Electrospun natural proteins for cardiac tissue engineering

ELECTROSPUN NATURAL PROTEINS FOR CARDIAC TISSUE ENGINEERING PREETHI BALASUBRAMANIAN (B.E, ANNA UNIVERSITY) A THESIS SUBMITTED FOR THE DEGREE OF MASTER OF ENGINEERING DEPARTMENT OF MECHANICAL ENGINEERING NATIONAL UNIVERSITY OF SINGAPORE 2012 ACKNOWLEDGEMENT I would like to express my sincere appreciation to those who have helped and contributed to this thesis. I would like to express my sincere thanks to Professor Seeram Ramakrishna who has shown faith in me and given me tremendous encouragement and excellent supervision throughout my tenure. I would like to express my heartfelt gratitude to Dr. Molamma Prabhakaran, who has provided unmatched guidance and support, throughout this project. I would also like to thank all Prof Seeram’s lab members for their assistance in the completion of this project. I would like to thank the Department of Mechanical Engineering and Faculty of Science for their constant support. Last but not the least I would like to thank my family and friends for their profound love and support. I TABLE OF CONTENTS ACKNOWLEDGEMENTS I TABLE OF CONTENTS II LIST OF FIGURES VI LIST OF TABLES IX LIST OF ABBREVIATIONS X SUMMARY XII Chapter 1: Introduction 1.1 Background 1 1.2 Myocardial Infarction 2 1.3 Hypothesis and objectives 5 Chapter 2: Literature Review 2.1 Introduction 6 2.2 Tissue Engineering 7 II 2.3 Biomaterials 8 2.4 Natural Proteins 11 2.4.1 Collagen 11 2.4.2 Gelatin 12 2.4.3 Fibrinogen 14 2.5 Nanofibers by electrospinning 15 2.5.1 Electrospinning – Basic principle 15 2.5.2 Advantages of electrospun nanofibers for tissue engineering 17 Chapter 3: Human cardiomyocyte interaction with electrospun fibrinogen/gelatin nanofibers for myocardial regeneration 3.1 Introduction 18 3.2 Materials and Methods 19 3.2.1 Materials 19 3.2.2 Fabrication of scaffolds by electrospinning 19 3.2.3 Glutaraldehyde cross-linking 20 3.2.4 Morphological, chemical and mechanical characterization of the electrospun natural polymeric nanofibrous scaffolds 20 III 3.2.5 Cell culture on the electrospun scaffolds 3.3 Results 21 24 3.3.1 Morphological, chemical and mechanical characterization of the electrospun natural polymeric scaffolds 3.3.2 Cell culture studies 3.4 Discussion 24 30 36 3.4.1 Importance of the blood plasma protein – fibrinogen in myocardial regeneration 3.5 Conclusion 36 41 Chapter 4: Electrospun collagen/fibrinogen nanofibrous scaffolds enhance cardiomyogenic differentiation of adipose-derived stem cells for myocardial tissue engineering 4.1 Introduction 42 4.2 Material and methods 43 4.2.1 Materials 43 4.2.2 Fabrication of scaffolds by electrospinning 43 4.2.3 Glutaraldehyde cross-linking 44 4.2.4 Cell culture on the electrospun scaffolds 44 4.3 Results and Discussion 47 IV 4.3.1 Importance of collagen in myocardial regeneration 47 4.3.2 Myocardial regeneration potential of adipose-derived stem cells 53 4.4 Conclusion 62 Chapter 5: Conclusions and Recommendations 5.1 Conclusion 63 5.2 Recommendations 64 REFERENCES 66 V LIST OF FIGURES Figure 1.1: Schematic diagram illustrating the damage caused by MI 4 in human hearts Figure 2.1: Schematic electrospinning setup. (A) Polymer melt taken 16 in syringe (B) Nozzle (C) High voltage transformer (D) Electrospun jet from nozzle (E) Collector (rotatory or stationary) Figure 3.1: Morphology of electrospun (A) Fib/Gel(1:4)-CL (B) 24 Fib/Gel(2:3)-CL (C) Fib-CL nanofibers Figure 3.2: FTIR spectra of the electrospun nanofibers 26 Figure 3.3: Stress-strain curves of electrospun and dry scaffolds of 28 Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL nanofibers Figure 3.4: Stress-strain curves of electrospun PBS-soaked Fib/Gel(1:4)-CL, 29 Fib/Gel(2:3)-CL and Fib-CL nanofibers Figure 3.5: Human cardiomyocyte proliferation on electrospun nanofibers 30 measured by MTS assay. *Significant against proliferation on Fib-CL nanofibers at p≤0.05. VI Figure 3.6: Interaction of Human cardiomyocytes on various substrates 31 (A) Fib/Gel(1:4)-CL (B) Fib/Gel(2:3)-CL (C) Fib-CL (D) TCP Figure 3.7: Cardiac-specific-protein expressions of actinin on 33 Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, Fib-CL and TCP: (A, D,G, J); cell nuclei stained blue; (B, E, H, K); merged images of cell nuclei and actinin. (C, F, I, L) Figure 3.8: Cardiac-specific-protein expressions of troponin I on 34 Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, Fib-CL and TCP: (A, D,G, J); cell nuclei stained blue; (B, E, H, K); cell troponin stained green and merged images of cell nuclei and troponin (C, F, I, L) Figure 3.9: Cardiac-specific-protein expressions of connexin on 35 (A) Fib/Gel(1:4)-CL and (B) Fib/Gel(2:3)-CL; expression of MHC on (C) Fib/Gel(1:4)-CL and (D) Fib/Gel(2:3)-CL Figure 4.1: SEM morphology of electrospun (A) Fib/Coll(1:4)-CL 48 and (B) Fib-CL nanofibers Figure 4.2: FTIR spectra of the electrospun Fib/Coll(1:4)-CL and 49 Fib-CL nanofibers Figure 4.3: Stress-strain curves of electrospun and dry scaffolds of 51 Fib/Coll(1:4)-CL, and Fib-CL nanofibers Figure 4.4: Stress-strain curves of electrospun PBS-soaked 52 Fib/Coll(1:4)-CL, and Fib-CL nanofibers VII Figure 4.5: Adipose-derived cells are capable of direct differentiation 54 and paracrine actions upon damaged tissue Figure 4.6: Cell proliferation study of human ADSCs- co-cultured with 56 human cardiomyocytes. *Significant against proliferation on Fib-CL nanofibers at p≤0.05. Figure 4.7 SEM images showing the morphology of ADSC-cardiomycytes 57 co-culture cells on (A) Fib/Coll(1:4)-CL (B) Fib-CL and (C) TCP Figure 4.8: Cardiac-specific-protein expressions of MHC on Fib/Coll(1:4)-CL, 59 Fib-CL and TCP: (A, D,G, J); cell nuclei stained blue; (B, E, H, K); merged images of cell nuclei and MHC (C, F, I, L) comprising of human cardiomyocytes Figure 4.9: ADSC-specific protein expression of CD105 & MHC on Fib/Coll(1:4) 60 -CL (A, D), Fib-CL (B, E and TCP (C, F) comprising of human ADSCs. Figure 4.10: Dual immunofluorescent analysis for the expression of cardiac 61 -specific marker protein MHC (A, D, G) and ADSC-specific marker protein CD 105 (B, E, H) and the merged image showing dual expression of both MHC and CD 105 (C, F, I) on Fib/Coll(1:4)-CL (A – C), Fib-CL (D – F) and TCP (G – I) on the co-culture system VIII LIST OF TABLES Table 2.1 Overview of Biomaterials for cardiac tissue engineering 10 Table 3.1 Fiber Diameters and water-contact angles of the electro- 26 spun Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, and Fib-CL nanofibers Table 3.2 Tensile strength and stiffness properties of the electro- 27 spun Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, and Fib-CL scaffolds compared to the native human myocardium Table 4.1 Fiber Diameters and water-contact angles of the electro- 50 spun Fib/Coll(1:4)-CL, and Fib-CL nanofibers Table 4.2 Tensile strength and stiffness properties of the 50 Fib/Coll(1:4)-CL and Fib-CL scaffolds IX LIST OF ABBREVIATIONS ECM Extracellular matrix MI Myocardial Infarction LV Left Ventricle TE Tissue Engineering CTE Cardiac Tissue Engineering VEGF Vascular Endothelial Growth Factor RGD arginine-glycine-aspartic HFP 1,1,1,3,3,3 – hexafluoro-2-propanol DMEM Dulbecco modified eagle’s medium FBS Fetal Bovine Serum PBS Phosphate Buffered Saline EDTA Ethylenediaminetetraacetic acid HMDS Hexamethyldisilazane RT Room Temperature X ATR-FTIR Attenuated total reflectance fourier transform infrared spectroscopy TCP Tissue Culture Plate SEM Scanning Electron Microscopy Act Alpha-actinin, Trop I Troponin I Con-43 Connexin-43 MHC Myosin Heavy chain BSA Bovine Serum Albumin DAPI 4’, 6-diamidino-2-phenylindole dihydrochloride ADSC Adipose-Derived Stem Cells SVF Stromal-Vascular cell Fraction w weight v volume XI Summary Heart attack is a common and deadly condition which is the major cause of death in most developed countries. Myocardial infarction, commonly known as heart attack, occurs when there is a disruption in blood supply to parts of the heart due to occlusion of a coronary artery leading to excessive cell death. Over the past few decades, heart transplantation has been the backbone of therapy for the treatment of heart failure and it has proved to be one of the most effective therapies for end-stage heart failure. But the major obstacles with heart transplantation are the limited supply of suitable donors and lifelong immune suppression which often causes serious consequences. Several other therapies are experimented for the treatment of the infarcted myocardium, of which the tissue engineering approach is gaining much attention. Cardiac tissue engineering promises to bring about a change in the treatment of patients with myocardial infarction and aims at providing cutting-edge solutions to end-stage heart failure. To date, there are no successful models of bioengineered cardiac implant that can suitably mimic the anatomy, physiology and biological stability of a healthy heart wall. To address this challenge, here we report the development and analysis of electrospun nanofibrous composite scaffolds using natural proteins that mimic the native myocardial environment. It is necessary to understand the pathological processes following infarction of which massive cell loss is very critical as the myocardial tissue lacks significant intrinsic XII regenerative capability to replace lost cells. Therefore, it is essential to provide a physiologically relevant milieu for the cells to proliferate. The novelty in our idea is the use of natural body proteins – fibrinogen, collagen and gelatin for the fabrication of a fully natural composite scaffold, supplemented with human cardiac cells which can act as a temporary extracellular matrix until repair. Fibrinogen is a globular blood plasma protein and plays significant roles in hemostasis, wound healing, inflammation, angiogenesis, neoplasia, etc. In the case of the myocardium, fibrinogen along with fibronectin is crucial for the formation of “primary matrix” (in developing the granulation tissue) which functions as a meshwork for the deposition and adhesion of other matrix proteins such as interstitial collagens post myocardial infarction. The use of this blood plasma protein along with another natural body protein such as collagen or gelatin will be ideal to act as a temporary extracellular matrix to support the regeneration of the infarcted myocardium. We have proven that the fabrication of a fully natural nanofibrous composite scaffold using electrospinning process is biocompatible, closely imitates the myocardial extracellular-matrix and offers the possibility to enhance cell proliferation. The optimal amount of fibrinogen in the natural composite for cardiac tissue engineering application is identified. To improve the mechanical integrity of the natural polymeric scaffolds they were cross-linked and the tensile properties of the cross-linked electrospun natural polymeric scaffolds were studied and found to be close to the mechanical properties of the heart. Our electrospun natural polymeric scaffolds were found to provide better biocompatibility, hydrophilicity, biodegradability as well as suitable mechanical properties as desired for cardiac tissue engineering. The cell-scaffold interactions of human cardiomyocytes with the electrospun natural polymeric scaffolds were analyzed in detail by cell proliferation, confocal microscopic XIII analysis for the expression of four cardiac-specific marker proteins like Actinin, Troponin-T, Connexin-43, Myosin heavy chain. We established the cardiomyogenic differential potential of human adipose-derived stem cells which is found to be largely enhanced by the fabricated electrospun natural polymeric composite substrates. Characterization and dual immunofluorescent analysis showed that our fabricated nanofibrous scaffolds greatly promoted differentiation of adipose-derived stem cells, integration with the co-cultured cardiomyocytes, cell attachment and growth because of their biological components. To our knowledge, the idea of using electrospun fully natural-protein composite nanofibrous scaffolds supplemented with human adipose-derived stem cells/cardiomyocytes