NANOFIBER BASED SCAFFOLD FABRICATION, CHARACTERIZATION

AND OPTIMIZATION FOR TISSUE ENGINEERING AORTIC HEART VALVE

EHSAN FALLAHI AREZOUDAR

A thesis submitted in fulfilment o f the

requirements for the award o f the degree o f

Doctor o f Philosophy (Mechanical Engineering)

Faculty o f Mechanical Engineering

Universiti Teknologi Malaysia

FEBRUARY 2017

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To my beloved family

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ACKNOWLEDGEMENT

First and above all, I praise God, the almighty for providing me this

opportunity and granting me the capability to proceed with this research successfully.

Further, there are no proper words to convey my deep gratitude and respect for my

thesis and research advisor, Professor Dr. Noordin Mohd Yusof. He has inspired me

to become an independent researcher and helped me realize the power of critical

reasoning. He also demonstrated what a brilliant and hard-working scientist can

accomplish.

My sincere thanks must also go to co-advisory, Professor Dr. Ani Idris for the

trust, the insightful discussion, offering valuable advice, for her support during the

whole period of the study, and especially for her patience and guidance during the

writing process. She generously gave her time to offer me valuable comments toward

improving my work.

Besides, I would like to thank the authority of Universiti Teknologi Malaysia

(UTM) for providing me with a good environment and facilities. I also greatly

appreciate the excellent assistance and spiritual supports of my family and my friends

during my PhD study.

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ABSTRACT

The four valves in a mammalian heart provide a unidirectional, unobstructed blood flow pathway as a result of synchronic movement of valves’ leaflets during cardiac cycle. When one of the valves malfunctions, the medical choice is to replace the original valve with an artificial one. However, the inability to grow or to remodel an artificial valve leads to the innovation of tissue engineering heart valve (TEHV). The previously tissue engineered heart valve tends to be rigid, have low degradation rate and adverse structure which leads to TEHV failure. This study presents the design and fabrication of an aortic heart valve (AOHV) based on tissue engineering (TE) principle via electrospinning method. In TE, a three-dimensional (3D) scaffold with proper design, structure, and mechanical properties that resembles the original tissue is required as an initial template for tissue regeneration. For this purpose, materials’ ratio tuning and process optimization as well as the 3D scaffold design were considered. Initially, five different ratios of poly-L-lactic acid (PLLA)/thermoplastic polyurethane (TPU) blends containing 1% (w/v) maghemite (y-Fe2O3) nanoparticles were electrospun and characterized in terms of morphology, degradation rate, biological compatibility and mechanical properties. The existence of three components in the mats was confirmed by Fourier transform infrared and energy-dispersive X-ray spectroscopy. Scanning electron microscopy images illustrated well fabricated nanofibers with smaller diameter distribution for PLLA. The MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide) assay using human skin fibroblast cell indicates desired proliferation on the samples. Blood biocompatibility results in terms of clotting time, fibrin formation, and hemolysis were almost in the normal range. Samples’ degradation rate was investigated over 24 weeks where the PLLA shows 47.15% loss in mass versus 6.7% loss for TPU. High tensile strength and an extremely low elongation-at-break were determined from the stress-strain curve for PLLA, while TPU exhibits high elasticity. Overall, 50:50% of (1% y-Fe2O3) loaded PLLA/TPU mats are the most appropriate. Next, a two-level Taguchi (L8) experimental design followed by the response surface methodology (RSM) were used to optimize the fabrication process where the elastic modulus is the response while the factors investigated were A-flow rate (2-3 ml/h), B-voltage (20-30 kV), C- maghemite% (1-3% w/v), D-solution concentration (10-15 wt.%) and E-collector rotating speed (1000-2000 rpm). From the signal-to-noise ratio values, the influences of the factors were ranked as: D>B>C>E>A. The empirical quadratic model obtained consists of the voltage-B and second order effect of flow rate-(A)2, voltage-(B)2, maghemite %-(C)2 and concentration-(D)2. The optimum elastic modulus of the scaffold was found to be 35.24±0.64 MPa. Finally, an AOHV template was designed and installed as the electrospinning collector to fabricate the 3D scaffold based on the optimum ratio and settings. Later, the human aortic smooth muscles cell migration and proliferation, as well as the elastic modulus loss percent of the optimum 3D scaffold after cell seeding were checked during 34 days of incubation. Overall, the structural, biological and mechanical specifications of the fabricated TEHV have successfully proved that it can be a potential alternative in AOHV replacement surgery.

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ABSTRAK

Empat injap yang terdapat di dalam hati mamalia menyediakan laluan aliran darah yang searah, tidak terhalang disebabkan pergerakan daun injab yang diselarikan semasa kitaran jantung. Apabila salah satu daripada injap rosak, pilihan perubatan adalah menggantikan injap asal dengan injap tiruan. Walaubagaimanapun, injap tiruan tidak mempunyai kemampuan untuk tumbuh atau dimodel semula. Ini telah membawa kepada inovasi injap jantung kejuruteraan tisu (TEHV). Injap jantung kejuruteraan tisu sebelum ini adalah tegar, mempunyai kadar penurunan yang rendah dan struktur yang tidak sesuai yang membawa kepada kegagalan TEHV. Kajian ini membentangkan reka bentuk dan fabrikasi injap jantung aortik menggunakan prinsip kejuruteraan tisu (TE) melalui kaedah elektropintal. Dalam TE, perancah tiga dimensi (3D) dengan reka bentuk yang sesuai, struktur dan sifat-sifat mekanik yang boleh menyerupai tisu asal akan digunakan sebagai pencontoh permulaan untuk pertumbuhan semula tisu. Untuk tujuan ini, penalaan nisbah bahan-bahan utama yang digunakan dan pengoptimuman proses serta reka bentuk perancah 3D dipertimbangkan. Buat permulaan, lima nisbah berbeza poli-L-laktik asid (PLLA)/poliuretana termoplastik (TPU) dicampurkan dengan 1% (w/v) maghemite (y-Fe2O3) nanopartikel. Campuran ini telah melalui proses elektropintal dan pencirian dibuat dari segi morfologi, kadar penurunan, keserasian biologi dan sifat-sifat mekanik. Kewujudan tiga komponen dalam lapisan serat nano telah disahkan oleh jelmaan inframerah Fourier dan serakan-tenaga X-ray spektroskopi. Imej imbasan mikroskopi elektron menunjukkan bahawa serat nano yang baik terhasil dengan garis pusat yang lebih kecil untuk PLLA. Kajian MTT menggunakan sel fibroblas kulit manusia dan ia menunjukkan percambahan yang baik ke atas sampel. Keputusan bio-keserasian darah dari segi masa pembekuan, pembentukan fibrin, dan hemolisis hampir dalam julat normal. Kadar penurunan sampel telah diselidiki selama 24 minggu yang mana PLLA menunjukkan penurunan jisim sebanyak 47.15% berbanding dengan penurunan 6.7% bagi TPU. Kekuatan tegangan yang tinggi dan kadar pemanjangan sebelum putus yang amat rendah ditentukan dari lengkung tegasan-terikan untuk PLLA, manakala TPU mempamerkan keanjalan yang tinggi. Secara keseluruhan, lapisan serat nano PLLA/TPU yang mengandungi 50:50% daripada 1% (y-Fe2O3) adalah yang paling sesuai. Seterusnya, reka bentuk eksperimen dua aras Taguchi (L8) diikuti dengan kaedah gerak balas permukaan (RSM) telah digunakan untuk mengoptimumkan proses di mana modulus elastik merupakan sambutan manakala faktor yang dikaji ialah A- kadar alir (2-3 ml/h), B- voltan (20-30 kV), C-maghemite % (1-3% w/v), D-kepekatan larutan (10-15wt.%), dan E- kelajuan puteran pengumpul (1000-2000 rpm). Dari nilai nisbah isyarat-kepada-hingar (S/N), pengaruh faktor adalah: D>B>C>E>A. Model kuadratik empirikal yang diperolehi terdiri dari voltan-B dan kesan peringkat kedua kadar alir-(A)2, voltan-(B)2, maghemite %-(C)2 dan kepekatan-(D)2. Modulus elastik optimum perancah yang diperolehi adalah 35.24±0.64 MPa. Akhir sekali, AOHV telah direka dan dipasang sebagai pemungut kepada elektropintal untuk menghasilkan perancah 3D berdasarkan kepada nisbah optimum dan tetapan. Kemudian, penghijrahan dan perkembangan aortik sel-sel otot licin manusia serta peratusan kehilangan keanjalan modulus daripada perancah 3D optimum selepas pembenihan sel diperiksa semasa 34 hari pengeraman. Secara keseluruhan, spesifikasi struktur, biologi dan mekanikal bagi TEHV yang telah difabrikasi berjaya membuktikan yang ia boleh menjadi alternatif yang berpotensi untuk pembedahan penggantian AOHV.

