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1 IN VITRO ENDOTHELIALIZATION AND SHEAR CONDITIONING OF DECELLULARIZED VASCULAR ALLOGRAFTS By JOSEPH S. UZARSKI A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUI REMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY UNIVERSITY OF FLORIDA 2013
2 201 3 Joseph S. Uzarski
3 To my parents for their b elief in the value of education
4 ACKNOWLEDGMENTS There are numerous people who have assisted the completion of this dissertation through both direct involvement with the project and personal support provided to me whom I want to acknowledge for their contributions. I would first like to thank my mentor Dr. Peter McFetridge for offering me the opportunity to work on this innovative project in the summer of 2008 I thank him for his continuous guidance and research advisement over these years for always leaving his door open and for taking time to have impromptu discussions. I thank Dr. McFetridge for having a great sense of humor that made weekly lab meetings significantly less monotonous I also wish to thank the other members of my supervisory committee for their useful suggestions and critique of both my proposal and dissertation : Dr. Brandi Ormerod, Dr. Malisa Sa rntinoranont, and Dr. Edward Scott. I also thank my committee members for providing advice on venturing out into the unknown after graduation. I want to acknowledge all of my lab mates, research assistants, and other colleagues at both the University of Ok lahoma as well as the University of Florida for their cooperation I especially thank the past and present students of the McFetridge lab for their constructive criticism, collaboration, and for moral and intellectual support, for without them this work co uld not have been completed I individually thank the following students for direct advisement of this project: Salma Amensag, Cassandra Juran, Marc Moore, Kamil Nowicki, Dr. Zehra Tosun and Aurore Van d e Walle. I thank Jhon Cores for his contributions in the lab that directly assisted this work Alan Miles has provided invaluable consultation in the design, manufacture, and modification of parallel plate flow chambers, and for that I owe him my personal gratitude. Thanks to Dr. Phil Barish and Dr. Aaron T ucker for their advice on surviving graduate school.
5 Finally I would like thank my family and friends for always supporting me despite my having made minimal contact with many of them during my years in graduate school I thank m y parents Paul and Elizab eth, for their financial and invaluable moral support over the year s without which Lastly, I thank Elizabeth Haibel for her support and patience during this lengthy time period
6 TABLE OF CONTENTS page ACKNOWLEDGMENTS ................................ ................................ ................................ .. 4 LIST OF TABLES ................................ ................................ ................................ .......... 10 LIST OF FIGURES ................................ ................................ ................................ ........ 11 LIST OF ABBREVIATIONS ................................ ................................ ........................... 13 ABSTRACT ................................ ................................ ................................ ................... 14 CHAPTER 1 INTRODUCTION ................................ ................................ ................................ .... 16 Clinical Demand f or Vascular Grafts ................................ ................................ ....... 16 Current Vascular Graft Choices and Associated Limitations ............................ 17 Autografts ................................ ................................ ................................ ... 17 Prosthetic grafts ................................ ................................ ......................... 18 Allografts ................................ ................................ ................................ .... 19 Ex Vivo Derived Biomateri als and Vascular Tissue Engineering ............................ 20 Decellularization of Ex Vivo Derived Scaffolds ................................ ................. 21 The HUV as a Scaffold for Vascular Tiss ue Engineering ................................ 22 Anatomy of Blood Vessels ................................ ................................ ...................... 22 Physiological Roles of EC ................................ ................................ ....................... 23 EC Mechanotransduction of Shear Stress ................................ ........................ 23 Roles in Inflammation, Hemostasis, and Wound Healing ................................ 24 Roles in Vascular T one Regulation ................................ ................................ .. 25 Endothelialization of Vascular Grafts ................................ ................................ ...... 26 One stage Seeding ................................ ................................ .......................... 26 Two stage Seeding ................................ ................................ .......................... 27 Strategies to Improve EC Retention ................................ ................................ 28 History of EC Culture and Improvement of Perfusion Sy stems ............................... 28 Parallel Plate Flow Chambers (PPFC) ................................ ............................. 29 Application of SS Using Perfusion Systems ................................ ..................... 30 Summary and Present Objectives ................................ ................................ .......... 31 2 METHODS AND MATERIALS ................................ ................................ ................ 34 Experimental Methods ................................ ................................ ............................ 34 Human Tissue Collection ................................ ................................ .................. 34 Primary EC Isolation and Expansion ................................ ................................ 34 HL 60 Cell Cu lture ................................ ................................ ............................ 35 HUV Dissection, Decellularization, and Sterilization ................................ ......... 35 Computer aided Design ................................ ................................ ................... 36
7 Analytical Methods ................................ ................................ ................................ .. 37 Scanning Electron Microscopy ................................ ................................ ......... 37 Fluorescent Stains ................................ ................................ ............................ 37 Diamidino 2 Phenylindole (DAPI) ................................ ........................ 37 Rhodamine phalloidin (RP) ................................ ................................ ........ 37 Live/dead stain ................................ ................................ ........................... 38 Fluorescent Imaging and Analysis ................................ ................................ .... 38 Histology ................................ ................................ ................................ ........... 39 Tissue processing ................................ ................................ ...................... 39 Hematoxylin & eosin stain ................................ ................................ .......... 39 ................................ ................................ ........... 39 Color Imaging ................................ ................................ ................................ ... 40 Statistical Analysis ................................ ................................ ............................ 40 3 DECELLULARIZATION O F THE HUMAN UMBILICAL VEIN SCAFFOLD: EFFECTS ON ENDOTHELIAL CELL ADHESION AND VIABILITY ....................... 44 Methods and Materials ................................ ................................ ............................ 45 Decellulariz ation of HUV ................................ ................................ ................... 45 EC Seeding Assay ................................ ................................ ........................... 45 EC Viability Assay ................................ ................................ ............................ 46 Statis tical Analysis ................................ ................................ ............................ 46 Results ................................ ................................ ................................ .................... 47 HUV Isolation and Gross Morphology After Decellularization ........................... 47 Lumenal/ablumenal Surface Characterization ................................ .................. 47 EC Seeding and Culture ................................ ................................ ................... 48 EC Viability Assay ................................ ................................ ............................ 49 Lumenal Surface Morphology of Re endothelialized HUV ............................... 49 Discussion ................................ ................................ ................................ .............. 49 C onclusions ................................ ................................ ................................ ............ 52 4 IN VITRO METHOD FOR REAL TIME, DIRECT OBSERVATION OF ENDOTHELIAL CELL VASCULAR GRAFT INTERACTIONS UNDER SIMULATED BLOOD FLOW ................................ ................................ .................. 60 Methods and Materials ................................ ................................ ............................ 62 Particle Image Velocimetry ................................ ................................ ............... 62 Perfusion Culture System ................................ ................................ ................. 62 EC Seeding ................................ ................................ ................................ ...... 63 EC Viability Assay ................................ ................................ ............................ 63 EC Staining and Image Analysis ................................ ................................ ...... 63 Neutrophil Adhesion to Endothelialized HUV ................................ ................... 64 Results ................................ ................................ ................................ .................... 65 Flow Chamber Design and Assemb ly ................................ .............................. 65 Development of Neo endothelia on HUV Scaffolds and Adaptation to Flow .... 66 Shear Conditioning of Neo endothelia ................................ .............................. 67
8 Morphological Comparison of Endothelia Cultured Under Static or Flow Conditions ................................ ................................ ................................ ..... 68 Real time Observation of Neutrophil Rolling On Neo end othelia ...................... 68 Discussion ................................ ................................ ................................ .............. 69 5 ADAPTATION OF ENDOTHELIAL CELLS TO VARIABLE, PHYSIOLOGICALLY MODELED SHEAR STRESS STIMULATION: DISCRE PANCIES WITH FIXED MECHANICAL STIMULATION ........................... 79 Methods and Materials ................................ ................................ ............................ 81 EC Perfusion Culture ................................ ................................ ........................ 81 qPCR ................................ ................................ ................................ ................ 82 Nitrate/nitrite Quantification ................................ ................................ .............. 83 HL 60 Cell Adhesion Assay ................................ ................................ .............. 83 Statistical Analysis ................................ ................................ ............................ 83 Results ................................ ................................ ................................ .................... 83 PF Induces Morphological Adaptation In EC ................................ .................... 84 PF Induces Cardio Protective Gene Expression In EC ................................ .... 84 PF Does Not Significantly Modulate Expression of Coagulation/fibrinolysis Genes In EC ................................ ................................ ................................ .. 85 PF Induces Higher EC Expression of Chemotactic Factors Than SF ............... 86 PF Induces Sustained Endothelial Nitric Oxide Synthase Act ivity .................... 86 NO Production By EC Is Significantly Higher With Programmed Shear Changes Than Constant frequency Flow ................................ ...................... 88 PF Enhances E C Resistance to Activation Induced Leukocyte Adhesion ........ 88 Discussion ................................ ................................ ................................ .............. 88 6 ENDOTHELIAL CELL SEEDING AND FLOW PRECONDITIONING OF IM PLANTABLE HUMAN UMBILICAL VEIN ALLOGRAFTS ................................ 102 Methods and Materials ................................ ................................ .......................... 102 Perfusion Circuit Design ................................ ................................ ................. 102 Rotating Seeding Apparatus ................................ ................................ ........... 1 04 EC Seeding ................................ ................................ ................................ .... 105 Perfusion Culture ................................ ................................ ............................ 105 EC Staining and Image Analysis ................................ ................................ .... 106 Statistical Analysis ................................ ................................ .......................... 106 Results ................................ ................................ ................................ .................. 106 Rotational Seeding Improves the Circumferential Distribution of EC on the Lumenal HUV Surface ................................ ................................ ................ 106 Supplementation of Culture Media With Dextran Has No Effects on EC Proliferation or Viability ................................ ................................ ............... 108 Additional Maturation Time Improves Surface Coverage and EC Retention Under Flow ................................ ................................ ................................ .. 109 Flow Ramping Improves EC Resistance to Arterial Wall SS .......................... 110 Discussion ................................ ................................ ................................ ............ 111
9 7 CONCLUSIONS ................................ ................................ ................................ ... 124 Clinical Significance ................................ ................................ .............................. 124 Summary ................................ ................................ ................................ .............. 125 Future Directions ................................ ................................ ................................ .. 127 LIST OF REFERENCES ................................ ................................ ............................. 129 BIOGRAPHICAL SKETCH ................................ ................................ .......................... 144
10 LIST OF TABLES Table page 2 1 List of supplementary components included in EC culture medium. ................... 41 6 1 Bioreactor conditions during flow preconditioning. ................................ ............ 121
11 LIST OF FIGURES Figure page 1 1 Human umbilical cord structure ................................ ................................ .......... 32 1 2 EC mechanotransduction ................................ ................................ ................... 33 2 1 Automated dissection of the HUV ................................ ................................ ....... 42 2 2 Image analysis for co stained EC on the lumenal HUV surface ......................... 43 3 1 Decellularization of HUV scaffolds ................................ ................................ ...... 54 3 2 Morphological characterization of decellularized HUV scaffolds ......................... 55 3 3 Surface characterization of decellularized HUV scaffolds ................................ .. 56 3 4 EC adhesion to basement membranes of decellularized HUV scaffolds ............ 57 3 5 EC viability on decellularized HUV scaffolds ................................ ...................... 58 3 6 Lumenal surface morphology of re endothelialized HUV scaffolds ..................... 59 4 1 Flow chamber design ................................ ................................ ......................... 7 2 4 2 Flow chamber assembly ................................ ................................ ..................... 73 4 3 Perfus ion systems for imaging initial adhesion events or extended culture under flow ................................ ................................ ................................ ........... 74 4 4 Shear pre conditioning strategies for EC seeded HUV scaffolds ....................... 7 5 4 5 Mat uration of endothelialized HUV scaffolds ................................ ...................... 7 6 4 6 Morphological comparison of endothelia cultured under static culture or flow conditions ................................ ................................ ................................ ........... 7 7 4 7 Time lapse capture of neutrophil adhesion to endothelialized HUV scaffolds .... 7 8 5 1 Parallel plate culture system ................................ ................................ ............... 93 5 2 Physiologically modeled perfusion culture ................................ .......................... 94 5 3 Cytoskeletal morphology of flow conditioned EC ................................ ................ 9 5 5 4 Cardio protective gene expression in flow conditioned EC ................................ 9 6 5 5 Coagulation & fibrinolysis associated gen e expression in flow conditioned EC 9 7
12 5 6 Inflammation associated gene expression in flow conditioned EC ..................... 9 8 5 7 eNOS function in flow conditioned EC ................................ ................................ 9 8 5 8 NO profiling of EC exposed to additional flow conditions ................................ 100 5 9 HL 60 cell adhesion to flow conditioned EC ................................ ..................... 101 6 1 Bioreactor system for development of TE VG ................................ ................... 116 6 2 Rotating seeding apparatus design ................................ ................................ .. 117 6 3 Comparison of EC attachment after rotational or static seeding approaches ... 118 6 4 EC growth characteristics in normal or dextran supplemented culture media .. 119 6 5 Effects of maturation time on EC retent ion after exposure to flow ramping ...... 120 6 6 EC monolayer development under 3D perfusion culture ................................ .. 122 6 7 HUV scaffold morphology b efore and after endothelialization .......................... 123
13 LIST OF ABBREVIATION S CABG coronary artery bypass grafting DAPI 4',6 diamidino 2 phenylindole EA ethanol/acetone EC endothelial cell (s) ECM extracellular matrix ePTFE expanded polyte trafluoroethylene GFP green fluorescent protein HUV human umbilical vein (s) HUVEC human umbilical vein endothelial cell(s) NaCl sodium chloride NBF neutral buffered formalin NO nitric oxide PBS phosphate buffered saline PCI percutaneous coronary interventi on PET polyethylene terephthalate PF physiological flow PPFC parallel plate flow chamber(s) RP rhodamine phalloidin SDS sodium dodecyl sulfate SF steady flow SMC smooth muscle cell (s) SS shear stress TEVG tissue engineered vascular graft (s) TX Triton X 100
14 Abstract of Dissertation Presented to the Graduate School of the University of Florida in Partial Fulfillment of the Requirements for the Degree of Doctor of Philosophy IN VITRO ENDOTHELIALIZATION AND SHEAR CONDITIONING OF DECELLULARIZED VASCULAR ALLOGRAFTS By Joseph S. Uzarski May 2013 Chair: Peter S. McFetridge Major: Biomedical Engineering There is currently a severe shortage of functional conduit options available for r autologous grafts, one third of patients in need of this procedure lack suitable vessels for harvesting and the prosthetic grafts used instead have inferior patency rates due to thrombosis and intimal hyperplasia Ex vivo derived blood vessels, such as the human umbilical vein (HUV), have shown promise in vascular bypass applications due to their natural composition and appropriate mechanical properties. In native vessels, e ndothelial cells (EC) facilitate thrombus free blood circulation by shielding pro teins of the underlying extracellular matrix and through generation of paracrine factors (such as nitric oxide) that inhibit cell migration, adhesion, a ggregation and proliferation Development of an inner EC mono layer or endothelialization of tissue engi neered vascular grafts (TEVG) could maintain patency through inhibition of peripheral cell adhesion and excessive smooth muscle cell migration/ proliferation that are characteristic features of thrombosis and intimal hyperpla s ia respectively To date, e ndo thelialization of vascular grafts has had limited clinical success due to sub confluence of lumenally seeded EC poor resistance to arterial pressure/shear stress
15 (SS) and unfavorable phenotypic expression by ex vivo expanded EC. Given the importance of h emodynamic SS stimulation in the modulation of EC behavior we tested the hypothesis that details of the fluid mechanical environment are critical to ensure EC adopt a phenotype that is favorable for implantation of vascular grafts. This dissertation deta ils the progressive in vitro development of neo endothelia on decellularized human umbilical vein ( HUV ) allografts for the purpose of creating biologically compat i ble conduit s for vascular bypass applications First, the effects of several chemical decellu larization strategies on EC binding, viability, and repopulation of the HUV were comparatively assessed After selection of an appropriate processing strategy, rotational seeding was employed to improve the circumferential distribution of EC across the lum enal HUV surface. Custom designed parallel plate flow chambers were used to optimize the temporal adaptation of EC seeded on the HUV surface to arterial levels of SS. Finally, a physiological flow model was developed to emulate hemodynamic variations in vi vo and stimulation o f EC monolayers with this flow regime resulted in a more non interactive, quiescent phenotype characterized by increased nitric oxide synthesis. Taken collectively, r esults highlight the physiological importance of mechanical cues from the native arterial microenvironment on EC functionality Applying these cues to endothelialized vascular grafts within perfusion bioreactor systems could improve patency rates by eliciting more favorable interactions with host cells upon implantation
16 C HAPTER 1 INTRODUCTION Clinical Demand f or V ascular Grafts Cardiovascular disease is the leading cause of death in the United States, accounting for 1 out of 3 deaths in 2008 alone. 1 The annual n umber of deaths attri buted to cardiovascular disease outweighs the number of deaths caused by cancer, chronic lower respiratory disease, and accidents combined 1 In 2008, cardiovascular disease and stroke treatments cost the United Stat es an estimated $179 billion in direct expenditures as well as $118.5 billion in lost productivity related to mortality from these conditions. 1 Improving the outcomes of cardiovascular procedures will reduce the co s ts related to these treatments and improve the quality of life for patients suffering from cardiovascular disease. Atherosclerosis, a pathological thickening of the arterial wall due to accumulation of macrophages and lipids in the vascular intima is a ma jor disease of the arteries. 2 As atherosclerotic plaques form, they progressively narrow the arterial lumen, restricting blood flow and cutting off downstream tissu es from their nutrient supply. Rupture of these pl aques can result in rapid thrombotic occlusion of arteries causing ischemia, tissue death, and stroke Stenosis of the coronary arteries restricts blood flow to cardiac tissue and can eventually lead to ang ina and myocardial infarction. Coronary r evascula rization, or the restoration of blood flow using surgical intervention, can be accomplished mainly through two types of procedures: vascular bypass and percutaneous coronary intervention (PCI). 3 Vascular bypass surg ery is performed to redirect blood flow around a narrowed or occluded artery using a conduit vessel restoring blood delivery to
17 downstr eam tissues. PCI also known as angioplasty, is a procedure in which a balloon catheter is threaded through a narrowed a rtery, expanded to break up plaques, and a stent is left in place maintain patency of the narrowed segment PCI has been associated with higher morbidity than coronary artery bypass grafting (CABG), as demonstrated by higher rates of subsequent revasculari zation. 3 CABG which is usually indicated in more severely narrowed arteries, involves redirecting the blood flow around a stenosed area using a separate conduit. 4 While this procedure is more costly and invasive than PCI, it has been associated with lower incidence of major adverse cardiovascular or cerebrovascular events (e.g. stroke) 4,5 Vascular bypass surgery is also performed in patients suffering from peripheral artery disease, in which lack of blood flow can result in lower limb ischemia, claudication, and eventual amputation. 6,7 Patients suffering from cardiovascular disease often develop stenoses in multiple vessels, and preference is often given to tissues in critical need (e.g. CABG to treat ischemic cardiac tissue). This therefore results in increased implantation of prosthetic bypass grafts with inferior patency rates ( i.e. greater risk of occlusion) for peripheral artery bypass, which subsequently require repeat revascularizations. C urrent Vascular Graft Choices and Associated Limitations Autografts The gold standard conduit choice for vascular bypass surgery is an autologous vessel, such as the human saphenous vein or internal thoracic artery. 8 Segments from a suitable vein or artery are harvested and then surgically grafted to redirect blood flow around an occlusion. In patients with available vessels, autologous segments can be
