Blood Vessel Tissue Engineering

Biotextiles for Blood Vessel Tissue Engineering - Overview


Interest in the field of tissue engineering has grown tremendously as it provides a powerful tool for the development of a novel set of tissue replacement parts and technologies. The emergence of tissue engineering opens new possibilities in reconstructive vascular prosthesis. The most important aspects of an arterial graft are porosity, compliance, and biodegradability. The scaffold is expected to act as a substrate to promote cell growth, maintain differentiated cell function, and organize and direct the formation of new functional tissue. A fibrous scaffold has significant advantages over polymer films in the high level of porosity needed to accommodate large number of cells. Biotextiles, defined as structures composed of textile fibers and designed for use in a specific biological environment, can meet all the criteria currently considered necessary for the design of ideal scaffold structures and have an edge over the polymer scaffold. Textile scaffolds may be knitted, woven, non-woven, braided or of composite construction depending on the intended application. A review of the structures used in the vascular grafts will help clarify the underlying structure-property relationships in vascular grafts. The purpose of this review is to take a closer look at textile based structures application as tissue engineering scaffold for blood vessels. Based on the current research and existing products, the current scenario and approaches in this fascinating area are thoroughly discussed.


tissue engineering, scaffolds, blood vessels, biotextiles.

1. Introduction

Cardiovascular diseases are increasingly becoming the main cause of death all over the world. This has led to an increase in the economic and social burden of such diseases. According to the American Heart Association, 16.7 million people around the world die of cardiovascular diseases each year, which represents one third of all deaths around the globe [1]. In 2005, the direct and indirect cost of cardiovascular diseases estimated in the United States was about 393.5 billion dollars [2]. The most common among these is atherosclerosis, caused by a lesion of a raised focal plaque within intima [3]. The malfunctioning blood vessel can be replaced by autologous veins or arteries, but at the cost of other healthy tissues [4]. Arterial autografts are the ideal materials for the reconstruction of malfunctioning small diameter arteries. After implantation, the graft remains viable and possesses features similar to that of a normal artery, such as compliance, flexibility, and antithrombogenicity. Its applicability, however, is limited because the grafts of the required length and diameter are not always sufficiently available. Moreover, harvesting of these grafts may occasionally lead to tissue necrosis at the donor sites [5]. Further, complicating the use of autologous vessels is the fact that they are limited in number and a previous bypass surgery may have already required the vessel's use. One strategy to overcome these limitations applies tissue engineering approach to construct biologic substitutes of diseased native vessels. These artificial blood vessels should ideally be composed of viable tissue, able to contract in response to hemodynamic forces and chemical stimuli, and secrete normal blood vessel products [6].

The emergence of tissue engineering opens new possibilities in reconstructive vascular grafting. Tissue engineering is an interdisciplinary field that applies the principles of engineering and the life sciences toward the development of biological substitutes that restore, maintain, or improve tissue function [7]. The basic principle of tissue engineering involves the use of an appropriate cell source and a biocompatible and biodegradable scaffold to produce a construct that structurally and functionally mimics the target tissue. The requirements of scaffolds for tissue engineering are complex and depend on the structure and function of the tissue of interest [8]. The scaffold should be able to act as a substrate to promote cell growth, maintain differentiated cell function, and organize and direct the formation of new functional tissues. Thus, the successful tissue regeneration relies on the seeding cells, the scaffolds and the construction technologies [9, 10]. Functional TEBVs should be non-thrombogenic, non immunogenic, compatible at high blood flow rates and have similar viscoelasticity to native vessels [11-13]. Moreover, the grafts should be living tissues that could eventually integrate into the body and become indistinguishable from the native vessels. It has been accepted that the functional TEBVs cannot be achieved without ECs, SMCs, biodegradable scaffolds and the unique vessel-engineering techniques ( 1).

2. Blood Vessel Structure and Blood Vessel Tissue Engineering

Before considering design of tissue engineered arterial replacements, it is important to understand aspects of the anatomy and physiology of the normal arterial wall. The arterial wall is composed of three distinct tissue regions:

* the intima, with a non-thrombogenic endothelial cell lining,

* the media, with smooth muscle cells and elastin fibers aligned circumferentially for optimal mechanical properties, and

* the adventitia, composed primarily of fibroblasts and connective tissue. An ideal replacement would mimic the properties of each of the tissue layers [14].

