Biomaterials 31 (2010) 2701–2716
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Review
Electroconductive hydrogels: Synthesis, characterization and biomedical applications Anthony Guiseppi-Elie a, b, c, * a
ABTECH Scientific, Inc., Biotechnology Research Park, 800 East Leigh Street, Richmond, VA 23219, USA Center for Bioelectronics, Biosensors and Biochips (C3B), Clemson University Advanced Materials Center, 100 Technology Drive, Anderson, SC 29625, USA c Department of Chemical and Biomolecular Engineering, Department of Bioengineering, Department of Electrical and Computer Engineering, Clemson University, Clemson, SC 29634, USA b
a r t i c l e i n f o
a b s t r a c t
Article history: Received 1 November 2009 Accepted 18 December 2009 Available online 8 January 2010
Electroconductive hydrogels (ECHs) are composite biomaterials that bring together the redox switching and electrical properties of inherently conductive electroactive polymers (CEPs) with the facile small molecule transport, high hydration levels and biocompatibility of cross-linked hydrogels. General methods for the synthesis of electroconductive hydrogels as polymer blends and as polymer co-networks via chemical oxidative, electrochemical and/or a combination of chemical oxidation followed by electrochemical polymerization techniques are reviewed. Specific examples are introduced to illustrate the preparation of electroconductive hydrogels that were synthesized from poly(HEMA)-based hydrogels with polyaniline and from poly(HEMA)-based hydrogels with polypyrrole. The key applications of electroconductive hydrogels; as biorecognition membranes for implantable biosensors, as electrostimulated drug release devices for programmed delivery, and as the low interfacial impedance layers on neuronal prostheses are highlighted. These applications provide great new horizons for these stimuli responsive, biomimetic polymeric materials. Ó 2009 Elsevier Ltd. All rights reserved.
Keywords: Hydrogels Polypyrrole Polyaniline Blends Co-networks Electrically conductive hydrogel
1. Introduction Electroconductive hydrogels (ECHs) are polymeric blends or conetworks that combine inherently conductive electroactive polymers (CEPs) with highly hydrated hydrogels. First described by Guiseppi-Elie [1–3] in 1995 and later by Wallace et al. [4] and Guiseppi-Elie et al. [5], these polymeric materials portend the combination of the unique and enduring properties of their constituents. For the hydrogel component this implies a high degree of hydration, swellability, in vitro and in vivo biocompatibility and the high diffusivity of small molecules. For the inherently conductive electroactive polymer component this implies high
Abbreviations: ECH, electroconductive hydrogels; CEP, conductive electroactive polymers; IME, interdigitated microsensor electrode; MDEA, microdisc electrode array; QCM, quartz crystal microbalance; PPy, polypyrrole; OPPy, overoxidized polypyrrole; PAn, polyaniline; PTh, polythiophene; HEMA, 2-Hydroxyethylmethacrylate; NPD, neural prosthetic device; NRD, neural recording device; ESDRD, electro-stimulated drug release device; ISFET, ion-sensitive field effect transistors; PVP, polyvinyl pyrrolidone. * Center for Bioelectronics, Biosensors and Biochips (C3B), Clemson University Advanced Materials Center, 100 Technology Drive, Anderson, SC 29625, USA. Tel.: þ1 864 656 1712; fax: þ1 864 656 1713. E-mail address:
[email protected] 0142-9612/$ – see front matter Ó 2009 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2009.12.052
electrical conductivity, ON–OFF electrical and optical switching and electrochemical redox properties. Both polymer constituents are stimuli responsive materials and each on its own is a viable candidate for sensory and actuation applications. They share in common the potential for molecular engineering through copolymerization, crosslinking and/or grafting to tune the final hybrid material’s properties to targeted biologically relevant outcomes. They also share the fact that their syntheses are difficult to control and that the methods and conditions of synthesis can greatly influence the final material properties via subtle contributions of multi-length scale factors such as trace impurities, conformation, nano-structure and gross morphology. Since the early work of Guiseppi-Elie and Wallace, ECH have been the subject of growing attention [6]. This paper reviews the current status of these hybrid polymeric materials from the standpoint of their synthesis, characterization and key biomedical applications. Electroconductive hydrogels belong to the general class of multifunctional smart materials conceptually illustrated in Fig. 1. As an emergent class, these materials seek to creatively combine the inherent properties of constituent materials to give rise to technologically relevant properties for devices and systems. For example, electroconductive hydrogels have been synthesized, characterized and fashioned as a biorecognition membrane layer in
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Fig. 1. Illustration of the generalized concept of multi-functional smart materials that combine the properties of constituent materials to yield technologically relevant devices and systems of increasing complicatedness and possible complexity.
various biosensors. In one instance an electroconductive hydrogel that was synthesized from a poly(HEMA)-based hydrogel and poly(aniline) was fashioned into a biosensor by the incorporation of recombinant cytochrome P450-2D6 [7]. This device was shown to be responsive to fluoxetine, the active ingredient in ProzacÒ. The poly(HEMA)-based poly(aniline) electroconductive hydrogel was subsequently fully characterized for its electrical, switching and optical properties with demonstrated faster switching than its purely CEP counterpart [8]. In another instance, an electroconductive hydrogel fashioned from poly(hydroxyethyl methacrylate) [poly(HEMA)] and polypyrrole (PPy) was investigated for its potential application to clinically important biomedical diagnostic biosensors by the incorporation of analyte-specific enzymes [9–12]. Among the various devices for which electroconductive hydrogel polymers were investigated were; neural prosthetic and recording devices (NPDs and NRDs) [13–15] electro-stimulated drug release devices (ESDRDs) [16–20] and implantable electrochemical biosensors [21–23]. Neural prosthetic devices, drug eluting stents and implantable biosensors of the central and peripheral nervous system benefit from a tissue-to-electrode interface that is compliant, conformal, of low interfacial impedance and electrically diffuse [24–26]. The electro-stimulated drug release devices benefit from a high loading capacity and low voltage actuation. In all cases these polymeric materials, which are both electronically and ionically conductive, provide a non-cytotoxic interface between the device and native living tissue or cell culture medium [27]. The background on hydrogels and inherently conductive polymers as well as the materials rationale for the co-joining of hydrogels with conductive electroactive polymers is addressed. Details are provided on the methods and approaches for the synthesis of electroconductive hydrogels. The review then addresses three key biomedical applications of these materials; that of implantable electrochemical biosensors, electro-stimulated drug release devices and neural prosthetics. 2. Inherently conductive electroactive polymers [28] Conductive electroactive polymers (CEPs) are a family of highly conjugated polymers possessing spatially extended p-bonding that confers unique electrical, electrochemical and optical properties. These properties have been found useful for applications in a number of current and many emerging industrial applications.
