The nervous system plays a central and complex role in human biological processes interacting in physiological processes, such as cognition and individual cell function. Damage to the nerve may impose tremendous consequences and its recovery is difficult. Moreover, malfunctions in other parts of the body might also occur because mature neurons do not undergo cell division (Huang et al., 2006). Neurodegenerative disorders of the spinal cord and brain after injury, stroke or multiple sclerosis have increased over the past few years (Prabhakaran et al., 2009). Peripheral nerve lesions are also common, with serious injuries affecting 2.8% of trauma patients annually, leading to lifelong disability (Ciardelli and Chiono, 2006). In the USA, 360 000 people suffer from upper extremity paralytic syndromes on an annual basis and approximately 253 000 people in the USA live with the after-effects of spinal cord injury. Moreover, each year this number grows by an estimated 11 000 people in the USA (Willerth and Sakiyama-Elbert, 2007) and in Europe more than 300 000 cases of peripheral nerve injury are reported annually (Ciardelli and Chiono, 2006; Bueno and Shah, 2008).
Numerous strategies have been applied for the repair of peripheral nerve lesions, with the common goals of directing the regenerating nerve fibres into the proper distal endoneural tubes and improving the prospects of axonal regeneration and functional recovery (Schmidt et al., 1997). Implantation of autografts, allografts and xenografts, providing grafts from the patient, cadavers and animals, respectively, are a few strategies applied in this field. But the loss of function at the donor nerve graft site and mismatch of damaged nerve and graft dimensions are major disadvantages of using autograft nerve repair system (Schmidt et al., 1997; Schmidt and Leach, 2003). On the other hand, allogenic and xenogeneic tissues have the advantages of their availability, along with the benefit of not requiring harvesting from patients. However, their disadvantages include disease transmission and problems of immunogenicity (Schmidt and Leach, 2003).
Tissue engineering provides a new medical therapy as an alternative to conventional transplantation methods, which regulates the cell behaviour and tissue progression through the development and design of synthetic extracellular matrix (ECM) analogues of novel biomaterials to support three-dimensional (3D) cell culture and tissue regeneration (Yang et al., 2004a; Subramanian et al., 2009). The fundamental approach in neural tissue engineering involves the fabrication of polymeric scaffolds seeded with nerve cells to produce a 3D functional tissue suitable for implantation (Yang et al., 2004). Historically, tissue-engineering strategies have been used in an effort to develop therapies for peripheral nerve and spinal cord injury, combining biomaterials, cell therapy and drug delivery approaches (Li and Hoffman-Kim, 2008). Successful nerve regeneration requires tissue-engineered scaffolds that provide not only mechanical support for growing neurites and prevention of ingrowth of fibrous scar tissue, but also biological signals to direct the axonal growth cone to the distal stump (Huang et al., 2006). Recently, a synthetic nerve guidance channel has provided surgeons with an interesting option to bridge severed nerves instead of conventional methods including autografts, allografts and xenografts (Zhang et al., 2007), and numerous efforts have been devoted towards the development of synthetic guidance conduits for the repair of peripheral nerve defects (Hadlock et al.,2000; Kijima et al., 2009; Rooney et al., 2008a; Ellis and Chaudhuri, 2008; Jiang et al., 2008; Rooney et al., 2008; Patel et al., 2009; Houchin-Ray et al., 2009; Park et al., 2007). Nerve guidance channels are biomaterials-based devices that are designed to be transplanted for nerve repair, experimentally studied as a possible alternative to nerve grafts. Nerve guidance channels aim to provide a conduit through which regenerating axons can grow and connect to their appropriate targets (Li and Hoffman-Kim, 2008). An appropriate synthetic material for fabricating nerve guidance channels must readily shape to a conduit with the desired dimensions, must be sterilizable, tear-resistant, easy to handle and suture, biodegradable, and should maintain its shape and resist collapse during implantation over the course of nerve regeneration (Blacher et al., 2003). The physical, chemical and electrical properties of synthetic conduit affect the outcome of nerve regeneration. The inherent nature of neurons is to transmit electrochemical signals throughout the nervous system and, as a result, they are highly influenced by electrical stimuli (Yu et al., 2008b). Previous studies have shown that electrical stimulation is an effective cue in stimulating either the proliferation or differentiation of various cell types (Zhang et al., 2007; McCaig and Zhao, 1997; Sun et al., 2006; Ciombor and Aaron, 1993; Aaron and Ciombor, 1993; Goldman and Pollack, 1996; Dust and Bawornluck, 2006; Shi et al., 2008; Zhao et al.,1996, 1999; Yaoita et al.,1990; Guimarda et al., 2007; Rivers et al., 2002; Jeong et al., 2008; Whitehead et al., 2007; Ateh et al., 2006; Schmidt et al., 2003; Valentini et al., 1992). In this review we discuss the electrical properties of nerve cells, the most commonly utilized conductive polymers, namely polypyrrole (PPy) and polyaniline (PANI), the principle of electrical stimulation and the application of electrical stimulation through conductive scaffolds to nerve cells. A brief evaluation of the methods of electrical stimulation, electrospinning of conductive polymers and modifications carried out to conductive scaffolds, thus making them suitable substrates for tissue engineering, are also discussed.
2. Electrical properties of nerve cells
Recent interest in electrical stimulation arises from a growing knowledge of the electrical properties of tissues and cells (Snyder-Mackler, 1987). Living cells employ many of the properties of electrical systems. For example, they generate electromotive force, maintain a required difference in potential, increase or decrease the difference in potential as necessary, use varying resistances in series or parallel, switch current on and off, control current flow, rectify current flow and store charge, which is even more crucial (Kitchen, 2002). An electrical voltage exists across the plasma membrane, while the inside of the cell remains more negative than the outside. By convention, the potential outside the cell is called zero; therefore, the typical value of the membrane potential is −60 to −100 mV (Matthews, 2003). This potential difference is maintained at a steady level when excitable cells are inactive and is called the resting potential (Paul, 1975). Regarding the electrical properties of cells, electrical signals strongly affect cell behaviour, affecting ion influx across the cell membrane, altering the membrane potential and conditioning the intracellular signal transduction pathways (Mattioli-Belmonte et al., 2003). The transition of information from one place to another in the nervous system takes place along the axon. It is known that the activity in axons is accompanied by electrical changes in and around them and the specific electrical event that occurs is called the action potential, which is an active response generated by the neuron that appears on an oscilloscope as a brief (∼1 ms) change from negative to positive in the transmembrane potential. The action potential represents transient changes in the resting membrane potential. One way to elicit an action potential is to pass an electrical current across the membrane of the neuron. Several steps are involved during this process, the first being a stimulus received by the dendrite of a nerve cell. This causes the Na+ channels to open and, if the opening is sufficient to drive the interior potential from −70 mV to −55 mV, it reaches the action threshold. In this step, more Na+ channels are opened and the Na+ influx drives the interior of the cell membrane up to about + 30 mV. The process at this point is called depolarization. Further, the Na+ channels close and the K+ channels are opened, and since the K+ channels are much slower to open, the depolarization process is completed. The membrane begins to repolarize back towards its resting potential as the K+ channels open. The repolarization typically overshoots the resting potential to about −90 mV, when it is termed hyperpolarization and prevents the neuron from receiving another stimulus, or at least raises the threshold for a new stimulus. Hyperpolarization also assists in preventing any stimulus that has already been sent up an axon from triggering another action potential in the opposite direction. In other words, hyperpolarization assures that the signal is always proceeding in one direction. After hyperpolarization, the Na+/K+ pump eventually brings the membrane back to its resting state of −70 mV (Kandel, 2000).