for the purpose of cardiac tissue engineering is a novel idea and they have the immense potential to be used for in vivo animal studies. XIV Chapter 1 Introduction 1.1 Background The heart is the muscular organ of the circulatory system that constantly pumps blood throughout the body. The heart is made up of the cardiac muscle tissue which is very strong and has the ability to contract and relax rhythmically throughout the life span of a human. The fundamental concept of circulation was developed by Harvey who suggested that the circulation of the blood is caused by pumping of the heart. The human heart has four chambers. The upper two chambers are the left and right atrium which receives blood coming to the heart and delivers it to the lower two chambers – left and right ventricles which pumps blood away from the heart by rhythmic contractions. The human heart can be considered as two pumps – the right side pumps deoxygenated blood from the systemic veins into the pulmonary circulation to the lungs, and the left side of the heart receives the oxygenated blood from the lungs and pumps it into the systemic circulation through the aorta for circulation to the rest of the body. Systole and diastole are the contraction and relaxation of the cardiac muscle and there is increased pressure due to contraction in the ventricles (systolic pressure) and decrease in pressure due to the relaxation of the ventricles (diastolic pressure). The pumping motion of the heart is obtained due to the contraction and relaxation of the heart muscle which coordinated by a meshwork of nerve fibers. The heart is surrounded by a connective tissue layer which is the pericardium and there are three layers in the outer wall of the heart – epicardium or visceral pericardium, myocardium and endocardium. Epicardium is the outer layer and the innermost layer is the endocardium which is 1 in contact with the blood that the heart pumps and consists of blood vessels and heart valves. The myocardium is the thickest contractile middle layer of the heart wall and it is the ventricular wall between the epicardium and endocardium [1]. It is composed of cardiac muscle cells that form the bulk of the heart and consists of muscle fibers and blood vessels which are connected and interspersed by a network of connective tissue. The uniqueness of heart lies in its dynamic functionality which requires sophisticated tissue architecture with specialized cellular and extracellular components. Every excitation and contraction cycle of the cardiac muscle involves a number of mechanical events and recent studies have shown more significant association of the ECM in all aspects of the electromechanically active myocardium than it was previously believed. Proteins present in the myocardial extracellular matrix (ECM) include collagen subtypes, glycoproteins, elastin, etc. and these proteins play a vital role in the organization and support of myocytes and the capillary network. Proteoglycans and other glycoproteins aid in the hydration of the ECM and act as lubricants for the cardiac contractile machinery. Fibronectin is a glycoprotein which is localized homogeneously throughout the extracellular space in which cardiac myocytes and collagens are embedded. 1.2 Myocardial Infarction Heart disease is the leading cause of death and disability in both industrialized nations and the developing world, accounting for approximately 40% of all human mortality, more than all cancers combined, and it is becoming a leading global threat of the 21st century. Heart attack or myocardial infarction (MI) is a condition in which there is a blockage or deterioration in the pumping efficiency of the heart resulting in fluid congestion or inadequate blood flow to tissues. This is a progressive disorder in which damage to the heart causes weakening of the cardiovascular system and is induced by several underlying diseases, which includes ischemic 2 heart disease with or without an episode of acute myocardial infarction, hypertensive heart disease, valvular heart disease and primary myocardial disease. MI occurs when one or more of the blood vessels supplying the heart occlude and there is decrease in the supply of nutrients and oxygen to that portion of the heart. If the blood flow is not restored immediately, it will cause irreversible cell death and the adult heart cannot repair as the mature cardiomyocytes are unable to divide. MI leads to permanent loss of cardiac tissue, ultimately leading to end-stage heart failure and the pathological changes after MI are characterized by an initial inflammatory response and loss of cardiomyocytes. The death of cardiomyocytes initiates migration of macrophages, monocytes, and neutrophils into the infarct area, initiating the inflammatory response. Furthermore, matrix malleoproteases set off leading to infarct expansion and degradation of the ECM resulting in myocyte slippage. As a consequence of this, there is weakening in the collagen scaffold which eventually causes wall thinning and ventricular dilation. During the second phase, there is resistance to rupture and deformation due to the increase in the deposition of fibrillar, cross-linked collagen. Heart failure is generally accompanied by remodeling, a process which involves change in the ventricular shape and dimension. Remodeling occurs with an increase in myocardial and interstitial mass and negative left ventricular (LV) remodeling, and cause increased wall stress on the remaining viable myocardium. Various factors/processes such as myocyte hypertrophy, myocyte slippage and interstitial growth induce the remodeling process resulting in LV dilation. It is suggested that LV remodeling may contribute independently to the progression of heart failure. The consequence of MI is the formation of scar tissue which does not have contractile, mechanical and electrical properties as the native myocardium. The scar tissue decreases the pumping efficiency of the ventricles and certain compensatory mechanisms activate in response to the reduced cardiac 3 output which places an extra burden on the already weakened heart accelerating end-stage heart failure [2]. Heart failure or MI remains one of the biggest challenges in the field of medical sciences with great improvements are expected to progress in the tissue engineering (TE) field for regeneration of the heart, in the near future. Figure 1.1: Schematic diagram illustrating the damage caused by MI in human hearts [2] 4 1.3 Hypothesis and objectives Hypothesis This project is to develop an ideal substrate for myocardial tissue engineering using electrospun natural proteins. We hypothesize that the use of cardiac-specific proteins, which are of significance post MI such as fibrinogen, along with another natural protein (collagen) will provide a physiologically relevant environment for the cardiac myocytes to adhere and proliferate and also promote the possibility of mimicking the myocardial ECM in terms of biocompatibility, morphology, surface characteristics, tensile strength and stiffness, which are critical for the regeneration of the infarcted myocardium. Objectives  Develop a temporary cardiac-specific extracellular matrix to help in myocardial repair by fabricating a fully natural polymeric nanofibrous composite scaffold by electrospinning.  Optimize the composition of the natural proteins in the composite matrix and the processing parameters for electrospinning.  Improve the mechanical properties of the natural polymeric composite in order to suit for cardiac tissue engineering application.  Evaluate the morphological, chemical and mechanical properties of the fabricated natural polymeric scaffolds.  Determine the adhesion and proliferation of human cardiomyocytes and investigate the differentiation of ADSCs on the fabricated natural polymeric scaffolds by protein expression studies. 5 Chapter 2 Literature Review 2.1 Introduction Despite extensive surgical treatments, the number of patients with heart failure continues to increase. Over the past 30 years, heart transplantation has been the backbone of therapy for the treatment of heart failure and it has proved to be one of the most effective therapies for end-stage heart failure. Unfortunately, the major obstacle with heart transplantation is the limited supply of suitable donors. In addition, lifelong immune suppression often causes serious consequences [3]. Another surgical approach for the treatment of MI which is gaining significant attention in the past decade is the TE approach which aims in mimicking the myocardial ECM. The ECM, by definition, is the organic matter that is found between the cells in plants and animals. The ECM is a relatively stable structural material that lies under the epithelia and surrounds the connective tissue cells and provides a stable framework for multicellular organisms under gravity and physical loading, whereby it maintains the integrity of tissues enabling physiological functioning. Over the years, there is a gradual change in our understanding of the ECM as a static ‘connective tissue’ that binds everything together to one of the dynamic biomaterial that performs multiple functions such as providing strength and elasticity, activating growth factors during development and controlling their availability, cell-surface receptor interactions etc. Besides, ECM is essential for morphogenesis and assist in the regeneration of multicellular organisms and tissues. The integrin receptors on the cell surface along with the ECM can be pictured as intricate nanodevices that allow cells to physically organize their 3D environment, and sense and respond to various types of mechanical stress [4]. Although the composition of 6 ECM represents a complex alloy of variable members of diverse protein families such as the collagens, proteoglycans, glycosaminoglycans and elastin, their main function is to support the tissue with specific mechanical and biochemical properties. For example, the collagens are a source of strength to the tissues; elastin and proteoglycans provide matrix resiliency and other structural glycoproteins aid in inducing tissue cohesiveness. 2.2 Tissue Engineering TE is a significantly advancing multi-disciplinary field that engages the principles of engineering, biology and life sciences with the ultimate goal of restoring the native tissue function [5]. It enables the injured tissue to regenerate using a biomimetic approach with the help of three important factors such as biomaterial-based supporting scaffolds, functional cells and specific growth factors, among which the biomaterials play a dominant role as they can control and improve the cell retention, proliferation and differentiation. The objective of TE is to repair the injured tissues by supplementing functional cells, supporting scaffolds, growth-factors or genes and electric or physiologic signals to the organs when essential. Cardiac tissue engineering (CTE) has grown as an indispensable field of research considering the increase in the number of heart failures and it has shown tremendous improvement in the past 10 years. Although it is very difficult to produce a perfect replica of the native cardiac tissue, there are certain important factors to be satisfied for the development of such myocardial models [6]: (i) clinical viable cell source with physiological composition of different cell types for each tissue component; (ii) biomaterial properties that closely match the physiological composition of cardiac ECM; (iii) easily controllable biomaterial degradation kinetics with degradation products being safely removed from the body; (iv) tissue construct needs to be non-thrombogenic and immune tolerant; (v) functional, biological, and histological properties of the tissue constructs should be 7 comparable to normal mammalian cardiovascular tissue; and (vi) the tissue construct need to withstand physiological stresses within the complex fluid environment of the cardiovascular system. Current, TE modalities include (i) in vitro engineered cardiac tissue approach – culturing cells on a biomaterial scaffold in vitro and implanting the tissue into the epicardial surface (ii) Implementation of a cardiac patch – populating the designed patch with isolated cells in vitro and further implanted in vivo (iii) Injectable systems – injecting cells and/or scaffold directly into the infarcted wall to create in situ engineered cardiac tissue. A cardiac patch is a three-dimensional matrix comprised of natural or synthetic biomaterials that host the cells and provide mechanical and structural support for the injured heart. By culturing contractile cells onto a 3D scaffold, the formation of functional cardiac patches can be induced. Implantation of these patch materials involves an invasive open chest procedure, such as sternotomy or thoracotomy. Patch materials can also be sutured to the epicardial surface of the heart, limiting the region of therapeutic benefits. These tissue-engineered patches may be utilized in cases of acute MI, augmenting lost contractile function of the left ventricle. 