A fully implantable stimulator for use in small laboratory animals

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Abstract

This paper describes a low cost, fully implantable, single channel stimulator that can be manufactured in a research laboratory. The stimulator generates charge-balanced biphasic current pulses which are delivered to a bipolar electrode array for chronic stimulation of neural tissue in free-running laboratory animals such as rats and mice. The system is magnetically coupled and contains no batteries or external leadwires. The subject is placed in a chamber surrounded by three orthogonal coils of wire which are driven to generate a magnetic field. Currents are induced in wire coils in the implanted stimulator then regulated to produce biphasic current pulses with fixed amplitude of up to 500 μA. Phase duration is adjustable from 25 to 250 μs per phase. Charge balance is maintained by capacitive coupling and shorting of the electrodes between pulses. Stimulus rate can be continuously varied, and the temporal precision of the stimulus means that the stimulator can be used in behavioural experiments or for generating electrically evoked potentials. We describe the application of this stimulator for chronic electrical stimulation of the auditory nerve (i.e. a cochlear implant); however it will have application in other areas of neuroscience requiring controlled safe electrical stimulation of neural tissue over extended periods. Circuit diagrams and manufacturing details are provided as supplementary data.

Introduction

Chronic electrical stimulation of excitable tissue in experimental animals has important application in evaluating the safety and efficacy of neural prostheses (e.g. motor neural prostheses, cochlear implants, deep brain stimulators, respiratory muscle stimulators and visual prostheses), as well as more broad application in neuroscience wherever controlled stimulation of neural tissue is required.

We describe a low cost, fully implantable stimulator suitable for chronic stimulation in small laboratory animals. We have successfully used this stimulator to chronically stimulate the auditory nerve in deaf rodents (Widijaja et al., 2006), however given appropriate stimulation electrodes, this device could be used to safely stimulate any target neural population in the peripheral or central nervous system.

Our design criteria (Table 1) called for the development of a fully implantable stimulator for use in unrestrained small laboratory rodents including rats and mice. The stimulator must be capable of delivering precisely timed biphasic current pulses at variable rates suitable for chronic electrical stimulation, behavioural experiments and to generate electrically evoked potentials. The output of the stimulator must deliver charge-balanced biphasic current pulses to a bipolar electrode array; stimulus intensity is varied externally by varying the phase duration (i.e. varying charge/phase as neurons are charge sensitive; Grill, 2004), a voltage compliance of at least 5 V, and a charge recovery technique similar to that used in clinical devices to minimize the generation of DC (i.e. electrode shorting between pulses and capacitive coupling; Huang et al., 1999). Finally, the implantable stimulator must be encapsulated to ensure it is protected from body fluids over its required implant life (at least 10 weeks) and be biocompatible. While there are systems described in the literature which address some of these issues (Cools et al., 1978, Winter et al., 1998) the size of these implants precludes their use in small rodents such as mice. Moreover, the present stimulator is designed to drive implanted electrodes in the same way as a commercial cochlear implant by generating charge-balanced biphasic current pulses and shorting the electrodes between pulses. Other experimental stimulators lack one or more of these features. A comparison of the output characteristics of the present stimulator with those of typical commercial devices is summarised in Table 2.

We have attempted to design an implant that is small, reliable and may be manufactured in a laboratory at modest cost. It is powered and driven by a pulsed magnetic field generated by coils of wire wound around the outside of the animal enclosure. The ability to manufacture the device in-house is important as it allows researchers to readily alter the device to suit different experimental requirements (e.g. the incorporation of a neurotrophin delivery system into the electrode array; Shepherd and Xu, 2002).

We have developed two implantable stimulators; one for implantation in the mouse and a larger stimulator with additional options suitable for implantation in larger rodents such as the rat. The larger stimulator differs from the one designed for mice by including light-emitting diodes (LEDs) which are visible through the skin of the animal in order to provide confirmation that the stimulator is functioning. It also uses coils that enable a longer pulse period than the mouse design, and has an active electrode shorting circuit which may enable higher rates of stimulation (Huang et al., 1999).

Section snippets

Overview

A pulsed magnetic field is generated by applying a voltage to an external coil of wire wound around a chamber containing the implant stimulator (the “excitation” coil; Fig. 1a). In our application a 230 V biphasic voltage pulse is applied to the excitation coil (Fig. 1b). The current through this coil increases with time at a rate proportional to the applied voltage (Fig. 1c). Because the voltage is constant during the phase period of the pulse, the rate of change of current is also constant,

Discussion

We describe a low cost, fully implantable single channel stimulator that can be used to chronically stimulate target neurons using charge-balanced biphasic current pulses—a stimulus widely considered safe for use in chronic applications and used extensively in neural prostheses (Seligman and Shepherd, 2004). The design also incorporates charge recovery techniques used by neural prostheses, including capacitive coupling (Huang et al., 1999), to ensure long-term safety of the stimulus.

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Acknowledgements

We thank our following colleagues for their contributions to this work; David Perry, Helen Feng, Dr. Jin Xu, Pavel Prado-Gutierrez, Stephanie Epp, Anne Coco, Lauren Donley, Dr. Sandra Widjaja, Dr. Justin Tan, Prof. Dexter Irvine, Dr. James Fallon, Prof. Peter Seligman, Dr. David Lawrence and Ursula Shepherd.

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This work was funded by the NIDCD (NO1-DC-0-2109 and NO1-DC-3-1005), The Department of Otolaryngology, University of Melbourne, The Bionic Ear Institute, a Wagstaff Fellowship from the Royal Victorian Eye & Ear Hospital, and the Victorian State Government.

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