18 harvested and immediately implanted in vascular byp ass surgery to restore blood flow to ischemic tissues. Autografting has several notable disadvantages. First, a significant number of patients in need of bypass operations lack suitable vessels for autografting, due either to previous harvesting or diseas e of the target conduit. 8 In the case of CABG many patients develop atherosclerotic stenoses in multiple coronary arteries and multiple grafts are therefore required I n 2009 for example, 2 42 ,000 pat ients underwent a total of 4 16 ,000 coronary artery bypass procedures in the United States alone. 9 An obvious disadvantage associated with autografting is the fact that a second surgery is required to harvest the con duit vessel, causing add itional morbidity for the patient. Despite these limitations autologous arterial or venous segments remain the preferred graft choices for vascular bypass surgeries. 10 The lack of viable autologous vessels in a significant number of patients in need of vascular bypass surgery has inspired the search for alternative conduit materials. 11 Prosthetic g rafts Prosthetic grafts composed of synthetic polymers such as polyethylene terephthalate (PET) or expanded polytetrafluoroethylene (ePTFE) have been used successfully in large diameter bypass surgeries, but have had comparatively disappointing results in small diameter (<6 mm) applications, such as below the knee peripheral or coronary artery bypa sses. 12 Because autologous grafts are the preferred choice for CABG, surgeons are often forc ed to use allografts or prosthetics for peripheral artery bypass with less than ideal results 6 The main complications associated with prosthetic grafts in small diameter bypass surgeries are thrombotic occlusion and anastomotic intimal hyperplasia. 7 With regards
19 to the former, the low flow and high resistance environment of small diam eter arteries facilitates adsorption of plasma proteins from the circulation onto the graft surface which result s in platelet adhesion, activation and aggregation. 13,14 Thrombotic occlusion is the most prevalent me chanism of early graft failure (<1 month) and occurs in autografts (though less commonly) as well as synthetic conduits. 13,15 The compliance mismatch between rigid polymeric conduits and the more elastic adjacent h ost vasculature leads to disturbed blood flow patterns through the graft and adjacent vessels; this often results in anastomotic intimal hyperplasia. 7,16 While various vein cuff or patch designs have been employ ed t o overcome this problem they require separate harvesting of an autologous vein segment and skilled surgical technique. 17 19 Bacterial infection is a secondary complication associated with prosthetic vascular grafts 20 which are less infection resistant than autografts or allografts. 21,22 This may be due to the inability of pro sthetic materials, which are not cleavable by endogenous matrix metalloproteases, to be infiltrated by inflammatory/immune cells. A significant limitation in grafts composed of PET and ePTFE is that neither material is bioresorbable, and they are th erefore unable to be remodeled in order to adapt to changes in the vascular environment over time. The inability of synthetic cardiovascular defects The disappointing results of many synth etic graft materials in bypass surgeries have led to a paradigm shift away from inert biomaterials toward biologically interactive vascular grafts. 12 14,23 Allografts Due to the obvious shortcomings of synthetic mat erials in the vascular environment, vascular grafts derived from human cadavers have been used in bypass
20 surgery. 24 Owing to their increased resistance to bacterial infection when compared to prosthetic grafts, vascular allografts are primarily used in cases of arterial graft infection. 2 5 Preservation is necessary to prevent deterioration over lengthy storage periods, and superior performance in cryo preserved allografts has been reported when compared to fresh allografts. 24,25 The human umbilical vein (HUV) allograft has been used extensively in vascular bypass surgeries for over 30 years 26 28 In most cases, the graft is chemically fixed using glutaraldehyde before implantation, which masks most antigens a nd lowers the overall immunogenicity of the scaffold. 26 While glutaraldehyde tanning allows the HUV to be tailored in a site specific manner ( e.g. by fitting the vessel to an appropriate diameter before fixation), c rosslinked grafts cannot be infiltrated by host cells or remodeled. 29 In some cases this can result in fibrotic encapsulation and /or graft weakening over time, leading to aneurysm formation 27 Ex Vivo Derived Biomaterials and Vascular Tissue Engineering Natural biological scaffolds, whether autologous or donor derived, have several features that are favorable for vascular reconstruction. 30 The obvious advantages of e x vivo derived tissues are the inherent structural, mechanical, and compositional properties that natural scaffolds possess. Most importantly, these grafts are composed of proteins, glycosaminoglycans, and other components of the ECM necessary to support c ell adhesion, migration, and proliferation. 31 The basement membrane upon which EC reside in vivo for example, is a complex arrangement of proteins and proteoglycans that are important modulators of EC phenotype. 32,33 A n advantageous feature of biological scaffolds compared to synthetic polymers is their capacity to be infiltrated and remodeled by host cells through innate catabolic and
21 anabolic processes. For example cell der ived matrix metalloproteases can cleave particular bonds on strands of collagen, elastin, laminin, and other proteins of the native ECM allowing cells to penetrate into the scaffold. Vascular allografts can therefore be repopulated by autologous cells and appropriately adapted to the physiological demands of the local environment in a site specific manner. Decellularization of Ex Vivo Derived Scaffolds Despite the advantageous structural features, allogeneic /xenogeneic vascular grafts cannot be directly im planted due to immune response generated by the host. The cellular components in blood vessels of foreign origin contain antigens that are recognized and targeted for destruction through innate immunity. 34 D ecellularization of donor blood vessels has been explored as a means to produce an acellular, non immunogenic biological scaffold that can be repopulated ( in vivo or in vi tro ) by autologous vascular cells 35,36 Current in vitro analyses in tissue decellularization, however, focus mainly on the efficacy of various treatment techniques, including quantification of residual DNA 37,38 and extraction of phospholipids 39 proteins 39 and other immunogenic components. Analyses of scaffold deterioration ( e.g. by measuring free hydroxyproline 38 40 ) and alterations in mechanical properties 40,41 have been performed to assess tissue integrity after processing. An equally important yet often overlooked consideration is ho w tissue processing affects adhesiveness for cells in both the peripheral circulation and adjacent vasculature. C areful tissue specific analysis is necessary to choose an appropriate decellularization method ; t his is addressed in further detail in Chapter 3.
22 The HUV as a Scaffold for Vascular Tissue Engineering The human umbilical cord is a n abundant tissue source in labor and delivery wards that (due to regular availability) does not require preservation. Structurally, the umbilical cord is composed of two a rteries wrapped around a central vein in a helical fashion (Figure 1 1) 42 The glycosaminoglycan umbili cal vessels, protects them by provid ing the umbilical cord with resistance to compressive forces The HUV is a unique vessel compared to most veins in that it deliv ers fully oxygenated blood from the placenta to the fetal circu lation, reaching a n average diameter o f 5.40.6 mm at full term. 43 Th e HUV has been explored as a tissue engineering scaffold not only for vascular reconstruction 40,44,45 but also peripheral nerve regeneration 46 periodontal tissue reconstruction 47 musculoskeletal soft tissue regeneration 48 and vocal fold restoration. 49 These previous studies have demonstrated that decellularized HUV scaffolds maintain favorable interactive properties such as permit ting cell adhesion, migration, and proliferatio n within the ECM Anatomy of Blood Vessels B lood vessels ( with the exception of capillaries ) are composed of three layers: the tunica intima, tunica me dia, and tunica adventitia. The innermost layer, the tunic a intima is composed of a single layer of EC attached to a very thin ( ~100 nm) 50 tightly woven network of type IV collagen, laminin, and other structural proteins/polysaccharides known as the basement membrane 51 Lying in between the tunica intima and the tunica adventitia is the tunica me dia, which is composed of several layers of circumferentially oriented smooth muscle cells (SMC) dispersed in a matrix of collagen and elastin. The tunica media is considerably thicker in arteries, which are
23 exposed to higher press ures than veins (Figure 1 1) The media l layer provides blood vessels with strength and contractility, which are nec essary for resistance to arterial stres ses as well as modulation of local blood flow rates The outermost layer, the tunica adventitia, has the lowest cell density of all three layers, and is primarily composed of fibroblasts and connective tissue which integrates the blood vessel with the surr ounding extracellular matrix (ECM) Physiological Roles of EC Situated at the interface between the circulating blood and the surrounding vascular wall/interstitium, t he vascular endothelium is a continuous selectively permeable mono layer of cells that li nes the entire circulatory system from the inner chambers of the heart down to individual capillaries. EC are involved in a variety of physiological processes including angiogenesis 32 hemostasis 52 inflammation 53 regulation of blood pressure/vascular tone 54,55 and wound healing 56 Although the vascular endothelium is often collectively referred to as a distinct endocrine organ 55 EC are phenotypically heterogenous throughout the vasculature 57 with diversity in expression patterns that correlate with tissue specific functions. EC Mechanotransduction of Shear Stress The un ique microenvironment in which EC reside in vivo exposes them to several types of mechanical stresses (see Figure 1 2). On their apical surface, EC are directly exposed to hemodynamic shear stress ( SS ) the frictional force caused by blood flow. SS has bee n shown to directly activate a variety of membrane bound protein s including ion channels 58 G protein coupled receptors 59 tyrosine kinase receptors 60 and other molecules, leading to intracellular signaling cascades. Additionally, the changes in arterial pressure caused by the pulsatile flow of blood induce cyclic circum ferential
24 strain of the underlying ECM. This strain is sensed both by integrins 32,60,61 (which anchor EC to the underlying basement membrane) and cell adhesion molecules such as VE cadherin 62 and PECAM 1 63 which connect adjacent cell s. This variety of mechanosensors in distant spatial locations through the cell membrane activate various cell signaling cascades (JNK 61 ERK 64 MAPK 65 ) that collectively induce an intracellular response to mechanical stimulation Changes in gene expression are then mediated by selective activation of transcription factors ( for example, NF kB is a pro inflammation transcription factor 66,67 whereas KLF 2 is an athero protective transcription factor 68 70 ), although changes in enzymatic activity that are independent of transcription have also been shown to occur in response to application of SS. 71,72 Because of the large (and growing) number of known endothelial mechanosensors 54,73 the response of EC to mechanical forces is a comple x interplay of various signaling pathways. Th us, rather than activation through a single cascade the physiological response of EC to applied mechanical stimulation is a global coordination of multiple signaling mechanisms. Roles in Inflammation, Hemostasi s, and Wound H ealing EC normally express a non interactive, quiescent phenotype that facilitates thrombus free laminar blood flow. This phenotype is characterized by the constitutive production of nitric oxide (NO) 74 an important signaling molecule that inhibits adhesion of platelets 75 and leukocytes 76 and migration 77 / proliferation 78 of SMC. Through the secretion of NO and other inhibi tory paracrine signaling molecules, EC prevent the spontaneous formation of thrombi, as well as the excessive migration/proliferation of SMC within the vascular intima. Both thrombosis and intimal hyperplasia are pathological mechanisms of vascular graft f ailure attributed to inappropriate function of EC and the loss of quiescence.
25 Given their unique anatomical location as the primary barrier to circulating blood EC play critical roles in the early response to injury. When trauma occurs, EC rapidly exocyto se the contents of specialized storage ves icles referred to as Weibel Palade bodies that contribute to blood coagulation and thrombosis in order to stop blood loss. 79 Included in these storage organelles are von Willebrand factor, which serves as a exposed subendothelial ECM proteins 80,8 1 and P selectin, a cell adhesion molecule that allows recruitment of leukocytes from the circulating blood. 79 Act ivated EC then upregulate expression of cell adhesion molecules and inflammatory cytokines to allow circulating neutrophils and monocytes to hone in, adhere, and migrate through the vascular wall to sites of injury. 5 3,82 Through these mechanisms, EC prevent excessive blood loss and facilitate the recruitment of neutrophils and monocytes to destroy infiltrating pathogens. E C also promote tissue repair through angiogenesis, the formation of new blood vessels, in order to improve nutrient delivery to sites of injury 32 Roles in Vascular Tone R egulation The tunica media is mostly composed of SMC, which are responsible for blood vessel contraction/relaxation as well as vascular remo deling. 83 EC contribute to the regulat ion of vascular tone through paracrine interaction with SMC which allows for the direction of both local blood flow as well as the control of systemic blood pressure In response to stimulation (e.g. SS 71,84,85 hormonal activation 86 ), EC secrete signaling molecules, such a s NO 87 that induce relaxation of SMC. 87 This occurs through the binding of NO to soluble guanylate cyclas e which activates the enzyme resulting in the conversion of GTP (guanosine triphosphate) to cyclic GMP 88 In SMC, cGMP causes relaxation through a variety of protein targets, including protein kinases, ion channels,
26 and cyclic nucleotide phosphodiesterases that have been reviewed elsewhere. 89 Other signaling molecules produced by EC induce vasoconstricti on, which reduces local b lood flow 90 Whether EC induce vasodilation or vasoconstriction of blood vessels is dependent on local environmental factors such as downstream metaboli c demand. Endothelialization of Vascular Grafts It has been hypothesized that the superior ity of small diameter vein graft performance over prosthetic materials may be due to the presence of an inner EC layer. 91 Due to the naturally non thrombogenic properties of EC they have been targeted for lining the blood meter vas cular grafts. 92 96 The spontaneous 97 or directed 94 development of an EC layer on implanted prosthetic grafts is also sometimes refe However, while this often occurs in animal models, it rarely happens in humans, and therefore for this dissertation hereafter re fer to the development of an EC layer in vitro 97,98 An initial challenge for in vitro endothelial i zation of both prosthetic grafts as well as TEVG is efficiently seeding an EC monolayer on a three dimensional condu it Various methods to accomplish this have included electrostatic seeding 96 precoating with fibrin glue 93 or ECM proteins 99,100 biofunctionalization with pepti de sequences 94,101 and rotational seeding 93,102 R otational EC seeding of the HUV graft, which has previously improved endothelialization of other graft types 93,102 is addressed in Chapter 6 One stage Seeding One stage seeding refers to the seeding of EC at the time of or on a time frame within the scope of surgical graft implantation. 103 This technique was introduced by He rring et al. in 1978, in which EC were mixed with whole blood during pre clotting of
27 Dacron grafts before implantation in canines. 104,105 EC (or endothelial progenitor cells) we re obtained from a piece of explanted blood vessel or other tissue source, isolated, and then seeded onto the graft surface using fibrin glue, immobilized RGD peptides, or other adhesive medium. While this technique produced promising results in animal models 105 clinical trials were disappointing 92 This was partially attributed to sub confluent EC densities at the time of implantation, resulting in thrombus formation due to exposure of the graft surface S eeding using highly concentra ted EC suspensions or to improve surface coverage 103 EC sodding is not clinically practical for one stage seeding however, because a limited number of cells c an be obtained from explanted vein segments 103 Rap idly seeded EC, which have little time to adhere and mature on the graft surface before exposure to blood flow, also have poor resistance to SS resulting in denudation of the graft surface after implantation Two stage Seeding Two stage seeding refers to the ex vivo expansion of EC/endothelial progenitor cells prior to seeding and/or the allotment of additional culture time to allow maturation of endothelialized grafts before implantation. 95 Clinically, two stage e ndothelialization has shown some improvement for patency of prosthetic grafts. E ndothelialized ePTFE grafts had higher patency rates as compared to acellular grafts in infrainguinal 93 and coronary artery 100 bypass applications. The most significant disadvantage of t he two stage seeding method is the time delay required for in vitro cell expansion, as the limited number of EC that can be explanted is insufficient to achieve a confluent lining of the lumenal graft surface. This limitation has restrict ed the usage of su ch grafts to elective cases. 95,103
28 Strategies to Improve EC Retention The poor results of early clinical trials involving endothelialized prostheses were attributed to low EC retention on the lumenal surface. Strate gies have more recently been developed that aim to strengthen EC attachment Miyata et al showed that EC retention on fibronectin coated ePTFE grafts after exposure to 90 minutes of S S was dependent on both culture time as well as initial seeding density. 106 Additional cult ivation time improved resistance of seeded cells to SS and maximal retention occurred when flow was initiated shortly after confluence had been reached. Flow preconditioning is a strategy to improve the retention of EC on the graft surface through the in vitro application of fluid SS using vascular perfusion systems Yazdani et al demonstrated that endothelial progenitor cell seeded decellularized porcine carotid arteries that were preconditioned to pulsatile high SS ( 13.2 dynes/cm 2 ) maintained lumenal co nfluence after exposure (15 minutes) to blood flow in an arteriovenous shunt model 107 Improved endothelial progenitor cell retention correlated with decreased adhesion of platelets and white blood cells compared to grafts preconditioned under low SS (1.7 dynes/cm 2 ) The use of perfusion systems to precondition EC seeded vascular grafts to fluid forces is a promising strategy to improve clinical results for endothelialized prosthetic or tissue engineered vascular gra fts (TEVG) and is explored in Chapter 6 History of EC Culture and Improvement of Perfusion Systems In vitro studies on EC biology began with characterization of explanted cells on tissue culture plastics which has since become routine for culture of adh erent cells 108 Progressive adv ancements in EC culture techniques, including the use of culture media additives such as EC growth factor 109 and heparin 110 as essential co factors, have aided
29 the ex vivo e xpansion of primary human EC isolates Extended culture of EC in vitro has allowed detailed characterization of these cells as well investigation of the factors that control their behavior the predominant culture technique mainly due to low cost and simplicity. Parallel Plate Flow Chambers (PPFC) Early interest in studying endothelial mechanotransduction (the translation of mechanical stimuli into biological signals) in vitro in the 1 sparked the now widespread use of SS generating devices such as PPFC and cone and plate viscometers. 71,111 113 PPFC have been used extensively to study EC mechanotransduction, leukocyte rolling, platelet adhes ion, and other cellular interactions under fluid flow for over thirty years. 71,114,115 The advantages of this device are its low reagent requirements, uniform fluid mechanics, wide range of flow rates, and ability t o permit real time microscopic observations. 116 Flow chambers are often used investigate the effects of applied SS on EC in the absence of other confounding physiological variables. However, the PPFC has limited ability to truly emulate the natural environment of the vascular intima in vitro The composition of the basement membrane changes with the physiological state of the vessel 117 and ligation of EC integrins by specific pept ide sequences of the basement membrane has previously been shown to permissively affect the signaling pathways activated by fluid SS. 114,118 120 Furthermore, the micromechanical properties of the underlying culture substrate (e.g. rigidity or surface roughness) affect the way EC attach, spread, and adapt to fluid forces, all of which can influence cytoskeletal tensegrity and the way these cells transduce fluid SS 121,122 An
30 im proved PPFC design that incorporates a natural tissue scaffold as a substrate for EC attachment is introduced in Chapter 4. Application of SS Using Perfusion Systems EC perfusion systems have clarified the importance of hemodynamic SS in the regulation of EC behavior. Pioneering work by Dewey et al. highlighted the distinctions in cytoskeletal alignment, endocytosis, thrombogenicity, and wound repair in EC exposed to uniform, continuous shear compared to static culture 112 Frangos et al. further demonstrated the physiological relevance of flow pulsation in these model system s in reporting that prostacyclin production in EC cultured under pulsatile flow was twice as much as cells cultured under continuous flow, and sixteen times that of static cultured EC. 71 It has become increasingly clear that recapitulating the complex characteristics of physiological blood flow is critical to the utility of in vitro SS models. Endothelial mechanotransduction has traditionally been studied in vitro by subjecting monol ayers to SS of a defined magnitude, pulsatility, and frequency for a time period ranging from minutes to hours. 71 More recently, investigators have focused on how particular charact eristics of applied SS such as the effects of flow pulsatility 123 pulse frequency 124 and rate of change of SS 125,126 influence EC function. However, a limitation in many of these studies is that the cyclic SS p rofile to which the cultured EC are exposed is unchanging over time; this is in direct contrast to the ever changing hemodynamic patterns in vivo Chapter 5 addresses the influence of hemodynamic variation s on EC phenotype an understanding of which is of critical importance for the in vitro development of TEVG
31 Summary and Present Objectives Thrombotic occlusion and intimal hyperplasia are prevalent mechanisms of vascular graft failure related to the absence of a functional EC lining. While much has been l earned about the biology of these cells in recent years, endothelialization of implanted vascular grafts has not yet resulted in repeatable improve ments in clinical outcomes 92,127 Previous results show decellulariz ed HUV allograft s to possess favorable properties for vascular reconstruction though none to date have specifically addressed EC seeding 39,40,44,45 Given the multitude of physiological roles played by EC in vivo it is generally agreed that a functional small diameter TEVG will require an inner EC monolayer. 11,92,128 The refinement of graft processing techniques, seeding methods, and mechanical pre conditioning strategies us ing perfusion systems will likely result in improved patency rates of EC seeded small diameter vascular grafts. This dissertation details the investigation of several properties of the in vitro EC microenvironment that affect adhesion, morphological adapta tion, and appropriate behavior In particular focus are c haracteristics of EC that are of critical importance for healthy vessel function The objective of th is dissertation is to elucidate details of the vascular microenvironment that are relevant to ensu re EC adopt a phenotype that will maintain proper functionality of vascular grafts in vivo
32 Figure 1 1. Human umbilical cord structure. The human umbilical cord is composed of a single vein, which carries oxygenated blood from the placenta toward the fetus, and two arteries which return oxygen deficient blood to the placenta. Masson Trichrome stain shows smooth muscle of the vessel wall in red, collagen in blue, and cell nuclei in black. The vessels are enclosed in a GAG rich, compression resistant mat rix
33 Figure 1 2 EC mechanotransduction EC are subjected to hemodynamic SS at their apical surface and circumferential strain at the basolateral surfaces due to the pulsatile flow of blood. These mechanical stimuli are transduc ed through a variety of protein receptors, including ion channels, growth factor receptors, cell adhesion molecules, and integrins. Collectively, these mechanosensors stimulate intracellular signaling cascades that result in selective activation of transcr iption factors, leading to changes in gene expression, protein synthesis, and enzymatic activity.
