Bioresorbable elastomeric vascular tissue engineering scaffolds via melt spinning and electrospinning
Introduction
One of the more severe forms of heart disease is associated with atherosclerosis, which affects more than 8 million Americans, leading to a significant process that causes narrowing of the arteries [1], [2]. Recent statistics show that peripheral arterial disease increases morbidity and mortality [3]. Surgical replacement of vessel segments or bypass surgery is the most common intervention for coronary and peripheral atherosclerotic disease, with at least 550,000 bypass cases performed per year [4]. Synthetic vascular prostheses made of poly(ethylene terephthalate) or expanded polytetrafluoroethylene have been effective for large diameter grafts. However, in spite of many years of research using a wide variety of biomaterials, clinical success for small diameter (<6 mm) vessels has yet to be demonstrated due to complications such as occlusion, thrombosis and intimal hyperplasia [1], [2], [4], [5], [6], [7].
Several attempts have been made to construct a blood vessel replacement with biological functionality. The use of protein coatings and the seeding of endothelial cells on the lumenal surface have shown some promise to increase biocompatibility, but permanent synthetic materials are unlikely to be fully accepted, given the chronic inflammatory response they provoke and the increased risk of infection. On the other hand, implantation of native vessels is limited by the mismatch of dimensional and mechanical properties [2], [6]. Given the limitations of these current techniques, the desirability of a biocompatible engineered blood vessel is pushing research into the development of tissue-engineered small-diameter blood vessels.
For tissue engineering, the design of the scaffold plays a significant role since the matrix provides the cells with a tissue-specific environment and architecture. In particular, designing a scaffold structure that enables cells to proliferate and grow into a three-dimensional (3-D) tissue while allowing the necessary amount of oxygen and nutrient diffusion is important. The key factors in creating a 3-D scaffold include tailoring the degradation rate to meet the requirements of new tissue growth, providing interconnected pores, generating high porosity so as to promote cell–cell and cell–matrix communication, and having sufficient mechanical stability [8], [9], [10], [11], [12]. Developing scaffolds that can maintain their mechanical integrity while exposing cells to long-term cyclic mechanical strains is especially necessary in cardiovascular applications when engineering smooth muscle cellular constructs [13], [14], [15], [16]. To achieve this, scaffolds should be elastic enough to withstand cyclic mechanical strains without any significant permanent deformation or creep. So, in the selection of a scaffold material for the fabrication of engineered tissues, a candidate polymer should possess appropriate mechanical properties with a low elastic modulus or high compliance, which are suitable for the target applications. In addition, its degradation products during implantation should be nontoxic. Over the past years, hydrolyzable and biocompatible copolymers of ε-caprolactone and l-lactide have been of great interest for medical applications [13], [17]. Polylactide (PLA) is a crystallizable hard and brittle material, whereas poly(ε-caprolactone) is a semi-crystalline material with rubbery properties. Copolymers of l-lactide and ε-caprolactone (poly(l-lactide-co-ε-caprolactone), PLCL) exhibit a range of mechanical properties from rigid solids to elastomers, depending on their composition [18], [19]. There have been previous studies using PLCL copolymers to fabricate scaffolds by various methods, including extrusion, particulate leaching [13], [14] and electrospinning [19], [20], [21]. In particular, a 50:50 ratio of l-lactide and ε-caprolactone monomers has proven to have high elastomeric properties with breaking strains in excess of 100% [13], [14], [18], [19]. Compared to other bioresorbable polymers such as polyglycolide and PLA, this PLCL copolymer has a slow rate of degradation, with in vivo studies reporting 81% mass retained after 15 weeks of implantation [22].
The objective of this study was to design a porous biodegradable multilayered tubular construct that mimics the structural configuration and mechanical properties of a native vessel. To achieve this goal, the highly elastomeric 50:50 PLCL copolymer was selected to produce a biconstituent tubular scaffold with distinct layers with different geometries. Both melt spinning and electrospinning techniques were chosen so as to create separate layers with different fiber diameters (macrofibers vs. submicron fibers) and different fiber alignments within the scaffold’s architecture. Recently, fibrous structures have gained considerable interest as tissue engineering scaffolds due to their high surface-to-volume ratios, interconnected pores and high total porosity, while giving versatile material selection and better control over pore size distribution, as well as consistent and flexible processing. Melt spinning and winding were undertaken to achieve larger-diameter fibers and to control the angle between them so as to modulate the pore size. Scaffolds were electrospun from two different solvents, acetone and 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP), and were evaluated in terms of their mechanical performance, morphology, ease of processing, cytotoxicity and cell viability. The scaffolds were designed to have values for total porosity in excess of 75%, an interconnected network of pores, and values for mechanical strength and elongation exceeding those of natural arteries. The transverse tensile strength and strain as well as the initial tensile modulus were compared between the different types of scaffolds. Finally, based on the characterization of the individual layers of melt-spun and electrospun structures, a biconstituent structural concept was attempted by combining concentric layers of melt-spun macrofibers and electrospun submicron fibers within the same tubular scaffold without the need for an adhesive.
Section snippets
Materials
PLCL copolymer was received as a solid bulk polymer from the Biomaterials Research Center at the Korea Institute of Science and Technology. The polymer had been synthesized following the procedure described previously [23]. Briefly, l-lactide (100 mmol), ε-caprolactone (100 mmol) and 1,6-hexanediol (0.5 mmol) were polymerized at 150 °C for 24 h, using stannous octoate (1 mmol) as the catalyst. After reaction, the product was dissolved in chloroform and precipitated in methanol, filtered and dried
Morphological properties
The melt-spun monofilaments were wound up directly after extrusion on a rotating Teflon-coated mandrel and were removed as a tube after collection once they had cooled to room temperature (Fig. 2). The inner diameter of the tubular scaffolds was 5 mm. The tubes were then cut up in 5 mm lengths for SEM and mechanical testing. When removed from the Teflon-coated mandrel, the tubular construct maintained its shape and integrity because the translucent monofilaments had bonded well to each other.
Conclusions
A 50:50 PLCL copolymer was successfully melt-spun and electrospun to form individual and combined porous tubular scaffolds. This is the first time that this particular elastomeric biodegradable copolymer has been melt-spun. Using two alternative solvent systems – acetone and HFIP – for electrospinning provided different results in terms of fiber dimensions, mechanical properties and cytotoxicity, but the HFIP solvent was preferred as it gave a more stable threadline. The mechanical properties
Acknowledgements
The authors acknowledge the financial support from the Boston Scientific Corporation and the College of Textiles at North Carolina State University. We also thank Dr. Bhupender Gupta, Dr. Ajit Moghe and Hai Bui for their insights and technical assistance.
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