A nerve guidance conduit (also referred to as an artificial nerve conduit or artificial nerve graft, as opposed to an autograft) is an artificial means of guiding axonal regrowth to facilitate nerve regeneration and is one of several clinical treatments for nerve injuries. When direct suturing of the two stumps of a severed nerve cannot be accomplished without tension, the standard clinical treatment for peripheral nerve injuries is autologous nerve grafting. Due to the limited availability of donor tissue and functional recovery in autologous nerve grafting, neural tissue engineering research has focused on the development of bioartificial nerve guidance conduits as an alternative treatment, especially for large defects. Similar techniques are also being explored for nerve repair in the spinal cord but nerve regeneration in the central nervous system poses a greater challenge because its axons do not regenerate appreciably in their native environment. The creation of artificial conduits is also known as entubulation because the nerve ends and intervening gap are enclosed within a tube composed of biological or synthetic materials. Whether the conduit is in the form of a biologic tube, synthetic tube or tissue-engineered conduit, it should facilitate neurotropic and neurotrophic communication between the proximal and distal ends of the nerve gap, block external inhibitory factors, and provide a physical guidance for axonal regrowth. The most basic objective of a nerve guidance conduit is to combine physical, chemical, and biological cues under conditions that will foster tissue formation. Materials that have been used to make biologic tubes include blood vessels and skeletal muscles, while nonabsorbable and bioabsorbable synthetic tubes have been made from silicone and polyglycolide respectively. Tissue-engineered nerve guidance conduits are a combination of many elements: scaffold structure, scaffold material, cellular therapies, neurotrophic factors and biomimetic materials. The choice of which physical, chemical and biological cues to use is based on the properties of the nerve environment, which is critical in creating the most desirable environment for axon regeneration. The factors that control material selection include biocompatibility, biodegradability, mechanical integrity, controllability during nerve growth, implantation and sterilization.
Scaffold topography In tissue engineering, the three main levels of scaffold structure are considered to be:
the superstructure, the overall shape of the scaffold; the microstructure, the cellular level structure of the surface; and the nanostructure, the subcellular level structure of the surface.
Superstructure The superstructure of a conduit or scaffold is important for simulating in vivo conditions for nerve tissue formation. The extracellular matrix, which is mainly responsible for directing tissue growth and formation, has a complex superstructure created by many interwoven fibrous molecules. Ways of forming artificial superstructure include the use of thermo-responsive hydrogels, longitudinally oriented channels, longitudinally oriented fibers, stretch-grown axons, and nanofibrous scaffolds.
Thermo-responsive hydrogels In traumatic brain injury (TBI), a series of damaging events is initiated that lead to cell death and overall dysfunction, which cause the formation of an irregularly-shaped lesion cavity. The resulting cavity causes many problems for tissue-engineered scaffolds because invasive implantation is required, and often the scaffold does not conform to the cavity shape. In order to get around these difficulties, thermo-responsive hydrogels have been engineered to undergo solution-gelation (sol-gel) transitions, which are caused by differences in room and physiological temperatures, to facilitate implantation through in situ gelation and conformation to cavity shape caused, allowing them to be injected in a minimally invasively manner. Methylcellulose (MC) is a material with well-defined sol-gel transitions in the optimal range of temperatures. MC gelation occurs because of an increase in intra- and inter-molecular hydrophobic interactions as the temperature increases. The sol-gel transition is governed by the lower critical solution temperature (LCST), which is the temperature at which the elastic modulus equals the viscous modulus. The LCST must not exceed physiological temperature (37 °C) if the scaffold is to gel upon implantation, creating a minimally invasive delivery. Following implantation into a TBI lesion cavity or peripheral nerve guidance conduit, MC elicits a minimal inflammatory response. It is also very important for minimally invasive delivery that the MC solution has a viscosity at temperatures below its LCST, which allows it to be injected through a small gauge needle for implantation in in vivo applications. MC has been successfully used as a delivery agent for intra-optical and oral pharmaceutical therapies. Some disadvantages of MC include its limited propensity for protein adsorption and neuronal cellular adhesion making it a non-bioactive hydrogel. Due to these disadvantages, use of MC in neural tissue regeneration requires attaching a biologically active group onto the polymer backbone in order to enhance cell adhesion. Another thermo-responsive gel is one that is formed by combining chitosan with glycerophosphate (GP) salt. This solution experiences gelation at temperatures above 37 °C. Gelation of chitosan/GP is rather slow, taking half an hour to initially set and 9 more hours to completely stabilize. Gel strength varies from 67 to 1572 Pa depending on the concentration of chitosan; the lower end of this range approaches the stiffness of brain tissue. Chitosan/GP has shown success in vitro, but the addition of polylysine is needed to enhance nerve cell attachment. Polylysine was covalently bonded to chitosan in order to prevent it from diffusing away. Polylysine was selected because of its positive nature and high hydrophilicity, which promotes neurite growth. Neuron survival was doubled, though neurite outgrowth did not change with the added polylysine.
