Review For reprint orders, please contact: [email protected]
Therapeutic application of electrospun nanofibrous meshes
Fabricating tissue architecture-mimicking scaffolds is one of the major challenges in the field of tissue engineering. Electrospun nanofibers have been considered as potent techniques for fabricating fibrous scaffolds biomimicking extracellular frameworks. Therapeutic agent-incorporated nanofibrous meshes have widely served as excellent substrates for adhesion, proliferation and differentiation. Many drugs, proteins and nucleic acids were incorporated into the scaffolds for regeneration of skin, musculoskeletal, neural and vascular tissue engineering in aims to control the release of the therapeutic agents. In the current article, we focus on introducing various fabrication techniques for electrospun nanofiber-based scaffolds and subsequent functionalization of nanofibers for therapeutic purposes. We also detail how the therapeutic nanofibrous meshes can be employed in the field of tissue engineering.
Hye Sung Kim1 & Hyuk Sang Yoo*,1,2 1 $EPARTMENTOF"IOMEDICAL-ATERIALS %NGINEERING 3CHOOLOF"IOSCIENCE "IOENGINEERING +ANGWON.ATIONAL 5NIVERSITY #HUNCHEON 2EPUBLIC OF+OREA )NSTITUTEOF"IOSCIENCE"IOTECHNOLOGY +ANGWON.ATIONAL5NIVERSITY 2EPUBLIC OF+OREA
!UTHORFORCORRESPONDENCE 4EL &AX HSYOO KANGWONACKR
Keywords: DRUGDELIVERYsELECTROSPINNINGsNANOlBERsSCAFFOLDsTISSUEENGINEERING
Strategies of regenerative medicine are focused on the restoration of tissue architectures and biologic function in host-tissue lesions by implantation of supportive scaffolds containing biomolecules or cells [1–3] . Recently, considerable interest has been given to biomimetic and biologically active scaffolds that can replace the natural extracellular matrix (ECM) and provide a favorable environment to host cells to repopulate and resynthesize a new natural matrix. Among those scaffolds, electrospun nanofibrous meshes (NFs) have been used for tissue engineering because of their unique features that mimic the structure of fibrillar collagen (COL) and elastin (ELAS) in the natural ECM [4–6] . In addition, the high porosity of NFs facilitates the transport of gases, nutrients and regulatory factors for the survival and proliferation of cells [7] . Moreover, a high surface area-to-volume ratio of NFs not only provides wider substrate dimension for cell attachment, but also offers better incorporation of biomolecules on their surfaces compared with nonfibrous scaffolds [1–3] . Thus, various electrospun nanofibers have been
10.2217/NMM.13.224 © 2014 Future Medicine Ltd
biofunctionalized for incorporating bioactive molecules by pre- and/or post-electrospinning processes [4,5] . Here, we describe the fabrication methods for electrospun nanofiber-based scaffolds, an overview of the methods for functionalization of fibrous scaffolds and application of NFs for tissue engineering. Principle of electrospinning Electrospinning process
Electrospinning is a unique approach using electrostatic forces to produce fine fibers from polymer solutions [8] . The fabricated fibers feature a thin diameter in the range of a few nanometers to submicrometers, as well as a large surface area-to-volume ratio. Thus, the electrospun NFs show high porosity and interconnected fibrous networks that mimic ECM architectures. An electrospinning system largely comprises a high-voltage power supply, a spinneret (e.g., single/dual needle and pipette tip), and a grounded collector (usually a metal plate or rotating mandrel). The completely dissolved polymer solution is injected into the needle.
Nanomedicine (2014) 9(4), 517–533
part of
ISSN 1743-5889
517
Review
Kim & Yoo
A polymer droplet at the tip of a spinneret is electrically charged with high electrical potential. If the repulsive electrical forces of the droplet overcome surface tension forces, a charged jet of the polymer solution is ejected from the tip of the Taylor cone. The polymeric fibers deposit on the grounded collector with a rapid whipping motion and solvent evaporation [9] . In comparison with other methods for producing thin fibers, such as phase separation, drawing and self-assembly, electrospinning is advantageous in terms of economics and technical viewpoints for manufacturing nanofiber-based scaffolds for tissue engineering. Polymers for electrospinning Natural polymers
Various polymers have been used in electrospinning for tissue-engineered scaffolds. Typical natural polymers for fabrication of electrospun NFs include COL, gelatin, chitosan, chitin, ELAS, casein, cellulose acetate, silk protein and fibrinogen [10] . In general, natural polymers exhibit better biocompatibility and low immunogenicity compared with synthetic polymers. COL has been used in various tissue-engineering applications because it is the most abundant protein in the mammalian ECM, is relatively nonimmunogenic and a good sequester of many factors for the maintenance and regeneration of tissue. Most studies have used 1,1,1,3,3,3-hexafluoroisopropanol or 2,2,2,-trifluoroethanol as the solvent to electrospin COL fibers because of its high volatility. Unfortunately, those solvents not only disrupt the native structure of COL, but also decrease the denaturation temperature, which leads to 45% of COL being lost during electrospinning [11,12] . Thus, several researchers have employed 40–50% acetic acid, and they confirmed that diluted acetic acid was better for protecting the native helical structure of COL than 1,1,1,3,3,3-hexafluoroisopropanol [13,14] . However, denaturation of the secondary structure leads to rapid degradation of electrospun COL fibers in aqueous media. Thus, crosslinking reagents such as glutaraldehyde [15,16] , 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride [17] , N-hydroxysuccinimide [18] , and genipin [19] have been used to stabilize the morphology of electrospun COL fibers. However, glutaraldehyde produces toxic side products and subsequently calcifies COL substrates. With respect to genipin crosslinking, it requires long processing times (>48 h) [19,20] . Although these methods could delay the solubilization of COL fibers, fiber fusion or disappearance of fibrous morphology occurs during crosslinking. Chew et al. introduced photochemical crosslinking by mixing COL polymers with 0.1% (weight/volume) of rose bengal as a photo-initiator before electrospin-
518
Nanomedicine (2014) 9(4)
ning, and electrospun COL fibers were irradiated with an argon laser for crosslinking [13] . They found that fibrous morphology was well maintained after photochemical crosslinking, and that the low concentration of photoinitiator could minimize cytotoxicity. Chitosan is one of the most abundant polysaccharides and is derived from chitin. It has been employed for wound-dressing material because of its biocompatibility, biodegradability and bioactivity. In addition, chitosan promotes hemostasis and aids tissue regeneration [21,22] . However, electrospinning of chitosan is a challenge because of its limited solubility and high viscosity. Chitosan is soluble in organic acids such as dilute aqueous acetic acid, formic acid and lactic acid. However, it is partially soluble in inorganic acids and insoluble in water. Although chitosan is dissolved in the solvent, it is positively charged because of free amino groups in the chitosan structure, which make chitosan solutions highly viscous [23,24] . The cationic charge of the solution inhibits the chain entanglement, and the strong hydrogen bonds in chitosan solution prevent the movement of polymeric chains exposed to the electrical field, which stop fiber formation [25–27] . Thus, to improve spinnability, chitosan-based nanofibers have been prepared by blending chitosan with second polymers such as gelatin, silk fibroin, poly(ethylene oxide), and poly(vinyl alcohol) (PVA) [28,29] . Hyaluronic acid (HA), a linear polysaccharide composed of repeating glucuronic acid and N-acetylglucosamine, is the main component of the ECM in connective tissues. HA performs important roles as a molecular filter, shock absorber and support COL fibril structure. In addition, it has been used extensively in the biomedical field due to its excellent biocompatibility and biodegradability. It is difficult to create uniform size fibers from hyaluronan solution using electrospinning due to the high viscosity, surface tension, and strong water retention ability of the hyaluronan solution. Thus, Li et al. and Um et al. performed electrospinning in combination with air flow at 57°C with a 70 ft/h flow rate and successfully produced HA nanofibers [30,31] . Keratins are the major structural fibrous protein of hair, feathers, wool and nails, which are characteristically abundant in cysteine residues. Wool keratins have been demonstrated to be a favorable biomaterial of scaffolds for fibroblasts and osteoblasts due to their cell adhesion sequences, arginine-glycine-aspartic acid and leucine–aspartic acid–valine, biocompatibility and biodegradability. Unfortunately, keratin or keratin-based materials are difficult to electrospin due to their low molecular weight (9–60 kDa); thus, they are alternatively blended with synthetic polymers such as poly(lactic-co-glycolide) (PLGA), poly-l-lactide