Within this family are polymers such as polypyrrole (PPy), polyaniline (PAn), polythiophene (PTh), poly(phenylene) vinylene (PPv), their copolymers and derivatives. Listed in Table 1 are several CEPs with their known electrical conductivities. These polymers find applications as diverse as electrical conductors, nonlinear optical devices, polymeric light emitting diodes (LEDs), as electrochromic windows and displays, as photoresists, as antistatic coatings, as chemical and biological sensors, as corrosion protective coatings, as electrodes of batteries, as electromagnetic shielding materials, as sensory elements in electronic noses, as solar cells, as microwave absorbing materials, in nanoswitches, as optical modulators, and as valves in MEMs devices [29–32]. CEPs display controllable switching of their electrical and optical properties. Moreover, because these properties may also be altered by changes in ambient variables such as temperature, humidity, gas/vapor composition, medium ionic strength, pH, and the like, these materials are also regarded as stimuli responsive. The incorporation of CEPs as elements of the structure of biosensors, molecular recognition devices that employ biomolecules (peptide sequences, enzymes, antibodies, oligonucleotides, aptamers), promises to improve the response time of these devices and their compatibility with advanced microfabrication and miniaturization technologies for incorporation into MEMs devices [33]. CEPs have been incorporated into biosensors for the detection of several tens of chemical species of biological or medical importance (enzyme substrates, antigens, ssDNA fragments, neurotransmitters, drug metabolites). Biosensors based on CEPs may operate as electrochemical, optical or gravimetric detectors for measurements in discrete, low-volume samples and continuous flow systems for which biosensors with fast response times, high sensitivities, and detection limits in the mM range (for detection of enzyme substrates) and even several orders of magnitude lower (for detection of DNA) are required. For in vivo applications, biosensors have to meet additional biocompatibility specifications [34]. The chemistry of CEPs and the relative ease with which these materials may be synthesized and fashioned into functional devices makes them compatible with many of the chemical processes found in microelectronics manufacturing. Polypyrroles, for example, may be synthesized under mildly oxidative conditions from aqueous media. The syntheses may be via an oxidative wet chemical method using such mild oxidants as FeCl3, peroxy sulfate, or benzoyl peroxide. Hydrogen peroxide and potassium permanganate, while capable of initiating such polymerization reactions, may also oxidatively degrade the resulting polymer, but with appropriate controls, can also introduce new reactive functional groups. Polypyrroles may also be synthesized electrochemically from neutral to acidic media possessing an appropriate supporting electrolyte. Potentiostatic (chronoamperometry) oxidation may be initiated at 0.65 V vs. Ag/AgCl and some influence exercised over the kinetics by the application of potentials up to 0.8 V vs. Ag/AgCl. True kinetic control is however achieved galvanostatically (chronopotentiometry) by the application of a constant current (typically 1 mA/cm2). Likewise, polyaniline may be synthesized under similar conditions, although formation of the highly conductive form requires the presence of acid (pH 4). Electrochemical initiation of polyaniline synthesis may be initiated at 0.65 V vs. Ag/AgCl. Simple polythiophenes require a more extreme oxidation potential and these monomers are generally not soluble in water. However, the attachment of simple functional groups, e.g. an alkyl sulfonate, to the monomer overcomes this limitation. The consequence of initiation and propagation of polymerization under oxidizing conditions is the simultaneous incorporation of dopant counter anions drawn from the environment into the CEP leading to formation of its conductive form.
A. Guiseppi-Elie / Biomaterials 31 (2010) 2701–2716 Table 1 List of some conductive electroactive polymers, their structures and known electrical conductivities.
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and conductivity. An alternative approach is the post polymerization modification via electrochemical overoxidation to convert PPy to OPPy, with becomes enriched in hydroxyl groups [35,36]. The electronic, optical and redox properties of CEPs may be exploited in a form of information transfer wherein the chemical potential energy of an analyte is converted into a proportionate electronic, optical or electrochemical signal. In this sense these polymers are transducer-active as they allow information about the concentration of an analyte to be conveyed electronically (impedance; conductance, capacitance, switching), optically (reflectivity, absorbance, shifts) or electrochemically (amperometrically, potentiometrically) via biomolecular recognition events to produce an analytical signal. These aspects of CEPs have been thoroughly reviewed by Guiseppi-Elie et al. in the Handbook of Conductive Polymers [6]. CEPs, while possessing many highly beneficial properties for bio applications, do suffer some serious limitations. Amongst these are slow switching speeds in bioelectronic applications, the potential for unintended overoxidation leading to formation of reactive species, time-temperature drift of materials properties and questionable biocompatibility. The integration of CEPs with hydrogels promises to address these issues.
3. Hydrogels [37]
The mild conditions used for polymerization are ideal for the simultaneous synthesis of the CEP and incorporation of biological molecules, such as enzymes, antibodies, nucleic acids, sub-cellular fragments or even whole living cells. These conditions also support the incorporation of surfactant molecules and other polymers that may become entangled during the electrophoretic polymer deposition step. Such co-deposition is facilitated by a net or partial negative charge on the dopant molecule. The organic, polymeric nature of the CEP, as opposed to the inorganic, periodic nature of most semiconductors and metals, produces favorable molecular level interactions between the host CEP and biological molecules. These interactions may be further enhanced by the inclusion of specific functionalities or moieties on the CEP backbone. Amongst these are reactive functional groups such as carboxylic acids and alcohols, which allow direct covalent bond formation between the CEP and biomolecule, pendant anion groups such as sulfonates that allow ‘‘internal doping’’, or peptide sequences that facilitate receptor–receptand interactions of whole cells. Such reactive functional groups may be introduced via copolymerization; achieved through the inclusion of functionalized monomers into the reactions mixtures. For example, pyrrole may be copolymerized with 4-(3-pyrrolyl)butyric acid to introduce pendant acid functionality that does not adversely influence main chain conjugation
Polymer hydrogels are three-dimensional polymeric networks formed from highly hydrophilic monomers rendered insoluble by virtual, electrostatic or covalent crosslinking. Hydrogels imbibe large amounts of water. The result is an elastic network with water effectively filling the interstitial space of the network. When immersed and equilibrated in aqueous medium, cross-linked hydrogels assume their final hydrated network structure which brings into balance the forces arising from the solvation of the repeat units of the macromolecular chains that leads to an expansion of the network (the swelling force) and the counter balancing elastic force of the cross-linked structure (the retractive force) [38]. The imbibed water of a hydrogel may be free (freezable) or bound (non-freezable) [39]. The hydrogel can accordingly easily change its size and shape in response to environmental stimuli and this is one of its intrinsic characteristics; effectively expelling or imbibing free water. Moreover, in so doing, hydrogels can also imbibe other monomeric, reactive and potentially polymerizable species into its interstices, essentially occupying its void volume and interacting with chain segments or pendant moieties of the host hydrogel. Hydrogels have emerged since the early 50s as being of great importance in the biomaterials field [40]. Their unique characteristics, being of a soft elastomeric nature, serves to minimize mechanical and frictional irritation to the tissue bed, their low interfacial tension contributes to a reduction in protein adsorption and hence biofouling and cell adhesion, and their swelling capacity results in high permeabilities for low molecular weight drug molecules and metabolites [41]. These characteristics have allowed hydrogels to be used in biomedical applications that include biosensors, drug delivery systems, contact lenses, catheters, wound dressings and tourniquets. Of particular interest is their use as matrices for the immobilization and stabilization of enzymes [42– 46]. This interest has lead to their parallel development as the biorecognition layer of potentiometric, conductometric, amperometric and fiber-optic based enzyme biosensors [47,48]. Because of their high water content, hydrogel membrane layers and gel pads also find application as micro-bioreactors for the hosting and stabilization of biological molecules and for the conduct of biological reactions [49,50]. Hence, hydrogels have been used to host