Typically, the membrane must be depolarized by about 10–20 mV in order to trigger an action potential. Figure 1 shows a schematic illustration of the resting state, depolarization, action potential and repolarization of nerve cells. As can be observed in this figure, in response to the appropriate stimulus, the cell membrane of a nerve cell goes through a sequence of depolarization from its resting state, followed by repolarization to that resting state. In the sequence, it actually reverses its normal polarity for a brief period before re-establishing the resting potential.
3. Electrical stimulation and its effects
An action potential can be elicited artificially by changing the electrical potential of a nerve cell by inducing an electrical charge to the cells, and the process is termed ‘electrical stimulation’ (Kitchen, 2002). A variety of cellular responses to electric stimulation of different cell types, including fibroblasts, osteoblasts, myoblasts, chick embryo dorsal root ganglia and neural crest cells, have been reported (Schmidt et al., 1997; Kimura et al., 1998; Wong et al., 1994; Li et al., 2006; Bidez et al., 2006; Wood et al., 2006). The proposition related to electrical stimulation is based on the fact that bioelectricity present in the human body plays an integral role in maintaining normal biological functions, such as signalling of the nervous system, muscle contraction and wound healing (Shi et al., 2008). McCaig et al. (1997) reported the generation of electrical fields during major cellular events such as cell division, development and migration. They also found that the presence of a steady weak direct current (DC) electrical field in some biological systems affects cellular activities such as cell division, differentiation, migration and the extension of motile processes (Zhao et al., 1999). Endogenous electric fields in the form of voltage gradients have been observed to polarize the nervous system along the rostral–caudal axis (5–18 mV/mm) and to direct nerve growth (Li and Hoffman-Kim, 2008). Initial studies investigating the effect of electrical stimulation on neurons were performed on Xenopus neurons after exposing them to a steady direct current field. Extracellular electric fields (0.1–10 V/cm) applied in solution reversibly influenced the direction of neurite growth and increased the neurite initiation and length in Xenopus (Li and Hoffman-Kim, 2008). Applied electrical fields have been shown to influence the rate and orientation of neurite outgrowth from cultured neurons in vitro (Valentini et al., 1992). For example, applied electric fields influenced the extension and direction of neurite outgrowth from neurons cultured in vitro and pulsed electromagnetic fields stimulated sciatic nerve regeneration in vivo (Wang et al., 2004). Borgens (1999) demonstrated that 7 days of electric field imposed within a damaged adult guinea-pig spinal cord can both induce the regeneration of axons and guide their growth into the ends of a hollow silicone rubber tube inserted into the dorsal half of the cord. Borgens (1999) concluded that this was due to production of a DC voltage gradient within the injured spinal cord, with the cathode located within the experimental tubes. Electrical stimulation has been shown to influence the differentiation of stem cells. Yamada et al. (2007) showed that mild electrical stimulation strongly influences embryonic stem cells to assume a neuronal fate. Although the resulting neuronal cells showed no sign of specific terminal differentiation in culture, they showed potential to differentiate into various types of neurons in vivo, and contributed to the injured spinal cord as neuronal cells. The induction of calcium ion influx is significant in this differentiation system.
Several theories have been suggested to explain the effect of electric stimulation on nerve regeneration. Patel et al. (1982) suggested three possible ways by which electrical stimulation could act directly on a neuron, including the redistribution of cytoplasmic materials, the activation of growth-controlling transport processes across the plasma membrane due to change in cell membrane potential, and the electrophoretic accumulation of surface molecules responsible for neurite growth or cell–substratum adhesion. Changes in ionic currents around the growing fibre tips induced by electric fields have been suggested by Freeman et al. (1985) as one possible mechanism through which electrical stimulation can affect nerve cells. Sisken et al. (1989) suggested that electrical stimulation affects protein synthesis in transected sciatic nerve segments and stimulates neurite outgrowth in vitro. Kimura et al. (1998) postulated that gene expression for nerve growth factor (NGF) is electrically activated for rat neuronal pheochromocytoma cells (PC12 cells) by alternative potential, while Kotwal et al. (2001) showed that fibronectin adsorption increased with immediate electrical stimulation and explained enhanced neurite extension on electrically stimulated PPy films.
3.1. Method of electrical stimulation in the tissue engineering context
Electrical stimulation is a relatively simple, flexible and feasible method carried out for both in vivo and in vitro two-dimensional (2D) and 3D cultured cells (Sun et al., 2006). The behaviour of an excitable cell such as nerve cells can be modified by application of electrical current through two external electrodes. The current passing between the electrodes can cause depolarization of the membrane (Paul, 1975). Surface electrodes, usually made of silver plates, available in different sizes in the range 0.5–1.0 cm in diameter, are commonly used for clinical applications. Gold and platinum electrodes have also been used for electrical stimulation procedures (Kimura, 2001). Some researchers applied voltage to conductive materials, considering it as one of the electrodes, where another electrode was used separately as anode or cathode. For example, Schmidt and co-workers (1997) utilized PPy film as the anode and a gold wire as the cathode for the electrical stimulation of PC12 cells. In other cases, voltage was applied between two electrodes through conductive scaffolds seeded with cells (Shi et al., 2007, 2008; Ghasemi-Mobarakeh et al., 2009).
Shi et al. (2007) studied the effect of electrical stimulation in culture media and did not detect any measurable variations in pH or temperature during the period of cell culture. Moreover, no biologically significant ionic current was observed in the electrical cell culture system. To avoid cultures from potentially toxic electrode products, agar-gelled salt bridges have also been applied for connecting the metal electrodes within the culture medium, where one end of a salt bridge is connected to a scaffold placed in the culture medium, and the other end of the salt bridge is placed in a beaker of electrodes, along with their relevant salt (e.g. Ag/AgCl) and the electrodes in each beaker are attached to a DC power supply (Mccaig et al., 2005).
Salt bridges are prepared from flexible plastic tubing filled with 2% agarose in PBS. This provides a conducting pathway for applied current and prevents electrolysis products from contaminating the chamber by providing a pathway for current transfer from each media reservoir to the AgAgCl electrode (Tandon et al.,2009).
The amplitude of the stimulus is also very important and a direct relationship exists between stimulus amplitude and response amplitude within a small range, which is quite enough to induce the depolarization of nerve cells (Paul, 1975). If the amplitude of the electrical stimulus is too weak to produce a threshold depolarization, an action potential will not take place. If the applied current depolarizes the membrane to threshold, an action potential will result.
