Review on electrospun nanofibers scaffold and biomedical applications.
Abstract: Recently, nanotechnology as novel interdisciplinary sciences has been introduced among all fields and gets numerous attentions, due to its unique applications. In biomedical engineering, electrospinning exhibits a lot of advantages as a nanofibers scaffolds producer, which can make appropriate resemblance in physical structure with extra cellular matrix (ECM). This is because of the nanometer scale of ECM fibrils in diameter, which can be mimicked by electrospinning procedure as well as its porous structure. In this review, we attempt to explore the electrospun nanofibers scaffolds applications in biomedical engineering.
Subject: Tissue engineering
Authors: Kanani, A. Gholipour
Bahrami, S. Hajir
Pub Date: 08/01/2010
Publication: Name: Trends in Biomaterials and Artificial Organs Publisher: Society for Biomaterials and Artificial Organs Audience: Academic Format: Magazine/Journal Subject: Health Copyright: COPYRIGHT 2010 Society for Biomaterials and Artificial Organs ISSN: 0971-1198
Issue: Date: August, 2010 Source Volume: 24 Source Issue: 2
Accession Number: 308129448
Full Text: Introduction

Tissue engineering is an emerging multidisciplinary field involving biology, medicine, and engineering that is likely to revolutionize the ways which can improve the health and quality of life for millions of people worldwide by restoring, maintaining, or enhancing tissue and organ function. it has shown great promise in generating living alternatives for harvested tissues and organs for transplantation and reconstructive surgery. The overall goal of tissue engineering is to apply the foundations and innovations of biology, medicine, and engineering to develop and manipulate viable, three dimensional physiologic substitutes that are capable of reinstating, sustaining, or recovering the function of tissues and organs. Composition (i.e., biomaterials of synthetic or natural origin) and architecture of a tissue engineered scaffold result in cell environment interactions that determine the structure's fate. The ultimate goal is to enable the body (cellular components) to heal itself by introducing a tissue engineered scaffold that the body recognizes as "self", and in turn, uses to regenerate "neo-native" functional tissues. It has long been hypothesized that in order to duplicate all of the essential intercellular reactions and promote native intracellular responses, the ECM must be mimicked. These synthetic ECMs or scaffolds must be designed to conform to a specific set of requirements [1, 2].

In this review, we would briefly cover the Historical records of tissue engineering. As it comes from the "tissue engineering", it is directly related to tissue and living cells. Then it is opportune to reminisce over Hooke [3] (1635-1703) who discovered small holes in cross-sections, which he called cells. In 1858, Virchow's ideas was described which was about cell formation with the now famous words, "Omnis cellula e cellula ..." it means that cells arise from pre-existing cells [4]. He presented his ideas about regeneration stating that tissue regeneration is dependent on cell proliferation. In 1874, Thiersch [5] discovered the important influence of granulation tissue on wound healing, during his attempt to grow skin cells into granulating wounds. 23 years later, Loeb [6] reported the idea of growing cells outside the human body. From that point on many researchers experimented with growing cells in vitro. While experimenting with different media, survival was established, despite no growth. Harrison [7] (1870-1959) was the first who grow frog ectodermal cells in vitro in 1907, thus the first neuronal tissue culture line was developed. Lately, much attention has been paid to growing cells instead of complete tissues. In 1916, Rous and Jones [8] discovered that trypsin is capable of degrading matrix proteins, thus separating cells. Trends changed toward expansion of cell types throughout the 1940s and soon after, much research was performed, which led to the ability to grow tissue-specific cell lines in vitro. As recently as the 1980's, the term tissue engineering was loosely applied to the use of prosthetic devices and the surgical manipulation of tissues [9]. In 1998, the development of stem cell lines formed the basis of modern tissue engineering [10]. Research since 1998 has focused on identifying the differential capacities of embryonic and adult stem cells and how to influence and accompany this differentiating pathway. Furthermore, fundamental research is continually being performed to gain insight into the cellular interactions that are of great importance in tissue engineering. Dr. Vacanti [9] has been investigated the history of tissue engineering extensively in his recent review.

Tissue engineered scaffolds, due to their unique properties, have been utilized widely. It has been recognized that tissue engineering offers an alternative technique to tissue transplantation for diseased or malfunctioned organs. Millions of patients suffer from end-stage organ failure or tissue loss each year. In the United States alone, at least eight million surgical operations are carried out annually, requiring a total national healthcare cost exceeding US$400 billion per year [11,12].

Scaffolds play a critical role in tissue engineering. The function of scaffolds is to direct the growth of cells either seeded within the porous structure of the scaffold or migrating from surrounding tissue. The prerequisite physicochemical properties of scaffolds are many: to support and deliver cells; induce, differentiate, and channel tissue growth; target cell-adhesion substrates; stimulate cellular response; provide a wound-healing barrier; be biocompatible and biodegradable; possess relatively easy processability and malleability into desired shapes; be highly porous with a large surface/volume ratio; possess mechanical strength and dimensional stability; and have sterilisability, among others [12,13]. Generally, three-dimensional porous scaffolds can be fabricated from natural and synthetic polymers, ceramics, metals, composite biomaterials, and cytokine release materials. There are some common methods for fabrication an emulated scaffold to imitate the structure and functional biology of native ECM. Selfassembly, phase separation, and electrospinning have been utilized to improve the nanofibers diameters to be look like ECM fibrils diameter which is in the range of 50-500 nm.

Electrospinning is a versatile technique which recently has significant development in nanofibers production. In this facile technique, a high electric field is applied to the droplet of a fluid which may be a melt or solution coming out from the tip of a die, which acts as one of the electrodes. This leads to the droplet deformation and finally to the ejection of a charged jet from the tip of the cone accelerating towards the counter electrode leading to the formation of continuous fibers [14]. Due to special characterization of electrospun mat (high surface area to volume ratio, flexibility in surface functionalities, and mechanical properties superior to larger fibers), many research [15-22] have been performed to improve some potential applications of nanofibers include: tissue engineering scaffolds, filtration devices, sensors, materials development, electronic applications and etc..

