| Collection Mémoires et thèses électroniques | |
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| Chapter 1: Introduction | ||
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Table des matières
The human body is a complex assembly of organs and tissues. Like any living organism, it and its components are not meant to last indefinitely. These organs can become damaged or diseased affecting our quality of life. The arterial system is no exception. In fact, cardiovascular diseases are the single most important cause of death in the world[1]. When not treated, these diseases which affect the heart and blood vessels, end in heart attacks or strokes. Cardiovascular diseases take many forms including coronary heart diseases such as atherosclerosis and hypertension, heart muscle disease, arrhythmias, and valve disorders[2]. Atherosclerosis for example is characterized by the accumulation of lipids, cells and extra-cellular matrix molecules in the vessel wall[3]. Atherosclerosis of coronary or peripheral vascular arteries is the largest cause of mortality in Canada and in all developed countries[4]. Symptoms include angina, or chest pain, and vascular deficiencies. Often, cases of atherosclerosis are not detected until complete occlusion results in heart attack or stroke. In many cases, pharmaceutical treatments are not sufficient and more drastic measures must be taken in order to return adequate blood flow through the arteries and minimize the risk of occlusion and embolism.
For coronary arteries or peripheral arteries such as that below the knee, autologous graft transplantation is the option of choice. The grafts usually consist of either mammary artery or saphenous vein harvested from the patient. Unfortunately, all too often, autologous grafts are not always available due to their limited number and their insufficient quality. Repeated procedures and multiple bypasses are also problematic since all vessels appropriate for replacement may have been employed during previous interventions. Furthermore, saphenous vein in elderly patients is prone to thrombi, neointimal formation, atherosclerosis or aneurysm[5]. Veins also lack vasomotor tone. Moreover, harvesting veins and arteries leaves wounds that can break down and become infected. Another alternative used in the past was fresh allografts. However, these allograft arteries or veins are no longer used for coronary bypass surgery due to poor patency, rejection complications, endothelial cell sloughing and reaction with leukocytes, and loss of cellular reactivity[6].
For large diameter arteries, synthetic prostheses made of Teflon or Dacron have been a relatively viable solution for the past few decades. These prostheses perform reasonably well in high-flow and low resistance conditions[5]. But their use is limited to vascular conduits larger than 5-6 mm in diameter due to the high risk of thrombus formation, embolism and occlusion associated with their use in smaller vessels. Despite extensive research on synthetic vascular prostheses, these have yet to prove suitable for long term implantation. In fact, 65% of them must be explanted in the 10 years following implantation[7]. Synthetic prostheses pose a high risk of foreign body reaction and extensive use of anticoagulant/antithrombotic control is often required[5]. Synthetic prostheses also lack a confluent endothelium. This single cell layer is what confers haemostatic properties to natural arteries rendering them antithrombotic[8]. Therefore, the importance of creating a suitable endothelium on the luminal surface of any small diameter vessel substitute is paramount. Research into surface modifications with coatings of proteins, polymer materials or cells is being pursued to increase the bioactivity of prostheses in order to render then more suitable for endothelial cell seeding. These approaches have been met with mixed success[5]. Also, these linings do not provide vital vascular functions such as vascular responsiveness or other biological secretory functions seen with normal blood vessels[8].
Recently, stents have been in clinical use when performing balloon angioplasty in order to unclog blocked arteries using interventional surgery as opposed to classical surgery methods. The stent remains in the artery upon removal of the balloon catheter thereby offering structural support to the compressed layer of lipids and cells on the luminal surface of the artery[9]. Although, some success has been observed with this method, many complications remain, most notably, their limited use in extensively diseased blood vessels. Furthermore, in stent restenosis is possible. In fact, in-stent restenosis (stenosis diameter ≥50%) occurs in 20-30 % of implanted stents[9].
Because of the problems associated with native, synthetic and modified grafts and prostheses, much research has been done in order to create tissue-engineered vessels. Tissue Engineering has been classified as an interdisciplinary field that applies the principles of engineering and life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function[10]. The present chapter consists of a review of the techniques available to generate an arterial replacement. Various approaches are mentioned but one in particular, using reconstituted collagen as a scaffold, is discussed in detail.
Before going any further, it is important here to take a step back and get familiar with this organ that needs replacement. The blood vessels are components of the cardiovascular system. They are much more than an arrangement of pipes and tubes with blood flowing through. There are many types of blood vessels in the body, each with a distinct link between structure and function. The arteries transport blood from the heart to the other organs and the peripheries of the body. The blood returns to the heart via veins. The arterial and venous systems are linked through a capillary system. In addition to blood transport functionality, blood vessels also possess biochemical functions making them organs in themselves.
In the majority of cardiovascular malfunctions, the problematic section of the cardiovascular system is the artery. The size and function of arteries depends on their location in the body (Figure 1.1)[11]. There are two main types of arteries: elastic and muscular. The aorta and other larger vessels near the heart possess an important elastic component giving them adequate properties to deal with high pressure and flow. Their elastic component (elastin) can be as high as 40%. As the arteries are situated further and further from the heart, this elastic component becomes secondary to the muscular component. In the muscular arteries which range in diameter from 1mm to 1 cm, the elastic component makes up roughly 10%.
All arteries have the same basic structure despite variations in the proportions of their structural components. They are composed of three main layers, or tunicae, which are from the inside out: the intima, the media, and the adventitia[12].
Intima
The intima is composed of a confluent monolayer of endothelial cells (EC) arranged longitudinally, which form the endothelium. The endothelial cells are orientated parallel to the direction of blood flow. This layer of cells assures the hemocompatibility and anti-thrombogenicity of the artery.
Media
The media is the major component of muscular arteries. It is composed of smooth muscle cells (SMC) aligned concentrically along with collagen and elastin fibers and
proteoglycans. The media confers the majority of the mechanical properties to the artery and is responsible for the peristalsis effect which assists with blood transport to the peripheral organs. To this effect, smooth muscle cells contract and relax in response to bio-chemical and bio-mechanical stimuli.
Adventitia
The adventitia is mainly composed of collagen and fibroblast cells (FC) both arranged longitudinally. The function of the adventitia is basically the same in all types of arteries. It provides anchorage with connective tissues and maintains nutrient support and vascularisation to the internal layers of the artery.
These different layers are separated by two elastic laminas which act as barriers. The internal elastic lamina provides also a surface for endothelial cell attachment and supports the endothelium in its role as a barrier to blood contact with the media.