34 CHAPTER 2 METHODS AND MATERIAL S In this chapter, experimental and analytical methods common to multiple chapter s are listed. Chapter specific methods will be explained within individual chapter s Experimental Methods Human Tissue C ollection Experiments involved de identified human tissue samples were approved according to the Institutional Review Board 01 at the University of Florida (Gainesville, FL; IRB appr oval #64 2010). Because tissue s amples were indirectly obtained and de identified prior to collection, informed consent was not deemed necessary by the Institutional Review Board. Human umbilical cords were obtained from the Labor & Delivery department at Shands Hospital at the University of Florida. Primary EC I solation and E xpansion For EC extraction, h uman umbilical cords were obtained from Labor & Delivery at Shands Hospital at the University of Florida (Gainesville, FL) and processed within 12 hours of delivery. HUVEC (hereafter referred to as EC) were isolated from cords using collagenase digestion adapted from the process described by Jaffe et al 108 U mbilical cords were cannulated through the vein using a barbed luer fitting attached to a 60 mL syringe. Ve ins were flushed with ph osphate buffered saline ( PBS; pH 7.4) pre warmed to 37 C to remove residual blood. The open end of the vein was then ligated using a cable tie and a 1 mg/mL solution of pre warmed (37 C) type I collagenase ( Life Technologies # 17100 017 ; Grand Island, NY) was injected into the lumen. After a 25 minute incubation period, umbilical cords were gently massaged and the solution was collected in a 50 mL Falcon tube with an equal volume of VascuLife media (see
35 formulation below). C ells were pelleted via centrifuga tion (232 xg RCF) resuspended in fresh media, and plated on a T75 flask Media was replenished every 2 3 days with VascuLife basal medium supplemented with VEGF LifeFactors kit (LifeLine Cell Technologies #LL 0003 ; Frederick, MD) and 100 U/mL penicillin/st reptomycin ( Thermo Scientific # SV30010 ; Waltham, MA ). Table 2 1 shows a complete list of supplements included in this formulation. EC were expanded in an incubator kept at 37C and 5% CO 2 passaged every 2 3 days using Accutase with 0.5 mM EDTA (Innovativ e Cell Technologies # 21 201 0100V ; San Diego, CA), and split at a 1:4 surface area ratio. Unless otherwise specified, EC were used experimentally between passages 2 and 5 HL 60 Cell C ulture HL 60 promyelocytic leukemia cells transduced with a green fluore scent protein (GFP) expressing lentiviral vector were generously provided by Dr. Christopher Cogle (University of Florida Department of Medicine; Gainesville, FL). They were maintained at concentrations ranging from 5x10 5 2x10 6 cells/mL in Hyclone Dulbecco Eagle Medium (Thermo Scientific # SH30021FS ; Waltham, MA) supplemented with 20% FBS (Gemini Bio Products #100 106 ; West Sacramento, CA) HUV D issection Decellularization, and S terilization HUV were isolated from the surrounding arteries and Wha using an automated dissection procedure that has been described previously. 40 U mbilical cords were rinsed clean and cut into 1 0 cm lengths threaded though the vein and the tissue was progressively frozen down to 80C at a controlled rate of 1 C/minute using a Styrofoam insulating container After at least 24 hours, frozen umbilical cord sections were removed from the freezer and quickly
36 mounted between the headstock and tailstock of a CNC lathe (MicroKinetics #999 6300 XXX; Kennesaw, GA). After setting the spindle to a rotational speed of 2000 RPM, sections were automatically machined to a uniform wall thickness of 750 m using TurnMaster Pro software (Figure 2 1). The entire process required less than 2 minutes, from re moval of frozen umbilical cords from the 80 C freezer to transfer of lathed HUV sections to the 20 C freezer. Veins were progressively thawed at 20C for 2 hours then at 4C for 2 hours Veins were then decellularized by immersion in 1% (w/v) sodium dod ecyl sulfate (SDS) in DI water under orbital shaking (100 RPM) for 24 hours at a 1:20 tissue mass to volume ratio. SDS was used as the decellularizing reagent for all HUV scaffolds unless specified otherwise. Decellularized HUV were then sequentially rinse d in rinsed in fresh DI water solutions under orbital shaking for 5 minutes, 15 minutes, 40 minutes, 1 hour, 3 hours, 12 hours, and 24 hours. HUV sections were then incubated in 70 U/mL of deoxyribonuclease I (Sigma Aldrich #DN25 ; St. Louis, MO) in PBS agi tated on an orbi tal shaker for 2 hours at 37C. HUV were rinsed in DI water (2 times ) for 5 minutes and terminally sterilized in a solution of 0.2% (v/v) peracetic acid and 4% (v/v) ethanol in DI water on an orbital shaker for 2 hours. HUV were sequentiall y rinsed for 5 minutes, 15 minutes, 40 minutes, and 1 hour in DI water, and balanced in PBS (pH 7.40) for 24 hours Scaffolds were stored in PBS at 4C for a maximum of 2 weeks until use. Computer aided Design Computer aided design (CAD) models of flow cha mbers and bioreactor components were created using Google SketchUp software version 8.0.15158.
37 Analytical Methods Scanning Electron M icroscopy HUV samples were placed in the well bottoms of 48 well plates for tissue processing. A 45 second microwave cycle followed by a one minute bench top incubation w ere used for the following s olutions Samples were fixed in 2.5% glutaraldehyde, washed three times in PBS, fixed in 1% osmium tetroxide solution, washed twice in PBS, washed once in distilled water, and progr essively dehydrated in 2 5%, 50%, 75%. 85%, 95%, and 100% (three times) ethanol solutions. Samples were then critical point dried, sputter coated with gold/palladium, and imaged using a Hitachi S 4000 FE SEM (10.0 kV). Fluorescent S tains Diamidino 2 p h enylindole (DAPI) After 10 minute s fixation in 10% neutral buffered formalin (NBF), cells were washed in PBS three times for 5 minutes then stained using 300 nM DAPI dihydrochloride (Molecular Probes #D1306; Eugene, OR) in PBS to visualize cell nuclei Ce lls were washed in PBS three times for 5 minutes and imaged. Rhodamine p halloidin (RP) Cells were fixed for 10 minutes in methanol free 10% NBF After two 5 minute washes, cells were permeabilized using 0.1% (w/v) Triton X 100 in PBS. Cells were incubated in a blocking solution of 1% (w/v) bovine serum albumin (BSA) in PBS then stained using 5 U/mL RP (Molecular Probes #R415; Eugene, OR ) in the presence of 1% BSA to visualize filamentous actin, or F actin Cells were washed in PBS three times for 5 minutes and imaged.
38 Live/d ead s tain EC mammalian cells ( Molecular Probes #L 3224; Eugene, OR ) Cells were incubated for 30 minutes with 2 m calcein AM (which fluoresces when enzymatically modified by intracellular esterases active only in live cells ) and 2 m ethidium homodimer 1 (an intercalating dye that stains the nuclei in dead or dying cells) in standard EC media and imaged through both GFP and DsRed filters as described below. Fluorescent I maging and Analysis Fluorescentl y stained cells were imaged using a Zeiss AxioImager M2 upright fluorescence microscope coupled with an AxioCam HRm Rev. 3 monochromatic digital camera operated by AxioVision software version 4.8. Multidimensional acquisition was employed to ima ge cells at multiple wavelength ranges which could then be exported as separate channel images for analysis EC seeded HUV scaffolds co stained with RP/ DAPI as described above were analyzed using ImageJ software (Figure 2 2). Channel i mages obtained t hrough DsRed/ DAPI filters were separately exported for analysis using NIH ImageJ software DAPI images were analyzed using the ITCN Automatic Nuclei Counter plugin ver. 1.6 to quantify the num ber of cell nuclei in each image, which was divided by the image surface area to calculate cell density (Figure 2 2, B) Thresholding analysis was using to quantify the percentage of each DsRed image devoid of F actin, which was subtracted fro m 100 to calculate percent coverage by cells (Figure 2 2, C) The number of s amples imaged and the number of images obtained per sample are described in further detail in Chapters 4 and 6.
39 Histology Tissue processing HUV samples were immersed in OCT medium within plas tic base molds and rapidly snap frozen. A stainless steel beaker filled with isopentane was placed in a larger Styrofoam container with liquid nitrogen. The base mold was then lowered to the top of the chilled isopentane solution using a pair of forceps and held in place until the sample was completely frozen which occ urred in roughly 30 seconds Snap frozen tissue samples were stored at 80 C until cryo sectioning. Frozen HUV scaffolds were cut into 8 m sections using a cryostat, and collected on glass slides for staining. Tissue sections were stored at 8 0 C until staining. Hematoxylin & e osin stain Tissue sections were fixed for 10 minutes using 10% NBF. Sections were rinsed in DI water for 30 second s, stained with hemotoxylin for 1 minute, dunked in tap water and then rinsed in running tap water for 1 minute. Sections were immersed in bluing reagent for 30 seconds, and then rinsed in DI water for 30 seconds. Sections were stained with eosin for 20 s econds, dunked in tap water and rinsed in running tap water for 1 minute. Sections were progressive dehydrated using 70, 80, 90, and 100% EtOH for 30 seconds each before clearing and mounting. richrome stain Masson richrome stain was carried out according to kit instructions (Richard Allen Scientific #87019; Kalamazoo, MI). Tissue sections were fixed for 10 minutes using 10% NBF. Sections were rinsed in running DI water for 1 minute and incubated in C for 1 hour. Sections w ere rinsed in running water for 5 minutes,
40 rinsed in running DI water for 5 minutes and stained in Biebrich Scarlet Acid Fuchsin Solution for 5 minutes. Sections were th en rinsed in running DI water for 30 seconds, placed in Phosphotungstic Phosphomolyb dic Acid Solution for 5 minutes, and then stained in Aniline Blue for 5 minutes. Sections were placed in 1% Aceitc Acid, rinsed in running DI water for 30 seconds then deh ydrated in 100% EtOH twice for 1 minute before clearing and mounting. Color I maging Stained s amples were imaged using a Zeiss Axio Imager.M2 microscope coupled with a Zeiss AxioCam MRc camera operated by AxioVision software version 4.8 Statistical A nalysi s Statistical analyses were performed using SPSS 15.0 for Windows The particular tests performed, as well as the sample sizes, are described in detail within individual chapters.
41 Table 2 1. List of supplementary components included in EC culture medium. Abbreviation Component Name Concentration rh VEGF Vascular endothelial growth factor 5 ng/mL rh EGF Epidermal growth factor 5 ng/mL rh FGF basic Basic fibroblast growth factor 5 ng/mL rh IGF 1 Insulin like growth factor 1 15 ng/mL Ascorbic acid N/A 5 0 g/mL Hydrocortisone hemisuccinate N/A 1.0 g/mL Heparin sulfate N/A 0.75 U/mL L glutamine N/A 10 mM FBS Fetal bovine serum 2% (v/v) Pen strep Penicillin streptomycin 100 U/mL (each)
42 Figure 2 1. Automated dissection of the HUV Human umbilic al cords were mounted on 80 C. Mandrels were then secured at each end on a CNC lathe (A) and axially rotated at 2000 RPM while a computer controlled blade was moved to within a distance of 750 m from the mandrel (B) and moved along the long axis of the mandrel (C) This process resulted in a dissected HUV scaffold of uniform thickness (D).
43 Figure 2 2. Image a nalysis for co stained EC on the lumenal HUV s urface. EC seeded HUV scaffolds were fixed and co staine d with RP /DAPI to visualize cytoskeletal F actin (red) and nuclei (blue), respectively (A). Images were separately obtained through DsRed and DAPI filters using a monochrome camera and analyzed using ImageJ software. Each DAPI image was analyzed using the ITCN plug in to quantify the number of cell nuclei in each image (B). Thresholding analysis was using to quantify the percentage of each DsRed image devoid of F actin (C), which was subtracted from 100 to obtain percent coverage.
44 CHAPTER 3 DECELLULARIZATI ON OF THE HUMAN UMBI LICAL VEIN SCAFFOLD: EFFECTS ON ENDOTHELIAL CELL ADHESION AND VIABILI TY Immune mediated rejection is a significant complication associated with the use of vascular allografts. 29,129 Decellulariza tion of allo /xenografts is often performed to abate the immune response generated upon implantation of the scaffold by solubilizing and stripping out antigenic cellular components while preserving the native architecture of the ECM 36,130 Decellularized donor tissues have shown great promise in applications both as acellular, implantable biomaterials as well as scaffolding material for cell seeded tissue engineered constructs. 40,46 49,131,132 However, the chemical 41,99 (e.g. organic solvents, detergents, ionic solutions) and mechanical 39,131 (e.g. agitation, sonication, freeze thaw) methods used to strip immunogenic components out of scaffolds also affect scaffold integrity and therefore the biological responses elicited upon implantation. 36 Because the ECM composition varies greatly among different tissues, there is no universally applied decellula rization protocol. Due in part to their regular availability, the HUV and arteries have previously been characterized for their ability to function as vascular graft scaffolds. 40,99 Decellularization of the HUV has been described previously by a variety of chemical treatments, including salts, ionic detergents, and a mixture of organic solvents. 39,40,44,49,133 However, to date, there are no studies specifically addressing how these treatments affect surface interactions at the host graft interfaces. Adhesive interactions between implanted biomaterials and plasma proteins, circulating/vascular cells, and other components play key roles in the initiation of thrombosis and inflamm ation, and repopulation/remodeling by native vascular cells. 13,134,135 Minimizing deleterious responses to implantable materials is therefore of critical importance to graft survival.
45 This study was undertaken for f unctional assessment of the HUV scaffold after tissue decellularization through various strategies. As mentioned in Chapter 2, the HUV was chosen as a target scaffold due to its increasingly common use as an implantable allogeneic biomaterial in a plethora of applications. 47 49,133,136 Here we characterized the effects of an organic solvent, anionic surfactant, non ionic surfactant, and hypertonic salt solution on the interactive properties of the HUV with vascular E C 137 The objective of this chapter was to comparatively assess surface interactions between processed vascular grafts and EC, which play critical roles in tissue performance or graft failure. Methods and Materials D ecellularization of HUV HUV were isolated from the surrounding tissue as described in Chapter 2, and then decellularized using one of four chemical solutions. The four decellularization solutions are as follows: a solution of ethanol, acetone, and DI water in a 60:20:20 volum etric ratio (EA); a 2 M solution of sodium chloride in DI water (NaCl), 1% (w/v) solution of sodium dodecyl sulfate in DI water (SDS); and a 1% (w/v) solution of Triton X 100 in DI water (TX) Sections were exposed to one of these four solutions under orbi tal shaking (100 RPM) for 24 hours at a 1:20 mass to volume ratio. Further processing steps were performed as described in Chapter 2. EC Seeding A ssay dHUV discs were prepared as described above and placed in the well bottoms of a 48 well plate with the lu menal surface facing up. HUV scaffolds were soaked in VascuLife VEGF culture media with 2% FBS and incubated overnight at 37C EC were lifted, counted, and seeded at a density of 10 5 cells/cm 2 (assuming uniform distribution across the well bottom). After 24 hours, unattached cells were removed by carefully
46 rinsing in PBS three times before subsequent fixation and processing; rinses were pipetted into the wells along the walls at an angle to avoid shear mediated denudation of adherent cells. Cells were stai ned with DAPI as described in Chapter 2 Non seeded control scaffolds were similarly processed for each treatment method; no intact or fragmented nuclei were observed on the lumenal surface, confirming removal of native EC Images were captured as describe d above. Six fields of view were imaged per scaffold (10x magnification), and the number of cells in each image were manually counted using NIH Image J software Attachment (reported as number of cells per surface area) was calculated for five discs per tr eatment. EC Viability A ssay EC monolayers seeded on dHUV scaffolds were assessed for viability using the as described in Chapter 2 EC were seeded onto dHUV scaffolds at 10 5 cells/cm 2 as described above and allowed to a dhere and spread for 24 hours before staining Images were captured and cells were manually counted as described above in the previous section. The percent of viable cells on each scaffold (n=3 per treatment) was calculated using the following formula: %Viabilit y = 100*(#Live cells/surface area)/((#Live cells/surface area)+(#Dead cells/surface area). Statistical A nalysis Results are presented as meanS.E.M. Significance among treatments was determined by one way ANOVA followed by post hoc Tukey Kramer HSD analys is to determine significance between individual means. Asterisks indicate significance at the 0.05 level between individual means (p<0.05).
