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APPENDIX A: PAPERS PUBLISHED

BioMed Central

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Journal of Biomedical Science

Open AccessReviewDevelopment of biomaterial scaffold for nerve tissue engineering: Biomaterial mediated neural regenerationAnuradha Subramanian, Uma Maheswari Krishnan and Swaminathan Sethuraman*

Address: Center for Nanotechnology & Advanced Biomaterials, School of Chemical & Biotechnology, SASTRA University, Thanjavur, India

Email: Anuradha Subramanian - anuradha@bioengg.sastra.edu; Uma Maheswari Krishnan - umakrishnan@biotech.sastra.edu; Swaminathan Sethuraman* - swami@sastra.edu

* Corresponding author

AbstractNeural tissue repair and regeneration strategies have received a great deal of attention because itdirectly affects the quality of the patient's life. There are many scientific challenges to regeneratenerve while using conventional autologous nerve grafts and from the newly developed therapeuticstrategies for the reconstruction of damaged nerves. Recent advancements in nerve regenerationhave involved the application of tissue engineering principles and this has evolved a new perspectiveto neural therapy. The success of neural tissue engineering is mainly based on the regulation of cellbehavior and tissue progression through the development of a synthetic scaffold that is analogousto the natural extracellular matrix and can support three-dimensional cell cultures. As the naturalextracellular matrix provides an ideal environment for topographical, electrical and chemical cuesto the adhesion and proliferation of neural cells, there exists a need to develop a synthetic scaffoldthat would be biocompatible, immunologically inert, conducting, biodegradable, and infection-resistant biomaterial to support neurite outgrowth. This review outlines the rationale for effectiveneural tissue engineering through the use of suitable biomaterials and scaffolding techniques forfabrication of a construct that would allow the neurons to adhere, proliferate and eventually formnerves.

IntroductionThe human brain is analogous to a black box of informa-tion and unraveling its mysteries is essential to under-stand its complex relationship with the variouscomponents of the peripheral and central nervous sys-tems. This information is vital to probe the causes for var-ious neural disorders and arrive at a plausible therapy forthe treatment of ischemic, metabolic, congenital, ordegenerative disorders of the central or peripheral nervoussystems. Conventionally, autologous grafts are goldstandards and have been used to treat neural defects [1-3].

However, autografts have limitations that include short-age of nerves since it is taken from the patient. Moreover,there is a mismatch of donor-site nerve size with the recip-ient site, neuroma formation and lack of functional recov-ery [4,5]. Allogenic grafts, which are isolated fromcadavers, are not limited by supply but suffer from host-graft immune rejection [6]. To overcome immune rejec-tion, several studies have been conducted to examine thepotency of acellular nerve grafts [7,8]. However, as acellu-lar nerve graft lacks viable cells, nerve regeneration andremodeling of extracellular matrix have been delayed [8].

Published: 25 November 2009

Journal of Biomedical Science 2009, 16:108 doi:10.1186/1423-0127-16-108

Received: 26 September 2009Accepted: 25 November 2009

This article is available from: http://www.jbiomedsci.com/content/16/1/108

© 2009 Subramanian et al; licensee BioMed Central Ltd. This is an Open Access article distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/2.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.

The cost of publication in Journal of Biomedical Scienceis bourne by the National Science Council, Taiwan.

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The use of pre-degenerated nerve grafts having highmatrix metalloproteinase (MMP) expression shows somepotential as it degrades the inhibitory chondroitin sul-phate and proteoglycans thereby retaining the ability topromote nerve regeneration even in the absence of cells[8,9].

Recent advances in nanotechnology [10] and tissue engi-neering [11,12] have been found to cover a broad range ofapplications in regenerative medicine and offer the mosteffective strategy to repair neural defects. The major deter-minant in all tissue engineering research is to regulate thecell behavior and tissue progression through the develop-ment and design of synthetic extracellular matrix ana-logues of novel biomaterials to support three-dimensional cell culture and tissue regeneration. Idealproperties of a scaffold for nerve regeneration are biocom-patibility, less inflammatory, controlled biodegradabilitywith non-toxic degradative products, porosity for vascu-larization and cell migration and three-dimensionalmatrices with appropriate mechanical properties to mimicthe extracellular matrix [13-15]. Figure 1 shows the vari-ous characteristics desired for an ideal scaffold for neuralregeneration.

Polymeric biomaterials are widely preferred as scaffoldsfor peripheral and central nerve regeneration both in vitroand in vivo [16-19]. There is a wide choice of polymersavailable with programmable biodegradability, non-toxic/non-inflammatory nature, mechanical propertiessimilar to the tissue to be replaced, high porosity that pro-motes cell attachment and growth, economical and sim-ple manufacturing processes along with a potential forchemical modification leading to increased interactionwith normal tissue [20]. Several techniques such as

nanofiber self-assembly, solvent casting and particulateleaching, gas foaming, emulsification/freeze-drying, liq-uid-liquid phase separation, electrospinning and compu-ter aided design and manufacturing techniques have beenemployed to fabricate tissue engineering scaffolds withvarying degrees of success [21-24]. Various attempts madetowards repairing neural defects have been discussed inthe following sections along with the rationale behindselection of a suitable scaffold material for successful neu-ral tissue engineering.

Response to Injury & RepairRegeneration strategies for peripheral and central nervoussystem damage have not been very successful due to lackof knowledge about the mechanisms of nerve injury andrepair [25]. Nerve cells have been found capable of easilybridging gaps of less than 6 mm [26]. Thus, regeneratingnerves across larger gaps are medically challenging phe-nomenon since, injuries of less impact in peripheral nerv-ous system (PNS) heal by formation of fibrin cable acrossthe gap [4]. This eventually allows the Schwann cells tomigrate from both the nerve ends, thereby orienting thebungner bands to promote neurite outgrowth [27]. Pres-ently, researchers have investigated a suitable strategy toenhance the formation of Bungner bands and have iden-tified microstructured biomaterial filaments provide abetter topography, promoting bungner band formationeven in the absence of biological factors [28]. Addition-ally, after injury, the protein synthesis and degradationmachinery in axons play a major role in the initiation ofgrowth cone formation [29]. However, the myelination ofcentral nervous system (CNS) is distinct from that ofperipheral nervous system (PNS). Astrocytes and oli-godendrocytes are found in the CNS which also marks akey difference between CNS and PNS in their response toinjury [16]. In the case of CNS, spontaneous regenerationis impossible due to its own inhibitory environment[16,25]. This includes glial scar formation and accumula-tion of myelin-associated inhibitors such as chondroitinsulphate proteoglycans [30,31]. Moreover, the primaryinjury in CNS expands further by damaging the nerve cellsdue to secretion of free radicals from the blood via theblood-brain barrier resulting in secondary injury whichimpedes the regeneration potency of these cells [32].Many strategies have been attempted to improve theregenerative potency of neurons such as cell therapy,exogenous delivery of growth factors and tissue engineer-ing approaches; each in turn restores the function withvarying degrees of success.

Regeneration Potential of Neural CellsGuenard et al. have observed that the potential use ofautologous Schwann cells that aid in CNS regeneration[33]. This novel approach modifies the environment ofthe CNS by transplanting Schwann cells which enhances

Ideal properties of scaffoldFigure 1Ideal properties of scaffold.

IDEAL NEURAL SCAFFOLD

Mimics native Extracellular matrix Provide better contact guidance on neurite outgrowth

Promote cell adhesion, viability and proliferation

Biodegradability

Growth factor delivery Cell Delivery Vehicle

Multiple cues (topographical, mechanical, biochemical and

electrical cue)

Radical scavenging ability

Replace lost cells/tissues

Eliminate chronic inflammation

Stimulate endogenous stem progenitor cells

Minimize secondary progression of injury

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axonal regeneration. This may be due to the fact thatSchwann cells are the myelinating glial cells in the PNSand are known to play a key role in Wallerian degenera-tion and subsequent CNS regeneration [4,20,34,35].Schwann cells promote neural regeneration and remyeli-nation by secreting adhesion molecules L1 and neural celladhesion molecule (N-CAM), extracellular molecules(collagen and laminin) and a number of tropic factorssuch as nerve growth factor (NGF), brain-derived neuro-trophic factor (BDNF) and neurotrophin-3 (NT-3) [36].Though the Schwann cells have regenerative potential,studies have shown the undesirable effects such as inhibi-tion of Schwann cell migration into the CNS, delayedfunctional recovery and in certain cases, the Schwann cellsmay follow its original white matter pathway instead ofthe intended grey matter pathway [36].