future science group
Therapeutic application of electrospun nanofibrous meshes
(PLLA) and poly(3-hydroxybutyrate-co-3-hydroxyvalerate) to improve the poor mechanical properties of the electrospun scaffolds [32–34] . Synthetic polymers
Synthetic polymers offer several advantages over natural polymers. They can be tailored to give the requisite mechanical properties, thermal stability and desired degradation profile. In addition, the range of solvents available to dissolve synthetic polymers for electrospinning is relatively wider than that for natural polymers. Typical organic solvents used in electrospinning are chloroform, ethanol, dimethylformamide, dichloromethane and their mixtures [35] . The selection of the solvent for preparing polymer solutions is important because the morphology and size of electrospun nanofibers are strongly dependent upon solution properties such as viscosity and surface tension, which are attributed to polymer concentrations and solvent properties (including polarity, solubility and volatility) [36,37] . Biocompatible and biodegradable polymers such as polyglycolide, polylactide (PLA), poly(H-caprolactone) (PCL) and PLGA have been electrospun into nanofibrous scaffolds for biomedical applications. However, electrospun NFs composed of biodegradable polymers such as polyesters or poly(D-hydroxy acids) are mechanically unstable and cannot withstand implantation because of the rapid degradation rate of the mesh [38,39] . The highly porous structure of the mesh facilitates degradation, and the degradation rate of the mesh is much faster than the recovery rate of the tissue. Hence, these features limit the use of meshes for tissue engineering. Nonbiodegradable polymers such as polyurethane and polyesterurethane possess substantial mechanical stability. However, they might interfere with the turnover and remodeling of tissue owing to their slow degradation [40,41] . Thus, the mechanical stability and degradation rate of nanofiber-based scaffolds can be controlled by mixing various polymers exhibiting different degradation profiles [42] . In addition, PLGA (a copolymer composed of lactic acid and glycolic acid) shows different degradation profiles according to the ratio of lactic acid and glycolic acid [43] . The higher the content of glycolide units, the lower the time required for degradation. Shin et al. reported that the mechanical properties of nanofiber scaffolds prepared with PLGA composed of 75–85% glycolic acid are likely to be similar to those of cartilage [44] . In addition, these PLGA nanofiber scaffolds degraded dramatically after 7 weeks of incubation under physiological conditions. These findings suggested that nanofiber-based PLGA scaffolds could support cell proliferation and tissue development for t2 months.
future science group
Review
For therapeutic application of electrospun nanofibers, various copolymers have been synthesized and electrospun to modulate cell affinity, as well as drug encapsulation/release of nanofibers. Electrospun nanofibers composed of biodegradable hydrophobic polyesters generally have good mechanical properties, while cells show poor adhesion the polymeric nanofiber. Thus, a proper hydrophilic polymer segment is widely conjugated to the hydrophobic polymers to enhance the biocompatibility of the electrospun nanofibers by increasing hydrophilicity. Interestingly, cell adhesion can be inhibited by increasing the proportion of hydrophilic segment in copolymers [45] . Poly(l-lactide-co-glycolide) (PLLGA), PLLGA blended with PLLGA-poly(ethylene glycol) (PEG) copolymer or pure PEG(PLLGA/PEG) were electrospun onto nonwoven mesh, respectively. Upon the addition of PLLGA-PEG or pure PEG, the hydrophilicity of the electrospun NF significantly improved. In addition, PLLGA/PLLGA-PEG significantly inhibited the adsorption of proteins and proliferation of cells, unlike PLLGA/PEG nanofibers. By introducing stimuli-responsive moieties in copolymers, drug encapsulation/release of electrospun nanofibers can be controlled under specific environments, such as temperature, pH, electric field, magnetic field and light. Poly(N-isopropylacrylamide) (PNIPAAm) has been extensively studied for a thermosensitive drug delivery system. It is soluble in water below it has a lower critical solution temperature (32°C) and is gelated at higher temperature. Thus, electorspun nanofibers composed of the NIPAAm homopolymers are not stable in water and disperse easily; therefore, copolymerization with crosslinkable moieties is required to obtain stable nanofibers in an aqueous medium [46,47] . PNIPAAm-co-N-hydroxymethylacrylamide (HMAAm) nanofibers showed an ‘on–off ’ switchable release of dextran according to a heating–cooling cycle. Upon thermal curing of nanofibers, the methylol groups of HMAAm in the copolymers were crosslinked by self-condensation. The crosslinked nanofibers preserved their morphology and reversible volume changes in aqueous media in response to cycles of temperature alternation, while noncrosslinked nanofibers collapsed and dispersed quickly in the aqueous solution. In addition, dextran release from the crosslinked nanofiber accelerated during the heating cycle, whereas a negligible amount of dextran was evolved during the cooling process. Strategies to biofunctionalize electrospun nanofibers To apply electrospun nanofibers to biomedical uses, bioactive factors such as drugs, growth factors, nucleic
www.futuremedicine.com
519
Review
Kim & Yoo
acids, enzymes and antibiotics have been incorporated into fibrous scaffolds via various methods (Figure 1) . Blend electrospinning
A simple approach for drug incorporation into nanofibrous scaffolds is the blending of polymer solutions with drugs in pre-electrospinning. In single-nozzle electrospinning, hydrophobic drugs that are soluble in common organic solvents can be incorporated into nanofibers, whereas water-soluble bioactive agents such as proteins and nucleic acids, are hard to be blended in
homogeneously, which leads to an initial burst release within 24 h [48–50] . Thus, in order to improve protein incorporation and release profile, hydrophilic additives, such as PEG, gelatin and hydroxyapatite (HAp), have been loaded together with protein during blend electrospinning [51–53] . The addition of hydrophilic polymers or amphiphilic copolymers could enhance protein miscibility within nanofibers, as well as the hydrophilicity and water uptake of the nanofibers, which accelerates protein release from electrospun nanofibers. Unfortunately, the denaturation of bioactive molecules by
• Blend electrospinning
Electrospinning of drug/polymer blend through single nozzle
• Coaxial electrospinning
Dual-nozzle electrospinning or drug/polymer core and polymer shell
• Postelectrospinning: surface modification
Surface functionalization
Physical/chemical drug incorporation
Electrospun nanofibers Figure 1. Various techniques of fabricating drug-incorporated electrospun nanofibers; blend electrospinning, coaxial electrospinning and surface modification.
520
Nanomedicine (2014) 9(4)
future science group
Therapeutic application of electrospun nanofibrous meshes
exposure to organic solvents and electric fields during electrospinning is problematic. Thus, bioactive molecules are, in general, emulsified or encapsulated into nanoparticles to improve dispersion within the nanofibers while maintaining bioactivity [54,55] .