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bioactive layers in several forms of enzyme-linked antibody biosensors and DNA biochips [51–55]. There are multiple reports involving hydrogels that document their biocompatibility [56], biodegradability [57], bioadhesion [58], dielectric relaxation [59] and mass transport properties [60]. The theoretical framework for the kinetics of the swelling of polymeric hydrogel has been well developed [61]. The ionizable polymer hydrogel gives rise to important electrochemically modulated electroosmotic swelling and deswelling characteristics and this has been explained in terms of electrokinetic processes. Hence, an ionized hydrogel changes volume discontinuously as the solvent composition is continuously varied (i.e. a phase transition occurs). This phase transition is induced not only by a change in the solvent composition but may also be induced by a change in pH, ionic strength (salt concentration), temperature and the application of an electric field. For example, a polyelectrolyte hydrogel when placed between a pair of electrodes deswells under an applied DC voltage with concomitant exudation of water. This property is associated with the electrophoretic and electroosmotic transport of highly hydrated macromolecules and of its counterions. The shape change of hydrogels has also engendered applications as hydrogel actuators in bioMEMS devices and as artificial muscles due to the hydrogel’s dramatic changes in physical dimensions caused by changes in the polarity of the electric field. These chemo-mechanical devices can allow dynamic control of delivery of drug and other solute molecules across and from within hydrogel membranes. Poly(HEMA)-based hydrogels are hydrolytically stable, may be engineered to possess similar water content and elastic moduli as body tissues, and exhibit good in vitro and in vivo biocompatibility [62,63]. Consequently, these polymer hydrogels have emerged as one of the most widely researched, patented, and successfully commercialized biomedical polymers [64]. There are numerous studies that aim to modify the properties of p(HEMA). These studies aim to improve the mechanical properties [65], transport properties [66], temperature responsive characteristics [67] and the degree of hydration [68,69]. The degree of hydration and/or swelling is one of the important properties that allows for an understanding of the transport of small molecule solutes through the hydrogel matrix. However, hydration also influences the elastic modulus and surface properties such as wettability - and consequently, protein adsorption – and is, thus, strongly correlated with in vitro and in vivo biocompatibility [70]. Apart from the poly (HEMA)-based hydrogels, other types of naturally occurring, synthetic and hybrid hydrogels have been proposed and studied for the immobilization of biorecognition molecules and whole cells. Amongst these are agarose [71–73], alginates [74–76], polyvinylalcohol [74], poly(acrylate) [77], collagen [77], albumin-PEG [78] and gelatin [73], chitosan [79] and hyaluronic acid [80]. Poly(HEMA)-based hydrogels are generally synthesized via the free radical polymerization of acrylate and methacrylate monomers. The free radical initiator may be excited thermally or via UV light. Typically, the hydrogel matrices are prepared by mixing monomers, crosslinker, pre-polymer, distilled water and ethylene glycol to obtain a homogenous solution of controlled viscosity and density to which the initiator is added. Varying the mole percentage composition and variety of monomers allows for synthesis of unique copolymers. Monomers of equally reactivity and diffusivity are expected to yield random copolymers. However, this condition is rarely ever met and so blocks should be expected, particularly formed during the later stages of the polymerization reaction. Controlling the mole percentage of the cross linker relative to the reactive monomer allows control of the molecular weight between cross-links and hence the void volume of the final hydrogel. Pre-polymers of medium to high molecular weight may be added to control the viscosity of the cocktail.
The hydrogel cocktail may be spun-applied to electrodes, electronic devices such as ion-sensitive field effect transistors (ISFETs), or surfaces of other materials to produce adherent thin films of thickness that may vary from 1 to 10 microns. Producing an adherent hydrogel thin film is a non-trivial matter as the reversible swelling and deswelling leads to appreciable interfacial shear stress and its hydrophilic nature leads to interfacial accumulation of water; both of which can lead to interfacial failure. Alternatively, the hydrogel cocktail may be free-standing, cast into a mould, deposited by any of the multiple coating processes, or prepared as microspheres via suspension or emulsion polymerization. UV polymerization is typically carried out at room temperature with brief exposures to the appropriate wavelength of light, generally with minimal increases in temperature, and with little or no compromise of biological activity. Bioactive hydrogels reflect an emerging paradigm in the development of responsive [81], multifunctional [82], biorecognition membrane layers for implantable biosensors and deep brain stimulation devices. Bioactivity is derived from a combination of hydration levels, mechanical properties, surface chemistries, and micro-nano-topologies that render the hydrogel mimetic of the tissue bed within which it is to be implanted. The design and molecular engineering of bioactive hydrogels as the recognition membrane layer of biosensors requires that bioactive molecules such as enzymes, their co-factors, redox mediators and biomimetic moieties be purposefully incorporated at bio-functionally relevant levels within the hydrogel [83]. Retaining these bioactive and or functional entities, either by physical entrapment or covalent tethering, necessitates an understanding of the transport characteristic of these entities within the hydrogel matrix. 4. Electroconductive hydrogels The term electroconductive is a contraction of electroactive and conductive. An electroconductive hydrogel describes a polymer that combines the properties of hydrogels and conductive systems and appears to have first originated with Gong et al. [84] who described a conductive charge transfer salt complex of 7,7,8,8-tetracyanoquinodimethane (TCNQ)-loaded hydrogel. Since the first papers of Guiseppi-Elie and Wallace, several studies on composites formed from CEP and hydrogels have been carried out. An electrically conductive composite material, consisting of polyaniline nanoparticles dispersed in a polyvinyl pyrrolidone (PVP) hydrogel was prepared by water dispersion polymerization of aniline using PVP as a steric stabilizer, followed by g-irradiation which induced cross linking of the PVP component [85]. Moschou et al. [86] developed an artificial muscle material based on a hydrogel that was composed of acrylamide and acrylic acid that was doped with a polypyrrole/carbon black composite. Lira et al. [16] synthesized polyaniline–polyacrylamide composites by electropolymerization of the conducting polymer inside an insulating hydrogel matrix of different pore sizes. The resulting new material was electroactive due to the polyaniline present inside the pores. These composites were applied to electrochemically controlled drug delivery devices. The synthesis of a hydrogel composite in which polyaniline (linear) was entrapped within a cross-linked polyelectrolyte hydrogel, poly (2-acrylamido-2-methyl propane sulphonic acid) (PAMPS) was reported by Kumar and Gangopadhyay in 2005 [87]. Nikpour et al. [88] synthesized conducting polymer composites of PPy with poly (methylmethacrylate), demonstrating the resulting material as controlled delivery devices. The liquid porogen used was polypropylene glycol while sodium chloride powder was used as the solid porogen. Koul et al. [89] reported on the synthesis of a polyaniline (acrylonitrile–butadiene–styrene) composite membrane as a sensor material for aqueous ammonia. The resistance change of