Evidently arteries have a main goal of transporting blood from the heart to the rest of the body in order to assure adequate vascularisation of all organs. However, they do not act as simple inanimate pipes to this effect. Arteries can vary their diameter in response to various external stimuli, such as nerve impulses and hormonal signals, to regulate blood flow in order to control the supply of blood to individual organs. This is especially true for muscular arteries in which the cellular component confers the capacity to multiply by twenty-five times the blood flow to skeletal muscles during physical exertion[13]. For their part, the large arteries near the heart act as a pressure reservoir. During the systolic period, the elastic fibers temporarily store mechanical energy and, in returning to their normal state during the diastolic period, release this energy thereby maintaining continuous blood flow.
The extracellular matrix (ECM) in the vascular wall provides a structural framework for the structural and functional properties of the vessel walls[14]. The ECM therefore affects the elasticity, the resistance and the stretching capabilities of the blood vessel. The two main structural protein constituents of the vascular ECM are collagen and elastin. The collagens provide tensile stiffness and elastin the elastic properties. Another molecule, the proteoglycans, contribute to compressibility; these combined with collagen and elastin, are also responsible for the viscoelastic properties supplying the necessary elasticity to stretch and recoil. These proteins also play essential roles in other functional requirements of vessels such as hydration, ion filtration, growth factor bioavailability and cell-matrix interactions. In fact, the absolute and relative quantities of collagen and elastin have a major impact on the biomechanical properties of vessels. Collagen, with its very high tensile strength, maintains the structural integrity of the vessel. These aforementioned macromolecules are synthesized to some extent by the three vascular cell types (endothelial, smooth muscle and fibroblast).
Collagen is the main protein constituent of muscular arteries where it serves a major structural role. It is in fact, the most abundant protein in the animal kingdom. It is present in almost all tissues ranging from bones, cartilage, tendons, ligaments and all other soft tissues (skin, muscles and all other organs). It is characterized by great tensile strength in molecular and fiber form.
There are at least 20 types of collagen[15]. The human body is mainly composed of collagens type I, II and III, however many other types are present. Collagens type I and III are the major fibrillar collagens in blood vessels where they represent 60% and 30% of vascular collagens respectively[14]. The basic unit of fibrillar collagens is the triple helix formed by three intertwining amino-acid chains (Figure 1.3.3). Each chain is roughly 330 amino acids long and the overall molecule, called tropocollagen, is 300 nm long and has a diameter of roughly 1.5 nm. The most abundant amino-acids are glycine (Gly), proline (Pro) and hydroxyproline. These form a repeating pattern of Gly – X – Y where X is usually proline and Y is usually hydroxyproline. The repetition of glycine in every third position is the most essential factor determining collagens triple helical structure. Intra- and inter-molecular hydrogen bonds are responsible for the stability of the triple helix. Such bonds can be inter-chain hydrogen bonds coupled by NH groups of a glycyl residue with the CO group of a residue in a neighboring chain. Bonds are also formed via water molecules participating in the formation of additional hydrogen bonds with the help of collagen hydroxyl groups[16]. The great strength of collagen fibers, however, originates from the stable intermolecular covalent bonds between adjacent tropocollagen molecules[17]. These fibers can range in diameter from 50-200 nm.
Elastin, an insoluble protein produced specifically by SMC in the media, confers elasticity and supports transportation of metabolic substances in arterial walls[18]. Tropoelastin is a protein of 750 to 800 residues long. This soluble precursor of elastin contains large hydrophobic domains, dominated by aliphatic residues of proline, alanine, valine, leucine and glycine, and smaller alanine-rich domains. Elastogenesis leads to the construction of mature functional elastin within elastic fibers. Mature elastin is an insoluble polymer constituted by several tropoelastin molecules cross-linked together[19]. Elastic fibres are basically composed of amorphous elastin and insoluble microfibrils. During fiber formation, the microfibrillar compound acts as a scaffold onto which elastin is deposited[20]. Despite its hydrophobicity, elastin is highly hydrated in water causing it to swell. Mature elastin is very stable and possesses an extremely low turnover rate.
The arteries are complex organs. In order to replace them, a viable long-term replacement must possess many essential properties. According to Thomas[5], a replacement artery must:
possess a confluent, adherent and quiescent endothelium to resist thrombosis in vivo.
be infection-resistant.
be biocompatible (noninflammatory, non-toxic, noncarcinogenic, nonimmunogenic) and biostable.
have appropriate mechanical properties. They must also possess good suturability during implantation and be kink resistant.
possess appropriate vasoactive physiological properties enabling it to constrict and relax in response to neural or chemical stimuli.
be manufactured relatively cheaply, in a relatively short time period, and in sufficient quantities.
Biocompatibility is a general term used to describe the suitability of a material for exposure to the body or bodily fluids. The specific meaning is dependent upon the particular application or circumstances. In fact, there are no completely biocompatible materials. The success of many medical devices and implants is limited by the interaction of the device materials with the tissues that they contact. This interaction can be improved by various means, but the inflammatory reaction of the body to foreign substances has yet to be eliminated completely. In the case of vascular substitutes, the activation of the immune system and the coagulation cascade can slow the healing process and the host’s ability to integrate the graft into the natural circulatory system. Biocompatibility can be associated to characteristics such as: non-inflammatory, non-toxic, non-carcinogenic and non-immunogenic.
Mechanical attributes relate to burst strength, compliance and viscoelastic properties. That is, they must possess sufficient strength to resist pre-implantation manipulations and resist the stresses acted upon it once in the host. A typical burst strength considered satisfactory in tissue engineering is 2000 mmHg. This burst strength is slightly higher than that of human saphenous vein, a currently used graft, which is 1680 ± 307 mmHg[21]. However, high burst strength is not sufficient. Vessel substitutes must also be compliant. Compliance can be defined as the change in luminal volume in response to a change in pressure inside a tube. Replacement arteries must dilate in a similar manner to natural arteries upon application of luminal pressure. Furthermore, they must be viscoelastic in order to contract to their initial diameter during low pressure periods, without permanent deformation induced during high pressure periods induced by each heart beat.
Furthermore, according to Mitchell, a blood vessel replacement must have a highly organized collagen matrix to impart tissue strength and must contain an elastin network to provide compliance and recoil[22]. In order to meet these demanding characteristics, tissue engineering has been deemed an interesting approach. This approach may prove to provide the solution regarding the shortage of suitable small diameter arterial replacements.
The following is a list of the main vascular tissue regeneration approaches. One approach involves removing cellular components from native tissue and subsequently seeding these acellular tissues with vascular cells. A second approach involves using the body’s natural wound healing response to form a tissue-like structure around a synthetic mandrel inserted in the body. Thirdly, cells can be used to produce ECM in vitro. These ECM/cell sheets can form a tube which mimic the natural physiology of the artery. Lastly, various biological or synthetic materials can be used as scaffolds to guide cell growth into the desired structure.