Steady DC voltage has been applied in many previous studies for electrical stimulation, mainly due to the existence of DC electrical gradients of voltage within tissues (endogenous electrical fields) (Mccaig et al., 2005). Wood et al. (2006) investigated the influence of brief DC electric stimulation on neurite outgrowth and outgrowth rates after application. Their results showed that the presence of a 25 V/m electrical field for 10 min increased overall neurite outgrowth over controls for up to 48 h after stimulation. In the literature, both electrical field and current have been reported to be effective in modulating cell behaviour. There are a few reports indicating more effectiveness of DC electrical fields than DC currents in electrical stimulation. Shi et al. (2007) applied both electrical field (100 mV/mm) and current (2.5–250 µA/mm) for electrical stimulation and concluded that a wide range of surface current density (2.5–250 µA/mm) had no significant effect on cell adhesion or cell viability, while a constant electrical field of 100 mV/mm upregulated the mitochondrial activity of human cutaneous fibroblasts and enhanced their adhesion. Their results showed that electrical field plays a more substantial role than electrical current in modulating the activity of cells cultured on conductive polymeric scaffolds compared to the non-stimulated cell cultures. However, these researchers did not rule out the importance of current as potential gradient reaches a certain critical level. Some studies also showed that, on exposure to an electrical field (<100 mV/mm), stimulative effects on cell migration and cell growth occur. Many cell types respond to an applied electric field by a perpendicular orientation. The mechanism is not very clear but the asymmetric redistribution of charged cell surface receptors might be effective (Zhao et al., 1996). Electrical field plays an important role in certain processes that regulate the axis of cell division, and in the presence of electrical field cell divisions occurred in vivo (Zhao et al., 1999).
However, electrical stimulation in the form of a pulsed electromagnetic field has also been applied in tissue engineering as well as in clinical settings, as an alternative treatment to promote wound healing, reduce chronic pain and headaches, to treat diseases such as Parkinson's disease and to enhance nerve regeneration (Shi et al., 2008). In vivo studies carried out by Sisken et al. (1989) showed that stimulation of rat sciatic nerves with pulsed electromagnetic fields (PEMFs) for as short a time as 1 h/day for 3 days increased nerve regeneration rates. Various conditions have been described in the literature for electrical stimulation for tissue engineering applications. Table 1 gives a summary of various conditions of electrical stimulation which have been applied by various researchers.
Table 1. Electrical stimulation conditions applied by various research groups
4. Electrically conducting materials in nerve tissue engineering
Electrically conductive polymers have attracted much interest in the last 20 years because they simultaneously display the physical and chemical properties of organic polymers and the electrical characteristics of metals (Chronakis et al., 2006) and are very attractive materials for the construction of nerve guidance channels. The use of conductive polymers would allow one to locally deliver electrical stimulus, provide a physical template for cell growth and tissue repair and allow precise external control over the level and duration of stimulation (Skotheim and Reynolds, 2007). These materials show good capacity to support and modulate the growth of various cells, such as nerve cells and bone cells, and are widely used in biological systems (Zhang et al., 2010).
The importance of conductive polymeric composites is based on the hypothesis that such composites can be used to host the growth of cells, so that electrical stimulation can be applied directly to the cells through the composite, proved to be beneficial in many regenerative medicine strategies, including neural and cardiac tissue engineering (Bettinger et al., 2009). Conductive polymers show great promise in biomedicine and are stimulus-responsive polymers that can be synthesized to form composites that could serve as ‘smart’ biomaterials (Skotheim and Reynolds, 2007). Unlike exogenous electromagnetic fields, electrical stimulation through such polymers would be predominantly focused to the area around the polymer, allowing spatial control of stimulation (Schmidt et al., 1997).
4.1. Conducting polymers
Polymers with loosely held electrons in their backbones are often referred as conducting polymers. Each atom along the backbone is involved in a π bond, which is much weaker than the σ bonds that hold the atoms in the polymer chain together, and they characteristically have a conjugated backbone with a high degree of π-orbital overlap (Breads and Silbey, 1991). Through a process known as ‘doping’, the neutral polymer chain can be oxidized or reduced to become either positively or negatively charged (Wong et al., 1994). It is well known that conducting polymers in general are not conductive without doping, and doping of π-conjugated polymers results in a highly conducting state of the polymer. The doping process includes charge transfer from dopant molecules to polymer chains within an overall neutral system, and in this process charge carriers, polarons and bipolarons, are introduced into the conjugated chain. The doping process can be influenced by factors such as polaron length, chain length, charge transfer to adjacent molecules and conjugation length (Breads and Silbey, 1991). Different dopants have been used for the protonation of conductive polymers. For example, strong inorganic hydrochloric acid (HCl), organic and aromatic acids containing different aromatic substitution, such as p-toluene sulphonic acid (PTSA), dodecylbenzenesulphonic acid (DBSA), organic and aliphatic acids having a long hydrocarbon chain, such as lauric acid (LA) and amphiphilic dopants that belong to the family of sulphonic acids have been used as dopants for PANI. The properties of doped conductive polymers depend on the type and molecular size of the dopant (Sinha et al., 2009; Reena et al., 2009). Liu et al. (1994) showed that the surface energies of the doped conducting polymers vary greatly, depending on the choice of the dopants and doping level. For example, surface energies and polarities of the HCl-doped PANI increase with increasing HCl concentration in the doping solution. Sinha et al. (2009) also showed that the solubility of doped PANI depends on the molecular size of the dopant, and with increase in the dopant chain length the solubility increases. Jang et al. (2004) synthesized organic solvent-soluble PPy by using functional dopants (such as naphthalene sulphonic acid sodium salt, dodecylbenzenesulphonic acid sodium salt, butyl naphthalene sulphonic acid sodium salt and di-2-ethylhexyl sulphosuccinic acid sodium salt). Liu et al. (2008a) doped PPy with different naphthalene mono/disulphonic acids by in situ doping polymerization, and investigated the thermal stability of PPy doped by different dopants. Their results showed that the thermal stability of PPy was greatly affected by the type and concentration of the dopant. PPy doped with non-biologically active dopants, such as tosylate, have been characterized for biological interactions as they can trigger cellular responses in biological applications. However, the incorporation of more biologically active dopants has shown greater promise and can be used to modify PPy-based hybrid material for biomedical applications (Guimarda et al., 2007; Skotheim and Reynolds, 2007). Although dopants that are relevant to biomedical applications, such as hyaluronic acid, can be used to develop PPy-based hybrid material, the choice of dopants is limited to the size and charge (negative) of molecules (Sanchvi et al., 2005).
The biggest limitation of conductive polymers for in vivo applications is their inherent inability to degrade, which may induce chronic inflammation and require surgical removal (Huang et al., 2007). To address the drawbacks of existing conductive polymers, attempts to blend them with suitable biodegradable polymers have been carried out. Rivers et al. (2002) synthesized biodegradable conductive polymer from conducting oligomers of pyrrole and thiophene that were connected together via degradable ester linkages. Ester linkages can be cleaved by enzymes in vivo and the resultant polymer had the unique properties of being both electrically conducting and biodegradable (Rivers et al., 2002). Studies by Shi et al. (2004) also introduced the biodegradation property to conductive polymers by blending them with suitable degradable polymers. On the other hand, Huang et al. (2007) fabricated a novel electroactive biodegradable composite containing polylactide as the biodegradable segment and the low molecular weight aniline pentamer as electroactive segment. A novel block copolymer of polyglycolide and aniline pentamer that is electro-active and degradable was also synthesized by Ding and co-workers (2007). Zhang et al. (2010) synthesized a novel, electricallyconductive and biodegradable polyphosphazene polymer containing aniline pentamer with glycine ethyl ester as side chains and evaluated its biocompatibility using Schwann cells. Their results showed that the polymer exhibited no cytotoxicity, indicating suitability of this polymer as a scaffold material for peripheral nerve regeneration or other biomedical devices that require electroactivity. Huang et al. (2008) synthesized a multiblock copolymer (PLAAP) through condensation polymerization of hydroxyl-capped poly(L-lactide) (PLA) and carboxyl-capped aniline pentamer (AP) with excellent electroactivity and biodegradability, suitable for tissue engineering application.