47 Results HUV I solation and G ross M orphology A fter D ecellularization HUV scaffolds were isolated fr om the surrounding arteries and using an automated lathing procedure 136 resulting in tu bular scaffolds of uniform thickness (Figure 3 1, A B). After lathing, HUV were immersed in ethanol/acetone (EA), sodium chloride (NaCl), sodium dodecyl sulfate (SDS), or Triton X 100 (TX) under orbital shaking for 24 hours. dHUV sections derived from the same donor umbilical cord were imaged before subsequent rinse steps for visual comparison of the scaffolds after each treatment (Figure 3 1, C F). Immersion in organic solvents resulted in dehydration and compression of the HUV scaffold (Figure 3 1, C). In comparison, treatment with NaCl, SDS, or TX (aqueous solutions) caused progressively increasing swelling of the scaffold (Figure 3 1 D F). Alterations to the native ECM structure and composition were highly dependent on the decellularizing reagent used. Both H&E and Masson richrome were used to stain thin sections of dissected or variously decellularized HUV scaffolds (Figure 3 2). Decellularized scaffolds appeared far more swollen and structurally less dense than native tissue (Figure 3 2, A,F). SDS a nd TX dHUV caused greater swelling of the glycosaminoglycan compared to other treatments (Figure 3 2, I J ) Collagen fibers appeared more compressed in EA dHUV, despite the fact that these scaffolds were extensi vely rehydrated after decellularization (Figure 3 2, B,G) Lumenal/ablumenal S urface C haracterization Processed HUV scaffolds were examined using scanning electron microscopy to visualize topographical features, surface structure, and global integrity a fte r
48 decellularization. Figure 3 3 shows representative scanning electron micrographs of both the lumenal and ablumenal surfaces of HUV subjected to each treatment. All four methods achieved complete removal of the native endothelium, leaving behind the tight ly woven basement membrane (BM) structure on the lumenal surface (Figure 3 3, A D). Small areas were visible where the BM was partially denuded, exposed the underlying collagen fibrils of the medial layer (Figure 3 3 insets). Larger (5 10 m) gaps were observed on NaCl dHUV and TX dHUV than on other groups; the BM was left mostly intact in EA dHUV or SDS dHUV. The ablumenal surfaces of dHUV, in comparison, showed a much less dense structure, with increased pore sizes present in between loo sely wove n collagen fibrils (Figure 3 3, E H). Much larger topographical variations were observed on the ablumenal dHUV surface than on the lumenal surface. No consistent global differences in ablumenal microstructure were apparent among the four treatment groups. EC S eeding and C ulture Because an intact basement membrane structure was visualized on the processed HUV, we next sought to characterize the ability of EC to adhere to and repopulate this surface. DAPI nucleic acid staining revealed that EC were able to adhere to the lumenal surface of dHUV with a uniform dispersion (Figure 3 4, A D). When HUV were pre soaked in Vasculife media without FBS, EC were unable to attach ( data not shown ). EC were able to adhere with high efficiency (over 73% of the cell s s eeded in all cases, Figure 3 4, E). EA dHUV exhibited the highest EC adhesion (101,4333,498 cells/cm 2 ), which was significantly higher than SDS dHUV, the treatment which yielded the lowest EC adhesion (73,7464,426 cells/cm 2 p=0.016).
49 EC V iability A ss ay Viability of re endothelialized HUV scaffolds was assessed by costaining with calcein AM and ethidium homodimer 1. A high number of live cells was observed on the l umenal HUV surface (Figure 3 5, A D) with comparatively few d ead or dying cells (Figure 3 5, E H). Very high EC viability was observed regardless of decellularization technique (Figure 3 5, I), with the highest viability on SDS dHUV (99.110.19%). EC seeded on TX dHUV, however, demonstrated significantly lower viability (96.490.38%) than on S DS dHUV (99.110.19%, p=0.001), EA dHUV (98.580.05%, p=0.004), or NaCl dHUV (98.350.39%, p=0.008). Lumenal S urface M orphology of R e endothelialized HUV dHUV discs seeded with EC as described above were also processed for scanning electron microscopy Rep resentative images of re endothelialized dHUV from each treatment group are presented in Figure 3 6 24 hours after seeding, monolayers were near confluence, with few gaps present exposing the underlying basement membrane (Figure 3 6, A D). After allowing monolayers to develop further for 7 days, full confluence was observed, and the borders between adjacent cells were more di fficult to discern (Figure 3 6, E F). Discussion Decellularization, the extraction of immunogenic cellular antigens, is a processing technique that has been explored for allogeneic/xenogeneic organs and tissues as an alternative to glutaraldehyde fixation that allows for ECM remodeling and regeneration. 130 The ideal decellularization process would compl etely remove or nullify antigens within a graft, while preserving the native extracellular structure, composition, and mechanical properties of the tissue. 36,130 A variety of commonly used decellularizing
5 0 reagents w ere chosen for the present study, based on their frequent citation in the literature and their differing methods for removal of cellular components Decellularization of the HUV scaffold with these reagents has been performed previously, and the efficacies of these methods are well characterized. 39,40,44,49 Here, we compared how these treatments affected extracellular HUV structure and function as they specifically relate to EC adhesion with vascular grafts. Native s tructural properties of HUV grafts were affected by these treatments in different ways. Immersion in organic solvents ( EA mixture), for example, caused gross dehydration and compression of the scaffold in the radial direction, whereas detergents (SDS/ TX ) c aused significant swelling of the tissue. Although the variously treated scaffolds normalized in thickness (approximately 1 mm thickness) after subsequent rinses, the tissue deformation caused during chemical processing likely resulted in permanent changes to the ECM microstructure. The l umenal surface of the HUV is covered by a thin basement membrane, constituted of tightly woven laminin, fibronectin, type IV collagen, and proteoglycans. 30,138 Scanning electron micr ographs of the lumen facing surface showed that scaffolds processed using NaCl or TX had large (5 10 m) holes in the basement membrane. HUV decellularized using organic solvents appeared slightly more coarse on the lumenal surface, indicating mild protein degradation. SDS treatment seemed to best preserve the tightly woven structure of the HUV basement m embrane. Because chemical processing distorts structural tissue features, characterization of functionality is an important precursor to implantation. For allogeneic vascular grafts, no studies to date have specifically addressed how these chemical decellu larization
51 strategies affect binding interactions EC Binding events at the graft host interface are critical determining factors in vein graft patency or failure. Thrombosis, intimal hyperplasia, and accelerated atherosclerosis are significant complicatio ns in vein allografts, and all are related to excessive adhesion, migration, and proliferation of native cells. 15,16,139 On the other hand, infiltration by inflammatory cells has been shown to support vascular remod eling in implanted vascular grafts, and tissue ingrowth in aneurysm models. 140,141 A balance therefore exists in producing a scaffold that is resistant to early complications (thrombotic occlusion), yet will still p ermit appropriate infiltration and ECM remodeling by host cells over time. Here, we sought to characterize cell adhesion at both the endovascular and adventitial surfaces of processed HUV grafts. 13,14,134 EC showed comparatively higher adhesion on the lumenal HUV surface than platelets/leukocytes. 137 Organic solvent decellularized HUV permitted the highest retention of EC, which was significantly higher than SDS decellularized HUV. It should be no ted that EC adhesion was only observed when decellularized HUV were pre incubated in medium supplemented with fetal bovine serum (FBS; 2%) overnight. Therefore, it is likely that EC adhesion depended on the selective adsorption of FBS proteins onto the lum enal HUV surface, which varied by treatment method. Despite that the basement membrane appeared to be most intact in SDS dHUV, lowest EC adhesion was observed on this surface. SDS, an anionic detergent, may have induced an overall negative charge to the su rface of the HUV, which would repel adsorption of certain protein moeities. 13,142 This would explain the lower observed efficiency of EC adhesion compared to other treatments. Despite initial differences in cell adh esion,
52 scaffolds were lumenally covered in tight monolayers after 7 days culture, indicating that EC could spread, proliferate, and form tight junctions on the processed grafts. Alteration of the chemical and/or structural properties of the HUV basement me mbrane also had a significant effect on EC viability. Scaffolds decellularized using TX a non ionic detergent, had more holes visible in the basement membrane region, exposing the underlying collagen fibrils of the medial layer. We observed that EC seeded onto Triton X decellularized HUV displayed significantly lower viability when compared to all the other treatment methods. Given the disrupted basement membrane structure and resultant lower viability of EC observed in the present study, TX may not be an optimal decellularization method for the HUV. Conclusions In the present study, we have demonstrated that varying structural alterations induced by decellularizing reagents on ex vivo derived tissue scaffolds result in clinically relevant changes in vascul ar EC adhesion. Strategies examined here to remove native antigens from HUV allografts (organic solvents, osmotic stress, or ionic/non ionic detergents) were effective, and largely preserved the overall integrity of the HUV scaffold. While it is acknowledg ed that the bulk mechanical properties of the HUV scaffold are undoubtedly affected in various ways by these chemical treatments, the results presented here underscore that alterations in surface properties influence initial cell adhesion. The basement mem brane was mostly retained on the lumenal surface with all treatments, though variety in cellular adhesion among treatments indicates that subtle changes in protein structure/surface chemistry existed. Processed HUV scaffolds were capable of being repopulat ed by viable EC which attached with high efficiency and grew to confluence on the basement membrane. These results show much promise
53 for further development of the ex vivo derived tissues as implantable vascular biomaterials.
54 Figure 3 1. Decellulariz ation of HUV scaffolds. Human umbilical cords were mounted 80C, and lathed down to a thickness of 750 microns (B). HUV were then decellularized using one of four chemical treatments: EA (C), NaCl (D), SDS (E), or TX (F). Shown are representative images of HUV scaffolds immediately following decellularization and before subsequent rinsing/sterilization. Figure adapted with permission from the Journal of Biomedical Materials Research Part A.
55 F igure 3 2 Morphological characterization of decellularized HUV scaffolds. HUV were decellularized using one of four chemical treatments: EA ( B,G ), NaCl ( C,H ), SDS ( D,I ), or TX ( E,J ). Shown are representative cross sectional images of dissected (A,F ) or decellularized HUV stained using H&E (top) or Masson t richrome (bottom). All four treatments resulted in removal of resident cells from the HUV scaffold, leaving behind an intact ECM composed of HUV smooth muscle and collage n richrome (MT): black=cell nuclei, red=smooth mus cle /cytoplasm blue=collagen. Scale bar: 1 00 m.
56 Figure 3 3 Surface characterization of decellularized HUV scaffolds. HUV were decellularized using one of four chemical treatments: EA (A,E), NaCl (B ,F), SDS (C,G), or TX (D,H). Shown are representative scanning electron micrographs of the lumenal (A D) and ablumenal (E H) surfaces. All four treatments resulted in complete removal of resident cells from the surfaces of the HUV, while leaving behind the basement membrane (A D). Small areas where the underlying fibrillar collagen are exposed underneath denuded basement membrane are shown in the insets. Scale bar: 50 m. Figure adapted with permission from the Journal of Biomedical Materials Research Part A.
57 Figure 3 4 EC adhesion to basement membranes of decellularized HUV scaffolds. EC were seeded onto the lumenal surface of decellularized HUV. Shown are repre sentative images of DAPI stained EC monolayers after 24 hours on HUV scaffolds decellularized using EA (A), NaCl (B), SDS (C), or TX (D). EC adhesion was quantified on several donor scaffolds (n=5); results are presented per surface area (E). Asterisk indi cates significant differences in mean EC adhesion between treatment groups (p<0.05). Scale bar: 100 m. Figure adapted with permission from the Journal of Biomedical Materials Research Part A.
58 Figure 3 5 EC viability on decellularized HUV scaffolds. EC were seeded onto the lumenal surface of decellularized HUV and analyzed using a live/dead viabilit y assay. Shown are representative images of calcein stained live EC (A D) and ethidium stained dead EC (E H) after 24 hours on HUV scaffolds decellularized using EA (A,E), NaCl (B,F), SDS (C,G), or TX (D,H). EC viability was quantified on several donor sca ffolds (n=3); results are presented as percent viability (I). Asterisks indicate significant differences in mean EC viability between treatment groups (p<0.05). Scale bar: 50 m. Figure adapted with permission from the Journal of Biomedical Materials Research Part A.
59 Figure 3 6 Lumenal surface morphology of re endothelialized HUV scaffolds. HUV were decellularized using one of four chemical treatments: EA (A,E), sodium chl oride (B,F), SDS (C,G), or TX (D,H). Primary EC were seeded onto the lumenal surface of decellularized HUV. Shown are representative scanning electron micrographs of re endothelialized surfaces after 24 hours (A D) or 7 days (E H) of culture. Scale bar: 50 m. Figure adapted with permission from the Journal of Biomedical Materials Research Part A.
60 CHAPTER 4 IN VITRO METHOD FOR REAL TIME, DIRECT OBSERVATION O F ENDOTHELIAL CELL VASCULAR GRAFT INTERACTIONS UNDER S IMULATED BLOOD FLOW In C hapter 3, the effects of various chemical decellularization strategies on EC adhesion, maturation, and viability on the lumenal HUV surface were compared. Based on these results, SDS was selected as an optimal decellularizing reagent for EC growth on the HUV. 13 7 However, these experiments were performed under static conditions, and not taken into account were the important effects of hemodynamic SS on EC binding, maturation, and adaptation. This chapter will focus on efforts to create a n in vitro microenvironme nt more reminiscient of the native blood vascular interface. Cell adhesion is a critical consideration for implantable biomedical devices, particularly in vascular grafts where appropriate surface properties are critical to graft success. Thrombus formatio n, characterized by excessive platelet adhesion and aggregation, is a significant cause of occlusive failure of vascular grafts. 14,143 As such, persistent efforts have been made to produce materials resistant to pro tein adsorption and peripheral cell attachment. 7 As part of t he design process, an important precursor with blood. 13 Preclinical characterization of a and infiltration has typically been assessed in vivo using vascular bypass 144 or ex vivo shunt models. 107 While bypass surgeries are useful for predicting long term patency of implanted bi omaterials, a significant drawback is that assessment can only be performed terminally. Ex vivo shunt models were developed to allow real time assessment of blood cell accumulation using radiolabelled platelets. 145 Animal trials are expensive and time consuming, however, and more conservative strategies that offer a higher throughput while being cost effective have been sought.
61 Adhesion of circulating blood cells is a dynamic pr ocess mediated by the mechanical forces associated with blood flow. Due to the difficulties associated with i) accurately measuring variable hemodynamic forces across different vascular geometries and ii) imaging peripheral cell adhesion events in vivo PP FC have been used to simulate these processes in a more controllable in vitro setting. These devices produce a parabolic flow velocity profile between two planar surfaces, subjecting each surface to uniform fluid SS 71,116,146 Glass, polystyrene, or other optically transparent substrates can then be coated with cells or proteins, or micropatterned with selectively adhesive peptides/non adhesive monomers to observe cell adhesion under flow using conventional light mic roscopy. Much of our current understanding of the role shear plays in cell signaling, protein conformational changes, and other phenomena associated with peripheral cell adhesion has been obtained through the use of parallel plate chambers. 71,147 Given the practical advantages of real time, non invasive imaging of circulating cell adhesion without the need for expensive imaging systems, we sought to develop a device that could be used to monitor these events on the surfaces of clinically implantable vascular biomaterials. Here we detail the design of a novel PPFC for real time observation of peripheral cell adhesion and morphological adaptation on opaque vascular biomaterials under controlled SS Using labeled cell s uspensions, we demonstrate the capability of this device to permit live capture of dynamic cell adhesion events at the intimal surface of naturally derived blood vessels using time lapse fluorescence microscopy. We also validate the use of this flow chambe r for long term culture of neo endothelia and optimize a flow pre conditioning regime that yields
62 confluent EC monolayers resistant to arterial SS These assays demonstrate the versatility of this device for improved in vitro modeling of physiological cell adhesion events as they occur in vivo Methods and Materials Particle Image Velocimetry GFP+ HL 60 cells were passed through the flow chamber at a constant flow rate, while time lapse imaging was conducted. The objective (5x) was focused at a position hal fway between the imaging window and the lumenal HUV surface in the z direction. HL 60 cells passing through the flow field were imaged at various increments from the center of the channel to the side wall. Velocity was calculated using AxioVision software by measuring the distance travelled by the same cell in two consecutive images, and dividing by the time elapsed. N=10 20 cells were measured at each zone presented. Perfusion C ulture S ystem Assembled flow chambers were connected to a media reservoir fitte d with a 0.22 m air filter for gas exchange. The entire system was placed in a dry incubator maintained at 37C and 5% CO 2 The flow rate through the chamber was directly modulated by the rotational speed of the peristaltic pump, which was controlled by M asterflex Linkable Instrument Control Software V3.1. The mean wall SS to which EC were exposed was calculated according to the Hagen Poiseuille equation (assuming steady laminar flow): (4 1) where is viscosity of water at 37C Q is mean volumetric flow rate, and b and h are the base width and channel height, respectively At the maximum flow rate, the
63 Reynolds number was calculated to be 332, indicating laminar flow. The entrance length (Le) for fully developed flow, calculated according to: (4 2) where Re is the Reynolds number and h is the height of the flow field, was determined to be 11.96 mm. Because the viewing window was located 15 mm from flow field entrance, the area under analysis was subjected to fully developed laminar flow. EC S eeding Assembled flow circuits were sterilized with a solution of 4% ethanol, 0.2% peracetic acid for two hours and balanced with PBS (pH 7.4). Standard EC media ( with 2% FBS) was flowed through the system prior to see ding. EC suspensions (10 6 cells/mL) were inoculated into the flow field and allowed to settle, attach, and spread out on the scaffold surface for 5 hours before initiating flow. EC V iability A ssay EC monolayers were assessed for viability using the Live /D ead Viability/ Cytotoxicity Kit for mammalian cells as described in Chapter 2 EC were imaged in situ within flow chambers to avoid distortion secondary to scaffold manipulation. Images were captured through both the GFP and DsRed filters to visualize live and dead cells, respectively. EC S taining and I mage A nalysis At the end of each experiment, scaffolds were rinsed in PBS, formalin fixed, and co stained using RP /DAPI as described in Chapter 2 EC were imaged in situ within flow chambers to avoid distortio n secondar y to scaffold manipulation. Scaffolds were imaged
64 in 10 predetermined locations per experimental condition and analyzed for cell density and percent area coverage as described in Chapter 2. RP images obtained at 10x magnification were analyzed us ing the OrientationJ plugin to determine actin fiber orientation with respect to the flow direction. 148 DAPI images obtained at 10x magnification were analyzed using the built in particle analysis function of ImageJ to determine perimeter, area, and fit each nucleus with an ellipse to determine major axis orientation. The circularity (C) of each nucleus was calculated according to: (4 3 ) Where A is surface area and P is perimeter of each nuc leus. The angle of the major axis was also determined with respect to the flow direction. Cells were individually analyzed from each 10x flow field to determine mean values for each geometric parameter (n=5 7 images per condition) Neutrophil A dhesion to E ndothelialized HUV GFP+ HL 60 cells were differentiated into neutrophil like cells by exposure to 1.3% DMSO (Fisher Scientific; Hampton, NH) for 3 days, as previously described. 149 Endothelialized HUV scaffolds were activated for four hours by adding 1 U recombinant human TNF (Thermo Scientific # EN RTNFAI ; Hampton, NH) to the media reservoir. Flow chambers were then removed from the incubator and placed on the stage of a Zeiss AxioImager M2 upright fluorescence microscope (Figure 4 3, A). Suspensions (10 6 cells/mL) of differe ntiated HL 60 cells (dHL 60) were then drawn through the flow chamber at a calculated wall SS of 1 dyne/cm 2 using a programmable syringe pump (Harvard Apparatus). Time lapse imaging of dHL 60 adhesion to endothelialized HUV
65 was then performed over a 5 minu te period. Culture media without cells was flowed through for an additional minute in order to remove non adherent cells. Monolayers were fixed using 10% formalin and co stained with RP /DAPI as described above. Results Flow C hamber D esign and A ssembly A PP FC was designed to perfuse fluid (culture media or peripheral blood) across an opaque substrate in an observable environment. Three design criteria were considered paramount: 1) that live imaging can be conducted under high magnification (40x or greater) f or dynamic visualization of platelet adhesion, 2) that the device be capable of maintaining parabolic flow while accommodating scaffolds of various thicknesses, and 3) that the reagent volume required to fill the chamber be minimized to maximize use from l imited volumes of peripherally drawn blood. A significant challenge was decreasing the distance between the top of the chamber and the substrate to allow the focal plane of the microscope objective to reach the scaffold surface. To accomplish this, a thin acrylic slide was bonded over two lanes forming the walls of the flow channel using clear epoxy resin to create a viewing window (Figure 4 1 A ) with a total focal distance of 1.6 mm. With this design a 40x Zeiss objective (LD Plan Neofluar) with a maximum working distance of 2.9 mm was compatible for high magnification imaging of cell adhesion within the flow chamber. Decellularized HUV scaffolds of uniform thickness (750 m), prepared as previously described 7 were sliced open axially and affixed to acrylic base plates in a slightly tensed conformation using a compressible silicone gasket (Figure 4 2, A C). The main body of the flow chamber was the n screwed to the base plate, creating a sealed parallel plate flow channel with the bottom plate composed of the lumenal HUV surface
66 (Fig ure 4 2, D E). By minimizing the base width (b) and height (h) of the flow field ( b 6.35 x h 0.60 mm), the volume requi red to fill the flow channel was reduced to < 300 L, approximately 20% of the volume of the native HUV vessel (average diameter 5 6 mm). No difference in mean flow velocity was observed across >90% of the channel width, ind icating uniformity of flow ( Fig ure 4 1, B ). Development of N eo endothelia on HUV S caffolds and A daptation to F low Primary EC were seeded onto the lumenal surface of the HUV at an initial seeding density of 60,000 cells/cm 2 which correlated with a nearly confluent density for the same cells when cultured on polystyrene. After allowing t he cells to adhere and spread out over the basement membrane for five hours, e ndothelia were pre conditioned over a 48 hour period using one of four strategies as indicated in Figure 4 4, A In G roup 1, EC were exposed to low SS (0.3 dynes/cm 2 the minimum programmable rotational speed for media exchange) for the entire period; this resulted in a high density of EC at the end of the pre conditioning period (Figure 4 4, B). In G roup 2, seeded EC were immediately exposed to high SS (6 dynes/cm 2 representing approximately half the desired mean SS value), which resulted in a significant decrease in cell density compared to group 1 (126 4.92 83.86 vs. 179 .38 38.57 cells/m m 2 p=0. 000 ). Based on the initial seeding density (approximately 600 cells/mm 2 ), this sugges ted that immediate exposure to high SS caused the EC to strip off the lumenal surface. To allow the EC to more gradually adapt to SS, Group 3 was designed in which the rotational speed of the pump was steadily ramped from 0.3 to 6 dynes/cm 2 over the 48 hou r period. This resulted in improvement of the cell density when compared to Group 2, but the cell density was significantly lower than in Group 1 ( 1264.92 83.86 vs.