The nearly impossible task of rebuilding the nervous sys-tem has since undergone a dramatic transformation withthe discovery of stem cells [30,37]. The differentiation ofneural stem cells to specific cell lineages has been control-led by constructing scaffolds of composite biomaterialsthat consist of extracellular matrix (ECM) componentsand growth factors [38]. The embryonic stem cell-derivedoligodendrocytes have been recognized to myelinateaxons in culture and to replace lost myelin in the injuredadult CNS [39]. Human embryonic stem cell-derived oli-godendrocyte progenitors when transplanted into injuredspinal cords of rat proved to be a safe procedure, resultingin improved locomotor function [40]. The ideal cell cul-ture conditions for embryonic stem cell proliferation anddifferentiation in fibrin-based scaffolds has been identi-fied for neural tissue engineering applications [41].

Investigations on astrocytes have supported its role in theinduction of neurogenesis from adult neural stem cells[42]. The directed growth of astrocytes on polymer sub-strates provides an innovative approach to promote con-trolled outgrowth and differentiation of neural stem cells[43]. Though the astrocytes contribute the cytotrophiceffects in neural repair, at certain stages in response toinjury it inhibits neurite outgrowth by releasing signalsinhibiting neurite extension [44]. Researchers have triedto improve the cytotropic effects of astrocytes in regenera-tion by harnessing its cytotoxic effects [44].

In addition to these cells, olfactory ensheathing cells(OEC) [45] and trans-differentiated mesenchymal stemcells [46] have also been identified to promote axonalregeneration and functional recovery at the site of a spinalcord injury. The olfactory system promotes the axonaloutgrowth into the CNS from the PNS due to the presenceof both peripheral and central tissues [47]. Thus the phe-notype of OEC is closer to the Schwann cells and has theproperties of both Schwann cells and Astrocytes [36]. The

transplantation of OEC enhances its migration [48] andsecretion of extracellular molecules type IV collagen,tropic factors such as vascular endothelial growth factor(VEGF), NGF and BDNF [36]. The OEC are found toreduce the secondary neuron apoptosis and the degree offunctional recovery after implantation of OEC is rapid ascompared to Schwann cells [36,49].

Although the stem cell therapy seems to be promising, itis very hard to control the cell proliferation and differen-tiation into three-dimensional architectures of tissues[50]. Moreover, the use of stem cells to repair spinal cordhas been reported to cause adverse side effects such asallodynia in unaffected forepaws of the rats [51]. Hence, itis a challenge to provide optimized conditions for con-trolled and adequate differentiation of transplanted stemcells in order to develop a safe stem cell-based therapy.Certain factors in stem cell-based therapy such as carcino-genicity and ethical concerns in the use of embryonic stemcells remain unanswered thereby limiting their use.

Delivery of growth factors to promote regenerationGrowth factor signaling play a major role in tissue repairprocess. In addition to maximizing the intrinsic regenera-tive potency of endogenous progenitor cells, bioactive fac-tors are also used to manipulate the differentiation andgrowth of exogenous stem cells [52]. The discovery ofNerve Growth Factor (NGF) has been more helpful tosupport ailing neurons by promoting nerve regeneration[53,54]. However, targeting and retaining the requiredconcentration of these factors at the site of injury is quitecomplicated phenomenon. The nerve guidance channelsprovide a conduit for the diffusion of these growth factorsand reduce the glial scar formation [16]. The sustainedrelease of lipoplexes (complexes of liposomes and oligo-nucleotides) has shown promise to repair nerve injury byexpressing neurotrophic factor [55,56]. Moreover, restor-ing the activity of such proteins is pretty challenging.Although Poly (D,L-lactic-co-glycolic acid) (PLGA) micro-spheres has been demonstrated as a potential carrier ofgrowth factors, the main concern is the inactivation ofproteins due to the release of acidic products. This prob-lem can be overcome by using polyphosphoesters as adelivery vehicle [57]. Recently, Sun et al. targeted a colla-gen binding domain-nerve growth factor β (CBD-NGF β)to nerve extracellular matrix collagen to restore theperipheral nerve function in rat sciatic nerves [58]. Theyhave confirmed the functional recovery by performingwalking track, histological and electrophysiological anal-ysis. These discoveries of regenerative capacity in adultcentral nervous system hold promise to neural victims forcomplete recovery [58]. Currently electrospun nanofi-brous scaffold become a successful and safe delivery vehi-cle since it can simultaneously act as a scaffold and

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improve contact guidance while delivering the bioactivefactors in controlled and sustained way [59].

Biomaterials in Nerve RegenerationNatural polymers (chitosan, chitin, collagen, gelatin, algi-nate), synthetic non-degradable polymers (silicone), syn-thetic biodegradable polymers such as PLGA, poly (ε-caprolactone (PCL), poly L-lactic acid (PLLA) and con-ducting polymers (polypyrrole, polyaniline) have beenused in various nerve regeneration approaches. An idealnerve conduit should be thin, flexible, porous, biocom-patible, biodegradable, compliant, neuroinductive, neu-roconductive and with appropriate surface andmechanical properties [14]. Although these biomaterialspromise to fulfill some of the above stated criteria, theyhave some demerits which have to be overcome to meetthe specific tissue engineering applications. For example,a scaffold made from non-degradable materials should beavoided to prevent the chronic inflammation and com-pression of nerve over time and therefore it is preferableto use biodegradable materials [26]. Even in biodegrada-ble materials, surface erosion is desired over bulk erosion,since it permits the scaffolds to retain their structural sta-bility for a longer time after implantation. Therefore, asurface eroding polymeric scaffold is expected to providebetter contact guidance cues continuously for nerve regen-eration. This fact is reinforced by a recent report on the useof poly (glycerol sebacate) (PGS) as a nerve guide materialdue to its surface erodible and elastomeric properties [60].In general, researchers have attempted to improve theneural scaffold properties by several novel fabricatingtechniques such as polymer blending and electrospin-ning, incorporating nerve growth factors in the scaffold[61], and improving the wettability of the scaffold surfaceby surface modifications. Table 1[62-78] summarizes alist of biomaterials and techniques that have been used topromote nerve regeneration.

One of the most important properties needed for success-ful graft uptake by host tissues is mechanical stability andcompatibility of the scaffolds. For nerve tissue engineer-ing, the scaffold should be pliable, harmless to the sur-rounding tissues, resist structural collapse duringimplantation [79] which may lead to necrosis and inflam-mation. Many approaches have been attempted toimprove the properties of common biomaterials to makethem suitable for neural tissue engineering. A highly flex-ible PLGA scaffold was developed by microbraidingmethod to improve its flexibility and porosity [75]. Suita-bility of PLLA porous conduits fabricated by extrusiontechnique has been evaluated in vivo for scaffold applica-tions [80-82]. Biomechanical properties of electrospunPCL scaffolds were improved by thermal treatment whileretaining the structural stability, gross appearance, poros-ity and fiber diameter [66].