The extracellular matrix surrounding the vascular cells is complex and combines to provide the biomechanical properties of the tissue [15]. The complex mixture of molecules and their organization provide the blood vessels with their properties that allow them to function throughout life. 2 shows a schematic of the structure of a medium sized artery, with the main features, components, and environmental influences listed. The innermost layer called the intima consists of a monolayer of specialized endothelial cells (ECs), which forms a tight nonthrombogenic barrier between the lumen of the vessel and the rest of the vessel wall. This layer is critical not only in preventing unwanted clot formation, but also in preventing infection and inflammation of the underlying tissue, as well as in signaling to the muscular component of the vessel wall. Beneath the intimal layer, is a layer of basement membrane enriched in collagen IV and laminin, followed by the internal elastic lamina which is a fenestrated but acellular layer of elastin. The media is the muscular layer of the artery, and is composed largely of collagens Type I and III, as well as lesser amounts of other proteins and proteoglycans. Smooth muscle cells (SMCs) that have a specialized contractile function are also found in this layer. The collagen matrix and the SMCs are generally aligned circumferentially or in a spiral pattern along the axis of the vessel. In larger vessels, elastic laminae further subdivide the media. Another elastin layer, called the external elastic lamina, which separates the medial layer from the adventitia, the outermost layer of the vessel wall, surrounds the medial layer. The adventitia consists mainly of a loose collagen matrix with embedded fibroblasts. This layer provides a substrate for a vascular supply to the artery wall and serves to anchor the blood vessel to the surrounding tissue, as well as to provide additional structural support [16]. This structure varies along the arterial tree and the layer thicknesses are not constant. Therefore a precise quantitative study of the wall structure is difficult, and the actual wall structure is very complex [17].

Polymers used in tissue engineering serve mainly as a physical support for seeded cells and function as three dimensional (3D) scaffolds to enable cell adhesion, migration, proliferation, differentiation, and eventually tissue regeneration [18]. Although, it depends on specific application, in general, TEVG should [19]:

· be biocompatible (noninflamatory, noncarcinogenic, nonimmunogenic) and biostable.

· be non-toxic, where fiber polymer or fabrication techniques must be non-toxic and fiber should be free of contaminants.

· possess a proper degradation rate. Degradation behavior of scaffolds also plays a major role in engineering new tissues and affects the cell growth and viability.

· have suitable artificial surface for the body cells to easily adhere to and grow on.

· be microporous to provide a stable anchorage for vascular cells and stimulate cell ingrowth.

· be leak-proof and thrombo resistant, but with adequate porosity for healing and angiogenesis.

· possess appropriate vasoactive physiological properties including the ability to constrict or relax in response to neural or chemical stimuli and be able to be manufactured cheaply in a relatively short space of time and in sufficient numbers with differing specifications (diameter, length, etc.) to meet commercial demand and fit the graft [20-22].

· have appropriate mechanical properties. It should provide sufficient kink and compression resistance, a high degree of integrity that ensures an efficient transfer of force between the soft vascular tube and the reinforcing component, as well as ready-to-suture properties [23].

· maintain compliance after tissue ingrowth into the graft. Matching the compliance of implanted vascular segments to native vessel is related to the material elasticity, pore size, type and degree of tissue growth into the porous structure and anastomotic suture and technique.

By replacing small arteries with rigid grafts, a mismatch in compliance is introduced to the system [24]. This mismatch not only hinders patency initially, but also the discrepancy actually becomes more pronounced as the implant remains in vivo, as shown in Table I [25]. In addition, low compliance is directly linked to decreased patency [26]. Today's commercially available vascular grafts are classified according to their fabric construction and surface characteristics. Table II shows the classification of commercially available vascular grafts. Textile scaffolds may be knitted, woven, nonwoven, braided or of composite construction depending on the intended application. Most strikingly, these are some of the simplest structures found in textile products. Regardless of the simplicity of the structures, it is important to realize that these prostheses possess significantly different properties in terms of porosity, bursting strength and thickness [27]. A review of the structures used in the vascular grafts will help clarify the underlying structure-property relationships in vascular grafts.

3.1 Woven Scaffolds

Early experimental and commercial vascular grafts belonged to the woven category. Today, woven vascular grafts account for about 45% of the grafts being implanted each year [28]. A woven fabric is a textile structure formed when two sets of yarns are interlaced at right angles to each other. The longitudinal yarns are known as the warp and a single yarn of the warp is called an end. The widthwise yarns are known as the weft, or filling, and a single yarn of the weft is called a pick. The fabric weave is the order in which the warp and filling are interlaced. Plain weave is the simplest of all weaves. In plain woven fabrics, the warp yarns lie alternately over and under the weft threads. Despite their simplicity, plain woven fabrics have the highest number of interlacing and are, therefore, a most dimensionally stable structure. They exhibit high bursting strength, minimum tendency to fatigue, and they can be fabricated tightly enough to lower permeability to water (and blood). However, they are extremely difficult to handle and suture. Moreover, they are non compliant and have a tendency to fray at the edges.