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TABLE OF CONTENTS

CHAPTER TITLE PAGE

DECLARATION ii

DEDICATION iii

ACKNOWLEDGEMENT iv

ABSTRACT v

ABSTRAK vi

TABLE OF CONTENTS vii

LIST OF TABLES xii

LIST OF FIGURES xiv

LIST OF ABBREVIATIONS xix

LIST OF SYMBOLS xxii

LIST OF APPENDICES xxv

1 INTRODUCTION 1

1.1 Overview of the Research 1

1.2 Research Problem Statement 8

1.3 Research Questions 10

1.4 Research Hypothesis 11

1.5 Research Aim and Objectives 11

1.6 Research Scopes 12

1.7 Significance of Research 14

1.8 Organization of Thesis 14

2 LITERATURE REVIEW

2.1 Introduction

15

15

2.2 Overview of Heart Valves 15

2.3 Aortic Heart Valve 18

2.3.1 Microstructure and Function of Normal Aortic

Valve 18

2.3.2 Mechanical Properties of the Aortic Heart Valve 22

2.3.2.1 Uniaxial and Biaxial Mechanical

Properties 23

2.3.2.2 Flexural Mechanical Properties 28

2.3.3 Aortic Valves Pathology 29

2.4 Tissue Engineering Heart Valve (TEHV) 31

2.4.1 Three-dimensional Heart Valves Scaffold 32

2.4.1.1 Polymeric and Biological Materials 32

2.4.1.2 Scaffolds Fabrication Techniques 42

2.4.1.3 Synthetic Scaffolds Mechanical

Properties 47

2.4.2 Cells Sources and Cultivation Methods 48

2.4.2.1 Adipose-derived Cells 50

2.4.2.2 Valve Interstitial Cells 50

2.4.2.3 Bone Marrow Stem Cells 51

2.4.3 Development Condition 52

2.5 Summary 53

3 RESEARCH METHODOLOGY 55

3.1 Introduction 55

3.2 Research Framework 55

3.3 Materials Requirement 59

3.4 Solution Preparation 61

3.4.1 Synthesizing Maghemite (y-Fe2O3) 61

3.4.2 Poly-L-lactic Acid (PLLA) and

Thermoplastic Polyurethane (TPU) 61

3.5 Electrospinning Setup to Fabricate PLLA/TPU-(y-Fe2O3)

Nanofiber Mats 62

3.5.1 Chemically Characterization of Electrospun Mats 64

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3.6 Materials Ratio Tuning Process 64

3.6.1 Morphology and Porosity 65

3.6.2 Hydrophilicity and Surface Roughness 66

3.6.3 Degradation Rate of Electrospun Mats 67

3.6.3.1 Changes in Morphology 67

3.6.3.2 Mass Change of Electrospun Mats 68

3.6.3.3 Porosity (%) Change 68

3.6.4 Cell Biological Compatibility Tests 68

3.6.4.1 Fibroblast Cell Thawing, Plating and

Sub-culturing 69

3.6.4.2 Cell Cytotoxicity Assay and Cell

Viability (MTT Assay) 70

3.6.4.3 Cells Attachment 74

3.6.5 Blood Hemocompatibility 74

3.6.5.1 Blood Clotting Time (PT & TT Assay) 75

3.6.5.2 Fibrin Formation 76

3.6.5.3 Hemolysis Test 76

3.6.6 Mechanical Properties of Electrospun Mats 77

3.7 Fabrication Process (Electrospinning) Optimization 79

3.7.1 Two-level Taguchi Experimental Design 79

3.7.1.1 Test for Significance of the Regression

Model 83

3.7.1.2 Test for Significance on Individual

Model Terms 83

3.7.1.3 Test for Lack-of-fit 84

3.7.2 Steepest Ascent Method 85

3.7.3 Response Surface Methodology (RSM) 87

3.8 Design the 3D Semilunar Heart Valve Template 88

3.9 Characterizations of 3D Heart Valve Scaffold 91

3.9.1 Morphology, Hydrophilicity and Porosity

Characterization 91

3.9.2 Human Aortic Smooth Muscle Cell Viability

Evaluation 92

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3.9.3 Mechanical Evaluation of Seeded Scaffold 94