Longitudinally oriented channels Longitudinally oriented channels are macroscopic structures that can be added to a conduit in order to give the regenerating axons a well-defined guide for growing straight along the scaffold. In a scaffold with microtubular channel architecture, regenerating axons are able to extend through open longitudinal channels as they would normally extend through endoneurial tubes of peripheral nerves. Additionally, the channels increase the surface area available for cell contact. The channels are usually created by inserting a needle, wire, or second polymer solution within a polymer scaffold; after stabilizing the shape of the main polymer, the needle, wire, or second polymer is removed in order to form the channels. Typically multiple channels are created; however, the scaffold can consist of just one large channel, which is simply one hollow tube. A molding technique was created by Wang et al. for forming a nerve guidance conduit with a multi-channel inner matrix and an outer tube wall from chitosan. In their 2006 study, Wang et al. threaded acupuncture needles through a hollow chitosan tube, where they are held in place by fixing, on either end, patches created using CAD. A chitosan solution is then injected into the tube and solidified, after which the needles are removed, creating longitudinally oriented channels. A representative scaffold was then created for characterization with 21 channels using acupuncture needles of 400 μm in diameter. Upon investigation under a microscope, the channels were found to be approximately circular with slight irregularities; all channels were aligned with the inner diameter of the outer tube wall. It was confirmed by micro-CT imaging that the channels went through the entire length of the scaffold. Under water absorption, the inner and outer diameters of the scaffold became larger, but the channel diameters did not vary significantly, which is necessary for maintaining the scaffold shape that guides neurite extension. The inner structure provides an increase in compressive strength compared to a hollow tube alone, which can prevent collapse of the scaffold onto growing neurites. Neuro-2a cells were able to growth on the inner matrix of the scaffold, and they oriented along the channels. Although this method has only been tested on chitosan, it can be tailored to other materials. lyophilizing and wire-heating process is another method of creating longitudinally oriented channels, developed by Huang et al. (2005). A chitosan and acetic acid solution was frozen around nickel-copper (Ni-Cu) wires in a liquid nitrogen trap; subsequently the wires were heated and removed. Ni-Cu wires were chosen because they have a high resistance level. Temperature-controlled lyophilizers were used to sublimate the acetic acid. There was no evidence of the channels merging or splitting. After lyophilizing, scaffold dimensions shrunk causing channels to be a bit smaller than the wire used. The scaffolds were neutralized to a physiological pH value using a base, which had dramatic effects on the porous structure. Weaker bases kept the porous structure uniform, but stronger base made it uncontrollable. The technique used here can be slightly modified to accommodate other polymers and solvents. Another way to create longitudinally oriented channels is to create a conduit from one polymer with embedded longitudinally oriented fibers from another polymer; then selectively dissolve the fibers to form longitudinally oriented channels. Polycaprolactone (PCL) fibers were embedded in a (Hydroxyethyl)methacrylate (HEMA) scaffold. PCL was chosen over poly (lactic acid) (PLA) and poly (lactic-co-glycolic acid) (PLGA), because it is insoluble in HEMA but soluble in acetone. This is important because HEMA was used for the main conduit material and acetone was used to selectively dissolve the polymer fibers. Extruded PCL fibers were inserted into a glass tube and the HEMA solution was injected. The number of channels created was consistent from batch to batch and the variations in fiber diameter could be reduced by creating a more controlled PCL fiber extrusion system. The channels formed were confirmed to be continuous and homogeneous by examination of porosity variations. This process is safe, reproducible and has controllable dimensions. In a similar study conducted by Yu and Shoichet (2005), HEMA was copolymerized with AEMA to create a P(HEMA-co-AMEA) gel. Polycaprolactone (PCL) fibers were embedded in the gel, and then selectively dissolved by