Acellular tubes can be processed from allogenic and xenogenic arteries and veins[23,24], human umbilical cord[25], or heparinized porcine intestinal submucosal layers[26]. Cellular components can be removed from these tissues without extensive degradation of the extra-cellular matrix resulting in a naturally derived scaffold consisting of extra-cellular matrix molecules. These tubes can become colonised by host cells post-implantation or seeded and cultured prior to implantation. A scaffold processed in this manner from intestinal submucosa consists mainly of type I collagen without other proteins, lipids or nucleic acids[26]. This scaffold can be cross-linked with bovine collagen or synthetic elastomer tubes[27] and has shown SMC and endothelial cell infiltration and remained patent after 13 weeks of implantation in rabbits. A burst strength of 240 mmHg was obtained which is considerably less than the human saphenous vein standard (~1600 mmHg). Another study demonstrated that partially devitalised collagen/elastin matrices may be prepared, without denaturation of the protein, from blood vessels and that the success of this method depends on the segment and type of blood vessel used[23].
Another approach involves stimulating cells in the peritoneal or pleural cavities to coat inert tubing with bone-marrow-derived cells and mesothelium[28]. This approach, that uses the body as a natural bioreactor, attempts to adapt the body’s natural wound healing response to produce a hollow tube. The tubes inserted in the peritoneal cavity become encapsulated with layers of myofibroblasts, derived from bone-marrow-derived cells, and mesothelial cells, which possess antithrombotic properties similar to endothelial cells, to resemble an inverted blood vessel wall[5]. Granulation tissue formation can take up to 2-3 weeks followed by its removal from the tubing mould. The myofibroblasts, which are smooth muscle-like cells, have as much alpha-smooth muscle actin and desmin as a native artery. Autogenic implantation in abdominal rat aortas demonstrated patency of at least 4 months accompanied with a development of contractile responsiveness.
L’Heureux et al. used a completely different approach to constructing a tissue engineered blood vessel (TEBV) by taking advantage of the natural ability of cells to produce their own ECM[21]. Sheets of human SMC and sheets of FC were grown to post-confluence and subsequently wrapped around a perforated synthetic tubular mandrel. The construct was placed in a bioreactor providing luminal flow and cultured for extended time periods. EC seeding was performed one week prior to implantation or in vitro testing. The overall culture period of the constructs was 3 months. These constructs demonstrated adequate mechanical strength (2000 mmHg), blood compatibility and suturability. Short term implantation in a canine model demonstrated a 50% patency rate after 1 week implantation. However, this approach lacks the appropriate medial structure of circumferentially aligned collagen molecules and SMC with functional contractile abilities.
Although the disadvantages of using synthetic polymers for vascular tissue engineering scaffolds were mentioned, they provide adequate mechanical results and therefore may be a short-term option or at the very least another tool to study cell behaviour and in vitro tissue formation. It is to them that we must compare any other regenerated tissues.
Niklason et al.[29,30], used a scaffold of degradable polyglycolic-acid (PGA) seeded with SMC to construct a TEBV. PGA scaffolds were seeded with SMC and matured in bioreactor at 165 pulses per min for 6-10 weeks. Endothelial coating was performed after 8 weeks. Implantation in pigs showed patency up to 4 weeks. It was demonstrated that pulsations increased the production of collagen and improved patency as compared to static culture. These pulsed constructs also demonstrated contraction in the presence of appropriate chemical signals. Appropriate mechanical strength (2000 mmHg) as well as suturability was achieved with this method. After 8 weeks of culture, PGA fragments remained which resulted in a dedifferentiated SMC phenotype in the vicinity of these fragments.
Hoerstrup et al.[31] used a similar approach by seeding fibroblasts and endothelial cells on synthetic support of PGA scaffolds coated with poly-4-hydroxybutyrate (PHA). These constructs were matured for up to 1 month in a bioreactor producing pulsed flow directly through the lumen, thereby generating direct shear stress to the luminal surface as well as periodical radial distension of the vessel wall. Maximal burst strength of 300 mmHg was obtained. More recently, the same group[32] used human umbilical cord cells seeded on PGA/PHA scaffolds to generate a pulmonary artery conduit. They demonstrated the feasibility of using cells which exhibited myofibroblast-like characteristics.
With a similar approach, Jeong et al.[33] seeded vascular smooth muscle cells on tubular rubber-like elastic degradable polymer poly-(lactide-co-caprolactone)(PCLC). These constructs were matured in a pulsatile perfusion bioreactor incorporating shear stresses and radial distension for up to 8 weeks. It is believed that the more rubber-like scaffold was more beneficial than previous attempts in delivering mechanical signals to cells.
Biological support
The above in vitro seeded-scaffold attempts have based their tissue generation on synthetic scaffolds which pose certain threats and disadvantages in vascular tissue engineering. The following is a presentation of various attempts to replace synthetic scaffolds by biological ones such as collagen and fibrin.
Weinberg et al.[34] were the first to attempt a completely biological TEBV using animal collagen gels and cultured bovine endothelial cells and smooth muscle cells, and fibroblasts. A three layered structure with endothelial cells seeded on inside of SMC seeded collagen media with outer layer of adventitial fibroblasts was grown for 4 weeks. A Dacron mesh sleeve was used for support between media and adventitia and the construct displayed a burst strength of 300 mmHg after 3 weeks. Despite relatively weak mechanical results, this study demonstrated the feasibility of this approach.
Hirai et al.[35,36] also used a collagen gel-based approach for autologous canine venous substitutes. Smooth muscle cells were suspended in collagen and statically matured for 7 days. Following EC seeding, the construct was implanted into the posterior vena cava for 24 days after which 9 of 14 grafts remained patent. To prevent tearing, the grafts were supported by a Dacron support. No burst strength measurements were performed; however the effects of initial cell seeding density and initial collagen concentration on collagen contraction were studied.
Ye et al.[37] demonstrated that similar two-dimensional cell seeded gels can be made with from fibrin seeded with human myofibroblasts. Fibrin can be produced from blood plasma. This same group then demonstrated its use for heart valves[38]. Although these attempts were not aimed at blood vessel generation, they provided a basis for the following studies. Grassl et al.[39] examined the feasibility of replacing collagen by fibrin for vascular tissue replacements. They seeded fibrin gels with aortic rat SMC. SMC in these constructs displayed a higher synthesis of collagen. However, the synthesized collagen is not largely incorporated into the developing ECM. The same group pursued this further with SMC seeded tubular gels of fibrin[40]. They found that fibrin media-equivalents (ME) compacted more and were stronger and stiffer than collagen MEs.
