1 MODULATED CELL FUNCTION DURING THE REMODELING EVENTS OF ENGINEERED VASCULAR TISSUES By ZEHRA TOSUN A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FO R THE DEGREE OF DOCTOR OF PHILOSOPHY UNIVERSITY OF FLORIDA 2012
2 2012 ZehraTosun
3 To my family and friends who supported me no matter what, during my graduate studies
4 ACKNOWLEDGMENTS I want to express my thanks to Dr. Peter McFetridge for his invaluable contributions to this project and his constant support during my graduate studies at the University of Florida. I would also like to thank Dr. Ben Keselowsky, Dr. MalisaSarntinoranon, and Dr. Edward S cott for their contributions as well as for their willingness to serve as members of my supervisory committee. I very much appreciate all the help and coopera tion provided by the BME staff. Furthermore, I would like to thank the National Institute of Health for funding the grant to carry out these investigations as well as my assistantship. I want to express my gratitude to my current and past McFetridge lab friends, Salma Amensag, Rita Issa, Cassie Juran, Marc Moore, Joe Uzarski, Aurore Van de Walle and Carolina Villegas for their help and their words of encouragement. I thank my family, whose support and love gave me the strength I needed many times throughout these years. A very special thanks goes to Pedro Diaz for his words of encouragement, h is advice, his company, and most importantly for his patience.
5 TABLE OF CONTENTS page ACKNOWLEDGMENTS ................................ ................................ ................................ .. 4 LIST OF TABLES ................................ ................................ ................................ ............ 8 LIST OF FIGURES ................................ ................................ ................................ .......... 9 ABSTRACT ................................ ................................ ................................ ................... 11 CHAPTER 1 GENERAL INT RODUCTION ................................ ................................ .................. 13 Blood Vessel Structure ................................ ................................ ........................... 13 Clinical Significance ................................ ................................ ................................ 13 Vascular Graft History Highligh ts ................................ ................................ ...... 15 Current Vascular Grafts ................................ ................................ .................... 15 Biological grafts ................................ ................................ .......................... 15 Synthetic grafts ................................ ................................ .......................... 17 Tissue Engineered Blood Vessels ................................ ................................ .......... 17 ScaffoldTypes ................................ ................................ ................................ ... 18 ECM based materials ................................ ................................ ................. 18 Biodegradable polymers ................................ ................................ ............ 21 Intraperitoneal grafts ................................ ................................ .................. 21 Scaffold free approaches ................................ ................................ ........... 21 Decellularized tissues ................................ ................................ ................ 22 SMC Plasticity, and Its Relevance to Vascular Tissue Engineering ................. 24 Endothelial Cell Vascular Smooth Muscle Cell Interactions .......................... 26 Effects of Mechanical Stimuli on Vascular Cell/Tissue Growth ......................... 28 Cell Seeding Techniques ................................ ................................ .................. 32 Alternate Cell Sources ................................ ................................ ...................... 33 Human Umbilical Cord Structure ................................ ................................ ............ 33 Acellular Human Umbilic al Vein as an Engineered Vascular Scaffold .................... 34 2 MATERIALS AND METHODS ................................ ................................ ................ 55 Experimental Methods ................................ ................................ ............................ 55 Isolation of vSMCs from Human Umbilical Artery ................................ ............. 55 Human U mbilical Vein (HUV) Dissection ................................ .......................... 56 Decellularization ................................ ................................ ............................... 57 Bioreactors Assembly and Pretreatment ................................ .......................... 57 Preparation of Collagen Gels ................................ ................................ ........... 58 Seeding Protocol ................................ ................................ .............................. 59 Analytical Methods ................................ ................................ ................................ .. 59 Construct Characterization ................................ ................................ ............... 59
6 DNA Quantification ................................ ................................ ........................... 60 Cell Metabolic Activity ................................ ................................ ...................... 60 Histology ................................ ................................ ................................ ........... 60 Uniaxial Tensile Testing ................................ ................................ ................... 61 Statistical Analysis ................................ ................................ ............................ 61 3 HYDROGEL CELL SEEDING AROUND VASCULAR TISSUES ............................ 67 Materia ls and Methods ................................ ................................ ............................ 67 Experimental Methods ................................ ................................ ...................... 67 Hydrogel Contraction Time ................................ ................................ ............... 67 Perfusion Flow Profile Design ................................ ................................ .......... 67 Results ................................ ................................ ................................ .................... 68 Hydrogel Contract ion Time ................................ ................................ ............... 68 Perfusion Flow Profile Design ................................ ................................ .......... 68 Summary ................................ ................................ ................................ ................ 69 4 IMPROVED RECELLULARIZATION OF ENGINEERED SCAFFOLD USING DIRECTED TRANSPORT GRADIENTS ................................ ................................ 79 Materials and Methods ................................ ................................ ............................ 80 Experimental Methods ................................ ................................ ...................... 80 Perfusion Flow Profile Design ................................ ................................ .......... 81 Analytical Methods ................................ ................................ ........................... 81 Burst Pressure ................................ ................................ ................................ .. 81 Calculation of Scaffold Permeability ................................ ................................ 82 Results ................................ ................................ ................................ .................... 82 Vascular Construct Morphology and Thickness ................................ ............... 82 Scaffold Cellularity ................................ ................................ ............................ 83 Cell Distribution ................................ ................................ ................................ 83 Biomechanical Properties ................................ ................................ ................. 84 Permeability ................................ ................................ ................................ ...... 85 Summary ................................ ................................ ................................ ................ 86 5 CONTROLLED VASCULAR CELL FUNCTION BY USING MO DELED MECHANICAL STIMULATION ................................ ................................ ............... 98 Materials and Methods ................................ ................................ ............................ 99 Measurement of Contraction and Relaxation Forces ................................ ..... 100 Scanning Electron Microscopy ................................ ................................ ....... 100 Results ................................ ................................ ................................ .................. 100 Morphology of the Scaffolds ................................ ................................ ........... 100 Biomechanical Properties ................................ ................................ ............... 101 Pharmacological Response of the Constructs ................................ ................ 102 Scaffold Microstructure and Cell Distribution ................................ .................. 102 Summary ................................ ................................ ................................ .............. 103
7 6 RELATIVE GENE EXPRESSION ANALYSIS OF ENGINEERED SCAFFOLDSIN RESPONSE TO PHYSIOLOGICAL FREQUENCY S TIMULI ..... 113 Materials and Methods ................................ ................................ .......................... 115 RNA Extraction and Reverse Transcription ................................ .................... 115 Real Time PCR ................................ ................................ .............................. 115 Results ................................ ................................ ................................ .................. 116 Summary ................................ ................................ ................................ .............. 117 7 CONCLUSIONS AND FUTURE WORK ................................ ............................... 131 Conclusions ................................ ................................ ................................ .......... 131 Future Work ................................ ................................ ................................ .......... 135 Combinations of Physiological and Constant Frequencies as Mechanical Stimuli ................................ ................................ ................................ ......... 135 In Vivo Acellular Implantation Study ................................ ............................... 135 In Vitro Blood Vessel Maturation Followed by an In Vivo Study ..................... 136 Different Approaches for Scaffold Preparation ................................ ............... 136 Exploring Different Cell Sources ................................ ................................ ..... 136 vSMC Heterogeneity within the Scaffold ................................ ........................ 137 Economical Perspectives ................................ ................................ ............... 137 APPENDIX: A COMPOSITE SWNT COLLAGEN MAT RIX: CHARACTERIZATION AND PRELIMINARY ASSESSMENT AS A CONDUCTIVE PERIPHERAL NERVE REGENERATION MATRIX ................................ ................................ ..... 139 LIST OF REFERENCES ................................ ................................ ............................. 141 BIOGRAPHICAL SKETCH ................................ ................................ .......................... 165
8 LIST OF TABLES Table page 1 1 Scaffolds used for tissue engineered blood vessel studies to date ..................... 40 1 2 Summary of in vitro and in vivo vascular tissue engineering studies .................. 42 1 3 vSMC seeding techniques developed over the years ................................ ......... 53 1 4 Main cell sources that are used for vascular tissue engineering ........................ 54 4 1 Acronym, and the description of each bioreactor condition. ............................... 91 6 1 Selected vSMC marker genes and their protein function. ................................ 122 6 2 List of selected extracellular matrix genes and their general description .......... 123 6 3 Comparison of the adhesion receptors ................................ ............................. 124 6 4 Pubmed reference sequence number of each gene used in this study ............ 1 25 6 5 F old regulation and their standard error values of each gen e in PF 12h samples. ................................ ................................ ................................ ........... 126 6 6 F old regulation and their standard error values of each gene in PF 24h samples. ................................ ................................ ................................ ........... 127
9 LIST OF FIGURES Figure page 1 1 Cross section of a porcine carot id artery ................................ ........................... 37 1 2 Relative research interest and total number of publications. .............................. 38 1 3 Cross section of a human umbilical cord. ................................ ........................... 39 2 1 Autodissection methodology. ................................ ................................ .............. 63 2 2 De cellularized human umbilical vein. ................................ ................................ .. 63 2 3 Dynamic cell culture setup.. ................................ ................................ ................ 64 2 4 Hydrogel contraction around the HUV scaffold.. ................................ ................. 65 2 5 Characterization of the constructs ................................ ................................ ...... 66 3 1 The sequence of events during cell seeding that required optimization ............. 71 3 2 Flow conditioning profile. ................................ ................................ .................... 72 3 3 Flow conditioning profile.. ................................ ................................ ................... 73 3 4 Hydroge l contraction time prior to flow ................................ ............................... 74 3 5 After 2 hours of lumenal flow only.. ................................ ................................ ..... 75 3 6 After 2 ho urs of only ablumenal flow only ................................ ........................... 75 3 7 After 24 hours of cell seeding ................................ ................................ ............ 76 3 8 Comparison of the flow condition. ................................ ................................ ....... 77 3 9 Proliferation and migration of the cells to the lumen. ................................ .......... 78 4 1 Schematic drawing ofseeding and hydrogel compaction within the perfusion bioreactors ................................ ................................ ................................ ......... 91 4 2 Representative images of vascular constructs. ................................ .................. 92 4 3 T ensile stress strain relationship.. ................................ ................................ ..... 93 4 4 Cell density and metabolic activity ................................ ................................ ...... 94 4 5 Fluorescence RNA labeling ................................ ................................ ................ 95
10 4 6 Ultimate tensile strength (UTS) and Calc ulated Burst pressure values .............. 96 4 7 strain at max load (B) .................. 97 5 1 Acronym, the description, an d frequency profile of each bioreactor condition. 107 5 2 R epresentative images of vascular constructs after 6 weeks. .......................... 107 5 3 Comp arison of representative load (B) and % strain at max stress (C) and Ultimate tensile strength (D).. ......... 108 5 4 strain at maximum stress over the physiological range ................................ ................................ .................... 109 5 5 Maximum contractile and relaxation force of the constructs. ............................ 109 5 6 SEM images showing longitudinal cross section of the scaffolds ..................... 110 5 7 Microscopic images of constructs. Scale bar correspond to 200 m. ............... 111 5 8 Cell density and metabolic activity Cell density (above left) of cellular HUV constructs at week 6.. ................................ ................................ ....................... 112 6 1 Schematic representation of relative expression of cytoskeletal, co ntractile and transcription factor vSMC marker genes ................................ .................... 128 6 2 Schematic representation of relative expression of extracellular matrix and cellular remodeling marker genes. ................................ ................................ .... 129 6 3 Schematic representation of relative expression of cell cell adhesion, cell matrix adhesion related marker genes. ................................ ............................ 130 7 1 Progression of each ex perimental chapter ................................ ....................... 138
11 Abstract of Dissertation Presented to the Graduate School of the University of Florida in Partial Fulfillment of theRequirements for the Degree of Doctor of Philosophy MODULATED CELL FU NCTION DURING THE REMODELING EVENTS OF ENGINEERED VASCULAR TISSUES By ZehraTosun August 2012 Chair: Peter McFetridge Major: Biomedical Engineering T he use of ex vivo derived materials designed as implant scaffolds have increased significantly d u ring th e last decade. This is particularly so in the area of regenerative medici ne, or tissue engineering, in which the natural chemical and biomechanical properties have been shown to be advant ageous. By focusing on events that occur during remodeling processes, our objective was to detail progressive changes to further our understanding of these processes. A perfusion bioreactor system and acellular human umbilical veins were used as a model three dimensional vascular scaffold i n which human myofibroblasts or vascular smooth muscle cells were seeded and cultured under defined pulsatile flow conditions. Cells were seeded to the ablumenal periphery of an ex vivo derived vascular scaffold using a collagen/hydrogel cell delivery system. The aim of the present work was to modulate nutrient gradients using perfusion bioreactors in order to create a d irected growth environment in which cells can fully populate the scaffold and remain vi able over extended periods. T he following aim was to focus on the short term variat ion of pulse rate Short term variation such as that
12 caused by elevations in respiratory activity, as occur s during exercise was included a produce a more physiologically relevant modeled pulsatile flow and make comparisons to constant pulse pulsatile fl ow These investigations have shown that by modulating perfusion culture conditions during early remodeling events, cell distribution and extracellular matrix remodeling can be regulated resulting in pronounced effects on construct biomechanics. By furthe r understanding the details of graft development through control of transport and of mechanical stimuli conditions that modulate the cellular microenvironment, we are progressively moving toward the goal of developing clinically functional small diameter vascular grafts.
13 CHAPTER 1 GENERAL INTRODUCTION Blood Vessel Structure All blood vessels are composed of three layers : tunica intima, tunica media and tunica adventitia [1 ] [2 ] .i) The t unica intima layer is composed of a fine lining of endothelial cells and basement membrane s Endothelial cells are aligne d towards the blood flow direction and are attached to the base ment membrane that is approximately 100 nm in thickness [3 ] ii) The t unica media is composed of vascular s mooth muscle and e lastic connective tissue that consists of elastic sheets of fi brils and collagenous fibrils. Their c omposition varies according to the type of the artery. Tunica media is responsible for vasoconstriction/vasodilation and the mechanical st rength of the blood vessel. iii) Tunica adventitia is composed of the connecting collagen tissue surrounding the blood vessel. Cross section of a porcine carotid artery is shown in Figure 1 1 as an example. The t hickness of each layer varies depending on t he purpose of the vessel. Tunica adventitia is the thickest layer in the vei ns; it links the veins to the external tissues, whereas tunica media is the thickest layer in the arteries responsible f o r hemodynamic changes and the ability to withstand higher pressures compared to veins. Clinical Significance Pathological conditions and traumatic injuries may affect the function of blood vessels. Vascular disease, in particular the occlusion of coronary and peripheral blood vessels, is one of the most significa nt clinical issues facing western societies resulting in excessively high morbidity and mortality rates. Some cases require replacement or
14 bypass surgery. In the past years s ten ts are frequently used as a non invasive method for the treatment of vascular c logging. Stents have been improved in the past years from bare metal stents to drug eluting stents with improvements in the geometric design. However, stent insertion still has drawback s such as restenosis, thrombosis and inflammatory response s [4 ] Improvements s uch as antiproliferative drugs and anticoagulants improved the sh ort term patency by reducing endothelial injury. However anti thrombogenic modification of the stent surfaces is still under study that reduces late stent th rombosis. Improved long term results of stents will increase demand to this minimally invasive therapy [5 ] Vascular bypass grafts are designed to repair terminally occluded blood vessels by co nstructing a conduit that bypasses the diseased segment. The success of the latter procedure has been limited to large diameter vessels (>6 mm) such as the thoracic and abdominal aorta [ 6 ] There are more c omplications associated with s mall diameter vascular grafts. Therefore, small diameter blood vessel substitutes (<6mm) are still needed for coronary artery or bel ow the knee bypasses, as well as arteriovenous dialysis shunts [7 ] [8 ] Every year more than 500,000 coronary arteries are replaced with grafts due to atherosclerotic diseases, which are t he leading cause of death in the United States [9 11 ] In addition to atherosclerosis, diabetes can lead to cardiovascular complications [12 ] In particular, this group of patients may not have arteries or veins suitable for replacement [13 ] In addition, after an autovenous bypass, 10 20% of t he patients need reoperation [14 ]
15 Vascular Graft History Highlights The first vascu lar anastomosis (connection of blood vessels) was performed by Eck in 1877 [15 ] C.Gluck made the first experimental attempt at venous replacement in 1898, where an autovein was inserted into a carotid artery. Carrel and Guthrie optimized the vascular anastomo sis transplantation technique in 1912 and won the first Nobel Prize in physiology and medicine [16 ] However until the 1950s vascular prostheses were largely unsuccessful due to thrombosis and g raft preservation techniques. A femoropopliteal bypass with a reversed sa phenous vein by Kunlin in 1948 initiated a very succes sful era that has lasted until the presen t day [17 ] In 1951 Charles Dubost in France performed the f irst successful resection of a n aneurysm with a cadaver homograft [18 ] An artificial vascular prosthesis as an aortic replacement with a Vinyon [19 ] The first successful coronary artery bypass was performed by Kolesow (1967 ) [20 ] Current Vascular Grafts Commonly used co nduits for replacement of small diameter blood vessels include biological grafts (autografts, allografts, and xenografts), and synthe tic grafts [21 ] Biological g rafts Current biological grafts can be grouped in three categories. Autografts : A graft from one point to another of the same individual's body. Allografts: A graft transplanted from same species. Xenografts: A graft of tissue from different species than the recipient. Autologous grafts are the gold standard as a bypass graft. For coronary artery byp ass grafts the internal mammary artery and the radial artery are the primary choice s [22 ] For peripheral bypass grafts, the saphenous vein is preferred [23 ] However, the disadvantages of the autograft technique are the donor site mor bidity. In
16 addition, the autograft of choice may not be available due to prior vessel harvest, unsuitability or varicosities [24 ] Vein grafts are the most durable grafts for replacement. However when veins are implanted into arterial circulation, they become stiffer at arterial pressures and become exposed to forces that venous system s do not experience [25 ] Gradual dilation and pressurization exp oses the vein to altered forces such as shear stress Due to circumferential wall deformation these grafts may thicken overtime Vein wall damage has been as sociated with inflammatio n, smooth muscle cell migration an d proliferation that may ultimately cause intimal hyperplasia. Allografts lost popularity after the 1960s due to graft deterioration, and preservation problems. In the last years, preservation me thods of allograft s have gone through many modificati ons, an d improved preservation method s make allografts possible to use more frequently in limb th reatening situations, in repeated surgery or if infections are occurred [26 ] I nfection associ ated with bioprosthetic grafts can happen in up to 3% of implanted materials [27 ] [28 ] In a systematic early review by Johnson et al. venous and arterial allografts overall performance were compared. Even though there were no direct comparisons between the different clinical trials, the authors concluded that human umbilical vein allografts may have superior performance [29 ] [30 ] Human umb ilical vein graft history will be discussed in detail later in this chapter. Xenografts are generally used as vascular access grafts for hemodialysis The immune response to these materials has been found to b e higher than to allograft tissue By using a bo vi ne mesenteric vein bioprosthesis for hemodialysis access a consid erable reduction in thrombosis, and infection was observed compared to
17 synthetic implants [31 ] Decellularized no n c r o sslinked grafts will be discussed in the Blood section. Synthetic grafts Pol y(ethylene terephthalate) (Dacron) polytetrafluoroethylene (PTFE, Teflon, Gore Tex), and polyurethane are the three most widely used vascular synthetic graft s up to date [32 ] As small diameter vascular graft materials, they have high failure rates. For above the knee femoropopliteal bypass su rgery, prosthetic bypass grafts (PTFE) exhibited 39% patency rates, whereas autovenous bypasses have a 74% patency rate [32 ] The main reasons for poor biological functionality of synthetic grafts are the thrombogenicity of the artificial surface and their lack of compliance [33 ] Synthetic grafts do not endothelize spontaneously except in the anastomosis region. In order to improve patency Zilla et al. first made efforts move towards a thrombogenic surface by anchoring endothelial cells to synthetic materials [34 ] The following studies by many research groups enhanced the cell seeding method to synthetic materials that improved the patency of these g rafts to 66 % afte r 7 years of implantation [35 ] This approach can be seen where tissue engineer ing principles are applied clinically. The drawback in this method is that endotheli al cell extraction and expansion takes a long time which means it cannot be used in an emergency situation [36 ] Tissue Engineered Blood Vessels Early attempts were focused on the con struction of synthetic materials as a non thrombogenic, passive conduit [37 ] There has been a great effort to maintain endothelial cells at a non thrombotic and non inflammatory phenotype when seeded o n synthetic grafts [38 42 ] However, low success rates during clinical studies of such grafts shifted the attention of the scientific community towards creating a functional blood vessel.
