A multielectrode array for intrafascicular recording and stimulation in sciatic nerve of cats

Brain Research Bulletin, Vol. 51, No. 4, pp. 293–306, 2000 Copyright © 2000 Elsevier Science Inc. Printed in the USA. All rights reserved 0361-9230/00/$–see front matter

PII S0361-9230(99)00231-2

A multielectrode array for intrafascicular recording and stimulation in sciatic nerve of cats Almut Branner and Richard Alan Normann* Center for Neural Interfaces, Department of Bioengineering, University of Utah, Salt Lake City, UT, USA [Received 27 April 1999; Accepted 22 September 1999] ABSTRACT: The feasibility of implanting an array of penetrating electrodes into peripheral nerves is studied in acute experiments in the cat sciatic nerve. A novel, silicon-based array of microelectrodes, the Utah Electrode Array, was used, which contains 25 or 100 1-mm long electrodes that project out from a silicon substrate. Electrode arrays of this complexity, when inserted in the peripheral nerve, could cause significant compression of the nerve and block the conduction of action potentials. Using a high velocity insertion technique, the electrode array was implanted into the sciatic nerve. Compound action potentials were evoked by and recorded with cuff electrodes. Compound action potentials recorded 1 h after insertion were only slightly altered from those recorded before insertion. Single units were readily extracted from evoked multiunit neural recordings in response to cutaneous stimulation and limb rotation around joints. Current injections into the nerve through the electrodes evoked muscle twitches with currents in the 10 ␮A range. Recording and stimulation stability were demonstrated for periods of up to 60 h. We have shown that high density arrays of electrodes can be inserted into the peripheral nerve and can provide a stable recording and stimulating interface to individual peripheral nerve axons. Such an array may be useful in future neuroscience research and potential neuroprosthetic applications. © 2000 Elsevier Science Inc.

Peripheral nerve electrodes fall into two basic designs: extraneural electrodes [10,17,19,23,25,37,41,42] that electrically stimulate or record activity from large segments of the nerve, or intraneural electrodes [2,3,13,20,27,46,49,53] that electrically stimulate and record activity from small groups of neurons. Each of these two basic designs has unique advantages and disadvantages [15,16]. Extraneural electrodes are generally wrapped around the circumference of the nerve [9,26,42]. As these “cuff electrodes” do not penetrate the epineurium, they are easy to implant and are generally less invasive to the nerve than are intraneural electrodes. Long-term basic studies of the peripheral nervous system and most neuroprosthetic applications have relied upon the use of cuff electrodes. Unfortunately, the extraneural placement of these electrodes tends to preclude their use in applications requiring selective stimulation or recording. Conventional cuff electrodes generally only record compound action potentials (CAPs); as CAPs represent the activity of very large groups of neurons, selective information about small groups of neurons cannot be accessed. In order to increase the selectivity of the cuff electrode for peripheral nerve stimulation, researchers have placed a number of active electrodes around the circumference of the nerve, and the current “steering” that can be achieved with these “carousel” electrodes has improved selectivity to a certain degree [19,42]. Unfortunately, extraneural electrodes can also produce undesired pathological responses. The higher stimulating currents can cause damage if applied over a long time [1]. Cuff electrodes have to be custom fit to the nerve diameter or they will damage the nerve. Even with a well-fit cuff, nerve damage may still result due to swelling of the nerve as a result of the surgery [1]. The use of silicone and helical designs [23,43,50] in cuff electrodes has improved surgical access and reduced damage. Intraneural electrodes often have a needle structure. The great selectivity of intraneural electrodes has allowed researchers to understand how individual sensory receptors represent cutaneous sensation, heat, pain, and proprioception in trains of action potentials [14,48]. This sensory information can be used as a control and feedback signal in functional neuromuscular stimulation [17,54]. CAPs such as those recorded with cuff electrodes are not able to provide this selective information. Intraneural electrodes are also used to effect local stimulation of a select population of neurons. The proximity of the electrodes to the axons reduces the stimulation thresholds by factors of 10 to 100 over extraneural electrodes

KEY WORDS: Neuroprosthetic, Peripheral nerve, Silicon microelectrode, Compound action potentials.