2.3 Biomaterials A biomaterial is defined as a non-viable material used in a medical device, intended to interact with biological systems [7, 8]. Biomaterials were mostly studied for applications in orthopedics and for development of prosthetics. Recently, novel molecularly designed biomaterials which can deliver growth factors and control the environment of transplanted cells are being developed. Biomaterials are required for most of the TE approaches and the major functions of an ideal biomaterial are to enhance cell adhesion, proliferation and differentiation. The general requirements for a biomaterial are that it should be biocompatible - not induce an inflammatory response, biomimetic – reflect the ECM of the tissue, biodegradable – possess appropriate 8 degradation kinetics and degradation products should be non-toxic, and possess suitable mechanical integrity. It is also essential that the biomaterial aids in vivo revascularization as well as integration with the host tissue. There are additional specific requirements for a biomaterial to be applied for CTE. It includes that the biomaterial is able to tolerate the continuous stretching/relaxing motion of the myocardium that occurs at each heartbeat. It is favorable if the biomaterial exhibits a nonlinear elasticity of heart muscle such that it could reshape with the heart and thus provide mechanical support throughout the beating mechanism. Other specific criteria for CTE are that the biomaterial should encourage cardiomyocyte alignment and maturation in vitro before implantation or in vivo and also, the biomaterial should enable electrical integration of engineered graft with the native tissue to allow synchronized beating between the artificial construct and the heart. Many natural and synthetic polymeric biomaterials such as collagen, gelatin, alginate, fibrinogen, etc. and PGA, PLLA, PCL, PGS, PNIPAAM, etc. respectively, are used for CTE and these biomaterials are used in several forms such as nanofibers, microspheres, injectables, meshes, sponges, etc. Biomaterials used in the form of injectable without inclusion of cells can decrease remodeling after MI by increasing thickness of the infarct leading to decreased wall stress on surviving myocardium [9, 10]. Fetal rat myocardial cells seeded into alginate sponges produced by freeze-drying technique were developed to engineered heart construct and were implanted into rat hearts with myocardial infarct. Intensive neovascularization was found and the specimens showed almost complete disappearance of the scaffold and good integration into the host after 9 weeks [11]. It is, therefore, preferred to choose the biomaterial or tailor-design the properties in order to produce robust yet flexible, contractible, electrophysiologically stable, readily vascularized myocardial construct. Table 1 provides the overview of the biomaterials used in CTE [12]. 9 Table 2.1 Overview of biomaterials used in cardiac tissue engineering [12] 10 2.4 Natural Proteins The selection of biomaterials for scaffold design is a major criterion and should be carefully considered based on the biodegradability of the polymer, its ability to deliver and foster cells and their mechanical properties for appropriate applications. Natural, ECM-derived proteins are apparently the first choice for soft TE, since they provide a physiologically relevant and recognizable platform for the cells to attach, proliferate and differentiate. Moreover, they are biocompatible, non-toxic, possess appropriate degradation kinetics and generate mild inflammatory response. On the other hand, synthetic polymers are non-physiological and act mainly to provide a physical or mechanical support to hold the cells. A slight change in the environmental pH upon degradation of the synthetic polymers is also harmful to the surrounding cells and in addition, the polarity, water absorption and degradation properties raise questions against the medical suitability of these synthetic polymeric scaffolds [13]. 2.4.1 Collagen The most plentiful proteins in the ECM are the collagen family of proteins and collagens form the fundamental organic matrix of the bone, skin, arteries, ligaments, cartilage and most of the ECM in general. Contributing ~30% of the total protein mass in mammals, the collagenous proteins are a broad class of molecules found with extreme heterogeneity. To date, 29 types of collagen have been identified in the collagen superfamily and these 29 types of collagen are discriminated by considerable complexity and diversity in their structure, their splice variants and the presence of additional non-helical domains, their assembly and their function [14]. Collagen network contribute chiefly to the cardiac ECM and they play a vital role in the 11 myocardial structure and function. Collagen provides strength and stiffness to the myocardium and establishes a structural framework for the myocytes and provides myocyte-to-myocyte connections (collagen struts) that are vital in adhering the cells. Five collagen isoforms are present in the myocardium – types I, III, IV, V and VI with the most abundant forms of collagen being type I and type III collagens each contributing 80% and 12% respectively and they also constitute for the bulk scar tissue following MI [15, 16]. Type I and type III collagen molecules form aggregate struts of varying thickness and are widely distributed between myocytes and among muscle fibers [17, 18] whereas type IV collagen arrange themselves to form end to end aggregates characterized by a fishnet appearance in the basement membrane of cardiac myocytes and fibroblasts [19, 20]. Type VI collagen is present as fine filaments in the myocardium oriented perpendicularly opposite to other collagen fibers [21]. Collagen types present in the myocardium are relatively insoluble and are characterized by abundant inter- and intra-molecular covalent cross-linkage [22]. Also, a hierarchy of decreasing tensile strength exists among cardiac collagen (type I > type III > type VI and fibronectin > basement membranes) so that changes induced by contraction and relaxation could be effectively distributed throughout the heart [16]. Przyklenk et al [23] found that the stiffness and tensile strength of the myocardium correlated directly with collagen content and as such, a collagen-rich matrix is critical in maintaining the integrity of the cardiac. 2.4.2 Gelatin Gelatin is a natural glycoprotein and an irreversibly hydrolysed form of collagen and the protein pathogens are removed during denaturing hydrolysis. Gelatin is derived from collagen found 12 inside animals’ skin and bones. It is obtained mainly from the most abundant collagen (type I), present in bone and skin (main sources include pig skin, cattle bones, cattle hide etc). A negatively charged acidic gelatin or a positively charged basic gelatin can be produced by changing the isoelectric point. This exclusive property of tailoring the electrical nature nature of gelatin can be applied for the sustained release of proteins from polymer matrices [24]. Gelatin is commonly used in several forms such as nanofibers, hydrogels, microspheres, etc. for various TE [25 - 29] and drug delivery applications [30 - 32]. Gelatin is applied as a gelling agent the food industry (e.g. gelling and foaming agent), in the pharmaceutical industry (e.g. soft and hard capsules, microspheres), in the biomedical field for wound dressing and TE and also in photography and cosmetic manufacturing. Gelfoam, a commercially available gelatin mesh has been used as a biodegradable scaffold to support the 3-dimensional growth of seeded cells such as fetal rat ventricular cardiomyocytes, gastric smooth muscle cells, skin fibroblasts and adult human atrial cardiomyocytes [33]. Recently, a group in China used gelatin microspheres to deliver vascular endothelial growth factor (VEGF) to enhance the efficacy of bone marrow mesenchymal stem cell transplantation in swine model of MI [34]. Gelatin is considered more advantageous compared to collagen because of its (i) abundant availability at an affordable cost; (ii) lower antigenicity than collagen; (iii) biodegradability and biocompatibility in physiological environments; (iv) ability to tailor the electrical nature of gelatin and flexibility in processing to suit diverse applications; The arginine-glycine-aspartic (RGD) sequences of collagen is also present in gelatin making it highly effective for cell adhesion [35]. 13 2.4.3 Fibrinogen Fibrinogen is a ubiquitous glycoprotein present in the human blood plasma at a concentration of about 2.0 - 4.5 g/L [36]. It is a (slightly) bent and twisted trinodular molecule consisting of six polypeptide chains 2Aα, 2Bβ and 2γ which are bound together by 29 disulfide bonds [37, 38]. Fibrinogen plays a major role not only in blood clotting and wound healing but also in fibroblast proliferation and defense mechanisms against infection. Of the total fibrinogen content in the human body, 80 – 90% is found in the blood plasma and it is vital for hemostasis, wound healing, inflammation, angiogenesis, neoplasia and performs myriad other functions like serving as an essential co-factor for platelet aggregation, a determinant of blood rheology, etc. The degraded products of fibrinogen and fibrin stimulate the migration and proliferation of smooth muscle cells and fibrinogen is also associated with cultured endothelial cells and helps in migration [39]. It is also believed that the specific and tightly controlled intermolecular interactions of fibrinogen domains influence several other aspects of cellular function and developmental biology. When fibrinogen is exposed to thrombin, two peptides are cleaved to produce fibrin monomers which in the presence of Ca2+ and factor XIII lead to the assembly of stable fibrous clots (insoluble gels) and/or other fibrous structures. These stable structures function as nature’s provisional matrix, on which tissues rebuild and repair themselves. This process is similar to the TE objective and taking this cue, fibrinogen along with another natural protein will be an ideal choice among scaffold fabrication for regeneration of the infarcted myocardium. 14 2.5 Nanofibers by electrospinning 2.5.1 Electrospinning – Basic principle Electrospinning refers to the technique of drawing very fine fibers on a nano scale by applying high electric potential and this method is applicable for any fusible polymer. The technique involves pumping a polymer solution or melt through a thin nozzle. The nozzle is connected to a high voltage of the order of 30KV thereby serving as an electrode. A grounded collector is used to collect the electro spun polymer. On pumping the polymer melt out, through the nozzle, the high electric field causes the polymer droplet to separate due to high electrostatic forces of repulsion on its way to the collector. This electrified jet undergoes elongation and whipping, leading to the formation of a long thin thread. The solvent subsequently evaporates and the polymer solidifies producing fibers of nano to micro diameter [40]. Electrospinning process does not require high temperature and therefore, it is suited for the manufacture of fibers using large and complex molecules. The major parameters influencing this process include molecular weight of the polymer, solution properties, electric potential, flow rate, needle diameter, distance between the needle and the collector, motion of the target collector, etc. This technique enables the production of interesting aligned textures and porous structures and the electrospinning parameters can be modified in different ways for combining material properties with different morphological structures for desired, specific applications. 