18 Weinberg and Bell used collagen gels seeded with smooth muscle cells, fibroblasts and endothelial cells to produce a comp letely biological blood vessel [43 ] As a result of their creative work, starting in 1980, tissue engineering has become an alternative option to creat e small diameter blood vessels [44 ] To quantify the trend over the years, r elative research interest and total number of pu blications were computed using GoPubmed.com t ools. issue enginee r ed keyword s. Compared with 2000 the number o f research articles published in 2011 increased by up to 4 times as shown in Figure 1 2 The use of autologous cells on scaffolds may allow tissue engineered blood vessels t o acquire or assimilate structural and functional properties similar to autologous grafts. However, engineering a vessel that possesses native vascular composition, mechanics, and physiological function is a technical challenge. Current challenges include thrombogenicity, vasoactivity, compliance mismatch, and intimal hyperplasia [45 49 ] These limitations are rela ted to the manipulation of the endothelial cell and smooth muscle cell function wit hin an in vitro environment [50 ] [51 ] Each of these limitations is still a challenge and must be addressed to achieve clinical success. Scaffold T ypes Scaffolds th at are used for TEBV studies are summarized in Table 1 1 The T able 1 1 is grouped into three categories with respect t o scaffold type : ECM based materials, biodegradable polymers and decellularized tissues E ach description and its advantages and disadvantages are given with references. ECM based materials Collagen is the most abundant protein in the body. 30% of body protein is collagen. In m ost of the tissues, collagen is responsible for tensile strength. The
19 predominant collagen type is collagen type I that represents 90% of the collagen content [52 ] This most abundant structural protein was first tissue engineered vascular scaffold used As discussed earlier the first tissue engineered blood vess el was made by coculturing vSMCs within collagen gels and EC on the lumen [43 ] Howeve r low mechanical p roperties of these materials le d to improvement of the materials by crosslinking or using synthetic materials as support material s Sihibaahi and Matsuda used a s i milar method to Weinberg et al. and incorporated a knitted Dacron s upport and dermatan sulfate within the collagen gel [53 ] Contracting th e hydrogel around the rigid tubular mandrel further increased the mechanical propert ies of these material s but still bu rst in pressures less than 120 mmHg [44 ] Another improvement was using magnetic alignment during co llagen fibril formation [54 ] by Tranquillo et al Dynamic mechanical stimulation improved the strength of the collagen gels significantly and will be discussed furth er in detail in the C hapter 2 The main obstacle while using collagen gels is that cells exhibit little ECM synthesis when entrapped within collagen gels. Fibrin is plays a major role in wound healing. I t is the major structural protein in blood clots [55 ] Studies of fibroblasts entrapped in fibrin gel have suggested an increased amount of collagen synthesis in comparison with fibroblasts entrapped within collage n gel, and there is evidence of collagen fibrils and other ECMs accumulating in the fibrin network [56 ] O ne ECM protein is crucial for compliance: E lastin Elastin occurs as sheets in the medial layer. Even though the collagen gels with additives resulted with relatively higher mechanical properties, t cells were entrapped into collage n gels. Fibrin gels yielded enhanced elastin production during maturation of tissue engineered constructs [57 ] Fibrin is also incorporated into
2 0 biodegradable polymers to improve their biological properties of the biodegradable polymers, and to increase the mechanical properties of fibrin [55 ] Low c ompliance of synthetic materials is one of the biggest causes of failure along with thrombogenicity [58 ] In order to have a material that is stiff but distensible an elastin component is essential [59 ] Elastin prevents dynamic tissue creep and recoils back to the original configuration after the load is rel eas ed. L ow (physiological range) strain r eflects the behavior of the ela stic properties before stiffer collagen fiber s contributes. However, p urified elastin scaffolds do not have the sufficient strength to resist arterial pressures In a study by Wise et al. s caffolds were fabricated by the electronspinning technique incorporating elastin with various molecules such as collag en and biodegradable polymers [60 ] A p orous elastomeric scaffold with matching mechanical properties and mass transfer supported the coculture of baboon carotid artery SMCs and EPCs from baboon peripheral blood [59 ] Hyaluronic Acid (HA) is another polymer that is actively involved in wound healing processes. The microbia l fermentation technique made it possible for HA to be produced in large quantities [61 ] As a vascular graft HA was investigated by Rath et al. [62 ] Hyaluronic acid incorporated scaffolds are able to promote elastin biosynthesis [63 ] Chitosan is a derivative of chitin with a similar str ucture to glycoaminoglycans. I t is another biomaterial o btained from shrimps and crabs. In addition to its biocompatibility and bio degradability, it has low toxic and inflammatory properties. However, chitosan itself is not biologically sufficient material as a scaffolding material fo r vascular tis sue engineering applications [64 ] Chitosan needs to be modified
21 with biomimetic molecules in order to progress as a vascular tissue engineering scaff old [65 ] Biodegradable polymers Food and Drug Administration ( FDA ) cleared biodegradable polyme rs have received significant attention. However their main disadvantage is their degraded fragments initiating inflammation. Poly(glycolic acid) Poly(lactic acid) caprolactone) caprolactone) Polyurethan e poly(lactic co glycolic acid), and Polydioxanone are some of the most common examples [66 ] In a study by Shinoka et al a utologous bone marrow cells were seeded onto a PGA scaffold to replace the pulmonary artery N o graft mortality was observed over a 6 year period with 25 patients [67 ] [68 ] Intraperitoneal grafts Intraperitoneal grafts were an innovative approach first reported by Campbell et al. [69 ] A silastic tubing was inserted into the peritoneal cavit y of rabbits. When the implant w as harvested, the tissue around the tube consisted of a blood vessel like structure where the inn er lining was mesothelial cells (intima), myofibroblasts in the medial layer and collagenous advential layer. Scaffold free approaches A scaffold et al. us ing f ibroblast or smooth muscle cells sheets that were rolled ove r a ma ndrel without any scaffold [70 ] [71 ] These grafts were successfully used for hemodialysis access. The h emodialysis access graft produced by cell sheet multilayer method was implanted in 10 patients The patency rate was 78 % at the first month and 60% at the 6 th month [72 ]
22 Decel lularized t issues Decellularized constructs have the advantage of preserving extracellular matrix components. The goal of a decellularization protocol is to efficiently remove all cellular and nuclear material while minimizing any adv erse effect on the composition, biological activity, and mechanical integrity of the remaining ECM. T he se grafts are aimed to be decellularize d and recellularize d with patients own cells in vitro to be used as a completely biological blood vessel. One c oncern with decellularized tissues is the remnant s of c ellular antigens recognized as foreign by the host that induce s an inflammatory response after implantation. However, extra cellular matrix proteins are preserved and tolerated even across different spe cies Small intestinal submucosa (SIS), the porcine carotid artery the bovine carotid artery, and the canine carotid artery are widely studied. One of the most frequently characterized decellularized materials in the literature is the small intestinal sub mucos a. In a n ex vivo ovine arteriovenous shunt model by Shwartz et al. carotid arteries, jugular veins, and small intestinal submucosa based grafts were tested. SIS grafts demonstrated physical properties between those of carotid arteries and jugular vein s [73 ] Zhou et al. compa red heparin coated decellulariz ed canine ve ssels and progenitor endothelial cell seeded canine arteries [74 ] After 3 months e ndothelial cell seeded canine vessels exhibit ed greater patency in cell donor dogs. A n innovative approach was proposed by Niklason et a l : Vascular grafts with similar mechanical properties to native human blood vessels were produced culturing PGA and vascular smooth muscle cel ls in a bioreactor. Ce llular material was re moved by decellularization. Next, these grafts were implanted in dogs as a peripheral and coronary artery bypass
23 graft. Grafts showed excellent patency and resisted dilatation, calcification, and intimal hyperplasia [75 ] [76 ] For feasibility of decellularized tissues as vascular grafts tissue processing and especially method of decellulariza tion are important. Decellularization aims to remove non structural xeno or allogenic cellular components from tissue C ellular components elicit an immune response by lysing the cellular membrane and disrupt binding between cells and ECM components. Ove r the years many methods have been developed with various agents used for decellularization. The e ffectiveness of decellularizat ion relies upon tissue type and properties such as tissue cellularity, thickness and content. The decellularization methods can be mainly divided into three groups: i) C hemical, ii) E nzymatic and iii) P hysical. The most common decellularization protocols utilize a combination of these three treatments [77 79 ] The chemical agents include acid and bases hypotonic and hypertonic solutions non ionic detergents, ionic detergents, z witterionic detergents, and solvents. The enz ymatic ag ents include nucleases and trypsin. The physical treatments include freezing and thawing, pressure, perfusion, and agitation. Ev ery decellularization technique will disrupt ECM to a degree while removing cellular components. Ph ysical treatments mainly resu lt in the lysis of the cell membrane, while chemical treatments solubilize cyto p lasmic components and lipid lip i d interactions. Enzymatic treatments are mainly used to c atalyze the hydrolysis of ribonucleotide and deoxyribonucleotide chains Soluble protei ns and lipids remaining within tissue after decellularization have been related to graft rejection, and undesired secondary effects such as calcification [80 82 ]
24 SMC Plasticity a nd Its Relevance t o V ascular T issue E ngineering Tissue engineered scaffolds act as a conduit to provide structural support and recreate a natural environment where cells can adhere and proliferate However, regeneration of cell den se and fully functional vascular media and intima within the engineered scaffolds has been problematic. Important parameters related to the lack of the tunica media are: i) the inability to form layer s of functional circumferenti ally aligned smooth muscle ii) non native extracellular matrix composition, iii) inappropriate mechanical properties (requiring at least 1700 mm Hg burst pressure, and to be elastic with appropriate compliance 4 8%/mmHg 10 2 ) needs to be demonstrated. Creating a functional media layer is crucial for normal vasoactive behavior, as well as mechanical properties of the blood vessel such as tensile strength, compliance, and burst pressure [9 ] [48 ] [83 86 ] Obstacles arise from the manipulation of the smooth muscle cell phenotype dur ing the maturation period in vitro Therefore, vascular smoot h muscle cell plasticity is a key aspect [51 ] vSMCs should be able to migrate and proliferate to create a smooth muscle cell layer while secreting ECM components, and to remodel the scaffold. After constructing the medial cell layer, a shift to a contractile phenotype is important to possess vasoactivity and inhibit vSMC growth; which may cause narrowing of the blood vessel. The m edial layer is composed of vascular smooth muscle and elastic connective tissue that consists of elastic sheets or fibrils and collagenous fibrils. Tunica media is responsible for vasoconstriction/vasodilation and the mechanical strength of the blood vess el. vSMCs carry this function in its contractile machinery that re volves around ion channels, agonist receptors, contractile and cytoskeletal proteins, and intracellular signal transduction molecules [1 ]
25 vSMCs are plastic so that they can adjust to a variety of differe nt physiological and pathological conditions. The primary function of vSMC is contraction. v SMCs proliferate in extremely low rates and do not produce extracellular matrix proteins during vascular remodeling, and vascular injury induces cell proliferation and matrix production. Proliferation of vSMC in disease lesions is well studied. D uring these vascular remodeling processes, continuous proliferation of vSMCs can lead to intimal hyperplasia, atherosclerosis, and posta n gioplasty restenosis [87 ] vSMC phenotype shift from contractile to synthetic state can be characterized in a variety of ways [88 ] Changes in morphology, proliferation, and migra tion are closely associated with a phenotype shift. In addition to these broad changes some marker proteins have been identified to characterize the changes in the molecular level [89 ] The vSMC phenotype shift is a complex system that can be a ffected by many factors such as biochemical factors, mecha nical stimuli, extracellular matrix composition, endothelial cell influences, neurotransmitters, and vascular injury. In concert with what cell s sense in terms of extracellular space, integrins modify the intracellular pathways leading to changes in gene e xpression It is important to characterize the molecular signature of vSMC at a given differentiation stage [89 ] [90 ] However, many vSMC proteins are also being expressed in many other c actin is also the most abundant protein in vSMCs (up to 40% of total protein) and is also expressed in myofibroblasts during tumor progression [91 ] [92 ] Howe ver, proteins that are used as contractile marker s are downregulated when a phenotype shift occurs towards a more proliferative stage.
26 However expression levels of each gene vary along the synthetic to contractile spectrum 1 Endot helial Cell Vascular Smooth Muscle C ell I nteractions An intact endothelium layer in vivo serves as an antithrombotic/coagulant surface between blood vessels and the bloodstream. It prevents the direct contact of platelets and coagulation factors with the subendothelial surface and provides a nonthrombogenic surface for blood flow. In addition, endothelial cells selectively control the permeability of the blood vessel wall by regulating soluble substance s and macromolecules. Permeability can change during inflammation, and this mechanism is crucial for vascular remodeling processes [93 ] Permeability is regulated by cell cell contacts between adjacent endothelial cells [5 ] [94 ] Endothelial cells exhibit biochemical responses that regulate vascular smooth muscle. Nitric oxide (NO), endothelin (ET 1), and prostacyclin (PGI2) are the most important biochemical substances secreted from the endothelium. Endothelin is a vasoconstrictor where as nitric oxide and prostacyclin are vasodilators [94 ] Nitric Oxide inhibits the activation of g rowth factors and induces anti migratory/antiproliferative effects on underlying SMCs. E ndothelial cells are capable of affecting smooth muscle cell function through growth factors, such as PDGF and FGF. Furthermore, synergistically, they regulate vascular tone by vaso dilator/constrictor substances [95 ] Dy sfunction of endothelial cells during these pathologic processes or disease states leads to platelet adhesion, inflammation, and vSMC proliferation and migration. Rensen et al [84 ] presented expression levels of genes associated with a particular SMC phenotype The chart that explai ns these relationships be found here : http://www.ncbi.nlm.nih.gov/pmc/articles/PMC1847757/figure/F2/
27 EC in a pathological state are often characterized by reduced NO production, with an increase in endothelin. In addition, specific EC marker proteins can be used to characterize ECs, such as PECAM1 (platelet EC adhesion molecule 1), CD34,and von Willebrand factor (VWF) [96 ] on to the lumenal surface. However, e onto the luminal surface of the synthetic grafts has resulted in problems with the maintenance of the EC layer on limitations tissue engineers have studied the hemodynamic forces that blood vessels experience: shear stress and cyclic stretch [97 ] Changes o ccurring in biomechanical stimuli can directly affect EC cell morphology and function such as actin stress fibers formation [98 ] [99 ] At lower shear stress es stone morphology conformation. In addition to morphological changes, endothelial cells exhibit biochemical responses to defined flow stimuli. A variety of coculture sy stems have been utilized to study vSMC EC interactions. These studies include direct coculture [100 ] transmembrane culture [101 ] [102 ] [103 ] perfusion bioreactor systems [104 ] [105 ] culture of vSMCs within EC conditioned media The presence and absence of ECs have had controversial results on vSMC behavior because they are mainly dependent on the differentiation level of both cell types, culture length, and limitations of the culture system. Endothelial cell conditioned mediums initially inhibited growth, and delayed the phenotypic changes from contractile to synthetic states that occur in vitro overtime [106 ] Confluent EC layers wher e non confluent EC layers did
28 not Heparin has similar vSMC growth inhibitory effects to similar to EC cell c onditioned medium [107 ] In a direct coculture study by Lavender et a l. cell culture media formulation was optimized for the direct EC vSMC coculture systems, and a confluent and adherent endothelial layer was able to be cultured on vSMC cells [108 ] In another direct coculture study by Wallace et al. ECs were attached to quiescent SMCs, withstanding shear stresses up to 300 dyncm within these co culture sy stems with confluent endothelia maintained over 30 days [109 ] Remond et al. proposed that endothelial cell plasmin ogen activator inhibitor type I secreted from ECs reduces flow induced SMC migration in coculture systems [110 ] B ioreactors induce mechanical forces as well as improve the retention and function of ECs in vitro properties are important parameters. Williams et al. seeded vSMCs into porous poly(glycolic acid) tubular scaffolds and cultured them under pulsatile shear stress within a bioreactor fo r 25 days [105 ] Two conditions were compared: ECs were seeded either at day 10 or day 23. After 15 days of EC SMC co c ulture, cell proliferation was increased and collagen and proteoglyc an deposition was down regulated compared to 2 day cultures. Another study b y Yazdani et al. compared different preconditioning protocols and used these preconditioned grafts in an ex vivo shunt model. High steady shear stress and cyclic shear stress exhibited the highest degree of EC adherence and alignment compared to low steady shear stress [111 ] Effects o f Mechanical Stimuli o n Vascular Cell/Tissue Growth An important goal has been to engineer the vascular grafts with similar mechanical properties to natural vessels and as such a significant amount of research has been conducted to assess cult ure systems and environmental conditions that may improve
29 graft performance. In order to recapitulate these conditions, bioreactors and perfusion flow systems aim to emulate the in vivo environment in order to enhance tissue regeneration by modulating cell ular phenotype, improving mass transport limitations, and long term sterility [112 115 ] To overcome these limitations, many scientific groups have developed biorea ctors with conditions similar to the physiological vessel environment [49 [112 [113 [116 119 ] This effort resulted in increased ECM production [120 ] increased cell density [121 ] and greater contractility of SMCs compared to the grafts cultured under static conditions [41 ] [122 ] In a study by Seliktar et al ., exposure of rat SMC to 10% strain at 1Hz frequency led to the actin expression and SMC prolifer ation [123 ] Niklason et al. showed that under pulsatile flow (165 pulses per minute) alignment of v SMCs and extracellular matrix production can be enhanced [49 ] After 8 weeks of vessel culture, high burst strength is obtained. Using a pulsed perfusion bioreactor, the systems designed by Hoerstrup et al. have shown that constructs cultured for 1 month resulted in significant improvements of burst strength, from 50 mmHg for static controls to 326 mmHg for conditioned samples [124 ] et al. developed a well defined three layered vessel structure by adding concen tric layers of cellular sheets that require twelve weeks of preparation and culture This three layered ve ssel structure displayed burst strength over 2000 mmHg [71 ] More recently Syedain et al investigated the effects of circumferential strain amplitude ranging from 2.5% to 20% either constantly or incrementally to evalua te the effects on fibrin rem odeling. It was found that the incremental circumferential strain resulted in higher ultimate tensile strength (2295 467 kPa,50% increase) when compared to c onstant circumferential
30 stress. This demonstrates the complex nature of vessel remodeling and adaptation during regenerative processes [125 ] While we are beginning to understand the complex nature of tissue remodeling and broad effects of mechanical stimulation, little is currently known of how these early remodeling events influence the biomechanical behavior of these developing tissues. ECM remodeling, an d the factors that influence the remodeling process, are clearly important and require more detaile d investigation, particularly with tissue based materials where the primary ECM macromolecules provide the bulk of the structural support. With ex vivo mater ials modifies the ECM through balanced degradative and synthetic pathways that alter the tissues mechanical properties over time. A balance between ECM synthesis and degradation is clearly a critical fac tor, particularly with arterial grafts where ECM failure may have fatal consequences. Similarly, Stegemann and Nerem [126 ] have shown that embedding SMCs in 3D collagen scaffolds, without mechanical loading, leads to actin expression relative to monolayer cultures. When subjected to a low load environment (i.e. floating collagen m atrix) cells were inclined to become quiescent over time and as the mechanical load was increased, actin fiber formation. Mechanical stimulation has generally been shown to benefit c onstruct properties. However in vitro extended culture times may not result in beneficial remodeling. The investigations conducted by Seliktar et al. showed that while mechanically stimulated collagen constructs seeded with human aortic SMCs displayed enh anced ultimate tensile stress and increased stiffness values over shorter culture
31 periods (4 days), constructs subjected to prolonged mechanical stimulation (8 days) deteriorated. Based on their findings, it was suggested that the variation in collagen com paction and consequent mechanical properties is the result of a cell mediated remodeling that is largely regulated by mechanical stimulation [7 ] [123 ] [127 ] While these time points a re reduced, there is some correlation with a change in stiffness that is possibly attributed to a similar mechanism where the bal ance between remodeling degradation and synthesis may not always be in balance. This again demonstrates the importance of mechanical and biochemical cues that determine the cell ph enotype [126 ] In vivo graft dilatation may occur as early as the initial adaptation per iod where overexpression of MMP s by vascular smooth muscle cells and inflammatory cells can result in degr adation of elastin and collagen [128 ] To some extent weakening of the scaffold is expected during this early remodeling p eriod as the newly synthesized and deposited ECM is in a less structured orientation. It is then reorganized as the remodeling process matures and is naturally cross linked to enhance the strength of the collagen fibers and ECM as a whole. However a secon dary mechanism may explain a partial ECM deterioration that is more specific to acellular ( ex vivo ) scaffolds. MMP s that are typically associated closely with the cell membrane m ay have a more distant effect on ECM mechanics by freely diffusing through the acellular scaffold to more distant zones of th e scaffold. T his may explain an overall weakening that may occur, especially if cell migration is limited. Several other research groups attempted to develop a functional small diameter blood vessel replacemen t. However, an engineered blood vessel with native artery parameters has not been achieved. Thus, it is necessary to optimize the current small
32 diameter blood vessel development strategies for urgent need s in current clinical practice. Engineering vessels that possess similar biomechanical properties to natural arteries is an important research objective. Scaffold remodeling, the key process in effective regeneration, is a complex process that is clearly influenced by multiple factors including mechanical s timulation, scaffold composition and structure, cell ty pe, culture chemistry, and many other more subtle elements. A m ismatch between the graft and the host artery causes a variety of effects that can result in an abnormal endothelial and smooth muscle cel l phenotype that results in pathologies such as thrombosis and intimal hyperplasia, which are primary causes of graft failure [33 ] [51 ] A critical, yet poorly understood, aspect of the remodeling processes is the biomechanical transformation that occurs throughout the regenerative process. It has been shown that when material mechanical properties are mis mat ched to the host vasculature (at the implantation site) it is likely to lead to of graft failure, and as such is a cr itical parameter to assess [58 ] [129 ] From this perspective ex vivo based materials have an advantage over currently used synthetic conduits that are typically stiffer and less compliant than native vessels. There are comprehensive revie w publications on tissue engineered blood vessels [130 138 ] .The dynamic vSMC bioreactor culture systems are briefly summarized at the end of this chapter in T a b le 1 2 All studies were id entified by Pubmed with the following keywords: bioreactors, vascular tissue engineering, preconditioning. Cell S eeding T echniques A variety of cell seeding techniques have been developed over the years. Cost effective reliable and efficient culture techniques are needed. The p ros and cons of
33 various cell seed ing techniques are summarized i n T able 1 2 The unique cell seeding system developed in our lab will be discussed in Chapter 3 in detail. Alternate Cell Sources The choice of optimum cell source has been one of the main obstacles in the field. The r eplicative capacity of adult cells has also significantly limited the clinical applications of small diameter tissue engineered blood vessels In addition older cells deposit les s ECM matrix than younger cells. Autologous human stem and progenitor cells from a range of sources have received significant attention due to th eir potential to differentiate these cells into EC or SM C lineages. The m ain cell sources that are used for vas cular tissue engineering are summarized in T able 1 4 Some o f the cell lines used to date are adult endothelial vascular smooth muscle, and fibroblasts, as well as embryonic and adult stem cells The i deal cell source s should be functional in vitro non i mmunogenic, and easy to extract and expand. Human Umbilical Cord Structure The u mbilical cord links the fetus to the placenta. An Umbilical cord s physical dimensions change with factor s such as gestational age. G enerally at term of gestation, the umbilica l cord is about 50 60 cm long, and 1 2 cm in diameter [139 ] [140 ] It consists of one vein, and two arteries as shown in Figure 1 3 These vessels are surrounded by epithelium The ro le of the umbilical arteries is to return blood to the placenta with a high pulsatility The u mbilic al vein carries oxygenated blood to the fetus from the placenta with a smaller pulse [140 144 ] The u mbilical vein contains an elastic subintimal layer. Type I, III, IV, and V are the collagen types with type I and type III being the most abundant in the umbilical vein.