INTRODUCTION Much of our understanding of the organization of the motor and somatosensory systems that has evolved over the past 70 seventy years has been made possible, in large part, through continuously improving systems of electrodes. These systems have allowed researchers to both electrically stimulate peripheral nerves and muscles and record the electrical activity of the nerves, evoked by external sensory stimulation or volitional movements [7,11,13,46]. Interest in new types of multielectrode arrays has been motivated by two considerations: (1) the desire to study multiple representations of sensory and motor information in populations of peripheral nerves and (2) to apply this knowledge to the emerging field of neural prosthetics, the goal of which is to use human engineered devices to restore lost neural function. Application of a multielectrode array in a neuroprosthetic device could greatly enhance the quality of life for the many individuals with spinal cord injuries.

* Address for correspondence: Richard Alan Normann, Department of Bioengineering, University of Utah, 20 S. 2030 E., RM. 506, Salt Lake City, UT 84112-9458, USA. Fax: ⫹1-801-581-8966; E-mail: [email protected]

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FIG. 1. Electron microscopic picture of the Utah Electrode Array. The array consists either of 25 (shown) or 100 electrodes on a 2 ⫻ 2-mm or 4 ⫻ 4-mm (not shown) base. The spacing between electrodes is 400 ␮m and each electrode is 1 mm long.

[39,53] and very small groups of axons can be stimulated, which provides high topographical selectivity. However, because these intraneural electrodes must penetrate the relatively tough epineurium that surrounds the nerve as well as the perineurium that surrounds each nerve fascicle, intraneural electrodes can cause short- and long-term nerve damage during the electrode insertion procedure [8,31–33,47,51,52]. Nevertheless, histological results suggest that intraneural electrodes can be used chronically in peripheral nerve [2,12,21].

An ideal peripheral nerve electrode array would have the neural selectivity of the intraneural electrode and the ease of implantation, stability, and biocompatibility of the extraneural electrode. We have developed a novel, multielectrode array that may satisfy some of these conflicting criteria. The Utah Electrode Array (UEA) [5,18] was developed for use in the cerebral cortex. It consists of up to 100 1–1.5-mm long electrodes that project out from a 4 ⫻ 4 ⫻ 0.2-mm thick silicon substrate. To achieve implantation in cortical tissue, the array is rapidly inserted into cortex (in less than

FIG. 2. Comparison of CAP recordings from 30 min before to 60 min after implantation of the electrode array. Each waveform represents the waveform averaged over 5 min. The bold-faced line represents the last waveform before implantation.

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FIG. 3. (A) Changes in CAP amplitude at several times relative to the implantation of the electrode array (0, 5, 30, 60 min, and at the end of the experiment) observed in five different experiments. (B) Changes in the onset of the CAP (␮s). Each bar indicates the mean value and the whisker reflects the corresponding standard deviation.

200 ␮s) [35]. In a series of chronic experiments in feline auditory and visual cortexes, the structure of the electrode array and the implantation technique did not cause significant damage to the tissues [24,40]. Single-unit recording from visual and auditory cortexes of cat, motor cortex of monkey, and electrical stimulation of feline auditory cortex have been possible for over 2 years using these electrode arrays. This paper explores the use of this array as a multichannel peripheral nerve interface. We studied whether implantation of an array of penetrating electrodes through the tough epineurium of the nerve and the perineurium of the fascicles could be achieved without producing significant compression damage or other trauma. We also investigated the stimulation and record-

ing properties of this multichannel neural interface in peripheral nerve. In acute experiments, we implanted modified UEAs in cat sciatic nerve using a pneumatic insertion device as used in the central nervous system [35]. The impulse inserter consists of a compressor, electronics, and an insertion wand that rapidly accelerates the electrode array into the neural tissues (in under 200 ␮s). The wand is 20 cm long and 1 cm in diameter, and it is positioned stereotaxically over an electrode array, properly positioned over the neural tissues to be implanted. The inserter is lowered until it just touches the back surface of the UEA. During insertion there is a momentum transfer between the piston and the array, and the array is fully inserted. The travel of the tip of the inserter is