15 Figure 2.1: Schematic electrospinning setup. (A) Polymer melt taken in syringe (B) Nozzle (C) High voltage transformer (D) Electrospun jet from nozzle (E) Collector (rotatory or stationary) 16 2.5.2 Advantages of electrospun nanofibers for tissue engineering:  The fibers in the nano scale offer much larger scope for engineering and biological applications than the conventional fibers. The nano structuring effect of fibers imparts many new properties to the system that can be exploited commercially as bacterial filters, drug delivery agents, wound healing, etc. [41].  Recently, there has been a center of attention given to the applications of nanofibers in tissue engineering. TE refers to the efforts of to perform specific biochemical functions using cells with an artificially created support system that may compromise various materials and then interaction with the human body environment. Electrospun nanofiber matrices show morphological similarities to the natural extracellular matrix characterized by ultrafine continuous nanofibers, high surface to volume ratio, high porosity and variable pore size distribution and have properties that can modulate the cellular behavior. These properties make them ideal and best suited to produce biological scaffolds [42].  In order to guide and orient cells in soft TE, such as the myocardial tissue which has an aligned texture, scaffolds with an aligned texture are desirable which can be produced by the electrospinning process. The cells take cues from the electrospun nanofibrous topography and maintain their appropriate phenotypes and lay down the ECM in vivo better on the textured scaffolds. 17 Chapter 3 Human cardiomyocyte interaction with electrospun fibrinogen/gelatin nanofibers for myocardial regeneration 3.1 Introduction Myocardial Infarction leads to end-stage heart failure and it is the major cause of death in many industrialized nations. Tissue engineering approaches for treatment of the infarcted tissue has gained huge attention over the recent years and research in this direction mainly aims for the optimization of a biomaterial scaffold with cell-source for tissue regeneration. In this regard, we fabricated a composite, but absolutely natural polymeric scaffold, using the blood protein, namely fibrinogen and the denatured collagen glycoprotein or gelatin by electrospinning process. Scaffolds with two different weight ratios of fibrinogen:gelatin (Fib:Gel) was prepared and cross-linking (CL) of the electrospun scaffolds was carried out using glutaraldehyde vapors for a short period of time to improve the mechanical properties of the scaffolds [Fib/Gel(1:4)-CL; Fib/Gel(2:3)-CL]. We evaluated the morphological, chemical and mechanical properties of the fabricated the natural polymeric scaffolds and identified the optimal amount of fibrinogen in the fabricated natural composite. Further, we supplemented the scaffolds with human cardiomyocytes and studied the attachment and proliferation of the cardiomyocytes and analyzed the cell-scaffold interactions to determine the suitability of the fabricated scaffolds for myocardial TE. 18 3.2 Materials and methods 3.2.1 Materials Gelatin type A, Fibrinogen from bovine plasma, and 1,1,1,3,3,3 – hexafluoro-2-propanol (HFP) were purchased from Sigma-Aldrich (Singapore). Human cardiomyocytes and the Myocyte Growth Medium were purchased from Promocell, Germany. Dulbecco modified eagle’s medium (DMEM), fetal bovine serum (FBS), phosphate buffered saline (PBS), antibiotics and trypsinethylenediaminetetraacetic acid (EDTA) were purchased from Gibco, Invitrogen Corp., USA. Hexamethyldisilazane (HMDS) and polyvinylalcohol mounting medium were obtained from Fluka, Singapore and CellTiter96 Aqueous one solution was purchased from Promega (WI, USA). Monoconal anti-α-actinin, anti-troponin-T, anti-connexin-43, anti-cardiac myosin heavy chain produced in mouse were all purchased from Abcam, Hongkong. 3.2.2 Fabrication of scaffolds by electrospinning Fibrinogen was dissolved in HFP by stirring for a period of 24 hours and gelatin was further added to make a total of 10% (w/v) solution. Two different compositional ratios of fibrinogen: gelatin, namely 20:80 (Fib/Gel 1:4) and 40:60 (Fib/Gel 2:3) was prepared. The polymer solutions were electrospun from a 3-mL syringe using a 0.5 mm blunted stainless steel needle at a high voltage of 15 KV to obtain Fib/Gel (20:80 & 40:60) nanofibers. The flow rate of the polymer solutions was maintained at 0.85 mL/h using a syringe pump (KD Scientific, Holliston, USA) and the drawn fibers were collected on a flat aluminum foil wrapped around the collector or on 15 mm glass cover slips placed at a distance of 10 cm from the needle tip. The electrospinning process was conducted at room temperature (RT) and at a humidity of 50%. The collected electrospun Fib/Gel (1:4 and 2:3) nanofibers were vacuum dried to remove any residual solvent 19 present. Fibrinogen was dissolved in HFP at a weight ratio of 12% and electrospun at a high voltage of 20 kV to obtain pure fibrinogen fibers, which served as the control scaffold. The collected pure fibrinogen nanofibers were vacuum dried to remove any residual solvent present and were used as a control for characterization and cell culture experiments. 3.2.3 Glutaraldehyde cross-linking The electrospun Fib/Gel (1:4 and2:3) and pure fibrinogen nanofibers were cross-linked using glutaraldehyde vapors to improve their mechanical integrity and structural stability. In short, the electrospun nanofibers were placed in a petri dish and kept under 50 % glutaraldehyde vapors at RT in a closed container. The cross-linking time was limited to 30 minutes in order to maintain the non-cytotoxic nature of the natural polymeric nanofibers. The cross-linked nanofibrous scaffolds were vacuum dried for 24 – 48 hours. The composite scaffolds of Fib/Gel nanofibers after crosslinking were named as Fib/Gel(1:4)CL and Fib/Gel(2:3)-CL, respectively for Fib/Gel (1:4) and Fib/Gel (2:3) nanofibers. The crosslinked fibrinogen scaffolds were named as Fib-CL throughout this manuscript. 3.2.4 Morphological, chemical and mechanical characterization of the electrospun natural polymeric nanofibrous scaffolds The cross-linked electrospun nanofibrous scaffolds were sputter coated with gold (JEOL JFC1200 Auto Fine Coater, Japan) and the morphology was analyzed using scanning electron microscopy (JSM5600, JEOL, Japan) at an accelerating voltage of 10 – 20 KV. Image analysis software (Image J, National Institutes of Health, USA) was used to determine the diameter of the nanofibers from the SEM micrographs. 20 Attenuated total reflectance Fourier transform infrared (ATR-FTIR) spectroscopic analysis of the cross-linked electrospun nanofibrous scaffolds was done using Avatar 380 (Thermo Nicolet, Waltham, MA) over a range of 1000 – 3500 cm-1 at a resolution of 4 cm-1. The wettability or the hydrophilic/hydrophobic nature of the cross-linked electrospun nanofibrous scaffolds was measured by drop water contact angle measurement using VCA Optima Surface Analysis System (AST products, Billerica, MA). The cross-linked electrospun nanofibers on coverslips were placed on the testing plate under the needle and deionized water was used for drop formation. The droplet size was 0.5 µl. The tensile properties of the cross-linked electrospun nanofibrous scaffolds were obtained using Instron table-top tensile tester (Instron 5943, MA, USA) at a load cell capacity of 10 N, cross head speed of 5 mm/min and gauge length of 20 mm under ambient conditions of 24oC and 34% humidity. The scaffolds were soaked in PBS for a time period of 24 hours and the mechanical strengths of the PBS-soaked scaffolds were also evaluated during this study. The specimen preparation was done by cutting the cross-linked electrospun scaffolds to rectangular shape of dimensions 10 mm breadth x 30 mm length. Six specimens of each scaffold type were tested and the stress-strain curve was plotted. The tensile stress-strain values obtained from the instrument were plotted using an Excel sheet. 3.2.5 Cell culture on the electrospun scaffolds Four separate samples were prepared during this study: Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, FibCL and the control tissue culture plate (TCP). The electrospun fibers collected on 15 mm diameter glass cover slips were placed in a 24-well plate, pressed with a stainless steel ring and 21 sterilized under UV light for 2 hours. The electrospun scaffolds on the cover slips were washed thrice with PBS at an interval of 15 minutes each to remove any residual solvent or crosslinking agent, and soaked in DMEM overnight before cell seeding. Human cardiomyocytes were cultured in myocyte growth medium supplemented with 10% FBS and 1% antibiotics (penicillin 100 units ml-1 and streptomycin 100 µg ml-1) in 75 cm2 cell culture flask. The tissue culture flask was kept in an incubator at 37oC with 5% CO2, and the media was changed every alternate day. After the cells became confluent, they were detached from the flask using 1 x Trypsin, centrifuged, counted by using a hemocytometer and were seeded onto the scaffolds at a seeding density of 10,000 cells per well. The proliferation efficiency of the human cardiomyocytes seeded on the electrospun, crosslinked natural protein substrates was studied over an 8-day period using a colorimetric MTS (3(4, 5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium, inner salt) assay. On days 2, 4, 6 and 8, the media was removed from the 24-well plates and the cell-scaffold constructs were washed once with PBS. The samples were, further added with 20% of cell titre reagent and serum-free DMEM and, incubated for 3 hours at 37oC in a 5% CO2 incubator. After 3 hours, the contents were aliquot into a 96-well plate and the absorbance was measured at 490 nm using a microplate reader (FLUOstar OPTIMA; BMG Lab Technologies, Germany) and the results were exported to Excel sheet. The cell morphology of the cell-seeded electrospun construct was studied using SEM. After 8 days of cell seeding on the nanofibrous substrates, human cardiomyocyte-seeded scaffolds were processed for Scanning Electron Microscopy (SEM) evaluation. The cell-scaffold construct was washed with PBS and fixed using 3% glutaraldehyde in PBS for 3 hours. The samples were 22 dehydrated with a series of ethanol gradients ranging from 50%, 75%, 90% to 100%. Further, the samples were air dried with HMDS, gold coated and subsequently, taken for SEM analysis. Immunocytochemical studies were performed for evaluating the functional capability of the human cardiomyocytes after seeding them on the electrospun nanofibrous scaffolds. Alphaactinin (Act), Troponin I (Trop I), Connexin (Con-43) and Myosin Heavy chain (MHC) were the four different cardiac-specific proteins used as cardiac marker proteins of this study. After 8 days of cardiomyocyte seeding on the cross-linked electrospun scaffolds, the media was aspirated out and the scaffolds were washed with PBS, fixed using formalin for 20 minutes at RT and permeated using 0.1% Triton-X100 for 5 minutes at RT. Non-specific sites is blocked using 3% bovine serum albumin (BSA) for 90 minutes period. Monoclonal mouse primary antibodies (Act and Trop I) were used at a dilution of 1:100 and stained for a period of 90 minutes at RT. The samples were then washed thrice with PBS and the secondary antibody (Alexa 488 anti-mouse, green) at a dilution of 1:400 and 4’, 6-diamidino-2-phenylindole dihydrochloride (DAPI) at a dilution of 1:1000 was added and kept for 90 minutes at RT. The composite scaffolds were also immunostained for Con-43 and MHC, using the same protocol as stated above, except that the secondary antibody used was AlexaFluor 594 (anti-mouse, red). The cell-scaffold constructs were washed with PBS several times to remove the excess staining and the samples were mounted onto rectangular glass slides using fluromount and taken for laser scanning confocal microscopy analysis (LSCM, Fluoview FV300, Olympus). 