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the composite film upon exposure to different concentrations of aqueous ammonia showed its utility as a sensor material. Park and Park 2002 [90] investigated the electrical properties of the conducting composite poly (methylmethacrylate-co-pyrrolmethylstyrene)g-polypyrrole (PMMAPMS-g-PPy). The PMMAPMS-g-PPy was synthesized by the electrochemical reaction of PMMAPMS and pyrrole in the electrolyte solution containing lithium perchlorate and a mixture solvent of acetonitrile and dichloromethane. Enzymes entrapped within polypyrrole (PPy) films, prepared by electropolymerization from aqueous solution, have been commonly used to prepare electrodes. Brahim et al., 2002 [11] developed a glucose biosensor based on GOx entrapment within a composite p(HEMA)/PPy hydrogel membrane. A mixture of HEMA, tetraethylene glycol (TEGDA) as a crosslinker and enzyme was deposited on the platinum electrode surface and the polymerization of HEMA was performed by irradiating with UV under argon atmosphere. Subsequently, the pyrrole monomer entrapped within the hydrogel network was electrochemically polymerized. An amperometric biosensor for cholesterol analysis was prepared by entrapping cholesterol oxidase (ChlOx; E.C 1.1.3.6) onto the p(HEMA)/PPy matrix [9]. The bioactive composites of polypyrrole containing p(HEMA) hydrogels were incorporated into an amperometric biosensor for clinically important analytes (galactose [91], glucose and cholesterol). The biosensor formed from the overoxidized polypyrrole within the hydrogel showed excellent screening of physiological interferences, such as ascorbic acid, uric acid and acetaminophen. 5. Synthesis of electroconductive hydrogels Electroconductive hydrogels are composites, blends or conetworks of hydrogels and conductive electroactive polymers; for example, the hydrogels may be the continuous or dominant component within which the CEP is polymerized. Alternatively, the CEP may be the dominant or continuous component and the hydrogel polymerized within the CEP. To date experiments with the former type are more common. Polymer composites of inherently conductive polymers synthesized within other (host)
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polymers are not new. Early attempts to stabilize the environmentally labile optical and electrical properties of polyacetylene (CHx) were achieved by Zigler–Natta catalytic synthesis of CHx within host polymers, such as polyethylene [92,93]. Shown in Fig. 2 is a schematic illustration of the generalized synthetic routes to electroconductive hydrogels. The figure shows that reactive monomer, which may be a combination of both hydrogel precursors and CEP precursors, may be combined, along with photo or thermally labile free radical initiators, into a single cocktail or pre-polymer cocktail. The cocktail may be cast into film, prepared as microspheres, spun as fibers or spun-applied to electrodes and other solid-state electronic device substrates. Chemical oxidative polymerization does not require a substrate and can proceed when the hydrogel form (fiber, film or microsphere) is immersed within a suitable solution containing the initializing oxidant such as FeCl3 or peroxy sulfate. Electrochemical polymerization requires that the hydrogel form be applied to a metallic or semiconducting electrode to which a suitable potential may be impressed relative to a reference electrode and the ensuing current supported by a counter electrode. In both cases the bathing solution may also contain additional free monomer that may or may not be equilibrated with the electroactive monomer trapped within the hydrogel. Listed in Table 2 are the monomer components used in a typical synthetic scheme for an electroconductive hydrogel based on polypyrrole. The table gives the composition for three formulations that yield a prototypical tetraethyleneglycol diacrylate cross-linked poly(HEMA) hydrogel, a poly(HEMA)-based hydrogel of poly(HEMAco-PEGMA-co-HMMA-co-SPMA)-PPy and a poly(HEMA)-based hydrogel of poly(HEMA-co-PEGMA-co-HMMA-co-SPMA)-P(Py-coPyBA) where PyBA is 4-(30 -pyrrolyl)butyric acid. In this case the photoinitiator is 2,2-dimethoxy-2-phenylacetophenone (DMPA). These electroconductive hydrogels are based on 2-hydroxyethylmethacryate (HEMA), a water soluble monomer that may be UV polymerized at low temperature (20 to þ10 C) and may be readily copolymerized with other acrylate, methacrylate and acrylamide monomers e.g. poly(ethylene glycol) methacrylate
Hydrogel -PPy co-network
Electro polymerization of Py 0.7 V vs Ag/AgCl UV-polymerization of acrylates Chemical Oxidative polymerization of Py Monomer Solution
Hydrogel network
Electro polymerization of Py 0.7 V vs Ag/AgCl
FeCl3
Fig. 2. Schematic illustration of the generalized synthetic routes to Electroconductive hydrogels.
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Table 2 Components, their role and composition of the hydrogel cocktail for synthesis of an electroconductive hydrogel. Role
Compound in formulae
Mole % Poly(HEMA)
Poly(HEMA)-PPy
Poly(HEMA)-P(Py-co-PyBA)
Monomer 1 Monomer 2 Pendant PEG Cross-linker Dopant Counter Anion Prepolymer Photoinitiator Electroactive Monomer Electroactive Monomer Solvent Solvent
2-Hydroxyethyl methacrylate (HEMA) N-[Tris(hydroxymethyl)methyl] acrylamide (HMMA) Poly(ethyleneglycol)(200)monomethacrylate (PEG200MMA) Tetraethylene glycol diacrylate (TEGDA) 3-Sulfopropyl methacrylate potassium salt (SPMA) Poly-(2-hydroxyethyl methacrylate) (pHEMA) 300,000 2,2-Dimethoxy-2-phenylacetophenone (DMPA) Pyrrole (Py) 4-(3-pyrrolyl)butyric acid Water Ethylene glycol
84.0 5.0 5.0* 3.0 0 2.0* 2.0 0 0 20.0 wt% 20.0 wt%
63.0 5.0 5.0* 3.0 5.0 2.0* 2.0 15.0 0 20.0 wt% 20.0 wt%
61.5 5.0 5.0* 3.00 5.00 2.0* 2.0 15.0 1.5 20.0 wt% 20.0 wt%
*
Calculated on the basis of the repeat unit mole percent.
(PEGMA) and N-[tris(hydroxymethyl)methyl]-acrylamide (HMMA) to yield hydrogels of varying physical and chemical properties [94]. The presence of PEGMA and HMMA confers the antiprotein fouling properties of the pendant polyethylene glycol (PEG) chains and the temperature responsive properties of the acrylamide to the characteristics of the [p(HEMA-co-PEGMA-co-HMMA) hydrogels [95,96]. The polypyrrole component is synthesized by electropolymerization to yield a homopolymer of pyrrole or a copolymer of pyrrole and 4-(30 -pyrrolyl) butyric acid monomer that were physically entrapped within the p(HEMA)-based hydrogel while supplemented by pyrrole monomer that was in the bathing solution [97]. In this system, the 3-sulfopropyl methacrylate potassium salt (SPMA) of the hydrogel component serves as the counter anion to the positively charged, oxidized form of the conductive electroactive polypyrrole. In Scheme 1, details for the synthesis of 4-(30 pyrrolyl)butyric acid are illustrated. The 4-(30 -pyrrolyl) butyric acid (PyBA) of the conductive, electroactive component serves as a hydrophilic monomer that can potentially establish electrostatic interactions with the HMMA monomer of the hydrogel component. The molecular intimacy is designed into the two polymer system this way so to favor chemical compatibility and blend formation. Similar hydrogels may be formed from cocktails containing aniline
(i) Succinic anhydride in AlCl3/EtCl2, rt, 20 min (ii)
O2S N
(ii) 2 in xM HCl/toluene (iii) xM HCl, reflux, 36h
O HO
O
O
O2S N
(i) Zn + Hg2Cl2/xM HCl rt, 10min, decant
O2S N
, rt, 4h
and N-phenyl-1,4-phenylenediamine (dianiline) to enhance electropolymerization kinetics [7,8,98]. Shown in Fig. 3 are the necessary substrate surface modification, derivatizations and monomer casting to yield a poly(HEMA)based – PPy electroconductive hydrogel membrane that is adhered to an electrode surface via covalent tethering and multiple hydrogen bonding interactions with the PEG moieties. Because of the interfacial share stress that accompanies repeated swelling and deswelling of the electroconductive hydrogel, it is necessary to pay particular attention to the immobilization of the hydrogel onto electrode surfaces. This is particularly true for electrode structures and electronic devices that are heterogeneous; possessing exposed metal, semiconductor and insulator surfaces. The results given in Fig. 3 is applicable to a microfabricated interdigitated microsensor electrode (IME) or microdisc electrode array (MDEA) which consists of patterns of gold with critical dimensions of 1–20 mm, revealed through photolithographically defined windows in an insulating Si3N4 layer [99–102]. Fig. 4 is a schematic illustration of these device structures, useful in the development of electroconductive polymer sensor technology [103]. In the referenced example, the IME and MDEA devices were cleaned by sequential ultrasonic washing in boiling
HO OH O
(i) xM NaOH/MeOH, rt, 24 h (ii) 75°C, 4h N H 4-(3-pyrrolyl)butyric acid
Scheme 1. Synthesis and structure of 4-(3-pyrrolyl)butyric acid.
