Experimental and theoretical characterization of implantable neural microelectrodes modified with conducting polymer nanotubes
Introduction
Impedance characterization of microelectrode–electrolyte interface is important for the application of biosensors and bioelectronics such as neural prostheses, where low impedance small microelectrodes are required for high-resolution stimulation and recording [1], [2]. Microfabricated neural prosthetic devices facilitate the functional stimulation and recording from neurons of the peripheral and central nervous system. In the physiologic environment, bioelectric potentials are carried in electrolyte media in the form of ionic current and the purpose of the neural electrode is to transduce these bioelectric signals to and from electronic signals [3]. The interface between microfabricated neural microelectrodes and neural tissue plays a significant role in the long-term performance of these devices.
A certain charge density is required to generate neural activity during stimulation using microelectrode arrays (i.e. 0.08–1.91 mC/cm2 for human retina) [4], [5]. High impedance electrodes would result in a large applied electrode potential leading to undesirable electrochemical reactions that may be harmful to the tissue. During recording, the extracellular signals are low, on the order of microvolts for neurons [1], [3]; so the neural signals will be lost in the noisy, ion-based electric fluctuations of the surrounding electrolyte media if the electrode impedance is not low enough. Therefore, a low impedance electrode–electrolyte interface is critical in the design of bioelectrodes. In order to design an optimized low impedance interface a detailed understanding of the physical processes contributing to the impedance is required.
Conducting polymers (CPs) have been widely used for biosensors and biomedical applications [6], [7], [8], [9], [10], [11], [12], [13]. The main characteristic of a conducting polymer (CP) is a conjugated backbone with a high degree of π-orbital overlap that can be subjected to oxidation or reduction by electron acceptors or donors, resulting in p-type or n-type doped materials (mostly p-type), respectively. Electrical conductivities can be varied by as much as 15 orders of magnitude by changing dopant concentrations so that control is feasible over the entire range from insulator to semiconductor and then to metal [14]. This makes CPs good candidates for coating the electrodes in order to minimize the impedance of electrode–electrolyte interface.
Among the known CPs, we have been interested in the electrochemical polymerization of PPy, PEDOT, and PEDOT derivatives because of their promising electrical properties and biocompatibility [12], [15]. PEDOT has exhibited some very interesting properties. In addition to high conductivity (ca. 300 S/cm), PEDOT was found to be almost transparent in thin film and showed high chemical stability in the oxidized state [16], [17], [18]. We have found that soft, low impedance, and biologically active coatings can be prepared by the electrochemical deposition of these CPs on neural microelectrode arrays [10], [19], [20].
Equivalent circuit models have long been used to model the electrode–electrolyte interface impedance. In 1899 Warburg first proposed that a polarization resistance in series with a polarization capacitor could represent that interface [21]. Randle's model consisted of an interface capacitance shunted by a reaction impedance, in series with a solution resistance [22]. As the use of electrodes in medical applications became more extensive, research was dedicated to the understanding of the electrode–physiological solution interface [23], [24]. Kovacs presented an equivalent circuit model based on Randle's model, with an additional Warburg impedance due to the diffusion of faradaic current [3].
Several equivalent circuits have been evaluated in order to model electrode–electrolyte interfaces coated with organic films [25], [26], [27] and CPs [28], [29], [30]. Cui and co-workers used a simple model of interface that was proposed by Bobacka et al. [28] for a neural electrode coated with PEDOT. This model consisted of a double layer capacitance (interface capacitance) in series with a Warburg impedance and the solution resistance [29]. Yang and Martin used Kovacs' model for evaluating the microporous PEDOT coatings interface on the neural electrode [30]. One of the most important challenges in the analysis of these models is to relate the parameters to chemical and physical characteristics of the films themselves.
In the work presented here, PPy and PEDOT were electrochemically polymerized on the neural electrode sites in the form of film and nanotubes. Neural electrodes were assembled at the Center for Neural Communication Technology (CNCT) at the University of Michigan (Fig. 1). The fabrication process of CP nanotubes is illustrated in Fig. 2 and has been described previously [10]. Electrochemical impedance spectroscopy (EIS) has been used to characterize the electrode–electrolyte interface of the modified neural microelectrodes. A new equivalent circuit model has been developed and used where each parameter represents a macroscopic physical quantity contributing to CP-modified electrode interface. The model consists of a coating capacitance in parallel with a pore resistance and interface impedance in series (Fig. 3). The model parameters have been fitted to the experimental results by using a nonlinear least-squares method. To confirm that the model parameters represent reasonable physical quantities, theoretical equations have been used to calculate the parameter values thereby validating the model. The effect of the initial interface conditions on the charge transfer resistance has also been determined. We have already shown that PEDOT nanotubes can improve the signal quality of recording sites and improves the long-term performance of chronically implanted neural microelectrodes in rats up to seven weeks [31].
Section snippets
Materials
Poly(l-lactide) (PLLA, RESOMER® L 210) with inherent viscosity 3.3–4.3 dl/g was purchased from Boehringer Ingelheim Pharma GmbH & Co. (KG, Germany). 3,4-Ethylenedioxythiophene (EDOT, BAYTRON® M) with molecular weight 142.17 g/mol was received from H.C. Starck Inc. (Newton, MA). The pyrrole monomer (Py) and lithium perchlorate (LiClO4) were purchased and used as received from Sigma–Aldrich.
Neural microelectrode arrays
The microfabricated neural arrays were provided by the Center for Neural Communication Technology (CNCT) at
Surface characterization of conducting polymer nanostructures
CPs (PPy and PEDOT) doped with LiClO4 were electrochemically polymerized on the microelectrode arrays with and without nanofiber templates in galvanostatic (constant current) mode. After PPy and PEDOT deposition, the PLLA fibers were removed by soaking in dichloromethane for 10 m. The wall thickness of the PEDOT nanotubes varied from 50–100 nm, and the nanotube diameter ranged from 100–600 nm. By controlling the polymerization time, we could reproducibly prepare tubular structures with thin walls
Conclusions
We have developed a method for fabrication of extremely low impedance and high charge transfer capacity CP nanotubes on the surface of neural microelectrodes. We have quantified the electrode–CP–electrolyte interface on the neural microelectrodes using measurement techniques, equivalent circuit modeling and theoretical analysis. Equations describing the physical properties occurring at the interface were presented, and the calculated values from the equations were in good agreement with the
Acknowledgements
Helpful comments on the manuscript were provided by Eugene Dariush Daneshvar. The NIH-NINDS-NO1-NS-1-2338, a Rackham Pre-doctoral Fellowship, and an Army Research Office MURI on “Bio-Integrating Structural and Neural Prosthetic Materials” (proposal no. 50376-LS-MUR, grant no. W911NF-06-1-0218) all supported this work. Partial support for this work was also provided by the National Science Foundation. The authors acknowledge the University of Michigan Center for Neural Communication technology
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