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FIG. 4. The waveforms obtained in one experiment (cat 3, leg 1) were averaged over 5 min and reflect the mean waveform before implantation, and 5 min and 1 h after implantation.

mechanically limited to the length of the electrodes, so the momentum of the insertion tip is never transferred to the neural tissue. The high velocity insertion is a new approach for electrode implantation in peripheral nerve that eliminates the need to cut the epineurium and/or perineurium of the nerve. These unique aspects of the insertion appear to reduce the nerve damage associated with the insertion. This notion is supported by the fact that recordings of CAPs before implantation were very similar to the ones 1 h after implantation. We show that the array is capable of recording single-unit responses from mechanoreceptors, and that digit twitches can be evoked with current injections in the 1–10 ␮A range over the 36-h long periods used in this study. We conclude that the UEA could provide a new tool that will enable new directions in studies of parallel processing by the peripheral nervous systems. The electrode array can also be a vehicle for neuroprosthetic therapies for disorders of the peripheral nervous system. MATERIALS AND METHODS Structure of the Electrode Array Electrode arrays containing various numbers of electrodes were used in this work but all were built with the same basic architecture. The manufacturing process of the electrode array has been described elsewhere [18] and the UEA is now commercially available (Bionic Technologies, Inc., Salt Lake City, UT, USA). Electrical insulation of the silicon microelectrodes is accomplished with polyimide or with silicon nitride. The electrodes are regularly spaced on 400 ␮m centers and project 1 mm out of the plane of a 0.2-mm thick silicon substrate (Fig. 1). Each electrode is electrically isolated from its neighboring electrodes with a “moat” of glass around its base. The tip of each electrode is coated with platinum to provide an electrical interface to the extracellular space. Each electrode is approximately 80 ␮m wide at its base and tapers to a sharpened tip. A Teflon insulated Pt/Ir wire (10IR1T, Medwire, Mt. Vernon, NY, USA) is soldered to the bond pad

deposited on the back of each electrode on the array. The lead wires are soldered to a connector that plugs into a data acquisition system and these soldered interfaces are encapsulated with silicone elastomer (Silastic MDX4-4210, Dow Corning Corp., Midland, MI, USA). The impedances of the electrodes are measured with a 100-nA, 1-kHz sinusoidal current. Impedance generally varies between 50 and 200 k⍀. The arrays used in this study had 10 rows of 10 electrodes, each on a 4.2 ⫻ 4.2-mm2 base or a 5 ⫻ 5 pattern of electrodes cut from the 10 ⫻ 10 electrode array. Animal Preparation and Electrode Implantation Experiments were done on cats. Anesthesia was induced with Ketamine (10 mg/kg, Sanofi Winthrop Pharmaceuticals, Morrisville, PA, USA) and maintained with halothane gas (0.9%–1.5%, Halocarbon Laboratories, River Edge, NJ, USA) during the experiment. ECG, expired CO2, and rectal temperature were continuously monitored. The right leg was shaved and an incision was made along the thigh from the back to the knee. The biceps femoris muscles were separated and retracted to expose the sciatic nerve over a length of 7–9 cm. The array was positioned on the sciatic nerve from the lateral side of the animal 1–3 cm proximal to the nerve’s branching into the tibial and fibular nerve. The impact inserter was positioned to apply a slight pressure to the array to facilitate insertion direction during the impact. After implantation, a Pt/Ir reference wire (20IR2T, Medwire) was placed in the fluids surrounding the nerve or in a neighboring muscle. This reference wire was used in both recording and stimulation experiments. A 2 ⫻ 2-cm piece of plastic film covered the array and adjacent tissue, the muscle was put back over the nerve, and the skin was closed with a few sutures. At no point was the nerve or the leg supported or restrained. In some experiments, two cuff electrodes were placed around the sciatic nerve, one approximately 3 cm proximal and another approximately 3 cm distal to the array. The cuff electrodes were

