Moreover, the increased release of growth
factor from the clots of 55MMA correlated well with its increased fibrinolytic rate
over time. For other intact clots which previously released a considerable amount of
PDGF-AB, they generally showed a reduced release during clot lysis. Only the intact
clots of 65MMA which released the highest amount of PDGF-AB continued to
release a significantly higher amount compared to 65EMA and 65BMA in early lysis
period, suggesting the clots of 65MMA originally has a higher content of growth
factor.
On the other hand, we found that the release of TGF-β1 from degrading
clots increased gradually over time. The amount of TGF-β1 released during lysis
seems to correlate with fibrin thickness.
We found that the degrading clots of 65BMA released more TGF-β1
compared to 65EMA up to 4 h of lysis. This is associated with the significantly
thicker fibrin of 65BMA than 65EMA at both clot exterior and interior. Moreover,
among BMA surfaces, although the clots of 65BMA showed the most dense and the
thickest fibres at the clot exterior associating with the slowest fibrinolysis at early
lysis, it also released an amount of TGF-β1 as high as that of 55BMA, which
underwent a faster fibrinolysis due to its thin fibres at the exterior. This suggests that
the fibrin thickness has a major influence on the release amount of TGF-β1.
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In fact, fibrin fibres has been characterised to bind a range of growth factors
and adhesive proteins by FXIII-crosslinking, such as FGF-2, VEGF and fibronectin
(Standeven et al., 2005, Weisel, 2005, Mosesson, 2005). Fibronectin has been shown
to bind TGF- β1 with high affinity (Mooradian et al., 1989). Also, TGF- β1 binds
thrombospondin-1 (Murphy-Ullrich et al., 1992, Crawford et al., 1998, Blair and
Flaumenhaft, 2009), which in turn interacts directly with fibronectin,
fibrinogen/fibrin and plasminogen (Wencel-Drake et al., 1985, Panetti et al., 1999).
Hence, these findings imply that TGF- β1 is potentially bound to fibrin fibres. With
the assumption that the amount of TGF- β1 bound to fibrin is directly proportional to
the fibrin thickness, and the fact that fibrin fibres are transversally cleaved rather
than uniformly around during fibrinolysis (Veklich et al., 1998, Collet et al., 2000),
this might explain the correlation between the increased fibrin thickness and the
increased amount of TGF-β1 detected during lysis, and why no significant
differences were seen in the release of TGF-β1 from intact clots.
Taken together, our results showed that more growth factors were released
during clot formation than during clot lysis. From the intact clots, the release of free
growth factor (i.e. PDGF-AB) seems to be strongly influenced by fibrin density,
which alters clot retraction. On the contrary, this influence appeared to be less
significant on the release of fibrin-bound growth factor (i.e. TGF-β1). Instead, during
clot lysis, the release of TGF-β1 is more likely to be associated with the fibrin
thickness while the release of PDGF-AB appeared to be generally associated with the
fibrinolytic rate. Our findings suggest the clots formed on material-coated surfaces
serve as a natural system as a haematoma to provide localised and substantial
releases of growth factors.
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4.4 CONCLUSION
By using an incubation vial where the inner surface was coated with
material, the effect of surface chemical functionalities –COOH/–CH3, –CH2CH3 or –
(CH2)3CH3 and their ratios on whole blood response was studied in a three-
dimensional manner and in the context of clot formation.
The surface functionalities and their ratios showed a modulatory effect on
coagulation response. Surfaces with –COOH/–(CH2)3CH3) induced a faster rate of
coagulation response compared to –COOH/–CH3 and –CH2CH3 regardless of the –
COOH ratios. An increase in –COOH ratio on –COOH/–CH3 and –CH2CH3
surfaces decreased the rate of initiation. However, all material-coated surfaces
resulted in clot formation. All coated surfaces markedly reduced complement
response when compared to uncoated glass surfaces. An entirely similar pattern of
coagulation and complement response was observed on material-coated surfaces. All
coated surfaces resulted in thicker fibrin with a tighter network at the clot exterior
when compared to uncoated glass surfaces. For all surfaces, the interior of clots
showed thicker fibres with loose network when compared to the clot exterior.
Surfaces presenting same –COOH ratio but different alkyl groups showed a
consistent trend in fibrin density throughout the clots. Material-coated surfaces
produced more rigid clots with significantly slower onset of fibrinolysis when
compared to uncoated glass surfaces, which is consistent with the thicker fibres and
tighter network observed on the clot exterior of coated surfaces. Similarly, the
significant differences in fibrinolytic rate of coated surfaces after 1 h of lysis
correlated well with the surface-dependent differences in fibrin thickness and density
at clot exterior. Generally, more PDGF-AB and TGF-β1 were released during clot
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formation than during clot lysis. From an intact clot, 65MMA (40% –COOH/60%–
CH3) released the highest amount of PDGF-AB compared to those of other material
surfaces.
To our best knowledge, this is the first study which provides a more
comprehensive picture of how surface functional groups and their concentrations
considerably modulate blood cascade activation in the context of whole blood clot
formation; subsequent fibrin architecture, clot rigidity, susceptibility to fibrinolysis
and growth factor entrapping/release ability of the modified clots. Further studies are
required to determine whether these changes in fibrin clot structure and function
have any therapeutic relevance to bone regeneration.
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5Chapter 5: A Pilot Study of the Osteogenic
Properties of Ex vivo Blood Clots Formed on
Materials in a Rabbit Femoral Defect
5.1 INTRODUCTION
A blood clot formed at the injured bone is vital in initiating the natural bone
healing response (Tsiridis et al., 2006, Tosounidis et al., 2009). During the
implantation of synthetic bone implants, disruption of bone vasculature and tissue
injury causes blood contact with the implant and ultimately the formation of a peri-
implant clot (Anderson, 2001, Gorbet and Sefton, 2004). Despite its resemblance to
the clot formed normally on injured bone, the structure and properties of peri-implant
clots have not been studied for their effect on the healing capacity of artificial bone
implants. Rather, most work in bone tissue engineering focuses on using synthetic
scaffolds containing osteogenic factors or in combined with PRP gels (Sánchez et al.,
2003, Grageda, 2004, Stevens, 2008, Schliephake, 2009).
Results of the previous chapters have shown that surface functionalities and
their relative compositions on biomaterials can modulate the structural properties of
whole blood clots including fibrin architecture and rigidity, as well as biological
properties such as susceptibility to fibrinolysis and the release of PDGF-AB and
TGF-β1 from clots during intact and degrading stages.