There is a strenuous interest in electrospun tissue engineered scaffold due to its great potential to pattern the native ECM. The basic requirements for a material to be used for tissue engineering purposes is biocompatibility, although During the past two decades significant advances have been made in the development of biodegradable polymer so biodegradability is one of the most important properties, as the scaffold should degrade with time and be replaced with newly regenerated tissues. Nair and Laurencin classified the biodegradable polymer and investigated the most applicable ones. Poly(a-ester)s, especially lactic acids, glycolic acids and their copolymers with 3-caprolactone, are the most commonly known and used among all biodegradable polymers for fabrication of novel materials for medical use and for tissue engineering applications. Among natural polymers, collagen, chitin and its N-acetyled derivative (chitosan), fibronectin, gelatin,... are the most popular polymers for biomedical application[23]. Various synthetic and natural biopolymers have been electrospun to satisfy different clinical requirements.

Biopolymers (solution or melt) would be electrospun using high voltage (10-40 KV) source in average distance (10-20 cm) and collected on a conductive target (an electrically grounded metal sheet or a winder.) with mediocrity diameter 100-500 nm. Aligned nanofibrous scaffolds can be produced using two parallel conductive blades or an adjustable rotating drum as collector [24, 25].

This review briefly discusses nanofibers impressionable properties by changing electrospinning parameters and widely describes functional electrospun scaffolds in biomedical applications. There are many applications for electrospun tissue engineered scaffolds, that here we introduced some of them and investigated some recent works on each of them. Classification of electrospun nanofibers scaffolds based on their functional application, are: Wound dressing; Drug delivery; Blood vessel (vascular tissue engineering); Bone tissue engineering; Heart tissue engineering; Cartilage tissue engineering etc.

Wound dressing

Dressing for wounds is a vital function which led to protection, removal of exudates, inhibition of exogenous microorganism invasion, and improved appearance. In foretime, due to empirical methods, it has been reported that work on blockaded wound dressing began since human observed a blister healed faster if left unbroken. As mentioned above, dressings play a substantial role in Conglutination of certain types of open wounds (e.g., traumatic, thermal, or chronic wounds), since the moist, warm and nutritious environment of wound beds provides an ideal condition for microbial growth. Ideal antimicrobial dressings should have a number of key attributes, including provision of a moist environment to enhance healing, and broad-spectrum antimicrobial activity, including activity against antibiotic-resistant bacteria [26-29].

Nanofibrous wound dressing using electrospinning process potentially offers numerous advantages in comparison with conventional processes. Nanofibers innate properties include high surface area and microporous structure, afford quickly start signaling pathway and attract fibroblasts to the derma layer, which can excrete important extracellular matrix components, such as collagen and several cytokines (e.g. growth factors and angiogenic factors), to repair damaged tissue. The electrospun membrane is also important for cell attachment and proliferation in wound healing [30]. Nanofibers are usually obtained in non-woven form, which is very suitable for applications such as wound dressings. The nonwoven mats usually have pores which are small enough to prevent bacterial penetration. The high surface area is of importance for fluid absorption and dermal drug delivery. Although numerous polymers have been successfully electrospun into nanofibers [15-22] the reports on electrospun nanofibrous mats suitable for wound dressings are still scarce.

Electrospinning has been applied to prepare collagen [28,29] hydrophilic polymers (poly(vynil alcohol)(PVA), poly(etylen oxide)(PEO), gelatin, chitin and chitosan, polyesters and polyurethane [31] nanofibrous materials which might be used as wound dressings.

The majority of skin substitutes are comprised of fibroblasts and/or keratinocytes on collagen scaffolds thus, for the choice of proper biomaterials in wound dressing, the role of collagen at each phase of wound healing is well understood and appreciated. In this regard, Powell and coworkers [32] fabricated bovine collagen nanofibrous scaffolds using either freeze-dry (FD) or electrospinning (EC) processes (Fig 1). They reported that no significant differences were observed in cell proliferation, surface hydration, or cellular organization between the electrospun collagen skin substitutes (ECSS) and freeze-dry collagen skin substitutes (FCSS) of in vitro evaluations. After grafting to full thickness wounds in athymic mice, the rates of engrafting were 87.5% and 100%, in FCSS and ECSS, respectively. In comparison, histological evaluations of wounds revealed that bovine collagen persisted in the wound at week 8 in the FCSS group while no bovine collagen was seen in the ECSS group. At 8 weeks post-grafting, the ECSS grafts were 61.3 [+ or -] 7.9% original graft area whereas the FCSS grafts were 39.2 [+ or -] 8.8% original area (p0.01). They concluded that ES scaffolds can be better than FD scaffolds as skin substitutes due to its optimal cellular organization.

[FIGURE 1 OMITTED]

In another work, Rho and coworkers [33] reported producing of collagen type I nanofibrous matrix by the electrospinning process for the application of wound dressing. The average diameter of collagen nanofibers electrospun from 8% collagen solution in 1, 1, 1, 3, 3, 3-hexafluoro-2-propanol (HFIP) was 460nm (range of 100-1200 nm). The as-spun collagen nanofibrous matrix was chemically cross-linked by glutaraldehyde vapor, during this process, porosity of the collagen matrix decreased from 89% to 71%. Three groups of nanofibrous scaffold were investigated: uncoated collagen nanofibers, collagen nanofibers treated with collagen type I and treated with laminin. They examined the effects on cytocompatibility, cell behavior, cell and collagen nanofiber interactions, and open wound healing in rats. Relatively low cell adhesion was observed on uncoated collagen nanofibers, whereas collagen nanofibrous matrices treated with type I collagen or laminin were functionally active in responses in normal human keratinocytes. Powell and coworkers [32] investigated the differences between two common methods for collagen nanofibers scaffolds and reported that electrospun scaffold was a better choice for skin substitutes than freeze-dried one. On the other hand, Rho and coworkers [33] conclude that among 3 type scaffolds of electrospun collagen nanofibers, the treated ones with collagen or laminin showed better results. Some attempts have been performed to fabricate the blend nanofibrous mat of collagen with other biocompatible polymers such as chitosan, PEO... [30, 34-36]. Chen and coworkers [30] prepared composite nanofibrous membranes (NFM) of type I collagen, chitosan, and polyethylene oxide by electrospinning, which could be further crosslinked by glutaraldehyde vapor. Nanofibers with 134 [+ or -] 42 nm in diameter were obtained; the diameters get increased after crosslinking as well as Young's modulus, while the ultimate tensile strength, tensile strain, and water sorption capability decreased after crosslinking. NFM was subjected to in vitro biocompatibility examine and showed no cytotoxicity toward growth of 3T3 fibroblasts. The NFM was compared with gauze and commercial collagen sponge wound dressing in animal studies (test was done on SD rats and the surface area of wound was 2 cm x 2), and explored better result in wound healing rate.