Review
incorporation of various types of nucleic acids (including plasmid DNA and siRNA), but also liberated the poly(ethyleneimine)–DNA complex from the nanofibers, which was strictly controlled by digestion of the linker under an abundant MMP-2 environment such as diabetic ulcers [62,63] .
Coaxial electrospinning
Alternatively, coaxial electrospinning can be carried out through a dual nozzle (a spinneret composed of two coaxial capillaries) for electrospinning two separate immiscible polymer solutions into coreshell-structured nanofibers simultaneously. The aqueous-based biological molecules can be incorporated separately into the core of nanofibers, thereby minimizing interactions with organic solvents during electrospinning [56–58] . Therefore, coaxial electrospinning is suitable for the preparation of NFs for delivery of unstable bioactive molecules such as growth factors. It has been shown that the growth factors released from coaxial nanofibers maintain the same level of bioactivity as ‘fresh’ growth factors [57] . Choi et al. coelectrospun a PCL-PEG block copolymer with a basic FGF (bFGF) solution using a dual nozzle to incorporate bFGF into the core of nanofibers, and then EGF was conjugated chemically on the shell of coaxial nanofibers [58] . The bFGF/EGF nanofibers showed binary release profiles of each growth factor; the core-encapsulated bFGF showed a high initial burst in 24 h, whereas the EGF was released in 7 days. Eventually, the bFGF/EGF NFs could promote wound recovery in diabetic ulcers because different release profiles of growth factors offered precise control of cell proliferation during re-epithelialization. Surface modification
Surface modification of nanofibers with biomolecules is another promising approach for introducing biofunctionality into nanofibers. Biomolecules are absorbed physically on NFs via electrostatic interactions, hydrogen bonding, hydrophobic interactions and van der Waals interactions [59] . However, modulating the degree of drug loading and release kinetics is difficult. For a slow and prolonged release of a therapeutic agent, a target biomolecule is immobilized chemically on the surface of nanofibers [60,61] . However, in the case of gene delivery, which requires endocytosis of a therapeutic gene, covalent conjugation of nucleic acids onto the scaffolds does not lead to a therapeutic effect because of the limited release of the gene from the scaffolds. Thus, the rate of release of immobilized target molecules can be controlled precisely via introduction of a responsive moiety to local external cues. For example, poly(ethyleneimine)immobilized nanofibers via a matrix metalloproteinase-cleavable linker not only showed high efficiency of
future science group
Biomedical application of the electrospun nanofibers-based scaffold Tissue regeneration Skin
Electrospun NFs have received much attention as potential candidates for wound-dressing materials. As a result of their high surface-to-volume ratio and nanoporous structure, NFs can absorb wound exudate effectively and keep the wound site under a moisturerich environment during healing. In addition, oxygen can penetrate easily through the mesh, while bacterial infiltration is strictly excluded because of the nanoscaled pores of the NF [64,65] . To enhance the efficiency of skin regeneration, NFs are usually administrated in combination with various drugs: growth factors, antibiotics and hemostatic factors. In diabetic ulcers, EGF-immobilized PCL/PCL-PEG NFs had a faster rate of wound closure than EGF solutions or NFs without EGF [60] . The self-regeneration of skin after minor epidermal injury is remarkable. However, if the injury is severe (e.g., massive, deep loss of the epidermis and dermis in full-thickness skin wounds), the lost skin cannot regenerate spontaneously [66] . Thus, electrospun NFs as skin substitutes have been developed with supporting cells such as keratinocytes, dermal fibroblasts and mesenchymal stem cells (MSCs). COL/PLGA-blended NFs have been modified by conjugation with a CD29 antibody for rapid attachment of bone marrow (BM)derived MSCs [67] . Implanted BM-MSCs attached to COL/PLGA NFs promoted ingrowth of epithelial edges and COL synthesis. In addition, implanted BM-MSCs expressed keratin 10 and filaggrin, which suggested that BM-MSCs participated in re-epithelialization, with epidermal differentiation at early and middle stages in full-thickness wounds. To mimic the multilayers of native skin structures, on-site layer-by-layer scaffolds have been developed using electrospun nanofibers [68,69] . This approach allows coculture of multiple types of cells in the same scaffolds, and could precisely control the thickness of fiber layers, density of cell seeding, and the cell type for each cell layer, as well as the composition for each nanofiber layer [68] . A layer of electrospun PCL/COL nanofibers is deposited first followed by cell seeding on top of the fiber layer. By repeating these steps, 3D cell-fiber multilayered structures are prepared and
www.futuremedicine.com
521
Review
Kim & Yoo
the formed construct shows a layered structure with uniform distribution of cells between the layers of PCL/COL nanofibers. In one study, a bilayer skin construct composed of two assembled layers of keratinocytes on a layer of dermal fibroblasts with PCL/COL nanofibers showed similar morphology to the native skin structure after culture for only 3 days. Furthermore, the layers were tightly connected because of the active interaction of keratinocytes and fibroblasts, as well as strong cell-fiber binding. These results suggested that the time needed to create bilayer skin grafts can be reduced significantly to 90%, respectively [70] . In addition, Martins et al. found that the parallel or uniaxially aligned nanofibers induced osteogenesis of human MSCs and increased deposition of mineralized ECM along the predefined fiber direction [71] . For load-bearing applications of electrospun nanofibers, bioceramic materials such as tricalcium phosphate and HAp have been used to enhance bone mineralization and the mechanical strength of scaffolds. HAp is a suitable material for bone regeneration because of its biodegradability, bioactivity and osteoconductive properties, but it is too rigid to shape in bone-defect sites. Therefore, this fiber composite can provide biocompatibility, osteoconductivity and even sufficient mechanical strength to the implantation site. Biocomposite polymeric nanofibers composed of PLLA and HAp were coelectrospun via a single nozzle [72] . The tensile strength of PLLA scaffolds was higher (4.69 MPa) than those of PLLA/HAp (3.10 MPa), whereas the strain at the break of PLLA/HAp scaffolds was increased (43.66%) compared with PLLA scaffolds (25%). In addition, PLLA/HAp composites showed better bone mineralization than PLLA nanofibers. For engineering of cartilage tissue, cartilage-specific biological cues such as chondroitin sulfate (CS), HA, COL and gelatin have been introduced into nanofibrous scaffolds because they help support chondrocyte proliferation and promote/enhance the chondrogenesis of MSCs [73–75] . Cationized gelatin (CG), in which the carboxyl groups of gelatin are converted into amino groups, was grafted onto oxygen plasma-treated PLLA nanofibers using water-soluble carbodiimide as a coupling agent [73] . The CG-modified PLLA NF showed tight attachment of chondrocytes and cell ingrowth into the interior of the mesh, while maintaining its chondrocytic phenotype. In addition, secretion of
522
Nanomedicine (2014) 9(4)
glycosaminoglycan and COL from chondrocytes on CG-PLLA NF was enhanced significantly. In another study, PVA-methacrylate (PVA-MA) and CS-MA were coelectrospun into an ethanol bath to form lowdensity, fibrous scaffolds [74] . In the presence of a photoinitiator, the fibers were rendered insoluble by UVlight crosslinking between MA groups. After 42 days of MSC culture in PVA-MA/CS-MA fiber scaffolds under chondrogenic induction conditions, the scaffolds appeared as cartilage-like constructs similar to hyaline cartilage. Compared with pellet cultures (the conventional method for cartilage engineering), the fiber scaffolds enhanced chondrogenic differentiation of MSCs and, in particular, CS-containing fibers increased the synthesis of cartilage-specific type-II COL. Interestingly, the fate of stem cells to chondrogenic or osteogenic differentiation can be determined by control over the mechanical modulus of electrospun nanofibers [75] . The modulus of coaxial poly(ether sulfone)-PCL fibers (30.6 MPa) was higher than that of pure PCL fibers (7.1 MPa). Embryonic mesenchymal progenitor cells on the softer PCL fibers showed upregulation of expression of chondrocytic genes (Sox9, COL type-II and aggrecan) and chondrocytespecific glycosaminoglycan production in the ECM. By contrast, the stiffer poly(ether sulfone)-PCL fibers supported osteogenesis by promoting expression of osteogenic genes (Runx 2, alkaline phosphatase and osteocalcin). These results suggested that the difference in the mechanical moduli of the fibers caused distinct cytoskeletal organizations, which should be considered for appropriate application of nanofibrous scaffolds. Ligaments & tendons
Ligaments and tendons are dense connective tissues composed of aligned COL molecules, fibrils, fiber bundles and fascicles, and are responsible for the movements and stability of joints. These tissues are torn or ruptured by high physiological loads, and fail to heal because of their hypocellularity and hypovascularity [76] . Electrospun nanofibers can be fabricated into the desired sizes and shapes to meet specific mechanical and biological requirements for the regeneration of ligaments and tendons (Figure 2) . In addition, more complex 3D fibrous structures can be prepared by introducing well-developed textile methods such as weaving, braiding and knitting. Braided nanofibrous scaffolds have been employed for their high tensile strength but they often exhibit poor cellular integration [77] . Mechanical properties of braided fiber scaffolds can be altered by changing braiding parameters such as braided fiber bundle num-
future science group
Therapeutic application of electrospun nanofibrous meshes
Nonwoven nanofibers
Sheet
Review
Aligned nanofibers
Stacked sheet (layer-by-layer)
Tube
Braided
Knitted
Figure 2. Electrospun nanofibrous scaffolds showing the various shapes.