67 640 .17 110.84 cells/mm 2 p=0.000 ). To allow the EC additional time to mature on the HUV sur face, Group 4 was designed in which the cells were cultured under low SS (0.3 dynes/cm 2 ) for the first 24 hours, and then exposed to ramped flow (0.3 to 6 dynes/cm 2 ) over the second 24 hours. This strategy resulted in a similar cell density as Group 1, wit h no statistically significant difference between these two groups ( 1264.92 83.86 vs. 1236 .00 169.30 cells/mm 2 p=0.969 ) Further experiments were conducted using the shear pre conditioning strategy described in Group 4. S hear C onditioning of N eo endotheli a We next used the device to assess the potential of the decellularized HUV to support an EC monolayer under physiological SS c onditions for experimentally relevant time periods Endothelia were conditioned to SS by ramping the flow rate as indicated in Fi gure 4 5 A Cytoskeletal F actin filaments aligned parallel to the flow direction, and a significant decrease in the average filament angle relative to the shear direction was observed after flow ramping was initiated (Figure 4 5, C). This correlated with nucleus alignment in the flow direction (Figure 4 5, D). Progressive nuclear elongation was observed after physiological levels of SS were applied (see Figure 4 5, E) Two days after seeding, a confluent EC monolayer was developed that was able to withsta nd SS commonly found in small diameter arteries, confirmed by an additional day of perfusion culture Confluent neo endothelia were maintained under pulsatile perfusion (meanamplitude 1212 dynes/cm 2 at 80 pulses/min) for up to 7 days after seeding ( Figur e 4 5 B)
68 Morphological C omparison of E ndothelia C ultured U nder S tatic or F low C onditions Neo endothelia cultured on the opaque lumenal surface of decellularized HUV (as described above) were subjected to either pulsatile SS (1212 dynes/cm 2 at 80 pulses/ min) or static culture conditions for 24 hours. SEM taken from an oblique (45) angle relative to the HUV surface show EC cultured under flow to have a significantly more flattened morphology, with less variation in height than s tatic cultured cells (Figur e 4 6 A B). This finding correlated with cellular elongation i n the flow direction (Figure 4 6 E F), which accounts for the reduced topographical variation between cell cell junctions that is visible in SEM images. Calcein/ethidium co staining confirmed a high proportion of viable cells on endothelial monolayers cultured on the HUV basement membrane, and exposure to arterial SS levels had no deleterious effect o n cellular viability (Figure 4 6 C D). Real time O bservation of N eutrophil R olling O n N eo endo thelia For functional assessment of neo endothelia, we designed a system to image rolling adhesion and arrest of human leukocytes under flow. After 3 days of flow conditioning EC monolayers were activated with recombinant human TNF to stimulate cell adhe sion molecule expression. Time lapse imaging conducted at 10x magnification (for observing a larger surface area) captured rolling adhesion and arrest of GFP expressing neutrophil like differentiated HL 60 cells (dHL 60) over a 5 minute period (Figure 4 7, A). Quantification of adherent cells in each frame showed a roughly linear increase that reached a plateau after 4 minutes. RP /DAPI co staining confirmed the presence of a confluent EC monolayer beneath adherent dHL 60, which predominantly localized at ce ll cell junctions (Figure 4 7 B E). Firm adhesion lasted an average of
69 29.115.9 seconds per event, and dHL 60 rolling velocity was generally higher after fir m adhesion than before (see Figure 4 7 C D ). Discussion Several techniques have been developed f or visualizing the dynamic events associated with platelet aggregation/leukocyte adhesion both in vivo and in vitro Intravital microscopy allows real time observation of leukocyte trafficking to sites of injury, rolling adhesion, and extravasation 19 9 32 as well as thrombus formation in situ 6 17 While intravital microscopy is useful for visualizing local cell influx during inflammation, it can only be performed using small animal models due to limited penetration depth of the light source. 31 Other systems have been developed for analyzing vascular behavior ex vivo where biochemical factors, the mechanical stimuli, and interactions with isolated cell populations can be precisely controlled. 21 These types of systems are particularly suited for observing agonist induced contraction/dilation and other whole vessel responses 4 but have limitations for observing intimal phenomena due to an increas e in light scattering coupled with the loss of resolution associated with thicker, opaque vessels. Alternatively, harvested vessels can be opened longitudinally to directly image cells adhered to the lumenal surface, but this does not permit real time imag ing under controlled flow conditions. 3 PPFC have been used for several decades to study behavior of particular cell populations under flow. 10 These devices have several advantages over tubular ex vivo derived or eng ineered vessels: lower reagent requirements, more predictable SS and non invasive observation with enhanced clarity compared to intravital microscopy. In most parallel plate designs, a vacuum channel is incorporated to pull a glass slide over
70 a gasket of known thickness in order to seal the flow channel; for this reason, the substrate must be rigid. 10 25 The use of glass/polystyrene substrates in these systems has become a limiting factor as the dynamics of cell adhesion are significantly influenced by the physical properties of the underlying culture substrate (e.g. surface modulus, topogra phy), which affects the way cells attach, spread, and adapt to fluid forces. 11 13 Given the disparities in the physical properties of ECM proteins of blood vessels and tissue culture plastics, the use of these materials for studying cell behavior under flow may be less than ideal. To overcome the limitations associated with ima ging through thick, opaque vessels, we designed a novel flow chamber that holds a compliant vascular biomaterial in a predetermined conformation such that SS across the wall can be precisely controlled. By cutting a vessel open longitudinally and securing it between two plates, the intimal surface can be visualized while whole blood, select cell suspensions, or media with agonists are perfused across. This permits live imaging using conventional epifluorescence microscopy techniques. The specific protein co mposition of the vascular basement membrane regulates EC phenotype via ECM mediated integrin signaling. It has been previously shown that sub endothelial protein substrates influence activation of signaling pathways elicited by SS 24 The use of a natural basement membrane may therefore serve as a more physiologically relevant substrate for modeling hemodynamic SS patterns in vitro In the present study, we were able to temporally observe EC prolife ration and morphological adaptation to SS applied in situ To avoid stripping EC off the HUV surface by the sudden onset of flow, cells were shear conditioned by progressively increasing flow
71 rates until physiological levels of arterial wall SS were reache d. Adaptation of EC on other vascular biomaterials to physiological shear levels could similarly be optimized using this device in preparation for implantation in the aggressive arterial circulation. In summary, we have designed and characterized a novel f low chamber that uses ex vivo derived vascular tissues for in vitro modeling of cell adhesion events in an environment more reminiscent of the natural vasculature. We have demonstrated functional uses of this device for investigating cell adhesion events d uring normal vascular homeostasis or in response to injury. The design presented can be used for pre implantation screening of a wide range of other biomaterials to assess blood cell interactions.
72 Figure 4 1. Flow chamber design. The main body design of the flow chamber is shown in (A). Culture media/cell suspensions are delivered through inflow/outflow ports. An acrylic viewing window was bonded to the center of the flow field to accommodate high magnification imaging under flow. The flow field is com posed of an acrylic chamber and substrate (not shown) in parallel plate geometry, which allows for simple SS calculations. The inset schematic shows how wall SS ( w the flow field dimensions (equatio n 1). B : Velocity profile across the channel width. Particle image velocimetry was used to calculate mean flow velocity (presented as mean SD) as a function of distance from the center of the channel. Velocity measurements were grouped as a function of dis tance from the center of the flow field base. Uniform flow velocity was observed across >90% of the channel width.
73 Figure 4 2. Flow chamber assembly. Decellularized HUV scaffolds were cut open axially (A) and affixed to a base plate (B) using a compres sible gasket (C). The base plate was then tightly screwed to the flow chamber, creating a parallel plate flow profile across the HUV surface. A cross sectional view of the various components is shown in (D). Assembled flow chambers could then be placed on the stage of an upright epifluorescence microscope for high magnification real time imaging of adhesive interactions between
74 Figure 4 3. Perfusion systems for imaging initial adhesion events o r extended culture under flow. (A): Short term imaging schematic. Syringe pumps were used to inject HL 60 cell suspensions through the flow chamber under defined shear conditions. A high speed camera captured binding events of fluorescently labeled cells a t high magnification through a long working distance microscope objective. (B): Long term recirculating perfusion culture. Computer controlled peristaltic pumps were used to modulate the flow across endothelialized HUV membranes. A bubble trap/pulse dampen er was incorporated to abate EC denudation, while a pressure transducer was used to monitor pressure downstream of the flow chamber.
75 Figure 4 4 Shear pre conditioning strategies for EC seeded HUV scaffolds. Decellularized HUV grafts were cut open axia lly and placed in the flow chamber as described above. Primary EC were seeded onto the lumenal surface, after which they were adapted to flow using various shear pre conditioning strategies (A). After 48 hours, the average cell density was quantified from each group (B). Results are presented as mean+SD. Asterisks indicate significant differences in mean with respect to Group 1. N.S. indicates no significant difference. Representative images of the lumenal surface from each group are shown in panel C Scale bar : 100 m.
76 Figure 4 5. Maturation of endothelialized HUV scaffolds. Decellularized HUV scaffolds were affixed to flow chambers as described. EC were seeded onto the lumenal HUV surface at a high concentration and gradually adapted to flow using computer contr olled peristaltic pumps. Panel (A) shows the calculated mean wall SS along the lumenal HUV surface over the culture period. At various time points (indicated by the dashed lines), monolayers were fixed and co stained with RP /DAPI to visualize F actin/cell nuclei, respectively. Panel (B) shows representative images of EC morphology at 1, 2, 3, and 7 days after seeding, respectively. Actin filament angle (C) was quantified using the OrientationJ plugin of ImageJ, and nucleus major axis angle (D) and circulari ty (E) were quantified using the built in functions of ImageJ Results are presented as mean+ SD values for each parameter (n=5 7 10x images per time point). Asterisk indicates significant differences in mean as specified, and # indicates significant differ ences in mean relative to all other time points. Scale bars: 20 m.
77 Figure 4 6 Morphological comparison of endothelia cultured under static culture or flow conditions. Processed HUV scaffolds were fit in custom designed flow chambers as described previously. Primary EC suspensions were injected into the flow fiel d and either cultured under static conditions (A,C,E) or adapted to flow (B,D,F) over a 72 hour period. Scanning electron micrographs of endothelialized HUV scaffolds at oblique angle show the more flattened morphology under flow (B) than under static cult ure (A). Calcein/ethidium live/dead staining reveal s high cellular viability in static culture (C) or under SS (D). RP /DAPI staining shows cytoskeletal F actin (red)/cell nuclei (blue) (E,F). Arrow shows the flow direction in B, D, and F. Scale bars: 50 m.
78 Figure 4 7 Time lapse capture of neutrophil adhesion to endothelialized HUV scaffolds. Endothelia adapted to flow were activated with 1 U TNF A bolus of dHL 60 was then flowed across the lumenal HUV surface at a calculated SS of 1 dyne/cm 2 Tim e lapse imaging captured adhesion of GFP+ dHL 60 cells to the activated endothelia over 5 minutes (A). Shown are images at various time points throughout the experiment. After dHL 60 cells were flowed across the surface, non adherent cells were washed away and scaffolds were fixed and co stained with RP /DAPI. Multi dimensional images show a confluent endothelial monolayer with adherent GFP+ dHL 60 cells (B E). B: cell nuclei; C: F actin; D: GFP+ neutrophils; E: overlay. Scale bar: 25 m. Additional box plots (C) demonstrate the firm adhesion duration and rolling velocities measured before (upstream) or after (downstream) firm adhesion occurred. Box plots show the range, 1st 3rd quartiles (shaded region), and median (dark line) for each parameter.
79 CHAPTER 5 ADAPTATION OF ENDOTH ELIAL CELLS TO VARIABLE, PHYSIOLOGICALLY MODELED SHEAR STRESS STIMULATION: DISCREP ANCIES WITH FIXED MECHANICAL STIMUL ATION In Chapter 4, a temporal flow regime was designed to facilitate EC perfusion culture on the HUV surface in a directly observable environment. In this chapter, a traditional PPFC design was used to analyze the effect of variable SS patterns on EC gene expression. A glass substrate was used in place of the HUV scaffold so that donor to donor varia tions in basement membrane composition would not affect the analyses described. Vascular EC are directly exposed to hemodynamic SS the frictional force applied by blood flow, and this stimulus is a principal mediator of EC phenotype. 54,73 Acute changes in blood flow patterns, which occur in response to variations in cardiac output/downstream metabolic demand, also change the patterns of SS applied, thereby eliciting phenotypic adaptations ( e.g. changes in gene trans cription/protein expression) in EC. It is has previously been demonstrated using in vitro SS generating culture systems that EC behave significantly differently under SS than they do under static conditions. Applied SS causes changes in gene transcription (up/downregulation) as well as protein expression/function. 54,73 Short term adaptive changes to acute increases in SS (i.e. physiological increases in blood flow) include morphological reorientation of the cytoskel eton 112,150 and intracellular protein localization 151 and stimulation of enzymatic activity. 85,126 SS also stimulates metabolic production of endothelial derived paracrine factors that regulate the physiology of both cells of the vascular wall (e.g.