Berglund et al.[41] demonstrated the feasibility of supporting collagen gels seeded with neonatal fibroblasts with cross-linked collagen sheets. These constructs demonstrated burst strengths of 650 mmHg due to the cross-linked supports but this strength deteriorated during culture due to excessive degradation. The same group also used tubular collagen gels seeded with rat aortic SMC and matured in a bioreactor with various biochemical factors[42]. This group also studied the effect of a specific growth factor and mechanical strain on the phenotype of SMC. They found that TGF-β strongly inhibited cell proliferation and increased smooth muscle actin (SMA) expression, especially in the presence of mechanical strain.
Collagen can be employed to fabricate arterial replacements in many different ways. Cellular components of natural tissues such as blood vessels and intestinal tissue can be removed to obtain a collagenous constructs. These constructs can be formed into tubular structures and implanted in order to obtain in vivo cell in-growth. These constructs can also be cell-seeded prior to implantation or implanted without cells with the goal of achieving post-implantation cell migration. Another approach involves encapsulating tubes in vivo with granulation tissue by using the body’s natural healing response. Thirdly, in vitro cultured vascular cell sheets can be wrapped around a mandrel to form a tubular cell/ECM structure. Lastly, a more traditional tissue engineering approach utilizes a scaffold structure to provide a substrate onto which cells may adhere and proliferate into a tissue in vitro. These last two tissue engineering approaches are presented below and are summarized in Table 1-1.
As mentioned previously, one of main approaches to tissue engineering involves the use of a scaffold to provide a matrix onto which the cells can organize and develop in the proper environment prior to implantation (Figure 1.4). The scaffold provides an initial biochemical substrate for the novel tissue until the cells can produce their own extra-cellular matrix[43]. This scaffold not only defines the three-dimensional space for the formation of new tissues with appropriate structure, but serves also to provide tissues with appropriate functions[44]. Most cells are anchorage dependant and their growth is influenced by the substrate onto which they are adhered. The scaffold surface properties such as chemistry and wettability affect cell spreading and proliferation while scaffold structure affects cell spatial arrangement and the transmission of biochemical and mechanical signals[45]. Also, various signals provided by the scaffold may affect cellular gene expression. These signals may include cell adhesion molecules, growth factors and mechanical signals.
Tissue engineering scaffolds must meet the following criteria[45]:
be biocompatible and meet the various nutritional and biological needs for the specific cell populations.
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reproducible in three dimensional complex shapes.
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highly porous and structured to permit an adequate cell distribution for cell seeding and permit diffusion of nutritional elements during cultivation.
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potentially tuneable with respect to their chemical, physical and mechanical properties.
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controlled biodegradability.
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In order to meet these requirements, the most common materials range from biological scaffolds such as decellularized xenogenic matrices, small intestinal submucosa, fibrin and collagen to synthetic materials such as Polyglycolic-acid (PGA), Polylactid-acid (PLA), and the biologically derived polymer poly-4-hydroxybutyrate (P-4-HB).[44,43]
Different demands involved in various tissue engineering applications call for an educated choice of scaffold material. Synthetic materials are attractive due to their controlled manufacturability, mechanical properties and degradation times. For many tissues, synthetic scaffolds appear to function quite well.
The cost involved in the conception and production of new materials precludes their conception for very specific applications such as biomaterials. Often, research is conducted with materials conceived for completely different applications. Inevitably, these materials are not tailored to the specific requirements of biological applications. In vascular tissue engineering, the use of a non-biodegradable synthetic material implies a complete loss of vaso-activity and would also hinder the normal remodelling response of the vascular system, thus becoming a physical barrier to long-term implantation[8]. Therefore, the scaffold should be degraded or metabolized during the formation and organization of the newly generated matrix at a rate in accordance with the rate at which newly synthesized molecules are produced and incorporated in the regenerated tissue. This degradation should not produce large debris which would cause occlusion problems down-stream nor degradation products which would be toxic to the body. Although biodegradable synthetic materials are available, biological materials are often favoured. One reason for this is that vascular replacements that are composed in whole or in part of synthetic polymeric materials remain at risk for bacterial colonization and subsequent infection, and are capable of promoting a low-level, chronic inflammatory response that may lead to neointimal hyperplasia[46]. These materials may fracture, induce immunological responses and are difficult to anchor[47]. Also, the chemistry of the scaffold profoundly affects cellularity, cellular gene expression and the overall tissue composition[48]. Natural ECM molecules possess ligands to which cells adhere and are therefore more appropriate than synthetic surfaces which do not possess natural binding sites. A major mechanism by which cells bind to ECM is through integrins which exist as heterodimeric transmembrane glycoproteins, consisting of an α and a β subunit. The binding and activation of integrins promotes a signalling cascade within the cell that affects differentiation, activation, gene expression, and proliferation[44]. The response of cells to external stimuli such as mechanical stimulation has also been shown to be influenced by cell chemistry[49]. More specifically for vascular tissue, compliance mismatch has been associated with graft failure[50]. Hence, the compliance difference between synthetic materials and native arteries may contribute to the high failure rate of synthetic prostheses.
Collagen is a prevalent protein in the vessel wall and, in-situ, plays a structural role as the main load carrying element. It therefore constitutes a valid choice of scaffold material for vascular tissue engineering. Furthermore, it is versatile and can be processed in a variety of forms such as sheets, tubes, sponges, powder and fleece[51]. It can be solubilized by acidic aqueous solution and can be engineered to exhibit customized properties. Collagen in the form of thin sheets or gels has been shown to provide a suitable substrate for many different cell types such as renal[52], hepatocytes[53], epithelial[54], smooth muscle[34,55,56], endothelial[34] and fibroblast[57,58,59]. The following lists other advantages and inconveniences of using collagen as a biomaterial[60].
Advantages:
available in large quantities and relatively easily purified
non-antigenic
adjustable biodegradability and bioresorbability by cross-linking
non-toxic and biocompatible
good resistance and tensile properties
compatible with synthetic polymers
Disadvantages:
high cost for collagen type I
high extracted collagen variability (cross-linking density, fiber size, impurities, etc)
complex handling properties
variable enzymatic degradation
possible side effects and mineralization
As a scaffold for vascular tissue engineering, collagen is commonly used in the form of gel. Collagen in acidic solution can be neutralized and mixed with culture medium. When placed in an incubator at 37ºC for less than one hour, the solution forms a weak gel[57].
Various cell types can be suspended in this gel mould. Unfortunately, collagen in gel form does not possess high mechanical strength, resulting in constructs that are too weak for implantation. Although many research teams have supported collagen matrices with synthetic materials, it has been mentioned previously why this is not a desirable approach. Fortunately, this approach to vascular tissue generation is in its infancy and much can still be improved without resorting to synthetic supports. Some of the main areas in which improvements can be made with this approach include: cell phenotype and functionality, collagen concentration and cell seeding density, collagen fiber orientation, cross-linking, incorporating of other biological molecules, incorporating other biological support structures, mechanical conditioning and vascularisation.