[FIGURE 2 OMITTED]

The natural ECM proteins in the body are primarily composed of collagen and proteoglycans, whose compositions depend on tissue type. Chitin and chitosan has structural similarity to glycosaminoglycans (GAGs; main component of proteoglycans), on the other hand their antibacterial activity convert these polymers to appropriate choices for tissue engineering applications. Muzzarelli [37] in his recent review investigated chitin and chitosan effects on repairing the different wounds. There are variable works on chitin which electrospun in pure form [38] or blend [39] with other polymers. Noh and coworkers [38] used electrospinning method to fabricate chitin nanofibrous matrices for biodegradability and cell behavior test. In this study, they compared the resulted chitin nanofibers (Chi-N) with commercial chitin microfibers (Beschitin W[R]; Chi-M) and the result showed that, during in vitro degradation for 15 days, the degradation rate of Chi-N was faster than that of Chi-M. For in vivo study, Chi-N was grafted into rat subcutaneous tissue, it had almost degraded within 28 days, and no inflammation could be seen on the nanofiber surfaces or in the surrounding tissues (except in the early stage wound). Fig 2 shows Cell attachment and spreading of normal human keratinocytes and fibroblasts plated onto chitin nanofibrous and microfibrous matrices without ECM protein coating. Relatively high cell attachment and spreading of cells tested were observed on Chi-N in comparison to Chi-M, and Chi-N treated with type I collagen significantly promoted the cellular response. They indicated that chitin nanofibers have proper characterizations for wound healing application.

Chitosan as the most demotic derivative of chitin has got more attentions due to its good solubility in inorganic and some organic solvents such as acetic acid. Bhattarai and coworkers [40] produced chitosan-based nanofibers by electrospinning solutions containing chitosan, polyethylene oxide (PEO), and Triton X_100. They studied the potential use of the matrix for tissue engineering by examining its integrity in water and cellular compatibility. They found that that the matrix with a chitosan/PEO ratio of 90/ 10 retained excellent integrity of the fibrous structure in water. Experimental results from cell stain assay and SEM imaging showed that the nanofibrous structure promoted the attachment of human osteoblasts and chondrocytes and maintained characteristic cell morphology and viability throughout the period of study. Xu and coworkers [41] prepared chitosan/ Poly (lactic acid) (PLA) blend micro/nanofibers by electrospinning. The using co-solvent was trifluoroacetic acid (TFA). It was found that the average diameter of the chitosan/PLA blend fibers became larger, and the morphology of the fibers became finer with the content of PLA increasing. The authors believed that the spun micro/nanofibers are expected to be used in the native extracellular matrix for tissue engineering.

Gholipour et.al fabricated nanofibrous web made of chitosan-PVA blend solutions in different blend ratios. They characterized the web by SEM, FTIR and DSC and found that the 25/75 Cs/PVA is the best ratio in applying 15 Kv voltages and 15 cm distance. From in vitro studies, they demonstrated that resulted web shows good antimicrobial effect against Gram-negative bacteria (Pseudomonas aeroginosa) and it could be prepared for wound dressing applications [42].

Gelatin is another natural polymer which its sufficiency in biomedical application is undeniable. Rujitanaroj and coworkers [43] prepared ultrafine gelatin fiber mats with antibacterial activity against some common bacteria found on burn wounds, from a gelatin solution (22%w/v in 70 vol% acetic acid) containing 2.5 wt% AgNO3. Silver nanoparticles (nAg), a potent antibacterial agent, first appeared in the AgNO3-containing gelatin solution after it had been aged for at least 12 h, with the amount of nAg increasing with increasing aging time. The average diameters of the as-formed nAg ranged between 11 and 20 nm. Electrospinning of both the base and the 12 h-aged AgNO3-containing gelatin solutions resulted in the formation of smooth fibers, with average diameters of w230 and w280 nm, respectively. The antibacterial activity of these materials, regardless of the sample types, was greatest against Pseudomonas aeroginosa, followed by Staphylococcus aureus, Escherichia coli, and methicillin-resistant S. aureus, respectively. Electrospun gelatin fiber mats containing silver nanoparticles (nAg) are proposed to be used as antibacterial dressings. The nAg-containing gelatin fiber mats were characterized for release characteristics of the as-loaded silver, as well as for their antibacterial activity against some common bacteria found on burn wounds.

Polyurethane (PU) is frequently used in wound dressings because of its unique properties (good barrier properties and oxygen permeability). Khil and coworkers [31] fabricated electrospun PU membrane for wound dressing application. The nanofibrous PU membrane showed controlled evaporative water loss, excellent oxygen permeability, and promoted fluid drainage ability due to the nanofibers with porosity and inherent property of PU. Neither toxicity nor permeability to exogenous microorganism was observed with the nanofibrous membrane. Histological examination confirmed that epithelialization rate was increased, and the exudate in the dermis was well controlled by covering the wound with the electrospun membrane. Thus, nanofibrous PU membrane prepared by electrospinning could be properly employed as wound dressings. One of the variable forms of production of wound dressing is incorporating the drug on the nanofibrous substrate to release the drug in the wound position to enhance healing. Verreck and coworkers [44] selected PU as a nonbiodegradable polymer and electrospun it to a drug-laden (Itraconazole and ketanserin) nanofibrous mat for potential use in drug administration and wound healing. For both itraconazole and ketanserin, an amorphous nanodispersion with PU was obtained when the drug/polymer solutions were electrospun from dimethylformide (DMF) and dimethylacetamide (DMAc), respectively. They observed that at low drug loading, itraconazole was released from the nanofibers as a linear function of the square root of time suggesting Fickian kinetics with no initial drug burst. Ketanserin showed a biphasic release pattern. They concluded a number of potential applications of electrostatically spun fibers in controlled drug delivery based on the following observations: (1) release of poorly watersoluble drugs can be achieved from a waterinsoluble polymer and (2) the rate of drug release can be tailored by changing the drug/ polymer ratio. For both itraconazole and ketanserin, an amorphous nanodispersion with PU was obtained in the form of a nonwoven fabric.