34 Conversely, umbilical arteries do not contain an elastic subintimal layer, b ut possess little elastin in their medial layer [144 ] In recent years, several investigators published protocols for isolating MSCs from [145 ] Umbilical cord tissue from this perspective is a very important source of MSCs MSCs can be isolated from umbilical cord with high cell numbers relati vely quickly and easily accessible source compared to other MSC sources the umbilical vein and artery from external forces such as compression and torsion attributed to fetal movements. Hyaluronic acid is the most abundant GAG, and constitutes approximately 70% of GAG content. Also, several peptide growth facto rs were detected in WJ, such as fibroblast growth factor (FGF), insulin like growth factor (IGF I), platelet der It was reported that r times more collagen, and twice more glycoaminoglycans compared to umbilical arteries. The presence of the growth factors might help these cells to synthesize these molecules. [142 ] [143 ] Acellular Human Umbilical Vein a s a n Engineered V ascul ar S caffold A significant effort has been made to remove imm uno genic components from these scaffold s A promising alternative to xenogenic decellularized vascular scaffolds is the human umbilical cord vein and artery. the human umbilical vein (HUV) has been used as a glutaraldehyde cross linked allograft showing improvements in graft patency compared to many synthetic materials [145 151 ] Several changes were made in the manufacturing
35 process in the past years, including tanning with glutaraldehyde, incorporating an additional polyester fiber mesh and finally, storing grafts in a 50 % ethanol sol ution. Similar to other aldehyde treated materials the cross linking process used on current clinical versions of this material inhibit s cell migration and subsequent remodeling due to the inability of cells to degrade these induced bonds. Another study s howed that after 2 years of implantation, a significant percent of the patients that received glutaralde hyde tanned HUV presented aneury smal degeneration [152 ] Daniel et al (2005) developed an automated methodology to rapidly and uni formly dissect the HUV directly from the umbilical cord [153 ] Using this automated dissection procedure the HUV scaffold used in these investigations was decellularized to fo rm an acellular scaffold without artificially induced stabilization such as formaldehyde or glutaraldehyde treatments. The perceived advantage of this approach is a reduced immune response allowing adhered cells (seeded in vitro or naturally in vivo ) to po sitively remodel the scaffold without the inflammatory response rapidly degrading the tissue, which would otherwise lead to graft dilation and possible aneurysm formation. As such remodeling can occur in a more stable fashion leading to a fully cellular and functional graft. This approach made it possible to dissect HUV with preserved mechanical properties and uniform wall thickness a significant improvement compared to manual dissection In addition, the burst pressure and compliance of the material a re preserved with this automated method. C ell incorporation experiments in these grafts showed evidence that the acellular HUV is conducive to cell attachment and smooth muscle cells can migrate and remodel the extracellular matrix throughout the scaffold [154 ]
36 HUV as an acellular scaffold is advantageous because it is already in the shape of a blood vessel and consists of natural ECM components which give appropriate mechanic al properties and biocompatibility. Histological images after decellularization showed that HUV presents the three layers of a natural vessel even after cells h ave been removed.
37 Figure 1 1 Cross section of a porcine carotid artery The b ottom layer is the intima followed by the medial layer The outer layer is the adventitia followed by the connective tissue.
38 Figure 1 2. Relative research interest and t otal number of publications, engineer blood vessels
39 Figure 1 3. Cross sec tion of a human umbilical cord Scale bar corresponds to 200 micron s
40 Table 1 1 Scaffolds used for tissue engineered blood vessel studies to date Difficulties as a clinical product Specific Commen ts Biological Gels The low mechanical strength, structural stability Collagen [44 ] [53 ] [98 ] [155 ] [156 ] L ow antigenicity, L ow inflamm atory and cytotoxic responses, D esirable biological and hematolo gical properties. Fibrin [56 ] [57 ] [157 161 ] Act ively involved in wound healing Enhanced ECM production compared to collagen gels Hyaluronic acid Acid [63 ] [16 2 ] Higher elastin synthesis, Can be produced in large quantities, C an altered regulation of angiogenesis Chitosan [64 ] [163 [164 ] Similar structure to glycoaminoglycans. Needs to be modified with biomimetic peptides Decellularized tissues The remnants after decellularization within the tissue may trigger i mmune response Allogenic tissues [153 ] [154 ] [165 ] [166 ] Umbilical veins or arteries: Excellent potential for cellular integration and preservation of the mechanical properties of the native blood vessels. Xenograft s [73 ] [167 ] [168 ] [169 ] R isk of trans mission of nonhuman animal pathogens to humans Readily Available Tissue Engineered Grafts [75 ] [76 ] [170 ] A brand new innovative approach to decellularize tissue engineered blood vessels Still l acks elastin. Biocompatible in a canine mode w hen seeded with endothelial cell s prior to implantation Cell synthe sized ECM [71 ] [72 ] [138 ] [171 ] [172 ] [173 ] [174 ] [175 ] Long in vitro culture times First tissue engineered product in the market. Used as a hemodialysis access graft Intraperi oneal grafts [69 [176 [177 ] O pening the peritoneal cavity has the risk of damage F aster than in vitro approaches with minimized risk of immune response
41 Table 1 1. Continued. Difficulties as a clinical product Specific Comments Polymer Based scaffolds Low compliance, Fragments from degradation eliciting inflammation, and immune response. PGA [49 ] [178 ] [179 ] [180 ] [181 ] Good biocompatibility, M ost widely used, S implest polymer structure PLA [182 ] [183 ] More hydrophobic than PGA, O ften mixed with PGA for various applications PCL [184 ] [185 ] Slow degradation rate T unable mechanical properties Polyurethane [45 ] [186 ] [187 ] Elastic, low thrombogenicity Polydioxanone [188 ] [189 ] Bioabsorbable flexib le
42 Ta ble 1 2 Summary of in vitro and in vivo vascular tissue engineering studies Authors Year Pubmed id Scaffold type Cell type Culture Method Outcome Badylak et al. [190 ] 1989 27394 01 (SIS) Autologous cells In vivo canine study (infrarenal aorta) 100% graft patency for up to 52 weeks Kim et al [191 ] 1998 10099177 PGA Rat aortic vSMCs Static More uniform distribution of SMCs adherent to the matrices were obtained with dynamic versus static cell seeding methods L'Heureux et al [71 ] 1998 943841 0 Scaffold free approach Umbilical cord v SMCs and EC's In vitro maturation followed by in vivo canine study (femoral artery) First completely biological TEBV to display a burst strength comparable to that of human vessels. Niklason et al. [49 ] 1999 10205057 PGA Porcine carotid artery v SMCs and EC's Pulsatile flow bioreactor followed by in vivo porcine study 100% patency at 4 weeks for ( Preconditioned gra ft ), non preconditioned vessels occluded at 3 weeks Campbell et al [69 ] 1999 10590244 Silastic tubing Autologous cells In vivo rat and rabbit study: peri tonea l cavity implantation first s tudy to demonstrate peri tonea l cavi ty as a bioreactor develop blood vessel at any length and diameter from host cells Huynh et al. [156 ] 1999 10545913 Intestinal col lagen layer derived from SIS Autologous cells In vivo rabbit study (carotid artery) Within three months after implantation, the grafts were remodeled into cellularized vessels that responded to vasoactive agents. Shum Tim et al [192 ] 1999 10617020 PGA Ovine carotid v SMCs EC's and fibroblatss Passive cell seeding followed by in vivo ovine model Evaluated in vivo a new scaffold based on a new copol ymer of PGA and PHA. Bader et al. [193 ] 2000 10919568 P orcine aorta Human saphenous vein v SMCs and EC's Perfusio n bioreactor Decellularized animal tissues may eventually lead to the engineering of vessels immunologically acceptable to the host using a relatively short preparation period of 2 3 weeks.
43 Table 1 2. Continued. Authors Year Pubmed id Scaffold type Ce ll type Culture Method Outcome Seliktar et al. [123 ] 2000 10870892 type I collagen gel Rat aortic vSMCs Cylic strain bioreactor Dynamic mechanical conditioning during tissue culture and proper biochemical environment improves the mecha nical strength and histological organization of vascular constructs. Watanab e et al. [180 ] 2001 11506732 PGA canine femoral v SMCs 7 days static culture followed by in vivo canine model 100 % patency at the end of 6 month Niklason et al. [194 ] 2001 11241137 PGA Bovine aorta EC's v SMCs Pulsatile flow bioreactor Vessel morphology and mechanical characteristics improved at 8 weeks Seliktar et al. [127 ] 2001 11791675 C ollagen Human aortic v SMCs Cyclic mechanical strain 10% cyclic radia l distention for 4 days, resulted in an overall increase in the production of MMP 2, also mechanical strength and material modulus significantly increased. Hoerstrup et al. [124 ] 2001 11423291 Nonwoven polyglycolic acid mesh Ovine carotid myofibroblasts and EC's Pulsatile flow bioreactor The mechanical properties o f the grafts resulted with physiological burst strength and suture retention strength appropriate for surgical implantation. Hoerstrup et al. [195 ] 2002 12118802 Nonwoven polyglycolic acid mesh Human umbilical cord progenitor Ecs, and vSMCs Pulsatile flow bioreactor Ti ssue strength characteristics s imilar to human pulmonary artery. Yu et al. [196 ] 2003 12947277 PTFE EC,SMC(both Rabbit jugular v ein) Rabbit aorta shunt Retention rate of EC @ 1 hour is 65% and 1 day (51%), EC/SMC @1 hour (98%), and 1 day (90%)EC retention on PTFE grafts in vivo is improved if seeded over a layer of SMC. Naito et al. [197 ] 2003 12579118 PLLA/PGA EC, SMC (peripheral vein) Static seeding followed by in vivo human study 100% patency 4 month
44 Table 1 2. Con tinued. Authors Year Pubmed id Scaffold type Cell type Culture Method Outcome Nasseri et al. [198 ] 2003 12740091 polyglycolic acid/poly 4 hydroxybutyrat e Ovine carotid myofibroblasts Rotational bioreactor Dynamic rotational seeding and culturing in a hybridization oven is an easy, effective, and reliable method to deliver and culture vascular myofibroblasts onto tubular polymer scaffolds. Stegemann et al. [126 ] 2003 12723680 Collagen type I Rabbit aortic vSMCs Circumferential strain controlled bioreactor Cell phenotype can be mod ulated in engineered blood vessels by applying selected combinations of biochemical and mechanical stimuli Williams et al. [199 ] 2004 15265311 PGA Bovine aorti c vSMCs and EC's Pulsatile flow bioreactor Bioreactor supports sequential seeding of vSMCs and EC's. Pulsatile flow leads to ECM toward the development of engineered vascular constructs. Chue et al. [176 ] 2004 15071455 Silicon tubing Autologous cells Peritoneal and pleural cavities Peritoneal and ple ural cavities of large animals can function as bioreactors to grow vascular grafts. Bagineid et al. [200 ] 2004 15032735 Collagen gel on a polyester graft Porci ne aortic vSMCs and EC's Pulsatile flow bioreactor Low shear stress preconditioning promotes the development of a functional and stable endothelium on an SMC collagen matrix. Opitz et al. [201 ] 2004 15008369 P 4 HB scaffolds Ovine aortic vSMCs and Ecs Pulsatile flow bioreactor Dynamic constructs displayed synthesis, DNA content and vSMC marker expression compared to static conditions Shin'oka et al. [68 ] 2004 15373223 PLLA/PGA Human bone mononuclear cells Hum an pulmonary artey implantation 100% patency 32 month
45 Table 1 2. Continued. Authors Year Pubmed id Scaffold type Cell type Culture Method Outcome Williams et al. [105 ] 2005 16060532 PGA Bovine aortic EC's and vSMCs Pulsatile flow bioreactor After 25 days of culture, 15 day EC co culture constructs have a more uniform cell distribution across the construct thickness a nd SMC express a more contractile phenotype compared to 2 day EC co culture Jeong et al. [202 ] 2005 15482828 PLCL collagen gel Rabbit aortic vSMCs Pulsatile flow bioreactor Enhanced vSMCs proliferation and alig nment, collagen production compared to static condition. L'Heureux et al. [70 ] 2006 16491087 Scaffold free approach Human adult skin fibroblasts In vivo rat, canine, primate study C ellular TEBV was achieved with 85 % patency at day 225 LafLamme et al [203 ] 2006 16968167 Scaffold free approach Human umbilical vein EC's, vSMCs In vitro maturation Contraction in the construct via endothelin was s imilar to native artery Hoerstrup et al. [204 ] 2006 16820566 PGA Ovine carotid artery and jugular vein EC's and vSMCs Pulsa tile bioreactor followed by in vivo sheep study No evidence of thrombus, calcification, stenosis, or aneurysm more than 100 weeks. There was a significant increase in diameter by 30% and length by 45%. Schmidt et al. [205 ] 2006 16 996955 PGA Human umbilical cord derived progenitor EC's and myofibroblasts Pulsatile flow bioreactor A three layered tissue architecture and functional endothelia similar to native blood vessels can be successfully generated from human umbilical cord proge nitor cells. Butcher et al. [206 ] 2006 16806457 Collagen hydrogel Rat aortic vSMCs Cyclic equibiaxial strain vSMCs become more synthetic in comparison to static controls. Isenberg et al. [160 ] 2006 16783653 Fibrin gel Rat aortic vSMCs Pulsatile Flow Loop Design for EC seeding ECs possess adhesion strength sufficient to withstand physiological shear stress an d maintain a normal phenotype.
46 Table 1 2. Continued. Authors Year Pubmed id Scaffold type Cell type Culture Method Outcome Risberg et al. [178 ] 2006 16674296 PGA Human saphenous vein EC's and vSMCs Pulsatile flow bioreactor Similar macroscopic appearance to native vessels with no visible evidence of the original PGA. After 6 weeks, lower gene expression of SMC specific markers, such as actin, calde smon, and vimentin compared to that of native arteries. Buttafocco et al. [207 ] 2006 16649175 P(DLLA co TMC) and collagen Human umbilical artery vSMCs Pulsatile flow bioreactor After 7 da ys of dynamic culture vSMCs were homogeneously distributed throughout the constructs, resulting five times stronger and stiffer compared to the noncultured scaffolds McFetridge et al. [113 ] 2006 15672794 Porcine carotid artery Human umbilical artery EC's and vSMCs Pulsatile flow bioreactor Increased of EC retention from 5.1 to 634 cells/mm2. Seeding VSMCs as sheets resulting cells migrating to the medial/adventitial boundary. vascular perfusion system serves as a useful tool to analyze cell adhesion and retention by allowing controlled manipulation of seeding and perfusion conditions. Abiliez et al. [208 ] 2006 16542683 Matrigel Mouse embryonic stem cells Pulsatile flow bioreactor Undifferentiated mESC's can be grown in 3D and under pulsatile conditions. Engbers et al. [209 ] 2006 16343614 C rosslinked i nsoluble type I collagen and elastin Human umbilical artery vSMCs Pulsatile flow bioreactor Higher vSMC proliferation/distribution throughout the scaffolds as well as higher collagen expression compared to static conditions.