It is well known that cell growth and functionality is guided by the substrate. In mature arteries, the collagen network plays a structural and signalling role in the physiological response of SMC[56]. Scaffolds derived from natural biomaterials such as collagen have intrinsic cell adhesion properties[48]. Arterial SMC have very different characteristics when grown in culture than when found in normal vessels. The normal contractile phenotype changes to a more proliferating, protein secreting mode when grown in culture[61]. The morphology becomes less elongated, proliferation increases with each passage in culture and the expression of smooth muscle actin (SMA) expression decreases with each passage[62]. The factors that play a role in this phenotype change include the form of the substrate, seeding density, the presence of ECM proteins and growth factors, mechanical conditioning and the presence of endothelial cells.
The physical form of ECM molecules strongly influences gene expression of adherent cells[48]. In collagen gels, SMC appear much more elongated, appear to be more in a contractile state, and therefore grow more slowly than those grown on plastic or on two-dimensional collagen sheets. SMA expression of SMC is also down regulated in collagen gels. Cells seeded below confluence maintain a synthetic phenotype until confluence is reached[63].
The ECM is known to affect cell function through both biochemical and mechanical signalling pathways. Culture of SMC on different ECM molecules affects cell phenotype by modulating morphology, proliferation and protein expression[62]. Collagen type IV, elastin and laminin maintain cells in the contractile phenotype while fibronectin promotes a change to a more synthetic phenotype[64,63].
Phenotype control can also be modulated by exogenous biochemical stimulation such as growth factors. Growth factors are proteins which promote proliferation and migration of cells via interactions with specific cell-membrane receptors. To this effect, ascorbic acid, platelet-derived growth factor (PDGF), transforming growth factor β (TGF-β) and heparin have been studied[65,62]. Table 1-2 shows an overall picture of the various growth factors and their effects. SMC and FC reach confluence faster in the presence of ascorbic acid[65]. In gels, PDGF caused an increase in cell number and a decrease in SMA. These effects were the opposite with cultures in two-dimensions. Although TGF-β causes an increase in SMA and a decrease in proliferation on flat surfaces, this factor was found to have little effect in three dimensions. Heparin slightly increases SMA in flat culture conditions but has no effect in gels. It does, however, decrease the growth rate of cells in gels and on flat surfaces. Therefore, the presence of a three-dimensional collagen matrix has an effect on cell morphology, proliferation and SMA expression. This matrix also affects the cells reactions to various biochemical signals. There is also a marked difference in SMA between cells grown in disc-shaped gels as compared to tubular gels compacted around a mandrel[42]. It is possible that the higher mechanical stresses in the latter case are sensed by the cells which react accordingly. This may also explain the varied morphological behaviour of SMC depending on their position in the disc-shaped gels and tubes[66]. In discs, cells on the surface of gels are more flattened. Cells on the bottom of gels are more organized into a cellular network than those in the middle which are more elongated. This also applies to tubular gels in which the position of cells relative to the lumen may affect phenotype.
Table 1-2 : The effects of various chemical signals on smooth muscle cell seeded collagen gels or films. (a blank box indicates no relevant influence was found in the literature.)
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Cyclic distension induces SMC to a more pronounced contractile phenotype[67]. It has been shown that SMC in the synthetic phenotype can revert to the contractile phenotype after implantation[68]. Shear stress also affects EC functionality. These mechanical environmental conditioning effects are discussed in more detail further on.
The engineering of a TEBV in this manner infers the co-culture of EC and SMC. This co-culture adds increased complexity than simply referring to two simple cultures of each cell type. EC and SMC can interact by two mechanisms: by body fluids and by direct contact[61]. EC secrete both inhibitors and stimulants of SMC growth and SMA expression such as TGF-β and PDGF. However, in co-culture, EC release more SMC growth inhibitory factors such as a heparin-related glycosaminoglycan and transforming growth factor β (TGF-β). Co-culture of these two cell types also results in completely different endothelial properties. EC grown on SMC seeded collagen gels are more elongated than those grown in culture dishes.
It is also possible with the collagen gel method to culture all three vascular cell types together. It has even been demonstrated that it is unnecessary to separate SMC and FC prior to gel formation. Vessel replacement constructs with a homogeneous mixture of SMC and FC prior to implantation demonstrated a segregation of these two cell types after implantation with SMC accumulating on the subendothelial layer and FC accumulating on the outer layer[69].
Control over the phenotype of SMC is essential. Intimal hyperplasia caused by excessive growth of SMC at the TEBV-artery interface is a major impediment to long-term implantibility of TEBVs[18].
A major problem associated with TEBV without synthetic support is a lack of adequate mechanical integrity, more specifically stiffness, strength and elasticity. Collagen tubular constructs naturally stiffen during culture although this takes an extended period of time and may not be sufficient. Therefore, many approaches exist to attempt to improve mechanical properties of collagen vessel replacements. The main approaches to achieve this are discussed below.
When dealing with a collagenous tissue derived from a contracted gel, contraction primarily determines the initial mechanical strength[35]. The two most basic factors that affect collagen gel compaction are the collagen concentration in the initial solution and the SMC seeding density. A higher SMC seeding density accompanied with relatively low initial collagen concentration induced more rapid and prominent shrinkage[36]. Gel compaction can also be affected by various biochemical and mechanical factors. Gel compaction is increased with PDGF and TGF-β and decreased with heparin[62]. Mechanical stimulation in the form of cyclic distension caused an increase in compaction[42]. Gel compaction is also affected by collagen type with collagen type I promoting higher gel compaction than collagen type III[70].
Despite these findings, the overall effect of collagen concentration in initial solution is somewhat limited. Greater strength can be achieved to some extent by increasing post-compaction collagen concentration. A higher cell-produced collagen concentration would lead to increased mechanical strength. SMC are known to synthesize their own matrix when grown on or in an adequate substrate and when in the appropriate phenotype[63]. In this sense, an increase of endogenous collagen production by cells accompanied with an appropriate rate of collagenase activity and enzymatic degradation of reconstituted collagen would lead to improved mechanical properties. In vitro collagen synthesis can be influenced by various biochemical factors such as: cyclic strains, growth factors, ascorbic acid and amino acid supplementation[22]. Collagen synthesis is also increased by mechanical stimulation[71,60].