Polymer blending offers one of the most successful meth-ods to develop a suitable scaffold with all preferred prop-erties for specific tissue engineering applications [83,84].For example, a polymer blend of the brittle-natured PLGAand a soft, elastic polymer such as polyurethane or poly(ethyleneglycol) (PEG) have been found to exhibit muchgreater elasticity than PLGA itself [85]. Chitosan-gelatincomposite films show improved mechanical property andnerve cell affinity due to its softness and elastic properties[84]. Fabrication of biodegradable nerve guidance chan-nels based on chitin/chitosan has been carried out forimprovements in nerve tissue engineering [86]. An artifi-cial nerve graft of chitosan/polyglycolic acid (PGA) blendhas been used to bridge a 30 mm sciatic nerve defect in alarge animal model [87]. Repairing long-term delayedperipheral nerve defects is clinically very challengingwhich includes number of various factors such as availa-bility of surviving Schwann cells, worsening of growthpermissive environment by disintegration of Schwann cellbasement membrane [88]. Jiao et al. attempted to bridgethe long term delayed defects of rat sciatic nerves usingbiodegradable chitosan-PGA graft and measured the func-tional recovery using histological and electrophysiologi-cal assays [88]. They have observed very few functional(regenerated) nerve fibers and poor growth support indelayed repair groups can improve the potency of chi-tosan/PGA grafts in delayed repair [88].

Cell adhesion property of a scaffold mainly depends on itssurface characteristics such as charge density and wettabil-ity [66,89]. Most of the synthetic biodegradable materialsare hydrophobic (PLGA, PCL, PHB etc.,) which limitstheir use as tissue engineering scaffolds. Thus surfacemodification either by coating of the surface with ECMproteins such as laminin, fibronectin, collagen or byincorporating specific adhesion peptide sequences likeRGD and IKVAV, YIGSR can induce hydrophilicity whichin turn improves the cell adhesion property of the scaf-fold. Currently the cell adhesion property of the materialssuch as methyl cellulose [90], alginate [91], poly (hydrox-yethyl methacrylate) (PHEMA) [92], poly (hydroxybu-tyrate) (PHB) [65] has been improved by modifying theirsurface with specific peptide sequences.

Different approaches have been used to improve the bio-compatibility. The performance of neural implants hasbeen improved by using the layer-by-layer (LbL) tech-nique [93]. A photochemical method has been employedto make PHEMA bound neural growth factor (NGF) morebioactive [94].

Polymeric Scaffolds as Extracellular Matrix AnaloguesScaffolding is a temporary framework used to supportcells in the construction or repair of tissue. Surface chem-istry of scaffold materials is considered to be the most

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important parameter in tissue engineering [95]. The extra-cellular matrix (ECM) in biological systems holds the cellstogether and provides a medium for the cells to interactand migrate [96]. Thus it is desirable that the syntheticscaffold mimics the ECM in promoting cell adhesion, pro-liferation, and differentiation in vitro and in vivo [66,97].Two dimensional tissue cultures as the name suggestsoffer only a monolayer of cells as opposed to the three-dimensional nature of tissue in organisms [98]. Hence it

is an inadequate model for complex cellular interactionsand is prone to hydrodynamic damage in bioreactors[98]. Three-dimensional tissue cultures have receivedmuch consideration than two dimensional culturesbecause of their superior hydrodynamic protection,higher surface area per unit volume, better cell-cell inter-action and improved regeneration of the injured tissue[99]. Improved functional recovery and formation of neu-ral networks to bridge the gap following spinal cord injury

Table 1: Modified biomaterials attempted to promote nerve regeneration.

Biomaterials Modification/Method of fabrication Improved Properties Ref

Star-Poly(ethylene glycol) Incorporation of polysaccharide (Heparin) Tunable physical and mechanical properties to adopt specific tissue requirements

[62]

Chitosan Modified with (γ-glycidoxypropyltrimethoxysilane (GPTMS)

Mechanical strength [63]

Poly(sialic acid) Hydrogel modified with adsorbed poly-L-lysine or poly-L-ornithine or laminin or collagen

Mechano compatibility; Cell adhesive property [64]

Poly(β-hydroxybutyrate) Sheets impregnated with extracellular matrix molecules

Cell adhesion and proliferation [65]

Poly(ε-caprolactone) Electrospinning and Thermal fiber bonding Mechanical strength [66]

Poly(ε-caprolactone) Aligned fibers by Electrospinning Contact guidance [67]

Poly(lactic-co-glycolic acid) Modified immersion precipitation method Selective permeability; Hydrophilicity [68]

Poly(D, L-lactide-co-ε-caprolactone) [PDLLA/CL]

PPy coating substrate and PPy nanoparticle/PDLLA/CL composite

Electrical cue for multitude of cell functions [69]

Chitosan Polylysine-functionalised thermoresponsive chitosan hydrogel

Injectable scaffold; Mechano compatibility; Surface property (wettability, charge density)

[70]

Poly(ε-caprolactone) Electrospinning (Polymer blending with collagen) Biological property (schwann cell adhesion, migration and differentiation)

[71]

Collagen Hydrogel crosslinked with YIGSR peptide modified dendrimers

Biological function (promote the growth of corneal epithelial cells and

neurite outgrowth)

[72]

Poly(glycerol-sebacate) Replica molding Micropatterned substrates; Flexibility; Surface degradable; Strong contact guidance response

[73]

Poly(lactic-co-glycolic acid) Microbraiding method Flexibility; Porosity [75]

Poly(D,L-lactide-co-glycolide) Low pressure injection molding Porosity; Longitudinally aligned channels; Mimics the geometry of native nerves

[76]

Poly(2-hydroxyethyl methacrylate)

Fiber templating technique Oriented scaffold; Physical characteristics similar to soft tissue.

[77]

Poly(2-hydroxyethyl methacrylate)

Liquid-liquid centrifugal casting Mechanical property similar to spinal cord [78]

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has been reported in the transplantation of stem cells onpolymeric scaffolds than the transplantation of stem cellalone [30,100].

Geometric Cues - Structure of ScaffoldsWell defined nanostructured topographical cues such asgrooves, ridges, pores, nodes can influence cell-substrateinteraction by promoting cell adhesion, migration, prolif-eration and differentiation to new tissue [101]. For exam-ple nanopatterned gratings on poly(methyl methacrylate)(PMMA) and poly(dimethylsiloxane) (PDMS) were usedto induce alignment and elongation of smooth musclecells [101]. There are different scaffold fabrication tech-niques such as solvent casting, particulate leaching, meltmolding etc., to fabricate scaffolds of various geometrieswith the desired porosity and surface area for cell scaffold-ing [102]. Among the different forms of scaffold (nanofib-ers, sintered matrix, nanofoams, hydrogels, nanotubes,etc.,) hydrogels and nanofibers have been extensivelyinvestigated for use as a scaffold in neural regeneration.

In recent years, hydrogels have received considerableattention as a suitable scaffold material in neural tissueengineering [103,104]. Hydrogels provide appropriatechemical, mechanical and spatial microenvironment akinto the natural ECM to support the neurite extension forcell proliferation, differentiation and axon extension[95,105]. Moreover they are biocompatible and possesssimilar mechanical properties to soft tissue, low interfacialtension, and are good injectable scaffolds. A biocompati-ble polymeric hydrogel has been shown to induce recon-struction of the rat spinal cord after chronic compression-produced injury [55,106]. Neuroinductive and neurocon-ductive properties of a biocompatible heterogeneous poly[N-(2-hydroxypropyl) methacrylamide] (PHPMA) hydro-gel have been used extensively to repair tissue defects inthe central nervous system by promoting the formation ofa tissue matrix and axonal growth [107]. Poly (2-hydrox-yethyl methacrylate) (PHEMA) hydrogels have alsoproved to be useful in neural tissue engineering applica-tions [77]. Cells get introduced easily into the liquid pre-cursors of the gel due to its smaller mesh size. Productionof cell-based hydrogel polymer constructs has been envi-sioned for tissue replacement in the central nervous sys-tem with combined physico-chemical properties such asbiocompatibility, stability, porosity and hydrophilicityalong with biological recognition such as expression ofbiospecific surface receptors and synthesis of bioactivemolecules [108]. It has been reported that the use of pol-yethylene glycol (PEG) hydrogel as a cell carrier supportedthe neural precursor cell survival, expansion and differen-tiation in culture [109]. The evaluation of growth proper-ties of Schwann cells on chitosan proved itsbiocompatibility [110]. Implantation of three-dimen-sional polymeric hydrogel into the site of injury has been

attempted to enhance the axonal recovery [111]. Three-dimensional polylysine-functionalized polysaccharidehydrogel system promises to be a good scaffolding mate-rial for neural tissue engineering [70,112]. Three-dimen-sional peptide channels within hyaluronan (HA)hydrogel matrix modified with S-2-nitribenzyl cysteine(HA-SNBC) is expected to serve as a temporary scaffoldfor guided axonal regeneration in vivo [113]. Among thevarious polymeric hydrogel tubes that have been designedand studied for their suitability as neural scaffolds, rein-forced coil tube expressed excellent mechanical propertiesequivalent to soft tissues and they supported neural out-growth. Also it was observed that the conducting velocityof the construct exactly matched that of native axons[114].