To validate the concept that the blood clots formed on surface coatings
provide the essential microenvironment for bone healing, and that changes in clot
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structure and properties have any clinical implications, a rabbit femoral defect model
was established in this pilot study. In particular, an autologous clot formed on
65MMA surface (40% –COOH/60% –CH3) was employed. From our previous
results, compared to other surface chemistry, this combination of surface
functionalities triggered lower complement activation while also initiating
coagulation response, resulting in the formation of a clot. Moreover, the resultant
clots contained the highest density of thin fibrin fibres, resulting in a high clot
rigidity and increased resistance to fibrinolysis. Moreover, the highest amount of
PDGF-AB was released upon clot formation, and considerable amounts of PDGF-
AB and TGF-β1 were also released during clot lysis.
Ultraporous beta tricalcium phosphate (β-TCP) is a ceramic that has been
widely used as bone fillers in dentistry and in orthopaedics. Its similar chemical
composition to bone precursors, high biocompatibility, mechanical properties and
porous structure are believed to favour bone ingrowth by preferentially promoting
infiltration of cells, vascularisation, mineralisation and resorption for bone recovery
(Johnson et al., 1996, Annaz et al., 2004, Giannoudis et al., 2005, Walsh et al., 2008,
Van Lieshout et al., 2011).
In this chapter, we aimed to confirm the idea that a blood clot generated on
a controlled surface provides a beneficial microenvironment for bone regeneration
when compared to commercial β-TCP bone graft substitute or an empty defect. By
using a rabbit femoral defect model, the in vivo osteogenic potentials of different
treatments were assessed by the extents of calcification, new bone formation,
chondrogenesis as well as vascularisation and inflammation.
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5.2 MATERIALS & METHODS
5.2.1 Preparation of coatings on scaffold surfaces
Stainless steel mesh scaffolds were fashioned with an arm on each side of it
(Figure 5-1). The arms of scaffolds were used to clip around the femur bone for implant
anchorage. The scaffolds were sterilized by soaking in 70% ethanol for 10 min and dried
in a laminar flow hood. The scaffolds were then dip-coated with material solutions and
air-dried. This procedure was repeated three times to achieve multiple coating and
present surface functionalities as on glass surfaces. A visible film was formed across the
centre of the scaffolds. After drying, the film was punched with a hole with a sterile
needle (21G, 0.8x38 mm; TERUMO corporation). This assures clotting over the coated
surface area. The coated scaffolds were kept in tubes in which the internal surfaces were
also coated with the same type of material by solvent evaporation. This was done to
ensure a similar functionalised surface area to blood volume ratio as used in the previous
experiments.
Figure 5-1. Stainless steel scaffold coated with material solution. Arrow pointing to a hole in the middle of film formed across the scaffold. An arm was on each side of the scaffold.
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5.2.2 Animals
Experiments were performed in a total of three New Zealand white rabbits (6-8
month old, 3.9 - 4.6kg) of which the animal ethics were approved by the QUT Animal
Ethics Committee. All animals were housed in the Prince Charles Hospital animal
facility. General anaesthesia was induced after premedication with an intramuscular
injection of atrophine (0.04 mg/kg; AstraZeneca, Australia), by a slow intravenous
injection of ketamine (35mg/kg; Ketalar, Hospira Australia Pty Ltd., Australia) and
midazolam (2.5 mg/kg; Sandoz Pty Ltd., Australia). During the surgery ketamine and
xylazine (5 mg/kg; Troy Laboratories Pty Ltd., Australia) were used in maintenance of
the anaesthesia. After 4 weeks, animals were euthanized by intraperitoneal injection of
1.5 mL sodium pentobarbital. This method of euthanasia was in accordance with
National Health and Medical Research Council of Australia guidelines.
5.2.3 Ex vivo blood clot formation
To prepare ex vivo blood clots on the coated scaffolds, autologous blood was
collected from the marginal ear vein or artery of animals undergoing surgery (Figure 5-2
a). After the rabbits were anaesthetized, fur covering the ear was shaved and the ear was
sterilized with 4 % Chlorhexidine. Whole blood of 4 mL was taken from the tip of the
ear and immediately transferred to the tubes containing the scaffolds (Figure 5-2 b-c).
The tubes were incubated in a water bath at 37˚C for approximate 90-120 min whilst the
defects on the femur were induced.
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5.2.4 Surgical procedures in rabbit femur
An incision was made longitudinally along the midline of the shaved skin
area over the femur. Underlying musculature was dissected to expose the shaft.
Defects of diameter 3.2 mm were made using a bur driven by a drill. Care was taken
during drilling to avoid deep penetration into the marrow cavity. Excessive blood
was removed regularly from the defect sites.
Three defects were made on the femur approximately 1 cm apart and one
femur was operated for each animal (Figure 5-2 d). The defect holes were treated in
the following three ways: ex vivo blood clots formed on scaffolds, ChronOSTM
porous β-TCP granules (60% porosity, pore width 100-500µm, size of granules 0.7-
1.4 mm; Synthes GmbH, Australia) or Surgicel (NU-KNIT, absorbable haemostat,
oxidized regenerated cellulose, sterile; ETHICON, INC., Australia) (Table 5-1). For
the implantation of ex vivo blood clots on scaffolds, the clots were ensured to fill the
defect holes and secured to the position by clipping the arms of scaffolds around the
femur (Figure 5-2 e). Surgicel was served as negative control as it achieved
haemostasis but it prevented formation of complete clots in the defects. After the
implantations, the musculature and skin were closed by suture. Antibiotic therapy
(Alamycin, 200mg/mL; Norbook Laboratories, Australia) and analgesics
(Buprenorphine 324 µg/mL; Reckitt Benckiser Ltd, New Zealand) were
administrated in the immediate postoperative period. The animals were sacrificed
after 4 weeks.
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Figure 5-2. Implantation of ex vivo blood clots formed on material-coated scaffolds in rabbit femoral defects. a) Autologous blood was collected from the marginal ear vein or artery; b) Material-coated scaffold was placed in a tube in which the internal surfaces was coated by the same type of material solution. c) Collected blood was immediately transferred to the tube and incubated at 37˚C. d) Three defects were made on the femur by drilling, and e) clot-embedded scaffold was placed on the defect with clot filling the hole.
Treatment Groups
Ex vivo blood clots on scaffolds (65MMA clots)
ChronOSTM β-TCP granules
Surgicel
Table 5-1. Three treatment groups in the animal study.
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5.2.5 Examination of defects
5.2.5.1 Assessment of calcification by micro-computered tomography
Harvested femurs were fixed with 10% buffered formalin for one week.
Overlying tissue and stainless steel scaffolds were removed after preliminary
examination of the femur with a micro-CT scanner (μCT 40, Scanco Medical,
Switzerland). Cross sectional scans of the femurs were then performed and three
dimensional (3D) images were reconstructed from the scans using the micro-CT system
software package. To evaluate in vivo calcification in the defect areas, the total volume
and the average density against hydroxyapaptite (HA) of calcification within the defect
areas were measured and recorded for statistical analysis. Defect areas were defined by
selecting the area of interest between the lesion edges on each micro-CT slice.