MULTIELECTRODE ARRAY FOR SCIATIC NERVE RECORDING/STIMULATION made of silicone rubber tubing (0.125⬙ ID ⫻ 0.187⬙ OD) (Bio-Sil, Sil-Med Corporation, Taunton, MA, USA). The cuff electrode construction followed the description by Davis et al. [6]. The stimulating cuff electrode had two Pt/Ir electrode wires (10IR3T, Medwire) sutured into the silicone rubber tubing with 1-cm spacing. For recording of CAPs, three electrode wires were sutured into the silicone rubber tubing with 1-cm spacing. A tripolar recording setup was used to reduce the signal to noise (S/N) ratio of the stimulus artifact and muscle activity. All experiments were conducted according to NIH guidelines for the use of animals. Experimental Setup for Stimulation and Recording Stimulation and recording of CAPs with cuff electrodes. Two cuff electrodes were used for simultaneous stimulation of the nerve and recording of compound action potentials. The stimulation cuff electrode was placed proximally and the recording cuff was placed distally to the implantation site. A custom-built, computer-controlled, constant current stimulator was used to provide the current for electrical stimulation using both the electrode array and the cuff electrodes. Using cuff electrodes, the stimuli were single biphasic pulses with a width of 200 ␮s per phase and a 100-␮s interphase interval. The stimulation current was chosen to cause maximal amplitude of CAPs as determined by CAP amplitude to stimulus current curves. The current was limited to 255 ␮A by the stimulator. Compound action potentials were recorded using the distal cuff electrode, which was connected to a differential amplifier (fL ⫽ 1 Hz; fH ⫽ 3 Hz) (DAM 60, World Precision Instruments, Sarasota, FL, USA). The gain of the amplifier was 10,000. Responses were digitized with an analog to digital (A/D) board (Win-30D, United Electronics Industries, Inc., Watertown, MA, USA) installed in a PC with a Pentium processor. Recording and stimulation were triggered simultaneously by the computer, and recording was done in 50-ms segments at a sampling rate of 50,000/s. The stimulus was recorded for offline determination of its amplitude. The stimulus signal (second channel) was used to align the CAP data on the first channel. One recording of the stimulus served as reference signal and all other stimuli were shifted in time until the correlation coefficient between the two signals was maximal. That way, the whole waveform was used for

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alignment rather than one or more data points. The amplitude of the CAP will be referred to as the difference between the first maximum and the minimum of the waveform. Stimulation with the electrode array. Using the electrode array, biphasic pulses with a width of 200 ␮s per phase and a 100-␮s interphase interval were applied, delivered at 1 and 2 Hz. During each measurement, stimulation threshold was determined for each electrode. Stimulation was started at 10 ␮A and increased in steps of 5 ␮A if no muscle twitch could be detected. As soon as a muscle twitch was detected using visual and tactile cues, the current was decreased in 1 ␮A steps until threshold was reached. As intrafascicular electrodes have much lower thresholds than extraneural electrodes, and because we wanted to avoid any damage caused by stimulating currents, generally, the current delivered via the array was limited to 100 ␮A. Recording. The electrode array was connected to a 16-channel amplifier (fL ⫽ 250 Hz; fH ⫽ 7.5 kHz) (Bionic Technologies, Inc.). The gain of the amplifier was 25,000. Noise of the amplifier (referred to as the input) was 2 ␮Vrms. Responses were digitized with an A/D board (WIN-30D, United Electronic Industries, Inc.). The PC simultaneously saved the responses from the 16 recording channels. Up to 10 s of data were recorded in each run at a sampling rate of 20,000/s. To extract single-unit activity from a multiunit recording, individual action potentials were first extracted from the records of neural activity using a thresholding procedure. These action potential waveforms were peak-aligned and then passed to a template generation program. Templates were formed by a user utilizing time-amplitude information. An action-potential template consisted of a waveform description and a measure of the noise on the electrode for which the template was made. After classifying as many action potentials as possible and creating templates for them, these templates were then used to automatically classify the remaining action potential records. An action potential classification was accepted if its Euclidean distance from the template that it best fit was less than the noise present on that electrode. Histology After an experiment, the cat was either perfused through the heart for histology with a buffered one-and-one-eighth percent

TABLE 1 MEAN FRACTION OF THE AMPLITUDE OF THE CAP AT THAT TIME POINT COMPARED TO ITS PREIMPLANTATION VALUE (AMPLAFTER/AMPLBEFORE) (1ST LINE IN CELL) AND MEAN TIME SHIFT OF THE ONSET OF THE CAP COMPARED TO ITS PREIMPLANTATION VALUES (⌴S) (2ND LINE IN CELL) WITH THEIR STANDARD DEVIATIONS Preimplant.