Nanofiber Properties & Electrospinning TechnologyNanofiber seems to be most promising substrate in tissueengineering applications due to its resemblance to nativeextracellular matrix [67,115,116]. Nanofiber is a broadphrase generally referring to fibrous structures with adiameter less than 1 micron. Figure 2 shows the scanningelectron micrograph of electrospun PLGA-Polyanilinenanofibers with smooth defect-free morphology for neu-ral tissue engineering. Extraordinary mechanical strengthand high surface area to volume ratio makes the nanofibermore suitable for neural tissue engineering [10,117].Porous polymeric nanofibrous scaffold using biodegrada-ble poly (L-lactic acid) (PLLA) fabricated by liquid-liquidphase separation method resembles ECM of natural colla-gen to support neuron differentiation and neurite out-growth [22]. However, it is very difficult to maintain thefiber diameter and alignment using this technique. Like-wise, various techniques have been reported to develop

Electrospun PLGA-PANi nanofibersFigure 2Electrospun PLGA-PANi nanofibers.

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nanofibers namely, template synthesis, phase separation,self-assembly, drawing and electrospinning [115]. Amongthese techniques, electrospinning offers more advantagesdue to its ease of fabrication. Nanofibrous conduit com-prising poly (D,L-lactide-co-glycolide) and poly(ε-caprol-actone) (PCL/PLGA) were found to promote nerveregeneration across 10 mm nerve gap in rat sciatic nerve[118]. PLGA random nano and microfibers, alignedmicrofibers and films were investigated for C17.2 neuralstem cell culture and recognized the differentiation ofneurons along the fiber direction [119].

PLLA nanofibrous scaffolds were developed via electros-pinning and found to support neural stem cell (NSC)adhesion, outgrowth and differentiation [21]. Suitabilityof aligned electrospun PLLA nanofibers compared withrandom nanofibers was evaluated for neural tissue engi-neering in terms of their fiber alignment and dimension[21]. The aligned nanofibers were found to support theorientation of cells and improve the neurite outgrowthand contact guidance [21,67]. Based on the experimentalresults, this study recommended the aligned PLLA nanofi-brous scaffold as a potential cell carrier in neural tissueengineering [21,119]. The parameters such as viscosity,conductivity, surface tension of polymer solution, appliedelectric potential, flow rate, and distance between the elec-trodes are to be optimized while carrying on the electros-pinning process [120]. It is also observed that theorientation of fiber became disordered at the top layer ofthe electrospun mesh when the collecting time was longerthan thirty minutes due to the residual charges on the col-lecting fibers [21]. The desirable properties of electrospunnanofiber scaffolds seem to offer a promising alternativetowards the treatment of neural defects. Aligned electro-spun collagen/PCL fibers supports cell proliferation, glialmigration, orientation of neurite outgrowth, suggested itssuitability as nerve implants [71]. Aligned electrospunPCL fibers have been found to up regulate specific genessuch as P0 and down regulate NCAM - 1 on culturedSchwann cells thereby promoting the Schwann cell matu-ration [67]. Significantly higher Schwann cell migrationand neurite outgrowth was observed on uniaxially alignedfibers of poly (acrylonitrile-co-methylacrylate) (PAN-MA)developed by electrospinning than on random fibers[121].

Other ApproachesFabrication of a multi-channel scaffold using injectionmolding with solvent evaporation technique has beendemonstrated to promote spinal cord axon regeneration[122]. A new facile method named 'fiber stimulating tech-nique' used to fabricate oriented PHEMA scaffolds suc-cessfully for neural tissue engineering, promises to bemore effective and reproducible [77]. Melt compressionand melt extrusion are also considered to be viable tech-

niques to prepare nerve guides [14]. Innovative fabrica-tion techniques such as wire mesh method and mandreladhesion method are used to prepare multi-channel bio-degradable nerve guides without the requirement of com-plex instrumentation, acidic conditions or exposure toextreme temperatures [123]. Designing of biodegradablePLGA hollow fiber by wet phase inversion technique hasbeen attempted for the development of nerve tract guid-ance conduit [13]. Micropatterning is a novel patterningtechnique for biodegradable polymers and is reported toenhance peripheral nerve regeneration by controlling thealignment of Schwann cells [124-126]. Fabrication tech-niques continue to evolve novel routes to provide themost suitable nanostructure topography for adequateneural growth.

Electrical CuesHuman body responds to electrical fields and the keycomponent of neural communication in the body is theaction potential generated at the synapse. This impliesthat an ideal neural scaffold should also possess electricalconductivity to promote neurite outgrowth and therebyenhance nerve regeneration in culture. The use of electri-cally conducting polymers in biomedical applications hasbecome more attractive due to its tailor-made specificities[127]. Polypyrrole (Ppy), a well-known conducting poly-mer used in biomedical applications has been found toenhance the nerve regeneration by electrical stimulation[128,129]. Moreover, the antioxidant property of polypyr-role and polyaniline makes them more attractive sub-strates for tissue engineering applications [130,131] asthey could scavenge any free radicals at the site of injuryminimizing scar formation which is a bane of neuralregeneration. The structures of polythiophene, poly-aniline and polypyrrole are shown in figure 3.

A comparative study was made on the efficacy of two dif-ferent laminin fragments p20 and p31 as dopants in con-ducting polypyrrole surfaces for in vitro growth ofneurons. The results indicated that p20 as dopant sup-ported the highest neuronal density than p31 dopant[132]. Conducting Ppy/PDLLA/PCL composites havebeen implanted to bridge the gap of 8 mm in rat sciaticnerves and shown to promote the nerve cell proliferationand axon regeneration using electrical cues [69]. The ratswere gradually recovered the mobility in operated limbover the period of 2 months [69]. Moreover, the immuno-histological analysis and transmission electron micros-copy of harvested implants demonstrated the presence ofnewly formed myelinated axons and Schwann cells simi-lar to that of native nerve [69]. Recently the cell adhesionproperty of polypyrrole has been improved by chemicalconjugation of a functionalized carboxylic acid groupwith RGD peptide [133]. The suitability of poly ethylenedioxythiophene (PEDOT) as a biomaterial was evaluated

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by studying the adhesion and proliferation of epithelialcells and was demonstrated that the electroactive substratefavors cell adhesion [134]. Though the mechanism inwhich electrical stimulation promotes nerve regenerationis not clearly understood, several possible hypothesis havebeen postulated to elucidate their role in nerve regenera-tion. Some of the plausible reasons include electro-phoretic redistribution of cell surface receptors, activategrowth controlling transport process and altered adsorp-tion of adhesive proteins [135]. The later hypothesis hasbeen proved by stimulating the adsorption of fibronectin(ECM adhesive glycoprotein) from serum to polypyrrolesurface via electrical stimulation, which subsequentlyfacilitate the neuronal attachment and neurite outgrowth[128].

However, these conducting polymers are non-biodegrad-able and questions on their safety in biological systemshave delayed their wide-spread use in neural conduits.

Future ProspectiveMany Challenges still remain unwrapped. Though theresearchers have found different strategies to achieve thefunctional recovery to some extent, regaining the maximalor full function remains unexplored. There are someissues listed below have to be addressed in future; (1) thefirst inescapable conclusion arising over various reportson nerve tissue engineering by super positioning of allthese approaches is crucial for promoting the neuralregeneration on multiple levels (2) the probable hazardsof long term usage of such novel biomaterials on humanhealth yet to be revealed (3) the need for novel Biomate-rials and approaches has to be established in order to treatthe delayed nerve injuries in patients who have neurolog-ical disorders.