5.2.5.2 Assessment of new bone formation by haematoxylin and eosin staining
After micro-CT analysis, the femurs were segmented according to the defect
sites and decalcified in PBS buffered 10% EDTA (Ethylenediamine tetraacetic acid
Di-sodium salt; AJAX Finechem, Australia). Decalcified samples were washed with
PBS and dehydrated in ethanol as previously described. A midline cut was performed
longitudinally on the defect area on each segment prior to paraffin embedding.
Embedded samples were sectioned at 5 µm on the sagittal plane with a microtome
(Leica Microsystems GmbH, Germany). Sections near the central sagittal plane were
used for all examinations. Paraffin sections were dewaxed three times in xylene for 3
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min intervals, and rehydrated through 100%, 90% and 70% ethanol and distilled
water for 3 min intervals.
For haematoxylin and eosin (H & E) staining, sections were stained with
Mayer’s Haematoxylin (HD Scientific Supplies Pty Ltd., Australia) for 2 min and
excessive stain was rinsed away with running water. After dehydration through 70%,
90% and 100% ethanol the sections were exposed to Eosin stain (HD Scientific
Supplies Pty Ltd., Australia) for 15 s. Excessive stain was removed by dipping in
100% ethanol. Sections were then processed 3 times in xylene for 3 min intervals and
mounted with DePex mounting medium (VWR International Ltd., England). De novo
bone formation in the defect sites was observed under a microscope (Carl Zeiss Inc.,
USA) and the images were captured. Area of new bone tissue per mm2 was measured
in five random fields using AxioVision software.
5.2.5.3 Assessment of chondrogenesis by Alcian Blue staining
Dewaxed and rehydrated sections were first stained with 3% acetic acid for 3
min followed with Alcian blue solution (1% alcian blue in 3 % acetic acid, Fronine
Laboratory Supplies, Australia) for 1 h. After rinsing excessive stain, the sections were
stained with Nuclear fast red (Fronine Laboratory Supplies, Australia) for 5 min and
excessive stain was rinsed. The sections were dehydrated, rinsed with xylene, and
mounted as previously described before observed under the microscopy (Carl Zeiss Inc.,
USA).
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5.2.5.4 Assessment of new bone formation by immunohistochemistry
To identify osteogenic cells in bone newly formed in the defects, an alkaline
phosphatase (ALP) antibody (mouse anti-rat; Santa Cruz Biotechnology Inc., USA) was
used in immunohistochemical staining. Briefly, following dewaxing and rehydration of
the sections, endogenous peroxidise activity of the tissue was quenched by incubating
with 3% H2O2 for 20 min. The sections were blocked with 10% donkey serum (Sigma-
Aldrich) in PBS-Triton at room temperature for 1 h. The blocking solution was aspirated
and ALP primary antibody (1:100) in fresh block solution was added to the section for
overnight at 4˚C. After that, the sections were washed once in PBS-Triton for 3 min and
twice in PBS for 3 min intervals. Biotinylated donkey anti-mouse secondary antibody
(1:200) was then added to the sections for 20 min at room temperature. After washing,
horseradish peroxidase-conjugated avidin-biotin complex (ABC) was added to the
sections for 20 min with excessive complex washed. Antibody bindings were visualised
by the addition of diaminobenzidine (DAB) substrate solution for 3 min. Development
of brown colour was stopped by rinsing the sections with running water. The sections
were counterstained with Mayer’s haematoxylin (HD Scientific Supplies Pty Ltd.,
Australia) for 15 s followed by dehydration, xylene rinsing and mounting. Images were
taken with a microscope (Carl Zeiss Inc., USA).
5.2.5.5 Assessment of vascularisation and inflammation by immunofluorescence
To evaluate the vascularisation and inflammation in the defects, von
Willebrand factor (vWF) antibody (mouse anti-human; Upstate Cell Signaling
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Solution, Australia) and CD68 antibody (mouse anti-rat; Abcam, Australia) were
used to identify the endothelial cells and macrophages, respectively. In brief,
following dewaxing, rehydration and blocking the sections were incubated with vWF
(1:200) or CD68 (1:200) primary antibody at 4˚C overnight. After washing, the
sections were incubated with Alexa Fluor® 488 labelled donkey anti-mouse
secondary antibodies (1:200; Molecular Probes, Australia) overnight at room
temperature. Washed sections were mounted with Vectashield mounting medium
(Vector Laboratories, USA) and observed under Zeiss Z1 ApoTome fluorescence
microscope (Carl Zeiss Inc., USA). Images were captured and the number of vWF+
or CD68+ cells per mm2 were quantified from six random fields using AxioVision
software. Antibodies used in this study were summarized in Table 5-2.
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Primary Antibody Stain For Company Dilution
Mouse anti-rat alkaline phosphatase (ALP)
Osteogenic cells Osteoblasts
Santa Cruz Biotechnology
Inc 1/100
Mouse anti-human von Willebrand factor (vWF) Endothelial cells
Upstate Cell Signaling Solution
1/200
Mouse anti-rat CD68 (CD68) Monocytes/macrophages Abcam 1/200
Secondary Antibody Colour Company Dilution
Biotinylated donkey-anti mouse
Brown Santa Cruz
Biotechnology Inc
1/100
Alexa Fluor ® 488 donkey anti-mouse Green
Molecular Probes 1/200
Table 5-2. Primary and secondary antibodies used in immunohistochemistry and immunofluorescence studies.
5.2.6 Statistical analysis
Analysis was performed using SigmaPlot (version 11.0; Systat software Inc).
All data were analysed using one-way analysis of variance (ANOVA) for group
differences. The significance level was set at p ≤ 0.05.
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5.3 RESULTS
5.3.1 Micro-CT analysis for calcification in the defects
Segments of defects treated with ex vivo blood clots formed on 65MMA-coated
scaffolds, ChronOSTM granules or Surgicel were shown in Figure 5-3 a-c. Based on
general observation, Surgicel showed very limited healing in the defect compared to
others as demonstrated by red tissue fragments in the defects (Figure 5-3 a-c). Micro-CT
scanning and 3D reconstruction of images were performed to calculate the total volume
and average density of the calcified areas within the defects. In Figure 5-3 d-f, the
defects treated with 65MMA clots or ChronOSTM granules showed significantly higher
volume of calcified areas than that with Surgicel, in which hardly calcification was
found (p ≤ 0.001). However, no significant differences were found between 65MMA
clots and ChronOSTM granules (p=0.0557) (Figure 5-3 g). In addition, no significant
differences were found in the average density of calcified areas among all defects
(p=0.221) (Figure 5-3 h). This was likely due to the radio-opaque nature of ChronOSTM
granules.