Kasyanov et al. [29] used polyurethane monofilament yams with low modulus of elasticity and bulked polyurethane monofilament yams with high modulus of elasticity, in developing compliant structures. Weaving process was used to make tubular grafts diameter in the range of 4-6 mm. The stretch in the polyurethane thread caused the crimp in the weft threads which made the graft compliant in the transverse direction. The luminal surface was covered with gelatin for minimization of water permeability and raising of blood compatibility. The results of mechanical tests showed that grafts behavior found is similar in character to that noted for the human carotid artery. Three grafts were implanted in mongrel dogs. Morphological examination carried out after three month of implantation showed that the endothelial like cells appeared on the graft flowing surface. The hope for achieving long term patency in grafts lies in matching the compliance of a vascular graft with that of the artery and developing a thrombo resistant surface at the luminal wall. Moreover, by varying the sizes and properties of the individual yarns, the structure and properties of vascular grafts could be effectively engineered to suit the application.

The performances of textile composite vascular grafts, specially designed and constructed by Gupta et al. [30] were investigated by them. In developing compliant structures, textile threads of two widely different deformative characteristics, one matching nearly those of the elastin and other of the collagen fibres, were selected. Two types of woven grafts were developed. Polyester threads were used as warp, and the same polyester and pre-stretched polyurethane were used as the weft in the first type (HVG-1). In the second type (HVG-2), pre-stretched polyurethane thread combined with polyester were used as both the warp and the weft threads. 3 shows the structure of the outer and inner walls of HVG-1 before implantation. The loops noted are those of the bulked polyester yarns formed by the recovery of the polyurethane yarns. The grafts obtained were stretchable and thus compliant in both the transverse and the longitudinal directions. The polyurethane and the polyester threads in the graft seemed to play approximately the same roles as the elastin and the collagen fibres, respectively in the natural artery.

3.2 Knitted Scaffolds

Knitted structures possess special features which are of interest for structural anisotropic biomaterials, whenever anisotropy is required either for homoelasticity or for building scaffold structures [31-34]. The architecture of a knitted fabric is defined by a highly ordered arrangement of interlocked loops. Different structures can be formed by controlling the order in which the loops are made and are interlooped. Each different structure bears its own name. The single weft knit is most simple knit architecture, consisting of sequential left and right interlocks. The structure holds together (is coherent) because a single thread is used for knitting. The interior of the fabric (as produced) has somewhat rougher texture than the exterior. The same fabric turned inside out is known as reverse single jersey. In consequence, the interior surface (blood contacting surface) would be smoother and may therefore cause less turbulence in the blood flow [35]. The fibre orientation and the pore distribution are thus defined during the handling and shaping of the fabric, even in the case of large deformations. Warp knitting refers to the knitting of one or more sets of yarn. Warp knitted fabrics are extremely versatile in the sense that they can be especially designed to resemble woven or weft knitted fabrics in terms of mechanical performance. Wefts knitted fabrics are inherently more porous than wovens. In a knitted fabric, there are two porosities, one defined by the open space inside a loop and the second by the distance between filaments. Furthermore, a third kind of porosity can be induced by the method of assembling knit layers by folding, and rolling. The increased porosity in this class of grafts results in a greater degree of yarn mobility. In consequence, these grafts are more compliant, and easier to handle and suture [36].

Wang et al. [37] fabricated two-ply tubular chitosan conduits by combining the chitosan tube prepared using an industrial warp knitting process (inner layer) with the thermally induced phase separation process (outer layer). The resulting tubes have a biphasic wall structure, with a fibrous inner layer and a semi permeable outer layer as shown by the scanning electron micrographs ( 4). Inner and outer surface of a knitted chitosan tube was mantled with a layer of chitosan/gelatin complex solution, and then freeze-dehydrated by Zhang et al. [38] to develope a tubular scaffold. In vitro characterization showed that the scaffold had a wall of 1.0 mm in thickness with a sandwich structure, and a porosity of 81.2%. The scaffold possessed proper swelling property, burst strength of almost 4000 mmHg, and high suture-retention strength. Vascular smooth muscle cells could spread and grow very well on the scaffold ( 5). Scanning electron microscopy demonstrated that vSMCs spread and grew very well after 48 h on chitosan fibers and the chitosan- gelatin complex demonstrating the feasibility of the scaffold in the field of blood vessel tissue engineering.

Sarah Gundy et al. [39] studied the cytotoxicity of a polylactic acid (PLA) warp-knit textile was assessed with human coronary artery smooth muscle cells (HCASMC). After 3 weeks, there were no obvious cytotoxic effects observed as a result of the knitting process and the gene expression profile did not appear to be altered in the presence of the mesh in the fibrin gel. The addition of the PLA textile to fibrin gel does not seem to have a major effect on the biocompatibility of HCASMC suggesting that the knitting process did not adversely affect the cell response; there was no dramatic change in cell number or metabolic rate compared to the positive control. Overall, the results highlight the potential of a PLA textile/fibrin gel composite scaffold.