PPy and PANI remain the most extensively studied conductive polymers for tissue engineering purposes to date, and hence emphasis is given to these polymers in this review, in terms of their biocompatibility for tissue-engineering applications and especially for nerve tissue engineering.
Polypyrrole (PPy) is a conductive synthetic polymer that has numerous applications in drug delivery and nerve regeneration and it has also been used in biosensors and coatings for neural probes (Sanchvi et al., 2005). Figure 2 shows the chemical structure of PPy before and after doping. The high degree of conjugation in the molecular backbone of PPy makes it very rigid, insoluble and poorly processable. It is therefore very difficult to be used alone as a structural material and must be optimized and transformed into a mechanically manageable and processable form (Shi et al., 2004).
Wong et al. (1994) synthesized optically transparent PPy thin films and studied them for mammalian cell culture. In vitro studies demonstrated that PPy thin films supported endothelial cell attachment and growth. Wang et al. (2004) showed that PPy has good biocompatibility with rat peripheral nerve tissue and showed it to be a suitable substrate for bridging the peripheral nerve gap. They reported that PPy extraction solution showed no evidence of acute and subacute toxicity, pyretogens, haemolysis, allergens and mutagenesis, and the migration of Schwann cells and the neurite extension from dorsal root ganglia on the surface of PPy membrane-coated glass were found to be better than those on bare glass (Wang et al., 2004). Williams and Doherty (1994) investigated the biocompatibility of PPy in vitro and in vivo using mouse fibroblast (L929) and neuroblastoma cells. Their results demonstrated that PPy is cytocompatible and in vivo experiments showed that there was only a minimal tissue response after 4 weeks in situ. Jiang et al. (2002) coated polyester fabrics with PPy and investigated the in vivo biocompatibility and biostability of PPy-coated polyester fabrics. Their results showed in vivo biocompatibility of the PPy-coated and non-coated polyester fabrics. In vitro enhancement of nerve cell axonal extension has been reported using PPy with application of either constant current or constant voltage (Rivers et al., 2002). Chen et al. (2000) fabricated PPy tube as nerve guide for sciatic nerve regeneration. Their results showed light contraction of the gastrocnemius muscles and little ulceration, proving the biocompatibility of PPy. George et al(2009) also fabricated conductive PPy tubes using the electrodeposition of PPy onto wire templates and subsequent separation from the template after electrochemical reduction. Conduits made of this material did not show any acute or active chronic inflammatory infiltrate, or tissue damage in the surrounding tissues, for at least 8 weeks of in vivo implantation as sciatic nerve guides. Olayo et al. (2008) synthesized semiconductor biomaterials of iodine-doped PPy and PPy-polyethylene glycol. Their results showed good compatibility of these materials after implantation into the transectioned spinal cord tissue, indicating the suitability of the materials for repairing spinal cord damage.
To date, PPy has been reported to support cell adhesion and growth of a number of different cell types, which makes it a potential candidate for tissue engineering (Garner et al., 1999a, 1999b; Song et al., 2006; Stauffer et al., 2006; Gomez et al., 2007b; Castano et al., 2004; Seal et al., 2001; Evans, 2001; Ateh et al., 2006b).
PANI is the oxidative polymeric product of aniline under acidic conditions and is commonly known as aniline black (Nalwa, 1997). Depending on the oxidation level, PANI can be synthesized in various insulating forms, such as the fully reduced leucoemeraldine base, half-oxidized emeraldine base (PANI-emeraldine base) and the fully oxidized pernigraniline base. PANI emeraldine base is the most stable and widely investigated form of PANI (León, 2001). Figure 3 shows the reduced, oxidized and half-oxidized forms of PANI, while Figure 4 shows the chemical structure of PANI emeraldine base (the most important form of PANI) before and after doping (MacDiarmid and Epstein, 1994).
The exploration of PANI for tissue-engineering applications has progressed more slowly than the development of PPy for similar applications. However, recently there has been more evidence of the ability of PANI and PANI variants to support cell growth (Mattioli-Belmonte et al., 2003).
Wang et al. (1999) investigated the in vivo tissue response to PANI and found no characteristic features resulting from tissue incompatibility after PANI implantation. Studies have assessed the in vivo response to implants of different oxidation states of PANI. In general, no significant inflammation at the implant site and no signs of abnormality of muscle and adipose tissues in the vicinity of the implants were observed (Guimarda et al., 2007). Kamalesh et al. (2000) also investigated the biocompatibility of PANI in different states and found that these polymers are sufficiently biocompatible to be used for biomedical applications. Bidez et al. (2006) investigated the adhesion and proliferation properties of H9c2 cardiac myoblasts on a conductive PANI substrate. Both the non-conductive emeraldine base and the conductive salt forms of PANI were found to be biocompatible and to support cell attachment and proliferation. The demonstration of the biocompatibility of PANI in vivo has sparked much interest in tissue-engineering applications (Guimarda et al., 2007).
4.2. Other conducting polymers
Although PPy and PANI remain the most extensively studied conductive polymers for tissue-engineering purposes to date, the use of a few other conductive polymers, such as polythiophene, for tissue engineering application has also been investigated (Guimarda et al., 2007).
4.2.1. Poly (3, 4-ethylenedioxythiophene)
Poly (3,4-ethylenedioxythiophene) (PEDOT) is considered the most successful polythiophene derivative, with interesting properties. Valle et al. (2007) investigated the interaction of PEDOT films with epithelial cells and showed the biocompatibility of PEDOT with epithelial cells. Luo et al. (2008) synthesized PEDOT films and investigated their biocompatibility by seeding NIH3T3 fibroblasts on PEDOT films, and carried out subcutaneous implantation of PEDOT films by in vivo study. Their results showed that PEDOT films exhibit very low intrinsic cytotoxicity and that their inflammatory response upon implantation was good, making them ideal for biosensing and bioengineering applications. Bolin et al. (2009) fabricated electrospun poly(ethylene terephthalate) (PET) nanofibres and coated PET nanofibres with PEDOT doped with tosylate, using the vapour phase polymerization method. Their results showed excellent adhesion and proliferation of neuronal cells on PEDOT-coated PET nanofibres.
Modification of conventional platinum, gold or iridium oxide electrodes with conducting polymer coatings has the potential to significantly improve the long-term performance of neural implants, including cochlear implants, vision prostheses, neural regeneration devices and neural recording electrodes (Green et al., 2009). Implantable electrodes can be used for the treatment of different disabilities and neurological disorders and can be used either to electrically elicit neural impulses or to record neuron signalling (Asplund et al., 2009). Asplund et al. (2009) coated platinum electrodes with PEDOT and investigated the biocompatibility of resultant electrode. L929 fibroblasts and human neuroblastoma SH-SY5Y cell lines were used for in vitro cell culture study and further in vivo study was carried out using polymer-coated implants in rodent cortex. Their results indicated that platinum electrodes coated with PEDOT were non-cytotoxic and showed no marked differences in immunological response in cortical tissue compared to pure platinum controls. Green et al. (2009) used anionically modified laminin peptides, such as DEDEDYFQRYLI and DCDPGYIGSR, to dope PEDOT electrodeposited on platinum electrodes and assessed the bioactivity of incorporated peptides and their effect upon nerve cell growth (PC12 cells). Their results demonstrated that large peptide dopants produced softer PEDOT films with a minimal decrease in electrochemical stability, compared to conventional dopants, and longer neurite outgrowth was observed on PEDOT films doped with synthetic anionic laminin peptides than that on PEDOT films doped with conventional paratoluene sulphonate dopant. Peramo et al. (2008) investigated a method to chemically deposit PEDOT on acellularized muscle tissue constructs. Their results revealed that in situ polymerization occurs throughout the tissue, converting it into an extensive acellular, non-antigenic substrate which will be crucial for in vivo experiments related to nerve repair and bioartificial prostheses.