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Fig. 3. Substrate surface modification, derivatization and monomer casting to yield a poly(HEMA)-based – PPy electroconductive hydrogel membrane that is adhered to an electrode surface via covalent tethering and multiple hydrogen bonding interactions with the PEG moieties.
trichloroethylene (3 min, 86.7 C), acetone (1 min; 56.2 C), 2propanol (1 min; 82.4 C) and then washed profusely in room temperature (RT) deionized water. To remove adventitious chemisorbed organic residues, the MDEA devices were treated for 10 min
in the UV-Ozone cleaner, washed by ultrasonication in 2-propanol and then washed profusely in room temperature deionized water. To remove residual organic/ionic contamination and to produce a uniform, reproducible layer of –OH groups on the surface of the
Fig. 4. Left: An interdigitated microsensor electrode (IME) showing 10 mm fingered electrodes. Right: A microdisc electrode array (MDEA) showing individual 50 mm diameter microdiscs. Both shown modified via coating with poly(HEMA)-based PPy electroconductive hydrogels.
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Fig. 5. Molecular constituents of a poly(HEMA-co-PEGMA-co-HMMA-co-SPMA)/PPy electroconductive hydrogel membrane.
Si3N4 (activation), the electrodes were immersed in a (5:1:1, v/v/v) 60 C solution of deionized H2O:NH4OH:H2O2 (RCA Clean), held for about 10 s, quenched in deionized water for 1 min and then washed profusely with running deionized water, followed by exposure to radiofrequency (RF) glow discharge plasma for 10 min within which was periodically allowed humidified air to bleed into the chamber. The gold surface of the cleaned and activated devices was chemically modified with 0.01 M 3-mercapto-1-propanol (in anhydrous ethanol) and stored at room temperature overnight to introduce pendant alkylhydroxy surface functionalities. The pendant alkylhydroxy and surface –OH groups of the silicon nitride surface were subsequently
functionalized by treatment with 3-(aminopropyl)trimethoxysilane (g-APS; 0.01 M, in anhydrous toluene, RT, 2 h) in order to introduce pendant alkylamino surface functionalities. The devices were cured at 40 C for 20 min, then 110 C for 20 min, and then 40 C for 20 min, for a total of 1 h. For the final functionalization and to establish a continuous path of covalent bonding between the device surfaces and the hydrogel layer the devices were incubated for 2 h in acryloyl(poly(ethylene glycol))-N-hydroxysuccinmide (Acrl-PEG-NHS, Mw ¼ 3500) solution (0.001M) made up in 0.1 M HEPES (pH ¼ 8.5) – prepared under UV filtered conditions. Under UV-free conditions, the final hydrogel cocktail was sonicated, purged with nitrogen and applied evenly to the surface of the Acryl-PEG functionalized devices using a spin coater. The mixture was immediately irradiated with UV light (366 nm, 2.3 W/cm2, 5 min) in a UV crosslinker under an inert nitrogen atmosphere to effect polymerization of the hydrogel component. Finally, the electrodes with the base hydrogel were conditioned and un-reacted monomer extracted by sequential immersion in ethanol:deionized water mixtures (100% ethanol, 75:25; 50:50; 25:75; 100% deionized water; v/v) for a minimum of 1 h each. The electrodes with the pyrrole monomer containing hydrogel were immersed in a saturated pyrrole solution for subsequent electropolymerization of PPy or P(Py-coPyBA). Illustrated in Fig. 5 is the hypothetical chemical structure of the UV cross-inked hydrogel and its association with polypyrrole. Shown in Fig. 6 are the substrate surface modifications, derivatizations and monomer casting necessary to yield a poly(HEMA)based – P(Py-co-PyBA) electroconductive hydrogel membrane that is adhered to an electrode surface via covalent tethering and multiple hydrogen bonding interactions with the PEG moieties. While the hypothetical chemical structure of the UV cross-inked hydrogel and its association with poly(Py-co-PyBA) are illustrated Fig. 7. The electropolymerization of Py is favored within the hydrogel milieu. As shown in Fig. 8, the chronopotentiograms are produced by
Fig. 6. Substrate surface modification, derivatization and monomer casting to yield a poly(HEMA)-based –P(Py-co-PyBA) electroconductive hydrogel membrane that is adhered to an electrode surface via covalent tethering and multiple hydrogen bonding interactions with the PEG moieties.
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Shown in Fig. 9 is the electrochemical impedance magnitude, jZj, as a function of frequency for a basic poly(HEMA) hydrogel membrane (column 3, Table 2), a pristine PPy membrane and an electroconductive hydrogel poly(HEMA)-based–P(Py-co-PyBA) (column 5, Table 2) that were prepared on an MDEA 050 Au* device. The figure confirm that poly(HEMA) hydrogel is an insulator with a purely capacitate response. However, the electroconductive hydrogel shows similar interfacial impedance properties as the electropolymerized PPy whether within the hydrogel membrane or directly onto the electrode device surface. 6. Opportunities for electroconductive hydrogels 6.1. Biomaterials biocompatibility
Fig. 7. Molecular constituents of a poly(HEMA-co-PEGMA-co-HMMA-co-SPMA)/P(Pyco-PyBA) electroconductive hydrogel membrane.
galvanostatic (1 mA, fixed current) electropolymerization of Py at chemically modified and derivatized IME chips (IME*) to form IME*jPPy and IME*jGel-P(Py-co-PyBA) chips. To ensure adequate hydration of the hydrogel layer and to ensure equilibration of the pyrrole monomer and electrolytes between the solution and hydrogel phases, the devices were equilibrated in the electropolymerization bath of saturated pyrrole monomer (w0.4 M Py in 0.1 M Tris/0.1 M KCl, pH ¼ 6.1) for approximately 1 h prior to electropolymerization. The electrode potentials needed to support the kinetics at the uncoated and hydrogel coated IMEs, separately, is shown in Fig. 8. For electropolymerizations occurring directly at the electrode there is a decrease in potential from 0.85 V to 0.75 V within the first 25 s. Within this same time period the potential at the IME*jGel electrode falls more sharply from 0.80 V to 0.68 V. The physical evidence also supports a more rapid and uniform electropolymerization of PPy within the hydrogel membrane. Several factors are work in promoting more favorable kinetics (i) The Py monomer may partition more favorably into the hydrogel resulting in a higher concentration; (ii) The presence of the SPMA dopant anion on the hydrogel network, and (iii) the presence of the PyBA within the hydrogel may both favor uniform and rapid electropolymerization.
In order for biosensors, electro-stimulated drug release devices and neuronal prostheses to be successfully used in vivo, it is necessary to first minimize undesirable interactions between the indwelling device and the host’s foreign body response. The body responds to any foreign object by launching a series of physicochemical reactions triggered by the denaturation of adsorbed proteins, the receptor-mediated recruitment and attachment of cells, the production of a range of cytokines and the eventual fibrous encapsulation of the device [104,105]. Electroconductive hydrogels can be molecularly engineered to possess characteristics favorable for implantation, without eliciting the host immune response that so often leads to rapid device dysfunction. The in vivo and in vitro biocompatibility of hydrogels such as poly(HEMA) have previously been demonstrated by numerous authors [106,62]. Poly(HEMA) was developed around 1960 by Professor Otto Wichterle and has since found widespread use in biomedical applications, including soft contact lenses [107]. PEG has been shown to be a non-cytotoxic material, likely due to its well-known ability to inhibit protein adsorption [95,96], a likely precursor to the immune response. The biocompatibility of a hydrogel based on poly(HEMA) that incorporated PEG and MPC was shown to be effective at resisting protein adsorption (based on three extracellular matrix proteins), as well as supporting good cell proliferation [62]. Hydrogels can also be designed to possess hydration characteristics and mechanical properties similar to that of tissue. By varying the percent ratio of a crosslinker, tetraethylene glycol diacrylate (TEGDA), within a poly(HEMA)-based hydrogel, the water carrying capacity of the gel can be varied. The mechanical properties of the gel may also be varied with the degree of crosslinking. Although poly(HEMA) hydrogels, as materials, have the potential to provide implantable devices with an interface that
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time (s) Fig. 8. Kinetics of galvanostatic electropolymerization of PPy onto: (A) Surface modified and derivatized Pt IMEs and (B) Surface modified and derivatized Pt IMEs that were coated with hydrogel. Electropolymerization conditions were: 0.4 M pyrrole, Tris buffer, pH 6, 1 mA.