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Figure 5-3. Micro-CT scanning analysis on the femoral defects. Representative pictures of segmented femoral defects filled with a) an ex vivo blood clot formed on 65MMA-coated scaffold; b) ChronOSTM granules; c) Surgicel. The 3D reconstruction images of the defect sites corresponding to the treatments a-c) were shown as d-f). g) Measurement of the volume of calcified area in defect sites showed that treatments with 65MMA clots and ChronOSTM porous β-TCP granules resulted in a significantly higher volume of calcified area than with Surgicel, while there was no significant differences in average density among different treatments. Data was presented as the mean ± SD from three measurements. * p ≤ 0.001; Scale =1 mm
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5.3.2 Histological examination of de novo bone formation with H & E staining
The bone regeneration in defects evaluated by H&E staining differed
profoundly. More trabecular bone was formed and bridged the gap in the defects
treated with 65MMA clots (Figure 5-4 a). Much less reparative bone spicules were
identified in the defects treated with ChronOSTM granules. In the space of the
spicules, the myeloid elements of bone marrow were seen as bluish patches. The
positions of ChronOSTM granules that were removed during decalcification were
readily identifiable by the irregular empty gaps around spicules (Figure 5-4 b). The
defects treated with Surgicel remained empty. Bone healing response was restricted
to the defect edges (Figure 5-4 c).
The area of new bone tissues per mm2 in the defects was measured (Figure
5-4 d-f). The defects with 65MMA clots showed a significantly higher density of de
novo bone tissues than other treatments (Figure 5-4 g; p ≤ 0.001). These findings
suggested that the 65MMA clots led to formation of more reparative bones as a result
of faster bone remodelling.
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Figure 5-4. De novo bone formation in the defects shown by H&E staining. Representative images of defects treated with a) an ex vivo blood clot formed on 65MMA-coated scaffold; b) ChronOSTM granules; c) Surgicel (40x) (Scale=500µm). d-f) A higher magnification (100x) views of area seen in the box of a-c) respectively (Scale=200µm). g) Measurement of the de novo bone formation in the defects showed a significantly higher area of new bone per mm2 was formed in the defects treated with 65MMA clots compared to ChronOSTM granules and Surgicel; Data was presented as the average area ± SD from five measurements. * p ≤ 0.001
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5.3.3 Histological examination of chondrogenesis in the defects
The extent of chondrogenesis during new bone formation in the defects was
illustrated by Alcian blue staining (Figure 5-5 a-c). Chondrogenesis was stronger and
more extensive in the defects treated with 65MMA clots than other treatments. Zones of
hypertrophic chondrocytes were found in the top centre of the defects (Figure 5-5 f). The
edges of newly formed trabecular bone and bone spicules in the defects treated with
65MMA clots and ChronOSTM granules also showed some positive blue staining,
indicating the remnants of cartilage. Such chondrogenesic response was not seen in the
defects treated with Surgicel (Figure 5-5 d-g). Overall, these results indicated that the de
novo bone regeneration in defects treated with 65MMA clots and ChronOSTM granules
occurred through cycles of chondrocytes hypertrophy as in natural bone healing.
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Figure 5-5. Chondrogenesis in the defects shown by Alcian Blue staining. Representative images of defects treated with a) an ex vivo blood clot formed on 65MMA-coated scaffold; b) ChronOSTM granules; c) Surgicel (40x) (Scale=500µm). d-g) A higher magnification (100x) views of area seen in the box of a-c) (Scale=200µm). f) Zones of hypertrophic chondrocytes were observed in the defects with 65MMA clots, indicating a stronger response of chondrogenesis than ChronOSTM granules and Surgicel.
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5.3.4 Histological examinations of ALP expression
ALP staining was performed to identify osteoblasts within the defect sites
(Figure 5-6 a-c). In the defects with 65MMA clots, a greater abundance of ALP+
osteoblasts were lining the periphery of trabecular bone and zones of hypertrophic
chondrocytes when compared to those with ChronOSTM granules, in which less ALP+
osteoblasts were around the bone spicules (Figure 5-6 d-f). No ALP+ expression was
observed in the defects with Surgicel (Figure 5-6 g). Consistent with more newly formed
bone shown by H&E staining, these results further confirmed osteogenesis was
remarkably enhanced by 65MMA clots than other treatments.
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Figure 5-6. Osteoblasts in the defect shown by ALP staining. Representative images of defects treated with a) an ex vivo blood clot formed on 65MMA-coated scaffold; b) ChronOSTM granules; c) Surgicel (40x) (Scale=500µm). d-g) A higher magnification (100x) views of area seen in the box of a-c) (Scale=200µm). More ALP+ osteoblasts were shown surrounding the newly formed bone in the defects treated with d) 65MMA clots than e) ChronOSTM granules. f) ALP+ osteoblasts were also found lining the periphery of zones of hypertrophic chondrocytes in the defects with 65MMA clots, confirming an enhanced osteogenesis response. g) No ALP+ expression was detected in the defects with Surgicel.
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5.3.5 Vascularisation revealed by vWF
To determine the vascularisation of the defect sites after treatment, vWF+
endothelial cells and blood vessels were counted using immunofluorescence. vWF+
endothelial cells and blood vessels appeared as circular structure were identified in
the defects with 65MMA clots (Figure 5-7 a, d, g) and ChronOSTM granules (Figure
5-7, b, e, h). Unexpectedly, aggregates of activated platelets which display vWF were
also seen in defects with Surgicel (Figure 5-7 c, f, i). Quantitative analysis showed
that 65MMA clots resulted in a significantly higher density of vWF+ endothelial cells
and blood vessels than ChronOSTM granules (p ≤ 0.001), indicating a higher degree
of vascularisation in the defects. No obvious vascularisation was found in the defects
with Surgicel (Figure 5-7 j).
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Figure 5-7. Vascularisation of the defects shown by vWF staining. Representative immunofluorescence images stained with von Willibrand factor (vWF, green) and nucleus (DAPI, blue) of defects treated with a) an ex vivo blood clot formed on 65MMA-coated scaffold; b) ChronOSTM granules; c) Surgicel (100x) (Scale =200µm). Images at d-f) 200x (scale = 100 µm) and g-i) 400x (scale = 50 µm) magnification for a-c), respectively. vWF+ endothelial cells and blood vessels were seen (pointed by red arrows) in defects treated with g) 65MMA clots and h) ChronOSTM granules. i) No obvious vascularisation was seen in defects treated with Surgicel. Only clusters of activated platelets were positively stained with vWF. j) Measurement of the number of vWF+ endothelial cells and blood vessels per mm2 showed that a significantly higher degree of vascularisation was found in defects treated with 65MMA clots compared to ChronOSTM granules and Surgicel. Data was represented as the mean ± SD from six measurements. * p ≤ 0.001
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5.3.6 Inflammation response revealed by CD68
The extent of inflammation response induced by different implants was
determined by counting macrophages identified by marker CD68. The defects treated
with ChronOSTM granules showed significantly more CD68+ macrophages that were
merged together (red arrows in Figure 5-8 b, e, h) when compared to those with
65MMA clots (Figure 5-8 a, d, g) and Surgicel (Figure 5-8 c, f, i) (p ≤ 0.001). Given
that there was hardly or if any bone formation within the defects treated with
Surgicel, the presence of 65MMA clots resulted in a significantly lower degree of
inflammation response with bone formation compared to ChronOSTM granules
(Figure 5-8 j).