23 3.3 Results 3.3.1 Morphological, chemical and mechanical characterization of the electrospun natural polymeric scaffolds Figure 3.1 represents the SEM micrographs of Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and the Fib-CL scaffolds, obtained as a result of the optimized electrospinning parameters. SEM images of the electrospun natural polymeric scaffolds showed uniform randomly oriented beadless fibers in the nano-scale range. The fiber diameters of the Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL scaffolds were in the range of 309 ± 18 nm, 161 ± 24 nm, and 133 ± 12 nm, respectively. Figure 3.1: Morphology of electrospun (A) Fib/Gel(1:4)-CL (B) Fib/Gel(2:3)-CL and (C) Fib-CL nanofibers 24 ATR-FTIR images (Figure 2) of the Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL scaffolds revealed the presence of the characteristic bands of fibrinogen and gelatin. Fib-CL nanofibers showed peaks at 1656 and 1540 cm-1 representing amide I and amide II bands which are characteristic of proteins with high α-helix content. Fib/Gel(1:4)-CL scaffolds showed peaks at 1640, 1545, and 1237 cm-1 and the Fib/Gel(2:3)-CL scaffolds showed peaks at 1650, 1539 and 1230 cm-1. The peaks shown by the Fib/Gel(1:4)-CL and Fib/Gel(2:3)-CL scaffolds correspond to amide I, amide II and amide III bands representing C=O stretching, N-H bending, C-N stretching and N-H bending, respectively confirming the triple helical structure of gelatin. There was no distinct new peak which shows no particular chemical reaction between the fibrinogen and gelatin molecules. The hydrophilic/hydrophobic nature of the Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL scaffolds was measured using drop water contact angle studies. The water-contact angles of the Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL, along with their fiber diameters are listed in Table 3.1. 25 Figure 3.2: FTIR spectra of the electrospun nanofibers Table 3.1 Fiber Diameters and water-contact angles of the electrospun Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, and Fib-CL nanofibers Sample/Parameter Fib/Gel(1:4)-CL Fib/Gel(2:3)-CL Fib-CL Fiber Diameter(nm) 309 ± 18 161 ± 24 nm 133 ± 12 Water-contact angle 30.32o ± 2o 37.95o ± 3o 99.07o ± 6o 26 The mechanical properties of the electrospun as such cross-linked scaffolds, as well as the PBSsoaked (24 hours) scaffolds were determined by evaluation of their tensile strength and strain at break. The tensile strength and stiffness properties of the dry and PBS-soaked Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL scaffolds are listed in Table 3.1 and the mechanical properties of the native myocardium are also shown in the table for comparison purposes. Table 3.2 Tensile strength and stiffness properties of the electrospun Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, and Fib-CL scaffolds compared to the native human myocardium Dry Scaffolds PBS-Soaked Scaffolds (for 24 hours) Sample/ Human Parameter Fib/Gel(1:4)CL Fib/Gel(2:3)CL FibCL Fib/Gel(1:4)CL Fib/Gel(2:3)CL FibCL myocardium Tensile Strength(MPa) 1.2 0.98 0.061 0.0125 0.009 0.0002 0.003 – 0.015 Stiffness(MPa) 2.5 1.84 0.830 0.4600 0.380 0.0030 0.2 – 0.5 27 Figures 3.3 & 3.4 show the stress-strain curves of the dry and PBS-soaked fabricated electrospun fibers, respectively. The tensile strength and the elastic modulus for the dry Fib/Gel(1:4)-CL scaffolds were obtained as 1.2 MPa and 2.5 MPa respectively; and for the dry Fib/Gel(2:3)-CL scaffolds the values were 0.98 MPa and 1.84 MPa. Figure 3.3: Stress-strain curves of electrospun and dry scaffolds of Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL nanofibers 28 The tensile strength of the PBS-soaked scaffolds were found reduced in the order for Fib/Gel(1:4)-CL > Fib/Gel(2:3)-CL > Fib-CL scaffolds and the same order was also followed for their stiffness values (0.46 MPa, 0.38 MPa and 0.0030 MPa, respectively). However the tensile strength and elastic modulus for cross-linked fibrinogen scaffolds (0.061 MPa) was less than that of the composite scaffolds, which also highlight the non-feasibility of using pure fibrinogen scaffolds for tissue regeneration. Figure 3.4: Stress-strain curves of electrospun PBS-soaked Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL and Fib-CL nanofibers 29 3.3.2 Cell culture studies The proliferation efficiency of the cardiomyocytes on the electrospun scaffolds was determined by MTS assay which was studied for a period of 8 days. Figure 3.5 shows the results of the MTS assay, where the cell proliferation increased with culture time for all the scaffolds and TCP. Figure 3.5: Human cardiomyocyte proliferation on electrospun nanofibers measured by MTS assay. *Significant against proliferation on Fib-CL nanofibers at p≤0.05. 30 The morphology of the human cardiomyocytes grown on the fabricated scaffolds was studied using SEM analysis and the results are shown in Figure 3.6. Compared with the fibrinogen scaffolds, higher cell attachment was observed on the composite Fib/Gel nanofibers and it was found that the nanofibrous topography of the fibers allowed the cardiomyocytes to make extensive use of provided cues for isotropic or anisotropic growth. Figure 3.6: Interaction of Human cardiomyocytes on various substrates (A) Fib/Gel(1:4)-CL (B) Fib/Gel(2:3)-CL (C) Fib-CL (D) TCP 31 Immunocytochemistry studies were carried out in detail using four cardiomyocyte specific marker proteins – Act, Trop I, Con-43 and MHC. Act is a cardiac-muscle-specific intermediate filament protein and is an important constituent of the contractile apparatus [43]. Trop I is a calcium receptive protein complex and a key regulatory protein in cardiac muscle contraction and relaxation [44]. Con-43 is the main constituent of cardiomyocyte gap junctions and is essential for cell-cell coupling and normal cardiac function [45] and MHC is a cardiomyocytespecific gene [46]. Results of the immunocytochemical analysis are shown in Figures 3.7 & 3.8. We observed that the Fib-CL scaffolds supported fewer cell growth compared to the cell growth on Fib/Gel(1:4)CL and Fib/Gel(2:3)-CL matrices. Protein expressions using Con-43 and MHC were carried out on Fib/Gel(1:4)-CL and Fib/Gel(2:3)-CL scaffolds and the results which are shown in Figure 3.9, further strengthened the application of these composite scaffolds for CTE. 32 Figure 3.7: Cardiac-specific-protein expressions of actinin on Fib/Gel(1:4)-CL, Fib/Gel(2:3)-CL, Fib-CL and TCP: (A, D,G, J); cell nuclei stained blue; (B, E, H, K); merged images of cell nuclei and actinin. (C, F, I, L) 33 Figure 3.8: Cardiac-specific-protein expressions of troponin I on Fib/Gel(1:4)-CL, Fib/Gel(2:3)CL, Fib-CL and TCP: (A, D,G, J) cell nuclei stained blue; (B, E, H, K) cell troponin stained green and merged images of cell nuclei and troponin (C, F, I, L) 34 Figure 3.9: Cardiac-specific-protein expressions of connexin on (A) Fib/Gel(1:4)-CL and (B) Fib/Gel(2:3)-CL; expression of Myosin Heavy Chain on (C) Fib/Gel(1:4)-CL and (D) Fib/Gel(2:3)-CL 35 3.4 Discussion Electrospinning is a simple yet inexpensive method for preparation of fibers in the nano-scale range from polymeric solutions by supplying high electric voltage. The electrospun fibers in the nanoscale range offer larger scope for TE applications as they possess morphological similarities to the native ECM characterized by ultrafine continuous nanofibers, high surface to volume ratio, high porosity and variable pore size distribution and have properties that can modulate the cellular behavior. In the past decade, electrospinning has become one of the desirable processes for fabricating matrices for TE [47, 48]. However, the major requirement for CTE is the establishment and maintenance of physiologically high density of viable cells. In the human body, the cardiac myocytes are the most physically energetic cells contracting for more than 3 billion times in an average human lifetime and pumping over 7,000 liters of blood per day along 100,000 miles of blood vessels. Therefore, it is vital for tissue engineers to apprehend the intricacies involved in the cardiac muscle and choose a material which when supplied with beating cardiomyocytes does not (i) hinder the contractile effect (ii) increase the diastolic stiffness (iii) impede relaxation and (iv) increase the overall wall stresses. Therefore, it would be more appropriate to use natural protein based scaffolds, whose physical and mechanical characteristics lie within the range of the native myocardium as the basic strategy such that the construct provides support to the nascent cardiomyocytes and subsequently integrate into the host and biodegrade when assimilation gets completed. 3.4.1 Importance of the blood plasma protein – fibrinogen in myocardial regeneration Fibrinogen is a naturally occurring, globular blood plasma protein with a molecular mass of 340,000 Da. The pathophysiological process following an injury involve fibrinogen undergoing 36 several complex coagulation processes leading to the formation of fibrous structures which have the property to facilitate as nature’s provisional matrix based on which further repair of the tissue takes place. In the case of the myocardium, fibrinogen along with fibronectin is crucial for the formation of “primary matrix” (in developing the granulation tissue) which functions as a meshwork for the deposition and adhesion of other matrix proteins such as interstitial collagens post MI [49, 50]. It has been observed that the fibrinogen infiltrates to individual cardiomyocytes, approximately 12 to 24 hours post MI and construct the framework for the primary matrix. Also, the degradation products of fibrinogen like the smaller fragments of peptides are found to show very unique properties such as (i) improve healing promotion both in vitro and in vivo [51] (ii) protective effect against myocardial reperfusion injury in rats because of its ability to prevent leukocyte migration (iii) reducing infarct size (iv) preserving the endothelial barrier function and alleviating the myocardial reperfusion injury [52, 53]. Fibrinogen has the ability to bind to a wide array of molecules that are favorable from a CTE perspective and it also contains RGD integrin binding sites and has the capacity to bind with functional VEGF, fibroblast growth factor and several other cytokines [54, 55]. However, in the past two decades, researchers figured out that fibrinogen, when present in higher levels in the heart, represent a major and independent risk factor for MI [56 - 58] and other cardiovascular diseases. There was it was also essential to determine the optimum amounts of fibrinogen that could be utilized for the fabrication of composite scaffolds, for application in CTE. We obtained fibrinogen nanofibers with diameters much closer to that of the native fibrinogen present in plasma clots [59]. Being a natural protein, such smaller diameter fibers may encounter breakage during contraction/relaxation or degrade at a very early stage as they do not possess the mechanical integrity and structural stability when used as a myocardial TE substrate. However, 37 with incorporation of gelatin, an increase in concentration of gelatin in the polymer solution resulted in increased fiber size for the composite scaffolds. Fib/Gel(1:4)-CL scaffolds presented the appropriate fiber morphology and diameter in the range of 300 nm which was found similar to that of the native proteins in the ECM of nearly 300 nm [60]. The electrospun Fib/Gel nanofibrous composite scaffolds were cross-linked using glutaraldehyde vapors for a limited time. Though