ConclusionAn ideal nerve conduit requires a suitable porous, bio-compatible, biodegradable, neuroconductive, neuroin-ductive, infection resistant, compliant three-dimensional

biomaterial scaffolds. The engineered construct shouldalso mimic the ECM architecture and porosity, desirablefor cell attachment and other vital functions. Biomedicalnanotechnology, electrospinning techniques and tissueengineering methods give us exciting insights to thedesign of a scaffold with good electrical, mechanical, bio-logical properties and compliance match closely resem-bling the native ECM. Such scaffolds can also avoidinfections, multiple surgeries and additional cost to thepatient. An array of methods has been used for polymerscaffold preparation but electrospinning scores high dueto its ease of operation, better control of fiber propertiesand desirable results. Lot of synthetic biodegradable poly-mers has been used till date but at the same time sufferfrom the demerits of release of acidic degradation prod-ucts, hydrophobicity, poor processability and loss ofmechanical properties. Also they support elongation andpartial collapse of nerves. Hydrogels mimic soft tissueproperties but are very difficult to sterilize and handle dueto their fragile nature. A new strategy using polymers likepolypyrrole, polyaniline etc., having conducting proper-ties are being investigated for neural tissue engineering tostimulate neurite extension. However, the biocompatibil-ity of these polymers has not been conclusively proved tillnow. It is seen that though different classes of biomateri-als are available, no single material is enough to improvethe scaffold properties for nerve regeneration. Hence, themajor challenge in developing a scaffold lies primarily inthe choice of a blend of biomaterials with the correct com-bination of properties. The field is still wide open todesign the most appropriate polymer scaffold with all thevital conditions and properties for effective neural appli-cations in vivo.

Competing interestsThe authors declare that they have no competing interests.

Authors' contributionsAll authors have equal contribution in the preparation ofthe manuscript.

Chemical structure of conducting polymersFigure 3Chemical structure of conducting polymers.

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130. Kilmartin PA, Gizdavic-Nikolaidis M, Zujovic Z, Travas-Sejdic J, Bow-maker GA, Cooney RP: Free radical scavenging and antioxidantproperties of conducting polymer examined using EPR andNMR spectroscopies. Synth Met 2005, 153:153-156.

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134. Valle LJ, Aradilla D, Oliver R, Sepulcre F, Gamez A, Armelin E, AlemanC, Estrany F: Cellular adhesion and proliferation on poly(3,4-ethylenedioxythiophene): Benefits in the electroactivity ofthe conducting polymer. European Polymer Journal 2007,43:2342-2349.

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Fabrication of uniaxially aligned 3D electrospun scaffolds for neural regeneration

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Biomed. Mater. 6 (2011) 025004 (10pp) doi:10.1088/1748-6041/6/2/025004

Fabrication of uniaxially aligned 3Delectrospun scaffolds for neuralregenerationAnuradha Subramanian1,2, Uma Maheswari Krishnan1,2

and Swaminathan Sethuraman1,2,3

1 Center for Nanotechnology & Advanced Biomaterials, SASTRA University, Thanjavur 613 401,Tamil Nadu, India2 School of Chemical & Biotechnology, SASTRA University, Thanjavur 613 401, Tamil Nadu, India

E-mail: swami@sastra.edu

Received 28 October 2010Accepted for publication 18 January 2011Published 7 February 2011Online at stacks.iop.org/BMM/6/025004

AbstractNanofibrous scaffolds are very promising physical guidance substrates for regenerating nervesto traverse larger nerve gaps. In this study, we have attempted to develop 2D random and 3Dlongitudinally oriented nanofibers of poly(lactide-co-glycolide) (PLGA) by the modifiedelectrospinning process and characterized the surface morphology, mechanical properties,porosity, degradation and wettability. The orientation of aligned fibers was optimized byvarying the speed of the rotating mandrel in the electrospinning process. The mean diameterof random PLGA nanofibers was 197 ± 72 nm, whereas that of the aligned PLGA fiber was187 ± 121 nm. The pore size of aligned PLGA nanofibers (3.5 ± 1.1 μm) was significantlylower than their respective random nanofibers (8.0 ± 2.0 μm) (p < 0.05). However, thepercentage porosity of both scaffolds was comparable (p > 0.05). The mass loss percentageand molecular weight loss percentage due to degradation was higher in random PLGA fiberswhen compared to aligned PLGA after 5 weeks (p < 0.05). The tensile strength and Young’smodulus of random PLGA fibers were significantly higher than those of the aligned PLGAnanofibers (p < 0.05). Both random and longitudinally aligned scaffolds were used for thein vitro culture of Schwann cells. Morphology and cell proliferation results demonstrated thatthe aligned fibers assist the direction of Schwann cells and a better proliferation rate than theirrandom fibers. The results confirmed that aligned nanofibers have better deformability, slowdegradation, comparable porosity and orientation cues than random nanofibers. Hence thelongitudinally aligned nanofibers may be ideal scaffolds for nerve regeneration.

(Some figures in this article are in colour only in the electronic version)

1. Introduction

Nerve injury is a serious clinical complication which requiresimmediate intervention to restore the functions since thematured neurons do not undergo cell division. In a smallnerve defect the axons can bridge the gap through theformation of fibrin cable and Bungner bands [1]. However,in larger lesions, denuded peripheral nervous system axons

3 Author to whom any correspondence should be addressed.

are unable to regenerate inherently [2]. The current goldstandard approach of using autografts to bridge larger gapshas limitations such as painful neuroma formation, functionalloss at the donor site, compliance mismatch, etc [3]. Hence,development of artificial nerve grafts offers a viable alternativeto the conventional autografts to treat neural defects [4].

After peripheral nerve injury, glial cells migrate towardthe distal end of smaller gaps to direct the neurite outgrowth[5]. In larger nerve gaps, the fibrin cable and Bungner bandformation are compromised [6]. Therefore, it is desirable

1748-6041/11/025004+10$33.00 1 © 2011 IOP Publishing Ltd Printed in the UK

Biomed. Mater. 6 (2011) 025004 A Subramanian et al

to treat the larger nerve gaps using a scaffold having well-defined architecture to orient the glial cells longitudinallytoward the distal end while acting as physical template tobridge the gap. Thus the proposed scaffold should be tubularwith desired topographical cue (uniaxial alignment fibers),compliant match to the native nerves in order to orient theglial cells toward the distal end [7].

Electrospinning is a well-known technique for thefabrication of nanofibrous scaffolds for tissue engineeringapplications [8]. Electrospun nanofibrous scaffolds havereceived much attention in tissue engineering due tothe resemblance with collagen fibers in extracellularmatrix, high surface area to volume ratio, high porosity,and tailorable mechanical and degradable properties[7, 9–11]. Recent studies have demonstrated that the directionof the aligned nanofiber scaffold determines the path forneurite extension [12]. Poly(lactide-co-glycolide) (PLGA)random nanofibers and microfibers, aligned microfibersand films were investigated for the C17.2 neural stemcell culture and the differentiation of neurons along thefiber direction toward the treatment of neural defects [13]was recognized. Aligned electrospun collagen/(poly(ε-caprolactone)(PCL) fibers support cell proliferation, glialmigration, and orientation of neurite outgrowth, suggestingtheir suitability as nerve implants [7]. Aligned electrospunPCL fibers have been found to up-regulate specific genessuch as P0 and down-regulate NCAM-1 in cultured Schwanncells thereby promoting the Schwann cell maturation [12].Significantly higher Schwann cell migration and neuriteoutgrowth were observed in uniaxially aligned fibers ofpoly(acrylonitrile-co-methylacrylate) (PAN-MA) developedby electrospinning than in random fibers [14].