After 5 Min

After 30 Min

After 60 Min

After X Min

Cat 1

1.00 ⫾ 0.007 0.00 ⫾ 7.6

0.892 ⫾ 0.046 24.5 ⫾ 7.0

0.936 ⫾ 0.006 21.9 ⫾ 6.5

0.915 ⫾ 0.007 13.5 ⫾ 9.8

Cat 2, leg 1

1.00 ⫾ 0.044 0.00 ⫾ 7.4 1.00 ⫾ 0.013 0.00 ⫾ 4.7

0.727 ⫾ 0.042 24.5 ⫾ 9.00 0.856 ⫾ 0.023 10.2 ⫾ 5.0

0.007 ⫾ 0.001 — 0.836 ⫾ 0.012 ⫺2.3 ⫾ 5.3

— — 0.862 ⫾ 0.004 ⫺1.1 ⫾ 5.2

Cat 3, leg 1

1.00 ⫾ 0.003 0.00 ⫾ 5.1

0.932 ⫾ 0.005 74.4 ⫾ 10.6

1.004 ⫾ 0.002 46.9 ⫾ 8.9

1.028 ⫾ 0.003 21.9 ⫾ 7.1

Cat 3, leg 2

1.00 ⫾ 0.004 0.00 ⫾ 7.0

0.911 ⫾ 0.007 39.0 ⫾ 9.5

1.011 ⫾ 0.010 48.0 ⫾ 4.2

1.036 ⫾ 0.004 ⫺0.7 ⫾ 5.2

0.934 ⫾ 0.005 ⫺11.9 ⫾ 6.5 (X ⫽ 135 min) — — 0.899 ⫾ 0.010 ⫺7.3 ⫾ 0.0 (X ⫽ 120 min) 1.029 ⫾ 0.004 14.1 ⫾ 7.6 (X ⫽ 65 min) 1.052 ⫾ 0.004 ⫺15.7 ⫾ 9.8 (X ⫽ 90 min)

Cat 2, leg 2

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FIG. 5. Histogram of the stimulation thresholds. The data was obtained using a single, biphasic constant current pulse of 200 ␮s pulse width with a 100-␮s interphase interval in three feline experiments (data from 133 electrodes).

gluteraldehyde–1% formaldehyde solution or sacrificed with a 10 cc-IV injection of 3M KCl. The array was removed from the nerve after one day of soaking in fixative. The nerve was sectioned into pieces of 1 mm thickness at the implant site. The same was done with control tissue taken from the unimplanted leg. The nerve tissue was osmicated, dehydrated, and embedded in Epon. Semithin sections of 1 ␮m thickness were cut orthogonal to the fibers, stained with Richardson’s stain, and put on glass slides. Axon density was compared between the implantation site and control tissue, and the position of the electrodes was determined using light microscopy. RESULTS Acute recording and stimulation experiments were performed in the sciatic nerve of 12 cats. Two different classes of recording experiments were performed: experiments using the CAP as obtained with cuff electrodes and experiments using single units as accessed with the electrode array. Stimulation and Recording of CAPs with Cuff Electrodes One of the main goals of this study was to determine if a high-density array of penetrating electrodes could be inserted through the epineurium into the nerve without doing significant damage to the nerve. The CAP recorded from an undamaged nerve offers a global index of the nerve’s activity and can be used to monitor nerve damage. Any blockage in the conduction of action potentials along the nerve or changes in conduction velocity resulting from nerve injury caused by the insertion of the electrode array into the nerve would manifest as a decrease in the amplitudes of the CAP peaks and/or as a temporal dispersion of the CAP waveform [22,44]. Recording CAPs before and after array implantation should allow us to monitor any major consequences of the insertion but will be insensitive to very localized damage to axons. Five of these experiments were performed in 3 of the 12 cats. Two cuff electrodes were put around the nerve approximately 3 cm distal and proximal to the implantation site of the electrode array. The nerve was stimulated using a supramaximal current. The nerve was stimulated about twice a minute using the proximal cuff electrode starting 30 min preimplantation and until