Ascorbic acid is an essential cofactor for hydroxylation of proline and lysine residues in collagens synthesized by human FC. It is also known to modulate the growth properties of cells[37]. Ascorbic acid in low concentrations is essential for the production of collagen and has been shown to increase type I collagen production by SMC and FC[72,48]. However, it also produces negative effects on elastin accumulation and production. These effects are dependant on both the dose and the time of exposure. Combining ascorbate with TGF-β and insulin increases collagen incorporation. TGF-β increases overall protein synthesis while ascorbate increases the collagen fraction of protein synthesis[73]. Also, TGF-β, plasmin and insulin were shown to increase collagen production by human dermal fibroblasts[74].
The degradation of the matrix, involves enzymes known as matrix metalloproteinases (MMPs) which include collagenases, gelatinases, metalloelastases, etc. Each degrades a specific ECM component. These enzymes are produced by vascular cells and can be stored in latent form until required[12]. Their activities are regulated by growth factors and cytokines. These MMPs play a crucial role in vascular matrix remodeling and are essential for cell migration, synthesis of new ECM components and regulation of growth factors. MMP activity must be controlled and regulated to match that of collagen production in order to maintain the structural integrity of TEBVs. MMP production, like all aspects related cell function, may be modified by various biochemical and mechanical factors[42].
Fibers of 5-50 µm can orientate and guide cells. This phenomenon was discovered by Paul Weiss in 1945 and is called ‘contact guidance’. Collagen fibers have this ability to guide and orientate cells. Collagen fibers also contribute to cell remodelling and are absorbable over time as cells produce their own extracellular matrix[41]. The natural media is composed of circumferentially aligned collagen fibers and SMC and this alignment is important for both vasoactivity and structural integrity. Since cells can be orientated by collagen fibers, by first of all guiding fibers to a circumferential orientation, the overall media can aligned in this manner.
By using a cylindrical collagen gel seeded with smooth muscle cells, gel compaction can occur around an inner mandrel and results in a natural circumferential orientation of both collagen fibers and smooth muscle cells. This orientation can only occur if the gel is allowed to contract longitudinally as well as circumferentially. This natural phenomenon can be assisted by various means. Alignment can also be achieved using the effect of a strong magnetic field during collagen fibrillogenesis[75]. Magnetic field fiber orientation worked well with constructs without a central mandrel. However, mandrel compaction proved to be more beneficial than magnetic pre-alignment on construct integrity. This approach, however, may decrease the time to total contraction. More importantly, cyclic distension via forced peristaltic flow through the lumen induces circumferential collagen fibril/SMC orientation[76].
Collagen can be cross-linked in order to increase its mechanical properties and molecular stability. The most basic strategies introduce stable and covalent intermolecular cross-links between collagen fibrils[77]. Cross-linking influences the strength, resorption rate, and biocompatibility of biomaterials. Collagen molecules are endogenously cross-linked by cells and can be artificially cross-linked in many different ways, although not all are appropriate for vascular tissue engineering. There are two main conditions which must be met to provide suitable cross-linking[77].
Collagen fibrils must contain an amino-acid reactive group to be targeted by the chemistry of the cross-linking agent and coupling mechanisms.
The conditions such as temperature, pH and nature of solvent must not be detrimental to the collagen molecule, fibrils or properties of the bioprosthesis.
There are basically three main ways to induce collagen cross-linking. These include artificial methods, using biological compounds or inducing endogenous collagen cross-linking by cell.
Some chemical cross-linking agents include: formaldehyde, glutaraldehyde[78,79], polyepoxy compounds[78,80], isocyantes[77], chromium and carbodiimide[81]. Although the exact mechanisms may vary, these methods achieve cross-linking by forming covalent bonds between adjacent collagen fibrils. Although effective at this, they are themselves cytotoxic. Glutaraldehyde can autopolymerize and subsequently hydrolyze, releasing free glutaraldehyde which is cytotoxic. Also, glutaraldehyde causes calcification and an exaggerated increase of stiffness which is not beneficial for vascular tissue. Polyepoxy compounds have been suggested as a replacement since they do not cause calcification nor an exaggerated increase in stiffness[80]. In order to decrease cytoxicity, reagents such carbodiimide (EDC) and N-hydroxysuccinimide (NHS) can be used. These reagents act solely as catalysts for specific cross-linking reactions and are removed from the tissue following the treatment. These above methods have been used with collagen materials with which the goal is to perform subsequent in vitro or in vivo cell ingrowth. However, these chemical methods, which are cytotoxic, may not be used on collagen gels suspended with cells since the cells are present at the time of cross-linking.
Artificial cross-linking can also be achieved without the use of potentially harmful chemical reagents.
Ultraviolet irradiation is efficient for the introduction of cross-links. Telopeptides on each end of the collagen molecule are responsible for the photochemical reaction to ultraviolet light. Collagen can be cross-linked in this manner in the form of films, fibers or while in solution. In the latter, the viscosity of the solution increases with irradiation time and a gel may be formed. Collagen is degraded by prolonged irradiation and the fibril formation ability of collagen is easily deteriorated even after short periods of ultraviolet irradiation. This irradiation also modifies the properties of collagen in collagen-cell interactions by increasing cell growth. This effect can be deleterious for vascular tissue engineering. Therefore, these approaches are mainly appropriate for collagen films and fibers. Recently, a method involving visible-light photomediated cross-linking of collagen gels in the presence of SMC was investigated[46]. Collagen was derivatized, through lysine and hydroxylisine residues, with methacrylamide moieties. These moieties, in the presence of a photo-initiator, were photochemically cross-linkable. This method showed some efficiency in cross-linking while retaining collagen triple helical structure and high cell viability.
Dehydrothermal treatment (DHT) is a physical method of cross-linking collagen fibers that avoids potentially cytotoxic reaction products and provides moderate strength and resorption rate[82]. Cross-linking is dependant on exhaustive removal of bound water from collagen molecules which results in condensation reactions between the carboxyl and amino groups on adjacent amino acid chains. This method can take at least three to five days to complete and causes significant degradation to the collagen molecules. Evidently, this method is not appropriate for collagen gels.
Glycation is the nonenzymatic crosslinking of amine groups of collagen and other ECM proteins brought about by reducing sugars, such as glucose and ribose. It results in increased tissue stiffness and increased resistance to enzymatic degradation and is cell tolerated. In order to increase glycation, it is simply necessary to increase the sugar concentration in the medium, particularly that of ribose which is more effective than glucose because it is about 17 times more available in open form[83].