ber, fiber diameter and braiding angle. By braiding three, four or five aligned bundles of electrospun PLA nanofibers, braided nanofibrous scaffolds were fabricated and they showed similar mechanical behavior to native tendons and ligaments during loading [78] . When human MSCs were seeded on braided nanofibrous scaffolds, cells aligned parallel to the direction of the nanofibers and differentiated into the tenogenic lineage in the presence of tenogenic growth factors and upon stimulation with cyclic tensile strain. Knitting methods enable the fabrication of less dense structures. The tensile mechanical properties of knitted structures have been shown to be adequate for the engineering of ligament tissue [79,80] . However, it is difficult to achieve controlled and homogeneous cell seeding because of the high porosity of the knitted structure. When knitted scaffolds composed of PLGA and silk sutures were covered with aligned poly(l-lactic-co-Hcaprolactone) electrospun microfibers, the elastic modulus of composite scaffolds (150 ± 14 MPa) was similar to the modulus of human ligaments [80] . In addition, these scaffolds enabled easy cell seeding, and cells oriented spontaneously along the direction of fiber alignment. Interestingly, they also induced a high level of COL type I and type III, the main components in ligaments. Nerves
Typical therapeutic strategies for neural repair are to create aligned fibrous scaffolds to provide guidance for cell migration and directional axonal regeneration across the lesion site in nerve injuries. Several studies have focused on evaluation of the proliferation and differentiation of nerves, as well as neurite extension of various cell types on aligned fibrous scaffolds fabricated with different materials and fiber sizes. In general, aligned nanofibers enhance the differentiation of stem cells into neural progenitor cells in the
future science group
presence of signaling molecules such as retinoic acid (RA), NGF and brain-derived neural factor, which are involved in neuronal patterning, neural differentiation and axon growth [61,81,82] . Aligned PCL nanofibrous scaffolds containing RA support proliferation and differentiation of murine CE3 embryonic stem cells into neural progenitors by the controlled release of RA from scaffolds [81] . NGF-conjugated aligned PCL-PEG NFs facilitate the neuronal differentiation of MSCs and enhance directional neurite extension [61] . Randomly oriented NFs with physically absorbed NGF show low expression of neural-specific marker genes. For implantation of nanofibrous scaffolds as nerve guide conduits, NFs are often developed to cylindrical or half-cylindrical constructs. Aligned electrospun PCL and ethyl-ethylene phosphate containing human glial cell-derived neurotrophic factor showed maximal electrophysiological recovery in rat sciatic nerves [82] . Polysulfone nerve conduits filled with aligned poly(acrylonitrile-co-methylacrylate) fibrous mats were used as a tibial nerve bridge in rats [83] . The aligned construct significantly improved functional outcome in a 17-mm nerve defect compared with a random construct. Besides alignment, the degree of fiber alignment and fiber diameter can also influence the differentiation of neural progenitor cells and stem cells because those parameters affect the rate of cell elongation [84,85] . Electrospun PLLA fibers with nanoscaled diameters (300 nm) were shown to enhance the neural differentiation of neonatal mouse cerebellum C17.2 stem cells compared with those with a microscale diameter (1.25 Pm), because the cells on nanoaligned fibers showed a longer neuritis extension compared with micron-aligned fibers [84] . Electroconducting polymers such as polypyrrole (PPy), polythiophene, polyaniline and poly(3,4-
www.futuremedicine.com
523
Review
Kim & Yoo
ethylenedioxythiophene) have been used in electrospinning to prepare nerve-tissue engineering scaffolds and neural prostheses to provide electrical stimulation to neurons and nerve tissue [86] . Conductive fibrous meshes were fabricated by growing PPy on random and aligned PLGA nanofibers [87] . When PC12 cells derived from a pheochromocytoma of the rat adrenal medulla were cultured on PPy-coated PLGA nanofibers under electrical stimulation with a potential of 10 mV/cm, they exhibited 40–50% longer neuritis and 40–90% more neurite formation compared with unstimulated cells on the same scaffolds. Furthermore, under electrical stimulation, cells on aligned PPy-PLGA nanofibers showed longer neuritis compared with those on random fibers, which suggested a synergetic effect of electrical stimulation and topographical guidance for application of neural tissue. Blood vessels
Native blood vessels have unique arrangements as tissues with three concentric layers: intima, media and adventitia [88] . The intima consists of a layer of endothelial cells lining the internal surface of the vessel. Smooth muscle cells and COL fibrils have a circumferential orientation in the media, and this layer provides the mechanical strength to withstand the high pressures of the blood circulation. An internal lamina composed mainly of ELAS exists between those two layers, and confers elastic properties to the blood vessel. The adventitia consists of a collagenous ECM that mainly contains fibroblasts and perivascular nerve cells, and adds rigidity to the vessel. Multilayered electrospun constructs have been fabricated to mimic the morphological and mechanical properties of native blood vessels [89,90] . A threelayered electrospun conduit composed of PCL, type-I COL and soluble ELAS was prepared by sequential electrospinning with different polymer ratios [90] ; that is, blended in a ratio of 98:2:0 PCL-ELAS-COL as the intimal layer; 45:45:10, 55:35:10 and 65:25:10 PCL-ELAS-COL as the medial layer; and 70:0:30 PCL-ELAS-COL as the adventitial layer. The artificial intimal layer provided stiffness owing to a higher content of PCL fibers; the artificial medial layer had decreased stiffness owing to increased ELAS content; and the artificial adventitial layer prevented rupture and increased cell adhesion by increasing the number of COL fibers. Unlike artery grafts, small-diameter vascular grafts (diameter,
Therapeutic application of electrospun nanofibrous meshes
Fabricating tissue architecture-mimicking scaffolds is one of the major challenges in the field of tissue engineering. Electrospun nanofibers have been considered as potent techniques for fabricating fibrous scaffolds biomimicking extracellular frameworks. Therapeutic agent-incorporated nanofibrous meshes have widely served as excellent substrates for adhesion, proliferation and differentiation. Many drugs, proteins and nucleic acids were incorporated into the scaffolds for regeneration of skin, musculoskeletal, neural and vascular tissue engineering in aims to control the release of the therapeutic agents. In the current article, we focus on introducing various fabrication techniques for electrospun nanofiber-based scaffolds and subsequent functionalization of nanofibers for therapeutic purposes. We also detail how the therapeutic nanofibrous meshes can be employed in the field of tissue engineering.