3.10 Statistical Data Analysis 96

4 RESULTS AND DISCUSSION 97

4.1 Introduction 97

4.2 Maghemite (y-Fe2O3) Nanoparticles Characterization 97

4.3 Chemically Characterization of Electrospun

Nanofiber Mats 99

4.4 Materials Ratio Tuning Process 103

4.4.1 Nanofibers Morphology, Diameter Distribution

and Porosity 103

4.4.2 Hydrophilicity and Surface Roughness 107

4.4.3 Degradation Rate of Electrospun Mats 111

4.4.3.1 Changes in Morphology 111

4.4.3.2 Mass Changes of Electrospun Mats 115

4.4.3.3 Porosity (%) Change 117

4.4.4 Cells Biocompatibility 119

4.4.4.1 Cytotoxicity Assay and Cell Viability

(MTT Assay) 119

4.4.4.2 Cell Attachment 122

4.4.5 Blood Hemocompatibility 126

4.4.5.1 Blood Clotting Time (PT & TT Assay) 126

4.4.5.2 Fibrin Formation and Hemolysis

Percent (HP %) 127

4.4.6 Mechanical Properties 129

4.4.7 Conclusion on Materials Ratio Tuning

Properties 132

4.5 Fabrication Process Optimization 132

4.5.1 Two-level Taguchi Experimental Design 132

4.5.1.1 Confirmation of Taguchi Design 136

4.5.1.2 ANOVA Analysis 137

4.5.1.3 Linear Regression Model 139

4.5.2 Steepest Ascent Pathway 144

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4.5.3 Response Surface Methodology (RSM) 146

4.5.4 Confirmation Test 152

4.6 Three-dimensional Semilunar Heart Valve

Scaffold Fabrication 154

4.7 Characterization of 3D Heart Valve Scaffold 155

4.7.1 Morphology, Hydrophilicity and Porosity

(%) of 3D Scaffold 156

4.7.2 Human Aortic Smooth Muscle Cells Migration

and Proliferation 158

4.7.3 Blood Hemocompatibility Tests 160

4.7.4 Evaluate the Mechanical Properties of Seeded

Scaffold 163

4.8 Summary 166

5 CONCLUSION AND FUTURE DIRECTION 168

5.1 Introduction 168

5.2 Conclusions 168

5.3 Recommendations and Future Direction 171

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REFERENCES

Appendix A

173

202

LIST OF TABLES

TABLE NO. TITLE PAGE

2.1 Uniaxial mechanical properties of human and animal aortic valve 25

2.2 Summary of reported studies of biopolymers in TEHV 37

2.3 Techniques used to process biomaterials in TEHV 45

3.1 List of materials used in the research 60

3.2 Factors and levels for Taguchi experimental design 80

3.3 Layout of Taguchi experimental design orthogonal array 81

3.4 Normal human aortic valve dimensions (mm) at 80 mmHg pressure 89

4.1 Porosity measurement for different ratios of PLLA/TPU-(Y-Fe2O3) 106

4.2 Mean±SD results of water contact angles measurements in 60 sec 108

4.3 Mass change of PLLA/TPU-(y-Fe2O3) samples after immersion (mg) 116

4.4 Mechanical properties results of electrospun mats 131

4.5 Taguchi experimental results in terms of elastic modulus (MPa) 133

4.6 Taguchi design (L8) analysis for controllable factors and factors’ ranking 133

4.7 S/N ratio responses for controllable factors 136

4.8 List of aliased factors and interaction with intercept of ABC and ADE 137

4.9 ANOVA table (partial sum of square) for response surface model (response: elastic modulus) 138

4.10 ANOVA table (partial sum of square) for response surface after adding center points 140

4.11 ANOVA table (partial sum of square) for response surface after model reduction 141

4.12 Steepest ascent pathway design in terms of coded and actual variable 145

4.13 Responses of tensile strength, tensile strain and elastic modulus 146

4.14 Factors and levels for CCD design 147

4.15 Small fraction of CCD layout and responses of elastic modulus (MPa) 147

4.16 ANOVA table for quadratic model after model reduction 149

4.17 Confirmation experiments design and responses for elastic modulus 152

4.18 Water contact angle measurement for 3D optimum scaffold 157

4.19 Macro-indentation test result for 3D optimum scaffold after cell seeding 164

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FIGURE NO.

1.1

1.2

1.3

1.4

2.1

2.2

2.3

2.4

2.5

2.6

2.7

3.1

3.2

LIST OF FIGURES

TITLE

Schematic of (a) heart anatomy (b) valve position (c) single valve leaflet and its ECM structure

Artificial heart valves (a) mechanical (b) biological (c) TEHV

Concept of tissue engineering heart valves (TEHV)

Schematic diagram o f electrospinning setup

Pathway o f blood flow through the heart and lungs. Oxygenated blood flow (red colour) transfer from the lungs to the left ventricles and deoxygenated blood (blue colour) returns from the body to the right ventricle

Structures o f atrioventricular valve (mitral and tricuspid) include the leaflets, annulus, chordae tendineae, and papillary muscles

Schematic semilunar valves (a) top view o f lateral section o f outflow vessel during diastole to show the close valve,(b) by cutting longitudinally the vessel between two leaflets

Schematic diagram of aortic valve leaflet layers and components

Biaxial mechanical behaviour of aortic heart valve leaflets radially and circumferentially

Biological, mechanical and physiochemical properties of commonly studied biodegradable natural and synthetic polymer

Chemical structure formula of (a) poly-L-lactic acid (b) thermoplastic polyurethane

Research framework in (a) summary form (b) detailed form

Electrospinning setup to fabricate PLLA/TPU-(y-Fe2O3) nanofibers

PAGE

2

3

4

7

16

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18

22

27

37

39

56

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3.3

3.4

3.5

3.6

3.7

3.8

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3.10

3.11

3.12

3.13

4.1

4.2

4.3

4.4

4.5

4.6

4.7

4.8

4.9

4.10

Image of hemocytometer gridlines for under microscope cells counting

Layout of samples loading in 24-well plate

MTT assay procedure

Stress versus strain uniaxial tensile test

Linear graph for Taguchi L8 design

First order response surface and path of steepest ascent

Two-dimensional central composite design experiments with level (a)

Human aortic heart valve geometrical parameters

Design of aluminium based template to be used as collector

3D printed based design of aortic heart valve template

Schematic diagram of macro-indentation test

Characterization of (y-Fe2O3) nanoparticles (a) XRD pattern (b) TEM images (c) reference of maghemite XRD pattern

FTIR spectra of (a) TPU, PLLA, PLLA/TPU and PLLA/TPU-(y-Fe2O3) (b) references of PLLA and TPU polymers

EDX graphs of elements presented in PLLA/TPU and PLLA/TPU-(y-Fe2O3)

DSC heating curve of PLLA/TPU and PLLA/TPU-(y- Fe2O3)

FE-SEM images of nanostructure and diameter distribution of (a) 100:0% (b) 75:25% (c) 50:50% (d) 25:75% and (e) 0:100% PLLA/TPU-(y-Fe2 O3) scaffold

TEM image of 50:50% PLLA/TPU-(y-Fe2 O3)

Porosity (%) of different ratios of PLLA/TPU-(y-Fe2O3) electrospun mats

Water contact angle measurement of (a) 100:0% (b) 75:25% (c) 50:50% (d) 25:75% and (e) 0:100% PLLA/TPU-(y-Fe2O3) nanofibers