47 Table 1 2. Continued. Author s Year Pubmed id Scaffold type Cell type Culture Method Outcome Buttafocco et al. [210 ] 2006 16289328 C rosslinked insoluble type I collagen and elastin Human umbilical artery vSMCs Pulsatile flow bioreactor After 14 days a more uniform distribution of vSMCs and partially oriented collagen fibers compared to static conditions. McFetridge et al. [168 ] 2007 17885337 Porc ine carotid artery Human umbilical artery vSMCs Pulsatile flow bioreactor Developed a perfusion system to allow growth and proliferation of vascular constructs over extended culture periods. Matrix remodeling enzymes were assessed to investigate if nutrie nts or other essential supplements were unable to diffuse further into the scaffold, or the cells were not synthesizing the necessary enzymes to degrade the scaffold. Boccafoshi et al. [155 ] 2007 17457943 type I collagen gel Porcine aortic EC's and vSMCs Rotating device Custom made rotating device provides an adequate environment for cell growth. Future investigation needs to be done to enhance the mechanical properties of the scaffold. Jeong et al. [211 ] 2007 17112581 Marine collagen and PLGA fibers EC and VSMCs from rabbit aorta Pulsatile flow bioreactor Enhanced cellular alignment, and upregulated vSMC markers compared to static cultures, Fi gallo et al. [212 ] 2007 17269690 Hyaluronic acid Fibroblasts were isolated from human derma Pulsatile flow biore actor Lamination techniques allowed the fabrication of both planar and tubular multilayer scaffolds. Multilayer scaffolds were also tested in dynamic culture conditions.
48 Table 1 2. Continued. Authors Year Pubmed id Scaffold type Cell type Culture Method Outcome Hann et al. [117 ] 2007 17180465 PEG hydrogel 10T 1/2 mouse smooth muscle progenitor cells Pulsatile flow bioreactor Constructs subjected to mechanical conditioning had significantly higher collagen levels, and improved biomechanical properties compared to static conditions. Iwasaki et al. [213 ] 2008 18824769 PGA sheets Bovine aortic vSMCs and EC's Pulsatile flow bioreactor Similar elastic characteristics as native arteries were obtained Tschoeke et al. [214 ] 2008 18684200 Fibrin gel with PVDF mesh Ovine carotid myofibroblasts Pulsatile f low bioreactor Mean suture retenti on strength of the graft tissue was 6.3 N and the burst strength was 236 mm Hg within 14 days Syedain et al. [125 ] 2008 18436647 Fibrin gels Porcine valve interstitial cells Circumferential strain controlled bioreactor Incremental circumferential strain stimulated constructs had superior properties compared to constant circumferential strain Arrigoni et al. [162 ] 2008 18383121 Hyaluronic acid Porcine aortic vSMCs Rotating versus perfusion bioreac t or Data indicate that the rotating bioreactor is mo re efficient than the perfusion bioreactor Zhu et al. [215 ] 2008 18377984 Porcine carotid artery Human umbilical cord EPC's, and vSMCs Bioreactor culture followed by in vivo rat model Blood vessel xeno transplantation caused rejection; but A20 transfected tissue e ngineered blood vessels remained open for 6 months postoperatively. Gong et al. [181 ] 2008 18199698 PGA Adult human bone marrow MSCs Cyclic strain bioreactor hMSCs can serve as a ne w cell source of SMCs in vessel engineering. O'Cearbh aill et al. [216 ] 2008 18194813 Silicon tubing Adult human bone marrow MSCs Pulsatile flow bioreactor MSCs reorientate parallel with direction of flow and have adapted their morphology to be similar to that of ECs. Gene expression results show the cells exhibit greater levels of SMC associated markers.
49 Table 1 2. Continued. Autho rs Year Pubmed id Scaffold type Cell type Culture Method Outcome Xu et al. [119 ] 2008 18155136 PGA Canine carotid artery vSMCs Pulsatile flow bioreactor Dynamic constructs exhibited great mechanical properties compared to static culture Konig et al. [217 ] 2008 19111338 Scaffold free aproach Adult fibroblasts from skin In vitro maturation followed by human in vivo shunt model No differences at burst pressure and suture retention compared to IMA and after six months similar compliance values to IMA Li et al. [218 ] 2009 19812476 PCLA Canine bone marrow MSCs Pulsatile flow bioreactor Scaffolds were no ntoxic to cells and were favorable for the growth and migration of MSCs. Yazdani et al. [169 ] 2009 19290806 Porcine carotid artery Rat aortic vSMCs Pulsatile flow bioreactor Bioreactor preconditioning accelerates the formation of a significant muscular layer on decellularized scaffolds, in particular on adventitia denuded scaffolds. Yang et al. [219 ] 2009 19240529 Canine carotid artery Canine saphenous vein EC's and vSMCs Pulsatile flow bioreactor followed by in vivo canine study After 6 months of implantation constructs had 100% patency and similar histologica l organization to native artery grafts Zhang et al. [220 ] 2009 19232717 Electrospun silk fibroin scaffolds Human coronary artery vSMCs, and human aortic EC's Pulsatile flow bior eactor Constructs subjected to mechanical conditioning had significantly higher cellular alignment/proliferation, and improved biomechanical properties compared to static conditions. Tschoeke et al. [55 ] 2009 19125650 Fibrin gel with PVDF mesh Human aortic vSMCs Pulsatile flow bioreactor Significant increase in both cell number and collagen content over 21 days with burst pressures over 460 mmHg
50 Table 1 2. Co ntinued. Authors Year Pubmed id Scaffold type Cell type Culture Method Outcome Wang et al. [221 ] 2010 19819545 PGA Human a dipose derived stem cells Pulsatile flow bioreactor hASCs can serve as a new cell source for SMCs in blood vessel engineering Song et al. [177 ] 2010 20692791 Silicon tubes Au tologous cells Mouse peritoneal cavity The mouse peritoneal cavity of mice has the ability to function like a bioreactor to generate bio engineered tissues. Sharifpoor et al. [187 ] 2010 21463894 Polyurethane Embryonic rat aortic vSMCs Cyclic mechanical strain bioreactor Pore interconnectivity was optimized while maintaining its mechanical integrity Kelm et al. [222 ] 2010 20223267 Scaffold free approach Fibroblasts and EC's from umbilical cord Pulsatile flow bioreactor Self assembled tissues composed of fibroblasts displayed significantly accelerated E CM formation compared to monolayer cell sheets. Harris et al. [223 ] 2011 19959190 Acellular human saphenous vein Human adipose derived stem cells Pulsatile flow bioreactor Human adipose der ived stem cells exhibit variable expression of SMC molecular markers after differentiation, exhibit a contractile phenotype after differentiation proliferate on a vascular graft scaffold. Tosun et al. [154 ] 2011 21872418 Acellular human umbilical vein Immortalized SMCs Pulsatile flow bioreactor Dynamic studies enhanced cell proliferation, migration, and scaffolds become more elastic overtime. Wang et al. [224 ] 2011 21671960 Collagen Fibrin gel/PLGA Adipose derived stem cells Pulsatile flow bioreactor Dynamic Stimulation and growth facto rs induced transformation of stem cells to vSMCs Sharifpoor et al. [187 ] 2011 21463894 Polyurethane Human coronary artery smooth muscle cells Cyclic mechanical strain bioreactor Better cellular organization/proliferation, and increase in expression of vSMC marker genes compared to static constructs
51 Table 1 2. Continued. Authors Year Pubmed id Scaffold type Cell type Culture Method Outcom e Ferdous et al. [225 ] 2011 21284997 Type I Co llagen gel Human aortic vSMCs human aortic valvular interstitial cells (HAVIC) Cyclic mechanical strain bioreactor Constructs were cultured in regular vs osteogenic media as a model to investigate cell mediated differences in early markers of calcificatio n. Lee et al. [226 ] 2011 21282618 PGS Adult primary baboon vSMCs Pulsatile flow bioreactor Constructs contained mature elastin equivalent to 19% of the native arterie, withstand up to 200 mmHg burst pressure and exhibited compliance comparable to native arteries. Heine et al. [227 ] 2011 21189071 Acellular porcine carotid artery Human EC's and vSMCs Pulsatile flow bioreactor ECs a nd SMCs with lower CD 31 actin signaling than controls. Constructs exhibited vasomotoricity, but in a lower rate compared to native arteries. Neff et al. [228 ] 2011 20934837 Acellular porcine carotid artery Ovine autologous EC's from circulating progenitor cell s, ovine femoral artery vSMCs Pulsatile flow bioreactor followed by sheep in vivo study Enhanced in vivo wall maturation and contractile function of TEBVs coseeded with autologous SMCs and ECs compared with EC seeding alone. Syedain et al. [229 ] 2011 20934214 Fibrin scaffold Human dermal fibroblasts Pulsatile flow bioreactor After 9 we eks of bioreactor culture, grafts were extensively remodeled into cir cumferentially aligned tubes of extracellular matrix with burst pressures in the range of 1400 1600 mmHg and compliance comparable to native arteries Song et al. [23 0 ] 2011 20807005 PTMC Human umbilical vein vSMCs Pulsatile flow bioreactor Better cellular organization/proliferation, and vSMC markers compared to static constructs
52 Table 1 2. Continued. Authors Year Pubmed id Scaffold type Cell type Culture Method Outcome Park et al. [231 ] 2011 21943800 PLCL collagen gel Rabbit aortic vSMCs, EC's Pulsatile flow bioreactor Engineered vessels acquired similar elastin contents as native arteries. Dahan et al. [167 ] 2012 21919798 Acellular porcine carotid artery Human umbilical artery EC's, and vSMCs Pulsatile flow bioreactor Dynamic co culturing led to a higher infiltration, migration and prolifer ation of vSMCs toward the media compared to static cultures
53 Table 1 3 vSMC seeding techniq ues developed over the years Seeding technique Description Specific Comments Static Seeding [36 ] [232 ] [233 ] Pipetting cell suspension directly onto lumen or albumen Most widely used, Least efficiency Used in vivo Dynamic Seeding Rotational systems [186 ] [198 ] Graft is rotated in a cell suspension Reduced cell seeding in lower cell concentrations Longer seeding times Vacuum systems [234 ] [235 ] Vacuum pressure forces the cell suspension to the graft No in vivo studies that discusses the viability of this approach Cell sheet based [71 ] [236 ] Two different techniques were developed: 1) Cells were seeded on temperature re sponsive polymers a sheet of cells easily with this technique 2) Cell sheets are grown in a regular culture flask, and wrapped around a cylindrical mandrel Used in vivo Prolonged culture times Magnetic [237 ] Nanoparticles are placed inside of the cells, and magnetic forces are used for cell infiltration 90% cell seeding efficiency. Cell viability and adverse effect of the nanoparticles should be investigated Photopolymerizable gel based [238 ] Photopolymerization of a cell suspension in an aqueous polymer solution UV light exposure is a concern for in vivo applications
54 Table 1 4 Main ce ll sources that are used for vascular tissue engineering [66 ] [239 242 ] Cell type Source Specific Comments Autologous SMCs Small vessel biopsies. Limited proliferation potential Immortalization, or insertion of human telomorase reverse transcriptase subunit Safety when implanted in vivo is a concern May activate oncogenes Allogeneic cells Immun e rejection Embryonic Stem cells Pluripotent Capable of differentiati ng into every somatic cell type Ethical concerns Limited availability Adult stem cells Mesenchymal Multipotent Derived mainly from bone marrow, and adipose tissue Easy to isolate, and expand compare to embryonic cells Capable of differentiating to vascular cells Progenitor cells Vascular progenitor cells may exist in the circulation, bone marrow, vasc ulature, and peripheral tissues A viable cell source Advantage for patient specific therapies Evaluated in human subjects Induced pluripotent stem cells Transfected with stem cell associated genes reprogrammed to become pluripotent stem cells Exhibits the ess ential characteristics of hESCs Potential for the future clinical applications Potential tumor risks Low throughput
55 CHAPTER 2 MATERIALS AND METHOD S Experimental M ethods M ethods that were common for the Chapter 3,4,5, and 6 w ill be described here. The methods section has been divided in two groups, experimental methods an d analytic al methods. Specific m ethods are described separately in the mate rials and methods section of the associated chapter. Isolation o f v SMCs f rom Human Umbilical Artery Cell isolation protocol was previously described [243 ] Briefly, vascular smooth muscle cells were isolated from the human umbilical artery. Umbilical arteries were isolated from the cord by removing the surrounding extensive tissue. Dissected tissue was washed with HBSS mediu m. The artery was kept moist during following steps. The artery was cut open l ongitudinally and the endothelial layer was removed by scraping the intimal layer. Then the arter y was cut into about 2mm cubes u s ing a scalpel blade and tweezers. Before the cubes were placed into the 25 cm 2 ce ll culture flask, a drop of cell culture medium was added on each cube. About 15 20 cubes were placed inside of each 25 cm 2 cell culture flask, with 2.5 ml cell culture medium directly in the vertical bottom of the flask. Cell cu lture flasks were kept upright (vertically) f or 2 h in the incubator. Then flask s were placed horizontally in their proper orientation. Cell culture media was changed every two days. cultures after two to three we eks. After three weeks, explanted tissues were removed from the culture flask by using a Pasteur pipette. Cells were cultured in low glucose (2 s medium (Hyclone, Carlsbad, Calif), supplemented with
56 10% fetal bovine serum, 1% L glutamine, and 1% penic illin streptomy cin (Invitrogen) with 5% CO 2 at 37C. Cells were frozen at p=3 to be used in the experiments. Cell freezing medium contained 5% Dimethyl sulfoxide ( DMSO), 25% fetal bovine serum, and low glucose (2 s medium (Hyclone, Carlsbad Calif). Human U mbilical Vein (HUV) D issection H uman umbilical co rds were collected from Shands Hospital (Gainesville, FL) within one day after delivery (IRB approval #64 2010). The hu man umbilical vein (HUV) was dissected automatically as described by D aniel et. al [153 ] Briefly, a 10 mm segment was discarded from both ends to eliminate any end defects. Umbilical c ords were cleaned by rinsing the outer surface with cold water and the inside of the vein was cleaned using a stainless steel cannula The umbilical cords were cleaned to prevent the remaining blood clot s from rupturing later in the process Subsequently, segments of 120 mm were cut out of the cords and a stainless steel m andrel (250" OD x .028") was inserted thr ough the vein. The cords were stored in a 86C freezer for at least 48 h ours to ensure a uniform temperature in a styrofoam container to allow progressive freezing of the vein at the rate of 2.5C/m in. For required experiments, frozen cords were removed from the 86 C freezer individually and promptly fixed between the headstock and the tailstock of the lathe. An automated program was used for the operation of the lathe (MicroKinetics, GA, USA) to give a co nsistent cutting thickness throughout the samples. Simultaneously, the cu tting tool was moved along the axis of the cord and therefore all the external tissue around the vein wall was removed with a final cutting thickness of 750 isolated vein had a smooth and uniform surface, which consists of the vein wall and as shown in F igure 2 1 After
57 lathing, veins kept on the mandrel were stored at 20 C for 2 hour s and fin ally placed in distilled water and allowed to rehydrate for 2 hours. The veins were carefully removed from the mandrel and cut to the desired lengths for the experiments. Decellularization After veins were disse cted from the cord, they were gradually tha wed and then carefully removed from mandrels. Decellularization steps were conducted at room temperature and under continuous agitation at 100 rpm. HUV segments were placed in glass bottles containing a 1% S odium dodecyl sulphate ( SDS ) solution. The HUV sam ples in the surfactant solution were agitated for 24 h ours using an orbi tal shaker. The solution was discarded and all detergent traces w ere removed by washing the veins in distilled water for 10 min utes 20 m inutes, 1 hour, 6 hours, and 24 hours. This was followed by incubating samples for 3 hours in 70 U/ml DNase (Sigma Aldrich Inc., St. Louis, MO) solution prepared in phosphate buffered saline. Figure 2 2 illustrates decellularized HUVE scaffolds. Sterilizati on and neutralization steps were conducted in the bioreactors after the system was assembled. Bioreactors Assembly and P retreatment Materials needed 3 bioreactors Medium reservoir + tubing mandrels Pipet tips (all types) Beaker (2) Silicon and Pharmed Tubing Connectors Cable ties Cable tie gun Sil icone stoppers Screw caps Air filter Paper towels
58 Dissection tools and trays Perfusion bioreactors used were made from hand blown glass, with dimensions of 100 mm in length and 12.9 mm internal diameter which were designed to have two ports on the shell side to facilitate ablumenal flow. Stainless steel tubing connectors with silicone stoppers and screw caps were used to load the 80 mm long HUV. As seen in Figure 2 3 three bioreactors w ere connected lumen to lumen side, and albumen to albumen side in ser ies with standard silicone tubing. Pharmed tubing (VWR, GA) was used around pump heads. A peristaltic pump (Cole Parmer L/S digital Stan dard Drive, model EW 07551 10) dro ve the solution through the lumen side of the vessel at a constant flow rate. Aft er as sembly, bioreactors were transferred to an incubator with 5% CO 2 and 37 C. The system was sterili zed in a solution of 0.2% perace tic acid (v/v) for 2 hours, and washed four times in d istilled water. Sections w e re finally washed in sterile PBS for 24 hours to stabilize the pH of the HUVs at 7.2 7. 4. Afterwards, the scaffolds were pretreated with c ell culture media overnight. Preparation o f Collagen Gels Acid soluble collagen hydrogels can gelate by neutralzing the pH. All the reagents were chilled to 4 Acid soluble Collagen that kept between 4 Sterile 10X phosphate buffered saline solution (0.2 M Na2HPO4, 1.3 M NaCl, pH 7.4) or 10X solu tion of buffered cell culture media 0.1 M HCl 0.1 M NaOH Phenol Red or pH Paper
59 Collagen gels were prepare d with a final volume of 10 ml as described in the following. 8 ml of c hilled acid solubilized bovine type I collagen (Inamed Biomaterials, Fremont, C A) in 0.01 N HCl was mixed with 1 ml of 10X buffered cell culture media. Then, pH was was adjusted to 7.4 using 1.17 ml of 0.1 M NaOH. Chilled acid solubilized bovine type I collagen (Inamed Biomaterials, Fremont, CA) in 0.01 N HCl was adjusted to pH 7.4 using 0.1 M NaOH. Cells were mixed with the chilled collagen solution and poured into the void space of the bioreactor at 37 C to initiate fibrilogenesis. Cell seeding protocol to bioreactors were de scribed in detail at C hapter 3 Seeding P rotocol Vascul ar smooth muscle cells isolated from the human umbilical artery 63 or myofibroblasts ( CRL 2854, ATCC VA, USA ) were seeded on the ablumen of the acellular HUV using hydrogel as the cell delivery method. Figure 2 4 illustrates the sequence of h ydrogel contr action around the HUV scaffold Under sterile conditions, 1.9 mg/ml collagen gels containing 110 6 cells/ml were prepared f rom acid soluble collagen (Devro Medical, CA, USA ) and loaded in the ablumena l void space of the bioreactors Under a laminar flow ho od, a total of 10 ml of gel solut ion containing the cells was prepared and inoculated individually into the void space of the bioreactors from one of the shell port s. Then, the bioreactors were pl aced in a 0% CO 2 incubator for 3 h at 37 C allowing for the collagen to gel and contract. Analytical Methods Construct C haracterization After the completion of the exp eriment, the HUV constructs were removed from the bior eactors. Each HUV construct was characterized. Histological analysis was done
60 with 5 mm ringle ts from proximal, middle and distal positions. Three other 5 mm ringlets w ere cut from the same positions to perform uniaxial tensile testing. Samples used for tensile testing were further analyzed for GAG and elastin content quantification. Two 10 mm ring lets will be used from the middle position for mRNA quantification (for relative gene expression) and dsDNA for total DNA content of the ringlet as in Figure 2 5 DNA Q uantification 10 mm HUV ringlets wer e snap frozen in liquid N 2 Tissue was turned into a po wder form by using a custom tissue crusher. The double stranded DNA was quantified using the PicoGreen Assay (Invitrogen, CA, USA) by using a fluorescence dye at wavelengths of 485nm (excitation), 535nm (emission) DNA concentration was determined by plotting a calibration curve by known amounts of DNA. Cell Metabolic A ctivity The metabolic activity of 10 mm lon g HUV construct ringlets was determined using the AlamarBlue (Invitrogen, CA, USA) assay. Samples were transferred to a new 24 well plate and 1 mL of supplement ed medium containing 10% (v/v) A lamarBlue (AB) reactant were a dded to each well. Plates w e re incubated at 37 C for 6 h to allow the reduction to occur. The reduction related to metabolic activity causes the indicator to change from oxidize d (blue) to reduced (red). Plates were removed from the incubator and absorbance w plate reader (Synergy plate reader, Bio Tek) Histology After expe riment termination, samples were f rozen and slice d using a cryostat (Microm HM 550,Thermo Scientific, USA) Samples were stained wi th H&E or RNA select dye
61 Gross distribution of cell viability and distribution of RNA was achieved by labeling permeable nucleic acid stain, exhibiting a very weak signal when bound to DNA and a strong signal when bound to RNA (absorption/emission maxima ~490/530 nm) [244 ] Sections for RNA staining were C) for 10 min, washed with PBS (2 x 5 min). Samples were then incubated for 20 minutes in 500 nM of RNASelect green fluorescent stain (Invitrogen, USA) in PBS to label RNA. Slides were then washed (2 x 5 min). Samples were viewed using a Zeiss Axio Imager M2 microscope with Axiovision soft ware release 4.6. The source images for RNA quantification were processed to measure pixel intensity of fluorescent areas by using a threshold command function script. Each image was divided into equally spaced sections. Uniaxial Tensile T esting 5 mm ringl ets were cut into tissue strips and then rin sed in PBS. The analysis was made at room temperature using a tens ile testing rig. Samples were clamped in with the use of sand paper to enhance grip ping of the samples. Stress was applied in a circumferential di r ection; first, specimens were preloaded to a stress of 0.005 N at a rate of 5 mm/min and then elongated until failure, with f orce and extension recorded over time. Statistical Analysis All data are presented as the mean standard deviation. Statistical s ignificance (p<0.05) is assessed with a multiple comparison test (Tukey test) via one way ANOVA. Statistical analysis of real time PCR data is performed by the Relative Expression
62 Software Tool. Differences in gene expression level between experimental gro ups is considered significant when p <0.05.