Fig. 9. Electrochemical impedance magnitude, jZj, as a function of frequency for a hydrogel membrane, an electroconductive hydrogel and pristine PPy electropolymerized on an MDEA 050 Au* device showing the equivalent reduction in interfacial impedance following electropolymerization of PPy whether within the hydrogel membrane or directly onto the electrode device surface.
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confers some measure of tissue-to-device biocompatibility, the issue becomes somewhat complicated when the surface of the device to be implanted also has a specific function, which may be electrical or electrochemical in nature. Hydrogels are typically non-conducting, and as a consequence, the insertion of an implantable neural tissue electrode into the subthalamic nucleus of the brain, or the insertion of a glucose biosensor into muscle tissue, having been previously coated with a hydrogel membrane, may suffer from high impedances that reduce the signal-to-noise ratio (SNR) of the devices. For neural implantable electrodes, the non-conducting hydrogel can greatly diminish the strength of the detected or generated electric potential signal. For in vivo electrochemical biosensors, the hydrogel on the electrode surface contributes to a reduction in the amperometric or impedance signal because of the low diffusion coefficient of the hydrogel compared to a physiologically relevant electrolyte solution. To overcome these challenges, an electrically conducting polymer that can be incorporated into the hydrogel seems promising. In vitro and in vivo studies have concluded that PPy is non-cytotoxic to mammalian cells and tissues [24,108,109]. Wong [110] and Shastri [108] have shown that through manipulation of the dopant anion and polymer oxidation state, PPy can exert influence on protein adsorption and conformation and so influence progress in the cell cycle. Shastri and Langer have reported the synthesis of a conductive bioerodible PPy formed from pyrroles that were 3substituted with ionizable (butyric acid) or hydrolyzable (butyric ester) groups [111]. By controlling the ratio of these pendant groups, the authors could engineer the bioerosion time from a few weeks to a year. In addition, they were able to demonstrate that PPy formed from these monomers supported the growth and differentiation of human-derived mesenchymal progenitor cells into osteoblasts. Schmidt et al. [112] have shown that PC-12 cells and primary chicken sciatic nerve explants were equally attached and extended equal neurites on PPy films in the absence of an electrical stimulation. However, in the presence of an electrical stimulus PC 12 cells showed a ca. two-fold increase (from 9.5 mm to 18.14 mm) in the lengths of neurite outgrowths. Upon animal implantation, PPy invoked little adverse tissue response when compared with poly(lactic acid-co-glycolic acid). PPy has been shown to improve the regeneration of damaged peripheral nerves in rats and it is a suitable matrix for the growth and proliferation of cell cultures. PAn films, subcutaneously implanted into laboratory rats did not cause inflammation or infection in the surrounding tissue, suggesting that PAn is not likely to produce adverse effects when implanted in the dorsal region of the skin. However, because these polymers are nondegradable, there is still the concern of their longterm impact in the in vivo environment such as inducing chronic inflammation. This problem is overcome by the use of biodegradable electrically conducting polymers (BECPs). The first example of such materials is a polymer composed of monomeric sequences of alternating units of pyrrole and thiophene capped on either side by degradable ester linkages and aliphatic linkages. Human neuroblastoma cells cultured in vitro on BECP films demonstrated significant cell proliferation after a few days, indicating good cell compatibility. Subcutaneous implantation into rats for several weeks resulted in only mild inflammation and tissue infiltration beyond the tissue–polymer interface. Recently, Schmidt has demonstrated biodegradable adipic acid polyesters possessing electroactive quaterthiophenes that confer biocompatibility [113]. These polymers were shown to biodegrade to produce the intact quaterthiophene subunit and were shown to be non-cytotoxic to Schwann cells. Fig. 10 illustrates the in-vitro biocompatibility through growth and proliferation of PC-12 cells on electroconductive p(HEMA)-PPy hydrogels.
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Fig. 10. Comparison of PC12 cell densities post-incubation (for 4 days) on Au*, Au*jhydrogel, Au*jPPy, Au*jhydrogel-PPy (5, 25 and 50 s electropolymerization times). Initial seeding cell density was 4.9 0.7 105 cells/ml (broken line). * Indicates a pvalue greater than 0.05.
While various Bio-microelectromechanical systems (BioMEMs) applications using PPy have been demonstrated, there has yet been no reported use of conducting polymers solely for implantable biosensor devices. However, recent material advances such as those described above suggest that the time is right for the use of electroconductive hydrogels for in vivo biosensors. 6.2. Implantable biosensors The utility of short and long term implantable biosensors for a plethora of clinical applications, particularly with respect to point of care medicine, surgery and intensive care is evident. There is tremendous value in obtaining real time data concerning the metabolic and physiological status of a patient in order for clinicians to make medically important decisions that could impact the immediate as well as the long-term outcome of patients. For example, an electrochemical biosensor that can interrogate lactate concentration in muscle tissue would be useful as an indirect measure of oxygen debt associated with hemorrhaging. Through the continuously updated analysis of increasing or decreasing trends in lactate concentration (increasing lactate being indicative of blood loss and increasing oxygen debt), it may be possible to predict patient outcome and/or inform the resuscitation strategy. If an unfavorable outcome is evident from such data, clinical intervention to avert that particular outcome may be possible or scarce resources may be alternatively deployed. Such a device could potentially save countless lives in the emergency room or in triage on the battle field, where trauma induced hemorrhaging is associated with high incidences of patient mortality. Temporary implantable biosensors for the detection and quantification of a diverse range of metabolites and biomarkers have been widely used in research and have achieved some measure of commercial success [114]. Many biosensors rely on immobilized enzymes that catalyze the formation of an electrochemically active byproduct from the metabolite or biomarker of interest. The presence of immobilized enzymes contributes to device specificity. The ability to achieve long-term implantation of an artificial device in the body comes with numerous challenges that greatly restrict the effectiveness of implantable medical devices for clinical applications. The primary challenge has been identified in the previous section –
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biocompatibility. Biofouling has been determined to be the most important factor contributing to the decreased sensitivity of implantable glucose biosensors within the first 24 h of implantation [115]. Any material that finds itself in a physiological environment will be identified by the organism as ‘‘non-self’’ subsequent to which the organism would immediately seek to expel or contain the implanted material. For many devices, such as the cardiac pacemaker, isolation via fibrous encapsulation is not entirely detrimental. This is because the metal surface making up the outer container of the pacemaker itself does not provide any other function than to provide a hermetically sealed environment to protect its internal components, the battery and associated microelectronics. In the case of an implantable electrode; however, the situation is entirely different. Active electrodes for electrical stimulation of tissue, or passive electrodes used for detection of glucose or lactate, require a non-fouling surface, that is, a surface that is resistant to protein adsorption in the presence of complex media. Once implanted, a subcutaneous or intramuscular electrochemical glucose or lactate biosensor, for example, such as described by Guiseppi et al. [62], Schuvailo et al. [116], Petrou et al. [117], and Moussy et al. [118–121], would inevitably come into contact at some point with bodily fluids, including blood. Tissue damage arising either out of surgical implantation, or micromovements once implanted can lead to the inflammatory cascade described previously. If it is understood that adsorptive denaturation of ECM proteins is a precursor to the inflammatory cascade [122,123], it therefore becomes imperative to develop materials that can effectively inhibit this crucial step. Increasingly, however, the evolving view is that it is protein denaturation pursuant to adsorption that may be the carrier of molecular cues (appropriate amino acid sequences) that provoke the inflammatory cascade [124,125]. Silencing the local adverse effects of the inflammatory response may be achieved through material–tissue interactions