AFM results of surface roughness, 3D micrograph and region profile for (a) 100:0% (b) 75:25% (c) 50:50% (d) 25:75% and (e) 0:100% of PLLA/TPU-(y-Fe2 O3) scaffold

Correlation between porosity, hydrophilicity and surface roughness

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4.11

4.12

4.13

4.14

4.15

4.16

4.17

4.18

4.19

4.20

4.21

4.22

4.23

4.24

4.25

4.26

FE-SEM images of degradation rate of different ratios of PLLA/TPU-(y-Fe2O3) scaffolds in weeks 1, 4, 6, 12 and 24 of immersion 113

Change of mass (mg) as a function of degradation time 116

EDX graph of 50:50% PLLA/TPU-(y-Fe2O3) after 4 weeks of immersion in SBF 117

Change in porosity (%) as a function of degradation time 118

The well plate analysis of (a) cytotoxicity assay (b) MTT assay of different ratios of PLLA/TPU-(y-Fe2O3) scaffold during 1, 3, 5 and 7 days of incubation 119

Cytotoxicity result during 72 h of incubation 120

Results of relative cell viability versus the incubation time (days). Each value is the mean±SD of three independent experiments. (*,p<0.05) Indicates significant differences compared to the control analyzed by one-way ANOVA 121

Results of relative cell viability for (50:50%) neat PLLA/TPU and PLLA/TPU-(y-Fe2O3) electrospun mats.(*) Significant difference compared to the control analyzedby unpaired t-test (p<0.05). (**) Indicates Significantdifferent compare to the PLLA/TPU-(y-Fe2O3) 122

FE-SEM images of HSF-1184 fibroblast cells attachmentresults on PLLA/TPU-(y-Fe2O3) electrospun mats 124

Results of (a) thrombin time (b) pro-thrombin clotting time for different ratios of PLLA/TPU-(y-Fe2O3) and neat PLLA/TPU scaffolds. (*) Significant difference compared to the control analyzed by unpaired t-test (p<0.05). (**)Indicates the significant difference between neatPLLA/TPU and PLLA/TPU containing maghemite(p<0.05) 127

Results of the first fibrin detection point by using Clot Sp. instrument in terms of (a) time (s) (b) concentration (g/L) 128

Hemo-lytic test records for different ratios of PLLA/TPU-(y-Fe2O3) electrospun mats. (*) Represents significantdifference compared to the negative control analyzed byunpaired t-test (p<0.05). (**) Indicates the significantdifference between neat PLLA/TPU and PLLA/TPUcontaining maghemite (p<0.05) 129

Mechanical stress vs. strain curve for uniaxial tensile testof different ratios of PLLA/TPU-(y-Fe2O3) electrospunmats 131

Main effect plots for mean S/N ratio of elastic modulus 134

Interaction plots data means for elastic modulus 135

Contribution percent of each parameter on elastic modulus 139

4.27

4.28

4.29

4.30

4.31

4.32

4.33

4.34

4.35

4.36

4.37

4.38

4.39

4.40

4.41

4.42

4.43

4.44

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Plot of (a) normal probability (b) residual vs. predictedvalue 142

Plot of (a) 3D surface response (b) counter plot of elastic modulus 143

Pathway of steepest ascent method 144

Trend of steepest ascent steps responses of elastic modulus 146

Plot of (a) normal probability (b) residual vs. predictedvalue for CCD 150

Plot of (a) 3D surface response (b) counter plot ofquadratic model for elastic modulus 151

Plot of trend between the predicted and actual elasticmodulus 153

Exterior and interior images of 3D scaffold (aluminiumbased design) 154

Images of fabrication procedure of 3D scaffold of aorticheart valve 155

Images of (a) FE-SEM (b) diameter distribution of 156optimum 3D scaffold leaflet

Water contact angle for optimum scaffold 157

Procedure of porosity measurement using liquiddisplacement method 158

FE-SEM and CLSM images of valve’s leaflet and root in 15, 20 and 34 days of incubation. The different colours represented the live cells (Green), dead cells (Red) and scaffold (Dark) area 159

Results of 3D scaffold relative cell viability versus theincubation time (days). Each value is the mean ± SD of allthe experiments 160

Blood clotting time in terms of PT and TT assay 161

Results of the first fibrin detection for optimum scaffold in terms of concentration (g/L) and time (sec). (*) Significant difference compared to the control analyzed by unpaired t- test (p<0.05) 162

Hemo-lytic test records for PLLA/TPU containing 3.80% and 1% of (y-Fe2O3) scaffold. (*) Represents significant difference compared to the negative control analyzed by unpaired t-test (p<0.05) 163

Load vs. extension curve for optimum scaffold after 15, 20and 34 days of cell seeding 165

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4.45 Elastic modulus loss as a function of incubation time (days). (*) Represents significant difference compared to the day zero (for dry scaffold) and day 15 (for wet scaffold) analyzed by unpaired t-test (p<0.05).

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2D - Two-dimensional

3D - Three-dimensional

Adj-R - Adjusted R-Square

ADSCs - Adipose derived stem cells

AFM - Atomic force microscopy

ANOVA - Analysis of variance

AOHV - Aortic heart valve

AV - Atrioventricular valve

BMSCs - Bone marrow stem cells

C - Carbon

CAD - Computer aided design

CCD - Central composite design

CLSM - Confocal laser scanning microscopy

CT - Computed tomography

DCM - Dichloromethane

DMF - Dimethylformamide

DMEM - Dulbecco’s modified Eagle’s medium

DMSO - Dimethyl sulfoxide

DOE - Design of experiments

DSC - Differential scanning calorimetry

ECM - Extracellular matrix

EDX - Energy dispersive X-ray

EGFP - Enhanced green fluorescent protein

LIST OF ABBREVIATIONS

xx

FBS - Fetal bovine serum

FDA - Food and drug administration

Fe - Iron

FE-SEM - Field-emission scanning electron microscopy

FeCl2 - Ferric (Iron (II) chloride)

FeCl3 - Ferrous (Iron (III) chloride)