The Weilin Xu et al. [40] developed tubular fabric by knitting polyester filament yarns on a specially designed weft-knitting machine. The fabric was used to reinforce polyurethane vascular graft and thus a kind of composite vascular graft was fabricated with a small inner diameter of 4 mm. Strength of the reinforced composite vascular grafts was almost 5-10 times of the strength of the pure PU vascular grafts. Microporous structure can also be fabricated in the wall of the reinforced composite vascular grafts.

In vitro hydrolytic degradation study of bicomponent vascular fabrics made from Dacron and polyglycolic acid (PGA) yarns by Tarng-Jenn Yu et al. [41] showed that the most dramatic changes in the bicomponent fabric characteristics and properties occurred 30 and 60 days of hydrolysis. This schedule coincided with the hydrolytic degradation rate of PGA absorbable sutures. Structural integrity of these fabrics was retained at the end of hydrolytic degradation study. The data obtained could be used to correlate with the subsequent in vivo performance of these bicomponent vascular grafts. If correlations exist, they could be used to improve the design of future bicomponent vascular grafts for improved performance.

3.3 Braided scaffolds

Braiding is a technique that has been used to create products designed to bear axial loads, supply reinforcement, or serve as protective covers. The simplest braids are composed of sets of yarns that follow circular paths in opposite directions with a sequence of crossovers that cause the yarns to interlace forming a fabric [42]. These structures can transfer large loads and provide extension, their design makes them shear resistant and conformable [43, 44].

Bini et al. [45] developed the tubular structure by braiding the required number of PLGA or chitosan fibers onto the spindles of the microbraiding machine. The diameter of the microbraided tubes can be controlled by controlling the diameter of the mandrel. After the microbraided tubes were cut into the required length, the Teflon mandrel was removed leaving the fibrous conduit. 6 shows the schematic of a microbraided conduit. The porosity of the tubular structure obtained by microbraiding can be varied by changing any of the following parameters, the braiding angle, the number of fibers in a spindle, the number of monofilaments in a fiber and the number of spindles. The pore morphology of these tubular scaffolds was mainly due to the cross alignment of fibers.

PCL-PGLA composite was developed by Moet al. [46] by coating a porous layer of PCL on the outside of a PGLA (with GA:LA =90:10) fiber braided tube. The PGLA tube was fabricated using a braiding machine by inserting a Teflon tube with the desired diameter in center of the 20 spindles, which are the carriers of PGLA fibers. Thermally induced phase separation method was used for PCL solution coating on the surface of the PGLA braided tube. The fibroblast cells were seeded on the tubular scaffold and cultured in vitro for the biocompatibility investigation. Histology results showed that the fibroblast cells proliferated on the interconnected pore of the PCL porous layer in 1 week.

3.4 Nonwoven Scaffolds

The most frequently used textile based scaffolds are nonwoven structures which are biodegradable to eliminate any permanent foreign body tissue reaction toward the scaffolds and to generate, over time, more volume space for the engineered tissue to grow [47]. Nonwoven fabrics are broadly defined as sheet or web structures bonded together by entangling fiber or filaments (and by perforating films) mechanically, thermally or chemically. They are flat, porous sheets that are made directly from separate fibers or from molten plastic or plastic film. Fibre bonding methods such as nonwoven technologies could make possible precisely engineered structures, with control over parameters such as fibre orientation and porosity [48].

Nonwovens possess many properties due to which they became famous in medical field, as various parameters such as porosity, weight of fabric, thickness, can be controlled, easy to sterilize nonwovens, various manufacturing technique options exist according to applications and economical manufacturing process etc. The properties of nonwoven fabrics are determined by those of the constituent polymer or fiber and by the bonding process. However, nonwoven porous matrices currently used in tissue engineering have a relatively large porosity and pore size, in the range of several hundred micrometers, and have not been structurally optimized for specific applications. There is a need for a reliable method that can be easily used to modify the microstructure of a nonwoven porous matrix as the scaffold for tissue engineering applications.

Kim et al. [49] fabricated and characterized a new tubular, macroporous, fibrous scaffold using a very elastic biodegradable copolymer, poly(L-lactide-co-caprolactone). The porosity and median pore size of the fibrous PLCL scaffolds were 55-75% and 120-150 μm, respectively. The scaffolds exhibited 500-600% elongation at break. However, the tensile strength and modulus of fibrous PLCL scaffolds, produced from 5% solutions were approximately 4 and 5 times higher than those of extruded PLCL scaffolds suggesting that the fibrous PLCL scaffolds were very elastic and mechanically strong. SEM imaging showed that the scaffolds were well interconnected between the pores. In addition, the cell-seeding efficiency was 2-fold higher using gel-spun scaffolds than using extruded scaffolds.