Although not in current use in works related to vascular tissue engineering, there exist various biological compounds capable of cross-linking collagen. Genipin, which is a compound isolated from fruits of the gardenia plant (Gardenia jasminoides), cross-links collagen in manner similar to gluraldehyde. The dialdehydes from Genipin react with ε-amino groups on lysine side chains of neighboring collagen molecules[77]. Another compound, Nordihydroguaiaretic acid (NDGA), isolated from the creosote bush, produces a different fixation mechanism. Cross-links are not formed between side chains of collagen molecules but rather, NDGA polymerizes and is interpolated in the collagen network. Fibers produced in this manner exhibit mechanical characteristics superior to those obtained with all of the chemical reagents mentioned above. Furthermore, this method and that using Genipin are less cytotoxic than the chemical reagents and create little foreign body response or inflammatory reaction[77]. Enzymatic cross-linking methods, which can be used for binding peptides and proteins, may also constitute a valid means for cross-linking collagen.[84]
Cross-linking of collagen can be achieved by cells and their natural biochemical products. In this sense, it is also possible to enhance the cells ability to cross-link. Endogenous cross-linking can be enhanced by ascorbic acid and ribose[35]. Lysly oxidase-mediated cross-linking is a significant contributor to stiffening. Enzymatic cross-linking of free amine groups of lysine and hydroxylisine residues in collagen can be achieved by SMC-produced lysyl oxidase[83]. Mechanical conditioning also influences the cells ability to cross-link and organize collagen.
The overall objective of any cross-linking approach is to obtain adequate biomechanical properties as well as a rate of degradation of collagen in tune with that of the repair process.
Other natural proteins have been assessed in their ability to enhance the mechanical properties of collagen-gel based blood vessels.
As mentioned previously, compliance mismatch has been associated with graft failure. This would also be the case with regenerated vessels. Elastin contributes to compliance by allowing stretch and recoil of the arterial wall with each pulse. Elastin fibers act in the physiological range of deformation and provide the necessary resilience to recover from deformations associated with pulsatile flow[85]. Consequently, an elastic network in a TEBV would prevent vascular dilation in response to the continuous pressures exerted by blood flow in vivo. Also, soluble elastin inhibits platelet aggregation induced by collagen[22]. Therefore, incorporation of some form of elastin or stimulation of insoluble elastin production in vitro would be greatly beneficial for a successful TEBV. Attempts to do so are being pursued in one form or another.
Elastin is very difficult to process, due to purity and chemical contamination issues and a strong tendency to calcify[20]. It is also very hydrophobic making it difficult to use in matrix fabrication techniques such as collagen gels. Other forms of elastin may be used. It was shown that incorporation of soluble α-elastin into collagen gels inhibits SMC proliferation and migration limiting SMC hyperplasia without significant effects on endothelial cell formation[18]. This soluble elastin protein also inhibits platelet aggregation induced by collagen. Recombinant human elastin polypeptides may be expressed, produced and purified. These polypeptides may prove to be a valuable component of a tissue engineered vascular conduit[86]. A different approach combines the traditional cell-seeded collagen gel approach with an auxiliary elastin scaffold isolated from porcine carotid arteries[85]. This structure increased mechanical properties, especially creep resistance, as compared to control gels. Collagen was also combined with elastin by lyophilizing various mixtures of both molecules followed by cross-linking of the dried sponge[87]. An increase of collagen ratio improved stiffness while increasing the elastin ratio improved elasticity. Collagen appeared to act as glue, incorporating elastin into the scaffold and holding the elastin fibers together. Pore sizes could be varied by changing the freezing rate prior to lyophilization.
Due to the difficulties in processing elastin prior to incorporation in a TEBV, approaches have been directed at enticing cells to increase their natural elastin production. SMC and FC are known to produce elastin in vitro. ECM molecules likely influence elastogenesis. Fibrillar and non-fibrillar glycoproteins and proteoglycans have been shown to increase elastin synthesis[73]. Various biochemical factors also influence these cells to produce more elastin. It was shown that TGF-β and insulin-like growth factor-1 increase tropoelastin mRNA and protein synthesis by cultured cells. TGF-β also increases lysyl oxidase activity which acts to crosslink tropoelastin into its mature fully functional form[73]. However, ascorbic acid and ascorbate, which was shown to increase collagen synthesis, inhibits elastin synthesis. It is important to note here that, elastin production by SMC is maximal for neonatal cells and almost inexistent for adult cells[88]. This effect is less severe with fibroblasts which retain their ability to produce elastin mRNA for longer time periods. However, fibroblasts deposit little insoluble elastin.
Most glycosaminoglycans (GAGs) are found in the form of proteoglycans on cell surfaces and in the extracellular matrix[89]. The interactions between GAGs are essential for the adhesion, migration, proliferation and differentiation of cells[90]. Incorporation of GAGs into TEBV may therefore allow exploitation of their biocharacteristics and valorize biomaterials like collagen. It was shown that various GAGs such as chondroitin sulfate (CS), dermatan sulfate, heparan sulphate (HS) and heparin can be covalently cross-linked to collagen using EDC-NHS[90]. This cross-linking method was mentioned above as suitable for collagen cross-linking resulting in low cytotoxicity and adequate biocompatibility. This study showed that CS in particular decreased the tensile strength of collagenous matrices[91] forgoing it’s utilization to increase mechanical properties of gels. However, other GAGs may prove to be more beneficial to this effect. The interest in incorporating GAGs is not solely for mechanical integrity. For instance, heparin sulfate proteoglycans, which are abundantly present in basement membranes, regulate cellular activities and therefore may be used to achieve appropriate cellular adhesion and differentiation. Furthermore, heparin, heparan sulphate and dermatan sulphate have been shown to exhibit anticoagulant activity[92] and may reduce the need of systemic heparinization. Heparin also improves endothelial cell proliferation and reduces proliferation of SMCs, thereby helping to prevent the formation of intimal hyperplasia[93]. Various GAGs can also be used as slow-release vehicles for growth factors and other biochemical molecules[94].
The influences of mechanical forces on cells in the vascular system have long been recognized. Cyclic mechanical strain on two-dimensional substrates has been shown to increase growth factor and matrix production and increase the expression of contractile proteins[42]. These effects are also present in three-dimensions. It was demonstrated that flow and cyclic stretch both have an effect on vascular biology[67]. Tension as a specific form of mechanical stress also affects cell function and behaviour[76]. Also, circumferential alignment of cells is extended throughout the media equivalent in conditioned constructs. Unconditioned constructs demonstrated cell alignment mainly in the immediate vicinity of the lumen.[95]. Dynamically stressed constructs have a higher potential for reorganization and a higher orientation of SMC and collagen. Cell-mediated remodelling of the tissue is thereby accelerated. The amplitude of strain also affects the remodelling response with 15% distension having a larger effect than 5%[95]. Mechanical stress enhances synthesis and secretion of proteins and mechanical stiffness, and has effects on the level of gene expression. Tension as a specific form of mechanical stress affects the cell function and behaviour[76]. Dynamic cyclic stretch improved mechanical integrity of TEBV as compared to static flow and completely static culture[96]. Pulsed flow induces increased production of collagen. The rate of pulsing also has an effect with 90 beats per minute (bpm) inducing higher structural integrity than 165 bpm[97]. The presence of MMP-1 was found to be greater in pulsed vessels and higher at 90 bpm than 165 bpm. It is possible that pulsatile culture conditions induce a higher rate of collagenolysis.