FTIR - Fourier transform infrared spectroscopy

GAG - Glycosaminoglycan

HASMCs - Human aortic smooth muscles cells

HBSS - Hank's buffered salt solution

HCl - Hydrochloric acid

HNO3 - Nitric acid

HRBCs - Human red blood cells

HSF-1184 - Human skin fibroblast

HA - Hydroxyl apatite

LV - Left ventricles

MRI - Magnetic resonance image

MTT - 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide

NH 3 - Ammonia solution

O - Oxygen

P4HB - Poly-4-hydroxybutyrate

PAN - Polyacrylonitrile

PBS - Phosphate buffer saline

PCL - Polycaprolactone

PF - Polyfumaroles

PGA - Polyglycolic acid

PGS - Polyglycerol sebacate

PLA - Polylactid acid

PLGA - Poly (lactic-co-glycolic) acid

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PLLA - Poly-L-lactic acid

PDLLA - Poly-DL-lactide

Pred-R - Predicted R-Squares

PT - Pro-thrombin time

PTT - Partial thrombin time

PVA - Polyvinyl alcohol

RSM - Response surface methodology

SBF - simulated body fluid

SD. - Standard deviation

SFF - Solid free form fabrication

S/N - Signal to noise

SV - Semilunar valve

TCP - Thricalcium phosphate

TE - Tissue engineering

TEHV - Tissue engineering heart valve

TEM - Transmission electron microscopy

TPU - Thermoplastic polyurethane

Tris - Tris hydroxymethyl amino methane

TT - Thrombin time

UTM - Universal testing machine

VICs - Valve interstitial cells

XRD - X-ray diffraction spectroscopy

Y-Fe2O3 - Maghemite

xxii

LIST OF SYMBOLS

A - Flow rate — First factor

B - Voltage — Second factor

C - Maghemite content — Third factor

D - Solution concentration — Fourth factor

E - Collector rotating speed — Fifth factor

El - Stiffness index

F - Force

H - Valve height

HP - Hemolysis percent

K - Number of factors

L - Gap length between grids

Ms - Mean of squares

S - Square area

SS - Sum of squares

T - Thickness

v/v - Volume/Volume

w/v - Weight/Volume

% - Percent

°C - Degree Celsius

Y - Gamma

Xc - Crystallinity

AHm - Melting enthalpy

W f - Weight fraction of reference polymer

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AH°m - Melting enthalpy of the reference polymer

a(E) - Tensile stress

8 - Tensile strain

P - Regression coefficient

S - Depth of indentation

5max - Maximum depth

p - Density

AL - Change in length

AX - Step size

Y - Regression response

R2 - R-square

A 0 - Cross section area

Cd - Commissure diameter

Ch - Commissure height

d f - Degree of freedom

Db - Base diameter

Dc - Absorbance value of control

DNegC - Absorbance value of negative control

DposC - Absorbance value of positive control

Ds - Absorbance value of sample

Dt - Top diameter

Ee - Elastic modulus

Er - Relative elastic modulus

L0 - Origin length

Lf - Leaflet free-edge

Lh - Leaflet height

M1 - Average of counted cells via hemocytometer

M2 - Number of cells to be seeded

Pr - Porosity

xxiv

Ra - Surface roughness average

Re - Contact surface area between ball probe and sample

Rpv - Peak to valley

Rq - Root-mean square of surface roughness

Rz - Five lowest valley and five highest peak averages

Tg - Glass transition temperature

Vc - Required volume of cell suspension

Vt - Total required solution

Vi - Known volume

V2 - Total of new volume after immersion of sample

Vs - Total of new volume after sample removal

Vf - Poisson’s ratio

Wd - Dry weight

W0 - Original weight of sample

Wt - Mass change

Ww - Wet weight

2 - Predicted response

xxv

APPENDIX

A

LIST OF APPENDICES

TITLE PAGE

List of publications 202

CHAPTER 1

INTRODUCTION

1.1 Overview of the Research

The four valves represented in the mammalian hearts are responsible to

maintain unidirectional, non-hinder blood flow from heart to the other parts of body.

The heart valves synchronically open and close approximately 40 million times a

year and more than 3 billion times during the life (75 years average life expectancy)

(Aagaard, 2004; Rabkin-Aikawa et al., 2005). The four valves are namely; (I):

Aortic, (II): Pulmonary, (III): Mitral (bicuspid) and (IV): Tricuspid. The aortic and

pulmonary valves are in the arteries leaving the heart and known as semilunar valves

(SV), the mitral and tricuspid are between the atria and ventricles which known as

atrioventricular valves (AV) (Gallyamov et al., 2014; Saito et al., 2016). Valvular

heart dysfunction is a significant cause of morbidity and mortality around the world.

The prevalence of heart disease in adult US population in the early of 21st century

has been estimated at more than 5 million (Schoen, 1997; Basso et al., 2013). In the

meantime; heart valves (especially aortic and mitral) dysfunction is a significant part

of heart disease, which leads to death of approximately 20,000 people around the

world annually. Approximately more than 290,000 heart valve surgeries is required

around the world each year and according to the increase in average age of the

population, it is estimated to reach 850,000 by the year 2050 (Yacoub and

Takkenberg, 2005; Sewell-Loftin et al., 2011). A heart valve consists of two or three

semicircular shape moving flaps which are called leaflets and comes together in the

center of the valves to close it. These leaflets are attached to the walls of a cylindrical

conduit which is called the valve root. The histology of the valve leaflets (cusps)

2

exhibits an extracellular matrix (ECM) structure of three distinct layers: fibrosa,

spongiosa and ventricularis. The fibrosa, the surface away from the blood flow,

compose of parallel, dense collagen which is associated with the mechanical

properties such as stiffness and strength of the leaflets. The spongiosa, the middle

surface, compose of proteoglycans and lower abundance of collagen which facilitate

the movement, and finally, the ventricularis, the adjacent surface, compose of aligned

fiber of elastin interspersed and short collagen fiber, which provides the elasticity

properties of leaflets. The aortic valves leaflets comprise of 45% collagen (types, I

(74%), III (24%) and V (2%)), 20% elastin in dry weight and 35% of glycosamino

acid (Gross and Kugel, 1931; Garcia-Martinez et al., 1991; Cox, 2009; Falk et al.,

2011; Gallyamov et al., 2014; Roberts et al., 2015) Figure 1.1 illustrates the

anatomy/position of the four valves and the structure of ECM of leaflets.

Pulmonary Valve

(b)

(c)

Figure 1.1: Schematic of (a) heart anatomy (b) valve position (c) single valve leaflet

and its ECM structure (Schoen, 1997)

3

Heart valve dysfunctions may arise due to development regulation,

mechanical properties shortage (such as leakage and overlaps of the leaflets),

absence or abnormality of tissue in congenital cases and even calcification causes by

the deposition of mineralized calcium and genetic defect in the matrix protein

structure (Nasuti et al., 2004; Ng et al., 2004). When one of the valves malfunctions,

the end stage of medical choice may be to replace the original valves with an

artificial one. Generally, the cardiac surgeries to replace the heart valves are common

around the world and improve the life expectancy. Currently, the mechanical valves

and the biological (glutaraldehyde xeno-grafts or cryopreserved homo-grafts) are

used clinically as the state-of-the-art of artificial valves, despite the occurrence of

prosthesis side effects such as the need to anticoagulation treatment and durability of

valves (Iung et al., 2003; Nkomo et al., 2006; Thom et al., 2006). In order to resolve

the drawbacks of prostheses, the tissue engineering concept is introduced (Tanaka et

al., 2005; Jain et al., 2010; Lueders et al., 2014; Cheung et al., 2015). Figure 1.2

depicts the artificial heart valves.