In the another work by Buttafoco et al. [50], P(DLLA-co-TMC) was processed by melt spinning and fiber winding into porous tubular structures. This approach obviates the need of organic solvents that may compromise subsequent cell culture. The tubes were dipped in a suspension of collagen type I and subsequently freeze-dried, to improve cellular interactions. A collagen microsponge with interconnected pores was formed in the pores of the P (DLLA-co-TMC) porous tubular structure. Dynamically cultured scaffolds were slightly stronger than scaffolds cultured under static conditions.

Dynamically cultured constructs showed SMCs homogeneously distributed throughout the wall. These constructs had improved mechanical properties compared with specimens cultured without mechanical stimulation. Buttafoco et al. concluded that the dynamically cultured hybrid constructs can be used as base structures for tissue engineering of small diameter blood vessels, since their mechanical properties were comparable to those of the human artery mesenterica.

Williamson et al. [51] developed small-diameter vascular grafts which promoted strong attachment of endothelial cells. Composite scaffolds were produced by wet spinning PCL fibres which form the luminal surface, followed by electrospinning of polyurethane (PU) onto the back of the PCL fibres to form the vessel wall substitute. HUVECs showed good attachment to the unmodified PCL-PU composite scaffold (PCL surface uppermost) ( 7). SMC had good attachment to the anti-luminal surface of the scaffold. Human endothelial cells demonstrated strong attachment to the composite PCL-PU scaffold, and proliferated to form a monolayer with strong PECAM-1 expression and cobblestone morphology and the luminal PCL surface of the scaffold supported the formation of stable functional EC monolayers.

K. Tuzlakoglu et al [52] proposed a new route for producing fiber mesh scaffolds from a starch-polycaprolactone (SPCL) blend. Argon plasma treatment was applied to the scaffolds to enhance the cell attachment and proliferation. Surface morphology and chemical composition were significantly changed because of the etching and functionalization of the fiber surfaces. High viability and alkaline phosphatase enzyme activity was shown by the cell seeded on wet-spun SPCL fiber mesh scaffolds with those values being even higher for the cells seeded on the plasma-treated scaffolds.

3.5 Electospun scaffolds

The process of electrospinning, well known for many years in the textile industry and in organic polymer science [53-56], has recently drawn strong attention in biomedical engineering, providing the basis for the fabrication of unique matrices and scaffolds for tissue engineering [57-62]. The potential of applying electrospinning in vascular tissue engineering is enormous since it cannot only mimic the nanosized dimension of natural ECM but also its spatial organization on the mesoscopic scale (control over fibre orientation and spatial placement) [63]. The development of electrospinning technology has enabled the creation of nanofiber-based scaffolds that possess surface topography characteristics that facilitate cell growth. In addition, the electrospinning method consistently produces scaffold materials with defined composition, which allows for controlled degradation during remodeling [64-66]. Electrospinning has been used as an effective method to fabricate biomimetic nonwoven scaffolds that are comprised of a large network of interconnected fibers and pores. This high porosity allows the efficient exchange of nutrients and metabolic waste between the scaffold and its environment, and provides a high surface area for local and sustained delivery of biochemical signals to the seeded cells [67-72]. To date, electrospinning has been used for the fabrication of scaffolds from numerous biodegradable polymers, such as PCL, poly-(lactic acid) (PLA), poly(glycolic acid) (PGA), poly(lactide-co-glycolide) (PLGA) and PU [73-81].

Apart from the simplicity of its setup, the versatility of the electrospinning process has also made it ideal for tissue engineering. The process is compatible with a wide array of polymers, both natural and synthetic in origin, as well as the combination of any number of different polymers. A large number of biopolymers have been electrospun for e.g. collagen, elastin, fibrinogen, hemoglobin and myoglobin [82]. Polymer solutions can be combined in a single reservoir, spun side by side from separate spinnerets, or layered sequentially to provide a large number of options for altering scaffold properties. In addition, electrospun scaffolds can be created in nearly any shape required, and can consist of fibers in an assortment of orientations [83]. Electrospinning provides researchers with the ability to create, not only flat sheets, but also the seamless tubes required for vascular applications. These tubes can be created in any diameter to fit the specifications of any vascular conduit [84].