Recently a tubular scaffold has been proved as a promisingsubstrate for use in tissue engineering due to its uniformthickness, properties and mechanical strength all over the tube[15]. However, the orientation of electrospun nanofibers isa challenge due to their intricate three-dimensional whippingmotion of a jet [16]. A high-speed rotating drum has beenused to align the fibers circumferentially [17]. However, thecomplaint mismatch of the graft and partial alignment of fiberslimit their use in neural regeneration [18]. A thin disk with asharp edge collector has been found to improve the orientationof fibers [19]. However, fibers were collected only at the edgeof the disk. The use of auxiliary electrodes as a collector has aprofound effect on the orientation of fibers [18]. A charged jetexperiences a strong stretching force between the electrodeswhen it reaches the collector. This facilitates the formation ofsuspended arrays of aligned fibers across the gap. However,developing a thicker layer and longer length of aligned fibersis very difficult [18].

Among the synthetic biodegradable polymers, PLGAhas been chosen for this study as it is most extensivelyinvestigated for neural regeneration and also because of itsestablished biocompatible and biodegradable properties [13].In addition to this, PLGA has shown better performancein terms of cell viability, apoptotic activity, mitochondrialmetabolic activity, proliferation and differentiation rate ofHiB5 cells than poly(L-lactic acid) (PLLA) and (poly(lactide-co-ε-caprolactone) (PLCL) [20]. In this study, we have

proposed a simple electrospinning setup to obtain a tubularstructure with longitudinally aligned nanofibers. Using suchsetup, we have developed a longitudinally aligned PLGAnanofibrous scaffold. The properties of a 3D alignedscaffold such as porosity, mechanical strength, wettability andbiodegradability were investigated and compared with respectto a 2D random mat to explore its applications on axonalregeneration. The rat Schwann cell adhesion and proliferationon both 2D and 3D scaffolds were investigated to investigate3D scaffold applications in neural tissue engineering.

2. Materials and methods

2.1. Materials

PLGA (Mw-118 kDa) was purchased from LakeshoreBiomaterials, Birmingham, USA. Dichloromethane (DCM)and N,N-dimethyl formamide (DMF) were purchased fromMerck, India and used without any further purification.

2.2. Fabrication of random and aligned nanofibrous scaffolds

The 2D random and 3D aligned PLGA nanofibers weredeveloped through electrospinning under optimum conditions.Briefly, 12% (w/v) of PLGA was dissolved in a mixtureof DCM and DMF (9:1). This solution was loaded into a5 mL glass syringe and the flow rate of 0.001 mL min−1

was controlled by syringe pump (KD Scientific 200, USA). Ahigh voltage (25 kV) was applied using a high-voltage powersupply (Zeonics, India). The 2D random fibers were collectedon aluminum foil fixed onto a grounded metal collector.

The 3D axially aligned nanofibers were collected usinga grounded rotating mandrel (3 mm in diameter) with smallinsulating gaps (figure 1). Briefly, a small insulating sleeve(1.5 cm in length) was introduced into the rotating metalrod (4 mm in diameter) to create an alternating metal andinsulating part, and the metal rod was connected with ground.The insulating gap between the grounded metal rod behavesas an auxillary electrode which allows the deposition oflongitudinally oriented fibers onto the insulating portion ofthe rod. The speed of the mandrel was varied from 1000 to2500 rpm using a motor (Remi, India) to improve the fiberorientation, while other parameters such as concentration,flow rate, applied voltage and tip–target distance weremaintained at 12% (w/v), 0.001 mL min−1, 25 kV and12 cm respectively. These electrospun nanofibers were storedin a vacuum desiccator for further characterization.

2.3. Surface morphology

The surface morphology, diameter of the fibers and orientationof aligned fibers were determined using a scanning electronmicroscope (SEM, JEOL FE-SEM 6701F, Japan). Theelectrospun nanofibers were sputter coated with platinum andexamined under SEM at an accelerating voltage of 3 kV. Thedistribution of fiber orientation was analyzed by calculating therelative angle of individual fibers from +90◦ to −90◦ where 0◦

was taken as the axial direction.

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Biomed. Mater. 6 (2011) 025004 A Subramanian et al

(A)

(B)

Figure 1. (A) Electrospinning setup used for the fabrication of thelongitudinally oriented tubular scaffold. (B) Tubular scaffold.

2.4. Contact angle measurement

The water contact angle of both 2D random and 3D alignednanofibers was measured using a goniometer (Rame-Hart,USA). Approximately 5 μL of deionized water was carefullyplaced on to the surface of electrospun fiber and contact anglevalues were calculated (n = 5).

2.5. Mechanical testing

Mechanical properties of 2D random and longitudinallyaligned 3D PLGA scaffolds were determined using an Instronuniaxial tensile testing machine (INSTRON, USA). The 2Dspecimens were cut into dumbbell shape with a width of10 mm, gauge length of 50 mm and thickness ranging from0.17 to 0.2 mm (n = 6). Three-dimensional samples of 50 mmlength and 5 mm outer diameter were taken for measurement(n = 6). Load deformation data were recorded at a deformingspeed of 1 mm s−1, and the stress–strain curve of nanofiberswas constructed from the load deformation curve.

2.6. Porosity measurement

The porosity and distribution of pore size of both 2D randommats and 3D aligned scaffolds were measured using mercuryintrusion porosimeter (Micromeritics Autopore IV, USA). The2D random PLGA scaffold (0.2 mm thickness) was cut into2 cm × 5 cm rectangular shapes and weighed (n = 3). Forthe 3D aligned scaffold, samples of 3 mm inner diameter and5 cm length were taken for analysis (n = 3). The pressure was

varied from 0.1 to 50 psi with the equilibration time of 60 sfor each intermediate data point. The pore size and porosityof all scaffolds were determined from the amount of mercurythat penetrated into the sample at different pressures.

2.7. In vitro degradation study

The 2D random electrospun membranes were cut into a squareshape with dimensions of 40 mm × 40 mm for in vitrodegradation (n = 3). The 3D aligned fibers with 3 mm innerdiameter and 20 mm length were used for degradation study(n = 3). The weight of the 2D random and 3D alignedfibers was measured and the fibers were immersed in30 mL phosphate buffered saline (PBS) at pH 7.4 and 37 ◦Cwith constant shaking using a water bath shaker (LABLINE,India), and the PBS solution was changed every 3 days. Atpredetermined time points (1, 2, 3, 4 and 5 weeks) the sampleswere removed and rinsed with deionized water, dabbed withtissue paper to remove surface water and dried under vacuumfor 1 week. The dry weights of both the 2D random and 3Daligned nanofibers were recorded and the surface morphologyof degraded sample was characterized using a SEM. Theweight loss percentage of each sample was calculated by thefollowing formula:

Weight loss percentage

= initial weight of sample − dry weight of degraded sample

initial weight of sample×100.

The molecular weight loss of both polymeric scaffolds atdifferent time points was characterized using gel permeationchromatography (GPC, Agilent 1200 Series, USA) byinjecting 50 μL of each sample. The samples were dissolvedin THF to make a 1% solution and eluted at a flow rate of1.5 mL min−1 through a size exclusion column calibratedusing a polystyrene standard.

2.8. Cell seeding

Cell adhesion and proliferation on the surface of the randomand aligned PLGA nanofibers were evaluated to access theviability of Schwann cells. The rat Schwann cell line (CRL-2765) was purchased from ATCC, USA, and was cultured inDMEM (ATCC, USA) supplemented with 10% FBS (Gibbco,USA) and 1% penicillin/streptomycin (Himedia, India) andmaintained at 37 ◦C and 5% CO2. Prior to cell seeding,each side of the random and aligned PLGA nanofiber matwas sterilized by exposure to UV light for 1 h. Five thousandcells were seeded on each scaffold to study the cell adhesionat various time points and twenty five thousand Schwann cellswere seeded to evaluate the cell proliferation.