acetone with sonication to create channels. It was found that HEMA in mixture with 1% AEMA created the strongest gels. When compared to scaffolds without channels, the addition of 82–132 channels can provide an approximately 6–9 fold increase in surface area, which may be advantageous for regeneration studies that depend on contact-mediated cues. Itoh et al. (2003) developed a scaffold consisting of a single large longitudinally oriented channel was created using chitosan tendons from crabs. Tendons were harvested from crabs (Macrocheira kaempferi) and repeatedly washed with sodium hydroxide solution to remove proteins and to deacetylate the tendon chitin, which subsequently became known as tendon chitosan. A stainless steel bar with triangular-shaped cross-section (each side 2.1 mm long) was inserted into a hollow tendon chitosan tube of circular-shaped cross-section (diameter: 2 mm; length: 15 mm). When comparing the circular-shaped and triangular-shaped tubes, it was found that the triangular tubes had improved mechanical strength, held their shape better, and increased the surface area available. While this is an effective method for creating a single channel, it does not provide as much surface area for cellular growth as the multi-channel scaffolds. Newman et al. (2006) inserted conductive and non-conductive fibers into a collagen-TERP scaffold (collagen cross-linked with a terpolymer of poly(N-isopropylacrylamide) (PNiPAAm) ). The fibers were embedded by tightly wrapping them on a small glass slide and sandwiching a collagen-TERP solution between it and another glass slide; spacers between the glass slides set the gel thickness to 800 μm. The conductive fibers were carbon fiber and Kevlar, and the nonconductive fibers were nylon-6 and tungsten wire. Neurites extend in all directions in thick bundles on the carbon fiber; however with the other three fibers, neurites extended in fine web-like conformations. The neurites showed no directional growth on the carbon and Kevlar fibers, but they grew along the nylon-6 fibers and to some extent along the tungsten wire. The tungsten wire and nylon-6 fiber scaffolds had neurites grow into the gel near the fiber-gel interface in addition to growing along the surface. All fiber gels except Kevlar showed a significant increase in neurite extension compared to non-fiber gels. There was no difference in the neurite extension between the non-conductive and the conductive fibers. In their 2005 study, Cai et al. added Poly (L-lactic acid) (PLLA) microfilaments to hollow poly(lactic acid) (PLA) and silicon tubes. The microfiber guidance characteristics were inversely related to the fiber diameter with smaller diameters promoting better longitudinally oriented cell migration and axonal regeneration. The microfibers also promoted myelination during peripheral nerve repair.
Stretch-grown axons Mature axon tracts has been demonstrated to experience growth when mechanically stretched at the central portion of the axon cylinder. Such mechanical stretch was applied by a custom axon stretch-growth bioreactor composed of four main components: custom-designed axon expansion chamber, linear motion table, stepper motor and controller. The nerve tissue culture is placed within the expansion chamber with a port for gas exchange and a removable stretching frame, which is able to separate two groups of somas (neuron cell bodies) and thus stretch their axons. Collagen gel was used to promote the growth of larger stretch-grown axon tracts that were visible to the unaided eye. There are two reasons for the growth enhancement due to the collagen coating: 1) the culture became hydrophobic after the collagen dried, which permitted a denser concentration of neurons to grow, and 2) the collagen coating created an unobstructed coating across the two elongation substrates. Examination by scanning electron microscope and TEM showed no signs of axon thinning due to stretch, and the cytoskeleton appeared to be normal and intact. The stretch-grown axon tracts were cultured on a biocompatible membrane, which could be directly formed into a cylindrical structure for transplantation, eliminating the need to transfer axons to a scaffold after growth was complete. The stretch-grown axons were able to grow at an unprecedented rate of 1 cm/day after only 8 days of acclimation, which is much greater than the 1 mm/day maximal growth rate as measured for growth cone extension. The rate of 1 mm/day is also the average transport speed for structural elements such as neurofilaments.