In this work by Drilling et. al., [85] electrospun tubes of interest for vascular tissue engineering were fabricated and evaluated for burst pressure and suture retention strength (SRS) in the same context as tensile strength providing a direct, novel comparison. Tubes could be fabricated displaying average burst pressures up to 4000 mmHg, well above the standard of 2000 mmHg and SRS values matching those of relevant natural tissues.

In another study, Tillman et al. [86] investigated whether PCL/collagen scaffolds could support cell growth and withstand physiologic conditions while maintaining patency in a rabbit aortoiliac bypass model. . Histological analysis of the retrieved implants suggests an absence of inflammatory infiltrate ( 8). Their results indicated that electrospun scaffolds support adherence and growth of vascular cells under physiologic conditions and that endothelialized grafts resisted adherence of platelets when exposed to blood.

Sang et al. [87] fabricated various scaffolds using the electrospinning technique with blends of collagen, elastin, and several biodegradable polymers. Biocompatibility, dimensional stability in vitro and mechanical properties were evaluated. Materials were blended at a relative concentration by weight of 45% collagen, 15% elastin, and 40% synthetic polymer to mimic the ratio of collagen and elastin in native blood vessels. Addition of synthetic polymers to the collagen/elastin blend improved the mechanical properties and resulted in differing mechanical behavior for vascular substitutes. SEM images of the resulting fibers showed nanoscale fiber diameters and a random orientation of fibers ( 9). It was suggested that the introduction of synthetic biodegradable polymers enabled tailoring of mechanical properties of vascular substitutes and improving compliance matching for vascular tissue engineering.

In another study, fabrication of P(LLA-CL) tubular nanofiber scaffolds and seeding of human coronary artery endothelial cells (HCAECs) onto the lumens of scaffolds by a customized seeding technique was studied by Wei et al. [88]. HCAECs were rotationally seeded onto the lumen of the scaffolds through a customized seeding device, followed with static culture. Results showed HCAECs maintained phenotypic expression of PECAM-1 and the scaffolds sustained the surgical process, kept the structure integrity, and showed the patency for 7 weeks.

Novel Collagen generally has relatively low mechanical strength for the applications in blood vessels. To improve the mechanical properties of porous collagen scaffolds, tubular scaffolds of marine source collagen and PLGA fibers were fabricated by freeze drying and electrospinning processes for vascular grafts by Jeong et al. [89]. As shown in Figs. 18(a)-(d), tubular jellyfish collagen scaffolds prepared by a freeze-drying method showed a highly porous structure. The average pore size and porosity of collagen scaffolds were about 210750 mm and 93%, respectively. The pulsatile perfusion system enhanced the SMCs and ECs proliferation. In addition, a significant cell alignment in a direction radial to the distending direction was observed in tissues exposed to radial distention, which is similar to the phenomenon of native vessel tissues in vivo. These results indicated that the co-culturing of SMCs and ECs, using collagen/PLGA hybrid scaffolds under a pulsatile perfusion system, leads to the enhancement of vascular EC development, as well as the retention of the differentiated cell phenotype.

3.6 Aligned Nanofibrous Scaffolds

Since, it is the nature of the solution jet to follow a chaotic path, it is difficult to control the electrospinning process. Thus most applications of electrospun fibres have been limited to areas where controlled aligned fibres have not been necessary. While ordinary textile methods of obtaining aligned fibres are achieved by winding a single fibre on a rotating drum [90], this is not possible for electrospinning given the chaotic motion of the electrospinning jet and the nanometre size of the fibres. Fibre deposition of significant thickness is restricted to circumferentially aligned fibres. These were obtained by collecting the electrospun fibres on a mandrel rotating at a high speed [91] and using a parallel grid target, which is given a negative charge, behind a rotating mandrel [92].

The copolymer P(LLA-CL) was chosen by Xu et al. [93] as a model polymer due to its lack of toxicity and suitable degradation period for making a unique aligned structure. Favorable interactions between the SMCs and the unique scaffold as well as a directional growth of the cells along the fiber orientation were demonstrated by cell morphology, adhesion and proliferation studies. The SMCs on the aligned nanofibrous scaffold expressed a more functional contractile phenotype while those on the TCPS assumed a synthetic phenotype with few myofilaments inside the cytoplasm. In addition, it can also be found that the organization of the cytoskeleton proteins inside the SMCs was in an orientation parallel to the direction of nanofibers.

Teo et al. [94] proposed a simple method of obtaining a tubular structure with well aligned electrospun nanofibres in a circumferential direction as well as at an angle to the longitudinal axis of the tube. Using a knife-edged counter-electrode, they were able to control the electrostatic field in such a way that we could influence the electrospinning jet. With the ability to obtain circumferentially aligned and diagonally aligned electrospun nanofibres to form an angleply laminate composite, they were able to develop a tube made of nanofibres with a greater versatility in itsmechanical strength and possibly a blood vessel scaffold closer to that of its natural extracellular matrix. When comparing the degree of alignment of the electrospun nanofibres collected on the Teflon tube using a parallel grid made of aluminium strips with those collected using a parallel grid made of knife-edged aluminium bars, it was found that fibres collected on the Teflon tube using knife-edged aluminium bars showed a greater degree of circumferential alignment.