2.9. Cell adhesion and proliferation

After the first and third days, the scaffolds were rinsedwith PBS twice and fixed with 2% glutaraldehyde for 1 h.The scaffolds were further rinsed with distilled water anddehydrated with graded concentrations of alcohol (50%, 75%,95% and 100%) and were air dried. The dried specimens were

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Biomed. Mater. 6 (2011) 025004 A Subramanian et al

sputter coated with gold and observed under a SEM with anaccelerating voltage of 3 kV.

The cell proliferation on both aligned and randomscaffolds was determined at different time points 1, 3, 5and 7 days using a colorimetric MTS assay (CellTiter 96 R©

AQueous One solution, Promega, USA). At each time point,the scaffolds were washed with PBS twice and 1 mL of freshserum-free medium and 200 μL of MTS reagent were addedand incubated for 2 h in a CO2 incubator at 37 ◦C. After 2 h,250 μL of the SDS reagent was added to stop the reaction. Theabsorbance was measured at 490 nm using a Tecan microplatereader (Tecan infinite M200, Australia). The absorbance wasconverted into cell number using a standard curve. Tissueculture polystyrene (TCPS) was used as control.

2.10. Statistical analysis

All values were statistically analyzed and expressed as mean± standard deviation. One way analysis of variance (one-wayANOVA) was performed to assess the effect of two differentscaffolds on scaffold properties such as pore size, porosity,mechanical strength and wettability. Two-way analysis ofvariance (two-way ANOVA) was used to evaluate the effectof two different scaffolds and time points on the degradationrate as well as on the cell proliferation rate. In both cases thestatistical significance was evaluated at 95% confidence limits(p < 0.05). If statistically significant, a post-hoc Tukey testwas performed to determine which means were different fromthe others.

3. Results

3.1. Surface morphology of random and aligned nanofibers

In this study, we have systematically optimized the variousprocess parameters such as flow rate, applied voltage andtip–target distance, and the solution parameters (viscosity,molecular weight, solution conductivity) to obtain defect-free random fibers of PLGA. The surface morphology ofPLGA nanofibers (figure 2) was smooth and the diameter ofPLGA nanofibers was 197 ± 72 nm. For the fabricationof 3D aligned nanofibers, the speed of rotating mandrel wasoptimized for PLGA while keeping the other parameters suchas concentration, flow rate, needle gauge, applied voltage andtip–target distance constant and the speed of the mandrel wasvaried from 1000 to 2500 rpm for maximum orientation.

At 1000 rpm, the randomness of the fiber orientationwas apparent (figures 3(A1) and (A2)) indicating that1000 rpm was not sufficient to orient the fibers axially.With the increase in speed to 1500 rpm (figures 3(B1) and(B2)) most of the fibers were in a diagonal direction whichconfirmed that the speed influences the orientation of fibers.At 2000 rpm, most of the PLGA fibers showed high-order axialalignment (figures 3(C1) and (C2)). However, at higher speeds(2500 rpm) multiple breakages in the collected fibers wereobserved thereby disrupting the orientation (figures 3(D1)and (D2)). Therefore, 2000 rpm was selected as an idealspeed for PLGA to achieve maximum longitudinal orientation.The mean diameter of the aligned PLGA was found to be

Figure 2. Scanning electron micrograph of two-dimensionalrandom PLGA. Defect-free fibers with average diameter of 192 ±72 nm were formed.

Table 1. Mechanical properties of random and aligned electrospunscaffolds. Aligned PLGA scaffolds have significantly higherelongation and lower ultimate tensile strength and Young’s modulus(∗p < 0.05).

Ultimate tensile Young’s ElongationMaterials strength (MPa) modulus (MPa) (%)

Random PLGA 2.6 ± 0.3∗

134.5 ± 22.7∗

14.5 ± 4.6∗

Aligned PLGA 2.1 ± 0.4∗

40.4 ± 18.2∗

26.0 ± 9.6∗

187 ± 121 nm. The 2D random and 3D aligned fibers ofPLGA were used for further characterization.

3.2. Contact angle

To study the wettable property of the scaffold, we havemeasured the water contact angle using a goniometer. Therandom PLGA nanofibers showed the contact angle of about99.22 ± 3.6◦, whereas the contact angle of aligned fibers was112.40 ± 2.0◦.

3.3. Mechanical testing

The ultimate tensile strength, percentage elongation andYoung’s modulus were calculated and shown in table 1for 2D random and 3D aligned nanofibers. The randomPLGA nanofibers exhibited significantly higher ultimatetensile strength, higher Young’s modulus and lower elongationpercentage than aligned PLGA (p < 0.05). However, the lowerYoung’s modulus and higher percentage elongation of alignedPLGA proved its higher flexibility and elastic property thanrandom PLGA.

3.4. Porosity and pore size distribution

The porosity of 2D random and 3D aligned nanofibers wasdetermined using a mercury porosimeter (table 2). The

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Biomed. Mater. 6 (2011) 025004 A Subramanian et al

(A1)

(B1)

(C1)

(D1)

(A2)

(B2)

(C2)

(D2)

Figure 3. Scanning electron micrographs showing the orientation and histograms of the fiber angle of the PLGA nanofiber at variousrotation speeds (A) 1000 rpm; (B) 1500 rpm; (C) 2000 rpm; and (D) 2500 rpm.

porosity of the random and aligned scaffolds was compared(p > 0.05). Mercury porosimeter analysis demonstrated thatthe pore size of random PLGA (8.0 ± 2.0 μm) nanofiberswas significantly higher than that of the aligned PLGA(3.5 ± 1.1 μm) (p < 0.05).

3.5. In vitro degradation

The in vitro degradation of both 2D random and 3D alignedelectrospun fibers was carried out in PBS at pH 7.4 and at37 ◦C with constant shaking for 5 weeks. At the end of

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Biomed. Mater. 6 (2011) 025004 A Subramanian et al

Random aligned

(A1)

(A2)

(A3)

(A4)

(A5) (B5)

(B4)

(B3)

(B2)

(B1)

Figure 4. In vitro degradation of 2D random of PLGA after (A1) 1 week, (A2) 2 weeks, (A3) 3 weeks, (A4) 4 weeks and (A5) 5 weeks; and3D aligned PLGA nanofibers after (B1) 1 week, (B2) 2 weeks, (B3) 3 weeks, (B4) 4 weeks and (B5) 5 weeks.

each degradation period, the samples were characterized formorphological change using a SEM and percentage massloss and percentage molecular weight loss. There was no

apparent change in the morphology of 2D PLGA randomfibers up to 3 weeks (figures 4(A1)–(A3)). After 4 weeks,there was presence of tiny pores on the nanofiber surface

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Biomed. Mater. 6 (2011) 025004 A Subramanian et al

(A)

(B)

Figure 5. In vitro degradation of 2D random and 3D alignednanofibers: (A) percentage mass loss percentage and (B) molecularweight loss percentage.

Table 2. Percentage porosity, pore size distribution and contactangle for random and aligned electrospun scaffolds (

∗p < 0.05).

Electrospun Porosity Pore size Contact anglescaffolds (%) (μm) (◦)

Random PLGA 74.9 ± 10.3 8.3 ± 2.0∗

99.2 ± 3.6∗

Aligned PLGA 75.7 ± 3.2 3.5 ± 1.1∗

112.4 ± 2.0∗

(figures 4(A4) and (A5)) indicating the onset of degradation ofPLGA fibers. It was apparent that the pore size increased withthe time indicating the bulk degrading behavior of randomPLGA. Interestingly, aligned PLGA fibers did not show anymorphological change up to 4 weeks (figures 4(B1)–(B4)).Only after 5 weeks, there was a visible thinning of fibers(figure 4(B5)), confirming the slow degradation rate ofthe aligned PLGA scaffold. This was also confirmed bydetermining the mass loss percentage. The PLGA randomfibers showed significantly greater mass loss percentage(31.5%) after 5 weeks than the mass loss percentage of thealigned PLGA scaffold (p < 0.05) (figure 5(A)). At the endof 5 weeks, random PLGA nanofibers were found to havesignificantly higher molecular weight loss percentage whencompared to aligned PLGA nanofibers (p < 0.05) (figure 5(B)).