Thakur and coworkers [45] reported that they successfully fabricated a dual drug release electrospun scaffold based poly-l-lactic acid (PLLA) containing an anesthetic, lidocaine, and an antibiotic, mupirocin with a dual spinneret electrospinning apparatus. Lidocaine hydrochloride exhibited an initial burst release (80% release within an hour) followed by a plateau after the first few hours, while Mupirocin exhibited only a 5% release in the first hour before experiencing a more sustained release to provide antibacterial action for over 72 h. For comparative purposes, both drugs were spun from a single spinneret and evaluated to determine their release profiles. it was found that the scaffold maintained its antibiotic activity throughout the processes of electrospinning and gas sterilization and supported cell viability. They demonstrated that the presence of the two drugs in the same polymer matrix altered the release kinetics of at least one drug. Based on the release profiles obtained, the dual spinneret technique was the preferred method of scaffold fabrication over the single spinneret technique to obtain a prototype wound healing device. In a recent attempt, Lee and coworkers [46] have electrospun Very elastic poly (l-lactide-co-a-caprolactone) (PLCL) (50:50) copolymer blended with gelatin into microfibers from a hexafluoroisopropanol solution. Their observations included that PLCL fiber sheet exhibited the unique soft and flexible behavior while gelatin fiber was hard and brittle. As the gelatin content of PLCL/gelatin fibers increased, Young's modulus was increased, but the elongation was decreased compared to those of PLCL. However, fibers containing 10-30 wt% gelatin demonstrated an enhanced tensile strength with still high elongation to be beneficial for tissue engineering scaffolds. Fibroblasts (NIH-3T3) cell culture was done for evaluating the cytocompatibility of electrospun fiber. Results showed that the cell proliferation exhibited a completely different and strong dependence on the fiber composition: a very high proliferation rate on PLCL90/gelatin10, followed by PLCL > gelatin > PLCL70/gelatin30. The authors concluded that due to enhanced effect of gelatin, especially at 10 wt% content, on strength and cytocompatibility of PLCL/ gelatin fibers would be very preferable for tissue engineering scaffolds.

Kumbar and coworkers [47] prepared Electrospun poly (lactic acid-co-glycolic acid) (PLAGA) scaffolds with fiber diameters of 150-225, 200-300, 250-467, 500-900, 600-1200, 2500-3000 and 3250-6000 nm for skin tissue engineering. Their investigations demonstrated that all fiber matrices have a tensile modulus from 39.23 [+ or -] 8.15 to 79.21 [+ or -] 13.71 MPa which falls in the range for normal human skin. Further, the porous fiber matrices have porosity between 38 to 60% and average pore diameters between 10 to 14 im. The efficacy of these matrices as skin substitutes was evaluated by seeding them with human skin fibroblasts (hSF). Human skin fibroblasts acquired a well spread morphology and showed significant progressive growth on fiber matrices in the 350-1100 nm diameter range. Collagen type III gene expression was significantly up-regulated in hSF seeded on matrices with fiber diameters in the range of 350-1100 nm. Based on the need, the proposed fiber skin substitutes can be successfully fabricated and optimized for skin fibroblast attachment and growth.

Drug delivery

Electrospinning supplies great flexibility in selecting materials for drug delivery applications. Either biodegradable or non-degradable materials can be used to control whether drug release occurs via diffusion alone or diffusion and scaffold degradation. Electrospun fibers can be oriented or arranged randomly, giving control over both the bulk mechanical properties and the biological response to the scaffold. Nowadays, all kinds of drugs such as antibiotics, anticancer agents and proteins, DNA, and RNA can be incorporated into electrospun scaffolds. Using the various electrospinning techniques the applications of electrospinning in tissue engineering and drug delivery are almost unlimited. There are a number of different drug loading methods like coatings, embedded drug, and encapsulated drug (coaxial and emulsion electrospinning). These techniques can be used to give precious control over drug release kinetics [48]. Controlled drug delivery systems have gained much attention in the last few decades. This is due to the many advantages compared with the conventional dosage forms such as improving therapeutic efficacy, reducing toxicity by delivering them at a controlled rate. The main advantages of the fibrous carriers are that they offer site-specific delivery of drugs to the body. Also, more than one drug can be encapsulated directly into the fibers.

Sill and Recum [48] in their recent review mentioned electrospun nanofibrous scaffolds and their applications in drug delivery and tissue engineering. There are also some review attempts regarding surface engineering and drug releasing scaffolds [49] and suitable polymers in which drug could be loaded [50].

There are various examples of using biodegradable and non-biodegradable polymers for drug delivery in electrospinning process. In this way, Kenway et.al [51] prepared nanofibers which were made either from polycaprolactone (PCL) as a biodegradable polymer or polyurethane (PU) as a non-biodegradable polymer or from blends of the two. Ketoprofen was embedded in polymer solutions, and its release profile was followed by UV-visible spectroscopy in phosphate buffer of pH 7.4 at 37 and 20 [degrees]C. regardless the differences in inherent properties of the polymers, the result showed similar release rates of the drug form the polycaprolactone, polyurethane and their blend nanofibers. The significant difference was in visual mechanical properties which get improved in the blend of PCL with PU. They also investigated the release profiles from electrospun mat compared to cast films. In their previous work Kenawy et al. [52] electrospun nanofibrous mats composed of PLA, poly (ethylene-co-vinyl acetate) (PEVA), and a 50:50 blend of the two polymers. The release of tetracycline hydrochloride from electrospun mats was examined. They pointed that both polymer composition and drug loading affected the rate of drug release with PEVA demonstrating quicker release than either the 50:50 PEVA/PLA blend or PLA. They also compared the release profile obtained from electrospun mats to those obtained from corresponding cast films. Due to their larger surface areas, electrospun mats tended to give greater release of drug than did the corresponding films.

In another work, Zong and coworkers [53] used an electrospinning method to fabricate bioabsorbable amorphous poly (D, L-lactic acid) (PDLA) and semi-crystalline poly (L-lactic acid) (PLLA) nanofiber non-woven membranes for biomedical applications. They apply one of the most popular antibiotics, Mexofin, in polymeric solutions and electrospun into nanofibrous non-woven mats. The resulting membrane exhibits very uniform structures with an average diameter of 160 nm. The release profile of the drug from membranes was determined with the UV spectroscopic technique by measuring the absorbance at 234 nm as a function of time. It was found over 90% typical loading efficiency of Mefoxin in the PDLA sample by electrospinning. Finally, the authors concluded that the drug functionality seems to be completely unaffected by the gentle electrospinning process.