63 Figure 2 1 Autodissectionmethodology. A B) Mandrel mounted cords wer e frozen to 86C and lathed C) HUV were dissected to yield a final thickness of 750 D) HUV se gments were thawed gradually and removed from mandrel Scale bar corresponds to 1c m. Figure 2 2. Decellularized human umbilical vein. Right panel shows the same material that was insert ed th r ough a glass mandre l. Scale bar corresponds to 1c m.
64 Figure 2 3. Dynamic cell culture setup. Three bioreactors (A, B, C) were connected in series. Two flow circuits, lumen and ablumen with respective medium reservoirs (D, E) and rotary pumps (G, F). Decellularized HUV samples were seeded with myofibroblasts and culture d for a period of 21 days.
65 Figure 2 4. Hydrogel contraction around the HUV scaffold. The sequence of hydrogel contracting from the bioreactor glass wall and around the HUV is sh own. A) The hydrogel was loaded into the bioreactor void space. B) After 5h of gel reaction, the collagen has polymerized and detached from the bioreactor wall. C) The medium was initially perfused around the void space and circulated freely on the ablumen side. D) After the selected time periods, the vascular constructs were removed from bioreactors and characterized.
66 Figure 2 5. After the completion of the exp eriment, the HUV constructs were removed from the bior eactors. Each HUV construct was characte rized. Histological analysis was done with 5 mm ringlets from proximal, middle and distal positions. Three other 5 mm ringlets w ere cut from the same positions to perform uniaxial tensile testing. Samples used for tensile testing were further analyzed for GAG and elastin content quantification. Two 10 mm ringlets was used from the middle position for mRNA quantification (for relative gene expression) and dsDNA for total DNA content of the ringlet.
67 CHAPTER 3 HYDRO GEL CELL SEEDING AROUND VASCULAR TISSUES Pri or to the perfusion studies in C hapters 4 and 5, c ontraction of hydrogels around porous soft tissue required optimization as scaffold porosity and hydrogel swelling can be problematic due to the rate of transmural and abl umenal flow (in the bioreactor) in the initial culture sequence. After the initial seeding, two parameters were studied to optimize the seeding protocols. i) The optimal start time of the flow circulation within the perfusion bioreactor, and ii) Perfusion flow design to reach physiological parameters overtime are discussed in this chapter. The optimized parameters are shown in Figure 3 1. Materials and Methods Experimental Methods Human umbilical vein (HUV) dissection, decellularization, seeding protocol and bioreactors assembly and treatme nt were previously described in Chapter 2. Hydrogel Contraction T ime In order to find the optimal time for the hydrogel to solidify without the environment becoming acidic 4 different time periods were investigated after initial cell seeding. Time periods for this experiments were t=0 hours t=3 hours t=5 hours t=7 hours Perfusion Flow Profile Design After gel polymerization culture medium (F igure 3 2 and F igure 3 3 ) was circulated in two different ways : i ) lum enal and ablumenal flow combined, ii) only lumenal flow. Lumenal and ablumenal flow: After the gel contracted, the culture medium was circulated on the ablumen void space at a flow rate of 1.5 ml/min while no medium was
68 run i n the lumenal side as shown in F igure 5. After 48 h ours lumen flow was slowly increased by increments of 2.5 ml/min every 2 h ours until 10 ml/min was reached. Lumenal flow: through the lumenal flow circuit at a minimal flow rate of 10 ml/min for 24 hours to maintain hydrogel stability. There was no media directed through the ablumenal flow circuits during culture. In both conditions, from day 2 to day 10 lumenal flow rate was incrementally increased to reach a maximum of 120 ml/min. The same initial flow conditioning profile was used in all bioreactor sets from seeding to day 10. Figure 3 2, and 3 3 show the flow conditions used from day 10 to 30, with a total culture time from seeding to harvest of 30 days, under standar d conditions of 5% CO 2 at 37C. Results Hydrogel Contraction T ime Hydrogel contraction times prior to flow a re shown in Figure 3 4. t=3 hours was sufficient time for hydrogel to solidify without the environment becoming acidic (gel with phenol red changes color) in the absence of the lumenal flow. Hydrogel was contracted from 6mm to less than 0.5 mm at the end o f 10 days by cellular remodeling of the gels. Perfusion Flow Profile Design The hydrogel stability was investigated for both flow conditions after flow was initiated. The time periods were 2 hours and 24 hours. After two hours, hydrogel was no longer inta ct at the ablumenal flow only circuit, as shown in Figure 3 6. Conversely hydrogel was intact after a period of 2 hours at the lumenal flow only circuit as shown in Figure 3 5. After 24 hours, the hydrogel began to detach from the surface and support itse lf around soft tubular tissue without sa gging, as shown in Figure 3 7.
69 The outcome of the flow perfusion design was investigated after 30 days of culture. By limiting the ablumenal flow, there were more cells at the end of the perfusion culture. Most cells were found to migrate, and proliferate toward lumen. Cells are relying on hydraulic perme ability through lumen, as demonstrated at Figure 3 8. Distance of the cells to the lumen was 522 45 m while there is ablum enal, and lumenal flow and 133 40 m w hile there is only lumenal flow, as seen in Figure 3 9. Summary Having a cell populated scaffold as fast as possible is a clinical research objective for the immediate use of vascular tissue engineering products. Some of the cell seeding strategies wer e r eviewed in the C hapter 1 In some of the decellularized ex vivo tissues, recellularization can be a problem both in vitro and in vivo because of the dense extracellular matrix. To overcome these problems, Neff et al removed the adventitial layer of the ar teries and obtained higher initial cell attachment, and more contractile elements compared to the porcine arteries with adventitial layer [228 ] I n this study our method combined the two approaches, using hydrogels as a method of cell delivery and decellularized tissu e as a main material. Hydrogels were used as a cell delivery vehicle to inoculate cells into the ablumenal void space of the bioreactor, and were contract ed around the scaffold. This has been shown be a promising approach to efficiently deliver high concen trations of cells directly to a 3D structure. It was initially thought that the flow in the ablumen allowed the cells to obtain enough nutrients without causing the hydrogel to deform or to be flushed away. hydrogel to form uniformly around non adhesive acellular HUV. Previous investigations found that a wall shear stress of 1.9
70 dynes/cm 2 was found not to effect the hydrogel structure [245 ] In our study, by modulating the initial perfusion conditions, cells were migr ated up to 100 micr ons from the luminal scaffold periphery Transmural flow was enough to nurture the cells over this time period. Previous hydrogel biopolymer techniques in the literature used distensible but rigid tubes where hydrogels were contracted and remodeled over ti me [159 ] [246 ] Compared to this technique, c ontraction of hydrogels around porous soft tissue required opt imization S caffold porosity and hydrogel swelling can be problematic due to the rate of transmural and abl umenal flow (in the bioreactor) in the initial culture sequence Under dynamic culture conditions the hydraulic conductivity (transmural) reduced sig nificantly after 4 days, with no flow detected over the remaining time course. This was likely the result of several factors : i) compaction of the hydrogel around the HUV reduces porosity of this layer, as well as a general compaction noted of the collagen fibers due to the mechanical forces imparted by the pulsatile flow, ii) as media serum binds the scaffold the pore size was progressively reduced, much like membrane fouling, and iii) the progressive increase in cell density and ECM production forms a de nser outer layer of the scaffold that further reduces h ydraulic permeability. O ptimization of the lumenal and ablumenal flow profile during the initial c ulture was an important factor f o r enhan ced uniform hydrogel compaction as cell delivery is fundamental for all studies in this dissertation. This m ethod was used in the Chapter 4,5, and 6
71 Figure 3 1. The sequence of events during cell seeding that required optimization. After the initial seeding, two parameters were studied to optimize the seeding proto cols: i) the optimal start time of the flow circulation within the perfusion bioreactor, and ii) the perfusion flow design to reach physiological parameters. Collagen hydrogel cell inoculation Hydrogel contraction time Needs optimization Flow profile design Needs optimization
72 Figure 3 2 Flow conditioning profile. After gel polymerization ,the culture medium was circulat ed from the lum enal and ablumenal flow circuits.Lumenal flow was gradually increased from 0 ml/min on day 1 to 120 ml/min at the end of the process, while the ablumenal flow rate remained constant at 2 ml/min.
73 Figure 3 3 Flow conditioning profile. After gel polymerization culture medium was circulated through lumenal circuit Lumenal flow was gradually increased from 2 ml/min on day 1 to 120 ml/min at the end of the process, while the ablumenal flow rate remained constant at 0 ml/min.
74 Figure 3 4 The hydrogel was loaded into the bioreactor void space. Gel reaction color changes at t=0 was observed at t=3, 5, and 7 hours. 3 hours was sufficient time for the hydrogel to solidify without the environment becoming acidic (where the gel with phenol red c hanges color).
75 Figure 3 5 Stability of the hydrogel after the initiation of the flow where media from the ablumenal circuit was excluded After 2 ho urs of flow only in the luminal circuit, hydrogel became slightly detached as it contracted around the HUV Figure 3 6 Stability of the hydrogel after the initiation of the flow where media from the lumenal circuit were excluded At t=2h, the hydrogel was disengaged from the HUV
76 Figure 3 7 24 hours after cell seeding. M edia from the ablumenal circuit we re excluded The collagen hydrogel remained intact but was detached from the bioreactor shell surface.
77 Figure 3 8 Comparison of flow conditions. Conditions that excluded media from the ablumenal circuit (cell seeded surface) provided a strong chemot actic drivi ng force for cellular migration, as seen in the right panel. Most cells in constructs cultured with both lumenal and ablumenal circuitsare located in the mid to outer layers of the scaffold, near the seeded surface as seen in the left panel 150mm
78 Figure 3 9 Proliferation and migration of the cells to the lumen. The distance between cells and the lumen was 522 45 m when both perfusion circuits were used, and 133 40 m when the ablumenal circuit is excluded. 0 100 200 300 400 500 600 Lumenal/Ablumenal flow Lumenal Flow Only Distance of the cells to the lumen ( m)
79 CHAPTER 4 IMPROVED RECELLULARIZA TION OF ENGINEERED SCAFFOLD USING DIRECTED TRANSPORT GRADIENTS Tissue engineering or de novo regeneration of small diameter blood vessels has shown significant promise. However, regeneration of a functional vascular media and intima in a timely fashion has been problematic [9 ] [48 ] [83 ] [85 ] [247 ] As such, creating a cell dense scaffold is a crucial step towards recapitulating a functional medial layer with vasoactive behavior and suitable mechanical properties. As an initial step, vascular smooth muscle cells (vSMC), are required to migrate and proliferate to create [51 ] vSMC migration and proliferation have been a n obstacle for decellularized tissue scaffolds, as cell seeding is limited to the outermost surfaces, resulting in extended culture periods to create cell dense scaffolds. This poses a barrier for passive cellular migration and ultimately nutrient deprivat ion in areas most distant to fresh media [169 ] In our previous investigations, dual circuit perfusion biorea ctors (lumen and ablumenal flow perfusion) provided a nutrient rich environment to promote cell proliferation and matrix deposition to assess early remodeling conditions of acellular ex vivo scaffolds [154 ] While the overall culture period was limited to 3 weeks, cell migration was limited to approximately 150 microns from the seeded surface primarily due to mass transfer limitations. Ex vivo derived vascular tissues preserve s the essential extracellular matrix components. However, cell infiltration from albumen to lumen still remains a problem in these materials. The investigations described herein aimed to enhance cellular migration in order to generate more uniformly popula ted vascular constructs. We report here a model system to drive cell migration, with myofibroblasts seeded on acellular human umbilical vein. These are the initial steps to populate an ex vivo tissue where
80 cells were uniformly distributed over a total cult ure time of four weeks. In order to achieve this, using perfusion bioreactors, pressure and flow conditions between the lumen and ablumenal circuits were modulated to control hydraulic conductivity. Lumenal perfusion conditions included 120/80 mmHg (high p ressure) or 50/30 mmHg (low pressure) either with or without media in the ablumenal flow circuit. Our hypothesis was that cells seeded on the ablumenal surface of the scaffold would respond to the chemotactic nutrient gradient (driven by higher lumenal pre ssure), and would display a more uniform distribution across the scaffold wall, rather than being limited to the scaffold periphery. Conditions that excluded media from the ablumenal circuit (cell seeded surface) were designed to create an aggressive nutri ent gradient across the porous scaffold, providing a strong chemotactic driving force for cellular migration (lumen toward albumen). The goal of the present work was to utilize a directed nutrient gradient to modulate micro environmental conditions ultimat ely to enhance the scaffold recellularization processes. Materials and Methods After the completion of each experimental group HUV constructs were removed from the bioreactors. Samples (5 mm ringlets) from proximal, mid and distal positions were analyze d to assess scaffold morphology, cell density, metabolic activity, cell distribution, and biomechanical properties, as follows. Common experimental methods were previously described in Chapter 2. Experimental M ethods Human umbilical vein (HUV) dissection, decellularization, seeding protocol and bioreactor assembly and treatment were previously described in Chapter 2.
81 Perfusion Flow Profile D esign After gel polymerization culture media (F igure 3 1) was circulated through the lumenal flow circuit at a minim al flow rate of 10 ml/min for 24 hours to maintain hydrogel stability. From day 2 to day 10 the flow rate was progressively increased to reach a maximum of 120 ml/min with a pulse frequency of 3 Hz. There was no media directed through the ablumenal flow ci rcuits during this period. The same initial flow conditioning profile was used in all bioreactor sets from seeding to day 10. Table 3 1 shows the flow and pressure conditions used from day 10 to 30, with a total culture time from seeding to harvest of 30 d ays, under standard conditions of 5% CO 2 at 37 C. Analytical M ethods After the completion of each experimental group HUV constructs were removed from the bioreactors. Samples (5 mm ringlets) from proxima l, mid and distal positions were analyzed to assess s caffold morphology, cell density, metabolic activity, cell distribution and biomechanical properties as follows Burst P ressure Using the ultimate tensile stress values (from the circumferential tensile testing) construct burst pressure was estimated usi (2 1) Equation 2 1: Laplace relationship Where P is the pressure; r, the radius; the Cauchy stress; and h, the wall thickness (25)
82 Calculation o f Scaffold P ermeability Permeability was assessed under low ( 50/30 mmHg ) and high (120/80 mmHg) of the flow rate Q (m/s) and pressure drop P across a matrix: (2 2) Equation 2 Where A is the scaffol d area (m 2 ), L is the matrix thickness (m), and is the dynamic solution viscosity (Pas). Pressure ( P ) was monitored using a digital pressure gauge (model # EW 68332 06 Cole Parmer, Vernon Hills, IL) immediately downstream from the bioreactor. R esults V ascular Construct Morphology a nd T hickness Gross morphology of cellular and acellular vascular constructs can be seen in Figure 4 2. Both mechanical data and observation show the scaffold to have two distinct zones, as seen in Figure 4 3 (image post tensi le testing), showing the inner and outer in proximal, mid and distal positions, wi th the initial thickness (prior to seeding) being 0.75 mm. The total thickness of scaffold and hydrogel composite (prior to gel contraction) was 6.0 mm. Constructs exposed to dynamic culture (DP Hp, SP Lp, and SP Hp) over the 30 day culture period decrease d in wall thickness from initial values, displaying significant variation between each condition, as follows: DP Hp: 1016 300 m; SP Lp: 1176447 m; SP Hp: 543431 m; and DP ACHp: 1560406 m. A
83 statistical multiple comparison analysis showed that the o verall thickness of the single perfusion high pressure set (SP Hp) was significantly reduced compared to single perfusion low pressure (SP Lp) and dual perfusion acellular high pressure bioreactor (DP ACHp) sets, but was similar to the dual perfusion hig h pressure set (DP Hp). Scaffold C ellularity After 30 days of dynamic conditioning, cell density increased up to eight times relative to the initial seeding density of 110 6 cm/linear scaffold. Terminal cell densities increased to 4.81 2.0710 6 cm/linear scaffold (DP Hp), 4.27 1.8410 6 cm/linear scaffold (SP Lp), and 8.12 4.8010 6 cm/linear scaffold (SP Hp). While the SP Hp constructs displayed higher average cell density values, both cell density and metabolic activity were statistically similar betwee n all groups, as in Figure 4 4 Cell Distribution U nlike DNA; mRNA is unstable and has a short half li fe [248 ] Fluorescence RNA labeling was used to localize cells within the scaffold and as an indirect method to assess cell viability. Figure 4 5 (left column) displays the scaffold under standard fluorescent conditions to highlight the scaffol ds ECM structure and to localize cells by labeling RNA within each construct (Figure 4 5. A,D,G). All images in Figure 4 5 show the construct orientation with the cell seeded surface (ablumen) on the right hand side, and the lumen surface on the left. The ima ges in center panel (Figure 4 5. B,E,H ) were ob tained using fluorescent thres holding to isolate RNA staining from background non specific ECM binding. Each image of the vessel wall was divided into three equal zones (1, 2, 3) to quantify cell distribut ion a cross the scaffold. Figure 4 5.C,F,I represents pix el intensity from labeled RNA within each section and a intensity value given as a percentage.