Hye Sung Kim1 & Hyuk Sang Yoo*,1,2 1 $EPARTMENTOF"IOMEDICAL-ATERIALS %NGINEERING 3CHOOLOF"IOSCIENCE "IOENGINEERING +ANGWON.ATIONAL 5NIVERSITY #HUNCHEON 2EPUBLIC OF+OREA )NSTITUTEOF"IOSCIENCE"IOTECHNOLOGY +ANGWON.ATIONAL5NIVERSITY 2EPUBLIC OF+OREA
!UTHORFORCORRESPONDENCE 4EL &AX HSYOO KANGWONACKR
Keywords: DRUGDELIVERYsELECTROSPINNINGsNANOlBERsSCAFFOLDsTISSUEENGINEERING
Strategies of regenerative medicine are focused on the restoration of tissue architectures and biologic function in host-tissue lesions by implantation of supportive scaffolds containing biomolecules or cells [1–3] . Recently, considerable interest has been given to biomimetic and biologically active scaffolds that can replace the natural extracellular matrix (ECM) and provide a favorable environment to host cells to repopulate and resynthesize a new natural matrix. Among those scaffolds, electrospun nanofibrous meshes (NFs) have been used for tissue engineering because of their unique features that mimic the structure of fibrillar collagen (COL) and elastin (ELAS) in the natural ECM [4–6] . In addition, the high porosity of NFs facilitates the transport of gases, nutrients and regulatory factors for the survival and proliferation of cells [7] . Moreover, a high surface area-to-volume ratio of NFs not only provides wider substrate dimension for cell attachment, but also offers better incorporation of biomolecules on their surfaces compared with nonfibrous scaffolds [1–3] . Thus, various electrospun nanofibers have been
10.2217/NMM.13.224 © 2014 Future Medicine Ltd
biofunctionalized for incorporating bioactive molecules by pre- and/or post-electrospinning processes [4,5] . Here, we describe the fabrication methods for electrospun nanofiber-based scaffolds, an overview of the methods for functionalization of fibrous scaffolds and application of NFs for tissue engineering. Principle of electrospinning Electrospinning process
Electrospinning is a unique approach using electrostatic forces to produce fine fibers from polymer solutions [8] . The fabricated fibers feature a thin diameter in the range of a few nanometers to submicrometers, as well as a large surface area-to-volume ratio. Thus, the electrospun NFs show high porosity and interconnected fibrous networks that mimic ECM architectures. An electrospinning system largely comprises a high-voltage power supply, a spinneret (e.g., single/dual needle and pipette tip), and a grounded collector (usually a metal plate or rotating mandrel). The completely dissolved polymer solution is injected into the needle.
Nanomedicine (2014) 9(4), 517–533
part of
ISSN 1743-5889
517
Review
Kim & Yoo
A polymer droplet at the tip of a spinneret is electrically charged with high electrical potential. If the repulsive electrical forces of the droplet overcome surface tension forces, a charged jet of the polymer solution is ejected from the tip of the Taylor cone. The polymeric fibers deposit on the grounded collector with a rapid whipping motion and solvent evaporation [9] . In comparison with other methods for producing thin fibers, such as phase separation, drawing and self-assembly, electrospinning is advantageous in terms of economics and technical viewpoints for manufacturing nanofiber-based scaffolds for tissue engineering. Polymers for electrospinning Natural polymers
Various polymers have been used in electrospinning for tissue-engineered scaffolds. Typical natural polymers for fabrication of electrospun NFs include COL, gelatin, chitosan, chitin, ELAS, casein, cellulose acetate, silk protein and fibrinogen [10] . In general, natural polymers exhibit better biocompatibility and low immunogenicity compared with synthetic polymers. COL has been used in various tissue-engineering applications because it is the most abundant protein in the mammalian ECM, is relatively nonimmunogenic and a good sequester of many factors for the maintenance and regeneration of tissue. Most studies have used 1,1,1,3,3,3-hexafluoroisopropanol or 2,2,2,-trifluoroethanol as the solvent to electrospin COL fibers because of its high volatility. Unfortunately, those solvents not only disrupt the native structure of COL, but also decrease the denaturation temperature, which leads to 45% of COL being lost during electrospinning [11,12] . Thus, several researchers have employed 40–50% acetic acid, and they confirmed that diluted acetic acid was better for protecting the native helical structure of COL than 1,1,1,3,3,3-hexafluoroisopropanol [13,14] . However, denaturation of the secondary structure leads to rapid degradation of electrospun COL fibers in aqueous media. Thus, crosslinking reagents such as glutaraldehyde [15,16] , 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride [17] , N-hydroxysuccinimide [18] , and genipin [19] have been used to stabilize the morphology of electrospun COL fibers. However, glutaraldehyde produces toxic side products and subsequently calcifies COL substrates. With respect to genipin crosslinking, it requires long processing times (>48 h) [19,20] . Although these methods could delay the solubilization of COL fibers, fiber fusion or disappearance of fibrous morphology occurs during crosslinking. Chew et al. introduced photochemical crosslinking by mixing COL polymers with 0.1% (weight/volume) of rose bengal as a photo-initiator before electrospin-
518
Nanomedicine (2014) 9(4)
ning, and electrospun COL fibers were irradiated with an argon laser for crosslinking [13] . They found that fibrous morphology was well maintained after photochemical crosslinking, and that the low concentration of photoinitiator could minimize cytotoxicity. Chitosan is one of the most abundant polysaccharides and is derived from chitin. It has been employed for wound-dressing material because of its biocompatibility, biodegradability and bioactivity. In addition, chitosan promotes hemostasis and aids tissue regeneration [21,22] . However, electrospinning of chitosan is a challenge because of its limited solubility and high viscosity. Chitosan is soluble in organic acids such as dilute aqueous acetic acid, formic acid and lactic acid. However, it is partially soluble in inorganic acids and insoluble in water. Although chitosan is dissolved in the solvent, it is positively charged because of free amino groups in the chitosan structure, which make chitosan solutions highly viscous [23,24] . The cationic charge of the solution inhibits the chain entanglement, and the strong hydrogen bonds in chitosan solution prevent the movement of polymeric chains exposed to the electrical field, which stop fiber formation [25–27] . Thus, to improve spinnability, chitosan-based nanofibers have been prepared by blending chitosan with second polymers such as gelatin, silk fibroin, poly(ethylene oxide), and poly(vinyl alcohol) (PVA) [28,29] . Hyaluronic acid (HA), a linear polysaccharide composed of repeating glucuronic acid and N-acetylglucosamine, is the main component of the ECM in connective tissues. HA performs important roles as a molecular filter, shock absorber and support COL fibril structure. In addition, it has been used extensively in the biomedical field due to its excellent biocompatibility and biodegradability. It is difficult to create uniform size fibers from hyaluronan solution using electrospinning due to the high viscosity, surface tension, and strong water retention ability of the hyaluronan solution. Thus, Li et al. and Um et al. performed electrospinning in combination with air flow at 57°C with a 70 ft/h flow rate and successfully produced HA nanofibers [30,31] . Keratins are the major structural fibrous protein of hair, feathers, wool and nails, which are characteristically abundant in cysteine residues. Wool keratins have been demonstrated to be a favorable biomaterial of scaffolds for fibroblasts and osteoblasts due to their cell adhesion sequences, arginine-glycine-aspartic acid and leucine–aspartic acid–valine, biocompatibility and biodegradability. Unfortunately, keratin or keratin-based materials are difficult to electrospin due to their low molecular weight (9–60 kDa); thus, they are alternatively blended with synthetic polymers such as poly(lactic-co-glycolide) (PLGA), poly-l-lactide