various physical (UV-irradiation, dehydrothermal treatment) and chemical (formaldehyde, glutaraldehyde, carbodiimide, dextran dialdehyde) cross-linking methods are currently available, glutaraldehyde is still considered as the most effective cross-linking agent for collagenous proteins [61]. Glutaraldehyde cross-linking is found to be advantageous in decreasing the bio-degradation, improving the thermal stability, biological and mechanical integrity, strength and flexibility [62]. The risk of inducing toxicity can be lowered by decreasing the concentration of glutaraldehyde and the cross-linking time [63]. Pure gelatin nanofibers were fabricated by Zhang et al [64] and the glutaraldehyde cross-linked gelatin nanofibers were described to maintain the fiber morphology. These researchers evaluated the cytotoxicity of the scaffolds by culturing human dermal fibroblasts and they found a linear increase in cell growth over the culture time. In our study, we minimized the slightest possibility of introducing toxicity to the natural polymeric scaffolds by containing the cross-linking time to a very limited time period. It was also found from our experiment that cross-linking did not affect the surface properties of the Fib/Gel scaffolds. The pure fibrinogen scaffolds were hydrophobic and the Fib/Gel(1:4)-CL substrates with the least concentration of fibrinogen (20%) showed highest hydrophilicity. By reducing the content of fibrinogen within the composite (Fib/Gel), we were able to improve the hydrophilicity of the composite scaffold. Hence the Fib/Gel(1:4)-CL scaffolds were more hydrophilic than the Fib/Gel(2:3)-CL scaffolds . Gelatin contains hydroxyl 38 groups which can form hydrogen bonds with the water molecules thereby providing higher hydrophilicity to Fib/Gel(1:4)-CL scaffolds. The mechanical properties of the scaffolds are very critical, in particular, for scaffolds used for CTE. The objective of this study being to imitate the myocardial ECM, which means the fabricated substrate should possess some of the intrinsic mechanical properties of the native cardiac tissue. It is understood that one of the major consequences of MI is heart wall thinning and ventricular dilation which tremendously increases the heart wall stresses. In such cases, the use of a biomimetic polymeric scaffold should help in simple changes in ventricular geometry resulting in the reduction of elevated local wall stresses. Synthetic polymers have very high stiffness and when used as a myocardial construct it will result in diastolic dysfunction causing overstress of heart walls leading to progressive structural and functional changes in ventricles, prompting the end stage of heart failure. Natural polymers are often ignored accounting to their inferior mechanical strengths. But, for the purpose of CTE, it is essential to understand the innate mechanical properties of the heart wall and use an appropriate polymer to avoid such serious consequences. The stiffness of a healthy heart muscle is in the range of 10 – 20 KPa at the early stage of diastole and rise upto 50KPa at the end of diastole [65, 66]. These values should be used as the upper and lower limits of stiffness when choosing a suitable biomaterial for application as a cardiac construct. From Table 3.2, we can see that the fabricated Fib/Gel scaffolds showed tensile and stiffness properties comparable to that of the native myocardium. Since the ECM of myocytes is collagen, it is presumed that the cardiomyocytes could beat adequately in a scaffolding material with mechanical properties analogous to that of myocardial collagen, which also demonstrate the feasibility of using our electrospun Fib/Gel(1:4)-CL scaffolds for CTE. Human cardiomyocyte 39 cell proliferation was found significantly higher in the Fib/Gel scaffolds, in particular, the Fib/Gel(1:4)-CL, demonstrating their ability to cause better cell adhesion and proliferation compared to the Fib/Gel(2:3)-CL scaffolds. The Fib-CL scaffolds showed minimal proliferation which might be due to the inability of pure fibrinogen to support cell proliferation. SEM images of the matrices populated with cardiomyocytes showed that cell growth and cell spreading was more predominant on the Fib/Gel(1:4)-CL scaffolds. This can be attributed to the fact that gelatin has several integrin binding sites for cell adhesion and proliferation [67] and also that fibrinogen when present in optimized levels might positively assist the cell growth. Immunocytochemistry results showed that the Fib/Gel scaffolds contained greater number of cells with the cell nuclei surrounded with mature cytoskeleton and the functional integration of the human cardiomyocytes was found to be more pronounced on the Fib/Gel(1:4)-CL scaffolds. Also, from our work, we have established the fact that fibrinogen is an indispensable factor before and after MI. Moreover, we cannot neglect the use of fibrinogen as they favor the development of granulation tissue post MI and their potential advantage can be understood from the synonymity displayed by the Fib/Gel(1:4)-CL scaffolds to the myocardial ECM. We established the importance of using pure natural protein nanofibers as suitable substrates for restoring the myocardium after an injury/infarction. 40 3.5 Conclusion An impairment in the functioning of the heart due to an infarction, demand the cardiac muscle for a simple, fundamental necessity, manifested by an artificial cell-construct/matrix that can support the native tissue provide cues for heart tissue regeneration. We found that electrospinning a composite scaffold of fibrinogen and gelatin, which contains an optimized amount of fibrinogen, can act as a substrate suitable for cell growth and proliferation, and might enable the regeneration of the myocardial tissue. Our work highlight the outstanding resemblance of the fabricated natural polymeric Fib/Gel(1:4)-CL composite cell-constructs to that of the native myocardial ECM in terms of morphology, surface characteristics, tensile strength and stiffness properties and emphasize their compatibility and efficiency for the proliferation of human cardiac cells placing them in the forefront as the best candidates for myocardial TE. 41 Chapter 4 Electrospun collagen/fibrinogen nanofibrous scaffolds enhance cardiomyogenic differentiation of adipose-derived stem cells for myocardial tissue engineering 4.1 Introduction Human adipose-derived stem cells (ADSCs) possess the ability to differentiate along multiple lineage pathways and they can be easily isolated by a simple, minimally invasive method and in large numbers. ADSCs represent a beneficial source of cardiomyocyte progenitors and they are promising candidates to be used in CTE. ADSCs are gaining rapid focus among tissue engineers mainly because of their several useful properties. In this study, we investigated the in vitro cardiac differentiation of human ADSCs in co-culture with human cardiomyocytes on electrospun natural polymeric nanofibrous scaffolds. We fabricated a completely natural polymeric nanofibrous composite scaffold using the ubiquitous body protein, collagen and the blood plasma protein, fibrinogen. The optimized amount of fibrinogen in the natural composite was identified from our previous experiment (Chapter 3). Cross-linking using glutaraldehyde vapors was carried out for a short period of time to enhance the mechanical integrity of the fabricated matrices. Morphological, chemical and mechanical properties of the developed natural polymeric composite scaffolds were studied. The cell-scaffold interactions were analyzed by evaluating the cell proliferation using MTS assay, and SEM analysis was done to determine the cell morphology. Dual immunofluorescent staining was performed to further confirm the cardiogenic differentiation of ADSCs by using cardiac-specific marker proteins MHC and CD105. In vitro cardiomyogenic differentiation of human ADSCs in co-culture with human 42 cardiomyocytes on the natural polymeric composite scaffolds proved the suitability of our electrospun fibrinogen/collagen nanofibrous matrices for CTE. 4.2 Materials and Methods 4.2.1 Materials Atellocollagen powder was obtained from Kokimo, Japan. Human ADSCs were got from Lonza, Singapore and cardiac-specific marker protein CD 105 was purchased from Abcam, Hong Kong. 4.2.2 Fabrication of scaffolds by electrospinning Fibrinogen was dissolved in HFP by stirring for a period of 24 hours and collagen was further added to make a total of 10% (w/v) solution. From our previous experiment, the optimum amount of fibrinogen in the natural composite was identified and according to that result, the fibrinogen:collagen (Fib/Coll) solution was prepared to a ratio of 20:80. The polymer solution was electrospun from a 3-mL syringe using a 0.5 mm blunted stainless steel needle at a high voltage of 15 KV to obtain Fib/Col (20:80) nanofibers. The flow rate of the polymer solution was maintained at 0.85 mL/h using a syringe pump (KD Scientific, Holliston, USA) and the drawn fibers were collected on a flat aluminum foil wrapped around the collector or on 15 mm glass cover slips placed at a distance of 10 cm from the needle tip. The electrospinning process was conducted at RT and at a humidity of 50%. The collected electrospun Fib/Coll (20:80) nanofibers were vacuum dried to remove any residual solvent present. Pure fibrinogen nanofibers were prepared as described in Chapter 3 (Section 3.2.2). The collected pure fibrinogen nanofibers were vacuum dried to remove any residual solvent present and were used as a control for characterization and cell culture experiments. 43 4.2.3 Glutaraldehyde cross-linking The electrospun Fib/Coll (20:80) and pure fibrinogen nanofibers were cross-linked using glutaraldehyde vapors (described in Section 3.2.3) to improve their mechanical integrity and structural stability. The composite scaffolds of Fib/Coll (20:80) nanofibers after crosslinking were named as Fib/Coll(1:4)-CL and the cross-linked fibrinogen scaffolds were named as FibCL throughout this manuscript. The morphological, chemical and mechanical characterization of the electrospun Fib/Coll(1:4)CL nanofibers were carried out as described in Section 3.2.4. 4.2.4 Cell culture on the electrospun scaffolds The electrospun fibers collected on 15 mm diameter glass cover slips were placed in a 24-well plate, pressed with a stainless steel ring and sterilized under UV light for 2 hours. The electrospun scaffolds on the cover slips were washed thrice with PBS at an interval of 15 minutes each to remove any residual solvent or crosslinking agent, and soaked in DMEM overnight before cell seeding. Human cardiomyocytes were cultured in myocyte growth medium supplemented with 10% FBS and 1% antibiotics (penicillin 100 units ml-1 and streptomycin 100 µg ml-1) in 75 cm2 cell culture flask. The human ADSCs were cultured in low-glucose DMEM supplemented with 10% FBS and 1% antibiotic in a 150 cm2 cell culture flask. The tissue culture flasks were kept in an incubator at 37oC with 5% CO2, and the media was changed every alternate day. After the cells became confluent, they were detached from the flask using 1 x Trypsin, centrifuged and counted by using a hemocytometer. The scaffolds were separated into three groups: Fib/Coll(1:4)-CL, Fib-CL and TCP. In order to perform co-culture, human ADSCs and human cardiomyocytes were seeded onto the scaffolds in 44 the ratio of 1:1 at a seeding density of 4, 000 cells per well (2000 ADSCs:2000 cardiomyocytes). The proliferation efficiency of the human ADSCs and cardiomyocytes (co-culture) seeded on the electrospun, cross-linked natural protein substrates was studied over a 21-day period using a colorimetric MTS (3-(4, 5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4- sulfophenyl)-2H-tetrazolium, inner salt) assay. On days 7, 14 and 21, the media was removed from the 24-well plates and the cell-scaffold constructs were washed once with PBS. The samples were, further added with 20% of cell titre reagent and serum-free DMEM and, incubated for 3 hours at 37oC in a 5% CO2 incubator. After 3 hours, the contents