4. Surface modification of scaffolds

The interaction between biomaterials and physiological environment is very important in tissue engineering. The cell surface has a variety of receptors that bind with other cells or specific proteins, which compose the surrounding of the cell. In order to develop the biomaterials that can promote ideal cellular responses, the molecular interaction of cells with extracellular components must be understood. In tissue engineering, cell adhesion to implant surface is critical because cell adhesion occurs before other biological events including cell spreading, cell migration and differentiation and cell function. Cell adhesion is closely related to the surface properties of biomaterials. It is commonly accepted that the adhesion of cells to solid substrata is influenced by several substratum surface properties, such as wettability, surface charge, roughness and topography. Many surface modification techniques, as summarized in table III have been used to produce various surface properties of polymers. Most conventional materials do not meet the criteria for serving as tissue engineering scaffolds. Surface modification is an effective approach to alter biological interactions of a particular material to develop appropriate scaffolds. Since surface modification only changes the outermost surface composition of a biomaterial, its bulk properties do not change. In addition, surface modification can provide accessible and chemical functional groups for the immobilization of drugs, enzymes, antibodies or other biologically active species for a variety of biomedical applications. The main reason for the long-term failure of the small-diameter vascular grafts is the incomplete cover of endothelial cells (ECs) on the vascular graft surfaces and the subsequent myointimal hyperplasia [95,96]. Surface modifications that have been used to enhance endothelial cell seeding on vascular prostheses include immobilization of fibronectin, laminin, collagen, and peptides [97-100]. Because most synthetic biodegradable polymers are hydrophobic, extensive efforts have thus been devoted towards increasing the biomaterial's hydrophilicity. One convenient measure is plasma treatment, which can easily introduce polarized groups such as hydroxyl, carboxyl, amino and sulfate groups on polymer surfaces using different reaction gases such as air, NH3, SO2, CO2 or other organic compounds.

Plasma treatments can modify the surface of vascular prostheses without dramatic changes in physical properties and graft dynamics [101,102]. Amide and amine plasma (butylamine) were applied to graft surfaces by Tseng et al. [103] using radio frequency glow discharge. Bovine aortic endothelial cells were seeded on amide and amine plasma coated ePTFE vascular grafts and placed inside an artificial circulatory system under well-defined flow conditions. Plasma coatings increased endothelial cell adhesion to PTFE to an extent virtually identical to the effects on water contact angle and surface tension. Fluorescence nuclear staining, scanning electron microscopy, and histological staining indicated the formation of an endothelial cell monolayer on the plasma coated graft surfaces.

In another study, by Jui-Che Lin et al. [104] the inner surface of small diameter LDPE tubing was modified by sulphur dioxide (SO,) plasmas in order to incorporate acidic sulphur-containing functionalities, such as sulphonate (-SO3-), onto the surface. The inner surface of the small diameter LDPE tubing was uniformly modified by SO2 plasma, HMDSO plasma and SO,-HMDSO gas mixture plasmas. Ex vivo blood compatibility studies indicated that the most hydrophilic surface, created by the SO2 plasma modification, is more thrombogenic than the untreated control. In contrast, surfaces modified by the plasma polymerized HMDSO and SO,-HMDSO gas mixtures exhibited thrombogenicity similar to that of the control polyethylene sample. These observations may be caused by the effects of both the surface chemical functionality (e.g. -SO3 or -G02H) and the surface hydrophilicity.

Another widely used method is to graft hydrophilic polymers on biomaterial's surfaces through grafting copolymerization of hydrophilic polymers. The introduction of initiators (radicals or peroxide groups) on the chemically inert surfaces can be realized by argon plasma treatment, ozone oxidation, gamma-ray, electron beam and laser treatment [105-108]. Cerium (IV) induced grafting polymerization is worthy of notation in that it is a pure wet chemistry method and does not need any special equipment like irradiation source and plasma or ozone generator. The reaction mechanism is

−CH2OH + Ce4+ =−C*HOH + H+ + Ce3+,

Where the asterisk stands for a carbon free radical. A limitation of the cerium (IV) induced grafting polymerization is that it requires polymers possessing alcoholic hydroxyl groups, such as polyvinyl alcohol [109] and cellulose [110]. However, it has been reported that the highly oxidative Ce(IV) can also attack organic substrates without hydroxyl groups to undergo the electron transfer reaction to give radicals [111-112].