(a) (b) (c)

Figure 1.2: Artificial heart valves (a) mechanical (b) biological (c) tissue

engineering heart valve (Fallahiarezoudar et al., 2015c)

Tissue engineering (TE) is an integrated science between the engineering

principle and life science to overcome the limitation of prostheses. The principle of

TE is to provide a three-dimensional (3D) scaffold (that resembles the original tissue

properties) for a specific tissue to develop the neotissue from their cellular

components. The scaffold provides an initial environment for cell attachment and

tissue growth. The cell can be either seeded onto the scaffold matrix in vitro (pre­

implementation) or in vivo (post-implementation) to develop a neotissue for

replacement or repair the damaged tissue. The fabrication of a neotissue from cellular

4

combination (which depicts most of the characteristics of the original tissue such as

non-inflammation and non-immunogenic reaction, adequate mechanical properties

and durability) is the ultimate goal of TE. The concept of tissue engineering heart

valve (TEHV) was introduced by Shinoka et al. (1995). The TEHV principle can be

summarized in three main phases: (I) 3D biocompatible scaffold fabrication, (II) cell

cultivation and seeding over the scaffold, and (III) development conditions of the

TEHV before implantation (Sheridan et al., 2000; Teebken et al., 2000). Figure 1.3

illustrates the concept of TEHV.

Figure 1.3: Concept of tissue engineering heart valve (TEHV) (Fallahiarezoudar et

al., 2015c)

In order to design and fabricate a proper 3D heart valve scaffold, the initial

and probably one of the most important phases of TEHV concept is to investigate the

valve structure and function. Each layer (fibrosa, spongiosa and ventricularis) which

has a particular property, forms the valves structure. In the scaffold fabrication phase,

the appropriate materials selection and fabrication techniques are debatable. The

scaffold architecture (matrix) is very important as it is the basis of TEHV concept. In

order to ensure a successful scaffold: (I) The utilized materials should be biologically

5

compatible, biodegradable and cover the requirements for mechanical properties, (II)

the structure should provide a hierarchical extensive network of interconnecting

pores to facilitate the cells attachment and provides the oxygen and nutrients to those

cells which are far away from the surface (usually more than 1 mm) and (III) the

shape and the size of the scaffold should resemble the native tissue (Aikawa et al.,

2006; Baaijens et al., 2010).

The usable materials for scaffolds fabrication purpose can generally be

classified into two groups: the polymer based and the natural based materials. The

basic of scaffold mechanical properties and hemocompatibility with in vivo are

highly dependent on the materials selection. The required properties for TEHV

scaffold can be translated into biocompatibility, biodegradability, thermo-plasticity,

elasticity and stiffness characteristics (Argento et al., 2012; Chen et al., 2013; Alves

et al., 2014). The main advantages of a synthetic polymer-based scaffold are the fact

that biomechanics and degradation properties can be chemically controlled according

to the requirements. Against the cytotoxicity, low degradation rate and inflammation

are the main drawbacks of the synthetic polymeric scaffold. Although no

biodegradable polymeric materials have been proven to be a desirable substitute for

the native valves, work continues to be promising (Zhai et al., 2010; Eckert et al.,

2013; Masoumi et al., 2013a). Furthermore, the utilization of the natural materials

such as collagen has also been reported as the raw material for scaffold fabrication.

Collagen is the main extracellular matrix protein of the origin heart valves. The

fibrosa is considered to be the main load carrying structure and is primarily

composed of circumferentially arranged densely packed bundles of collagen fibres

and a micro-fibrillar network of elastin. The use of collagen regarding to low

antigenicity and less immunogenicity can be the advantages of collagen based

scaffold. However, poor handling, low mechanical properties, and less controlled

biodegradability are the defects of collagen based scaffold (Chevallay and Herbage,

2000; Levy et al., 2004; Balguid et al., 2007; Chen et al., 2013).

Poly-L-lactic acid (PLLA) is one of the preferred biomaterials that are widely

used in different fields of TE. PLLA due to high tensile strength, good compatibility

in vivo and biodegradability is considered to be used in this research. However, the

6

rigidity and non-flexible matrix of PLLA limited the development of soft tissue by

this polymer (Kang et al., 2009; Sakai et al., 2013; Qiao et al., 2016). The PLLA

scaffold was fabricated through electrospinning technique for bone tissue

engineering purpose that indicated elastic modulus value of 35 MPa in such a way

that the tensile strain was around 5-10% (Luu et al., 2003). Application of pure

PLLA scaffold (fabricated by different techniques) was mostly reported as a template

for bone tissue engineering purpose. The biocompatibility and tissue formation using

nanofiber based PLLA scaffold was confirmed in vivo. PLLA scaffold was well

colonized with cells after implantation, but only showed marginal ossification

(Schofer et al., 2011). Besides, bio-grade thermoplastic polyurethane (TPU) exhibits

superb elasticity (more than 220% of strain) with non-inflammation behaviour in

vivo (Chen et al., 2010; Jing et al., 2016). The fabricated electrospun scaffold using

pure TPU exhibited the samples were deformed easily in a low stress value which

may not be appropriate in this case (Chen et al., 2009a; Jing et al., 2016). Fabricated

TPU electrospun nanofibers indicated super hydrophobicity that made difficulties for

cell culturing (Wang et al., 2011). Therefore, the mixture of these two polymers can

lead to a composite with desired tensile strength, elasticity and biocompatibility for

soft TE purpose (Mi et al., 2013; Jing et al., 2014; Marycz et al., 2016).

Maghemite (y-Fe2O3) nanoparticle, which is a novel biocompatible material,

has recently been used in biomedical applications such as magnetic cell seeding, cell

expansion and drug delivery and the results are quite promising. The outstanding bio­

behaviours (mechanical and biological) of maghemite have been reported in previous

researches (Tartaj et al., 2003; Wei et al., 2011; Ngadiman et al., 2015). Maghemite

nanoparticles filled nanofibers such as polyvinyl alcohol (PVA) were used

previously for composite reinforcement purpose (Fallahiarezoudar et al., 2015b).

Furthermore, maghemite filled polyvinyl alcohol was reported as a potential

materials for bone tissue engineering purpose which resulted in higher tensile

strength and better cell proliferation (Ngadiman et al., 2017).

Scaffold fabrication can be classified into two main techniques. Conventional

techniques such as solvent casting in combination with particulate leaching and

phase separation in combination with freeze drying, and fashionable techniques such

7

as electrospinning and solid-free-form (3D printing) fabrication. The ECM structure

and subsequently the mechanical properties and biological compatibility are highly

dependent upon the parameters which can be modified in fashionable fabrication

methods. The microstructure parameters such as the fiber diameter, interconnectivity,

porosity properties and stiffness, which ultimately shape the macrostructure

properties of the scaffold, should be investigated. In perspective of scaffold

manufacturing, each technique has its own pros and cons (Hutmacher and Cool,

2007; Hoque et al., 2012).