The interactions between neural cells and PEDOT for the development of electrically conductive biomaterials intended for direct and functional contact with electrically active tissues, such as the nervous system, heart and skeletal muscle, was also studied by Richardson-Burns et al. (2007). They polymerized PEDOT around living cells and described a neural cell-templated conducting polymer coating for microelectrodes and a hybrid conducting polymer–live neural cell electrode.
4.2.2. Carbon nanotubes
Carbon nanotubes (CNTs) are another group of conducting polymers incorporated into non-conducting polymers to provide structural reinforcement and impart novel properties, such as electrical conductivity, into the scaffolds and to direct cell growth (Harrison and Anthony, 2007).
Some studies have indicated that carbon nanotubes are cytotoxic, while others have shown nanotubes to be excellent substrates for cellular growth (Harrison and Anthony, 2007). A closer look at the published literature on this issue reveals the resultant positive or negative reports on cytotoxicity to be due to the manner in which CNTs have been used in the experiments by different researchers. CNTs are shown to be toxic to cells when used as a suspension in cell culture media in any given experiment, while they appear to be non-toxic if immobilized to a matrix or to a culture dish (Hussain et al., 2009). Potential cytotoxic effects associated with carbon nanotubes may be mitigated by chemically functionalizing their surfaces (Harrison and Anthony, 2007). Chemically functionalized CNTs have been used successfully as potential devices to improve neural signal transfer while supporting dendrite elongation and cell adhesion. These results strongly suggest that the growth of neuronal circuits on a CNT grid is accompanied by a significant increase in network activity. The increase in the efficacy of neural signal transmission may be related to the specific properties of CNT materials, such as the high electrical conductivity controlling the neuronal extracellular–molecular interactions (Lovat et al., 2005; Hu et al., 2004, 2005; Mattson et al., 2000). In fact, CNTs represent a scaffold composed of small fibres or tubes that have diameters similar to those of neural processes such as dendrites (Lovat et al., 2005). Interest in carbon nanofibres has also been growing exponentially due to their unique electrical, mechanical and surface properties. Webster et al. (2004) developed a carbon nanofibre-reinforced polycarbonate urethane composite in an attempt to determine the possibility of using carbon nanofibres as either neural or orthopaedic prosthetic devices. Their results showed that this composite supports neural cell function and has the ability to tailor electrical properties for polyurethane composites containing carbon nanofibres and represent acceptable parameters for further investigation of carbon nanofibres as a neural probe. However, there are several limitations to the application of carbon nanotubes being non-degradable (Harrison and Anthony, 2007).
4.3. Piezoelectric polymeric materials
In addition to conductive polymers, piezoelectric polymeric materials have also been considered in tissue-engineering applications. Piezoelectric polymeric materials generate transient surface charges by tiny mechanical deformations of the material under minute mechanical strain and do not require additional energy sources or electrodes (Schmidt et al., 1997; Valentini et al., 1992). Poly(vinylidenefluoride) (PVDF) is a synthetic, semi-crystalline polymer with piezoelectric properties that generate transient surface charges due to their unique molecular structure (Valentini et al., 1992). Tschoeke et al. (2008) fabricated a textile scaffold composed of PVDF which was integrated into the wall of vascular composite graft to provide sufficient mechanical support. A two-step moulding technique was carried out to integrate a PVDF mesh in the wall of a fibrin-based vascular graft seeded with carotid myofibroblasts. Cell growth and tissue development within the fibrin gel matrix surrounding the PVDF fibres was excellent, and the tissue structure demonstrated much similarity to the native tissue. Yang et al. (2010) demonstrated that the epithelial–mesenchymal interaction of salivary tissue could be mediated using PVDF membrane in a serum-free condition. Their result demonstrated that it is possible to establish a serum-free system that is competent in facilitating epithelial–mesenchymal interaction of salivary tissue and facilitated by PVDF, the submandibular gland recombinant was able to sprout new branches without serum (Yang et al., 2010). Lu et al. (2003) designed a PVDF surface coated with galactose-tethered pluronic polymer and investigated the efficacy of attachment of rat hepatocytes on PVDF membrane. Their results showed the potential of galactose-immobilized PVDF membrane as a suitable substrate for hepatocyte culture. Young et al. (2008) fabricated microporous PVDF membranes for nerve tissue engineering by immobilizing L-lysine covalently on the surface of PVDF membrane. PC12 cells cultured on L-lysine/PVDF membranes showed good cell adhesion and proliferation on L-lysine/PVDF membranes, suggesting its usefulness in the development of strategies to promote the regrowth and regeneration of nervous tissue. Previous studies showed that neurite extension is significantly enhanced on piezoelectric materials such as PVDF and are comparable to those obtained using electrically conducting polymers (Schmidt and Leach, 2003). Aebischer et al. (1987) fabricated piezoelectric nerve guidance channels using PVDF and evaluated it in a transected mouse sciatic nerve model. Their results showed that piezoelectric nerve guidance channels enhance peripheral nerve regeneration and provide a tool to investigate the influence of electrical activity on nerve regeneration. In yet another study, Fine et al. (1991) synthesized tubular nerve guidance using a vinylidenefluoride–trifluoroethylene copolymer synthesized by a melt–erosion process. Piezoelectrically active vinylidenefluoride–trifluoroethylene copolymer tubes were found to significantly enhance the nerve regeneration process. Enhanced neurite outgrowth has been attributed to the presence of their surface charges and transient charge generation (Schmidt et al., 1997; Aebischer et al., 1987). The use of intrinsically charged piezoelectric polymers as tissue culture substrates provide a means of exposing cells directly to local time-varying charges, such as those produced by elements of ECM (Valentini et al., 1992).
Conductive polymers, however, exhibit many advantages over piezoelectric materials. Compared to piezoelectric materials, conductive polymers generate electrical signals by electron transfer between different polymer chains and allow external control over the level and duration of stimulation, which is beneficial for biomedical applications. Moreover, conducting polymers are inexpensive, easy to synthesize, versatile and their properties can be readily modulated by a wide range of molecules that can be entrapped and do not require extensive processing to render them electroactive (Dust and Bawornluck, 2006; Rivers et al., 2002). Although piezoelectric materials such as PVDF can aid in localized stimulation, the external control over stimulation is limited. The possibility that electrical stimulation can be externally controlled and precisely regulated with electrically conducting polymers, along with the flexibility of synthesizing electrically conducting polymers, offer many advantages of conductive polymers compared to piezoelectric materials (Schmidt et al.,1997).