future science group
Therapeutic application of electrospun nanofibrous meshes
(PLLA) and poly(3-hydroxybutyrate-co-3-hydroxyvalerate) to improve the poor mechanical properties of the electrospun scaffolds [32–34] . Synthetic polymers
Synthetic polymers offer several advantages over natural polymers. They can be tailored to give the requisite mechanical properties, thermal stability and desired degradation profile. In addition, the range of solvents available to dissolve synthetic polymers for electrospinning is relatively wider than that for natural polymers. Typical organic solvents used in electrospinning are chloroform, ethanol, dimethylformamide, dichloromethane and their mixtures [35] . The selection of the solvent for preparing polymer solutions is important because the morphology and size of electrospun nanofibers are strongly dependent upon solution properties such as viscosity and surface tension, which are attributed to polymer concentrations and solvent properties (including polarity, solubility and volatility) [36,37] . Biocompatible and biodegradable polymers such as polyglycolide, polylactide (PLA), poly(H-caprolactone) (PCL) and PLGA have been electrospun into nanofibrous scaffolds for biomedical applications. However, electrospun NFs composed of biodegradable polymers such as polyesters or poly(D-hydroxy acids) are mechanically unstable and cannot withstand implantation because of the rapid degradation rate of the mesh [38,39] . The highly porous structure of the mesh facilitates degradation, and the degradation rate of the mesh is much faster than the recovery rate of the tissue. Hence, these features limit the use of meshes for tissue engineering. Nonbiodegradable polymers such as polyurethane and polyesterurethane possess substantial mechanical stability. However, they might interfere with the turnover and remodeling of tissue owing to their slow degradation [40,41] . Thus, the mechanical stability and degradation rate of nanofiber-based scaffolds can be controlled by mixing various polymers exhibiting different degradation profiles [42] . In addition, PLGA (a copolymer composed of lactic acid and glycolic acid) shows different degradation profiles according to the ratio of lactic acid and glycolic acid [43] . The higher the content of glycolide units, the lower the time required for degradation. Shin et al. reported that the mechanical properties of nanofiber scaffolds prepared with PLGA composed of 75–85% glycolic acid are likely to be similar to those of cartilage [44] . In addition, these PLGA nanofiber scaffolds degraded dramatically after 7 weeks of incubation under physiological conditions. These findings suggested that nanofiber-based PLGA scaffolds could support cell proliferation and tissue development for t2 months.
future science group
Review
For therapeutic application of electrospun nanofibers, various copolymers have been synthesized and electrospun to modulate cell affinity, as well as drug encapsulation/release of nanofibers. Electrospun nanofibers composed of biodegradable hydrophobic polyesters generally have good mechanical properties, while cells show poor adhesion the polymeric nanofiber. Thus, a proper hydrophilic polymer segment is widely conjugated to the hydrophobic polymers to enhance the biocompatibility of the electrospun nanofibers by increasing hydrophilicity. Interestingly, cell adhesion can be inhibited by increasing the proportion of hydrophilic segment in copolymers [45] . Poly(l-lactide-co-glycolide) (PLLGA), PLLGA blended with PLLGA-poly(ethylene glycol) (PEG) copolymer or pure PEG(PLLGA/PEG) were electrospun onto nonwoven mesh, respectively. Upon the addition of PLLGA-PEG or pure PEG, the hydrophilicity of the electrospun NF significantly improved. In addition, PLLGA/PLLGA-PEG significantly inhibited the adsorption of proteins and proliferation of cells, unlike PLLGA/PEG nanofibers. By introducing stimuli-responsive moieties in copolymers, drug encapsulation/release of electrospun nanofibers can be controlled under specific environments, such as temperature, pH, electric field, magnetic field and light. Poly(N-isopropylacrylamide) (PNIPAAm) has been extensively studied for a thermosensitive drug delivery system. It is soluble in water below it has a lower critical solution temperature (32°C) and is gelated at higher temperature. Thus, electorspun nanofibers composed of the NIPAAm homopolymers are not stable in water and disperse easily; therefore, copolymerization with crosslinkable moieties is required to obtain stable nanofibers in an aqueous medium [46,47] . PNIPAAm-co-N-hydroxymethylacrylamide (HMAAm) nanofibers showed an ‘on–off ’ switchable release of dextran according to a heating–cooling cycle. Upon thermal curing of nanofibers, the methylol groups of HMAAm in the copolymers were crosslinked by self-condensation. The crosslinked nanofibers preserved their morphology and reversible volume changes in aqueous media in response to cycles of temperature alternation, while noncrosslinked nanofibers collapsed and dispersed quickly in the aqueous solution. In addition, dextran release from the crosslinked nanofiber accelerated during the heating cycle, whereas a negligible amount of dextran was evolved during the cooling process. Strategies to biofunctionalize electrospun nanofibers To apply electrospun nanofibers to biomedical uses, bioactive factors such as drugs, growth factors, nucleic
www.futuremedicine.com
519
Review
Kim & Yoo
acids, enzymes and antibiotics have been incorporated into fibrous scaffolds via various methods (Figure 1) . Blend electrospinning
A simple approach for drug incorporation into nanofibrous scaffolds is the blending of polymer solutions with drugs in pre-electrospinning. In single-nozzle electrospinning, hydrophobic drugs that are soluble in common organic solvents can be incorporated into nanofibers, whereas water-soluble bioactive agents such as proteins and nucleic acids, are hard to be blended in
homogeneously, which leads to an initial burst release within 24 h [48–50] . Thus, in order to improve protein incorporation and release profile, hydrophilic additives, such as PEG, gelatin and hydroxyapatite (HAp), have been loaded together with protein during blend electrospinning [51–53] . The addition of hydrophilic polymers or amphiphilic copolymers could enhance protein miscibility within nanofibers, as well as the hydrophilicity and water uptake of the nanofibers, which accelerates protein release from electrospun nanofibers. Unfortunately, the denaturation of bioactive molecules by
• Blend electrospinning
Electrospinning of drug/polymer blend through single nozzle
• Coaxial electrospinning
Dual-nozzle electrospinning or drug/polymer core and polymer shell
• Postelectrospinning: surface modification
Surface functionalization
Physical/chemical drug incorporation
Electrospun nanofibers Figure 1. Various techniques of fabricating drug-incorporated electrospun nanofibers; blend electrospinning, coaxial electrospinning and surface modification.
520
Nanomedicine (2014) 9(4)
future science group
Therapeutic application of electrospun nanofibrous meshes
exposure to organic solvents and electric fields during electrospinning is problematic. Thus, bioactive molecules are, in general, emulsified or encapsulated into nanoparticles to improve dispersion within the nanofibers while maintaining bioactivity [54,55] .