80 SMC /fibrobla sts) as well as those in the circulation (e.g. platelets, leukocytes, and stem cells). 73,85 In vivo the SS waveform to which the endothelium is exposed is dependent on blood flow conditions that vary by cardiovas cular load, downstream metabolic demand, and local vascular geometry. As a result, EC phenotype is spatially heterogeneous throughout the vasculature. 57 Supporting clinical evidence exists in the focal development of atherosclerotic lesions in areas of the vasculature th at experience disturbed (e.g. oscillatory or reversing) blood flow patterns, which have been linked to endothelial dysfunction. 152 156 The focal development of cardiovascular disease states in areas of the vascular wall exposed to disturbed blood flow underscores the sensitivity of the endothelium to variations in applied SS patterns. To better understand the molecular signals underlying these phenotypic discrepancies, a number of useful computational models have be en developed to recreate in vitro the atheroprotective/atherogenic SS profiles to which EC are exposed in various locations throughout vascular wall. 67,70,157 The adaptation of EC to deleterious SS patterns, such as shear gradients or flow oscillation, has been characterized by increased expression of atherogenic transcription factors, such as NF kB, leading to a sustained pro inflammatory state. 67,68,114,157,158 In contrast, exposure of EC to unidirectional, laminar flow downregulates inflammatory cell adhesion molecules and cytokines, and increases production of relaxing factors such as NO that inhibit cell adhesion, migration, and proliferation. 68,70,156 An equally important consideration in the regulation of EC phenotype by hemodynamic SS is the dynamic nature of blood flow rate with respect to temporal
81 demand. In vivo homeostatic changes in cardiac output vary blood flow and pulse frequency to meet metabolic demand (e.g. during exercise), and so SS patterns at the blood vascular endothelium interface is dynamic and in a constant state of change. 159 Lo cal hemodynamic shear patterns, whether atherogenic or atheroprotective, are therefore not fixed mechanical stimuli but highly dynamic in terms of magnitude, amplitude, and duration. Here we tested the hypothesis that temporal cha nges in applied SS patterns have an important influence on EC phenotypic expression. In this study, we designed an in vitro model of physiological blood flow primarily intended to mechanically stimulate EC across a variable range of SS, rather than a fixed or steady state stimulus, which has been common in most model systems. Results highlight the significant phenotypic differences between primary human EC cultured under temporally modulated and steady pulsatile flow in vitro as relevant to their critical r oles in thrombosis, hemostasis, and inflammation. 72 Methods and Materials EC P erfusion C ulture EC were seeded onto glass cover slips and allowed to grow to confluence over 48 hours before initiating flow. Monolayers were affixed to PPFC using a vacuum pump and connected to a media reservoir fitted with a 0.22 micron air filter for gas exchange. The entire system was placed in a dry incubator maintained at 37C and 5% CO 2 Masterflex Linkable Instrument Control Software V3.1 was used to control digital peristaltic pump drives (Masterflex) that generated pulsatile flow of media through each chamber (Figure 5 1 ) The mean wall SS ( ) to which EC monolayers were exposed
82 was calculated according to the Hagen Poiseuille equation (assuming steady laminar flow): (5 1) where is media viscosity, Q is mean volumetric flow rate, and b and h are the base width and channel height, respectively. Mean volumetric flow was measured empirically, and system pressure was monitored using TruWave pressure transducers wit h Sonometrics SonoLAB digital acquisition system. qPCR After 24 hours, EC were collected and mRNA was purified using the RNAqueous 4PCR Kit ( Life Technologies #AM1914; Grand Island, NY ) according to instructions. mRNA was then normalized to an equivalent a mount (400 ng per sample) and reverse transcribed to cDNA using the RT 2 First Strand kit ( Qiagen # 330401 ; Valencia, CA ) according to instructions. cDNA was combined with RT 2 SYBR Green qPCR Master Mix (Qiagen # 330501 ) and loaded onto human EC biology PCR a rrays ( Qiagen # PAHS 015ZE 4 ). A CFX384 Real Time PCR Detection System (Bio Rad; Hercules, CA) was used to perform quantitative PCR. Amplification was performed at 95C for 10 minutes, followed by 40 cycles of (95C for 15 seconds and 60C for 60 seconds). The comparative C T method 160 was used to quantify gene expression relative to the housekeeping genes GAPDH RPL13A B2M ACTB and HRP ; EC cultured under static conditions were used as calibrating samples. Gene expression was reported as fold ch anges relative to calibrating samples; downregulated genes were reported inversely as negative fold changes.
83 Nitrate/nitrite Q uantification Total nitrate/nitrite content (stable salt derivatives of NO ) in conditioned media was quantified using the Nitrate/ Nitrite Fluorometric Assay Kit (Cayman Chemical #780051; Ann Arbor, MI ) according to kit instructions 4 6 replicate samples were analyzed per flow condition, and results were normalized by cell number. HL 60 C ell A dhesion A ssay Monolayers were activated d uring the final four hours of flow with 1 unit (0.16 ng/mL) recombinant human TNF ( Thermo Scientific #EN RTNFAI; Hampton, NH ). At hour 24, cover slips were removed from flow circuits, rinsed in media, and incubated with a bolus (10 6 cells/mL) of GFP expr essing HL 60 cells within Petri dishes for 10 minutes. Monolayers were rinsed 3 times in PBS and co stained with RP /DAPI as described in Chapter 2 The number of adherent HL 60 cells were quantified in 15 specified locations throughout the flow field for e ach condition (n=6). Statistical A nalysis Results are presented as meanSEM. One way ANOVA followed by post hoc Tukey Kramer HSD analyses (with the significance level set at 0.05) were conducted to compare gene expression by experimental group or NO produc tion by group or time. Alternatively, when equal variances could not be assumed NO test (significance set at 0.05) was used to compare HL 60 cell adhesion. Results We developed a flow regime that mimics constant physiological variability associated with cardiac output to mechanically stimulate endothelial monolayers across a physiological range of arterial SS Real time recordings of human heart rate in a
84 healthy male subject were obtained over a single 12 hour period and programmed into computer driven peristaltic pump drives so that rotational speed corresponded to the observed changes in cardiac output. The calculated mean wall SS within PPFC downstream of the pumps (Fig ure 5 2 ) ranged from a minimum of 9.2 14.2 (mean peak) dynes/cm 2 at 56 pulses/min (Figure 5 2, B ) to a maximum 19.3 43.1 (mean peak) dynes/cm 2 at 142 pulses/min (Figure 5 2, C ) For comparison, steady pulsatile flow was applied at a fixed rate ( 11.7 24 .0 (mean peak) dynes/cm 2 at 80 pulses/min ; Figure 5 3, A ); this was the calculated mean flow rate/pulse frequency from the modeled flow program averaged over the entire 12 hour cycle. Thus, both flow regimes applied the same total magnitude of SS, albeit a t different rates. PF I nduces M orphological Adaptation I n EC Primary human EC monolayers were grown to confluence on glass cover slips and cultured under one of three experimental conditions for 24 hours: static culture, steady flow (SF), or physiological flow (PF). For PF, two cycles were applied back to back to normalize the 24 hour perfusion culture period. EC maintained under static culture conditions grew to a higher density than those cultured under flow, with no global cytoskeletal organization (Fig u re 5 3 A ). As expected, culturing monolayers under laminar flow (SF or PF) induced alignment of cytoskeletal F actin fibers in the flow direction ( Figure 5 3, B C ). No significant global differences in cytoskeletal morphology were apparent between EC cult ured under SF or PF after 24 hours, indicating similar adaptation to both shear regimes. PF I nduces C ardio P rotective G ene E xpression I n EC After 24 hours of conditioning under static culture, SF, or PF, EC mRNA was extracted, reverse transcribed to cDNA, and amplified using qPCR. We assessed
85 expression of several EC genes which, due to their importance in vascular physiology, are standard clinical diagnostic markers of cardiovascular health or disease (Fig ure 5 4). Maintaining EC under 24 hours of either S F or PF resulted in upregulation of superoxide dismutase 1 ( SOD 1 ), an enzyme protective against oxidative stress (SF: 1.4930.064 fold, p=0.001; PF: 1.4950.065 fold, p=0.001). Furthermore, the vasopressive gene endothelin 1 ( ET 1 ), which causes SMC contr action and vasoconstriction and is implicated in atherosclerosis 90 was downregulated by both types of flow (SF: 7.7100.999 fold, p=0.002; MF: 1.775 0.258 fold, p=0.048), with no significant difference between flow groups (p=0.055). EC expression of prostacyclin synthase, a potent vasodilator and inhibitor of platelet aggregation, and characteristic marker of endothelial quiescence, was dependent on th e flow regime imposed. Interestingly, PTGIS was dramatically downregulated by SF ( 4.4890.809 fold, p=0.004), but unaffected by PF (1.0320.104 fold, p=0.171). PF D oes N ot S ignificantly M odulate E xpression of C oagulation/fibrinolysis G enes I n EC Distinct expression trends were found in EC genes associated with coagulation and fibrinolysis ( Figure 5 5 ) between static culture and SF. Annexin A5 ( ANXA5 ), a plasma protein with anticoagulant properties, was upregulated by SF (1.4290.089 fold, p=0.015), but not significantly affected by PF (1.3170.073 fold, p=0.062) Anti clotting proteins tissue factor pathway inhibitor ( TFPI ) and thrombomodulin ( THBD ) did not significantly vary by experimental group. Fibrinolytic enzyme urokinase plasminogen activator ( PLAU ) was upregulated by SF alone (1.8440.223 fold, p=0.031), and plasminogen activator inhibitor ( PAI 1 ) expression was significantly higher in cells conditioned under PF than SF (1.5410.227 fold vs. 1.4110.145 fold, p=0.034),
86 though neither deviated signif icantly from static culture. Overall, SF lowered expression of procoagulant genes and increased expression of fibrinolytic proteins, and while PF evoked similar trends, none significantly deviated from static culture. PF I nduces H igher EC E xpression of C he motactic F actors T han SF EC expression of genes associated with inflammation varied greatly depend ing on flow conditions (Figure 5 6 A B). Culture under SF resulted in upregulation of the cell adhesion molecules ICAM 1 (4.4670.660 fold, p=0.006) and PECA M 1 (3.1800.152 fold, p=0.000), and downregulation of VCAM 1 ( 9.1040.116 fold, p=0.006) compared with static culture. No PF conditioned EC adhesion molecules significantly varied in expression relative to static controls, though PECAM 1 expression was s ignificantly lower (1.5660.196 fold, p=0.000) when compared to SF Expression of monocyte chemoattractant protein 1 ( MCP 1 ) was significantly downregulated under either flow condition, though significantly more so in EC conditioned under SF ( 9.6201.352 fold) than PF ( 2.7150.246 fold, p=0.049). Fractalkine ( CX3CL1 ) was upregulated in PF conditioned cells only (2.5220.437 fold, p=0.028), and was significantly higher than in cells cultured under SF ( 1.8820.611 fold, p=0.004). ADAM17 an enzyme that med iates TNF shedding from the cell membrane, was upregulated by PF only (2.1070.297 fold, p=0.035). However, cytosolic phospholipase A2 gamma ( PLA2G4C ) expression was upregulated irrespective of flow (SF: 3.1900.501, p=0.015, PF: 2.966+0.293, p=0.036). P F I nduces S ustained E ndothelial N itric O xide S ynthase A ctivity Endothelial nitric oxide synthase (eNOS) converts L arginine to L citrulline and NO the latter of which is an important inhibitor of leukocyte adhesion, platelet aggregation, and smooth muscl e proliferation. This enzyme is constitutively expressed,
87 yet highly regulated in EC. 161,162 NO is an important signaling molecule in vascular physiology, as its inhibition of leukocyte adhesion and platelet aggrega tion facilitates laminar, uninterrupted blood flow. EC expression of eNOS mRNA was significantly upregulated after 24 hours of flow conditioning (SF: 2.0490.086 fold, p=0.000; PF: 1.7150.069 fold, p=0.003, Figure 5 7 A). However, there was no significan t difference in eNOS gene expression between the two flow groups (p=0.076). eNOS enzymatic activity was assessed by quantifying the total amount of nitrates/nitrites, stable salt derivatives of NO in conditioned culture media after 0, 12, or 24 hours of c ulture and normalized by EC number (Fig ure 5 7 B). Under static culture conditions, NO content did not significantly vary by culture time: (0 hours: 38.7458.144 nmol/10 5 EC, 12 hours: 41.55210.308 nmol/10 5 EC, 24 hours: 33.6802.614 nmol/10 5 EC, p=0.689 ). After 12 hours of perfusion culture, significant increases in NO byproduct accumulation were observed (SF: 88.3087.608 nmol/10 5 EC, p=0.025; MF: 84.99311.233 nmol/10 5 EC, p=0.030), with no significant difference in means observed between flow groups ( p=0.969). However, noteworthy deviations in NO output occurred between flow regimes during the last 12 hours of perfusion. No significant increase in nitrates/nitrites was observed between 12 and 24 hours of SF (12 hours: 88.3087.608 nmol/10 5 EC, 24 hours : 115.03912.460 nmol/10 5 EC, p=0.186). PF, however, induced significant temporal increases in NO output between 12 (84.99311.233 nanomoles/10 5 EC) and 24 hours (176.011.860 nanomoles/10 5 EC, p=0.000) of conditioning. After 24 hours, the highest nitrate/ nitrite concentrations were measured in media from EC cultured under PF, which was significantly higher than that from EC cultured under SF (p=0.002).
88 NO P roduction B y EC I s S ignificantly H igher W ith P rogrammed S hear C hanges Than Constant frequency F low Fo r added clarity, two additional flow groups were included to profile NO production under physiologically dynamic conditions First, another SF group (SF 160) was included in which the flow was applied at twice the rate of the SF (160 pulses/min, 20.6 dynes /cm 2 ). Similar trends were observed under SF 160 as normal SF; no significant differences in nitrates were observed between the two groups (Figure 5 8 ). Second, the PF model was applied in reverse chronological order. Reversed physiological flow (rPF) indu ced a significant increase in nitrates (relative to static culture, SF 80, and SF 160) after 6 hours that was maintained through 24 hours of culture (Fig ure 5 8 ). However, no significant differences in nitrates/nitrites were observed between PF and RF at a ny of the time points assessed. PF E nhances EC R esistance to A ctivation I nduced L eukocyte A dhesion Given the dramatic increase in NO production noted in PF conditioned EC compared with SF we next performed a functional assay to characterize the extent to which temporal changes in shear influenced EC adhesiveness for leukocytes. After four hours of activation with TNF flow conditioned monolayers were transferred to Petri dishes and incubated with a bolus of promyelocytic GFP + HL 60 cells ( Figure 5 9 ). Conditioning EC under PF significantly reduced HL 60 leukocyte attachment compared to conditioning under SF (PF: 376 HL 60 cells/mm 2 ; SF: 11118 HL 60/mm 2 p=0.003). Discussion Vascular EC phenotype is a function of multiple physiological factors, of which appropriate shear stimulation is critical for maintaining EC in a quiescent, non thrombogenic state. Defining how hemodynamic shear patterns mediate EC function, as
89 well as pathological dysfunction, is critical to our understanding of vascular physiology. 113,152,154,163 As reviewed in Chapter 2, i n vitro systems designed to ex pose EC to controlled levels of fluid SS have greatly advanced our understanding of the molecular mechanisms underlying the transduction of hemodynamic SS into biological signals. 54,164 It is now recognized that spe cific details of the shear waveform have important implications for EC adaptation/functionality. A significant gap lies between the majority of in vitro perfusion culture systems and physiological blood flow patterns seen in vivo 163 Although experimental flow models accurately recreate the shear waveforms experienced at specific locations on the vascular wall 67,70,157 the waveform is most often applied cyclically, without modulation of frequency or amplitude for the duration of the experiment. The temporal variations in blood flow rate that continuously occur in vivo are not taken into account. Given that EC phenotype has been shown to be a function of pulse frequency 124 SS magnitude 165,166 and pulse amplitude 123 all of which vary in vivo throughout the course of the day, blood flow dynamics warrant considerati on in these systems. Here we tested the hypothesis that temporal variability in applied SS would induce changes to EC phenotype as relevant to their critical roles in circulatory/vascular biology. The PF model in this study was developed to simulate phys iological changes in pulse rate and volumetric flow rate as they occur in vivo over the course of a single day. The model encompassed a range of SS commonly found in small diameter arteries. 167 169 SF with an equiva lent time averaged SS and fixed pulse frequency was used to represent the conditions used in current systems.
90 Transcriptional profiling of EC exposed to steady pulse frequencies versus physiologically modeled flow demonstrated significant differences in g ene expression. Prostacyclin synthase, the enzyme responsible for production of a potent vasodilator/inhibitor of platelet aggregation, was significantly downregulated in cells conditioned under SF while no change was induced by PF The leukocyte adhesion molecule/chemokine fractalkine was upregulated under PF alone. In some cases, gene expression trends (relative to static cultured cells) were statistically similar regardless of flow regime (protective enzymes eNOS, SOD 1 were upregulated; leukocyte chemo kine MCP 1 was downregulated). Elucidating the various signaling mechanisms involved in gene transcription regulation by SS is an area of intense interest. The transcription factor KLF 2, for example, has been found to play a role in eliciting the atheropr otective effects of laminar SS in EC, both through induction of eNOS expression as well as inhibition of agonist induced expression of E selectin and VCAM 1. 68 Boon et al have published a review on several known SS responsive transcription factors 69 and the role of these factors in the presently observed results will be the topic of future investigations. Because only endpoint gene expression analysis was performed, temporal trends were not observed and comparisons between the PF model and the steady control at 24 hours alone limits the number of conclusions that may be drawn. Taken collectively, however, the data show a more quiescent EC transcriptome elicited by both flow regimes compared to statica lly cultured cells. Shear stimulation of EC has important implications for metabolic production of NO an important inhibitor of platelet aggregation 75 leukocyte adhesion 7 6 and SMC mitogenesis 78 Diminished NO bioavailability is a hallmark of endothelial dysfunction,
91 which is a prevalent symptom in a wide variety of cardiovascular disorders, includi ng thrombosis 52 intimal hyperplasia 170 p ulmonary hypertension 171 and atherosclerosis. 172 In the present study, notable increases in NO production were observed when EC were exposed to PF, a regime characte rized by temporal increases (and decreases) in SS magnitude/pulse frequency. In agreement with previous studies, we found that application of SS caused upregulation of eNOS gene expression. 173,174 Though eNOS expres sion between cells exposed to each flow regime did not statistically vary, the metabolic activity of the enzyme changed dramatically. While endothelial NO production tapered off after 12 hours of SF PF induced dramatic increases in NO synthesis between 12 and 24 hours. Previous research has shown that endothelial NO production is dependent on not only the overall magnitude but also the rate at which shear stimulation is applied. 125,126,173 While rapid increases i n SS induce a transient G protein coupled receptor dependent burst in NO production, SF results in a state of sustained NO production. 175 We postulate that the tapering off in NO output is due to physiological adaptation of EC to the steady state conditions associated with fixed pulse flow. SS modulates eNOS activity in EC through a vari ety of mechanisms: calmodulin binding, phosphorylation state of certain residues, and association with membrane bound proteins such as PECAM 1. 84,87,161 In particular, acute SS increases have been shown to elevate e NOS activity through phosphorylation of Ser1177 residue by PI 3K. 176 The rapid increases in shear within the PF regime likely stimulated periodic bursts in NO production that were not induced under SF The significa ntly lowered leukocyte adhesiveness in PF
92 conditioned EC observed correlates with higher NO production, as well as significantly lower expression of PECAM 1 when compared to SF 177 The present evidence suggests EC quiescence can be induced by dynamic stimulation across a physiological range of arterial SS es. It is therefore plausible to consider that temporal alterations in mechanical stimulation may affect functionality of other cell types as well. Sydeain et al p reviously demonstrated increased ERK signaling and greater collagen synthesis by human dermal fibroblasts cultured in a cyclic distension bioreactor when incremental increases in stretch were applied, rather than fixed distension. 178 Though disparities exist between the two model systems, both show evidence that the response of cells to mechanical stimulation is dependent on the overall magnitude/frequency of the stimulus. Application of variable mechanical stimuli t o cells cultured in other bioreactor systems that may influence regeneration of blood vessels, tendon, bone, cartilage, or other tissues in vitro In conclusion, we have demonstrated the inherent sensitivity of critical EC functions to temporal changes in applied SS. The steady state flow applied in current culture systems, while useful for close examination of molecular signaling events, may not accurately represent the physiological hemodynamics to which EC are exposed in vivo These investigations show a clear EC phenotype modulation toward a quiescent arterial state and improved functionality in a simulated wound environment. Understanding the conditions that regulate EC function and having the capacity to model these appropriately in vitro has significa nt implications for treatment of clinical pathologies associated with endothelial dysfunction.
93 Figure 5 1. Parallel plate culture system. EC monolayers were grown to confluence on glass coverslips, then affixed to PPFC (A) using a vacuum pump Peristal tic pumps (B) we re used to impose pulsatile flow of culture media (C) over the endothelial monolayers The rotational speed of the pumps wa s controlled by an external computer (D) via RS 232 linkage. Figure adapted with permission from PLOS ONE.