that include the electro-stimulated release of bioactive agents at the site of implantation, through interfacial properties that confer resistance to adsorptive protein denaturation, and through topographies and mechanical properties that match the local extracellular matrix. For implantable electrochemical biosensors, it is important that the electrode surface resists protein adsorption and denaturation (fouling) and eventual fibrous capsule formation, as this can lead to an appreciable reduction in the current response and consequently a low signal-to-noise ratio [126]. A hydrogel co-network comprising highly hydrophilic and biomimetic monomer components, poly(HEMA-co-PEGMA-MPC), was shown to confer significant resistance to protein adsorption but to evolve improvements in these properties over a five-day period following immersive conditioning in water. The temporal emergence of the PEG and PC moieties at the surface was shown to be responsible for the time dependent properties. Such hydrogels are useful coatings for electrochemical biosensors. However, such hydrogels are ionically but not electronically conductive, and as a consequence demonstrate high interfacial impedances. Therefore, blends and co-networks of hydrogels and conductive electroactive polymers (CEP) are promising, stimuli responsive, multi-functional materials with its CEP component molecularly engineered to reduce interfacial electrical impedances. Blends and co-networks of poly(HEMA-co-PEGMA) and a crosslinker, tetraethylene glycol diacrylate (TEGDA), have previously been formulated for electrochemical biosensor applications [62,94]. By varying the ratios of the individual monomers within the hydrogel composition, one can influence the hydration characteristics, as well as the mechanical strength or rigidity of the final polymerized material. It is expected that appropriate ECH hydration and modulus that simulate or mimic the characteristics of living tissue are both amenable to reduced fibrous encapsulation
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once implanted. Polypyrrole was selected as the CEP of interest to form the co-network with the acrylate-based hydrogel. Polypyrrole has been previously studied and was selected on the basis of its known biocompatibility [109,127,128] and its relative ease of polymerization (chemical oxidation or electropolymerization). In addition to conferring biocompatibililty, hydrogels also offer a receptive mileu for the physical or covalent immobilization and stabilization of biomolecules such as peptides, enzymes, glycosaccharides and whole cells. One key feature regarding the immobilization of enzymes within electroconductive hydrogels is the potential for mediated electron transfer between the redox active sites of oxidoreductase enzymes and metalloproteins with the conductive polymer component. Many biosensors have relied on an amperometric response produced due to the formation of hydrogen peroxide that results from the enzymatic conversion of substrate in the presence of molecular oxygen. Peroxide can be damaging to biomolecules, cells, and the hydrogel itself, therefore, the formation of such a byproduct within living tissue is not desirable. In addition, peroxide can be detrimental to the extended pi-conjugation and thus to the conductivity of polypyrrole. Belanger [129] proposed that the reaction of hydrogen peroxide and polypyrrole decreases the electrical conductivity. Electron mediators, such as ferrocene monocarboxylic acid [130,131] may be used; however, an important problem is the fact that such mobile electron mediators (or electronophores) can eventually diffuse out of the hydrogel if not covalently immobilized [132]. Some researchers have suggested that direct electron transfer can occur between certain CEPs and enzymes. There are some dehydrogenases that may be capable of directly transferring electrons to conductive polymers without producing hydrogen peroxide resulting in direct biosensing [133]. 6.3. Electro-stimulated drug release devices Electro-stimulated drug release devices (EDRDs) are engineered devices that produce a programmed drug release profile influenced by the application of an impressed voltage or current [134,135]. Controlled release of a bioactive agent occurs when the bioactive agent is astutely combined with a polymeric material such that the active agent is released in a pre-designed way [136]. The design and use of polymeric materials for controlled delivery systems has become one of the most promising advanced technology areas of polymers as well as in modern-day medical, agricultural, and pharmaceutical sciences [137]. Conductive electroactive polymers possess substantial porosity and delocalized charge centers to allow counter-ion diffusion and electromigration inside the polymeric electrode body in response to oxidation or reduction [138–141]. In electro-stimulated molecular release mechanisms, electroactive polymeric films or supported membranes are used to deliver, exchange, or gate the movement of specific molecules, usually ions, on demand and in a programmed manner. A form of electropumping is possible when a concentration gradient is maintained on either side for the supported membrane and the concentration is allowed to cyclically accumulate within the membrane [142]. A form of electroactive gating is possible when an electric field is maintained on opposite sides of the membrane and its oxidation state is altered. Burgmayer and Murray [143,144] were the first to demonstrate this principle in the gated anion transport through PPy films linked to its oxidation state. Reynolds et al. [145] subsequently confirmed that anion transport and anion exchange were largely charge compensating, was linked to the density of polarons on the onedimensional conductor, and had diffusivities that ranged from 1010 to 1012 cm2/s. The work of Pickup et al. [146] showed that repeated, non-equilibrium cycling caused morphological changes in the PPy films that affected anion diffusivity, with diffusivities
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that were low in the anodically as-electropolymerized films but subsequently made high following reduction. Krishna et al. used the luminescence of pyrene and naphthalene sulfonates to in situ probe the anion transport characteristics of these dopant anions into and out of the PPy films and used the luminescence of cationic Ru(bpy)2þ 3 and acridine orange to probe cation exchange within the PPy–Sulfonate films [147]. Wallace et al. [148–150] established factors that influence the release of large anions such as anthraquinone disulphonic acid and later established the relative selectivity based on charge density and size dependence using the mono-valent cations Liþ, Naþ, Kþ, Rbþ and Csþ and four membrane types prepared using different dopant anions. Miller et al. [151–153] were one of the first to show that PPy–PSS [poly(styrene sulfonate sodium salt)] could be an effective ion exchange material for the redox switching release of biologically relevant cations. In the oxidized state, the PPy–PPS possesses a net negative charge that is balanced by Naþ drawn from the bathing solution. Reduction of the PPy–PPS results in ingress and further binding of Naþ ions, which upon reoxidation results in the release of those cations. This principle was used to investigate the cathodic binding and anodic release of dopamine, a cationic neurotransmitter, important in diseases such as Parkinsons. Later, Prezyna and coworkers [154] used PPy–PSS films for the irreversible binding and immobilization of cationic polypeptides such as polylysine and histones. This possibility that polypyrrole (PPy) may serve as a gating membrane for the controlled release of anionic drugs was further developed and studied by Kontturi et al. [155] using several model substances that possess therapeutic value. Recently, Wadhwa et al. [156] reported controlled local drug release of the anti-inflammatory drug dexamethasome from electroactive PPy coated electrodes using cyclic voltammetry (CV). In a twist to the use of sacrificial anodes familiar in corrosion protection of less noble metals, Moulton et al. demonstrated the autonomous controlled release of the antiinflammatory drug dexamethasome from a PPy–Dex composite through the galvanic coupling of a sacrificial magnesium alloy and the PPy$Dex polymer [157]. However, the continued development and exploitation of PPy, PAn, PTh and other CEPs as drug release platforms suffer from the following major limitations: (i) Passive loss of loaded drug molecules by diffusion, (ii) loss of loaded drug molecules as a result of anion exchange with the bathing environment, (iii) poor oxidative stability of the pristine CEP, particularly under repeated cycling and in the presence of weakly oxidizing agents, (iv) a type of ‘‘membrane fatigue’’ that is the result of the fast switching of the CEP and the generally slower diffusion of the drug molecule that gives rise to an accumulation of salts within the polymer and hence a reduction in its drug eluting capability, (v) inherently low drug loading levels wherein the equilibrium payload is governed by the density of redox active sites on the CEP component, (vi) poor control and unpredictability of the binding and release kinetics, and finally, and finally, (vii) low diffusion coefficients of the released drug through dense semi-crystalline polymer fibrils resulting in poor release kinetics and necessitating the application of slow cycling conditions that seek to match ON–OFF rates with attendant transport properties. Clearly, CEPs with increased molecular porosity (void volume) and increased diffusion coefficients for counter anions are required if electro-stimulated release is to be successful for anything other than very small molecules under low dosage conditions. Several of these limitations are addressed by the development of electroresponsive hydrogels [17,158,159]. These limitations are also addressed by the development of electroconductive hydrogels that offer the use of lower electro potentials and faster switching [16,98]. The electro-stimulated retention of Ca2þ ions from an electroconductive hydrogel formulated from a co-network of a poly (HEMA)-based hydrogel and polyaniline has been demonstrated by