were aliquot into a 96well plate and the absorbance was measured at 490 nm using a microplate reader (FLUOstar OPTIMA; BMG Lab Technologies, Germany) and the results were exported to Excel sheet. After 21 days of cell seeding on the nanofibrous substrates, the cell-seeded scaffolds were processed for SEM evaluation to study the morphology. The cell-scaffold constructs were washed with PBS and fixed using 3% glutaraldehyde in PBS for 3 hours. The samples were dehydrated with a series of ethanol gradients ranging from 50%, 75%, 90% to 100%. Further, the samples were air dried with HMDS, gold coated and subsequently, taken for SEM analysis. Immunocytochemical studies were performed for evaluating the functional capability of the human ADSCs and cardiomyocytes after seeding them on the electrospun nanofibrous scaffolds. The three groups of samples - Fib/Coll(1:4)-CL, Fib-CL and TCP, were further separated into co-culture scaffolds onto which human ADSCs and human cardiomyocytes were seeded in the ratio of 1:1 at a seeding density of 4, 000 cells per well (2000 ADSCs:2000 cardiomyocytes); positive control scaffolds, onto which cardiomyocytes were seeded at the same seeding density of 4, 000 cells per well; and the final set consisted of negative control scaffolds, onto which human ADSCs were seeded at the same seeding density of 4, 000 cells per well. 45 MHC and CD 105 were the two different cardiac-specific proteins used as cardiac markers in this study. After 21 days of seeding on the cross-linked electrospun scaffolds, the media was aspirated out and the scaffolds were washed with PBS, fixed using formalin for 20 minutes at RT and permeated using 0.1% Triton-X100 for 5 minutes at RT. Non-specific sites is blocked using 3% bovine serum albumin (BSA) for 90 minutes period. Monoclonal mouse primary antibodies (MHC) were used at a dilution of 1:100 and stained for a period of 90 minutes at RT. The samples were then washed thrice with PBS and the secondary antibody (Alexa 488 anti-mouse) at a dilution of 1:400 and 4’, 6-diamidino-2-phenylindole dihydrochloride (DAPI) at a dilution of 1:1000 was added and kept for 90 minutes at RT. Double immunofluorescent staining was further performed on the co-culture scaffolds to confirm the differentiation of human ADSCs into cardiomyocytes. The ADSC-cardiomyocyte co-culture cells plated on Fib/Coll(1:4)-CL, Fib-CL and TCP were stained with cardiac-specific marker MHC in the dilution 1:100 for 90 minutes at RT, before which the nonspecific sites were blocked with 3% bovine serum albumin. This was followed by the addition of secondary antibody Alexa Fluor 488 (anti mouse, green) in the dilution 1:400 and DAPI at a dilution of 1:1000 and kept for 90 minutes at RT. The samples were washed with PBS thrice to remove the excess staining. The samples were then treated with cardiac-specific marker protein CD 105 in the dilution 1:100 for 90 minutes at RT. This was followed by the addition of the secondary antibody Alexa Fluor 594 (anti mouse, red) in the dilution 1:400 for 90 minutes at RT. The samples were washed with PBS thrice to re-move the excess staining. The samples were then removed and mounted over a rectangular glass slides using fluromount and taken for confocal microscopic analysis. The pure culture scaffolds of human ADSCs were stained using the marker protein CD 105 and secondary antibody Alexa Fluor 594 (anti-mouse, red) and the pure culture scaffolds of human 46 cardiomyocytes were stained using the cardiac-specific protein MHC and the secondary antibody Alexa Fluor 488 (anti mouse, green) following the same procedure. 4.3 Results and Discussion 4.3.1 Importance of collagen in myocardial regeneration After MI, the collagen network is damaged and this weakened collagen network leads to wall thinning and ventricular dilation. Infarct expansion is closely associated with the impairment of the collagen network of the heart which is due to apparent loss of collagen struts which provide the additional damage essential for significant expansion. Further, when collagen has been deposited in the infarct to form the scar tissue, the infarcted myocardium becomes capable to resist further expansion to a certain extent suggesting that damaged collagen network is vulnerable while normal collagen is able to withstand expansion [68]. Therefore, efficient and rapid deposition of type I collagen is fundamental for healing since the collagen matrix provides a mechanically strong network, minimizes infarct expansion and resists maladaptive remodeling [69]. It is suggested that in spite of the deleterious consequences of degrading collagen fibrils, cleavage of type I collagen is essential for effective fibrotic healing after MI [70]. We fabricated fully natural, nanofibrous composite scaffolds using collagen and fibrinogen and cross-linked them using glutaraldehyde for a short period of time to enhance their mechanical properties. Figure 4.1 represents the SEM micrographs of Fib/Coll(1:4)-CL and the Fib-CL scaffolds, obtained as a result of the optimized electrospinning parameters. SEM images of the electrospun natural polymeric scaffolds showed uniform randomly oriented beadless fibers in the nano-scale range. The fiber diameters of the Fib/Coll(1:4)-CL and Fib-CL scaffolds were in the 47 range of 287 ± 10 nm and 133 ± 12 nm, respectively which is comparable to that of the natural proteins in the ECM [60]. Figure 4.1: SEM morphology of electrospun (A) Fib/Coll(1:4)-CL and (B) Fib-CL nanofibers The ATR-FTIR images of the Fib/Coll(1:4)-CL and Fib-CL scaffolds are shown in Figure 4.2. The peaks seen at 3200 and 3048 cm-1 are characteristic of collagen and represent the N-H stretching and C-H stretching, respectively. The peaks at 1650, 1587 and 1230 cm-1 correspond to amide I, amide II and amide III bands representing C=O stretching, N-H bending, C-N stretching, respectively confirming the triple helical structure of collagen. There was no distinct new peak which shows no particular chemical reaction between the fibrinogen and collagen molecules. 48 Figure 4.2: FTIR spectra of the electrospun Fib/Coll(1:4)-CL and Fib-CL nanofibers 49 Drop water contact angle studies were performed to determine the hydrophilic/hydrophobic nature of the Fib/Coll(1:4)-CL and Fib-CL scaffolds. Table 4.1 lists the water-contact angle measurements of the fabricated scaffolds. Table 4.1 Fiber Diameter & water-contact angles of the Fib/Coll(1:4)-CL, and Fib-CL nanofibers Sample/Parameter Fib/Coll(1:4)-CL Fib-CL Fiber Diameter (nm) 287 ± 10 133 ± 12 Water-contact angle 22.15o ± 2o 99.07o ± 6o The tensile strength and stiffness properties of the dry and PBS-soaked Fib/Coll(1:4)-CL and Fib-CL scaffolds were determined and the values are listed in Table 4.2. It can be observed from the table that the fabricated Fib/Coll(1:4)-CL scaffolds show tensile strength and stiffness properties comparable to that of the heart muscle. Figures 4.3 & 4.4 show the stress-strain curves of the dry and PBS-soaked, electrospun Fib/Coll(1:4)-CL and Fib-CL fibers, respectively. Table 4.2 Tensile strength and stiffness properties of the Fib/Coll(1:4)-CL and Fib-CL scaffolds Dry Scaffolds PBS-Soaked Scaffolds (24 hours) Sample/ Parameter Fib/Coll(1:4)-CL Fib- CL Fib/Coll(1:4)-CL Fib-CL Tensile Strength(MPa) 1.56 0.061 0.0143 0.0002 Stiffness(MPa) 2.90 0.830 0.5300 0.0030 *Human myocardium: Tensile Strength: 0.003 – 0.015 MPa; Stiffness: 0.2 – 0.5 Mpa 50 Figure 4.3: Stress-strain curves of electrospun PBS-soaked Fib/Coll(1:4)-CL, and Fib-CL nanofibers 51 Figure 4.4: Stress-strain curves of electrospun PBS-soaked Fib/Coll(1:4)-CL, and Fib-CL nanofibers 52 4.3.2 Myocardial regeneration potential of adipose-derived stem cells MI is considered to be one of the most challenging heart diseases for treating because it is irreversible and the cardiomyocytes do not possess the intrinsic regenerative capability to replace the lost cells. In principle, stem cells are the optimal cell source for myocardial tissue regeneration because they possess the ability to differentiate into one or more specialized cells, self-renew and produce unlimited number of stem cell progeny with similar properties throughout the lifetime [71, 72]. ADSCs demonstrate to differentiate along various mesenchymal lineages – adipogenesis, chondrogenesis, osteogenesis, myogenesis, etc. and also nonmesenchymal lineages like skeletal myogenesis, cardiogenesis, neurogenesis, etc. [73]. In addition to their good regeneration potential, ADSCs are also preferred for practical reasons. They are free of ethical, oncological and immunological concerns present in pluripotent stem cells. They are present in abundance, less invasive and less time-consuming to obtain compared to other tissue sources such as heart, skeletal muscle or bone marrow [74, 75]. Also, adipose tissue possess a higher stem cell density (5%) than the bone marrow (0.001%) [76]. The adipose tissue consists of mature adipocytes, preadipocytes, fibroblasts, vascular smooth muscle cells, endothelial cells, resident monocytes/macrophages and lymphocytes [77, 78]. ADSCs can be obtained in large amounts using liposuction procedure. The stromal-vascular cell fraction (SVF) of the adipose tissue is a rich source of pluripotent ADSCs and has become the focus of research for TE applications [79, 80]. Mature adipocytes are produced from the SVF, when the adipose tissue is collagenase digested. SVF shows phenotype characteristic of various populations including multipotent stem cells, hematopoietic, endothelial, smooth muscle cells, etc. [81]. The ADSCs are potent paracrine mediators and they are found to have significant effect on the 53 evolution of the ischemic myocardium [82 – 85]. The direct differentiation of transplanted cells and the paracrine actions upon damaged tissue is shown in Figure 4.5 [86]. Figure 4.5: Adipose-derived cells are capable of direct differentiation and paracrine actions upon damaged tissue [86] 54 Several animal studies have demonstrated the differentiation potency of ADSCs into cardiomyocytes [87] and reported that ADSCs can be useful to alleviate cardiac function in animals in vivo [88]. ADSCs co-cultured with rat cardiomyocytes demonstrated better differentiation expressed cardiac markers such as Act, Trop I and MHC. The ADSCs also showed contractions synchronous with the rat cardiomyocytes (upto 106 beats/min) [89]. In vivo differentiation study demonstrated that human ADSCs co-cultured with neonatal rat cardiomyocytes at 1:10 ratio in a vascularized cardiac tissue engineering model showed contractions of 140 beats/min. The ADSC-cardiomyocytes co-cultured group grew nearly twice as much vascular tissue in the construct than the cardiomyocyte group55% and 26% larger, respectively [90]. Immunocytochemical studies reported differentiation into cardiomyocytes, smooth muscle cells, adipocytes, recruitment into vascular structures, paracrine recruitment of endothelial cells and integration with the co-cultured rat cardiomyocytes. In order to prove that the fabricated natural polymeric scaffolds enhance differentiation of stem cells, we performed co-culture of human ADSCs along with human cardiomyocytes. The proliferation efficiency of the human ADSCs on the electrospun Fib/Coll(1:4)-CL, Fib-CL scaffolds and TCP after 7, 14 and 21 days was determined by MTS assay. The cell proliferation on Fib/Coll scaffolds was higher than that on the Fib-CL scaffolds and TCP. This is because of the presence of the natural myocardial ECM protein collagen and also the presence of fibrinogen (which helps in the formation of granulation tissue post MI) in optimal amount. Fig 4.6 shows the results of the MTS assay, where the cell proliferation increased with culture time for all the scaffolds and TCP. 55 Figure 4.6: Cell proliferation study of human ADSCs- co-cultured with human cardiomyocytes. *Significant against proliferation on Fib-CL nanofibers at p≤0.05. 