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Figure 5-8. Inflammatory response evaluated by the number of macrophages at the defects using CD68 staining. Representative immunofluorescence images stained with CD68 antibody (green) and nucleus (DAPI, blue) of defects treated with a) an ex vivo blood clot formed on 65MMA-coated scaffold; b) ChronOSTM granules; c) Surgicel (100x) (Scale =200µm). Images at d-f) 200x (scale = 100 µm) and g-i) 400x (scale = 50 µm) magnification for a-c), respectively. j) Measurement of the number of CD68+ macrophages (red arrow) per mm2 showed that a significantly higher density of macrophages was found in defects treated with ChronOSTM granules than with 65MMA clots and Surgicel. Data was represented as the mean ± SD from six measurements. * p ≤ 0.001
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5.4 DISCUSSION
In this proof of concept study, we have demonstrated that ex vivo blood
clots formed on 65MMA-coated scaffolds significantly enhanced bone regeneration
through normal healing process. The 65MMA clots led to increased new bone
formation, stronger chondrogenesis and more osteoblasts activity compared to the
commercial bone substitute: ChronOSTM β-TCP granules. Moreover, we did observe
a beneficial effect of 65MMA clots on increasing vascularisation while reducing
inflammatory response compared to other treatments. When blood clot formation
was disrupted by the presence of Surgicel, the defects remained empty after 4 weeks.
This affirmed that a blood clot provides the vital microenvironment for bone healing.
We observed faster and abundant woven bone ingrowth in defects treated
with 65MMA clots verse less bone formation with bone spicules in defects filled
with ChronOSTM granules. Moreover, it was showed that ChronOSTM granules
resulted in some empty cavities around the bone spicules whereas Micro-CT
scanning showed calcification as high as 65MMA clots. These findings indicate that
ChronOSTM granules (60% porosity) remained present at 4 weeks after implantation.
The superiority of 65MMA clots over ChronOSTM granules in accelerating bone
regeneration is probably related to its fibrin structure and growth factor contents.
In a rabbit tibial defect, ChronOSTM granules were found to be present up to
26 weeks and its slower resorption was associated with decreased new bone
formation compared to other β-TCP granules with same chemical composition but a
faster resorption profile (Walsh et al., 2008). Reduced porosity and interconnection
of β-TCP granules were demonstrated to correlate with decreased granule
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degradation and subsequent bone formation (Hing et al., 2005, Knabe et al., 2008).
On the contrary, we previously showed that 65MMA clots contain a heterogeneous
fibrin network structure, suggesting the clots would be relatively porous and
interconnected. Moreover, 65MMA clots which are physically rigid and resistant to
early fibrinolysis would therefore provide a stable osteoconductive scaffold for
infiltration of osteoprogenitor cells into the cavity. In addition, the interior of
65MMA clots are less dense in fibrin, which is a fully resorbable biological matrices.
This suggests that 65MMA clots would be readily dissolved in vivo as implicated by
our previous lysis study when compared to ChronOSTM granules. Hence, the
differences in implant porosity and interconnectivity in relation to in vivo resorption
profile likely explains the faster and higher quantity of bone formation with 65MMA
clots whereas ChronOSTM granules appeared to lag behind due to its long-lasting that
potentially limited new bone ingrowth.
On the other hand, ChronOSTM β-TCP granules, like all ceramic scaffolds,
does not have any osteoinductive properties (Hing, 2005, Hak, 2007). In contrast, a
blood clot is well characterised with its inherent chemotactic, angiogenic and
osteogenic factors (Street et al., 2000, Sánchez et al., 2003, Frechette et al., 2005,
Blair and Flaumenhaft, 2009). We have previously revealed that of the clots
generated in vitro, 65MMA clots released the highest level of PDGF-AB and
considerable amount of TGF-β1 upon clot formation than clots generated on other
surfaces. Substantial release of both growth factors were also observed during clot
lysis.
In fact, the ability of PDGF-AB and TGF-β1 to promote bone healing has
been demonstrated in many animal studies (Lind et al., 1996, Geiser et al., 1998,
Vehof et al., 2002, Ehrhart et al., 2005). In the presence of artificial implants, PDGF-
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AB has also been shown to play a crucial role in bone healing. An increased release
of PDGF-AB as a result of enhanced platelet activation on titanium are believed to
be attributed to the enhanced osteogenic properties of titanium (Hong et al., 1999).
As PDGF-AB contributes to half of the mitogenic activity for mesenchymal cells
(Andrae et al., 2008), PDGF-AB released from 65MMA clots may contribute to the
extensive chondrogenic differentiation observed in the defects compared to
ChronOSTM granules by stimulating mitogenesis of bone marrow stem cells and
mesenchymal cells from periosteum of broken bone ends. Moreover, the stronger
chondrogenesic response with 65MMA clots indicates that these clots are capable to
direct normal bone healing sequence similar to a physiological haematoma, to
endochondral ossification through a cartilaginous stage. Furthermore, the higher
number of vWF+ endothelial cells and vessel-like structures present in the defects
with 65MMA clots compared to ChronOSTM granules may be resulted from the
chemotactic and angiogenic properties of PDGF-AB on fibroblasts, smooth muscle
cells and endothelial cells (Lieberman et al., 2002c, Carano and Filvaroff, 2003).
For TGF-β1, its major mechanism on enhancing bone healing is achieved by
mediating chemotaxis, mitogenesis and differentiation of osteoblasts precursor at the
injured sites. This may be attributed to more osteoblasts found in the defects with
65MMA clots than ChronOSTM granules (Hosgood, 1993, Oprea et al., 2003, Celotti
et al., 2006, Bosetti et al., 2007). Moreover, the potential of TGF- β1 in enhancing
implant osteointegration has also been demonstrated on titanium surfaces through
stimulating differentiation and mineralization of human osteoblasts on the surfaces
(Zhang et al., 2005). Given that PDGF-AB acts synergistically with TGF-β1 to
promote osteoblasts growth (Kells et al., 1995, Lieberman et al., 2002b), and
thrombin generated from coagulation activation is able to stimulate osteoblasts
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proliferation (Frost et al., 1999), it is conceivable that the defects filled with entirely
65MMA clots would present more residual thrombin and growth factors for
osteoblast activity than that with ChronOSTM granules. Therefore, our demonstration
of altered structure and growth factor release from 65MMA clots may contribute to
the advanced osteogenesis and angiogenesis in the defects compared to ChronOSTM
granules.