Nanofibers scaffolds Research on nanoscale fibers attempts to mimic the in vivo extracellular environment in order to promote directional growth and regeneration. Three distinct methods for forming nanofibrous scaffolds are self-assembly, phase separation and electrospinning. However, there are many other methods for forming nanofibrous scaffolds. Self-assembly of nanofibrous scaffolds is able to occur only when the fibers themselves are engineered for self-assembly. One common way to drive the self-assembly of scaffold fibers is to use amphiphilic peptides so that in water the hydrophobic moiety drives the self-assembly. Carefully calculated engineering of the amphiphilic peptides allows for precise control over the self-assembled matrix. Self-assembly is able to create both ordered and unordered topographies. Phillips et al. (2005) developed and tested in vitro and in vivo a self-aligned collagen-Schwann cell matrix, which allowed DRG neurite extension alignment in vitro. Collagen gels have been used extensively as substrates for three-dimensional tissue culture. Cells are able to form integrin-mediated attachments with collagen, which initiates cytoskeleton assembly and cell motility. As cells move along the collagen fibers they generate forces that contract the gel. When the collagen fibers are tethered at both ends, cell-generated forces create uniaxial strain, causing the cells and collagen fibers to align. The advantages of this matrix are its simplicity and speed of preparation. Soluble plasma fibronectin can also self-assemble into stable insoluble fibers when put under direct mechanical shearing within a viscous solution. Phillips et al. (2004) investigated a new method of shear aggregation that causes an improved aggregation. The mechanical shearing was created by dragging out a 0.2 ml bolus to 3 cm with forceps; fibronectin aggregates into insoluble fibers at the rapidly moving interface in an ultrafiltration cell. The proposed mechanism for this fiber aggregation is protein extension and elongation under mechanical shear force, which leads to lateral packing and protein aggregation of fibers. Phillips et al. showed that mechanical shear produced by stretching a high viscosity fibronectin gel causes substantial changes in its structure and that when applied through uniaxial extension, a viscous fibronectin gel forms oriented fibrous fibronectin aggregates; additionally, the fibrous aggregates have a decreased solubility and can support the various cell types in vitro. Phase separation allows for three-dimensional sub-micrometre fiber scaffolds to be created without the use of specialized equipment. The five steps involved in phase separation are polymer dissolution, phase separation and gelation, solvent extraction from the gel, freezing and freeze drying in water. The final product is a continuous fiber network. Phase separation can be modified to fit many different applications, and pore structure can be varied by using different solvents, which can change the entire process from liquid–liquid to solid–liquid. Porosity and fiber diameter can also be modified by varying the initial concentration of the polymer; a higher initial concentration leads to less pores and larger fiber diameters. This technique can be used to create networks of fibers with diameters reaching type I collagen fiber diameters. The fibrous network created is randomly oriented and so far work has not been done to attempt to organize the fibers. Phase separation is a widely used technique for creating highly porous nanofibrous scaffolds with ease. Electrospinning provides a robust platform for development of synthetic nerve guidance conduits. Electrospinning can serve to create scaffolds at controlled dimensions with varying chemistry and topography. Furthermore, different materials can be encapsulated within fibers including particles, growth factors, and even cells. Electrospinning creates fibers by electrically charging a droplet of polymer melt or solution and suspending it from a capillary. Then, an electric field is applied at one end of the capillary until the charge exceeds the surface tension, creating a polymer jet that elongates and thins. This polymer jet discharges as a Taylor cone, leaving behind electrically charged polymers, which are collected on a grounded surface as the solvent as the solvent evaporates from the jets. Fibers have been spun with diameters ranging from less than 3 nm to over 1 μm. The process is affected by system parameters such as polymer type, polymer molecular weight, and solution properties and by process parameters such as flow rate, voltage, capillary diameter, distance between the collector and the capillary, and motion of the collector. The fibrous network created is unordered and contains a high surface-to-volume ratio as a result of a high porosity; a large network surface area is ideal for growth and transport of wastes and nutrients in