94 Figure 5 2 Physiologically modeled perfusion culture. A 12 hour perfusion regime (hollow line) was developed by programming a series of real time recordings in heart rate obtained in a healthy male subject into peristaltic pumps as ramp changes. For comparison the average SS magnitude calculated over the entire 12 hour cycle ( 11.7 dynes/cm 2 at 80 pulses/min) was programmed as steady pulsatile flow (dark line). Mean wall SS and pulse frequency experienced by the EC within PPFC over time are shown on the left an d right axes, respectively. Insets show pressure waveforms experienced at various time points during SF (A ) or PF ( B C ). Figure adapted with permission from PLOS ONE.
95 Figure 5 3 Cytoskeletal morphology of flow conditioned EC EC monolayers were grown to confluence and cultured under static conditions (A), SF (B), or PF (C) for 24 hours. Monolayers were subsequently fixed and co stained with RP (middle row) and DAPI (top row) in order to visualize f actin and cell nuclei, respectively. Shown are represe ntative images (40x) from each condition. Applying SS resulted in cytoskeletal alignment of EC in the flow direction (horizontally left to right). Scale bar: 20 microns. Figure adapted with permission from PLOS ONE.
96 Figure 5 4 Cardio protective gene expression in flow conditioned EC Shown is fold mRNA expression (with respect to static cultured EC) of genes that promote or inhibit cardiovascular disease progression. Results are presented as meanSEM. An embedded asterisk indicates a significant diffe rence with respect to static controls; an asterisk over a bracket indicates a significant difference between flow groups. Abbreviations: EDN1: endothelin 1; PTGIS: prostacyclin synthase; SOD1: superoxide dismutase 1. Figure adapted with permission from PLO S ONE.
97 Figure 5 5 Coagulation & fibrinolysis associated gene expression in flow conditioned EC Shown is fold mRNA expression (with respect to static cultured EC) of genes associated with hemostasis (A) and fibrinolysis (B). Results are presented as meanSEM. An embedded asterisk indicates a significant difference with respect to static controls; an asterisk over a bracket indicates a significant difference between flow groups. Abbreviations: ANXA5: annexin V; TFPI: tissue factor pathway inhibitor; TH BD: thrombomodulin; PLAT: tissue plasminogen activator; PLAU: urokinase plasminogen activator; PAI 1: plasminogen activator inhibitor 1. Figure adapted with permission from PLOS ONE.
98 Figure 5 6 Inflammation associated gene expression in flow condition ed EC Shown is fold mRNA expression (with respect to static cultured EC ) of cell adhesion molecules (A) and genes with roles in recruitment of inflammatory cells (B). Results are presented as meanSEM. An embedded asterisk indicates a significant differen ce with respect to static controls; an asterisk over a bracket indicates a significant difference between flow groups. Abbreviations: PSGL 1: P selectin glycoprotein ligand 1; ICAM 1: intercellular adhesion molecule 1; VCAM 1: vascular cell adhesion molec ule 1; PECAM 1: platelet endothelial cell adhesion molecule 1; ALOX5: arachidonate 5 lipoxygenase; PLA2G4C: cytosolic phospholipase A2 gamma; ADAM17: ADAM metallopeptidase domain 17; MCP 1: monocyte chemoattractant protein 1; CX3CL1: fractalkine. Figure ad apted with permission from PLOS ONE.
99 Figure 5 7. eNOS function in flow conditioned EC (A): eNOS mRNA expression (presented as meanSEM ) was upregulated under either perfusion condition, but no statistical difference between flow groups was observed. (B): After 0, 12, or 24 hours of conditioning, media was collected and samples analyzed using a fluorometric assay. Total NO byproduct accumulation was normalized by the mean cell count at the end of each period. Results are displayed as meanSEM (n = 4 6) Asterisks denote significant differences in individual significance with respect to t=0 and t=12 hours. Figu re adapted with permission from PLOS ONE.
100 Figure 5 8. NO profiling of EC exposed to additional flow conditions. EC were cultured under the conditions previously described (static, steady flow [SF 80], or physiological flow [PF]) as well as two addit ional control groups. SF 160: The rotational speed of the pump was doubled, so that the pulse frequency (160 pulses/min) and mean SS (20.6 dynes/cm 2 ) applied were twice that of every other flow group. RF: Additionally, the PF cycle (PF; Figure 2) was appli ed chronologically in reverse. Media was collected and samples analyzed using a fluorometric assay. Total NO byproduct accumulation was normalized by the mean cell count at the end of each period. Results are displayed as meanSEM (n = 4). Asterisks denote significant differences in individual means between groups at each time point. Figure adapted with permission from PLOS ONE.
101 Figure 5 9 HL 60 cell adhesion to flow conditioned EC EC monolayers were grown to confluence and cultured under SF (A,C,E,G ) or PF (B,D,F,H) for 24 hours. During the last four hours, EC were activated with 1 U TNF to stimulate adhesion molecule expression. At hour 24, monolayers were removed from flow chambers and incubated for 10 minutes with a bolus of GFP+ HL 60 cells (1000 cells/mm 2 ) and stained as described (A,B: DAPI; C,D : F actin; E,F : GFP+ HL 60 cells; G,H : overlay). Shown are representative images (40x) from each condition Scale bar: 20 microns ( I) : HL 60 cell adhesion in 15 predetermined locations per monolayer was quantified. Results are displayed as meanSEM (n=5 6). Asterisk denotes significant diff erence in means test. Figure adapted with permission from PLOS ONE.
102 CHAPTER 6 ENDOTHELIAL CELL SEE DING AND FLOW PRE CONDITIONING OF IMPLANTABLE HUMAN UMBILICAL VEIN ALLOGRAFTS Seeding cells onto three dimensional biomaterials with an even surface distribution is an initial, yet significant challenge in tissue engineering. In C hapters 3 and 4 HUV scaffolds were cut open and spread out in to a planar conformation to expose the lumenal surface; this permitted static gravitational seeding using direct pipetting This technique is insufficient f or seeding tubular HUV grafts with EC as sedimentation of the cells toward the graft bottom results in an uneven distribution across the lumenal surface. For i n vitro endothelia lization of vascular grafts to be clinically successful (e.g. to prevent thrombus formation and hyperplastic SMC proliferation ), the blood contacting surface must be covered with a confluent monolayer of cells. Current goals for improvement of two stage en dothelialization include efficient adhesion of ex vivo expanded EC isolates, growth to full confluence, and appropriate conditioning to resist arterial SS in as short a time frame as possible. In this chapter we present the design and implementation of a rotational seeding method for improved circumferential distribution of EC on the lumenal surface of decellularized HUV allografts After addressing initial seeding efficiency, we present efforts to develop a confluent, arterial SS and pressure resistant en dothelium on d ecellularized HUV grafts within a clinically acceptable time frame. Methods and Materials Perfusion C ircuit D esign SDS d ecellularized HUV scaffolds were connected to flow circuits in custom designed glass bioreactors as shown in Figure 6 1. E ach scaffold was slipped over stainless steel adapter mandrels at each end secured using cable ties and then
103 inserted into a bioreactor (Figure 6 1, A) 179 Additional ports fitted onto the body of the bioreactor allowed for access to the ablumenal void space surrounding each scaffold. This design therefore facilitates separate flow circuits through the lumen or albumen, which are separated by the HUV sca ffold. Because of the permeability of the tissue, some media loss into the void space occurs during perfusion. In these experiments, the distal ablumenal ports were connected to the perfusion circuit and left open so that as pressure increased in the void space, media was driven back into the lumenal perfusion circuit (Figure 6 1, B). This ensured that the lumenal flow rate was equivalent through all three scaffolds Three assembled bioreactor s w ere connected in series within each perfusion circuit (Figure 6 1, B) Bioreactors were connected via silicone tubing to a media reservoir fitted with a 0.22 m filter for gas exchange. Media was drawn from reservoirs by peristaltic pumps and driven through the lumen of the scaffold. A pulse dampener was incorporated in between the pump and the bioreactor to reduce pressure spikes and to trap gas bubbles; d owns tream pressure was monitored using a pressure transducer. T he diameter of the tubing downstream from the bioreactor s was reduced and the length increased until approximately arterial pressures (120/80 mmHg) were achieved. The entire system was placed in a dry incubator maintained at 37C and 5% CO 2 The flow rate through the bioreactors was directly modulated by the rotational speed of the peristaltic pump, which was controlled by Masterflex Linkable Instrument Control Software V3.1. The mean wall SS ( ) to which seeded EC were exposed was calculated
104 according to the Hagen Poiseuille equation (assuming steady laminar flow and a constant cross sectional area ) : (6 1) where is viscosity, Q is mean volumetric flow rate, and R is inner HUV radius At the maximum flow rate, the Reynolds number was calculated to be 299 indicating laminar flow. The entrance length (Le) for fully developed flow, calculated according to: (6 2) where D is HUV scaffold diameter and Re is the Reynolds number was determined to be 89.7 mm. Therefore, the tubing (matched to the inner diameter of the HUV and adapter mandrels) upstream of each HUV scaffold was extended so that flow was fully developed before reaching the seeded EC. In order to achieve physiological wall SS commonly found in arteries in vivo high molecular weight dextran was added to the culture media. Addition of dextran (1 7%, MW 70,000) to culture systems was previously shown to have no significant effect on prolife ration, enzymatic activity, or osteogenic differentiation of bone marrow derived mesenchymal stem cells. 180 6% dextran (MW: 70,000; Sigma Aldrich #31390 ; St. Louis, MO) was added to the EC culture media to make the viscosity similar to that of blood (~3 mPa*s) Live/Dead analysis was p erformed as described in C hapters 2 and 3 to assess whether addition of 6% dextran had any effect on EC growth. Rotating S eeding A pparatus An apparatus was constructed in order to rotate HUV scaffolds during sedimentation of EC The cylindrical body of the device is capable of holding up to 6
105 bioreactors at one time (Figure 6 2) The apparatus then rotates around the central axis of the apparatus, which is also parallel to the axis of each attached bioreactor The rotational speed of the device is controlle d by a computer driven peristaltic pump drive operated by Masterflex Linkable Instrument Control Software V3.1. EC S eeding Perfusion circuits were sterilized by two hours perfusion of 0.2% peracetic aci d/4% ethanol for two hours at 10 RPM (~56.7 mL/min) f ollowed by 2 hours of PBS (pH 7.4) at 10 RPM to balance the pH of the system. Standard EC culture media was perfused through the circuits at 37 C and 5% CO 2 for two hours at 10 RPM to enhance cell adhesion to the HUV surface. Decellularized HUV scaffolds w ere seeded with EC either under static conditions or axial rotation. For seeding, triplicate bioreactors were separated from each perfusion circuit within a biological safety cabinet by closing the three wa y stopcocks located at each end and then capping t he luer fittings to maintain sterility. EC (passage 4) suspensions (10 6 cells/mL) were injected into the lumenal space, the stopcocks were closed, and the bioreactors were attached to the rotational seeding apparatus located inside an incubator. EC were al lowed to settle out of suspension and attach to the HUV scaffold surface for 5 hours during rotation before the bioreactors were reconnected to the perfusion circuit and flow was initiated. Perfusion C ulture Five hours a fter seeding, low flow (0.5 RPM) w a s initiated for media exchange. Masterflex Linkable Instrument Network software (v3.1) was used to control the rotat ional speed of the pumps, thereby driving the flow rate of media through the HUV lumen. Standard Vasculife VEGF culture media was replenishe d every 3 days, and
106 dextran supplemented Vasculife media was added immediately before flow ramping was initiated EC S taining and I mage A nalysis At the cessation of each experiment, HUV scaffolds were fixed with 10% NBF for 10 minutes, sliced open axially and co stained with RP /DAPI as described in Chapter 2. Each scaffold was imaged in 15 predetermined locations that were divided into 3 circumferential zones (n=5 images per zone). Each image was analyzed using ImageJ as described i n Chapter 2. For rotatio nal seeding, the mean cell density/% coverage for each zone was ranked, so that zone 1 corresponded to the minimum, zone 2 to the median, and zone 3 to the maximum values for each scaffold. The mean cell density and % coverage for each zone were compared b etween experimental seeding groups (n=5 scaffolds per group). For all other experimental groups (after seeding), cell density/% area coverage values for each HUV scaffold were averaged between all 15 locations (n=3 6 scaffolds per group). Statistical A naly sis One way ANOVA followed by Tukey HSD post hoc testing with the significance level set at 0.05 was used to compare individual means. For cases of unequal variances between experimental groups variance) Tam HSD. Results Rotational S eeding I mproves t he C ircumferential D istribution o f EC o n t he L umenal HUV S urface Endothelialization of small diameter prosthetic grafts has had disappointing clinical results due in part to the lack of a continuous EC lining In the absence of rotation,
107 gravitational seeding ( injecting a cell suspension into the graft lumen ) results in an uneven distribution of cells on the inner surface le aving a large area of the graft exposed to blood fl ow. Improving the distribution of EC seeded onto the graft surface could reduce acute vascular graft failure caused by thrombus formation. On a custom built rotational seeding device ( Figure 6 2 ), dHUV scaffolds were rotated either continuously or alternat ely reversed af ter one complete axial rotation at two angular velocities (1 RPH or 100 RPH). The efficiencies of various seeding methods were compared among five different regimens: no rotation, 1 RPH continuous, 1 RPH alternating, 100 RPH continuous, and 100 RPH alternating. circumference. The mean cell density and percent surface area coverage from each zone were arranged within each seeding group, and the zones were ranked according t o the number of cells attached: zone 1 (minimum), zone 2 (median), and zone 3 (maximum). Each zone was separately compared among these five experimental seeding groups. Under static conditions, the majority of EC adhered to the bottom zone of the HUV surfa ce, where the cells settled gravitationally following inoculation. However, this did not correlate with a significant increase in the total % area coverage in zone 3 (the maximum cell density zone) relative to other seeding groups ( Figure 6 3, B). No rotat ion did result in a significantly lower cell density in zone 1 than 100 RPH continuous (4.042.07 vs. 241.22161.81 cells/mm 2 p=0.004) or alternating rotation (4.042.07 vs. 250.9975.99 cells/mm 2 p=0.003).