Review
incorporation of various types of nucleic acids (including plasmid DNA and siRNA), but also liberated the poly(ethyleneimine)–DNA complex from the nanofibers, which was strictly controlled by digestion of the linker under an abundant MMP-2 environment such as diabetic ulcers [62,63] .
Coaxial electrospinning
Alternatively, coaxial electrospinning can be carried out through a dual nozzle (a spinneret composed of two coaxial capillaries) for electrospinning two separate immiscible polymer solutions into coreshell-structured nanofibers simultaneously. The aqueous-based biological molecules can be incorporated separately into the core of nanofibers, thereby minimizing interactions with organic solvents during electrospinning [56–58] . Therefore, coaxial electrospinning is suitable for the preparation of NFs for delivery of unstable bioactive molecules such as growth factors. It has been shown that the growth factors released from coaxial nanofibers maintain the same level of bioactivity as ‘fresh’ growth factors [57] . Choi et al. coelectrospun a PCL-PEG block copolymer with a basic FGF (bFGF) solution using a dual nozzle to incorporate bFGF into the core of nanofibers, and then EGF was conjugated chemically on the shell of coaxial nanofibers [58] . The bFGF/EGF nanofibers showed binary release profiles of each growth factor; the core-encapsulated bFGF showed a high initial burst in 24 h, whereas the EGF was released in 7 days. Eventually, the bFGF/EGF NFs could promote wound recovery in diabetic ulcers because different release profiles of growth factors offered precise control of cell proliferation during re-epithelialization. Surface modification
Surface modification of nanofibers with biomolecules is another promising approach for introducing biofunctionality into nanofibers. Biomolecules are absorbed physically on NFs via electrostatic interactions, hydrogen bonding, hydrophobic interactions and van der Waals interactions [59] . However, modulating the degree of drug loading and release kinetics is difficult. For a slow and prolonged release of a therapeutic agent, a target biomolecule is immobilized chemically on the surface of nanofibers [60,61] . However, in the case of gene delivery, which requires endocytosis of a therapeutic gene, covalent conjugation of nucleic acids onto the scaffolds does not lead to a therapeutic effect because of the limited release of the gene from the scaffolds. Thus, the rate of release of immobilized target molecules can be controlled precisely via introduction of a responsive moiety to local external cues. For example, poly(ethyleneimine)immobilized nanofibers via a matrix metalloproteinase-cleavable linker not only showed high efficiency of
future science group
Biomedical application of the electrospun nanofibers-based scaffold Tissue regeneration Skin
Electrospun NFs have received much attention as potential candidates for wound-dressing materials. As a result of their high surface-to-volume ratio and nanoporous structure, NFs can absorb wound exudate effectively and keep the wound site under a moisturerich environment during healing. In addition, oxygen can penetrate easily through the mesh, while bacterial infiltration is strictly excluded because of the nanoscaled pores of the NF [64,65] . To enhance the efficiency of skin regeneration, NFs are usually administrated in combination with various drugs: growth factors, antibiotics and hemostatic factors. In diabetic ulcers, EGF-immobilized PCL/PCL-PEG NFs had a faster rate of wound closure than EGF solutions or NFs without EGF [60] . The self-regeneration of skin after minor epidermal injury is remarkable. However, if the injury is severe (e.g., massive, deep loss of the epidermis and dermis in full-thickness skin wounds), the lost skin cannot regenerate spontaneously [66] . Thus, electrospun NFs as skin substitutes have been developed with supporting cells such as keratinocytes, dermal fibroblasts and mesenchymal stem cells (MSCs). COL/PLGA-blended NFs have been modified by conjugation with a CD29 antibody for rapid attachment of bone marrow (BM)derived MSCs [67] . Implanted BM-MSCs attached to COL/PLGA NFs promoted ingrowth of epithelial edges and COL synthesis. In addition, implanted BM-MSCs expressed keratin 10 and filaggrin, which suggested that BM-MSCs participated in re-epithelialization, with epidermal differentiation at early and middle stages in full-thickness wounds. To mimic the multilayers of native skin structures, on-site layer-by-layer scaffolds have been developed using electrospun nanofibers [68,69] . This approach allows coculture of multiple types of cells in the same scaffolds, and could precisely control the thickness of fiber layers, density of cell seeding, and the cell type for each cell layer, as well as the composition for each nanofiber layer [68] . A layer of electrospun PCL/COL nanofibers is deposited first followed by cell seeding on top of the fiber layer. By repeating these steps, 3D cell-fiber multilayered structures are prepared and
www.futuremedicine.com
521
Review
Kim & Yoo
the formed construct shows a layered structure with uniform distribution of cells between the layers of PCL/COL nanofibers. In one study, a bilayer skin construct composed of two assembled layers of keratinocytes on a layer of dermal fibroblasts with PCL/COL nanofibers showed similar morphology to the native skin structure after culture for only 3 days. Furthermore, the layers were tightly connected because of the active interaction of keratinocytes and fibroblasts, as well as strong cell-fiber binding. These results suggested that the time needed to create bilayer skin grafts can be reduced significantly to 90%, respectively [70] . In addition, Martins et al. found that the parallel or uniaxially aligned nanofibers induced osteogenesis of human MSCs and increased deposition of mineralized ECM along the predefined fiber direction [71] . For load-bearing applications of electrospun nanofibers, bioceramic materials such as tricalcium phosphate and HAp have been used to enhance bone mineralization and the mechanical strength of scaffolds. HAp is a suitable material for bone regeneration because of its biodegradability, bioactivity and osteoconductive properties, but it is too rigid to shape in bone-defect sites. Therefore, this fiber composite can provide biocompatibility, osteoconductivity and even sufficient mechanical strength to the implantation site. Biocomposite polymeric nanofibers composed of PLLA and HAp were coelectrospun via a single nozzle [72] . The tensile strength of PLLA scaffolds was higher (4.69 MPa) than those of PLLA/HAp (3.10 MPa), whereas the strain at the break of PLLA/HAp scaffolds was increased (43.66%) compared with PLLA scaffolds (25%). In addition, PLLA/HAp composites showed better bone mineralization than PLLA nanofibers. For engineering of cartilage tissue, cartilage-specific biological cues such as chondroitin sulfate (CS), HA, COL and gelatin have been introduced into nanofibrous scaffolds because they help support chondrocyte proliferation and promote/enhance the chondrogenesis of MSCs [73–75] . Cationized gelatin (CG), in which the carboxyl groups of gelatin are converted into amino groups, was grafted onto oxygen plasma-treated PLLA nanofibers using water-soluble carbodiimide as a coupling agent [73] . The CG-modified PLLA NF showed tight attachment of chondrocytes and cell ingrowth into the interior of the mesh, while maintaining its chondrocytic phenotype. In addition, secretion of
522
Nanomedicine (2014) 9(4)
glycosaminoglycan and COL from chondrocytes on CG-PLLA NF was enhanced significantly. In another study, PVA-methacrylate (PVA-MA) and CS-MA were coelectrospun into an ethanol bath to form lowdensity, fibrous scaffolds [74] . In the presence of a photoinitiator, the fibers were rendered insoluble by UVlight crosslinking between MA groups. After 42 days of MSC culture in PVA-MA/CS-MA fiber scaffolds under chondrogenic induction conditions, the scaffolds appeared as cartilage-like constructs similar to hyaline cartilage. Compared with pellet cultures (the conventional method for cartilage engineering), the fiber scaffolds enhanced chondrogenic differentiation of MSCs and, in particular, CS-containing fibers increased the synthesis of cartilage-specific type-II COL. Interestingly, the fate of stem cells to chondrogenic or osteogenic differentiation can be determined by control over the mechanical modulus of electrospun nanofibers [75] . The modulus of coaxial poly(ether sulfone)-PCL fibers (30.6 MPa) was higher than that of pure PCL fibers (7.1 MPa). Embryonic mesenchymal progenitor cells on the softer PCL fibers showed upregulation of expression of chondrocytic genes (Sox9, COL type-II and aggrecan) and chondrocytespecific glycosaminoglycan production in the ECM. By contrast, the stiffer poly(ether sulfone)-PCL fibers supported osteogenesis by promoting expression of osteogenic genes (Runx 2, alkaline phosphatase and osteocalcin). These results suggested that the difference in the mechanical moduli of the fibers caused distinct cytoskeletal organizations, which should be considered for appropriate application of nanofibrous scaffolds. Ligaments & tendons
Ligaments and tendons are dense connective tissues composed of aligned COL molecules, fibrils, fiber bundles and fascicles, and are responsible for the movements and stability of joints. These tissues are torn or ruptured by high physiological loads, and fail to heal because of their hypocellularity and hypovascularity [76] . Electrospun nanofibers can be fabricated into the desired sizes and shapes to meet specific mechanical and biological requirements for the regeneration of ligaments and tendons (Figure 2) . In addition, more complex 3D fibrous structures can be prepared by introducing well-developed textile methods such as weaving, braiding and knitting. Braided nanofibrous scaffolds have been employed for their high tensile strength but they often exhibit poor cellular integration [77] . Mechanical properties of braided fiber scaffolds can be altered by changing braiding parameters such as braided fiber bundle num-
future science group
Therapeutic application of electrospun nanofibrous meshes
Nonwoven nanofibers
Sheet
Review
Aligned nanofibers
Stacked sheet (layer-by-layer)
Tube
Braided
Knitted
Figure 2. Electrospun nanofibrous scaffolds showing the various shapes.