84 Across all perfusion systems, cells were shown to adhere, proliferate and migrate into the scaffold over the culture period, providing evidence for active remodeling. Constructs exposed to complete media on both surfaces under physiological pressure conditions (DP Hp) displayed an RNA distribution significantly denser in zone 3 adjacent the seeded surface an d a reduced density in zone 2. No cells were detected within the inner zone 1, adjacent to the lumen of the vessel ( Figure 4 5. A,B,C ). Constructs cultured with single perfusion systems at low pressure (SP Lp) resulted in a more uniform distribution with h igh concentrations in zones 1 and 2 with a reduced dens ity in zone 3 ( Figure 4 5. D, E ,F ). Single perfusion, high pressure systems (SP Hp) resulted in similarly distributed RNA in the middle an d outer layers ( Figure 5. G,H,I ), with reduced RNA in the inner zone. Biomechanical Properties Analysis of the vascular constructs tensile properties displayed two distinct failure peaks, as shown in Figure 4 3, associated with the inner tissue zone composed of the original vascular media and outer zone consisting of h ydrogel. Results describing the ultimate tensile strength (UTS) after 30 days in culture show the inner zone of the constructs, exposed to the single perfusion higher pressure system (SP Hp 120/80 mmHg), to have signific antly higher UTS of 734210 kPa relative to the low pressure system (SP Lp 50/30 mmHg) with a UTS of 30562 kPa and acellular control sets (DP ACHp). The acellular HUV exposed to dynamic perfusion (DP ACHp) exhibited a significant decrease in UTS after 3 0 days from 55041 to 314
85 pressure systems trending higher relative to controls or low pressure constructs, seen in Figure 4 6A Burst pressure values calculated from Laplace equation are DP Hp: 542263 mmHg, SP Lp: 27642 mmHg, SP Hp: 705242, and DP ACHp: 1177135, seen in Figure 4 6B 4 7. ups to be significantly stiffer than constructs cultured under low pressure conditions in the inner tissue zone, as follows: DP Hp: 4578 1718 kPa, SP Lp:920 294, and SP Hp: 4078 elastic properties, as follows DP Hp: 277122 kPa, and SP Lp: 9835 kPa, SP Hp: 396 144 Hp, and DP Hp were statistically similar, while SP Lp had statistically lower values. There was no statistica l difference in the strain at maximum load values. The inner hydrogel remodeled scaffold, failed around 175% strain which is similar to acellular HUV values. Permeabil ity P ermeability of the HUV was calculated for each pressure, using a mean thickness value of 0.75 m. The average permeability coefficients were 1.20 0.33 10 12 m 2 at low pressures (50/30 mmHg), and 2.23 0.4 10 12 m 2 at high pressures (120/80 mmHg). H igh pressures resulted in a significantly higher permeability coefficient than low pressures.
86 Summary Engineering functional tissues has become a reality with clinical use of several regenerated organ systems [249 252 ] While earlier clinical successes were limited to relatively less complex tissues such as skin, more recent work has shown potential for more challenging organ systems to be regenerated. Ideally the time frame from diagnosis to implantation should be limited only by patient readiness for the procedure; in vitro from base components (scaffold + cells). Custom gro wn engineered tissues require an extended culture period lasting upward of 2 months, often significantly longer [253 255 ] During the regenerative process the distan ce between seeded cells and a suitable nutrient source increases as cells migrate to populate the scaffold. The distance between seeded cells and a suitable nutrient source creates an unfavorable nutrient gradient, which has been identified as the primary limiting factor to create functional tissues [112 ] The often results in the formation of a cell capsule around the construct, which further impedes migration as a progressively more hypoxic and nutrient deprived environ ment is created. Due to clinical requirements for rapid graft development, improving the rate at which scaffolds are repopulated is a key research objective. This is particularly important with bypass surgery where the time frame of development needs to be days, not weeks or months [169 ] [256 ] [257 ] The primary o bjective of this work was to study key parameters that influence cell migration in order to design systems that improve cellular infiltration and reduce the in vitro culture phase of construct development. These investigations compare the influence of dire cted nutrient gradients by precisely contr olling perfusion conditions to a ffect cell migration and function during early remodeling events. We hypothesized that
87 if nutrient gradients were applied across the vessel wall, a strong chemotactic response would lumenal surface, improving cell distribution and ECM interactions in a more controlled fashion. As cells migrate into and through an acellular scaffold, remodeling processes ar e in continual flux with both catabolic and anabolic pathways activated. These processes play a key role not only in the biological functionality of the construct, but also h for a vascular graft is critical. As such, these investigations also detail variation in scaffold mechanics as a function of each perfusion condition. A controlled nutrient gradient was generated using either single or dual perfusion circuits, under eit her low (50/30 mmHg) or high (120/80 mmHg) relative pulsed pressure profiles to assess the effects of nutrient gradients. The higher pressure profile was chosen to emulate physiological conditions, while the lower values were chosen to create more aggressi ve nutrient conditions by reducing hydraulic permeability. Results show that between each of the perfusion conditions, overall construct cellularity was constant, but distinct differences were seen in cell distribution that was attributed to each perfusion characteristic. Constructs cultured under dual perfusion at high relative pressure (DP Hp 120/80 mmHg) show most cells to be located in the mid to outer layers of the scaffold, near the seeded surface where initial conditions provided adequate nutrient s. As cell density increased and progressively reduced the availability of nutrients from the ablumenal flow circuit, the nutrient source from the lumen began to play a more dominant role. As such, cell migration toward the lumen was initially slower
88 due t o nutrient availability, but as nutrients became limited, cells migrated toward the lumen populating the mid region of the graft by the end of the culture period. Conversely, the single perfusion circuit running at low pressure (SP Lp 50/30 mmHg) create d a more aggressive nutrient gradient due to zero nutrient availability from the ablumen, and a reduced transmural flow relative to the higher pressure systems. This aggressive nutrient gradient resulted in cells behind the leading migration edge receiving less nutrients as cell density increased and migrated toward the lumen. Nutrient limitation made the seeded ablumenal surface less favorable and immediately drove cells toward the lumen. By the end of 30 days, these conditions resulted in the highest aver age densities at the lumen interface and the lowest on the seeded surface. By increasing the lumenal pressure (SP Hp 120/80 mmHg), the transmural supply increased (increased hydraulic permeability) and was able to supply nutrients to areas more distant fr om the lumen as supply exceeded demand by cellular consumption at the migration front. This resulted in reduced migration towards the lumen as metabolic demands were met. These data clearly show the importance of hydraulic permeability as a significant dri ving force in cell migration when nutrients are limited. Tensile analysis shows the initial acellular construct to display an anisotropic stress strain relationship that became more pronounced after seeding and culture. The int was associated predominantly with the medial layer of the HUV, while the secondary fracture point was the outer layer of the construct, dominated by the collagen hydrogel that was used to deliver cells to the scaffold surface, as in Figure 3. Compariso n of cell localization and construct mechanics show that cell distribution as a function of nutrient transfer has a significant effect on construct
89 performance. While the total cell density was similar in all groups, biomechanical characterization has show n that each of the culture conditions resulted in significant differences in construct remodeling, as shown by variation in scaffold mechanics. Single perfusion circuits at high pressure (SP Hp 120/80 mmHg) show significantly higher UTS compared to all ot to the initial decellularized HUV and acellular controls. Conversely, the single perfusion system with low pressure (SP values, showing the c onsequence of inefficient nutrient distribution across the scaffold. The dual perfusion system (DP Hp) shows cells to have migrated more slowly toward the lumen even with increased nutrient delivery from this circuit. This result indicates that having abu ndant nutrients on all sides of the scaffold is not sufficient to recellularize the graft, and may in fact hinder cell migration in early stages of remodeling. Our study and previous studies have shown that perfusion systems play an important role to modul ate vascular tissue growth. Several groups have demonstrated that v SMCs can grow, infiltrate, and remodel within decellularized scaffolds [167 ] [169 ] [227 ] [256 ] Yazdani et al. removed the adventitial layer of the arteries and obtained higher initial cell at tachment, and more contractile elements compared to the porcine arteries with adventitial layer [169 ] Althou gh these events have been studied, full cell infiltration remains a problem. We report here a model system to drive cell migration, with myofibroblasts seeded on acellular human umbilical vein. There has not been any publication that reports full migration in an ex vivo tissue after three weeks of culture. We report here some initial steps to populate the ex vivo tissues where cells uniformly distributed in a total culture time of four weeks. Further studies will try to enhance the
90 functional characteristic s of this model system to obtain a vasoactive engineered scaffold with a continuous endothelial cell layer. However, in order for v SMCs to co to be a maximum 2 50 micron s far from the lumen initally [258 ] Continued understanding of perfusion culture models provides a more detailed perspective of graft remodeling both in vitro and i n vivo where diffusion limits nutrient transfer and th us graft performance. L'Heureux et al. [171 ] and Guillemette et al [259 ] reported that the integration of a tissue engineered vascula r graft with a functional vasovasorum was established within three months after implantation. While the in vivo environment is clearly more complex, the conditions generated with the SP Hp circuit would be similar to those experienced by grafts when implan ted in the arterial vasculature. As such, these systems may provide further insight into early vascular remodeling events based on limited nutrient supply. These investigations have shown that by modulating perfusion culture conditions, the density and the distribution of the viable cells can be regulated to improve graft tissue biomechanics and graft maturation. In addition, results demonstrated the potential of the HUV as a 3D scaffold for guided vascular tissue regeneration.
91 Table 4 1. Acronym, and th e description of each bioreactor condition. Acronym Description DP ACHp Dual perfusion Acellular control (no seeded cells) DP Hp Dual perfusion seeded, with complete media in both flow circuits (lumen pressure at 80/120 mmHg) SP Lp Single perfusio n seeded, with complete media only in the lumenal circuit (Low pressure 30 50 mmHg) SP Hp Single perfusion seeded, with complete media only in the lumenal circuit (High pressure 80/120 mmHg) ***Flow rates were maintained 120 ml/min with 3 Hz frequenc y for all conditions. Figure 4 1 Schematic drawing of s eeding and hydrogel compaction within the perfusion bioreactors: Cells are suspended within the hydrogel prior to polymerization and inoculated into the void space of the bioreactor (A). The hydro gel/cell composite begins a contraction and compaction stage around the HUV scaffold (B), ultimately forming cell dense zone that is progressively remodeled around the HUV (C). Schematic representation of a single bioreactor is shown in F igure (D).
92 Figu re 4 2. Representative images of vascular construc ts. Cellular: SP HP (A B) and acellular: DP ACHp (C D) constructs shown with corresponding sections (ringlets) used for tensile testing. The more substantial structure of the cellula r construct is shown com pared to the acellular material after 30 days perfusion culture. Scale bars on le ft column images correspond to 5 m m.
93 Figure 4 3. Representative circumferential tensile stress strain relationship (A) of the cellular constructs, images of a cellular con struct ringlet after failure (B,C) for the clarification of inner and outer zones. D. Construct thickness. Thickness of the constructs of inner and outer zones, and cumulative thickness for all conditions after 30 days of perfusion culture. The initial a verage value for the wall thicknesses are represented with the dashed line. denotes significant difference in thickness b etween SP HP, and DP ACHP (p<0. 05).
94 Figure 4 4. Cell density and metabolic activity. Cell density (above left) of ce llular HUV constructs at Day 30 Cell seeding density is shown with a dashed line (1x10 6 / linear cm scaffold). Values indicate mean standard deviation. Metabolic Activity/cells (above right) of the vascular constructs with Alamar Blue dye. No statistical difference of metabolic activity /cell value was observed and values were consistent wit h our previous work.
95 Figure 4 5 Fluorescence RNA labeling was used to localize cells within the scaffold and as an indirect method to assess cell viability. Each image was divided into sections (1,2,3) from lumen to ablumen side ( B ,E, H ), an d fluorescen ce pixel intensity was measured for each region ( C,F,I ). Scale bars (center panel images) correspond to 100 m.
96 Figure 4 6. Ul timate tensile strength (UTS) of inner and outer zone of the constructs(A).Calculated Burst pressure values derived from the Lap lace's Law( B ). denotes significant higher UTS compared to acellular control set (DP ACHp), and SP LP (p<0.001) denotes Burst pressure of SP Hp significantly higher than SP Lp (p<0.05) # denotes that burst pressure of DP ACHP was statistically higher than all of the conditions, ( DP Hp, and SP L p<0.001 SP H p < 0.05 ) n = 9.
97 Figure 4 7. DP Hp significantly h Lp p < ACHp set was higher than outer zone of DP Hp (p<0.005) SP Lp ( p<0.001 )
98 CHAPTER 5 CONTROLLED VASCULAR CELL FUNCTION BY USI NG MODELED MECHANICA L ST IMULATION Scaffold remodeling is a complex multi factorial process in which chemical and mechanical signaling directs cell function. In order to assess (and modulate) these progressive remodeling changes perfusion systems are used to emulate physiological conditions. Current studies have focused on mechanical stimulation based on constant pulse frequencies and flow rates rather than the temporal variation seen in vivo [49 ] [154 ] [160 ] [167 ] [230 ] Recently it has been shown that cell adaptation is dependent not only on the rate of the mechanical stimuli but also its continuous change [125 ] In order to assess thi s temporal variation in pulse frequency, a model 3D scaffold derived from the human umbilical vein (HUV) was used to test the hypothesis that short term variation in heart rate, such as elevations in respiratory activity, can modulate vascular function and remodeling. If so, the var iability in vascular pulse should produce a more physiologically relevant system to investigate vascular cell function, tissue regeneration, and the effects of pathological conditions. Using vascular perfusion bioreactors and pro grammable process flow circuits, human vSMCs were dispersed within a collagen hydrogel and inoculated into the shell side void of the bioreactor to deliver cells directly to the ablumenal surface of the HUV scaffold. The pulse rate of a healthy volunteer w as logged over a 12 hour period to model the effect of daily activities on heart rate. Using masterflex linkable instrument control software (v3.1) programmable peristaltic pumps ran the script that delivered pulse frequencies that emulated the variable h eart rate. Three different mechanical condition s were assessed to compare the e ffect of variability in pulse frequency over a total culture time of 6 weeks: i) Constant pulse frequency, 1.3 Hz (CF), ii) Repeating 12h
99 of variable pulse frequency (physiologi cal pulse frequency) followed by 12h of constant (1 Hz) frequency (PF 12), and iii) Repeating 24 h variable physiological pulse frequency (PF 24h). All conditions averaged 1.3 Hz pulse frequency (78 bpm) over each 24h repeating period. Materials and Metho ds Cell culture media within the bioreactors consisted of low glucose (2.0 g/l) Bovine Serum (FBS), 1% l glutamine, and 1% penicillin streptomycin (Invitrogen, CA), and ascorb ic acid with 5% CO2 at 37 C. The same initial flow conditioning profile was used in all bioreactor sets from seeding to day 7, as described in C hapter 3 After gel polymerization culture medi a was circulated through the luminal flow circuit at a minimal f low rate of 10 ml/min for 24 hours to maintain hydrogel stability From day 2 to day 7 the flow rate was progressively increased to reach a maximum of 120 ml/mi n Figure 5 1 shows the frequency stimuli profile of each condition used from week 1 to week 6. In this study, t he short term variation of pulse rate such as that caused by elevations in respiratory activity, as occur during exercise was included a produce a more physiologically relevant frequency profile. Pulse rate of an individual measured in dai ly activities for a12 hour period was used. The pulse rate of an individual was converted to revolutions per minute (RPM). Calculated (RPM) measurements were used as the input parameters to program the software of the peristaltic pump. Three different mech anical conditions were assessed to compare the affect of variability in pulse frequency over a total culture time of 6 weeks: i) Constant pulse frequency, 1.3 Hz (CF), ii) Repeating 12h of variable pulse frequency (physiological
100 pulse frequency) followed b y 12h of constant (1 Hz) frequency (PF 12), and iii) Repeating 24 h variable physiological pulse frequency (PF 24h). All conditions averaged 1.3 Hz pulse frequency (78 bpm) over each 24h repeating period. Measurement o f Contraction a nd Relaxation F orces V aso activity testing was performed on three ringlets from each bioreactor. Each ring was placed in a homemade organ bath containing Krebs Henseleit solution, and then mounted to stainless steel hooks connected to the force transducer, as previously describ ed [262 ] All ring segments were preloaded to 0.15 g and stimulated with 60 mm KCl. Afterwards, the rings were stimulated with the ATP disodium salt. After contraction, the rings were stimulat ed with papaverin. Changes in force overtime were recorded. Scanning Electron Microscopy Samples were rinsed with phosphate buffered saline (PBS) three times for 5 minutes to remove excess cell culture media, and then fixed with 2.5% glutaraldehyde for 4 h ours and then 1% osmium tetroxide in PBS. Afterwards, samples were progressively dehydrated in graded ethanol solutions (25%, 50%, 75%, 85%, 95%, and 3x100% EtOH). A critical point drying (CPD) process was required to remove all liquids, briefly samples we re maintained in a pressurized liquid carbon dioxide vessel (above 700 psi). After a final CPD step samples where sputter coated before being analyzed using a Hitachi S 4000 FE SEM at 15kV. Results Morphology o f t he Scaffolds Gross morphology of the cellular constructs after 6 weeks of culture is shown in Figure 5 2. There was no visual difference between different conditions. Figure 5 2.A shows the construct before harvesting within the bioreactor. Figure 5 2.B D shows the
101 bioreactors just after harvesting, with the connector adaptors attached, and after the adaptors are removed, respectively. Biomechanical P roperties Two regions that represent the elastic behavior of the vessel wall are used to calculate the tensile properties of the scaffold: i) low (phys iological range) strain region which reflects the behavior of the material over the normal physiological pressure range, and ii) high strain region (failure range) where collagen fibers extend with increased strain levels prior to failure. Our previous a nd cu rrent investigations have shown that the constructs have a biphasic stress and strain behavior in the failure range. As seen in Figure 5 3A, this behavior was consistent with our previous investigations for constant frequency stimuli conditions (CF, C F Myo). When PF stimuli were applied, single failure peak s with higher stress values were observed progressively from PF 12 to PF 24h conditions. Most of the constructs exhibited a single peak at PF 24h condition, where constructs from PF 12 conditions exh ibited both behaviors. The u ltimate failure peak was used to calculate After 6 weeks of culture, tensile testing results show evidence that the constructs exposed to the continuous physiological frequ ency stimuli system (PF 24h) have Results for the UTS (0.790.25 MPa) di splayed a similar trend, higher relative to all other conditions, as see in Figure 5 3(B C ). Stiffer P F 24h constructs were found to have a similar extension at the maximum stress peak as presented in Figure 5 3 D There was no statistical difference at % strain at maximum load values between CF Myo, PF 12, and PF 24h. However, % s train at
102 maximum stress v alues f or CF condition had significantly higher % strain values compared to all other conditions. Over the physiological range, the variability between the conditions was found to be independent of mechanical sti muli but related to the cell lineage CF My o condition exhibited significantly stiffer behavior both with Y and % strain at 120 mmHg comp ared to all other conditions ( 3.3 % 0.88), as see n in Figure 5 4. Pharmacological Response o f t he Constructs The maximum contracti le and relaxation force of the constructs is shown in Figure 5 5. Significant differences were found at the contraction and the relaxation force of the constructs after 6 weeks of culture. The con tractility was not correlated to the distribution or the amo unt of the cells within the scaffold. The maximum contraction and the relaxation force diminished significantly when 24 h of physiological pulse stimuli (PF 24h) was applied. The CF Myo condition had lower contractility as expected compared to vSMC in the same mechanical stimuli (CF). Scaffold Microstructure a nd Cell Distribution SEM images in Figure 5 6 represent the longitudinal cross section of the construct at 1k times magnification. Images reveal a more organized collagen fiber orientation of the PF s amples compared to CF that is more similar to initial acellular HUV. Figure 5 7 displays the scaffold under standard H&E stain. All images in Figure 7 show the construct orientation with the cell seeded surface (ablumen) on the right side, and lumen surfa ce on the left. Cell location within HUV constructs showing that cell distribution was dependent on culture approach. After 6 weeks, cells density increased
103 embedded within matrix pores. The gross morphology and overall histological structure of starting m aterial was changed showing more densely packed organization. Across all conditions, cells were shown to migrate into the scaffold, and actively remodeled overtime. Constructs seeded with myofibroblasts (CF Myo) and vSMC (CF) exhibited completely differen t histological organization. CF Myo showed higher proliferative capacity forming clusters towards the inner layers of the scaffold, and vSMC were dispersed as individual cells across the scaffold When constructs were cultured under PF 12h, cells migrated towards lumen, but individual cells were still present on the ablumenal side. PF 24h had the highest cel l density across all conditions Quantification of dsDNA to determine cell density confirmed qualitative histological cellular differences, with no sig nificant difference in cellularity between CF /PF 12h, and PF 24h/CF Myo. Cellular metabolic activity was similar across all conditions. In response to 24 hours of physiological pulse frequency (PF 24h) cell growth increased significan tly, as seen in F igur e 5 8 Summary ECM synthesis and degradation is a continuous process during remodeling phase of the tissue engineering constructs. From a physiological perspective, improper balance of these processes result s in graft deterioration or thickening An ideal remodeled vascular graft should contain a functional media l layer which is crucial for normal vasoactive behavior and should have similar mechanical properties as a blood vessel such as tensile strength [9 ] [48 ] [83 86 ] [263 ] Obstacles arise from the manipulatio n of the smooth muscle cell phenotype during maturation period. Therefore, vascular smooth muscle cell plasticity is a key consideration vSMCs should be able to migrate and proliferate to create a smooth muscle cell layer while secreting ECM
104 components, a nd remodel the scaffold. After constructing the medial cell layer a shift to a contractile phenotype is important to possess vasoactivity and inhibit vSMC growth which may cause narrowing of the blood vessel. In these investigations v SMCs were seeded ont o an ex vivo scaffold and cultured under physiological and constant mechanical frequency. Our goal while comparing these different profiles was to understand how variability in the frequency affects vSMC phenotype even though average frequency values over each profile are similar. I nteresting ly, even though the average frequency is similar for both conditions PF 24h exhibited the highest vasoactivity was demonstrated. Contractility of CF (0.910.12 g) and PF 12 (0.76 0.06 g) were similar to some recent studies in the literature [256 ] [262 ] A n in vivo study by Neff et al de monstrated t he contractile function of vSMC and endothelial cell co seeded engineered blood vessels after 4 months of implantation in a sheep model. This was the first demonstration of enhanced contractile function of these grafts compared to only endothel ial cell seeded grafts in a sheep model [256 ] Similarly to our study Syedain et al. found that circumferential strain applied in incremental increases resulted in a nearly six fold increase in ultimate tensile strength compared to constant circumferential stress, demonstrating that cell adaptation is more dependent on continuous change rather than the rate of the strain [125 ] What are the mechanisms mediating the observed changes in response to variability in mechanical stimuli ? vSMC s have various r eceptor s that respond to shear, tensile, and compressive forces [264 266 ] Pulsatile stimulation is the dominan t force exerted on the medial layer of blood vessels It has been demonstrated by several
105 studies that i ntracellular pathways such as ERK phosphorylation, Akt/PKB phosphorylation, and Rho/ Rho kinase have well known roles in inducing synthetic vSMC pheno type as reviewed earlier [87 ] [267 ] [268 ] Mechanical forces can be transduced by mechanosensing machinery of the cell into intracellular signaling pathways and mechanical forces also modify the local chemical environments around the cell. Modification of the local chemical environment is an indirect mechanism of stress response [264 ] [269 ] Changes in flow modifies the biochemical environment around the cell and influences agonist concentrations such as ATP by in fluencing mass transport transiently from and to the cell surface [270 ] Therefore, changes in flow can modulate cellular processes by physically altering mechanosensing machinery as well as changing the agonist concentration around the cell with an indirect fashion. These investigations show that the temporal variability associated with normal physiological heart function (pulse frequency) has a profound effect on scaffold remodeling and vasoactive fun ction. T he development of tissue engineered grafts may require modulation in mechanical conditioning to optimize construct development. For example, physiological pulse frequency might be useful at the development stage of engineered constructs to induce proliferatio n, while conditions that promote a contractile phenotype can be used towards the end of the maturation period. Importantly, the continued cell proliferation and diminished vasoactivity of the PF 24 model has implications with the clinical pathology intimal hyperplasia, where uncontrolled vSMC proliferation and migration can result in graft occlusion. This is the first in vitro model that highlights vSMCs hyperplasticity in a 3D perfusion model. It is clear that the details
106 of the perfusion conditions, in pa r ticular pulse variability, play a significant role in both disease and regenerative medicine and as such warrant continued investigation.