Yang et. al [54] fabricated Gelatin/PVA bicomponent nanofibers via electrospinning, and investigated its control release of Raspberry ketone(RK). The result showed that the burst release of drug was observed in the first hour, and reached a plateau after two hours. There were three parameters which could adjust RK release rate: GEL/PVA ratio, the content of loaded RK, and the crosslinking time of glutaraldehyde vapor. The authors found that The RK loaded GEL/PVA electrospinning nanofibers should have potential application in controlled drug delivery. They documented their hypothesis based on following observations: (1) the addition of PVA helped to enhance both the tensile strength and elongation at break of the membrane and (2) the RK release rate could be tailored by changing the content of RK in GEL/PVA matrix, the ratio of GEL and PVA, and the crosslinking time by glutaraldehyde vapor.

Electrospinning of emulsions composed of an organic poly (L-lactide) solution and an aqueous protein solution was reported by Maretschek and coworkers [55]. Cytochrome C was chosen as a hydrophilic model protein for encapsulation. It was shown that the protein release was dependent on the surface tension of the release medium. Electrospinning of emulsions consisting of an organic solution of PLLA and an aqueous solution of hydrophilic polymers yielded fibers composed of a polymer blend. The resulting nanofibers nonwovens exhibited a less hydrophobic surface, which gave us the opportunity to tailor the release profile via this technology. Furthermore it was explained how the addition of different amounts of hydrophilic polymer to the aqueous phase influenced the morphology of the resulting nanofibers nonwovens.

Blood vessel (vascular tissue engineering)

As it is known, vascular structures are composed of three common layers named tunica intima, tunica media and tunica adventitia. The tunica intima forms the innermost lining with non-thrombogenic monolayer endothelial cells and due to secreting specific molecules like nitric oxide, endothelial cells inhibit platelet activation and prevent thrombus formation. The tunica media is generally composed of a dense population of concentrically organized smooth muscle cells and is separated from the tunica intima by an internal elastic lamina. The tunica adventitia forms the external layer and contains a collagenous extracellular matrix and fibroblast cells [56, 57]. Similar to other tissues, the main part of a vascular is its ECM. In summery the basic materials of the extracellular matrix surrounding the vascular cells consist of collagen (types I and III), elastin, some proteoglycans and glycoproteins. In vascular tissue engineering, tensile stiffness, elasticity and compressibility are the major mechanical properties. Collagen provides the tensile stiffness for the resistance against rupture, elastin dictates the elastic properties, proteoglycans contribute to the compressibility; and combined with collagen, elastin prevents irreversible deformation of the vessel against pulsatile blood flow [58].

First in 1986 Weinberg and Bell [59] reported about fabrication of artificial artery based on collagen scaffold. This report was the outset of vascular tissue engineering which has been followed by many researches till now. Nowadays nanotechnology development improves the tissue engineering applications due to unique similarity of resulted nanostructure to native ECM. It is the main reason for utilizing Electrospinning method to mimic the tissue structures such as blood vessel. In this way, Mo et al. [60] fabricated Poly(l-lactide-co-a-caprolactone) [P(LLA-CL)] elecrospun nanofibers and investigated the relationship between electrospinning parameters and fiber diameter. From SEM photographs, the fiber diameter decreased with decreasing polymer concentration and with increasing electrospinning voltage. Cell culturing on the nanofiber scaffold has been done for investigation the biocompatibility of P(LLA-CL) scaffold. Both smooth muscle cells (SMC) and endothelial cells (EC) adhered and proliferated well on the P(LLA-CL) nanofiber scaffolds. It was shown that the normal phenotypic shape of EC has been maintained on the nanofiber. The authors believed that the resulted structures have great potential in the application of tissue engineering.

As it mentioned above collagen is one of the most important materials in blood vessel ECMs. Matthews and coworkers [61] described how electrospinning can be adapted to produce tissue-engineering scaffolds composed of collagen nanofibers. They demonstrated that electrospun collagen promotes cell growth and the penetration of cells into the engineered matrix.

One of the major problems of blood vessel engineering is graft occlusion and failure. Luong-Van et al. [20] embedded heparin in poly (a-caprolactone) electrospun fibers to prevent the vascular smooth muscle cell (VSMC) proliferation that can lead to graft occlusion. In this study nanofibers with smooth surfaces and no bead defects were prepared from polymer solutions with 8%w/v PCL in 7:3 dichloromethane/methanol as solvent using electrospinning method. Form Fig3, it can be seen that fiber diameters decreased significantly with increasing heparin concentration. The effect of heparin concentration on the diameter and water contact angles of electrospun PCL fibers is also presented in table1.

[FIGURE 3 OMITTED]

The result showed that a sustained release of heparin from the fibers could be achieved over 14 days with controlled release diffusion during this time. From biocompatibility testes nanofibers did not induce an inflammatory response in macrophage cells in vitro. It suggested that electrospun PCL fibers are a promising candidate for delivery of heparin to the site of vascular grafts.

In another study a biological vascular substitute was fabricated from collagen, elastin and PLGA blend nanofibers using electrospinning (Fig4) [62]. From in vitro biocompatibility and mechanical testing data, the scaffolds possess tissue composition and mechanical properties similar to native vessels. Stitzel and coworkers concluded that the electrospun vessel matrix is biocompatible and does not elicit local or systemic toxic effects when implanted in vivo and it has potential application as a functional vascular graft for clinical use.

They enclosed that controlling the ratio of collagen, elastin and PLGA results in improved electrospinning characteristics and physical strength of the scaffolds, which resist bursting at nearly 12 times normal systolic pressure. The compliance characteristics of the vascular scaffolds were similar to native arteries.

[FIGURE 4 OMITTED]

[FIGURE 5 OMITTED]

Williamson et al. [63] developed a new scaffold suitable for small-diameter vascular grafts that advances strong attachment of endothelial cells. They produced the scaffold in two steps, first the luminal surface was formed by wet spinning polycaprolactone (PCL) fibers, and then the vessel wall substitute was prepared by electrospinning porous polyurethane (PU) onto the back of the PCL fibers. In conclusion as it is clear in Fig5, the luminal surface (Fig. 5a) exhibited orientated PCL fibres with an estimated gap between fibers ranging from 1 to 5 Mm. The antiluminal surface (Fig. 5b) showed a porous morphology with pore diameters of 10-30 Mm. gravity-spun PCL fibers have been combined with elastic electrospun PU fibers to produce a compatible PCL-PU composite scaffold. Cell culturing demonstrated strong attachment between human endothelial cells and the composite PCL-PU scaffold, and human endothelial cells proliferated to form a monolayer with strong PECAM-1 expression and cobblestone morphology. It was found that the composite scaffold may also deliver bioactive molecules. One of the test molecules was active trypsin, it had a defined 48 h pattern of release from luminal PCL fibers. The authors conclude that these data confirm the potential of this novel composite scaffold in vascular tissue engineering.