Wilson et al. [98]. The release profile of passively loaded Ca2þ ions as measured using an ion-specific Ca2þ electrode from the electroconductive hydrogel is shown in Fig. 11. The electroconductive hydrogel is seen to release up to 65% of its loaded Ca2þ ions under open circuit (OC) conditions (or in the absence of an applied potential) and to do so with considerable time lag. However, the electroconductive hydrogel instantaneously releases the loaded Ca2þ ions upon application of an oxidizing (þ200 mV vs. Ag/AgCl) or reducing (200 mV vs. Ag/AgCl) potential. The amount released under electrostimulated conditions is now ca. 25% reflecting the preferential ingress of Cl ions from the bathing solution. 6.4. Neural prostheses Neural prostheses (NP) are engineered assistive devices that present an engineering solution to restoring function lost to neural damage. Examples of such devices include optoelectronic prostheses [160] and artificial cochlear [161] for restoration of lost visual and auditory sensory functions, respectively, as well as prostheses for the motor control of artificial limbs [162,163], among other brain computer interfaces (BCI). Neural prostheses also include neurostimulation devices for pain relief [164], bladder control [165] and for the reduction of spontaneous tremors associated with Parkinson’s disease [166]. Each of these devices relies on effective electrical communication between the implanted device and the central nervous system. This is crucial in order to ensure that both the strength and the fidelity of the incoming and/ or outgoing signals are preserved. There is a tremendous amount of recent research which has gone into the development of motor prostheses to compensate for lost motor ability. One of the major problems associated with practical applications of these devices is the lack of long-term stability and eventual loss of function with time [167]. The reliability of 100 microelectrodes to obtain neural recordings from the primary motor cortex (MI) of monkeys for at least three months and up to 1.5 years has been previously reported [168]. The mechanism behind the loss in function is not completely understood, but fibrous encapsulation of the implanted devices have been observed. Electroconductive hydrogels based on a poly(HEMA)-based hydrogel component and polypyrrole polymer co-network may be useful in providing an electrically conductive, low-impedance and
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Fig. 12. Left – Picture and schematic of the Medtronic model 3389 DBS electrode showing the four 1.50 mm long platinum–iridium contacts with three 0.5 mm long separations on a 1.2 mm diameter lead and coated uniformly by a PES (polyethersulphone) of 0.25 mm thickness. Right – Electron micrograph of the freeze- fractured surface of an electroconductive hydrogel formed via electropolymerization and showing the dense PPy rich layer beneath a gel rich outer surface.
non-cytotoxic interface for long term, implantable neural prosthetic devices. The feasibility of employing the electrically conductive polymer, polypyrrole, as a non-cytotoxic, low-impedance material for the surface modification of neural implantable electrodes has been demonstrated [14,24,109]. Polypyrrole is easily electropolymerized in the presence of an appropriate dopant anion onto metal surfaces. However, the final polymer material itself has rather poor mechanical properties and can become easily detached from the metal surface or simply break apart when exposed to various forces. Moreover, the switching speed of the pure polymer is limited by the diffusivity of dopant counter anions and its intercalation into the semi-crystalline polymer. Such damaging forces may be expected during the insertion of electrodes into nervous tissue or over time once implanted due to micromotions and device migration. Detachment from the metal is understandable as the formation of the polymer film on the surface is not associated with any formidable covalent interactions. Incorporation of polypyrrole into a hydrogel network offers major advantages over the use of polypyrrole on its own. The mechanical strength of hydrogels can be influenced by the degree of crosslinking (based on the tetraethylene glycol diacrylate molar ratio). By incorporating the conductive electroactive polymer into a hydrogel matrix, the strength of the final material – the electronconductive hydrogel – would be that as designed for the native hyrogel itself. However, the new hybrid material would demonstrate virtually all the electrical characteristics – electrical impedance and electrochemistry – of the original polypyrrole conductive electroactive polymer. In addition, surface chemistries can be implemented to provide a strong and enduring covalent link as well as concerted hydrogen bonding interactions between the hydrogel and the electrode surface through a series of modification steps. Electrochemical/electrical studies conducted of polypyrrole on gold and of the ECH on gold reveal that little differences exist in the impedance plots for the two materials across a frequency range of 0.1 Hz–1 MHz [169]. Fig. 12 shows a graded electroconductive PPyhydrogel layer grown on the metallic conductor of a neural implant device. For many implanted devices, such as the artificial hip and cardiac pacemaker, scar tissue formation at the site of implantation is usually not a major hindrance to the biomedical device function. The formation of the tough fibrous tissue can actually be an advantage, as it would serve to anchor the implanted device in place; reducing motion and migration that could damage surrounding non-fibrous tissue. In the case of neural probes; however, fibrous encapsulation would eventually lead to device dysfunction. The high impedance associated with the surrounding
scar tissue would render the device inefficient and ineffective for electrical stimulation or for neural recording. Inclusion of components into the hydrogel, such as poly(ethylene glycol) monomethacrylate (PEGMA), phosphorylcholine methacrylate (MPC), can be used to reduce adsorptive protein denaturation and cell adhesion that can initiate the inflammatory response that eventually results in device encapsulation. The key to enhancing implant device function in vivo may lie in the identification and specific deactivation of important initiation factors that contribute to the inflammatory host response while activating and promoting those factors that contribute to the wound healing response. 7. Conclusions Electroconductive hydrogels represent a unique vista for combining the responsive properties of both materials within an aqueous milieu that is hospitable to biological molecules such as peptide sequences, enzymes antibodies and DNA. The combination of hydrogels and inherently conductive electroactive polymers allows both materials to retain their unique responsive properties. In addition, the ECH engenders a new class of devices with low interfacial impedances suitable for neural prosthetic devices such as deep brain stimulation electrodes, low voltage actuation for electrically stimulated drug release devices and potential for in vivo biocompatibility in implantable biosensors. Acknowledgments This work was supported by the US Department of Defense (DoDPRMRP) grant PR023081/DAMD17-03-1-0172, by the Consortium of the Clemson University Center for Bioelectronics, Biosensors and Biochips (C3B) and by ABTECH Scientific, Inc. References [1] Guiseppi-Elie A, Sheppard Jr. NF, editors. Conferring biospecificity to electroconductive polymer-based biosensor devices. ACS Northeast Regional Meeting (NERM): Rochester, NY; 1995 October 22–25. [2] Guiseppi-Elie A, Wilson AM, Sujdak AR, Brown KE. Electroconductive hydrogels: novel materials for the controlled electrorelease of bioactive peptides. Polym Prepr 1997;38(2):608. [3] Guiseppi-Elie A, Sujdak A, Wilson AM, editors. Electroconductive hydrogels: Electrical, electrochemical and impedance properties. Fall MRS Meeting Symposium J. Boston: Materials Research Society; 1997 December 1–5. [4] Small CJ, Too CO, Wallace GG. Responsive conducting polymer–hydrogel composites. Polym Gels Networks 1997;5(3):251–65. [5] Guiseppi-Elie A, Wilson AM, Sudjak AS. Electroconductive gels for controlled electrorelease of bioactive peptides. Washington, DC: ACS Symposium Series; 1998. pp. 185–202.
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