107 Figure 5 1. Acronym, the description, and frequency profile of each bioreactor condition. Figure 5 2. Rep resentative images of vascular constructs after 6 weeks. A B C D
108 Figure 5 3. Comparison of representative load (B) and % strain at max stress (C) and Ultimate tensile strength (UTS)(D) of the constructs after 6 weeks o f culture. Arrow points out the difference between CF to DF 24h. DF 12h constructs exhibited trends similar to both CF and DF 24h 24h are significantly higher than all conditions (p<0.001, n=12). denote s significant difference compared to all other conditions (p<0.005,n=12).
109 Figure 5 4. the physiological range: After 6 weeks, CF Myo p resented increased stiffness (p< 0.025 ) and less extension at 120 mmHg compared to all other conditions cultured with vSMC (p < 0.001). Figure 5 5. Maximum contractile and relaxation force of the constructs. Continuous physiological frequency conditioning (PF 24h) significantly diminished con tractil e and relaxation force response (p<0.05).
110 Figure 5 6. SEM images showing longitudinal cross section of the scaffolds
111 Figure 5 7 Microscopic images of constructs. Scale bar correspond to 200 m.
112 Figure 5 8 Cell density and metabolic activity Cell density (above left) of cellular HUV constructs at week 6. Values indicate mean standard deviation. Metabolic Activity/cells (above right) of the vascular constructs with Alamar Blue dye. No s tatistical difference of metabolic activity /cell value was observed and values were consistent with our previous works
113 CHAPTER 6 RELATIVE GENE EXPRESSION ANALYSIS OF ENGINEERED SCAFFOLDS IN RESPONSE TO PHYSIOLOGICAL FREQUENCY STIMULI Chapter 5 highlight s t he short term variation of pulse rate such as elevations observed in response to increased respiratory activity This short rate modulation closely resembling the changes that occur during physical exercise was included in o ur perfusion system to prod uce more physiological relevant stimuli. The broad changes in scaffold architecture, biomechanics, vascular smooth muscle cell function ality, and cell distribution were investigated in the C hapter 5 .This chapter further examines the changes that occur in t he cellular phenotypic level. The relative ch anges in the mRNA expression were quantified and compared to CF (constant mechanical stimuli). The u nique p henotypic shift capacity of vSMCs and its relevance to vascular tissue engineering was described in the introduction section. The biological function of each selected gene and its relevance is described in this chapter. Table 6 1 displays se lected vSMC marker genes and their protein function. R elative expression of these proteins is used to characterize v SMC s in this study The purpose of this work is to study the effect of physiological pulse frequency on vascular smooth muscle plasticity with an emphasis on extracellular matrix remodeling, cell adhesion, and cytoskeletal, contr actile vSMC marker genes The smooth muscle myosin heavy chain and smoothelin are the only two markers that are closely associated with the vSMC phenotype [90 ] These gene s have not been detected in non smooth muscle cells. However the degree of the phenotype shift should be investigated with a wider variety of genes that are associated with these changes and also scaffolds should be investigated for vasoactive functional ity (i.e. response to pharmacological drugs) [91 ]
114 The e xtracellular matrix has pro found effects on the vSMC phenotype. For instance, while collagen hydrogels are promoting a more synthetic phenotype, fibrillar collagens are found to promote a more contractile phenotype [271 ] Matrix met alloproteinases (MMPs) are proteases that are able to cleave and have the potential to degrade all matrix proteins. They play central roles in vascular pathologies, and also tissue repair/remodeling. T heir activities are regulated by tissue inhibitors of metalloproteinases (TIMPs) [127 ] [272 ] Along with collagen type I and type IV, laminin and fibronectin indu ces a contractile phenotype [51 ] [271 ] Osteopontin is also only observed during vascular pathologies such as during neotintima formation [273 ] List of selected extracellular matrix genes and their description that are used in this study can be seen in Table 6 2. vSMCs interact with other cell types, and extracellular environment via adhesion mol ecules [273 ] Adhesion molecules such as integrins are affected by the ECM composition of the scaffold that cells are surrounded by When the ECM composition changes, expression of integrins and their composition change; these changes ultimat ely affect the vSMC phenotype. Integrins are a group of heterodimer proteins consis ting of two sub and They play a key role in cell to cell and cell to matrix adhesions The i ntegrin family consists of the major cell matrix adhesion molecule s that provide a link between cytoskeletal proteins and are dominantly expressed in smooth muscle [274 ] [275 ] is associated with the are laminin binding integrins that are found to be closely associated with the contractile phenotype [276 ] [277 ] I CAM and V CAM are known for their important role when expressed in endothelial cells. However,
115 these markers of inflammation are only known to be expressed in a disease state with in v SMCs enabling macrophage infiltration [278 ] [279 ] .Syndecans are adhesion transmembrane receptor s which belong t o the proteoglycan family. They also play an import ant role in cell cell and cell matrix adhesion. All syndecans ( 1, 2, 3, 4) are found in vascular smooth muscle cells However, synde can s 1 and 4 are upregulated in response to vascular injury [273 ] These selected adhesion molecules closely associated with vSM C plasticity are shown in T able 6 3 Materials and Methods RNA Extraction a nd Reverse Transcription RNA was extracted by using a fi brous RNA tissue kit (Qiagen, CA ). Each RNA sample was mixed with the genomic DNA elimination mix and incu bated for 5 min utes at 42C. Afterwards, the samples w ere revers e transcribed to cDNA using RT2 qPCR Master mix (SA Biosciences, Austin,TX). The r everse transcription mix was incubate d at 42C for exactly 15 min utes, then the reaction was immediately stopp ed by incubating at 95C for 5 min utes These thermal cycling procedures for reverse transcriptase reactions were Real Time PCR Quantification of the relative mRNA expression was performed using a Bio Rad CFX system. Custom designed PCR arrays were purchased from SA Biosciences (CA, USA). For each analysis, cDNA samples were mixed with the RT2 qPCR Master mix and distributed across the PCR well plates. After cycling with Bio Rad CFX system, the obtained am plification data was analyzed using the Relative Expression Software Tool (REST). The following protocol for PCR was used: denaturation at 95C for 15 s and annealing/extension at 60C for 30 s for 40 cycles. After amplification, a melting curve
116 was be obt ained by slowly heating at a 10 minute ramp time from 60C to 95C with continuous fluorescence detection. Pubmed reference sequence number of each gene used in this study can be seen in T able 6 4. Results Fold regulation and their standards errors of each gene use d in this study can be seen in T able 6 5 and 6 6 Schematic representations of each gene group are displayed in Figure 6 1, 6 2, and 6 3. Schematic representation of relative expression of cytoskeletal, con tractile and transcription factors of vSM C marker genes are shown in Figure 6 1 The expression by PF 12, and PF 24h constructs were assessed relative to CF. All the genes were down regulated in PF 24h samples, with only MYH11 be ing up regulated. TAGLN VCL, and ACTA2 were down regulated in PF 12h constructs, as see n in Figure 6 1 The contractility t esting results in C hap ter 5 have shown that PF 24 did not exhibit any contractile response and PF 12h conditions contractile response was similar to CF. A s chematic representation of relative expression of extracellular and remodeling genes is shown in Figure 6 2 In PF 24 samples almost all the extracellular matrix and remodelin g genes were upregulated, though COL4A1 was downregulated, and TIMP2 and TI MP 3 were statistically similar to constant stimuli controls. Conversely in the PF 12 sample, MMP 2,3, and 9 were upregulated and MMP1 was down regulated and all ECM genes did not statistically deviate from constant stimuli controls. ECM synthesis and degradation is a continuous process that occurs at the remodeling phase of the tissue engineering constructs. From a physiological perspective, improper balance of these processes result s in graft deterioration or thickening. When MMPs, TIMPs are downregulated, this downregulation might have similar effects wh ere these genes are
117 upregulated as in PF 24h sample. Therefore, interaction of these molecules is not always in a stoichiometric ratio. However, it might be hypoth esized that upregulation of the ECM and remodeling genes PF 24 might be correlated with activ e scaffold remodeling, in whic h these conditions resulted in the highest mechanical properties. The PF 12h condition broadly resulted with similar ECM expression to controls and upregulated TIMP and MMP expression. A s chematic representation of relative ex pression of cell cell, and cell matrix adhesion genes is shown in Figure 6 3 Reduced associated with a proliferative phenotype. In our study, the 1 and 7 subunit had a downregulation at PF 24 h. Similarly upregulation of SDC1 and vCAM might be a sign of a synthetic phe notype of these cells. SPP1 was upreg ulated to a higher level in the PF 24 samples compared to PF 12h. SPP1 encodes for osteopontin which is found only in diseased blood vessels (exhibited to a degree in calcification and atherosclerosis). Summary The present study evaluated the effect of phy siological frequency stimuli within engineered constructs and compared it with constant frequency stimuli at the end of 6 weeks. The objective wa s to study the effect of physiological frequency on vascular smooth muscle plasticity with an emphasis on : i) E xtracellular matrix remodeling, ii) Cell adhesion, and iii) C ytoskeletal, contractile and transcription al level of vSMC marker genes The background regarding the influence of mechanical stimuli on v SMC s and vSMC plastic ity were reviewed in the C hapter 1 The expression patterns of a wide range of vSMC s markers have been characterized to describe the phenotypic state of these cells. Contractile v SMCs which are dominant in blood vessels exhibit a mature contractile apparatus including smooth
118 muscle alpha a ctin, myosin heavy chain, calponin, SM22, and smoothelin. Relative expression of these markers can be used to localize v SMCs on the contractile to synthetic range. In this study, it has been found that physiological frequency has profound effects on the vS MC phenotype. The general trend in the gene expression supports the functional cellular data acquired and discussed in the C hapter 5 Data regarding biomechanical properties and the ECM remodeling genes might support the relative extensive remodeling of th e PF 24 condition. For instance, compared to constant stimuli, physiological constructs had a higher expression of VCAM1, and SPP1 (osteopontin) expression and lower expression of 1 integrin, which may indicate a disease or synthetic state of v SMCs Chang es in these cell adhesion related genes in relation to mechanical stress may regulate the internal signaling cascades differently. Functional contractility is the most important indicator of the vSMC phenotype I n PF 24h samples ECM associated genes such a s collagen type I, III were significantly actin and smoothelin, were downregulated compared to CF. This was shown with direct analysis of vessel contractility, where maximum contraction and the relaxation f orce diminished significantly when 24 h of physiological pulse stimuli (PF 24h) was applied. The relative gene expression results presented here have demonstrated the big picture of remodeling events at a particular point in time. As the scaffold remodels overtime, it induces an evolving mechanical environment which rapidly changes the cell behavior through time. It s hould be considered that this study does not provide a detailed account explaining the variat ion in gene expression patterns; we considered
119 th e data a valuable component that help s describe these initial interactions to be used as a reference for future studies working on similar topics. While making statements about the relationship between gene expression data and protein expression important differences should be considered: r egulation of some of these genes may occur at multiple levels, i.e. though gene transcription, posttranslational modification and protein level (e.g secreted TIMPs inhibit MMPs ) [280 ] [281 ] Our study and pr evious studies have shown that the m echanical environment plays an important role in the modulation of the vSMC pheno type. Several groups have demonstrated that v SMCs can infiltrate, grow within and remodel within decellularized scaffolds [113 ] [154 ] [167 ] [169 ] [227 ] [228 ] Although bioreactors and several other cell culture systems with various scaffolds have been developed to study these events, the precise big picture events from outer cellular sensing via cell receptors to cytoskeleton to translation to intracel lular pathways have not been investigated extensively. To add depth, this study included different group s of genes closely related with the vSMC phenotype and ECM remodeling. As an example, we studied the transcription factors that regulate vSMC genes as w ell as cytoskeletal components and cel l adhesion receptors. Myocardin and SRF are genes that regulate v SMCs in the transcription level and are central regulators of the vSMC specific expression [281 [282 [283 ] Furthermore, myocardin itself can regulate the expression of most contractile marker genes [282 ] v SMCs in the medial layer of m ature blood vessels proliferate and secrete ECM at a very low rate, whereas v SMCs in vitro are in a more synthetic state. The degree of being s ynthetic is variable in the contractile synthetic spectrum because the response
120 to mechanical stimuli of v SMCs resulted in differences in gene expression profiles and ultimately different vascular construct properties [280 ] These properties of engineered blood vessels are dependent on v SMCs to breakdown ECM and depos it newly synthesized proteins. The d escriptive nat ure of this work results from a better understand ing of cellular remodeling mechanisms in response to adapt ing mechanical stimuli. Progressive remodeling of the scaffold to a more complex natural state is a dynamic process that will result in continuous changes in gene expression and the structural characteristics of the ECM These structural changes a ffect th e mechanical forces imposed on cells which also regulate the cell ular phenotype [284 ] Cells can adapt to their environment by remodeling their surrounding matrix. This occurs by xternal forces from reorganization of the initial scaffold This process occurred because constructs have t otally different macroscopic appearances between initial culture sequences to the culture termination time. As such, it is likely that the process of remodeling also changes gene expression profiles due to changin g forces exerted on cells and als o mass transfer of key nutrient change s during these processes. In addition, one thing that was not considered in this study is the evaluation of the crosslink between collagen fibrils. Crosslinking stabilizes the collagen fibrils and ultimately makes the construct stronger, and less prone to be attacked by proteases. crosslinking within our constructs may explain a further degree of remodelin g, and the differences between the biomechanical properties of the scaffold [85 ] Another element
121 that was not a part of this study is the vascular smooth muscle heterogeneity which is a very important factor for remodeli ng and disease progression in order to identify the subpopulations within this broad spectrum of phenotypes. This topic will be addressed in the future directions section Molecular regulation of v SMCs will be an important factor since its phenotype shift is not bla ck and white, but more likely has a broad spectrum of various possible phenotypes in between [90 ] Even though the great regenerative potential of v SMCs is an advantage to engineer ing vascular t issues, conditions that lead to a disease state such as intimal hyperplasia must be considered in experimental settings. While seeding v SMCs on 3D scaffolds in vitro and implant ing these scaffolds in relevant animal models, we need to expand our understand ing of how combinations of these external factors exerted on v SMCs affect their own molecular regulation machinery.