In term of inflammatory response, less CD68+ macrophages were found in
the defects filled with 65MMA clots than with ChronOSTM β-TCP granules. Study of
Hing et al. (2007) indicated the dissolution of β-TCP granules releases degradation
products which subsequently provoked an inflammatory response in rabbit
osteochondral defects. In an analogous investigation by Arvidsson et al. (2011) they
also found that calcium phosphate compound did increase inflammatory response.
Furthermore, using a rat subcutaneous model, Jansson et al. (2002) suggested that
cells at the titanium surface with a plasma clot layer were in a different stage than
those without clot and that these cells were not in a phagocytotic phase as shown by
reduced production of reactive oxygen. It is proposed that the presence of a plasma
clot-coated surface will likely preserve the radical-mediated degradation or killing
capacity for a longer period than those without the clot. Furthermore, Barbosa et al.
(2004, 2005) who used a rat air pouch model to study surface functionalities on in
situ inflammatory response found that –CH3 SAMs recruited similar number of
inflammatory cells as –OH SAMs, but most of the cells were neutrophils and only a
very low density of cells adherent to the surface in contrast to –OH SAMs where
most of the cells were monocytes and a higher density of adherent cells were
detected. These results indicated that the –CH3 groups likely induce neutrophils-
dominated local acute inflammation but is not associated with a significant leukocyte
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adhesion to implant. In line with this, we previously found that 65MMA surfaces
induces a low extent of complement response. This may be accounted for the acute
inflammation which is necessary in normal healing process but not a significant
chronic inflammation up to 4 weeks when compared to ChronOSTM β-TCP granules.
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5.5 CONCLUSION
Using a rabbit femoral defect model, we demonstrated that the ex vivo
65MMA clots provide the essential microenvironment for new bone formation. The
implantation of 65MMA clots led to advances in vascularisation, chondrogenesis,
osteoblasts activity and increased new bone formation compared to ChronOSTM β-
TCP granules. The impact of surface functionalities and their ratios on modulating
ultimate fibrin clot structure and growth factor release were likely contributed to
these differences through mediating infiltrations and activities of reparative cells in
the defect. Moreover, we observed a lower extent of inflammatory response with
65MMA clots compared to ChronOSTM granules, which were probably associated
with lower complement activation in vitro upon clot formation, as well as particulate
products released upon granules degradation. These findings emphasize that 65MMA
serves as a prothrombogenic and immunocompatible surface that alters structural and
biological properties of a consequent clot, which in turn did impact on new bone
regeneration in vivo.
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Chapter 6: General Discussion and Future
Directions
6.1 GENERAL DISCUSSION
It is commonly accepted that the formation of a blood clot at injured bones
is the first step in bone healing process. During bone implantation, blood contacts
with biomaterials and subsequently the formation of a clot around the implant occurs
prior to bone regeneration. Despite both situations being similar in forming a blood
clot at the defect site, there has been no attempt in controlling the biomaterial-blood
interactions and clot formation, such that the result may be beneficial in bone healing
approaches. It is theoretically suggested that applications of PRP gels could improve
de novo bone formation due to its increased contents of growth factors (Roukis et al.,
2006, Griffin et al., 2009, Cenni et al., 2010). While this was borne out in preclinical
experimentation, effective translation to the clinic has been unsuccessful. There has
been limited work addressing the fact that PRP gels contain different cell populations
(i.e. platelets only), and fibrin structure dramatically altered from a normal
haematoma, not to mention any assessment of such differences related to conflicting
effect of PRP gels on bone healing. Indeed, changes in the cellular content and fibrin
structure of a peri-implant clot can greatly affect osteoconduction, which is a vital
phase of peri-implant endosseous healing (Davies, 2003b). The clot architecture
which depends on thrombin concentration plays an important role in various
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thrombotic disorders considering its impact on clot viscoelasticity and degradability
(Collet et al., 2000, 2006, Mills et al., 2002, Wolberg and Campbell, 2008).
In addition to the issues with integration and bone formation, in vivo
function of many implants are often compromised by the blood-based immune
response leading to FBR (Anderson et al., 2008). As such, a control of blood-
biomaterial interactions by modifying biomaterial surface properties has been the
focus in developing an improved biocompatible device. Surface roughness,
hydrophobicity and chemical composition have been shown to influence blood
protein adsorption, cellular interactions, and subsequent immune complement
responses to biomaterials (Albrektsson and Wennerberg, 2004, Anderson et al.,
2004a, Ma et al., 2007, Thevenot et al., 2008, Nilsson et al., 2009). In particular, it
was demonstrated that surface presenting mixtures of –COOH and –CH3 functional
groups at varied ratios modulated profoundly coagulation and complement
activations, as well as extent of fibrin deposition on the surfaces, compared to the
surfaces with only either groups (Sperling et al., 2005a, 2009, Fischer et al., 2010b).
Furthermore, it has been shown that different chain length of alkyl groups provides a
further level to regulate these biomaterial-blood interactions (Berglin et al., 2004,
2009). Theoretically, these factors could modulate healing events and consequently
the extent of new bone formation in injured site. Hence, the rationale of this study
was to create surfaces with –COOH/–CH3, –CH2CH3 or –(CH2)3CH3 functionalities
at an optimal ratio to promote coagulation activation, produce a blood clot with
appropriate cellular content and structure properties while diminish adverse immune
response to the implants, and thus improve the new bone formation in the presence of
synthetic bone grafts.
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To scrutinise this hypothesis, we first fabricated a series of materials
composed of acrylic acid and alkyl methacrylate (MMA, EMA or BMA)) to create a
prothrombogenic and immunocompatible coating of artificial bone implants (Chapter
3). Unlike SAMs on gold substrate that is not feasible clinically, the materials used in
this work have been specifically selected due to their biological relevance. For
instance, PMMA and AA have been widely used as bone cement and protein-
immobilising agent, respectively (Kang et al., 1993, Lu, 2004). The resultant
materials are also shown to be non-toxic while possessing desired functionalities
(Yan and Gemeinhart, 2005). Our results clearly indicated that the inner surface of
incubation vial was modified effectively by material coating. Such coatings based on
materials can be adapted to most of the current biomedical materials without
affecting bulk properties. Our strategy of varying the types and mole fractions of
comonomers in forming materials was ascribed to different surface functional
groups: –COOH/–CH3, –CH2CH3 or –(CH2)3CH3) at different compositions. An
increase in AA proportion generally increased the surface content of –COOH groups
but the content of –COOH groups was lower than the expected AA fraction. Similar
observations were also seen in other studies and might be attributed to different
degrees of copolymerisation, polymer chain mobility and functional group
reorientation (Gupta et al., 2002, Berglin et al., 2009, Xu et al., 2009, Hermitte et al.,
2004, Ozcan and Hasirci, 2007).