Electrospinning is a versatile and straightforward technique for

cardiovascular scaffold fabrication. In the electrospinning setup, a high electric field

is responsible for transforming the emerging solution supplied via syringe pump into the

fibers. Figure 1.4 depicts a schematic diagram of the electrospinning setup. High

surface area to volume ratio and high porosity are the advantages of electrospinning

process. In fashionable fabrication techniques such as electrospinning the

microstructure can be modified according to the requirements by varying the

parameters involved such as types of polymers and solvents, flow rate, voltage,

needle-collector distance and others. The fibrous scaffold matrix prepares an

auspicious layout for the cell attachment, migrant and growth. However, the lack of

mechanical properties, time consuming and problems with residual solvent which

may stimulate the possibility of toxicity are the drawbacks (Ahmadipourroudposht et

al., 2015; Fallahiarezoudar et al., 2015a).

Figure 1.4: Schematic diagram of the electrospinning setup

8

The second phase in tissue engineering concept is the cell cultivation and

seeding over the scaffold. The successful cell seeding process is highly related to the

scaffold architecture and biocompatibility of materials that were used for the 3D

scaffold fabrication. The inter-related, hierarchical scaffold structure will stimulate

the cell proliferation/migration and subsequently shape the ECM. Three types of

cells are reported to be useful in TEHV: (i) The vascular smooth muscles cells

(Colazzo et al., 2011), (ii) the valves interstitial cells (Masoumi et al., 2013b) and

(iii) the bone marrow stem cells (Hadju et al., 2011). Two methods of cell seeding

are proposed as dynamic and static. In dynamic method the scaffold is in constant

motion during cell seeding in the incubator an opposed to the static environment.

Various types of cell adhesion molecules such as integrin have been used to coat the

scaffold matrix prior to cell seeding. Interaction between the integrins on the cell

membrane and the receptors on ECM are largely responsible for cell attachment. The

integrin consist of a and fi chains (18 subunits for a and 8 subunits for fi) which have

the responsibility for cell attachment to ECM and signal translation from the ECM to

the cells (Lam et al., 2002; Taylor et al., 2006).

Once the fabricated scaffold was seeded, the neo-tissue starts to develop. The

environment in which the scaffold grows is one of the criteria for successful TEHV.

The scaffold development environment will influence the formation of the ECM.

Two perspectives have been proposed by the researchers: (I) the black box approach

where the scaffold is implanted in vivo shortly after cell seeding and used as cell

delivery to the native tissue (Vacanti et al., 1988; Fong et al., 2006), (II) bioreactor

approach such as pulse duplicator which provides the physiological pressure and

flow to the developed TEHV and promotes both cell function and mechanical

properties (Sodian et al., 2001; Engelmayr et al., 2003; Mol et al., 2005).

1.2 Research Problem Statement

Currently the usable clinical prostheses of heart valves (mechanical or

biological) have a non-viable structure and have no capacity to grow, remodel or

9

repair. As a result, almost 50-60% of patients experienced problems with artificial

valves and require reoperation (Chevallay and Herbage, 2000; Aagaard, 2004; De

Heer et al., 2013; Grzymala-Lubanski et al., 2016). Mechanical valves as a foreign

material may cause inflammation, infection and thromboembolic complication due to

high shear stresses of blood flow which requires an anti-coagulation medication such

as a vitamin K antagonist (e.g. warfarin) along the life. Although warfarin could be

efficacious to alleviate coagulation, it has the risk of hemorrhagic and also the

embryo toxicity in fertile women (Lip et al., 2015; Roberts et al., 2016). On the other

hand, despite the lower thrombotic risk of biological valves compared to the

mechanical one (0.87% to 1.4% per annum, respectively) and no need to undergo

anti-coagulation treatment for the biological valves, the durability of these prostheses

are approximately 10-15 years due to the progressive tissue deterioration and this is

almost half the lifetime of the mechanical valves (20-30 years) (Potter et al., 2004;

Pibarot and Dumesnil, 2009; Tillquist and Maddox, 2011). These limitations forced

scientists to further investigate of other possible methods of creating a neo-tissue

similar to the original tissue that can fully solve the drawbacks of the artificial

valves. However, TEHV principle is introduced as a possible method of resolving

these limitations (Tanaka et al., 2005; Jain et al., 2010; Lueders et al., 2014; Cheung

et al., 2015). A biodegradable and biocompatible 3D scaffold with adequate

characteristics is fabricated, seeded with the appropriate cells source and developed

in vitro in a bioreactor to create a biomimetic tissue which resembles the original

tissue characteristics (Argento et al., 2012; Cui et al., 2016).

Previous researches on TEHV were performed using poly lactic acid (PLA)

(Armentano et al., 2013; Sakai et al., 2013), polyglycerol sebacate (PGS) (Masoumi

et al., 2013a; Sanz-Garcia et al., 2015), polyglycolic acid (PGA) (Matsumura et al.,

2003; Frese et al., 2016), polycaprolactone (PCL) (Vaz et al., 2005; Marei et al.,

2016) to fabricate the scaffold. Most of these researches result in a non-viable

structure or toxicity due to the use of inorganic solvent. In addition, the degradation

rate and mechanical properties of polymers also plays a critical role in TE concept.

Fabrication of scaffold using PLA indicated a desirable biocompatibility and

biodegradability (around 6 months), but much thicker and less flexible (roughly 1­

2% tensile strain) which was in conflict with dynamic mechanism of original tissue

10

(Armentano et al., 2013; Sakai et al., 2013). Despite good mechanical properties of

PCL (3.7 MPa tensile stress and 90% tensile strain), low degradation rate (more than

2 years) causes the failure of TE concept (Vaz et al., 2005; Yao et al., 2016). Also

the fabricated PGS-based heart valve scaffold exhibited very fast degradation rate (2­

3 weeks) with 0.5 MPa of elastic modulus but it could not provide the required

elastic modulus for heart valves (Masoumi et al., 2013a; Lin et al., 2015). On the

other hand, the scaffold adverse structure (that can be the result of inappropriate

fabrication method) may result in the aggregation of seeded cells just over the

scaffolds surface. As reported by Taylor et al. (2002) an extremely low expansion of

valvular interstitial cells was observed on micromould injection PLGA scaffold.

Colazzo et al. (2011) reported that the fabricated PLA scaffold using freeze drying

technique resulted in cell adhesion only on the surface that can be attributed to

disconnected pores. Furthermore, scaffold overgrowth, leakage and rupture before

implementation (Van Lieshout et al., 2006; Muylaert et al., 2014) or inflammation

due to low degradation rate after implementation (Choi and Park, 2002; Loger et al.,

2016) can happen due to the improper selection of polymeric materials source. Thus,

the selection of a proper source of polymeric based biomaterials in terms of

biocompatibility, biodegradability and mechanical properties as well as an applicable

fabrication technique that can satisfy the heart valve characteristics is the major

concern.