5. Modification of conductive polymers for tissue-engineering applications
Although conducting polymers offer many advantages over other materials because of their electrical properties, opportunities to further optimize these materials when targeting tissue engineering application remains a challenge (Guimarda et al., 2007). In the field of biomaterials the design of bioactive surfaces is of particular interest, since biological systems interact with biomaterials via the interface and a variety of surface modification and immobilization methods have been developed to create surfaces having bioactive ligands to interact with biomolecules and cells (Cen et al., 2002). Several attempts have been made to combine the conductivity and biocompatibility of conducting polymers, such as doping by biological dopants, incorporation of bioactive molecules for increasing cell adhesion and proliferation, patterning of conductive scaffolds for improvement of their surface topography and surface modification of conducting polymeric materials with biological moieties (Cen et al., 2002; Lee et al., 2006; Song et al., 2006; Stauffer and Cui, 2006; Gomez et al., 2007). Figure 5 summarizes the different modification methods that have been carried out for the modification of conductive biomaterials.
5.1. Modification of conductive polymers with bioactive molecules
At present, ECM molecules are mainly chosen as dopants under the premise that the resulting polymer will have a higher binding affinity for cell adhesion molecules (Gelmi et al., 2010). By selecting bioactive molecules as dopants, the polymers can be modified for specific functionality through the incorporation of proteins, peptides or ECM components (Gilmore et al., 2009). Doping of conductive polymers using bioactive molecules such as heparin, dextran sulphate, hyaluronic acid, chitosan, collagen, growth factors, oligodeoxyguanylic acids and ATP are being carried out for the modification of conductive polymers (Ateh et al., 2006; Stauffer and Cui, 2006; View et al., 2008; Richardson et al., 2007). However, the doping of biomolecules has some limitations, such as low loading and decrease in conductivity, and for induced release cases, with polymer reduction the supply becomes limited and the release occurs faster (Gomez and Schmidt, 2007b). Gelmi et al. (2010) and Gilmore et al. (2009) also showed that doping with a bioactive polymer could cause an inverse effect on the physical properties, such as surface roughness of the final composite material, depending on the kind of doping molecules. To overcome these limitations, surface immobilizations of macromolecules have been explored by a few researchers. Immobilization of biomolecules into conductive scaffolds can be achieved via methods such as adsorption, entrapment and covalent binding (Cen et al., 2002). To modify a surface by covalent modification, a reactive group must be present on the polymer, which is not always the case in many polymers (e.g. PPy; Cen et al., 2004). Lee et al. (2006) synthesized carboxylic acid-functionalized conductive PPy and further grafted RGD on the surface of PPy scaffold. Human umbilical vascular endothelial cells cultured on PPy-RGD scaffolds demonstrated improved attachment and proliferations. Cen et al. (2002, 2004) fabricated an electroactive PPy surface functionalized with biologically active hyaluronic acid for the development of new tissue-engineering strategies, such as wound healing and angiogenesis, and further investigated the in vitro bioactivity of functionalized PPy film using PC12 cells. Their results showed that cell attachment on a hyaluronic acid functionalized PPy film surface was significantly enhanced in the presence of nerve growth factor. Li et al. (2004) immobilized heparin on the surface of PPy film covalently via poly(ethylene glycol) methacrylate graft copolymerization. The immobilization of heparin on the PPy surface had a significant effect on the selective adsorption of plasma proteins, albumin and fibrinogen and consequently the formation of thrombus. Functionalized PPy with heparin has the highest ratio of albumin:fibrinogen adsorption, the lowest level of plasma adsorption and the lowest amount of thrombus formation, which increased the blood compatibility of PPy film.
Sanchvi et al. (2005) used phage display to select peptides that specifically bind to PPy and which can subsequently be used to modify the surface of PPy. The use of selected peptides for PPy by phage display can be extended to encompass a variety of therapies and devices, such as PPy-based drug delivery vehicles, nerve guidance channel conduits and coatings for neural probes. Selection of peptides using phage display thus represents a simple alternative to methods based on electrostatic and hydrophobic interactions between two moieties to achieve adsorptive modification of surfaces.
Addition of bioactive factors to conducting polymers is considered a major strategy in improving cell–tissue interactions (Green et al., 2008). Electrically conductive and biologically active scaffolds are desirable for enhancing the adhesion, proliferation and differentiation of various cell types, including neurons. The incorporation of bioactive molecules such as neuroactive molecules into conductive scaffolds has also been applied for modification of conductive scaffolds (Lee et al., 2009a). Kim et al. (2007) incorporated nerve growth factor (NGF) as a co-dopant in the electrochemical deposition of PPy and PEDOT. An in vitro study using nerve cells showed that the incorporation of NGF can modify the biological interactions of the conductive polymers. However, direct covalent attachment of moieties to the backbone of a conducting polymer usually has an adverse effect on the electrical properties of the polymer (Thompson et al., 2010). Lee et al. (2009a) developed copolymers of N-hydroxyl succinimidyl ester pyrrole (PPy-NSE) and further modified it, using immobilization of NGF molecules on PPy-NSE films. Immobilized NGF on PPy copolymer was studied under physiological conditions and with the application of an external electrical potential. Their results showed that PC12 cells extended neurites similar to cells cultured in NGF-containing medium. Application of an external electrical potential to NGF-immobilized PPy films did not cause any significant release of NGF or reduce their neurotrophic activity. Such novel scaffolds, providing electroconductive and neurotrophic activities, has potential for neural applications, such as tissue engineering and biosensor applications. Thompson et al. (2010) incorporated neurotrophin-3 (NT-3) and brain-derived neurotrophic factor (BDNF) into PPy during electrosynthesis and evaluated neurite outgrowth from cochlear neural explants grown on PPy containing NT-3 and BDNF. Neurite outgrowth from explants grown on polymers containing both NT-3 and BDNF showed significant improvement over PPy doped specifically with NT-3, due to the synergistic effect of the two neurotrophins. Lee et al. (2006a) synthesized carboxy-endcapped polypyrrole (PPy-α-COOH), composed of a polypyrrole (PPy) surface modified with pyrrole-α-carboxylic acid, and further grafted it with the cell-adhesive Arg–Gly–Asp (RGD) motif. Human umbilical vein endothelial cells (HUVECs) cultured on RGD-modified PPy-α-COOH demonstrated significantly higher adhesion and spreading on the RGD-modified PPy-α-COOH than negative PPy-α-COOH without RGD motif and unmodified PPy. Cullen et al. (2008) fabricated small diameter (<400 µm) fibres consisting of a blend of electrically conductive PANI and PPy and carried out bio-adhesive surface modification using either poly-L-lysine (PLL) alone or PLL combined with collagen type I.