[FIGURE 6 OMITTED]

In another study, Inoguchi and coworkers [64] electrospun a tublar scaffold made of elastomeric poly(L-lactide-co-ecaprolactone) fabrics at different wall thicknesses to design a "mechano-active" small-diameter artificial vascular graft (Fig 6). They observed that the wall thickness of the fabricated tube (inner diameter; approximately 2.3-2.5mm and wall thickness; 50-340 [micro]m) increased proportionally with electrospinning (ELSP) time. They exposed the scaffold under static and dynamic flow conditions to determine the wall thickness dependence of mechanical responses including intraluminal pressure-induced inflation. Results under static condition showed that the smaller the wall thickness, the more compliant the tube. Under dynamic flow condition (1 Hz, maximal/minimal pressure of 90 mmHg/45 mmHg) produced by a custom-designed arterial circulatory system, strain, defined as the relative increase in diameter per pulse, increased with the decrease in wall thickness, which was close to that of a native artery. Therefore a tublar scaffold with mechano-active property was produced from an elastomeric PLCL and ELSP technique.

[FIGURE 7 OMITTED]

Buttafoco et al. [65] fabricated meshes of collagen and/or elastin successfully by means of electrospinning from aqueous solutions. To achieve continuous and homogeneous fibers, it was necessary to add PEO (Mw = 8 x [10.sup.6]) and NaCl. Electrospinning parameters and polymeric solution parameters determined the morphology of the obtained fibers. For example, increasing the elastin content determined an increase in fibres diameters from 220 to 600 nm. They stabilized the collagen/elastin (1:1) scaffolds by crosslinking with N-(3 dimethylaminopropyl)-N'-ethylcarbodiimide hydrochloride (EDC) and N-hydroxysuccinimide (NHS). For cell compatibility smooth muscle cells (SMCs) were successfully cultured on crosslinked scaffolds and grew as a confluent layer on top of the crosslinked meshes after 14 days of culture. It also was found that electrospinning solutions of the two proteins separately from each other can also give the possibility to produce multilayered scaffolds with controlled morphology and/or good mechanical properties.

Jeong et al. [66] prepared a hybrid scaffold from marine source collagen and PLGA using freez drying and electrospinning as a vascular graft. The hybrid scaffolds were fabricated by freeze drying collagen to form a porous collagen matrix and using electrospinning method to form a fibrous PLGA layer, this fibrous layer improved the mechanical strength of the collagen scaffolds in both the dry and wet states. The SEM images of the resulted scaffold are shown in Fig 7. The average pore size is about 150 [+ or -] 50 [micro]m. For biological study, cell cultures were done on hybrid scaffolds. It was observed that smooth muscle cells (SMCs)--and endothelial cells (ECs)--cultured collagen/PLGA scaffolds exhibited mechanical properties similar to collagen/PLGA scaffolds unseeded with cells, even after culturing for 23 days. There were four separate effects of a mechanical stimulation on the proliferation and phenotype of SMCs and ECs, cultured on collagen/PLGA scaffolds. 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. On the other hand, cells in tissues engineered in the static condition were randomly aligned.

[FIGURE 8 OMITTED]

Mechanical properties of the resulted scaffolds are shown in table II. 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.

This study also demonstrated that the co-culturing of SMCs and ECs on collagen/PLGA hybrid scaffolds under a pulsatile perfusion system induced the cellular alignment, the enhancement of vascular EC development, and the retention of differentiated cell phenotype.

In a recent study, Mo and coworkers [67] fabricated collagen-chitosan and P (LLA_CL) nanofibers by electrospinning method for tissue engineering application such as blood vessel grafts. Results of the experiments showed that the mechanical properties of the collagen-chitosan complex nanofibers varied with the collagen content in the complex. It was also found that the biodegradability of P (LLA-CL) nanofibers was faster than its membrane and that smooth muscle cells (SMC) grow faster on collagen nanofibers than on P (LLA-CL) nanofibers.

Zhang et al. [68] investigated the in vitro evaluation of electrospun silk fibroin scaffolds for vascular cell growth. In this study Human aortic endothelial (HAEC) and human coronary artery smooth muscle cell (HCASMC) were cultured on electrospun silk fibroin scaffolds. SEM and confocal images showed alignment and elongation of HCASMCs on random nonwoven nanofibrous silk scaffolds was observed within 5 days after seeding. Results demonstrated formation of ECM for the HCASMCs based on the quantification of collagen type I expression at protein and transcription levels. The authors believed that results indicate a favorable interaction between vascular cells and electrospun silk fibroin scaffolds and the electrospun silk fibroin scaffolds supported the growth and expansion of human aortic endothelial cells and human coronary artery smooth muscle cells based on cell proliferation, morphology and phenotype studies in vitro.

Finally, a recent work on vascular tissue engineering using electrospinning method was done by Tillman and coworkers [69]. They prepared electrospun polycaprolactone-collagen scaffolds and evaluated the in vivo stability of the resulted scaffolds in a rabbit aorta-iliac bypass model (Fig8). Results concluded 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. On the other hand it was observed that when the scaffolds implanted in vivo, they were able to retain their structural integrity over 1 month of implantation as demonstrated by serial ultrasonography. Authors suggested that electrospun scaffolds combined with vascular cells may become an alternative to prosthetic vascular grafts for vascular reconstruction.

There are also some other methods to obtained porous structure for vascular tissue engineering[70] such as fabricating under vacuum suction and lyophilization methods[71-73].

Bone Tissue Engineering

Natural bone is a biocomposite compose of inorganic (mainly hydroxyapatite crystals) and organic (mainly Type I collagen matrix) materials. To mimic the matrix, electrospinning is known as a promising technique due to its facile method for producing ultrafine and continuous sub-micron fibers and/or nanofibers.