3.6. Cell adhesion and proliferation

The morphology of Schwann cells was observed using a SEMafter the first and third days. After 1 day cell culture, thecells on the aligned fibers (figure 6(A1)) follow the fiberdirection, whereas the random fibers (figure 6(B1)) showedrandomly oriented cells. After 3 days, cells on aligned

fibers (figure 6(A2)) acquired spread morphology and coveredthe maximum surface of the nanofibers than random fibers(figure 6(B2)). The cell proliferation of Schwann cells on bothrandom and aligned nanofibers was quantitatively evaluated.It was observed that significantly higher number of cells werefound on the aligned scaffold when compared to random andTCPS (control) (figure 7).

4. Discussion

Nerve tissue is composed of bundles of nerve fiberspacked together by connective tissue to transmit the signalunidirectionally [21]. Moreover, across a smaller nerve gap,the regeneration is promoted by the formation of a fibrincable which in turn orients the Schwann cells longitudinally[6]. Electrospinning has been used to develop longitudinallyoriented 3D nanofibers to bridge larger gaps, resemble thenative architecture of ECM and orient the axons across thelesion. In this study, we have optimized the process parameterssuch as flow rate, applied voltage and tip–target distance, andthe solution parameters such as viscosity, molecular weightand solution conductivity to fabricate defect-free randomand aligned fibers of PLGA. The choice of the solvent,concentration of polymers, flow rate, tip–target distance,and needle diameter will greatly influence the fiber size andmorphology [11].

Figure 1 illustrates the electrospinning setup (figure 1(A))used for fabricating a tubular scaffold (figure 1(B)) withlongitudinally oriented nanofibers. It is basically the sameas the usual setup except for the use of a rotating metalrod comprising a number of insulating gaps of 1.5 cm in itsmiddle as a collector. Such a collector is fabricated simplyby introducing an insulating sleeve into the metal rod. Afterapplying high voltage, the electric field in close proximity tothe collector was split into two portions pointing toward theopposite edges of the insulating gap [22]. It is assumed thatthe charged fibrous jet ejected from the needle experiencedelectrostatic forces. First electrostatic force is used to pullthe charged fibers toward the target. Due to Coulombicinteractions, the charged fiber in the vicinity of each electrodecreates another strongest electrostatic force used to stretchthe nanofibers across the insulating gap in the rod [23]. Themajor advantage of this setup is that the fibers deposited acrossthe gap can slowly lose their charge which can create anelectrostatic repulsion between the collected fibers and thedepositing fibers for few seconds. This repulsive force furtherimproves the orientation of fibers. Thereby we were fabricatedtubes made of longitudinally aligned fibers by varying thespeed of the rotating collector.

In this study, we have used a dynamic collector to generatehighly oriented fibers by varying the rotation rate. Similar tothe findings of Zussman et al, we observed that the rotatingspeed of the metal rod influenced the fiber orientation forPLGA [15, 24]. The axial deposition of fibers onto the rotatingrod is mainly through stretching of fibers across the insulatinggap. The wettability of the scaffold was increased in alignedrather than in random fibers. The increase in the contact angleof axial fibers may be due to the decreased pore size havingdirect influence on smoothness of the scaffolds.

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Biomed. Mater. 6 (2011) 025004 A Subramanian et al

(A1) (B1)

(B2)(A2)

Figure 6. Scanning electron micrograph of Schwann cells on scaffolds after (A1) first day on aligned PLGA, (B1) first day on randomPLGA, (A2) third day on aligned PLGA and (B2) third day on random PLGA.

The mechanical properties of 2D random and 3D axiallyaligned electrospun fibers were measured using a uniaxialtesting machine with a deforming speed of 1 mm s−1 atroom temperature to evaluate the effect of fiber alignment andmaterials on mechanical strength. The tensile strength of thePLGA of random fibers was significantly greater than that ofthe aligned PLGA fibers. This is because the entwined fibersin the random scaffold offer more surface resistance, therebyimproving the tensile strength than aligned fibers. However,the tensile strength of the aligned PLGA scaffold (2.1 MPa)closely matches with the tensile strength (1.4 MPa) of acellularrat peripheral nerve [25]. The significantly lower Young’smodulus and greater elongation percentage of aligned fibersdemonstrated the flexibility of aligned scaffolds than that ofrandom scaffolds [26].

Porosity is an important characteristic of the scaffoldsin tissue engineering since it provides structural space forcell lodging, exchange of nutrients and metabolic waste andangiogenesis [27]. Yang et al fabricated nanostructured porousPLLA scaffolds for nerve tissue engineering and observed thatrandom orientation increased the pore size in the nanofibersas compared to aligned fibers [28]. In the present study, asignificant reduction in the pore size was observed in axiallyaligned PLGA fibers as compared to random fibers. Thismay be attributed to the fact that the axial alignment improvesthe compactness of the fibers through the speed of the rod,

Figure 7. Number of Schwann cells on the surface of randomPLGA, aligned PLGA and TCPS (control) (∗p < 0.05).

thereby reducing the pore size [29]. Gupta et al reported thatthe 1–2 μm pore size of the scaffold was found to be desirablefor cell migration and nutrient flow [26]. The aligned scaffoldshowed suitable pore size (3.5 μm) for cell supporting capacitywhich is close to the range of 1–2 μm. The scaffold withlow degradation behavior and swelling property acquires moreattention to withstand nerve compression [26]. In this study,we have observed the influence of longitudinal orientation onthe manipulation of degradation behavior. The significantreduction in the degradation in axial fibers rather than in

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Biomed. Mater. 6 (2011) 025004 A Subramanian et al

random 2D fibers may be due to the decreased pore size ofthe aligned fiber which decreases the diffusion rate of PBSthereby controlling the degradation rate.

Schwann cells are the glial cells present in the PNSthat surround the neuronal axons to form a myelin sheathand helps in aiding the axonal regeneration by providingneurotropic factors and cell adhesion molecules [30, 31]. Atwo-dimensional aligned PCL/gelatin nanofibrous scaffoldfabricated by Gupta et al showed lesser proliferation ofSchwann cells than random fibers [26]. Currently three-dimensional constructs have received more attention tomaximize the topographical directional cue for neuriteoutgrowth and to minimize the dead space, cross-sectionalarea occupied by the polymer [14]. In our study, Schwanncells were adhered on the surface of the aligned 3D PLGAfibers and the cell proliferation study confirmed that the cellson the aligned fibers have a significantly higher proliferationthan those on random and TCPS. This observation couldbe attributed to the fact that the orientated as well as thethree-dimensional architectures improve the arrangement ofcells which assists the proliferation of a higher number ofcells. SEM micrographs of Schwann cells after days 1 and3 on aligned scaffolds supported the same by covering amaximum surface of the scaffolds. These results indicatedthat the fabricated oriented scaffolds have better flexibilityand prolonged degradation behavior with lesser pore size thanthe random scaffolds. The aligned scaffolds were also ableto guide the neural cells along their length and may provideenough physical substrate to enable regenerating fibers totraverse the larger gaps.

5. Conclusions

We have established a novel approach to producing atubular scaffold with longitudinally aligned structure bymodifying the electrospinning process. Such a techniquealso has an advantage of developing a thicker layer oforiented fibers. In this study, PLGA of both random andlongitudinally oriented nanofibers was fabricated successfullythrough electrospinning. It was confirmed that the orientationof the fiber plays a major role in the tailoring of thescaffold properties such as porosity, mechanical property,degradability and wettability. The longitudinally orientedscaffold demonstrated good flexibility, slower degradation,less pore size with comparable porosity of random fibers and abetter cell proliferation rate than its random fibers. In additionto this, the similar topography of natural ECM indicates thatthe longitudinally aligned scaffold is a suitable substitute fororienting the neural cells across the lesion.

Acknowledgment

The authors wish to acknowledge the infrastructuralgrant provided by the Nano Mission Council, DST(00/IFD/5634/2006–7) and financial assistance provided bythe Nano Science & Technology Initiative, DST, Govt of India(35/12/2009-NAN/BMS).

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