neural tissue engineering. The two features of electrospun scaffolds that are advantageous for neural tissue engineering are the morphology and architecture, which closely mimics the ECM, and the pores, which are the correct range of sizes that allows nutrient exchange but prevents in growth of glial scar tissue (around 10 μm). Random electrospun PLLA scaffolds have been demonstrated to have increased cell adhesion, which may be due to an increased surface roughness. Chemically modified electrospun fiber mats have also been shown to influence neural stem cell differentiation and increase cell proliferation. In the past decade, scientists have also developed numerous methods for production of aligned nanofiber scaffolds, which serve to provide additional topographic cues to cells. This is advantageous because large scale three-dimensional aligned scaffolds cannot be created easily using traditional fabrication techniques. In a study conducted by Yang et al. (2005), aligned and random electrospun poly (L-lactic acid) (PLLA) microfibrous and nanofibrous scaffolds were created, characterized, and compared. Fiber diameters were directly proportional to the initial polymer concentration used for electrospinning; the average diameter of aligned fibers was smaller than that of random fibers under identical processing conditions. It was shown that neural stem cells elongated parallel to the aligned electrospun fibers. The aligned nanofibers had a longer average neurite length compared to aligned microfibers, random microfibers, and random nanofibers. In addition, more cells differentiated on aligned nanofibers than aligned microfibers. Thus, the results of this study demonstrated that aligned nanofibers may be more beneficial than nonaligned fibers or microfibers for promoting nerve regeneration.
Microstructure and nanostructure Microstructure and nanostructure, along with superstructure are three main levels of scaffold structure that deserve consideration when creating scaffold topography. While the superstructure refers to the overall shape of the scaffold, the microstructure refers to the cellular level structure of the surface and the nanostructure refers to the subcellular level structure of the surface. All three levels of structure are capable of eliciting cell responses; however, there is significant interest in the response of cells to nanoscale topography motivated by the presence of numerous nanoscale structures within the extracellular matrix. There are a growing number of methods for the manufacture of micro- and nanostructures (many originating from the semiconductor industry) allowing for the creation of various topographies with controlled size, shape, and chemistry.
Physical cues Physical cues are formed by creating an ordered surface structure at the level of the microstructure and/or nanostructure. Physical cues on the nanoscale have been shown to modulate cell adhesion, migration, orientation, contact inhibition, gene expression, and cytoskeletal formation. This allows for the direction of cell processes such as proliferation, differentiation, and spreading. There are numerous methods for the manufacture of micro- and nanoscale topographies, which can be divided into those that create ordered topographies and those that create unordered topographies. Ordered topographies are defined as patterns that are organized and geometrically precise. Though there are many methods for creating ordered topographies, they are usually time-consuming, requiring skill and experience and the use of expensive equipment. Photolithography involves exposing a light source to a photoresist-coated silicon wafer; a mask with the desired pattern is place between the light source and the wafer, thereby selectively allowing light to filter through and create the pattern on the photoresist. Further development of the wafer brings out the pattern in the photoresist. Photolithography performed in the near-UV is often viewed as the standard for fabricating topographies on the micro-scale. However, because the lower limit for size is a function of the wavelength, this method cannot be used to create nanoscale features. In their 2005 study, Mahoney et al. created organized arrays of polyimide channels (11 μm in height and 20–60 μm in width) were created on a glass substrate by photolithography. Polyimide was used because it adheres to glass well, is chemically stable in aqueous solution, and is biocompatible. It is hypothesized that the microchannels limited the range of angles that cytoskeletal elements within the neurite growth cones could accumulate, assemble, and orient. There was a significant decrease in the number of neurites emerging from the soma; however, there was less decrease as the range of angles over which the neurites emerged was increased. Also, the neurites were on average two times longer when the neurons were cultured on the microchannels versus the controls on a flat