122 Table 6 1 S e lected vSMC marker genes and their protein function [50 ] [51 ] [90 ] [91 [280 ] [281 ] [285 ] [286 ] [287 ] [288 ] Gene name Protein name Function Notes ACTA2 actin Contractile M ajor constituen t in the contractile apparatus First identified and the m ost studied vSMC marker protein MYH11 Smooth muscle myosin heavy chain Contractile Major contractile protein that converts chemical energy into mechanical energy O ne of the most definitive marker of vSMC phenotype SMTN Smoothelin Cytoskeleton A structural protein that associates with stress fibers and const itute s part of the cytoskeleton E xpressed exclusively in c ontractile smooth muscle cells TAGLN SM Cytoskeleton A n actin stress fib e r associated protein VCL Vinculin Cytoskeleton A cytoskeletal protein that is associated with membrane actin filament attachment sites CALD1 Caldesmon Contractile Thin filament a ctin binding protein R egulates smo oth muscle contraction by increasing the stabilizat ion of actin filament structure CNN1 Smooth muscle calponin Contractile Smooth muscle specific, actin tropomyosin and calmodulin binding thin filament protein Modulates smooth muscle contraction MYOCD [282 ] [289 ] Myocardin Transcription Smooth muscle transcriptional coactivator that physically associates with SRF Can activate many smooth muscle specific genes via SRF SRF [282 ] [290 ] Serum binding factor Tra nscription Blocks expression of many smooth muscle specific proteins P romotes cell migration and proliferation
123 Table 6 2 List of selected extracellular matrix genes and their general description [1 ] [52 ] [272 ] [291 ] Gene Protein Function COL1A1 Type I Collag en E ncodes for the pro a lpha1 chains of type I collagen A fibril forming collagen found in most connective tissues COL3A1 Type III Collagen E ncodes the pro alp ha1 chains of type III collagen A fibril forming collagen Found inelastic tissues such as skin and blood vessels Fibronectin 1 Fibronectin E ncodes for fibronectin I nvolved in cell adhesion/migration, wound healing ELN Elastin ELN encodes for tropoelastin that isone of the components of elastic fibers P rovides flexibility to connective tissue s C OL4A1 Type IV Collagen, E ncodes for the major protein within the type IV alpha collagen M ainly found in basement membranes MMP1 Matrix Metalloproteinase 1 MMPs have distinct but alsotheir specificities overlap Initiates break down of the interstitial co l lagens, types I, II, and III K nown as collagenase MMP2 Matrix Metalloproteinase 2 Degrades mainly type IV collagen MMP3 Matrix Metalloproteinase 3 Mainly degrades fibrone ctin, laminin, collagens III, IV, IX, and X, and proteoglycans MMP9 Matrix Metalloproteinase 9 Mainly d egrades type IV and V collagens TIMP1 TIMP metallopeptidase inhibitor 1 I nhibitory role against most of the known MMPs TIMP2 TIMP metallopeptidase inhi bitor 2 Inhibitory role against most of the known MMPs TIMP3 TIMP metallopeptidase inhibitor 3 Inhibitory role against most of the known MMPs
124 Table 6 3 Comparison of the adhesion receptors involvement in vascular smooth muscle cell functions Adhesion receptor Contractile phenoype Synthetic phenotype SPP1 [292 ] Absent Expressed ITGB1 [274 [275 ] Expressed Downregulated ITGA1 [276 ] Expressed Downregulated ITGA7 [277 ] Expressed Downregulated ICAM1 [278 ] [279 ] Absent Expressed VCAM1 [278 ] [279 ] Absent Expressed SDC1 [273 ] Expressed Upregulated SDC4 [273 ] Expressed Upregulated
125 Table 6 4 Pubmed reference sequence number of each gene used in this study Gene Symbol Refseq # Off icial Full Name SPP1 NM_000582 Secreted phosphoprotein 1 ITGB1 NM_002211 Integrin, beta ITGA1 NM_181501 Integrin, alpha 1 ITGA7 NM_002206 Integrin, alpha 7 ICAM1 NM_000201 Intercellular adhesion molecule 1 VCAM1 NM_001078 Vascular cell adhesion mole cule 1 SDC1 NM_002997 Syndecan 1 SDC4 NM_002999 Syndecan 4 ACTA2 NM_001613 Actin, alpha 2, smooth muscle, aorta MYH11 NM_022844 Myosin, heavy chain 11, smooth muscle SMTN NM_006932 Smoothelin TAGLN NM_003186 Transgelin VCL NM_003373 Vinculin CALD1 NM_004342 Caldesmon 1 CNN1 NM_001299 Calponin 1, basic, smooth muscle MYOCD NM_153604 Myocardin SRF NM_003131 Serum response factor PDGFA NM_002607 Platelet derived growth factor alpha polypeptide COL1A1 NM_000088 Collagen, type I, alpha 1 FN1 NM_00 2026 Fibronectin 1 ELN NM_000501 Elastin COL3A1 NM_000090 Collagen, type III, alpha 1 MMP1 NM_002421 Matrix metallopeptidase 1 MMP2 NM_004530 Matrix metallopeptidase 2 MMP3 NM_002422 Matrix metallopeptidase 3 MMP9 NM_004994 Matrix metallopeptidase 9 COL4A1 NM_001845 Collagen, type IV, alpha 1 ANGPT2 NM_001147 Angiopoietin 2 TIMP1 NM_003254 TIMP metallopeptidase inhibitor 1 TIMP2 NM_003255 TIMP metallopeptidase inhibitor 2 TIMP3 NM_000362 TIMP metallopeptidase inhibitor 3 RPL13A NM_012423 Ribo somal protein L13a GAPDH NM_002046 Glyceraldehyde 3 phosphate dehydrogenase ACTB NM_001101 Actin, beta
126 Table 6 5 F old regulation and their standard err or values of each genein PF 12h samples. Schematic representation with statistical significance of ea ch gene group is displayed in F igure 6 1, 6 2, and 6 3 PF 12h Fold regulation Standard error SPP1 0.062581984 0.045757491 0.12969 ITGB1 0.693375151 0.255513713 0.932879 ITGA1 1.531989551 1.021189299 1.905218 ITGA7 0.139661993 0.237321436 0.04191 ICAM1 1.606241594 1.198574618 2.225263 VCAM1 1.147428969 1.005437923 1.374437 SDC1 0.450557009 0.215637563 0.81278 SDC4 0.075546961 0.106238238 0.199481 PDGFA 1.148170613 0.95650459 1.488607 ACTA2 1.537602002 1.585026652 1.4437 MYH11 0.002374691 0.018681334 0.104862 SMTN 0.12515583 0.000867722 0.329398 TAGLN 0.598599459 0.66756154 0.53018 VCL 0.412289035 0.519993057 0.24489 CALD1 0.159893906 0.28567024 0.001301 CNN1 0.130549 0.103122148 0.49153 MYOCD 0.002868824 0.029235154 0.071931 SRF 0.006893708 0.101823517 0.207365 MMP1 0.314258261 0.339388764 0.40044 MMP2 1.122674513 0.084014619 2.365464 MMP3 1.363104066 0.941829275 3.269671 MMP9 2.432677526 0.800817927 5.353415 TIMP1 0.728516105 0.036286269 1.485228 TIMP2 0.240549248 0 .186702821 0.58139 TIMP3 0.13815583 0.037727019 0.33258 COL1A1 0.003891166 0.041521831 0.084518 COL3A1 0.091514981 0.112605002 0.08719 COL4A1 0.126055432 0.221831767 0.045893 ELN 0.105349 0.091088497 0.313479 FN1 0.067526235 0.045419387 0.09255
127 Table 6 6. F old regulation and their standard err or values of each gene in PF 24h samples Schematic representation with statistical significance of e ach gene group is displayed in F igure 6 1, 6 2, and 6 3. PF 24h Fold regulation Standard error SPP1 0.638389408 0.61637 0.671265 ITGB1 0.240332155 0.27246 0.21396 ITGA1 0.512817759 0.3638 0.607777 ITGA7 2.267752649 3.86375 0.52996 ICAM1 0.059941888 0.1209 0.153205 VCAM1 2.565527881 2.430073 2.696463 SDC1 1.059108818 0.913655 1.237971 SDC4 0.769551079 1.01323 0.53018 PDGFA 0.246005904 0.202216 0.301898 ACTA2 0.761953897 0.78252 0.72125 MYH11 0.5289167 0.396896 0.663512 SMTN 0.571865206 0.75449 0.44612 TAGLN 0.262807357 0.30103 0.22403 VCL 0.049605613 0.003891 0.097604 CALD1 0.066006836 0.22768 0.031408 CNN1 0.906578315 1.00877 0.80688 MYOCD 1.431798276 1.72125 0.86646 SRF 0.155522824 0.25104 0.08566 MMP1 1.563267445 0.064309 3.208984 MMP2 1.254209597 0.105514 2.654679 MMP3 3.55337322 0.151246 7.267088 MMP9 0.876737297 0.219395 2.140137 TIMP1 0.992686039 0.027673 2.029554 TIMP2 0.144574208 0.094939 0.371947 TIMP3 0.115295805 0.034867 0.369736 COL1A1 0.492900011 0.033357 1.069702 COL3A1 1.584025754 0.085316 3.271831 COL4A1 0.681936665 0.194212 1.10791 ELN 2.340372718 0.334142 5.041713 FN1 0.131939295 0.042741 0.302494
128 Figure 6 1. Schematic representation of relative expression of cytoskeletal, contra ctile and transcription factor vSMC marker genes: Relative gene expression of PF 12h, Pf 24h anf CF Myo compared to CF. Mean values of fold change values were displayed on the left side of the panel. Rows denote statistical downregulation, rows + ( p < 0.05, n = 9 )
129 Figure 6 2. Schematic representation of relativ e expression of extracellular matrix and cellular remodeling marker genes. Relative gene expression of l PF 12h, Pf 24h anf CF Myo compared to CF. Mean values of fold change values were displayed on the left side of the panel. Rows denote statisti cal downregulation, rows + ( p < 0.05, n = 9 )
130 Figure 6 3. Schematic representation of relative expression of cell cell adhesion, cell matrix adhesion related marker genes. Relative gene expression of l PF 12h, Pf 24h anf CF My o compared to CF. Mean values of fold change values were displayed on the left side of the panel. Rows denote statistical downregulation, rows + ( p < 0.05, n = 9 )
131 CHAPTER 7 CONCLUSIONS AND FUTU RE WORK Conclusions The lo w patency of small diameter vasc ular grafts is the result of multiple factors. These factors can be broadly grouped into thrombogenic, immunogenic and hyperplasctic responses. Poor clinical outcomes shifted the attention of the scientific community towards creating functional blood vessel s rather than inert tubes However, engineering a vessel that possesses native vascular composition and mechanics, as well as physiological function is a significant technical challenge. These limitations are related to the manipulation of the endothelial cell and smooth muscle cell function within an in vitro environment. Each of these limitations is still a challenge and must be addressed in order to achieve clinical success. Otherwise, grafts that are seeded with vascular cells can fail due to similar negative biological reactions. From a vascular smooth muscle cell perspective, an ideal remodeled vascular graft should contain a functional medial layer, which is crucial for normal vasoactive behavior and should have simil ar mechanical properties, such as tensile strength as a blood vessel. Obstacles arise from the manipulation of the smooth muscle cell phenotype during the maturation period. Therefore, vascular smooth muscle cell plasticity is a key consideration. vSMCs sh ould be able to migrate and proliferate in order to create a smooth muscle cell layer while secreting ECM components that remodel the scaffold. After constructing the medial cell layer, shifting to a contractile phenotype is important because the vSMCs mus t possess vasoactivity and their growth should be inhibited. Failure to inhibit growth may cause narrowing of the blood vessel. A major challenge in developing in cell based vascular graft is reverting the phenotype
1 32 from proliferative to a more mature phen otype before implanting the graft. Therefore, studying the aspects of the in vitro conditioning of tissue engineered small diameter blood vessels is crucial to understanding how the phenotype reversion can be achieved. Many of the investigations discuss ed in Chapter 3, 4, 5, and 6focus on materials development and how biodegradable materials interact with cellular systems to regenerate blood vessels. The use of ex vivo materials as either direct implants or scaffolds for tissue regeneration has increased significantly over the last decade due largely to favorable tissue interactions. In this dissertation, I have assessed the use of the human umbilical vein as a decellularized scaffold to function as a potential small diameter vascular graft. Figure 7 1 sho ws how each experimental Chapter 3, 4, 5, and 6 are progressed. Current in vitro cultivation methods are limited by a number of factors as described in Chapter 1. These include cell source, inefficient cell seeding long culture periods, a nd scaffold type. In addition, the in vitro cultivation systems developed so far assume vascular smooth muscle cells have a proliferative ( wound healing like ) phenotype. In this dissertation, the specific factors that are related to these limitations have been studied. F irst, using ex vivo tissues as an acellular scaffold is advantageous because they are already in the shape of a blood vessel and consist of natural ECM components, which give appropriate mechanical properties and biocompatibility. The automated dissection system described in Chapter 2 made it possible to dissect HUV with mechanical properties and uniform wall thickness a significant improvement over manual dissection.
133 Chapter 3 focused on tissue engineered blood vessel limitations that occur in the initi al culture sequence due to low efficiency cell seeding methods that require a large cell number but results in low, non uniform cell attachment. In order to improve cell seeding, hydrogels were used as a cell delivery vehicle for inoculating cells into the ablumenal void space of the bioreactor, where they contracted around the scaffold. For this technique, the contraction of hydrogels around porous soft tissue required optimization. Transmural flow and hydrogel swelling can be problematic due to the rate o f transmural and ablumenal flow (in the bioreactor) in the initial culture sequence. Optimization of the lumenal and ablumenal flow profiles during the initial culture stage was an important factor in enhancing the uniformity of hydrogel compaction because cell delivery is fundamental to all of the studies included in this dissertation. Creation of a cell dense engineered construct has been an obstacle for decellularized tissue scaffolds as seeded cells are usually limited to the outer most surfaces, result ing in extended culture periods to overcome poor cellular infiltration. This is due largely to poor transport conditions that often limit cell migration. Chapter 4 focused on the key parameters that influence cell migration in order to design systems that i mprove cellular infiltration and reduce the in vitro culture time required for construct development. These investigations compare the influence of directed nutrient gradients by precisely contr olling perfusion conditions to drive cell migra tion during ear ly remodeling events. These investigations have shown that by modulating perfusion culture conditions, the density and the distribution of the viable cells can be regulated to improve graft tissue biomechanics and graft maturation.
134 In Chapter 5 acellular HUV was used to test the hypothesis that short term variation in heart rate, such as due to elevations in respiratory activity, can modulate vascular function and remodeling. If so, the variability in vascular pulse would produce a more physiologically rel evant system to investigate vascular cell function, tissue regeneration, and the effects of pathological conditions. These i nvestigations show ed that the temporal variability associated with normal physiological heart function (pulse frequency) had a profo und effect on scaffold remodeling and vasoactive function. While seeding vSMCs on 3D scaffolds in vitro we need to expand our understanding of how combinations of these external factors exerted on vSMCs affect their own molecular regulation machinery. In Chapter 6, the effect of the temporal variation in mechanical stimulation on vSMC phenotype was studied. In PF 24h samples ECM associated genes such as collagen type I, III were significantly upregulated and all vSMC gene markers for contractility such as actin and smoothelin, were downregulated compared to CF. This was shown with direct analysis of vessel contractility, where maximum contraction and the relaxation force diminished significantly when 24 h of physiological pulse stimuli (PF 24h) was appli ed Physiological pulse frequency might be useful at the development stage of engineered constructs to induce proliferation, while conditions that promote a contractile phenotype can be used towards the end of the maturation period. Importantly, the contin ued cell proliferation and diminished vasoactivity of the PF 24h model has implications with the clinical pathology intimal hyperplasia, where uncontrolled vSMC proliferation and migration can result in graft occlusion. It is clear that the details of the perfusion conditions, in particular pulse variability, plays a
135 significant role in both disease and regenerative medicine and as such warrant continued investigation. Future Work These findings show that studying the biomechanical properties of engineered blood vessels under conditions that broadly mimic in vivo mechanics is a useful for gaining better und erstand ing of in vitro remodeling processes. Although some parallels can be drawn with in vivo remodeling, these in vitro investigations, where single cel l populations can be assessed independently, are designed to be more progressive in nature. As such, the remodeling processes in vivo are considerably more complex, with contributions by other cellular systems, including the host immune response. The curre nt work can be taken in di fferent directions including in v itro or in vivo studies. The nex t section will summarize proposed future research objectives. Combinations o f Physiologic al a nd Constant Frequencies a s Mechanical Stimuli Physiological mechanical s timuli might be useful at the development stage of engineered constructs, while conditions that promote a contractile phenotype can be used to regulate the arterial tone (i.e. constant stimuli) By combining these two stimuli towards the end of maturation period, a recellularized in vitro graft but with a more quiescent phenotype and enhanced mechanical properties may be obtained. In V ivo A cellular Implantation S tudy Gluta raldehyde fixed umbilical veins are widely used in clinic s and we re reviewed in the introduction However an acellular, non crosslinked human umbilical vein has never been used in an in vivo study before. A further assessment of the acellular human umbilical vein, using an animal model, and following over a long term implantation is impor tant in terms of an acellular small diameter vascular graft in humans.
136 In V itro Blood Vessel Maturation F ollowed b y a n In V ivo S tudy In vitro maturation in the bioreactors should be followed by an in vivo implantation study in an animal model. Further in v ivo studies will provide insight into these more complex remodeling processes that may improve our unders tanding of both heterologous tran splant graft remodeling and engineered vessel regeneration Different Approaches f or Scaffold P reparation Freeze dryin g might be a good option f o r exploring product availability and reducing the cost and time in terms of commercialization. The downside is that the glycoaminoglycans will be diminished significantly due to the freeze drying process. However, from a clinical perspective it would reduce the time of preparation of the tissue engineered graft by approximately one week. This one week contains the time for HUV collection, freezing, lathing, and decellularization of the scaffolds. Exploring Different Cell S ources T he choice of optimum cell source has been one of the main obstacles in the field. Replicative capacity of adult cells has also significantly limited the clinical applications of small diameter tissue engineered blood vessels. Autologous human stem and prog enitor cells from a range of sources have received significant attention due to their potential to differentiat e into EC or SMC lineages. I believe that the model sys tem I developed during my PhD will work with stem cells (instead of vSMCs) to engineer sma ll diameter blood vessels. While the effects of different stem cell culture techniques have been assessed previously, few studies have detailed changes in the material s biomechanical properties. Further, no other study has yet looked at the long term cult ure effects on cell phenotype while cell mediated remodeling modifies the scaffold biomechanical properties.
137 vSMC Heterogeneity w ithin t he S caffold v SMCs in a vessel can cover the whole spectrum of phenotypes as v SMC s can modulate their phenotypes as a co nsequence of environmental factor s. In a vascular graft, vSMC heterogeneity is expected to occur e. g. due to nutrient gradients through the scaffold. While this study and many other s assess ed global gene expression within these grafts, cellular heterogenei ty might be assessed with respect to different locations. Therefore, l aser micro dissection ( LMD ) would be an effective tool in separating gene responses in different locations ( intima vs media ) of functioning vein grafts. Economical P erspectives As a tiss ue engineer, while we are doing the basic science to understand the interactions of various cells and scaffolds t he main goal is to carry these technologies to the hospital bench, and ultimately construct a completely biological vascular graft. One of the obstacles in carry ing the technologies to the clinic is economical feasibility. evolve to become more affordable with time (e.g. MRI), whereas an inexpensive technolog y that marginally works will be abandoned [138 ] I believe that the main focus still should be o n creating a cellular product with a superior p erformance compared to what has been used clinically. Tissue engineering problems can be overcome t h rough innovative thin king and m ultidisciplinary collaboration will be the key to these innovative approaches where engineers, biologists, physical scientists and clinicians work together [293 ]
138 Figure 7 1. Pr ogression of each experimental c hapter
139 APPENDIX A COMPOSITE SWNT COLLAGEN MATRIX: CHA RACTERIZATION AND PRELIMINARY ASSESSME NT AS A CONDU CTIVE PERIPHERAL NER VE REGENERATION MATRIX Unique in their structure and function, single walled carbon nanotubes (SWNT) have received significant attention due to their potential to create unique conductive materials. F or neural applications these condu ctive materials hold some promise as they may enhance regenerative process es. However, like other nano scaled biomaterials i t i s important to have a comprehensive understanding how these materials interact with cell systems and how the biological system re sponds to their presence These investigations aim to further our understanding of SWNT cell interactions by assessing the effect SWNT/collagen hydrogels have on PC12 neuronal like cells seeded within and (independently) on top of the composite material. T wo types of collagen hydrogel s were prepared: 1) SWNT dispersed directly within the collagen (SWNT/COL), and 2) Albumin to improve dispersion (AL SWNT/COL), and collagen alone serving as a control (COL). SWNT dispersion was significantly improved when using surfactant assisted dispersion. The enhanced dispersion resulted in a stiffer, a more conductive material with increased collagen fiber diameter. Short term cell interactions with PC12 cells and SWNT composites have shown a stimulatory effect on cell proliferation relative to to plain collagen controls. In parallel to these results, p53 gene displayed normal expression levels which indicate an absence of nanoparticle induced DNA damage. In This work was published at the Journal of Neural Engineering. Further deta ils of this work can be found in the journal website : http://iopscience.iop.org/1741 2552/7/6/066002/ 
140 summa ry, these mechanically tunable SWNT collagen scaffolds show the potential for enhanced electrical activity and have shown positive in vitro biocompatibility results offer ing further evidence that SWNT based materials have an important role in the biomedica l field. Figure A 1 Collagen gel preparations : A) COL, B) SWNT/COL, and C)Al SWNT/COL
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165 BIOGRAPHICAL SKETCH Zehra Tosun was b orn in Ankara, Turkey. She graduated from Suleyman Demirel Ankara University in 2006. After completing her bach ngineering S he enrolled in the Ph.D. program of the School of Chemical Biological Materials Engineering at the University of Oklahoma in 2006, focusing on tissue engineering, with a focus on peripheral nerve regeneration and contr olling nerve function with a conductive composite using single walled carbon nanotu be. In 2010 her advisor Dr. Peter S. McFetridge relocated his research program to the University of Florida, and she cal engineering program where her research focused on vascular regeneration. Zehra received her Doctor of Philosophy degree in Biomedical Engineering in August 2012.