108 Scaffolds that were rotated at 100 RPH had a si gnificantly higher cell density than statica lly seeded scaffolds (Fig ure 6 3, A) in zone 1, the zone with the minimum number of cells. The 100 RPH alternating rotation caused a significantly higher percentage of the lumenal surface area to be covered by ce lls than no rotation (32.289.67 vs. 1 .051.74 %, p=0.016) in zone 1. Using the present bioreactor system continuous rotation is only feasible if the tubing that connects the bioreactors to the pump head and media reservoir are temporarily disconnected ; o therwise, the tubing will continue to twist until separation occurs This could theoretically compromise the sterility of the system, which would complicate employing this strategy in the clinic. Alternating rotation which was included in order to allow t he scaffolds to be rotated without risking bacterial infection when bioreactors are disconnecte d from the flow circuit, had no significant difference in cell attachment compared to continuous rotation for either velocity. Given the more normalized circumfe rential distribution of EC 100 RPH alternating rotation was selected as the optimal seeding strategy and used for all further experiments. Supplementation of C ulture M edia W ith D extran H as N o E ffects on EC P roliferation or V iability The viscosity of blood is approximately three times that of normal culture media; for this reason, achieving wall SS levels in similarly sized vessels in vitro can be achieved by tripling the lumenal flow rate. An alternate strategy is to increase the viscosity of culture media via addition of a supplement. To achieve a viscosity similar to blood (~3 mPa*s) 6% (w/v) high molecular weight (70 kDa) dextran was added to culture media. This resulted in no significant differences in EC proliferation over three days of culture (Figur e 6 4, A). The doubling times for EC grown in normal and dextran
109 supplemented media during exponential growth were calculated to be 19.92 and 19.53 hours, respectively, indicating similar growth characteristics. The viability of EC cultured in 6% dextran w as significantly higher than in normal media only at 72 hours (98.52 0.41 vs. 94.05 1.39 % viability, p=0.001 ; Figure 6 4, B ), but this was attributed to increased detachment of cells due to over confluence at this time point (Figure 6 4, A) Thus, it was concluded that there was no deleterious effect of dextran supplementation of culture media on EC proliferation and viability and this formulation was used for subsequent perfusion experiments Additional M aturation T ime I mproves S urface C overage and EC R e tention U nder F low Rotational seeding resulted in a significantly more uniform distribution of attached EC compared to static seeding. However, the most efficient seeding method, 100 RPH alternating rotation, resulted in coverage of less than half of the t otal lumenal surface (462.23 cells/mm 2 at 48.61% coverage), indicating that additional maturation time was necessary for the adherent cells to grow to confluence. EC seeded using this method were allowed to mature for 1, 2, or 5 days after initial seeding on the HUV under very low SS (0.07 dynes/cm 2 ). After the specified interval, the lumenal flow rate was steadily ramped up to achieve a mean wall SS of 7.34 dynes/cm 2 after 24 hours (Figure 6 5 A). Initiation of ramped flow one day after seeding resulted i n a loss of cells at the lumenal surface, though the differences in cell density/percent coverage were not statistically significant (Figure 6 5 B C ) Two days of maturation allowed HUV grafts to retain similar cell densities (412.94 76.91 vs. 462.23 125. 13 cells/mm 2 p=0.987) and surface coverage (60.35 5.29 vs. 48.61 12.22 %, p=0.509) as initial seeding Finally, initiating flow after five days of maturation
110 resulted in a significant improvement in cell density (782.70 49.44 vs. 462.23 125.13 cells/mm 2 p=0. 016 ) and surface coverage (94.37 2.85 vs. 48.61 12.22 %, p=0. 005 ) compared to initial seeding. Flow R amping I mproves EC R esistance to Arterial W all SS After 5 days of maturation under low flow the lumenal surface of the HUV allograft was covered by a nearly confluent layer of EC. However, sudden exposure of these cells to arterial SS would likely result in denudation of the HUV surface (as occurred in Chapter 4) resulting in thrombus formation after implantation In vitro endothelialized HUV grafts wo uld likely benefit from flow pre conditioning to strengthen EC resistance to wall SS arterial pressures, and circumferential strain. At day 5, the lumenal flow rate was gradually ramped up according to a previously designed temporal shear regime (Chapter 4). Over a 48 hour period, the mean wall SS steadily increased from 0.3 dynes/cm 2 to 14.5 dynes/cm 2 while system pressure increased to a maximum of 111.1/79.6 mmHg ( Table 6 1, Figure 6 6 A). Onset of ramp flow resulted in a slight, though not statistical ly significant loss of cells at days 6 (702.13 94.76 vs. 737.55 177.69 cells/mm 2 p=1.000) and 7 (619.24 172.15 vs. 737.55 177.69 cells/mm 2 p=0.713) (Figure 6 6, B) Despite the decrease in cell density, after onset of ramped flow the lumenal coverage by EC increased to 97.33 3.55%, correlating with a 25.7% increase in the mean surface area per cell. This was also apparent morphologically, as cells appeared more spread out after onset of ramped flow compared to days 1 and 5 (Figure 6 6 C). EC progressivel y aligned cytoskeletal F actin filaments in the flow dir ection over time (Figure 6 6 C). After 48 hours of flow ramping, the flow was held constant at 80 pulsations per minute and a mean wall SS of 14.5 dynes/cm 2 for 3 days. Despite the early loss of cell s,
111 the cell density and surface coverage stabilized to final values of 644.77 cells/mm 2 and 88.88 % coverage (Figure 6 6 B). This indicated that the seeded endothelia were capable of withstanding arterial mechanical stresses After 10 days of perfusion cu lture, a con tinuous EC monolayer was observed attached to the basement membrane on the lum enal HUV surface (Figure 6 7 ). Discussion Small diameter prosthetic grafts in current clinical use fail to match the patency rates of autologous vessels due to occlus ive failure brought on via two main mechanisms: thrombus formation and anastomotic intimal hyperplasia. It has been suggested in numerous studies that the compliance mismatch between rigid prosthetic (e.g. Dacron or ePTFE) grafts and the adjacent host vess els causes alterations in hemodynamics throughout the graft that result in distal intimal thickening. 16,181 Compliance is an important mechanical attribute in n ative blood vessels that allows them to accommodate inc reases in blood pressure during systole while maintaining laminar blood flow Efforts have been made to synthesize vascular biomaterials with improved compliance relative to currently used prosthetic materials. 182 An alternative option that has become increasingly popular i n recent years is the development of TEVG from processed, ex vivo derived blood vessels that retain their native structural and mechanical properties making them appropriate for use as vascular conduits. 40,99,107 Th e HUV has recently been highlighted for use as a scaffolding material for vascular tissue engineering. 40,137,183 It has been reported that a t full term (37 40 weeks gestation) the HUV has compliance values similar to other autologous vascular grafts (e.g. greater saphenous vein, brachial artery) 40,184 The native structural properties
112 coupled with the regular availability of the tissue in hospital delivery wards, therefore m ake the HUV an appealing scaffolding material for vascular reconstruction. Another complication that contributes to early vascular graft failure is thrombotic occlusion which occurs in the weeks to months following bypass surgery 13,143 Thrombus formation has been attributed to the passive adsorption of plasma proteins on the blood contacting surface of prosthetic materials leading to platelet adhesion, activation, and aggregation ultimately resulting in graft fail ure. It has been hypothesized that the superior patency rates of autografts compared to prosthetic materials are largely due to the presence of an inner EC lining. 23 EC, which line the interior surfaces of native ve ssels, not only shield thrombogenic ECM proteins of the vascular wall from circulating blood, but actively inhibit cellular adhesion through the constitutive production of NO prostacyclin, and other biochemical factors. 52,138 Given the naturally non thrombogenic properties of EC, t he development of an inner EC layer, or endothelialization, would therefore theoretically improve the patency of prosthetic grafts 104 However, despite promising results in animal models 105 this strategy has not achieved widespread success in clinical trials. 92 The main issues associated with pre endothelialization are 1) the lack of confluence of EC seeded on the lumenal surface and 2) the denudation of EC seeded grafts caused by sudden onset of blood flow. Here we focused on two factors that relate to the efficacy of in vitro endothelialization of vascular allografts: improving the efficiency of the seeding technique and conditioning seeded cells to better resist mechanical stresses
113 Seeding three dimensiona l biomaterials with a uniform distribution of cells is a fundamental though significant challenge in tissue engineering. For vascular scaffolds, which are axially symmetric, rotation around the longitudinal axis (which passes through the vessel lumen) all ows lumenally injected EC settling out of suspension to freely attach to the entire lumenal surface Rotational seeding has previously been performed to improve the lumenal surface coverage of vascular grafts 93,102, 179 In agreement with these previous studies, axial rotation was shown to result in an improved distribution of ce lls on the lumenal HUV surface. A practical issue in many vascular perfusion systems that is associated with rotational seeding of vascular c onstructs is th e requirement that bioreactors are first disconnected from the adjacent perfusion circuit tubing in order for the scaffold to be continuously rotated. This additional exposure could compromise the sterility of the system, potentially limit in g the clinical application of the two stage seeding technique in which cell seeded grafts are cultivated in vitro prior to implantation. Alternating rotation was therefore implemented so that scaffolds could be freely rotated without disconnecting them fro m the circuitry ; t his resulted in no significant difference in seeding efficiency compared to continuous rotation. Overall, rotational seeding at high er speeds (100 RPH) resulted in the most homogenous circumferential distribution of EC on processed HUV gr afts. One stage seeding, in which EC are extracted and seeded onto prosthetic grafts at the same time that bypass surgery is perf ormed, appeals to clinicians for several reasons. 103 The additional costs associated with extended in vitro cultivatio n are avoided, and most importantly the graft can be quickly prepared for emergency vascular
114 reconstruction. However, several factors have limited the clinical success of this seeding technique. First, a limited number of venous or microvascular EC can be harvested, often resulting in implantation of vascular grafts with sub confluent cell densities 92 Second, EC allowed to atta ch, spread out, and mature for only minutes to hours are not resistant to hemodynamic shear forces, resulting in stripping off of seeded cells upon initiation of blood flow. For these reasons, two stage seeding may have more promise for future clinical app lications. The outcomes of the two stage seeding method can be improved upon by increasing the overall lumenal surface coverage by EC and increasing the resistance of seeded cells to mechanical stresses Here we noted that 5 days of maturation in the absen ce of significant mechanical forces improved the coverage of the HUV graft by EC. Gradual flow ramping to physiological pressure, circumferential stretch, and wall SS over a two day period allowed HUV grafts to retain high coverage by EC for an additional 3 days. This is consistent with other studies which showed improved EC retention when gradual flow ramping was employed compared to sudden onset of flow. 107,185 Using the aforementioned techniques, a functional nea rly confluent (~90% coverage) endothe lium could be developed on decellularized HUV allografts within 7 days of seeding. It should be noted that the cell source (primary HUVEC) and use of specialized culture media, which was supplemented with a cocktail of recombinant human growth factors that may not be clinically viable, resulted in significantly faster expansion (an average of 10 days from HUVEC extraction to graft seeding) than would likely be possible for EC derived from patients in need of bypass surge ry, who tend to be more elderly. However, the clinical application of rotational vascular graft seeding
115 would improve the efficiency of the seeding process, and significantly lower EC numbers may be required than was previously necessary. In conclusion, th e present study was undertaken to develop a functional, confluent neo endothelium on the lumenal surface of processed, implantable HUV allografts. A critical consideration in these investigations is the time frame required for these techniques to be deemed clinically acceptable. On one hand, implantation of a vascular graft in which seeded cells cannot withstand the aggressive hemodynamic forces of the arterial circulation, or are not functional is not likely to result in superior patency rates when compare d to currently available prosthetics. Alternatively, if months are required for expansion, maturation, and flow pre conditioning of harvested cells in order to produce a functional graft, this approach is unlikely to be considered as a feasible option by v ascular surgeons. It is therefore important to recognize that a balance must be struck between these two extremes for TEVG to reach clinical use.
116 Figure 6 1 Bioreactor system for development of TEVG A: Each HUV scaffold is housed in a custom desig ned glass bioreactor. The scaffold is affixed to adapter mandrels on each end which direct media flow through the lumen. Glass ports attached to the body allow media to enter/exit the ablumenal void space. B: This schematic shows the elements contained wit hin each flow circuit. A computer controlled peristaltic pump is used to draw culture media from a reservoir and drive it thro ugh the lumen of the scaffold. Three bioreactors were connected in series in this circuit to normalize the lumenal flow rate throu gh each scaffold. Transmural media flow is redirected to the lumenal circuit as pressure increases in the ablumenal void space.
117 Figure 6 2 Rotating seeding apparatus design This CAD drawing shows the seeding apparatus, which is attached to a rotating pump drive under computer control When bioreactors are attached, the axis of each HUV scaffold is parallel to the axis around which the device is rotated A: Oblique view showing the cylindrical seeding apparatus, fitted with 6 axially aligned HUV biorea ctors. B: End view demonstrating the arrangement of the bioreactors around the Not shown in this schematic are the tubing segments connecting the bioreactors in series.
118 Figure 6 3. Comparison of EC attachment after rotational o r static seeding approaches HUV scaffolds were seeded with EC overnight using various rotation methods then rinsed, fixed, and co stained with RP/DAPI to visualize cytoskeletal F actin and cell nuclei, respectively. Scaffolds were then cut open axially, i maged in 3 pre determined zones on the lumenal surface and analyzed in ImageJ. The mean c ell density and percent area coverage are reported in 3 zones (n=5 images each) along the graft circumference for each seeding method, ordered by number Zone 1: min imum; Zone 2: median; Zone 3: maximum. Asterisks indicate significant difference in means between seeding group within each zone (p<0.05).
119 Figure 6 4 EC growth characteristics in normal or dextran supplemented culture media HUV scaffolds were seeded with EC overnight using various rotation methods. Scaffolds were then rinsed, fixed, and co stained with RP/DAPI to visualize cell nuclei and cytoskeletal actin, respectively. Scaffolds were then cut open axially, imaged in 18 pre determined locations, and analyzed in ImageJ. Average cell density and percent area coverage are reported in 3 locations (n=6 images each) along the graft circumference.
120 Figure 6 5. Effects of maturation time on EC retention after exposure to flow ramping Decellularized HUV s caffolds seeded with EC under 100 RPH alternating rotation were allowed to mature for 1, 2, or 5 days before flow was steadily ramped up as described in (A) EC retention was then assessed as previously described using RP/DAPI co staining. Cell density (B) and percent area coverage (C ) were compared between initial seeding and various maturation times ; asterisks indicate significant differences in means (p<0.05) Five days of maturation resulted in a significant i ncrease in cell density (B) and percent area coverage (C ) when compared to initial seeding, indicating increased resistance to SS compared to two or five days of maturation. D : representative images of EC monolayers cultivated under each condition. Scale bar: 100 m
121 Table 6 1. Bioreactor conditio ns during flow preconditioning. Time After Seeding (days) Pulse Rate (pulses/min) Mean Shear Stress (dynes/cm 2 ) Minimum Pressure (mmHg) Maximum Pressure (mmHg) 0 5 1.5 0.3 2.80.4 9.01.6 5 5.1 0.9 1.21.5 9.61.1 6 40.5 7.3 27.91.3 54.14.0 7 10 80 14.5 79.63.5 111.112.7 Pressure values are presented as meanSD as measured downstream from each bioreactor (n=3 scaffolds per measurement). Flow ramping occurred from days 5 7, and the values shown were calculated at singular time points at days 5, 6, and 7, respectively.
122 Figure 6 6 EC monolayer development under 3D perfusion culture HUV scaffolds were seeded with EC overnight as described above at 100 RPH alternating rotation then cultured under low flow for five days to allow EC growth to confluence. At day five, the flow rate was progressively increased as shown in (A) until physiological wall SS levels were reached. At various time points, s caffolds were co stained with RP/DAPI and analyzed in ImageJ. Average cell densi ty (mean S.D.) and percent area coverage (mean+S.D.) over the ten day culture period are shown in (B) Representative images at high magnification (40x) show EC morphology at various time points. Scale bar: 50 m.
123 Figure 6 7 HUV scaffold morphology before and after end othelialization HUV scaffolds were seeded with EC and flow condit ioned as described in Figure 6 6 Scaffolds were then rinsed, fixed, and stained using Masson richrome to visualize scaffold morphology A continuous EC monolayer can be seen on the inner lumenal surface (arrow). Blue: collagen; red: cytoplasm ; black: cell nuclei. Scale bar: 100 m.
124 CHAPTER 7 CONCLUSIONS Clinical Significance The motivation for conducting the work presented in this dissertation stems from the critical need for functional c onduits to serve as vascular grafts. Despite the steady decline in the number of annual CABG surgeries performed in recent years 186 there is a consensus among clinicians that the outcomes in patients suffering f rom severe (three vessel or left main coronary artery disease ) are far superior when CABG is performed compared to PCI. 3 Prosthetic conduits in current use routinely fail to match the patency rates of small diameter autologous grafts 6,12,23 yet surgeons often have no alternative option and are forced to implant a vessel that will likely required repeat intervention. As reviewed in the Chapter 1, small diameter prosthetic graf ts typically fail due to thrombosis and anastomotic intimal hyperplasia 7 though both mechanisms can contribute to vein graft failure as well. 15 These complications are more pronounced in prosthetic grafts due to the innate material properties of synthetics in current use (ePTFE and PET). S pecifically, the propensity of these g rafts to allow plasma protein adsorption 13 and the lack of compliance 181 that results in disrupted hemodynamics through the graft and adjacent host vessels. 7 Prosthetic grafts were initially designed to withstand the mechanical stresses of the arterial circulation, yet remain completely non I t has become increasingly clear, however, that the appropriate interaction with the host is critical for small diameter vascular grafts to remain patent. T he design of a graft that is bio inert has been abandoned in favor of a graft that elicits an appropriate host re sponse The use of ex vivo derived biomaterials shows
125 significant promise in vascular reconstructive surgery due to the natural composition, structure, and capacity to be infiltrated and re populated by host cells. The ideal TEVG will be non thrombogenic; contractile; of appropriate size, structure, and mechanical properties; and exhibit favorable properties for vascular bypass surgery (e.g. suturability simplicity of handling ) 92,144,187 This dissertation was under taken in support of the development of a functional TEVG based on the use of the HUV as a scaffolding material Summary Thi s project specifically focused on the formation of a neo endothelium on the lumenal surface of decellularized HUV grafts in an effort to minimize early graft thrombosis as well as maintenance of seeded SMC in a contractile, non proliferative state. Previous efforts have focused on the development of a cell dense medial layer in HUV grafts through the use of directed nutrient gradients a nd appropriate mechanical stimulation 44,45 While the inclusion of a medial layer is critical for contractility and long term resistance to arterial pressures, i mplantation of SMC seeded HUV grafts without an endoth elium would likely result in acute failure due to either exposure of the lumenal surface to blood flow (thrombosis) or excessive SMC proliferation (intimal hyperplasia). Therefore the present work was undertaken to address the feasibility of in vitro endot helialization of processed HUV grafts. In Chapter 3 we first analyzed morphological features of the HUV scaffold after various decellularization treatments While all the methods used were effective at removing cells from the scaffold surface, the st ructu ral properties of the HUV we re greatly affected by processing technique T he use of surfactants to solubilize and extract cell ular components left an intact basement membrane and ECM structure most
126 reminiscent of the native HUV Given the acceptable EC bin ding properties and superior viability (compared to TX dHUV), SDS was selected as an optimal decellularizing reagent for the HUV. Chapter 4 details the design of a custom flow chamber for visualization of EC adaptation on the surface of vascular grafts Th e device was specifically designed to allow the lumenal surface of the HUV to be directly imaged while providing EC with a microenvironment more reminiscient of the native vasculature SDS d HUV scaffolds were seeded with EC within the flow chamber grown to confluence, and then adapted to physiological levels of SS using an effective pre conditioning strategy in which computer software was used to steadily ramp up the flow rate EC cultured on the lumenal surface of the HUV under SS exhibited high viabilit y that was similar to cells cultured under static conditions. Scanning electron micrographs showed these cells to be more flattened and spread out than static cultured cells. EC activated with TNF an inflammatory agonist, were able to maintain confluenc e while facilitating rolling HL 60 cell adhesion. That EC cultured on the HUV surface were capable of resisting physiological SS levels in the presence of a wound like activation state indicated that additional surface modifications, such as the use of fib rin glue, are not necessary for endothelialization of processed HUV grafts. A physiological perfusion model was developed in Chapter 5 that better represents the dynamic range of arterial blood flow. The adaptation of cultured EC to these variable SS patt erns was compared to adaptation to steady state SS, which is common in perfusion systems designed for either study of EC mechanotransduction events or the development of TEVG. Most significant in Chapter 5 was the finding that
127 EC cultured under variable SS patterns produced a significantly higher amount of NO than sister cells cultured under fixed SS; the increased NO production correlated with reduced leukocyte adhesion in a co culture assay. These findings suggest that TEVG may benefit from more dynamic ( e.g. real time changes in flow rate) perfusion conditions via the effects on EC physiology. Finally, Chapter 6 add ressed seeding tubular HUV grafts with EC In previous chapters, EC were seeded onto the HUV surface in two dimensional culture plates or flow chambers, in which the scaffold was fixed on a planar conformation. Seeding the entire lumenal surface of tubular grafts with EC was accomplished through axial rotation of the HUV during gravitational settling of EC out of suspension. Through the use of d extran to increase fluid viscosity, the temporal flow ramping approach designed in Chapter 4 was translated to a re create similar fluid SS on the lumenal surface of tubular HUV grafts. Additional maturation time (5 days) allowed EC to grow to confluence, which resulted in improved resistance to shear forces. The additional mechanical forces (physiological pressure and circumferential stretch) in this conformation necessitated that the ramp was extended to 2 days to prevent stripping away of EC from the lum enal surface. These results demonstrate the feasibility of EC seeding, maturation, and flow pre conditioning of implantable HUV grafts within a clinically acceptable time frame. Future Directions These investigations have focused on developing a functional neo endothelium the HUV graft The next most logical step is endothelialize HUV grafts in which a SMC dense medial layer has previously been engineered Significant culture time (~6 weeks) is required for SMC seeding, migration, and ECM remodeling of the HUV scaffold 44,45
128 and endothelialization should therefore be conducted after a cell dense medial layer has already been formed. The presence of SMC in HUV grafts may have significant e ffect s on EC adhesion, prolife ration and adaptation to mechanical stresses all of which will need to be assessed in detail. Analysis of SMC phenotype before and after endothelialization will be necessary to confirm that seeded EC can elicit a contractile (rather than synthetic) state The physiological SS model developed in Chapter 5 will next be translated to tubular HUV grafts. The inclusion of additional mechanical stimuli (pressure, circumferential strain) to the perfusion regime may result in other changes in EC gene expression, enzymatic activity, and paracrine interaction (both with peripheral blood cells and SMC) that will need to be assessed. This regime will also be applied to co cultured EC/SMC on the HUV graft to investigate the effects of SMC on EC adaptation to variable m echanical stimulation Understanding h ow dynamic perfusion conditions affect SMC (modulation of a contractile/synthetic state) in the presence or absence of EC will help to elucidate the role of the latter in transducing hemodynamic SS into physiological v ascular responses
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144 BIOGRAPHICAL SKETCH Joseph S. Uzarski began his research career in May 2006 as an undergraduate at Michigan Technological University, where he investigated the biological mechanisms of lymphatic vessel regeneration. In May 200 8, he graduated magna cum laude with a Bachelor of Science in biomedical engineering and a Minor in S panish from Michigan Technological University (Houghton, MI). Joseph began his graduate studies in August 2008 in the School of Chemical, Biological, and M aterials Engineering at the University of Oklahoma (Norman, OK) under the guidance of Dr. Peter McFetridge. In the fall of 2009, Dr. McFetridge accepted a faculty position in the J. Crayton Pruitt Family Department of Biomedical Engineering at the Universi ty of Florida (Gainesville, FL). To continue his doctoral research, Joseph transfer red to the University of Florida in January 2010 when he helped to set up a tissue engineering laboratory in the newly opened Biomedical Sciences Building. Joseph graduated with a Doctor of Philosophy in biomedical e ngineering in May 2013. His dissertation is focused on mechanical pre conditioning strategies for improved endothelialization of ex vivo derived vascular grafts.