Our results demonstrated that surface hydrophobicity of coated surfaces
correlates well with the chemical compositions, as in accordance to the literature
(Tsyganov et al., 2005, Ukiwe et al., 2005, Lai et al., 2006). At relatively the same –
COOH ratios, surfaces presenting –(CH2)3CH3) groups exhibited a higher
hydrophobicity than –CH3 and –CH2CH3 groups whereas the latter two did not differ
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significantly. This suggests that both the surface functional groups and their relative
compositions synergistically influence surface hydrophobicity (Tsai et al., 2007).
Moderate hydrophobicity of our surfaces is consistent with what is commonly
suggested to be the optimal than extremely hydrophilic or hydrophobic surfaces as
they support protein adsorption with preserved conformation (Hermitte et al., 2004).
Moreover, we observed that the material coatings were as smooth as
uncoated surfaces and without apparent cracks. The average surface roughness was
in nano-scale in which material coating (3.99 ± 0.54 nm) was slightly higher than
uncoated glass surfaces (2.22 ± 0.29 nm). Recently, in vivo studies have shown that
nano-scale roughness enhanced early stage of healing but micro-scale roughness
improved overall bone-implant contact and bone density of peri-implant endosseous
healing (Telleman et al., 2010, de Barros et al., 2011). Microstructure has also been
demonstrated to greatly increase fibrinogen adsorption, subsequent conformation
changes and platelet responses compared to sub-micron structure (Xie et al., 2009,
Koh et al., 2010a, 2010b). Another studies of Ferraz et al. (2008, 2010) demonstrated
that complement activation was stronger on surface with 200 nm pore sizes rather
than 20 nm whereas within this nano-range platelet response was influenced
differently over time. So far, it remains largely unknown whether nano-scale
roughness as low as 2-4 nm would have any impact on protein adsorption or cellular
interaction. Our results showed no significant differences in average roughness
among surfaces with same functionalities at different –COOH ratios nor among
surfaces with different alkyl groups but the same –COOH ratios. This suggests the
surface functional groups and their relative ratios do not influence the surface
roughness. Hence, these results imply that any difference in blood response is less
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likely a result of different surface roughness but surface functionalities and relative
compositions.
To link the surface chemical formulation of the material coating and blood
response, we carried out a comprehensive investigation of whole blood interaction on
the customised incubation vials (Chapter 4). The uses of the incubation vials and
whole blood were designed to mimic closely the formation of three-dimensional peri-
implant clots during surgery, allow blood cascades to cross-talk and avoid the impact
of anticoagulant in interrupting the measurement of blood response.
The prothrombogenic properties of surfaces was assessed by the rate of
coagulation initiation. Surfaces presenting –COOH/–(CH2)3CH3) groups showed a
faster rate of coagulation activation compared to those with –COOH/–CH3 and –
CH2CH3 groups, regardless of the –COOH ratios. Specifically, increasing –COOH
ratios on surfaces with –COOH/–CH3 and –CH2CH3 decrease the rate of the
activation. Analysis of complement activation revealed that all material-coated
surfaces induced a weaker response compared to uncoated surfaces, clearly
indicating these surfaces had a reduced immunogenic property. In addition, the
complement response followed an entirely similar pattern of surface-activated
coagulation, suggesting our in vitro incubation system allows an interaction between
these two cascades to take place as it is found in vivo (Markiewski et al., 2007,
Peerschke et al., 2008, Amara et al., 2010). This also reflects the acute inflammation
caused by the implanted biomaterials will occur naturally following thrombotic
events.
Examination of resultant blood clots showed that the material-coated
surfaces modulated the fibrin architecture resulting in a thicker fibre at tighter
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configuration at the clot exterior when compared to uncoated glass surfaces. In
addition, our observation of the clot interior revealed that fibrin architectures were
also different from the exterior for all surfaces in which fibres were thicker and with
very loose network. Differences in fibre thickness and density was also detected on
material-coated surfaces, reaffirming the specific effects of surface functionalities
and their relative ratios on controlling the kinetic of coagulation initiation and
subsequent fibrin structure throughout the clots. The heterogeneous morphology of
fibrin structures was also noted in other studies and was attributed to surface
functionalities-dependent differences in pattern of thrombin generation leading to
different kinetics of fibrinopeptides cleavage and fibrin polymerization (Geer et al.,
2007, Wolberg and Campbell, 2008), as well as different procoagulant potential of
entrapped cells in the clots in organising the fibrin bundles (Gugutkov et al., 2011).
To verify in vivo stability of the altered clot structure which is important for
physical support at the injured sites and subsequent new bone ingrowth, we measured
the clot rigidity and the rate of clot lysis. Material-coated surfaces were shown to
change the clot mechanical properties leading to more rigid clots than those formed
on uncoated glass surfaces. Also, it was shown that all coated surfaces resulted in a
slower initial fibrinolysis (i.e. first hour after lysis) when compared to uncoated glass
surfaces. These findings were consistent with the tighter network and thicker fibrin
observed on the clot exterior on material-coated surfaces. Similarly, surface-
dependent differences in fibrin thickness and density at the clot exterior were also in
good agreement with the difference in initial fibrinolytic rates observed among
coated surfaces. These results suggests that the material-coated surface modulates the
clot susceptibility to fibrinolysis by changing fibrin architectures in the clots.
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Hoi Ting Shiu Page 188
As an assessment of biological function of the altered clots for enhancing
bone regeneration, we measured the content of PDGF-AB and TGF-1 released upon
clot formation and lysis. Both growth factors were released in higher levels during
clot formation than during clot lysis. Our results revealed that from an intact clot the
release of PDGF-AB correlated with the fibrin density at the clot exterior and
interior, resulting in the highest level of PDGF-AB being released from clots formed
on 65MMA (40% –COOH/60% –CH3). The correlation of PDGF-AB released upon
clot formation with the surface-dependent difference in fibrin density is likely due to
fibrin-mediated clot retraction which expels the content entrapped within the clots
(Morgenstern et al., 1984, Carr and Zekert, 1994). During lysis, however, the release
of PDGF-AB seemed to correlate with fibrinolytic rate, whereas the release of TGF-
β1 was likely influenced by the surface-dependent fibrin thickness, implicating TGF-
β1 was fibrin-bound and PDGF-AB was soluble within the clots (Mooradian et al.,
1989, Murphy-Ullrich et al., 1992). Overall, these results suggested that the growth
factor release of the clots was modulated by surface-dependent fibrin architecture
and fibrinolysis. Next, we questioned whether such modification of resultant clots
have any functional impact in bone healing. This would implicate whether the clots
generated in vitro on material-coated surfaces can form the basis of truly therapeutic
agents.