5.2. Modification of conductive materials with biopolymers
Fabrication of hybrid material has also been investigated as a method for modifying the properties of conductive polymers. The biocompatibility of electroactive polymers is commonly improved by blending them with natural polymers (Huang et al., 2007). For example, Khor et al. (1995) prepared a chitosan–PPy hybrid biomaterial by chemical polymerization of pyrrole in the presence of 2% chitosan solution in aqueous acetic acid under anhydrous conditions. However, no cell culture experiments were carried out by these researchers, although the chitosan–PPy hybrid appeared a promising material for tissue-engineering applications. Collagen–PPy hybrid material prepared by the chemical polymerization of pyrrole, using FeCl3 as initiator, was also carried out by these researchers in the presence of soluble collagen (Li and Khor, 1994). Huang et al. (2010) fabricated a PPy–chitosan membrane, seeded Schwann cells on the scaffolds and showed that the PPy–chitosan membranes supported cell adhesion and proliferation, indicating its biocompatibility. We carried out the electrospinning of a blend of poly(ε-caprolactone) (PCL), gelatin and PANI to obtain conductive nanofibrous scaffolds for nerve tissue engineering. This scaffold had the advantages of utilizing both a synthetic polymer where the PCL domain provides mechanical strength and a natural polymer (gelatin) that supported cell adhesion and proliferations. Moreover, the incorporated PANI made the nanofibrous scaffold a conductive material suitable for the electrical stimulation required for enhanced nerve regeneration (Ghasemi-Mobarakeh et al., 2009). On the other hand, Li et al. (2006) blended PANI and gelatin, fabricated scaffolds and investigated the attachment and proliferation of H9c2 cardiac myoblasts on PANI–gelatin scaffolds. Their results showed that the incorporation of PANI into gelatin has the advantages of inducing electrical conductivity to the scaffolds with improved cell attachment and proliferation due to the presence of gelatin.
Akkouch et al. (2010) incorporated fibronectin and bovine serum albumin into PPy through water-in-oil emulsion polymerization to create bio-activated conductive PPy nanoparticles. Conductive biodegradable membranes were prepared by blending bioactive PPy nanoparticles with poly(L-lactide) and showed the ability of these membranes to modulate cell adhesion and proliferation due to the presence of the bioactive proteins on their surfaces.
5.3. Topographical modification of conductive scaffolds
The topographical features of a scaffold have a significant effect in tissue engineering and it has been demonstrated that the surface roughness of scaffolds affects cell behaviour, including cell attachment, proliferation and differentiation (Chen et al., 2007; Naji and Harmand, 1990; Meyer et al., 1993; Steele et al., 1993; Chu et al., 1999; Xu et al., 2004b). Moreover, the topographical features are described to have a positive effect on axonal orientation, where the physical guidance of axons is a vital component of nerve repair. Song et al. (2006) fabricated wide PPy microchannels using electron-beam (e-beam) lithography and electropolymerization, and further developed a method for directing the axonal guidance to a desired location in the PPy supporting matrix. Gomez et al. (2007) also synthesized PPy microchannels electrochemically to fabricate electroconductive, topographical substrates for neural interfacing and found that PPy microchannels facilitated axonal establishment of rat embryonic hippocampal neurons.
6. Electrospun conductive polymers for nerve tissue engineering
Scaffolds suitable for tissue engineering should mimic the structural and biological function of native ECM as much as possible (Ma et al., 2005; Ashammakhi et al., 2008; Ma, 2004; Smith and Ma, 2004). Collagen is a major natural ECM component and has a fibrous structure with fibre bundles varying in diameter (50–500 nm) (Ma et al., 2005; Ashammakhi et al., 2008; Ma, 2004; Smith and Ma, 2004). Nanoscale dimensions have been shown by various researchers to influence cell behaviour. Cells attach and organize around fibres with diameters smaller than that of the cells (Ma et al., 2005). Nanoscale features have the potential to improve the specificity and accuracy of the materials for a number of neural engineering applications, ranging from neural probes to guidance channels for nerve regeneration (Seidlits et al., 2008). Nanofibrous scaffolds having fibres in the nanometer scale can therefore serve as a potential substrate for cell attachment, function and proliferation rather than the traditional scaffolds (Ma et al., 2005; Ashammakhi et al., 2008; Ma, 2004; Smith and Ma, 2004). Polymeric nanofibres can be processed by a number of techniques, such as drawing, template synthesis, phase separation, self-assembly and electrospinning (Ashammakhi et al., 2008; Smith and Ma, 2004). Electrospinning is one of the most important techniques to fabricate nanofibres. An attractive feature of electrospinning is the simplicity and inexpensive nature of the set-up (Seidlits et al., 2008). The structure and surface morphologies of electrospun nanofibres can be controlled by adjusting various parameters during electrospinning, such as the solution properties, applied voltage and spinning conditions (Ramakrishna and Fujihara, 2000; Pham et al., 2006). During the electrospinning process, a strong electrostatic field is applied to a polymer solution and, when the electric forces overcome the surface tension of the solution, a charged jet of solution is ejected towards the collecting material screen, where a continuous stretch of nanofibres is obtained (Zong et al., 2002; Zeng et al., 2003; Yarin et al., 2001). Electrospun nanofibrous scaffolds of PCL, PCL–gelatin, poly(L-lactic acid), PCL–collagen, PCL–chitosan, poly(lactic acid-co-glycolic acid) have been used for nerve tissue engineering (Yang et al., 2004; Xu et al., 2004; Bini et al., 2006; Schnell et al., 2007; Yang et al., 2005; Murugan and Ramakrishna, 2007; Lee et al., 2009c). To take advantage of the nanofibrous structures together with the electrical stimulations in tissue engineering, conductive nanofibrous scaffolds were applied for nerve tissue engineering. However, the major obstacle concerning the electrically conductive polymers has been the difficulty associated with the processing of these materials (Pomfret et al., 2000). To overcome this problem, most researchers have electrospun conductive polymers by blending conducting polymers with other spinnable polymers, compromising the conductivity of the composite fibres (Yu et al., 2008; Norris et al., 2000; Veluru et al., 2007; Bishop and Gouma, 2005; Desai et al., 2004; Ju et al., 2007; Srivastava et al., 2007; Liu et al., 2008; McKeon et al., 2010). Blending of conductive polymers with other polymers affects the properties of the resultant nanofibres. For example, previous studies showed that the incorporation of conductive polymers to other polymeric materials on electrospinning decreased the fibre diameters of the resultant nanofibres, mostly due to increasing of conductivity of the electrospun solution. As mentioned earlier, we blended PANI with PCL–gelatin solution and fabricated conductive nanofibrous scaffolds suitable for nerve tissue engineering. Figure 6 shows the morphology of PCL–gelatin (PG) and PANI–PG at a ratio of 15:85 w/w nanofibrous scaffolds fabricated during our study.
Fibre diameter was found to decrease with the incorporation of PANI in the polymeric blend system, whereby PG and PANI/PG (15:85) were found to have fibre diameters of 189 ± 15 nm and 112 ± 8 nm, respectively, by SEM analysis (Ghasemi-Mobarakeh et al., 2009). We also fabricated poly(L-lactic acid) (PLLA) and PLLA/PANI nanofibres and observed a similar trend in their fibre diameters. Figure 7 shows the SEM images of PLLA and PLLA/PANI nanofibres. Fibre diameter decreased from 860 ± 110 nm for pure PLLA nanofibres to 175 ± 55 nm for PLLA/PANI nanofibres.
Li et al. (2006) also showed that the incorporation of PANI into gelatin solution decreased the fibre diameter from 803 ± 121 nm for pure gelatin fibres to 61 ± 13 nm for 60:40 PANI-gelatin blend fibres. In yet another study, Jeong et al. (2008) showed that by increasing the volume ratios of PANI in poly(L-lactide-co-ε-caprolactone) (PLCL) solution, the diameters of resultant nanofibres decreased from 430 ± 116 nm for pure PLCL to 382 ± 102 nm for PANI/PLCL nanofibres. Table 2 shows the different conductive nanofibrous scaffolds that have been fabricated for various tissue-engineering applications.