1.3 Research Questions

(i) What is the optimum ratio of polymer solutions blend according to the

required characteristics of an aortic heart valve?

(ii) What is the biological and mechanical effect of maghemite

nanoparticles presence in nanofibers matrix?

(iii) What are the biomechanical and structural characteristics of native

aortic heart valve leaflets?

(iv) What are the required specifications to design and fabricate a three­

dimensional scaffold that can resemble the origin tissue characteristics?

11

(v) How to perform the cell seeding procedure and evaluate the mechanical

performance after cell seeding?

1.4 Research Hypothesis

Design of complex structures such as heart valves need to be simplified prior

to fabrication and modelling. It this case some assumptions are applied to the design

procedure to make it tractable.

First, it is assumed that the valves cusps have identical characteristics in

dimensions and mechanical properties. Since the access to the human heart valve

information and dimension is difficult, in this research the specifications are

extracted from other previous works (Swanson and Clark, 1974; Labrosse et al.,

2006). Second, the leaflets lie at 120° from each other in the 3D circle plate. Third, it

is assumed that the top and bottom of the cylinder with the valve inside of it (aortic

root) are parallel. The final and noticeable hypothesis in designing the valve is to

consider that the valves component’s dimensions are not changed significantly

during the cardiac cycle.

1.5 Research Aim and Objectives

The aim of this research is to fabricate, characterize and optimize a nanofiber-

based scaffold of aortic heart valve with the extracellular matrix (ECM) structure and

appropriate biological and mechanical characteristics to assist the aortic smooth

muscles cell adhesion, migration and proliferation.

Objectives of this research are:

12

(i) To characterize different ratios of poly-L-lactic acid/thermoplastic

polyurethane polymers to tune the ratio that proposes the best

performance for tissue engineering heart valve.

(ii) To investigate the effect of maghemite nanoparticles on chemical,

biological and mechanical properties of electrospun mat.

(iii) To optimize the fabrication process (electrospinning) in terms of elastic

modulus that tailors the scaffold structure and mechanical properties.

(iv) To design and fabricate a three-dimensional nanofiber-based aortic

heart valve scaffold.

(v) To evaluate the migration and proliferation of aortic smooth muscles

cell over the scaffold and dwindle of mechanical properties as a

function of incubation time.

1.6 Research Scopes

The left heart typically achieves a peak pressure about six times of the right

over the cardiac cycle. So, the two valves on the left side of the heart are subjected to

much higher loads than those on the right heart (Hasan et al., 2014). Since the

majority of valve diseases involve the valves of the left heart, therefore, the scope in

this research is limited to investigating the anatomy and design microstructure of the

tissue engineering aortic heart valves. The heart valve tissue engineering concept can

be split into three main steps: (i) Fabrication of 3D scaffold with proper design and

properties, (ii) cell seeding over the 3D scaffold, and (iii) development of neo-tissue

in bioreactor prior to implementation. In this research the scope is limited to the

characterization of the novel mixture of poly-l-lactic acid (PLLA), thermoplastic

polyurethane (TPU) and maghemite (y-Fe2O3) nanoparticles as the potential

biocompatible and biodegradable materials for heart valve scaffold fabrication. A 3D

template associated to the semilunar heart valves was designed to fabricate a 3D

scaffold and the biological and mechanical properties of the fabricated scaffold are

characterized as well.

13

In order to fabricate a biomimetic scaffold of aortic heart valve with adequate

properties that can resemble the original tissue, two parameters of usable materials

and fabrication techniques are among the available options considered.

(i) The utilized materials are PLLA, TPU and (y-Fe2O3) nanoparticles

which are categorized in the synthetic, polymeric materials group and

their optimum proportions were determined.

(ii) The fabrication method was limited to the electrospinning technique.

(iii) The controllable parameters involved (which may have significant

effect on the microstructure) such as flow rate, voltage, maghemite %,

solution concentration and rotating speed, as well as noise parameters

were selected as the variables to optimize the microstructure

properties.

(iv) The characterization of electrospun mats in terms of morphology,

porous properties, surface roughness, hydrophilicity, cytotoxicity

assay, degradation rate, blood hemocompatibility, mechanical

behaviour, cells attachment, migration and proliferation over the

samples were investigated during this research.

(v) The characteristics mentioned with emphasis on the mechanical

properties (tensile stress, tensile strain and elastic modulus) of the

aortic heart valve leaflets were the major scope.

(vi) The optimization of scaffold fabrication technique was performed

(based on uniaxial tensile properties) in terms of elastic modulus

using Taguchi experimental design and response surface

methodology.

(vii) The MTT (3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium

bromide) assay using human aortic smooth muscles cells (HASMCs)

were used to verify the cell viability.

(viii) Blood hemocompatibility in terms of clotting time, fibrin formation

and hemolysis were investigated.

(ix) The biomechanics behaviours of seeded scaffold as a function of

incubation time were characterized using macro-indentation

bending/flexural tests.

14

1.7 Significance of Research

A novel combination of material PLLA/TPU containing maghemite is useful

in the design and fabrication of a synthetic biodegradable scaffold. The unique

material developed has both the extracellular matrix structure of the aortic heart

valve with interconnected networks (to provide the oxygen and nutrients to the cells

which are far away from the surface) and the required mechanical properties such as

tensile strength and flexural to resist against the blood pressure. In addition, the

material has the stiffness and elastic characteristics that can be useful on biomaterials

for scaffold development. Information on the electrospinning process parameters that

produces nanofibers with optimum fiber diameter distribution and porosity with

excellent mechanical properties and structure were also disclosed for the unique

material. The findings of the research can also increase the life expectancy of

patients experiencing valvular heart dysfunction. The developed synthetic bio­

polymeric TEHV has the possibility of reducing the number of times a patient needs

to undergo valve replacement surgery and this can reduce the complications after

surgery and also cost.

1.8 Organization of Thesis

In the first Chapter of this thesis, general information of the research,

objectives and scope are provided. In Chapter 2, the literature review on tissue

engineering and particularly tissue engineering heart valves is described including

the previous investigation on the microstructure, function and mechanical properties

of the origin porcine and/or human tissue. Chapter 3 presents the research framework

and detailed explanation of each phase to show the methods of experiments

conducted and initial finding is provided. Chapter 4 provides the results and a

detailed discussion on the findings of this research. This chapter is divided into three

main sections; materials ratio tuning, electrospinning process optimization using

Taguchi experimental design and 3D scaffold fabrication/characterization. In Chapter

5, the conclusion was made according to the assumptions made and the results

obtained.

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