To answer the question above, we implanted ex vivo 65MMA clots into the
rabbit femoral defect model and analysed after 4 weeks (Chapter 5). Implanting
65MMA clots into the defects led to significantly stronger chondrogenesis, increased
new bone formation, more osteoblasts activity as well as increased vascularisation
when compared to ChronOSTM porous β-TCP granules. Unlike β-TCP granules
which is known to be devoid of osteoinductive properties, 65MMA clots which was
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previously shown to release a high amount of growth factors such as PDGF-AB and
TGF-β1 were likely to be responsible for the enhanced osteogenesis and
angiogenesis. In contrast, negative control with Surgicel which inhibits blood clot
formation resulted in a empty defect, indicating the clots formed on material-coated
surfaces provide the essential healing microenvironment as normal haematoma does.
Previous studies indicated that the slow in vivo resolution profile associated with
ChronOSTM granules resulted in significantly lower amount of new bone formation
(Walsh et al., 2008, Hing, 2005, Knabe et al., 2008). This implies that the advance in
bone formation with 65MMA clots are possibly attributed to a better fibrin porous
structure and faster degradation profile compared to ChronOSTM granules. In
particular, 65MMA clots resulted in less macrophages in the defects compared to
ChronOSTM β-TCP granules over the time course of study. This affirms our previous
demonstration of 65MMA surfaces which had less immunogenic effect as it induced
a low extent of complement activation in vitro and did not show a potential in
promoting chronic inflammation nor FBR in vivo. In contrast to normal implant
coating which would be broken down by proteolytic enzymes from neutrophils in
initial inflammation and expose underlying synthetic materials that would evoke
FBR, our strategy of pre-formed blood clots around material-coated scaffolds could
completely hide the materials, mimic ECM for cell recognition and be relatively
resistant to premature enzymatic degradation.
Altogether, this project opens the new scope of blood clots generated on
various surface functionalities for treating severe bone injuries. Our studies on the
effect of surface chemistry on blood clots were initiated based on the notion that the
normal mechanism of bone healing could be useful for enhancing the healing
microenvironment in the presence of synthetic bone implants. As far as we can
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Hoi Ting Shiu Page 190
ascertain, this is the first study that demonstrates systematically surface functional
groups and their relative ratios on material coatings modulate initiation of blood
cascades in the context of whole blood clot formation; and subsequent fibrin
architecture, clot rigidity, susceptibility to lysis and growth factor entrapping/release.
Importantly, these clots generated on material-coated surfaces are in many ways
comparable to the natural haematoma, in both structural and functional aspects. The
provision of such pre-formed blood clots may recreate the healing microenvironment
and serve as therapeutic agents for improved bone regeneration.
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Hoi Ting Shiu Page 191
6.2 FUTURE DIRECTIONS
The present study raises many questions of phenomenological differences
observed and an extension of this study is to unravel the underlying mechanisms of
differences in clot structure on various compositions of implant surfaces. First question
is how different combinations of surface functional groups at varied ratios affects the
early mechanism of blood coagulation resulting in different fibrin thickness and density
as detected. Hence, future studies will apply circular dichroism spectropolarimetry to
investigate the dynamic interactions between surfaces and adsorbed protein in term of
surface coverage and conformational change. Furthermore, the blood transition of flow
to stasis will be characterised by using the technique of thromboelastography (TEG),
which records the overall coagulation profile, the time of initial fibrin formation, overall
clot strength and the dynamic properties between fibrin and platelet bonding via GP
IIb/IIIa (Lai et al., 2010).
In addition, the clinical implications of our material-coated surfaces on
supporting bone healing, apart from 65MMA, remain to be further defined. A larger
scale of animal study with critical sized defects is required to clarify any therapeutic
differences in the clots generated on various material surfaces, in order to establish the
corresponding surface chemical formulation for the development of a truly
prothrombogenic and immunocompatible synthetic bone grafts.
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Hoi Ting Shiu Page 192
6.3 GENERAL CONCLUSION
In summary, this thesis has demonstrated that materials composed of AA
and MMA, EMA or BMA at varied mole fraction provided solid coatings displaying
–COOH/–CH3, –CH2CH3 or –(CH2)3CH3 functionalities at varied ratios. Although
the material-coated surfaces yielded a lower content of –COOH groups compared to
the AA fraction fed in the material, the measured surface hydrophobicity correlated
well with the corresponding chemical composition. Surface functionalities and their
relative ratios on material coatings had no effect on surface average roughness. The
roughness of material-coated surfaces (3.99 ± 0.54 nm) was higher than uncoated
glass surfaces (2.22 ± 0.29 nm). Upon whole blood contact with coated surfaces on
incubation vials, the COOH/–(CH2)3CH3 functionalities significantly increased the
rate of coagulation initiation than other functionalities at all ratios. All material-
coated surfaces significantly reduced the complement activation than uncoated glass
surfaces. The similar pattern of material surface-mediated complement and
coagulation activation suggests that there is interaction of the two cascades.
Moreover, all material-coated surfaces produced clots with thicker fibrins and tighter
network at the exterior than uncoated glass surfaces, while the clot interior of all
surfaces contained thicker fibrins with very loose network than the clot exterior. This
affirms coated surfaces control the kinetics of coagulation initiation and alter fibrin
architectures in the clots. Material-coated surfaces resulted in more rigid clots with a
significantly slower onset of fibrinolysis than that of uncoated glass surfaces,
indicating the material coatings change clot mechanical property and stability.
Significant difference in fibrinolytic rate after 1 h of lysis among coated surfaces
were consistent to the surface-dependent differences in fibrin thickness and density at
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clot exterior, suggesting the surface chemistry influences clot susceptibility to lysis
by modifying the clot fibrin architectures. PDGF-AB and TGF-β1 were released in
greater amounts during clot formation than during clot lysis. In addition, the amount
of PDGF-AB released from intact clots correlated with the fibrin density at both
exterior and interior of the clots. During clot lysis, the release of PDGF-AB was
influenced by fibrinolytic rate whereas that of TGF-β1 was associated with the
surface-dependent fibrin thickness. Furthermore, in a rabbit femoral defect, the
implantation of ex vivo clots generated on 65MMA surfaces (40% –COOH/60% –
CH3) significantly increased vascularisation, chondrogenesis, osteoblast activity and
new bone formation but reduced chronic inflammatory response when compared to
ChronOSTM β-TCP granules. The empty cavities in defects treated with Surgicel
which inhibits proper clot formation versus abundant new bone ingrowth in defects
filled with 65MMA clots indicated that the blood clots provide the vital healing
microenvironment for bone regeneration. Overall, this study has emphasized the
important role of surface functionalities and their relative ratios on controlling the
initiations of blood cascades, the structural properties of resultant clots and the
ultimate effect of the clots on bone healing. These results explore the future potential
of applying blood clot regulation by various material coatings to improve the
efficacy of synthetic bone grafts.
Bibliography Page 194
Hoi Ting Shiu Page 194
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