ber, fiber diameter and braiding angle. By braiding three, four or five aligned bundles of electrospun PLA nanofibers, braided nanofibrous scaffolds were fabricated and they showed similar mechanical behavior to native tendons and ligaments during loading [78] . When human MSCs were seeded on braided nanofibrous scaffolds, cells aligned parallel to the direction of the nanofibers and differentiated into the tenogenic lineage in the presence of tenogenic growth factors and upon stimulation with cyclic tensile strain. Knitting methods enable the fabrication of less dense structures. The tensile mechanical properties of knitted structures have been shown to be adequate for the engineering of ligament tissue [79,80] . However, it is difficult to achieve controlled and homogeneous cell seeding because of the high porosity of the knitted structure. When knitted scaffolds composed of PLGA and silk sutures were covered with aligned poly(l-lactic-co-Hcaprolactone) electrospun microfibers, the elastic modulus of composite scaffolds (150 ± 14 MPa) was similar to the modulus of human ligaments [80] . In addition, these scaffolds enabled easy cell seeding, and cells oriented spontaneously along the direction of fiber alignment. Interestingly, they also induced a high level of COL type I and type III, the main components in ligaments. Nerves
Typical therapeutic strategies for neural repair are to create aligned fibrous scaffolds to provide guidance for cell migration and directional axonal regeneration across the lesion site in nerve injuries. Several studies have focused on evaluation of the proliferation and differentiation of nerves, as well as neurite extension of various cell types on aligned fibrous scaffolds fabricated with different materials and fiber sizes. In general, aligned nanofibers enhance the differentiation of stem cells into neural progenitor cells in the
future science group
presence of signaling molecules such as retinoic acid (RA), NGF and brain-derived neural factor, which are involved in neuronal patterning, neural differentiation and axon growth [61,81,82] . Aligned PCL nanofibrous scaffolds containing RA support proliferation and differentiation of murine CE3 embryonic stem cells into neural progenitors by the controlled release of RA from scaffolds [81] . NGF-conjugated aligned PCL-PEG NFs facilitate the neuronal differentiation of MSCs and enhance directional neurite extension [61] . Randomly oriented NFs with physically absorbed NGF show low expression of neural-specific marker genes. For implantation of nanofibrous scaffolds as nerve guide conduits, NFs are often developed to cylindrical or half-cylindrical constructs. Aligned electrospun PCL and ethyl-ethylene phosphate containing human glial cell-derived neurotrophic factor showed maximal electrophysiological recovery in rat sciatic nerves [82] . Polysulfone nerve conduits filled with aligned poly(acrylonitrile-co-methylacrylate) fibrous mats were used as a tibial nerve bridge in rats [83] . The aligned construct significantly improved functional outcome in a 17-mm nerve defect compared with a random construct. Besides alignment, the degree of fiber alignment and fiber diameter can also influence the differentiation of neural progenitor cells and stem cells because those parameters affect the rate of cell elongation [84,85] . Electrospun PLLA fibers with nanoscaled diameters (300 nm) were shown to enhance the neural differentiation of neonatal mouse cerebellum C17.2 stem cells compared with those with a microscale diameter (1.25 Pm), because the cells on nanoaligned fibers showed a longer neuritis extension compared with micron-aligned fibers [84] . Electroconducting polymers such as polypyrrole (PPy), polythiophene, polyaniline and poly(3,4-
www.futuremedicine.com
523
Review
Kim & Yoo
ethylenedioxythiophene) have been used in electrospinning to prepare nerve-tissue engineering scaffolds and neural prostheses to provide electrical stimulation to neurons and nerve tissue [86] . Conductive fibrous meshes were fabricated by growing PPy on random and aligned PLGA nanofibers [87] . When PC12 cells derived from a pheochromocytoma of the rat adrenal medulla were cultured on PPy-coated PLGA nanofibers under electrical stimulation with a potential of 10 mV/cm, they exhibited 40–50% longer neuritis and 40–90% more neurite formation compared with unstimulated cells on the same scaffolds. Furthermore, under electrical stimulation, cells on aligned PPy-PLGA nanofibers showed longer neuritis compared with those on random fibers, which suggested a synergetic effect of electrical stimulation and topographical guidance for application of neural tissue. Blood vessels
Native blood vessels have unique arrangements as tissues with three concentric layers: intima, media and adventitia [88] . The intima consists of a layer of endothelial cells lining the internal surface of the vessel. Smooth muscle cells and COL fibrils have a circumferential orientation in the media, and this layer provides the mechanical strength to withstand the high pressures of the blood circulation. An internal lamina composed mainly of ELAS exists between those two layers, and confers elastic properties to the blood vessel. The adventitia consists of a collagenous ECM that mainly contains fibroblasts and perivascular nerve cells, and adds rigidity to the vessel. Multilayered electrospun constructs have been fabricated to mimic the morphological and mechanical properties of native blood vessels [89,90] . A threelayered electrospun conduit composed of PCL, type-I COL and soluble ELAS was prepared by sequential electrospinning with different polymer ratios [90] ; that is, blended in a ratio of 98:2:0 PCL-ELAS-COL as the intimal layer; 45:45:10, 55:35:10 and 65:25:10 PCL-ELAS-COL as the medial layer; and 70:0:30 PCL-ELAS-COL as the adventitial layer. The artificial intimal layer provided stiffness owing to a higher content of PCL fibers; the artificial medial layer had decreased stiffness owing to increased ELAS content; and the artificial adventitial layer prevented rupture and increased cell adhesion by increasing the number of COL fibers. Unlike artery grafts, small-diameter vascular grafts (diameter,