surface; this could be due to a more efficient alignment of filaments. In electron beam lithography (EBL), an electron-sensitive resist is exposed to a beam of high-energy electrons. There is the choice of a positive or negative type resist; however, lower feature resolution can be obtained with negative resists. Patterns are created by programming the beam of electrons for the exact path to follow along the surface of the material. Resolution is affected by other factors such as electron scattering in the resist and backscattering from the substrate. EBL can create single surface features on the order of 3–5 nm. If multiple features are required over a large surface area, as is the case in tissue engineering, the resolution drops and features can only be created as small as 30–40 nm, and the resist development begins to weigh more heavily on pattern formation. To prevent dissolution of the resist, ultrasonic agitation can be used to overcome intermolecular forces. In addition, isopropyl alcohol (IPA) helps develop high-density arrays. EBL can become a quicker and less costly process by replicating nanometer patterns in polymeric materials; the replication process has been demonstrated with polycaprolactone (PCL) using hot embossing and solvent casting. In a study conducted by Gomez et al. (2007), microchannels 1 and 2 μm wide and 400 and 800 nm deep created by EBL on PDMS were shown to enhance axon formation of hippocampal cells in culture more so than immobilized chemical cues. X-ray lithography is another method for forming ordered patterns that can be used to investigate the role that topography plays in promoting neuritogenesis. The mask parameters determine the pattern periodicity, but ridge width and depth are determined by the etching conditions. In a study, ridges were created with periods ranging from 400 through 4000 nm, widths ranging from 70 through 1900 nm, and a groove depth of 600 nm; developing neurites demonstrated contact guidance with features as small as 70 nm and greater than 90% of the neurites were within 10 degrees of parallel alignment with the ridges and grooves. There was not a significant difference in orientation with respect to the feature sizes used. The number of neurites per cell was constrained by the ridges and grooves, producing bipolar rather than branching phenotypes. Unordered topographies are generally created by processes that occur spontaneously during other processing; the patterns are random in orientation and organization with imprecise or no control over feature geometry. The advantage to creating unordered topographies over ordered is that the processes are often less time-consuming, less expensive, and do not require great skill and experience. Unordered topographies can be created by polymer demixing, colloidal lithography and chemical etching. In polymer demixing, polymer blends experience spontaneous phase separation; it often occurs during conditions such as spin casting onto silicon wafers. Features that can be created by this method include nanoscale pits, islands, and ribbons, which can be controlled to an extent by adjusting the polymer ratio and concentration to change the feature shape and size, respectively. There is not much control in the horizontal direction, though the vertical direction of the features can be precisely controlled. Because the pattern is very unordered horizontally, this method can only be used to study cell interactions with specific height nanotopographies. Colloidal lithography is inexpensive and can be used to create surfaces with controlled heights and diameters. Nanocolliods are used as an etch mask by spreading them along the material surface, and then ion beam bombardment or film evaporation is used to etch away around the nanocolliods, creating nanocolumns and nanopits, respectively. The final surface structure can be controlled by varying the area covered by colloids and the colloid size. The area covered by the colloids can be changed by modifying the ionic strength of the colloid solution. This technique is able to create large patterned surface areas, which is necessary for tissue engineering applications. Chemical etching involves soaking the material surface in an etchant such as hydrofluoric acid (HF) or sodium hydroxide (NaOH) until the surface is etched away to a desired roughness as created by pits and protrusions on the nanometer scale. Longer etch times lead to rougher surfaces (i.e., smaller surface pits and protrusions). Structures with specific geometry or organization cannot be created by this rudimentary method because at best it can be considered a surface treatment for changing the surface roughness. The significant advantages of this method are ease of use and low cost for creating a surface with nanotopographies. Silicon wafers were etched using HF, and it was demonstrated that cell adhesion was enhanced only in a specified range of roughness (20–50 nm).