56 The cell morphology of the differentiated ADSCs was studied using SEM images (Figure 4.7). Nanotopography is critical in TE as nanoscale dimensions are capable of modulating cellular responses and it can influence cell morphology, adhesion motility, proliferation, protein expression and gene regulation [91]. Compared with the fibrinogen scaffolds, higher cell attachment was observed on the composite Fib/Coll nanofibers and it can be seen that the fabricated Fib/Coll(1:4)-CL nanofibrous topography provided the necessary cues for differentiation of the human ADSCs into cardiomyocytes, integration with the co-cultured human cardiomyocytes and also for their isotropic or anisotropic growth. Figure 4.7 SEM images showing the morphology of ADSC-cardiomycytes co-culture cells on (A) Fib/Coll(1:4)-CL (B) Fib-CL and (C) TCP 57 To observe the myocardiogenic differentiation of the seeded human ADSCs, immunofluorescence stains of specific cardiomyocyte marker proteins like MHC and CD 105 were analyzed. The cardiac-specific marker protein expression is predominant in the co-culture system than the positive (pure culture of human cardiomyocytes) and negative control (pure culture of human ADSCs) groups. Figure 4.8 shows the cardiac protein expression using MHC on cardiomyocytes (pure culture) and as can be seen from Figure 4.9, the ADSCs express the ADSC-specific marker CD105 (A – C) and they do not express the cardiac-specific protein, MHC in their undifferentiated state (D – F), but they express DAPI alone which stains the nucleus. Dual immunocytochemical staining of the scaffolds using ADSC-specific CD 105 and cardiacspecific MHC is shown in the Figure 4.10. As can be seen the expression of the cardiac-specific protein is predominant in the co-culture system and this is because, the ADSCs that have undergone cardiogenic differentiation owing to the presence of cardiomyocytes in close proximity also express the marker proteins. In Fib/Coll(1:4)-CL scaffolds, more number of cells are found to express both the CD105 and MHC marker indicating that the differentiation is more in these scaffolds than in the Fib-CL scaffolds/TCP. This ascertains the fact that our fabricated natural composite scaffolds of Fib/Coll(1:4)-CL are best suited for stem cell differentiation and can be applied for CTE. 58 Figure 4.8: Cardiac-specific-protein expressions of MHC on Fib/Coll(1:4)-CL, Fib-CL and TCP: (A, D,G, J); cell nuclei stained blue; (B, E, H, K); merged images of cell nuclei and MHC. (C, F, I, L) on the human cardiomyocytes pure culture system 59 Figure 4.9: ADSC-specific protein expression of CD105 and MHC on Fib/Coll(1:4)-CL (A, D), Fib-CL (B, E) and TCP (C, F) on the human ADSCs pure culture system. 60 Figure 4.10: Dual immunofluorescent analysis for the expression of cardiac-specific marker protein MHC (A, D, G) and ADSC-specific marker protein CD 105 (B, E, H) and the merged image showing dual expression of both MHC and CD 105 (C, F, I) on Fib/Coll(1:4)-CL (A – C), Fib-CL (D – F) and TCP (G – I) on the co-culture system 61 4.4 Conclusion In the present study, we fabricated natural polymeric nanofibrous composite scaffolds of fibrinogen and collagen using electrospinning and cross-linked them using glutaraldehyde vapors – Fib/Col(1:4)-CL. We tested the morphology, chemical properties and mechanical integrity of our scaffolds and found them comparable to the native myocardial ECM. We used human adipose-derived stem cells which are capable of myocardiogenic differentiation to prove that our fabricated Fib/Coll(1:4)-CL substrates supported the differentiation of human ADSCs into cardiomyocytes and better cell proliferation, adhesion and integration was also found. In short, the fabricated Fib/Coll(1:4)-CL nanofibrous composites, with native tissue-like mechanical properties and desirable nanotopographical and stem cell differentiation cues, enhanced the differentiation of ADSCs into cardiomyocytes when co-implanted and has the potential to serve as an ideal scaffold for myocardial regeneration. 62 Chapter 5 Conclusion and Recommendations 5.1 Conclusion After MI, the myocardial tissue is scarred, lacks regeneration capabilities and even though, compensatory mechanisms occur, the scar tissue is never fully repopulated leading to accelerated cardiac failure. Repair of the infarcted myocardium involves three objectives – (i) replacement of a new myocardial tissue (ii) formation of a functional vascular network and (iii) restoration of the ventricle to its proper geometry. All of these depend on the existence of a temporary matrix which could replicate/mimic the myocardial ECM with appropriate mechanical properties, thereby not inducing further stress on the ventricular walls. It is important to provide the cells with the suitable environmental cues and also, the use of tissue-specific biomaterials will promote the possibilities of the success of the tissue engineering approach. It is self-explanatory from this fundamental understanding that natural myocardial ECM proteins will be the appropriate choice for treating the infarcted myocardium. In the heart tissue, the complex structural matrix of the ECM primarily consists of fibrillar collagen type I (80%) and type III (11%), and smaller amounts of type IV and type V collagens along with GAGs. Fibrinogen is the blood plasma protein which aids the formation of granulation tissue post MI. The importance of fibrinogen and collagen for myocardial TE is explained in detail in Section 3.4.1 and Section 4.3.1, respectively. This work has proved the feasibility of employing a completely natural polymeric composite scaffold fabricated by the electrospinning process for cardiac regeneration. We have identified the optimal amount of fibrinogen in the natural composite and optimized the electrospinning 63 process parameters to obtain nanofibers of appropriate fiber diameter and morphology suitable for myocardial TE. The fabricated scaffolds were found to have identical morphological, chemical and mechanical properties as the myocardial tissue. Cell culture studies using human cardiomyocytes and stem cell differentiation studies were carried out and the results established that the fabricated natural protein substrates provided the necessary structural, chemical and biological cues for cell-matrix interactions that reflect the natural microenvironment. Our work attempts to emulate the myocardial extracellular environment in vitro by the use of biomaterials (collagen and fibrinogen) which play a significant role post MI. Materials which are currently used for CTE do not provide a cardiac-specific milieu. To our knowledge, the use of cardiac-specific protein composite along with human cardiomyocytes for the regeneration of the ischemic myocardium is a very novel idea. 5.2 Recommendations We have demonstrated the feasibility of fabricating a fully natural polymeric composite matrix by electrospinning process for CTE and from the above results, it can be seen that the fabricated natural protein scaffolds reflect the cardiac-specific extracellular environment for the cells to proliferate and differentiate in vitro. From this preliminary study, we can focus on naturally-derived biomaterial-based strategies to address and ascertain this novel concept. Future research using cardiac-specific proteins and cells can focus on – 64 (i) In vivo animal studies for the regeneration of the injured tissue with the help of cardiac-specific biomaterials in the form of an injectable (hydrogel). 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Tissue Eng. 2010:120623. 71 [...]... progress in the tissue engineering (TE) field for regeneration of the heart, in the near future Figure 1.1: Schematic diagram illustrating the damage caused by MI in human hearts [2] 4 1.3 Hypothesis and objectives Hypothesis This project is to develop an ideal substrate for myocardial tissue engineering using electrospun natural proteins We hypothesize that the use of cardiac- specific proteins, which... cross-linked electrospun natural polymeric scaffolds were studied and found to be close to the mechanical properties of the heart Our electrospun natural polymeric scaffolds were found to provide better biocompatibility, hydrophilicity, biodegradability as well as suitable mechanical properties as desired for cardiac tissue engineering The cell-scaffold interactions of human cardiomyocytes with the electrospun. .. to help in myocardial repair by fabricating a fully natural polymeric nanofibrous composite scaffold by electrospinning  Optimize the composition of the natural proteins in the composite matrix and the processing parameters for electrospinning  Improve the mechanical properties of the natural polymeric composite in order to suit for cardiac tissue engineering application  Evaluate the morphological,... function is to support the tissue with specific mechanical and biochemical properties For example, the collagens are a source of strength to the tissues; elastin and proteoglycans provide matrix resiliency and other structural glycoproteins aid in inducing tissue cohesiveness 2.2 Tissue Engineering TE is a significantly advancing multi-disciplinary field that engages the principles of engineering, biology... modified in different ways for combining material properties with different morphological structures for desired, specific applications 15 Figure 2.1: Schematic electrospinning setup (A) Polymer melt taken in syringe (B) Nozzle (C) High voltage transformer (D) Electrospun jet from nozzle (E) Collector (rotatory or stationary) 16 2.5.2 Advantages of electrospun nanofibers for tissue engineering:  The fibers... major cause of death in many industrialized nations Tissue engineering approaches for treatment of the infarcted tissue has gained huge attention over the recent years and research in this direction mainly aims for the optimization of a biomaterial scaffold with cell-source for tissue regeneration In this regard, we fabricated a composite, but absolutely natural polymeric scaffold, using the blood protein,... engineered cardiac tissue approach – culturing cells on a biomaterial scaffold in vitro and implanting the tissue into the epicardial surface (ii) Implementation of a cardiac patch – populating the designed patch with isolated cells in vitro and further implanted in vivo (iii) Injectable systems – injecting cells and/or scaffold directly into the infarcted wall to create in situ engineered cardiac tissue A cardiac. .. therefore, preferred to choose the biomaterial or tailor-design the properties in order to produce robust yet flexible, contractible, electrophysiologically stable, readily vascularized myocardial construct Table 1 provides the overview of the biomaterials used in CTE [12] 9 Table 2.1 Overview of biomaterials used in cardiac tissue engineering [12] 10 2.4 Natural Proteins The selection of biomaterials for. .. growth because of their biological components To our knowledge, the idea of using electrospun fully natural- protein composite nanofibrous scaffolds supplemented with human adipose-derived stem cells/cardiomyocytes for the purpose of cardiac tissue engineering is a novel idea and they have the immense potential to be used for in vivo animal studies XIV Chapter 1 Introduction 1.1 Background The heart is... fabrication of a fully natural nanofibrous composite scaffold using electrospinning process is biocompatible, closely imitates the myocardial extracellular-matrix and offers the possibility to enhance cell proliferation The optimal amount of fibrinogen in the natural composite for cardiac tissue engineering application is identified To improve the mechanical integrity of the natural polymeric scaffolds ... project is to develop an ideal substrate for myocardial tissue engineering using electrospun natural proteins We hypothesize that the use of cardiac- specific proteins, which are of significance post... other therapies are experimented for the treatment of the infarcted myocardium, of which the tissue engineering approach is gaining much attention Cardiac tissue engineering promises to bring about... optimal amount of fibrinogen in the natural composite for cardiac tissue engineering application is identified To improve the mechanical integrity of the natural polymeric scaffolds they were