Yoshimoto and coworkers [74] fabricated Microporous, non-woven poly (a-caprolactone) (PCL) scaffolds. For biological study, Mesenchymal stem cells (MSCs) derived from the bone marrow of neonatal rats were cultured, expanded and seeded on electrospun PCL scaffolds. Scanning electron microscopy (SEM), histological and immunohistochemical examinations were performed. Penetration of cells and abundant extracellular matrix were observed in the cell-polymer constructs after 1 week. SEM results demonstrated that the surfaces of the cell-polymer constructs were covered with cell multilayers at 4 weeks. In addition, mineralization and type I collagen were observed at 4 weeks. This suggests that electrospun PCL is a potential candidate scaffold for bone tissue engineering.

Ito et. al [75] used biodegradable and a biocompatible poly (3-hydroxybutyrate-co-3-hydroxyvalerate) (PHBV) copolymer from a microbial polyester and electrospun it to a nanofibrous web and composited with hydroxyapatite (HA) by soaking in simulated body fluid. In this work, nanofibrous film was compared with cast film and showed it was hydrophobic, although after HA depositions, both of the surfaces were extremely hydrophilic. Degradation rate of HA/PHBV nanofibrous films was investigated, it was found that in the presence of polyhydroxybutyrate depolymerase, the degradation rate was very fast. The surface of the nanofibrous film showed enhanced cell adhesion over that of the flat film, although cell adhesion was not significantly affected by the combination with HA.

[FIGURE 9 OMITTED]

In another study, electrospun silk fibroin-based fibers with average diameter 700 [+ or -] 50nm were prepared from aqueous regenerated silkworm silk solutions. Human bone marrow stromal cells (BMSCs) were cultured on these silk matrices and adhesion, spreading and proliferation were investigated. Scanning electron microscopy (SEM) and MTT analyses demonstrated that the electrospun silk matrices supported BMSC attachment and proliferation over 14 days in culture similar to native silk fibroin (B15 mm fiber diameter) matrices. The authors concluded that ability of electrospun silk matrices to support BMSC attachment, spreading and growth in vitro, combined with a biocompatibility and biodegradable properties of the silk protein matrix, suggest potential use of these biomaterial matrices as scaffolds for tissue engineering [76].

Kim and coworkers [77] electrospun silk fibroin (SF) and applied it as a device in bone and periodontal regenerative therapy because of its favorable biological properties. This study specially was done to evaluate the biocompatibility of the SF nanofiber membrane, and to examine its effect on bone regeneration in a rabbit calvarial model (Fig 9). Cell proliferation, morphology, and differentiation were investigated to examine the biocompatibility of the electrospun SF membrane.

Results showed that the cell numbers and osteocalcin production labels were significantly increased in accordance with culture period. Cells had a star-like shape and broad cytoplasmic extensions on the membrane. In in vivo tests, a complete bony union across the defects was observed after 8 weeks. At 12 weeks, the defect had completely healed with new bone. The authors believed that these results strongly suggest that the SF membrane should be useful as a tool for guided bone regeneration.

Ngiam et. al [78] prepared Poly-L-lactic acid (PLLA) and PLLA/collagen (50% PLLA+50% collagen; PLLA/Col) nanofibers using electrospinning. They mineralized these nanofibers using a modified alternating soaking method. X-ray diffraction (XRD), Fourier transform infrared spectroscopy (FTIR), scanning electron microscopy (SEM), and contact angle measurements were done to characterize the structural properties and morphologies of mineralized PLLA and PLLA/ Col nanofibers. Fig 10 shows the SEM micrograph of resulted nanofibrous mats.

[FIGURE 10 OMITTED]

To assess the biological properties of the nanofibrous composites, human bone-derived osteoblasts were cultured on the materials for up to 1 week. Results showed that, the bonelike nanohydroxyapatite (n-HA) was successfully deposited on the PLLA and PLLA/Col nanofibers. They observed that the formation of n-HA on PLLA/Col nanofibers was faster and significantly more uniform than on pure PLLA nanofibers. From cell attachment studies, n-HA deposition enhanced the cell capture efficacy at the 20-minute time point for PLLA nanofibers. Based on these observations, the authors demonstrated that n-HA deposition on nanofibers is a promising strategy for early cell capture [78].

Zhang and coworkers [79] have presented a two-step approach that combines an in situ co-precipitation synthesis route with electrospinning process to prepare a novel type of biomimetic nanocomposite nanofibers of hydroxyaptite/chitosan (HA/CS). A model HA/CS nanocomposite with the HA mass ratio of 30 wt% was synthesized through the co-precipitation method so as to attain homogenous dispersion of the spindle-shaped HA nanoparticles (ca. 100 x 30 nm) within the chitosan matrix. Continuous HA/CS nanofibers with a diameters of 214_25 nm had been produced successfully by using a small amount (10 wt %) of ultrahigh molecular weight poly(ethylene oxide) (UHMWPEO) as a fiber-forming facilitating additive. Further SAED and XRD analysis confirmed that the crystalline nature of HA remains and had survived the acetic acid-dominant solvent system. In vitro cell culture with human fetal osteoblast (hFOB) cells for up to 15 days was done to investigate biological properties of resulted scaffolds. They found that the incorporation of HA nanoparticles into chitosan nanofibrous scaffolds led to significant bone formation oriented outcomes compared to that of the pure electrospun CS scaffolds. The results obtained highlight the great potential of using the HA/CS nanocomposite nanofibers for bone tissue engineering applications.

[FIGURE 11 OMITTED]

[FIGURE 12 OMITTED]

Heart tissue engineering

Zong et.al [80] used electrospinning method to fabricate biodegradable non-woven poly (lactide)- and poly(glycolide)-based (PLGA) scaffolds for cardiac tissue engineering applications. In this study, the structural and functional effects of fine-textured matrices with sub-micron features were examined on the growth of cardiac myocytes. Post-processing was applied to achieve macro-scale fiber orientation (anisotropy). A dose-response effect of the poly (glycolide) concentration on the degradation rate and the pH value changes were confirmed through in vitro studies. After producing the scaffolds, cardiomyocytes (CMs) were cultured on the resulted webs to form tissue-like constructs. The nanofibrous structures of the non-woven matrix allowed the cardiomyocytes to make extensive use of provided external cues for isotropic or anisotropic growth, and to some extent to crawl inside and pull on fibers, these results were obtained from SEM studies. Structural analysis by confocal microscopy indicated that cardiomyocytes had a preference for relatively hydrophobic surfaces.