Table 2. Biocompatibility of conductive nanofibrous scaffolds utilized for tissue engineering applications
7. Application of electrical stimulation in nerve tissue engineering
Published reports on the application of electrical stimulation for nerve tissue engineering are limited. In vitro assays conducted by Schmidt et al. (1997) showed that PC12 cells cultured on PPy films and subjected to electrical stimulus showed a significant increase in neurite length (18.14 µm) compared to passive controls (9.5 µm). Subsequent studies by Schmidt and colleagues indicated that the longer neurites grown on PPy were due to increased protein adsorption from serum-containing medium mediated by electrical stimulation (Kotwal and Schmidt, 2001). We applied electrical stimulation to nerve cells through conductive nanofibrous scaffolds of PANI–PCL–gelatin (PANI–PG) and found significantly enhanced cell proliferation and neurite outgrowth compared with non-stimulated scaffolds (Ghasemi-Mobarakeh et al., 2009). The average neurite length for nerve cells grown on PANI–PG, with application of electrical stimulation by DC voltage of 100 mV/mm for 1 h and without electrical stimulation, was found to be 30 ± 1.1 µm and 22 ± 0.97 µm, respectively (n = 200), indicating that the application of electrical stimulus for 1 h to nerve cells cultured on PANI–PG significantly (p ≤ 0.05) enhanced neurite outgrowth. Figure 8 demonstrates the effectiveness of electrical stimulation on neurite length, showing cells exposed to electrical stimulation for 1 h with extended neurites, compared to control samples without electrical stimulation (Ghasemi-Mobarakeh et al., 2009).
Zhang et al(2007) produced poly(D,L-lactide-co-ε-caprolactone) membrane coated with PPy and the composite scaffolds supported the proliferation and differentiation of PC12 into neuronal phenotypes as well as sciatic nerve regeneration in rats. Their study demonstrated that PPy-coated poly(D,L-lactide-co-ε-caprolactone) membranes could successfully be used for electrical stimulation and it enhanced the neurite outgrowth in a current-dependent fashion. Kimura et al. (1998) applied alternative potential which generated rectangular pulse waves to PC12 cells cultured on the ITO electrode. Their results showed that PC12 cells were electrically induced to extend their neurites on the electrode surface, even in the absence of nerve growth factor (NGF). They concluded that electric stimulation induced c-fos expression, which is essential for cell differentiation, and alternative potential may stimulate cell differentiation through a protein kinase C (PKC) cascade. Lee et al. (2009b) produced PPy-coated electrospun PLGA nanofibres (PPy–PLGA) by nano-thick deposition of PPy on electrospun PLGA fibres for neural tissue applications and performed electrical stimulation of PC12 cells on these cytocompatible electroconductive nanofibres. In vitro cell culture using PC12 cells and embryonic hippocampal neurons demonstrated compatible cellular interactions, and the fabricated PPy–PLGA meshes were found appropriate for neuronal applications. Their results also showed that the electrical stimulation of PC12 cells on conducting nanofibre scaffolds improved neurite outgrowth compared to non-stimulated cells. Xie et al. (2009) fabricated a new type of scaffold comprised of conductive core-sheath nanofibres prepared by in situ polymerization of pyrrole on electrospun PCL or poly(L-lactide) (PLA) nanofibres, to investigate the synergistic effect of topographic cue and electrical stimulation on axonal regeneration from cultured neuronal populations. For electrical stimulation, a mat of random or random–aligned–random nanofibres was fixed to a custom-built culture plate with silver electrodes connecting the two ends. A constant voltage of 10 V was then applied across the mat for 4 h/day during cell culture. The effective direct current (DC) applied to the sample was estimated to be 250 mA (a value within the range 0.6–400 mA). Their findings indicated that the electrical stimulation enhanced the rate of neurite extension on both random and aligned nanofibres. Thompson et al(2010) also showed that neurite outgrowth from cochlear neural explants grown on the PPy-containing NT-3 and BDNF was significantly improved when the polymer containing both neurotrophins was electrically stimulated. Electrical stimulation of protein-loaded PPy has a dramatic effect on auditory nerve survival and growth from cochlear neural explants. They concluded that electrical stimulation using clinically acceptable waveforms has a significant effect on the rate of release of both proteins (NT-3 and BDNF). Electrical stimulation has also been applied to Schwann cells cultured on polypyrrole–chitosan membrane (Huang et al., 2010). Constant potential gradient (100 mV/mm) was applied to the cells through the membrane for 4 h and then the cells were incubated for an additional 12, 24 or 36 h. The results showed that electrical stimulation applied through conductive PPy–chitosan significantly enhanced the viability of Schwann cells and dramatically enhanced secretion of NGF and brain-derived neurotrophic factor (BDNF) and protein expression. Huang et al. (2008) carried out electrical stimulation of PC12 cells through conductive film fabricated from a copolymer of PLA and carboxyl-capped aniline pentamer. Their findings showed that the mean neurite length of PC12 cells on conductive film exposed to electrical stimulation were much higher than TCP exposed to electrical stimulation and conductive films without stimulation.
Electrical stimulation has also been used for other cell types. Shi et al. (2007) found the application of an electrical field to conductive biodegradable PPy–PLA membranes to be an effective approach to upregulate the mitochondrial activity of human skin fibroblasts. These researchers also fabricated a conductive biodegradable composite material made of PPy nanoparticles and poly(D,L-lactide), applied electrical stimulation after seeding fibroblasts on them and found similar upregulations (Shi et al., 2004). Jeong et al. (2008) fabricated nanofibrous scaffolds containing PANI and poly(L-lactide-co-ε-caprolactone) and investigated the behaviour of human dermal fibroblasts, NIH-3T3 fibroblasts and C2C12 myoblasts on this scaffold. The growth of NIH-3T3 fibroblasts was enhanced by stimulation of various DC flows. Sun et al. (2006) applied a non-invasive electrical stimulus to regulate the cell adhesion and orientation of bone marrow-derived mesenchymal stem cells (MSCs) and fibroblasts. Their results showed that the optimal application of electrical stimulus could offer a novel engineering technique to regulate the cell type-dependent cellular shape and orientation that are known to be involved in cell differentiation and growth.
8. Conclusion and future remarks
The nervous system responds to electrical fields and the key component of neural communication is the action potential generated at the synapse. This implies that for fabrication of an ideal neural scaffold, introduction of electrical stimulations through conductive scaffolds could promote neurite outgrowth and nerve regeneration.
The importance of conductive polymers for the fabrication of conductive scaffolds and their applications in electrical stimulation are known. Fabrication and modification of conductive nanofibrous scaffolds by the incorporation of biological molecules via different methods have been carried out to improve cell attachment and proliferation on these conductive scaffolds. Here, we report that electrical stimulation strongly influences nerve regeneration and describe some in vitro application of electrical stimulation through conductive scaffolds. Although several in vitro methods have been used for electrical stimulation, clinical application of electrical stimulation using conductive scaffolds has not been attempted successfully. Further study is required to fully understand, characterize and concurrently compare the effects of different conditions of electrical stimulation (intensity and duration) on neurons. Moreover the mechanisms caused by electrically stimulated cells need further investigation, and in vivo electrical stimulation using implanted conductive scaffolds must be carried out for deeper understanding and efficacy evaluations in nerve tissue engineering.