Chemical cues In addition to creating topography with physical cues, it can be created with chemical cues by selectively depositing polymer solution in patterns on the surface of a substrate. There are different methods for depositing the chemical cues. Two methods for dispensing chemical solutions include stripe patterning and piezoelectric microdispensing. Stripe-patterned polymer films can be formed on solid substrates by casting diluted polymer solution. This method is relatively easy, inexpensive, and has no restriction on the scaffold materials that can be used. The procedure involves horizontally overlapping glass plates while keeping them vertically separated by a narrow gap filled with a polymer solution. The upper plate is moved at a constant velocity between 60 and 100 μm/s. A thin liquid film of solution is continuously formed at the edge of the sliding glass following evaporation of the solvent. Stripe patterns prepared at speeds of 60, 70, and 100 μm/s created width and groove spacings of 2.2 and 6.1 μm, 3.6 and 8.4 μm, and 4.3 and 12.7 μm, respectively; the range of heights for the ridges was 50–100 nm. Tsuruma, Tanaka et al. demonstrated that embryonic neural cells cultured on film coated with poly-L-lysine attached and elongated parallel to poly(ε-caprolactone)/chloroform solution (1g/L) stripes with narrow pattern width and spacing (width: 2.2 μm, spacing: 6.1 μm). However, the neurons grew across the axis of the patterns with wide width and spacing (width: 4.3 μm, spacing: 12.7 μm). On average, the neurons on the stripe-patterned films had less neurites per cell and longer neurites compared to the neurons on non-patterned films. Thus, the stripe pattern parameters are able to determine the growth direction, the length of neurites, and the number of neurites per cell. Microdispensing was used to create micropatterns on polystyrene culture dishes by dispensing droplets of adhesive laminin and non-adhesive bovine serum albumin (BSA) solutions. The microdispenser is a piezoelectric element attached to a push-bar on top of a channel etched in silicon, which has one inlet at each end and a nozzle in the middle. The piezoelectric element expands when voltage is applied, causing liquid to be dispensed through the nozzle. The microdispenser is moved using a computer-controlled x-y table. The micropattern resolution depends on many factors: dispensed liquid viscosity, drop pitch (the distance between the centre of two adjacent droplets in a line or array), and the substrate. With increasing viscosity the lines become thinner, but if the liquid viscosity is too high the liquid cannot be expelled. Heating the solution creates more uniform protein lines. Although some droplet overlap is necessary to create continuous lines, uneven evaporation may cause uneven protein concentration along the lines; this can be prevented through smoother evaporation by modifying the dispensed solution properties. For patterns containing 0.5 mg/mL laminin, a higher proportion of neurites grew on the microdispensed lines than between the lines. On 10 mg/mL and 1 mg/mL BSA protein patterns and fatty-acid free BSA protein patterns a significant number of neurites avoided the protein lines and grew between the lines. Thus, the fatty-acid-containing BSA lines were just as non-permissive for neurite growth as lines containing BSA with fatty acids. Because microdispensing does not require direct contact with the substrate surfaces, this technique can utilitze surfaces with delicate micro- or nanotopology that could be destroyed by contact. It is possible to vary the amount of protein deposited by dispensing more or less droplets. An advantage of microdispensing is that patterns can be created quickly in 5–10 minutes. Because the piezoelectric microdispenser does not require heating, heat-sensitive proteins and fluids as well as living cells can be dispensed.
Scaffold material The selection of the scaffold material is perhaps the most important decision to be made. It must be biocompatible and biodegradable; in addition, it must be able to incorporate any physical, chemical, or biological cues desired, which in the case of some chemical cues means that it must have a site available for chemically linking peptides and other molecules. The scaffold materials chosen for nerve guidance conduits are almost always hydrogels. The hydrogel may be composed of either biological or synthetic polymers. Both biological and synthetic polymers have their strengths and weaknesses. It is important to note that the conduit material can cause inadequate recovery when (1) degradation and resorption rates do not match the tissue formation rate, (2) the stress-strain properties do not compare well to those of neural tissue, (3) when degrading swelling occurs, causing significant deformation, (4) a large inflammatory response is elicited, or (5) the material has low permeability.