Mechanical strain has been reported to stimulate extracellular matrix synthesis as well as secretion of enzymes responsible for matrix degradation and tissue remodelling[71,95,60,97]. Mechanical stimulation also has an effect on the SMC reaction to various biochemical factors and vice-versa[42]. Mechanical strain diminished the effect of PDGF on both gel compaction and cell proliferation. The increase in cell proliferation due to mechanical stress was suppressed by adding TGF-β.
Many teams have studied the effects of cyclic strain on a tubular tissue using an inner plastic compliant tube as a mediator between the flow and the collagenous tissue. This approach however, does not allow for shear stress which is an additional component of mechanical conditioning. Shear stress, the tangential force acting in the direction of blood flow on the surface of endothelial cells, also promotes the orientation and elongation of these cells. It also has an effect on biochemical factors. In the regulation of vascular tone, relaxation of SMCs is dependant on the integrity of the endothelium[98]. This is regulated by the endothelium-derived relaxing factor (EDRF), whose release is regulated by shear stress.
In order to simulate these mechanical factors that condition TEBV and improve their properties, adequate bioreactors must be used. These bioreactors must provide complete control over many different environmental factors such as temperature, culture medium, chemical factors, and mechanical environmental forces such as shear stress due to flow, cyclic radial distensions and perhaps even longitudinal tensile stress[99].
A major impediment to vascularisation and nutrient supply and waste removal to and from engineered tissues is the foreign body response which induces the formation of a fibrotic capsule surrounding the implant. Molecules must diffuse through this barrier. By using collagen as opposed to synthetic materials, this response is minimal, thereby favoring blood supply to the tissue.
Furthermore, although tubular gels of collagen are relatively thin, the incorporation of a microvascular network will insure adequate nutrient supply and excrement removal to and from vascular cells throughout the entire thickness. A thickness of less than 1 mm is required to block adequate diffusion of these elements without a microvascular network[100]. Fortunately, three-dimensional collagen scaffolds can support vascularisation processes[60]. Cultured endothelial cells on or within ECM substrates such as collagen, fibrin, fibronectin or laminin form capillary-like structures[101]. The ability to form these tubes is dependant on collagen concentration and type[102]. Mechanical factors such as cell adhesion and spreading also play a role. There also exist various pro-angiogenesis growth factors, such as acidic fibroblast growth factor (aFGF), basic fibroblast growth factor (bFGF), vascular endothelial cell growth factor (VEGF), angiopoietins and ephrins which may induce more-rapid and prominent capillary formation[101,98]. VEGF has the most critical influence on vascular formation since it is required to initiate the formation of immature vessels by vasculogenesis and angiogenic sprouting[98].
The human body has developed over many centuries of evolution into a highly complex system. We are no where near an understanding of the entire complexity nor, for that matter, able predict the result of various factors on tissue formation and growth. It is therefore all the more difficult to try and imitate evolution or the work of some higher being and generate a tissue from scratch. Although some success has been achieved with various tissues such as skin and cartilage, it is much more complicated when dealing with three-dimensional tissues implicating multiple cell types. Generating a pseudo-physiological artery is not a simple matter. A proper scaffold material must be found which mimics the extra-cellular matrix onto which cells adhere and proliferate in vivo. One of the most adequate materials to this effect is collagen. This collagen can be produced by cells or reconstituted collagen can be extracted from various sources. Reconstituted collagen can be processed into thin films or more appropriately cell-seeded gels which can be contracted by the cells around a mandrel to form a tubular collagenous tissue.
Furthermore, the choice of cells is important. Each cell type affects the others. In order to obtain adequate mechanical and functional tissue properties, all cell types must have their proper functionality in order to influence the other cells in the correct manner. In order to do this, the in vivo environment must be recreated as much as possible. To start with, the scaffold material is essential to provide adequate cell adhesion, proliferation and functionality. This is not sufficient however. The biochemical and mechanical environment must also be recreated. Perhaps most importantly are the mechanical stimulations, such as shear stress and cyclic distensions due to pulsed flow, which have great effects on both EC and SMC. These forces may induce proper phenotype functionality in both cell types. By doing so, EC and SMC can produce and transmit their proper biochemical signals. Similarly, biochemical factors can alter tissue structure or function in such a way as to alter the forces exerted on cells. Addition of other biochemical signals may be necessary in vitro to replace those found in vivo but not produced by cells.
To date, mechanical integrity of collagenous tubes is not sufficient for implantation in the arterial system. Burst strengths rarely exceed 2000 mmHg unless these constructs are enhanced by a synthetic support. Nor do the tissue engineered arterial replacements currently available possess adequate viscoelastic properties. Fatigue is also a factor as implanted TEBV are exposed to conditions involving cyclical stresses for long time periods. However, as discussed above, there are a multitude of factors at play that affect the strength of these tissues such as collagen concentration and fiber alignment, cross-linking, cell source and phenotype and biochemical and mechanical environmental factors. Progress is continually being made in this relatively young field. As we progress, our understanding of the requirements of cells increases which helps us to better improve the next generation of arterial replacements in order to meet the challenges of tissue engineering a vascular graft; that is providing a conduit that will have sufficient strength not to burst with changes in blood pressure, a vessel wall that is elastic and can withstand cyclic loading, matching compliance of the TEBV with the adjacent host vessel, and a lining of the lumen that is antithrombotic[103].
This approach takes a great deal of time and would therefore not be designed for use in emergency surgeries. However, with the increasingly efficient testing methods available today, problems may be more and more predicted in advance leaving sufficient time to culture a suitable TEBV with long-term capabilities.
The above chapter described the problem at hand concerning the lack of suitable blood vessel replacements as well as some of the major approaches to filling this lack. One approach in particular, involving using cell-seeded collagen gels, was discussed in detail with a review of the possible means to improve such tissue engineered blood vessels. The following chapter presents this approach as it was undertaken in our laboratory including the rationale, methodology and various unpublished result in order to work towards the overall project objective which is to design and develop a scaffold structure from collagen for a tissue-regenerated artery.
© Jason Habermehl, 2005
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 | Chapter 2: Rationale, Methodology and unpublished results |