In this work, by Zuwei et al [113] a conventional polymer used in vascular graft, PET, was processed into non-woven nanofiber mat (NFM) via electrospinning. To overcome the chemical and biological inertness of the PET surface, gelatin was covalently grafted onto the PET NFM surface. The electrospun PET NFM was first treated in formaldehyde to yield hydroxyl groups on the surface, followed by the graft polymerization of methacrylic acid (MAA) initiated by Ce(IV). Finally, the PMAA-grafted PET NFM was grafted with gelatin using water-soluble carbodiimide as coupling agent. Plane PET film was also surface modified and characterized for basic understanding of the surface modification process. The whole chemical reaction scheme for PET surface modification is shown in 10. ECs were cultured on the original and gelatin-modified PET NFM and the cell morphology, proliferation and viability were studied. The COOH density grafted on the PET film/NFM increased with the grafting time.

Nanofiber scaffolds with amino groups conjugated to fiber surface through different spacers (ethylene, butylenes, and hexylene groups, respectively) were prepared and the effect of spacer length on adhesion and expansion of umbilical cord blood hematopoietic stem/progenitor cells (HSPCs) was investigated by Chua et al. [114]. Electrospun polymer nanofiber scaffolds were functionalized with PAA grafting, followed by conjugation of amino groups with different spacers. HSPCs were expanded on aminated scaffolds for 10 days. Aminated nanofiber scaffolds with ethylene and butylene spacers showed high expansion. Both EtDA and BuDA modified nanofiber meshes mediated significant adhesion. This study demonstrated that aminated nanofibers are superior substrates for ex vivo HSPC expansion, which was correlated with the enhanced HSPC adhesion to these aminated nanofibers. These results highlighted the importance of scaffold topography and cell-substrate interaction to regulating HSPC proliferation and self-renewal in cytokine-supplemented expansion.

O-Carboxymethylchitosan (OCMCS) is one kind of blood-compatible chitosan derivative [115]. O-Carboxymethylchitosan (OCMCS) was covalently immobilized onto expanded poly(tetra-fluoroethylene) (ePTFE) vascular graft using a photosensitive hetero-bifunctional crosslinking reagent, 4-azidobenzoic acid by Zhu et al [116]. 4-Azidobenzoic-bonded OCMCS (Az-OCMCS) was prepared by reaction between an acid group of the crosslinking reagent and a free amino group of OCMCS. Immobilization was accomplished by irridiating the A3-OCMCS coating on the substrate surface with UV light. The reaction scheme for immobilization is shown in 11. The hydrophilicity of ePTFE was enhanced greatly by surface bonding with OCMCS. In vitro platelet adhesion results indicate that OCMCS immobilized on the ePTFE/OCMCS vascular graft surface can effectively reduce the fibrinogen adsorption and platelet adhesion and activation, demonstrating good biocompatibility.


Research in tissue-engineered vascular grafts has focused on improving scaffold design to impart mechanical properties that include tensile stiffness, compliance and elasticity. Textile scaffolds may be embroidered, knitted, woven, nonwoven, braided or of composite construction depending on the intended application. Most strikingly, these are some of the simplest structures found in textile products. Regardless of the simplicity of the structures, it is important to realize that these prostheses possess significantly different properties in terms of porosity, bursting strength, thickness, etc. Woven scaffold exhibit high bursting strength, minimum tendency to fatigue, and they can be fabricated tightly enough to lower permeability to water (and blood). However, they are extremely difficult to handle and suture. Moreover, they are noncompliant and have a tendency to fray at the edges. Weft knitted fabrics are inherently more porous than wovens. The increased porosity in this class of grafts results in a greater degree of yarn mobility. In consequence, these grafts are more compliant, and easier to handle and suture. Warp knitted fabrics are extremely versatile in the sense that they can be especially designed to resemble woven or weft knitted fabrics in terms of mechanical performance. Nonwovens possess many properties due to which they became famous in medical field, as various parameters can be controlled easily like porosity, weight of fabric, thickness; easy to sterilize nonwovens; various manufacturing technique options exist according to applications and economical manufacturing process etc. Surface modifications that have been used to enhance endothelial cell seeding on vascular prostheses include immobilization of fibronectin, laminin, collagen, and peptides. These biotextile structures, whether woven, knitted, or nonwoven, are believed to be uniquely suited to serve as tissue engineering scaffolds and to compare favorably with other fabrication techniques. Textile technology has provided several solutions for vascular surgery and a large number of textile vascular prostheses have been implanted in patients to revascularize districts downstream from diseased or injured arteries. The above review take a closer look at textile based composite structure applications tissue engineering scaffold for blood vessels.

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