Hydrogel Hydrogels are a class of biomaterials that are chemically or physically cross-linked water-soluble polymers. They can be either degradable or non-degradable as determined by their chemistry, but degradable is more desirable whenever possible. There has been great interest in hydrogels for tissue engineering purposes, because they generally possess high biocompatibility, mechanical properties similar to soft tissue, and the ability to be injected as a liquid that gels. When hydrogels are physically cross-linked they must rely on phase separation for gelation; the phase separation is temperature-dependent and reversible. Some other advantages of hydrogels are that they use only non-toxic aqueous solvents, allow infusion of nutrients and exit of waste products, and allow cells to assemble spontaneously. Hydrogels have low interfacial tension, meaning cells can easily migrate across the tissue-implant boundary. However, with hydrogels it is difficult to form a broad range of mechanical properties or structures with controlled pore size.
Synthetic polymer A synthetic polymer may be non-degradable or degradable. For the purpose of neural tissue engineering degradable materials are preferred whenever possible, because long-term effects such as inflammation and scar could severely damage nerve function. The degradation rate is dependent on the molecular weight of the polymer, its crystallinity, and the ratio of glycolic acid to lactic acid subunits. Because of a methyl group, lactic acid is more hydrophobic than glycolic acid causing its hydrolysis to be slower. Synthetic polymers have more wieldy mechanical properties and degradation rates that can be controlled over a wide range, and they eliminate the concern for immunogenicity. There are many different synthetic polymers currently being used in neural tissue engineering. However, the drawbacks of many of these polymers include a lack of biocompatibility and bioactivity, which prevents these polymers from promoting cell attachment, proliferation, and differentiation. Synthetic conduits have only been clinically successful for the repair of very short nerve lesion gaps less than 1–2 cm. Furthermore, nerve regeneration with these conduits has yet to reach the level of functional recovery seen with nerve autografts.
Collagen-terpolymer Collagen is a major component of the extracellular matrix, and it is found in the supporting tissues of peripheral nerves. A terpolymer (TERP) was synthesized by free radical copolymerization of its three monomers and cross-linked with collagen, creating a hybrid biological-synthetic hydrogel scaffold. The terpolymer is based on poly(NIPAAM), which is known to be a cell friendly polymer. TERP is used both as a cross-linker to increase hydrogel robustness and as a site for grafting of bioactive peptides or growth factors, by reacting some of its acryloxysuccinimide groups with the –NH2 groups on the peptides or growth factors. Because the collagen-terpolymer (collagen-TERP) hydrogel lacks a bioactive component, a study attached to it a common cell adhesion peptide found in laminin (YIGSR) in order to enhance its cell adhesion properties.
Poly (lactic-co-glycolic acid) family The polymers in the PLGA family include poly (lactic acid) (PLA), poly (glycolic acid) (PGA), and their copolymer poly (lactic-co-glycolic acid) (PLGA). All three polymers have been approved by the Food and Drug Administration for employment in various devices. These polymers are brittle and they do not have regions for permissible chemical modification; in addition, they degrade by bulk rather than by surface, which is not a smooth and ideal degradation process. In an attempt to overcome the lack of functionalities, free amines have been incorporated into their structures from which peptides can be tethered to control cell attachment and behavior.
Methacrylated dextran (Dex-MA) copolymerized with aminoethyl methacrylate (AEMA) Dextran is a polysaccharide derived from bacteria; it is usually produced by enzymes from certain strains of leuconostoc or Streptococcus. It consists of α-1,6-linked D-glucopyranose residues. Cross-linked dextran hydrogel beads have been widely used as low protein-binding matrices for column chromatography applications and for microcarrier cell culture technology. However, it has not been until recently that dextran hydrogels have been investigated in biomaterials applications and specifically as drug delivery vehicles. An advantage of using dextran in biomaterials applications include its resistance to protein adsorption and cell-adhesion, which allows specific cell adhesion to be determined by deliberately attached peptides from ECM components. AEMA was copolymerized with Dex-MA in order to introduce primary amine groups to provide a site for attachment of ECM-derived peptides to promote cell adhesion. The pepti