CMs on electrospun poly (L-lactide) (PLLA) scaffolds developed mature contractile machinery (sarcomeres). Functionality (excitability) of the engineered constructs was confirmed by optical imaging of electrical activity using voltage-sensitive dyes. They conclude that engineered cardiac tissue structure and function can be modulated by the chemistry and geometry of the provided nano- and microtextured surfaces. Electrospinning is a versatile manufacturing technique for designing biomaterials with potentially reorganizable architecture for cell and tissue growth.

In another work, Ishi et al. [81] prepared cardiomyocytes-electrospun nanofibrous poly (a-caprolactone) structures through seeding neonatal rat cardiomyocytes on biodegradable electrospun nanofibrous poly (a-caprolactone) webs. Between days 5 and 7 after culturing, the webs were overlaid to construct 3-dimensional cardiac grafts. On day 14 of in vitro culture, the engineered cardiac grafts were analyzed by means of histology, immunohistochemistry, and scanning electron microscopy. Results showed that the cultured cardiomyocytes attached well to the webs. From SEM images, it was found that the average fiber diameter of the scaffold was about 250 nm, well below the size of an individual cardiomyocyte. Constructs with up to 5 layers could be formed without any incidence of core ischemia. The individual layers adhered intimately. Synchronized beating was also observed. Morphologic and electrical communications between the layers were established, as verified by means of histology and immunohistochemistry. This report demonstrates the formation of thick cardiac grafts in vitro and the versatility of biodegradable electrospun meshes for cardiac tissue engineering. Cross-sectional view of a single graft, 5-layer graft are shown in Fig 11.

Articular Cartilage tissue engineering

Articular or hyaline cartilage tissue is comprised of at least four distinct zones, each with a specific cell and extracellular matrix (ECM) organization or orientation. The four zones known as superficial/tangential, intermediate/transitional, deep/radial, and calcified zone naturally function collectively to provide the low-friction, wear-resistant, load-bearing tissue on the end of bones in synovial joints [82, 83]. Cartilage defects are a major health problem. Tissue engineering has developed different strategies and several biomaterial morphologies, including natural-based ones, for repairing these defects [84]. Several attempts have been made to regulate the important cellular behaviors (Cell differentiation, adhesion, and orientation), but among all, electrospinning that can produce patterned extracellular matrix (ECM) features was one of the most efficient one.

Wise and coworkers [82] created electrospun oriented polycaprolactone (PCL) scaffolds (500 or 3000nm fiber diameter). To achieve a similar structure to articular cartilage, human mesenchymal stem cells (hMSCs) were cultured on oriented nano and microfibrous electrospun PCL scaffolds as well as a randomporous PCL film. For biological investigations cell viability, morphology, and orientation on the fibrous scaffolds were determined as a function of time. The findings suggest that engineering an oriented ECM environment to regulate tissue alignment could be optimized by oriented electrospun nanofibers. Creating the superficial zone of articular cartilage, may be significantly improved by a combination of stem cells and nanofibrous scaffolds. Fig 12 explores the apparent cell adhesion mechanism(s) between hMSCs and nanofibers that could potentially play an important role in stem cell differentiation.

In a recent study, electrospun polycaprolactone (PCL) and starch-compounded PCL (SPCL) nanofiber webs were used to evaluate extracellular matrix (ECM) formation by bovine articular chondrocytes (BACs)[84]. Alves Da Silva and coworkers [84] have done this to investigate the suitability of PCL and SPCL nanofiber webs in chondrocyte cultures, and their capability to produce ECM when seeded onto these nanofibrous materials. They also assessed the effect of culture conditions (static vs. dynamic) on ECM formation. Cell-scaffold constructs were characterized using scanning electron microscopy, histology, immunolocalization of collagen types I and II, and glycosaminoglycan (GAG) quantification. An extensive cell colonization of the entire nanofiber web, for both materials was found. Some degree of cell infiltration inside the nanofiber webs was noticeable for both materials. ECM formation and GAG were detected throughout the materials, evidencing typical construct maturation. PCL and SPCL nanofiber webs are suitable as supports for ECM formation and therefore are adequate for cartilage tissue-engineering approaches [84].

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A. Gholipour Kanani, S. Hajir Bahrami *

Department of Textile Engineering Amirkabir University of Technology Tehran, Iran, 15875-4413

* Corresponding author: haiirb@aut.ac.ir (S. H. Bahrami)

Received 30 November 2009, Accepted 27 April 2010, Published online 15 August 2010.
Table 1: Effect of heparin concentration on the
diameter and water contact angles of electrospun
PCL fibers [20]

Heparin        Loading          Fiber diameter
(%w/v)      ([micro]g/mg)            (nm)

0                 --           810 [+ or -] 150
0.05      0.51 [+ or -] 0.06   680 [+ or -] 150
0.5        5.8 [+ or -] 0.6    550 [+ or -] 110

Heparin         A                 R
(%w/v)

0         139 [+ or -] 6   113 [+ or -] 5
0.05      132 [+ or -] 7   105 [+ or -] 12
0.5       136 [+ or -] 3   108 [+ or -] 6

Table 2: Mechanical properties of tubular porous scaffolds[66]

Mechanical         PLGA scaffold (b)    Collagen scaffolds
properties (a)

Tensile
  strength (MPa)
Dry                0.25 [+ or -] 0.01   0.12 [+ or -] 0.01
Wet                0.24 [+ or -] 0.08   0.06 [+ or -] 0.01

Elongation at
  break (%)
Dry                20.0 [+ or -] 1.02   12.0 [+ or -] 0.65
Wet                20.0 [+ or -] 0.61   34.0 [+ or -] 1.40

Mechanical         Collagen/PLGA hybrid    Tissue-engineered
properties (a)          scaffolds            collagen/PLGAc
                                          hybrid scaffolds (c)

Tensile
  strength (MPa)
Dry                 0.49 [+ or -] 0.04     0.49 [+ or -] 0.02
Wet                 0.14 [+ or -] 0.02     0.14 [+ or -] 0.05

Elongation at
  break (%)
Dry                 17.0 [+ or -] 0.24     17.0 [+ or -] 0.72
Wet                 41.0 [+ or -] 1.89     41.0 [+ or -] 2.01

(a) The mechanical properties of the tubular porous scaffolds
were measured by a universal testing machine (UTM, INSTRON No.
4465). Three specimens were tested for each sample.

(b) Tublar porous PLGA scaffolds were prepared by a similar
method to collagen scaffolds using a tubular scaffolds mold.

(c) SMC and EC-cultured collagen/PLGA hybrid scaffolds.
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