60 –150 min postimplantation. Stimulation and recording were triggered simultaneously by the computer. Because a supramaximal stimulus was used, there was no need to average over many CAPs at each instant as demonstrated in Fig. 2. The CAPs recorded during 5-min intervals were averaged in all experiments to better illustrate the changes over time (Fig. 2). The changes in the amplitude of the CAPs (Fig. 3A) and the time they occurred are also shown (Fig. 3B). A sample of averaged waveforms before and 5 min and 1 h after implantation can be seen in Fig. 4. Table 1 summarizes the results obtained in all experiments. It can be seen that the postimplant waveform and amplitude of the CAP changed immediately after implantation of the electrode array. The amplitude decreased by up to 30% and it shifted in time by up to 75 ␮s. Usually, the CAP amplitude recovered to values that reflect changes of about 3–10% compared with preimplant values within 30 – 60 min after implantation. The electrode array broke during implantation in one experiment due to high insertion forces and defective array material. This caused heavy bleeding, which was followed by complete blockage of the action potential transmission. Such a severe reaction was not seen in any other experiment. No bleeding was apparent after successful implantation of the array. In some cases, there was bleeding after removal of the array at the end of the experiment, which suggests that the array sealed off impaled blood vessels. Stimulation and Recording of Single Units with the Electrode Array Stimulation experiments. The experiment with the CAPs provides insights pertaining to any major damage to the nerve produced by array implantation. For use on a chronic basis, it is not only important that the array does not cause damage but also that it does provide stability in terms of electrode position and, therefore, stimulation threshold and type of fibers excited by one electrode. Experiments were done using a stimulus train of 200-␮s wide biphasic pulses delivered either in a single pulse or at a rate of 1–2 Hz. Movements of the extremities were evoked by electrical stimulation of the sciatic nerve. Rotation was evoked around the ankle or the knee, but also more selective muscle activation could be

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FIG. 6. Changes in the stimulation threshold over time. (A) Stimulation threshold (␮A) over 30 h for 23 electrodes implanted into cat sciatic nerve. (B) Mean stimulation threshold and its standard deviation for each of the 23 electrodes over that same time course.

seen, such as movement of just one digit in a certain direction or a spreading of the digits. Since the leg was not restrained, motion caused by electrode stimulation ranged from twitches at threshold to full contraction of the stimulated muscle. Saturating stimuli were only used occasionally to investigate movement-induced displacement of the electrodes in the nerve. Fig. 5 is a histogram of the threshold currents delivered as a single biphasic pulse that evoked muscle contractions. Thresholds were mainly in the range of 1–20 ␮A, with a median value from 133 measurements of 9 ␮A. In terms of charge per phase, this translates to 1.8 nC/phase (approximately 650 nC/mm2/phase). The stimulation threshold did not depend on the impedance of the individual electrodes. In order to determine the stability of the implanted electrode array with respect to the individual fibers in the nerve, stimulation thresholds were determined for each electrode about every 3 h over the duration of each experiment (1–3 days). Fig. 6 has 2 different plots of the results from 1 cat experiment; the

stimulation thresholds of 23 different electrodes are plotted vs. time (array was implanted at time 0). In Fig. 6A, all 23 responses are plotted on a logarithmic ordinate. Fig. 6B is a plot of the means and standard deviations of thresholds on each electrode for the data of Fig. 6A. These results show that the UEA manifested exceptional stability in these acute experiments: for the 34 h over which this experiment was done, the mean coefficient of variation Cv was 0.09. To test the efficacy of the UEA in a stimulation application, we compared the number of potentially functional electrodes (electrodes that were electrically connected to our connector) with the number of electrodes that actually worked (twitches with less than 100 ␮A). The results varied between individual experiments from 44 –100% and generally reflected improvement in our experimental technique. The mean percent of functioning electrodes was 77%. As expected, this number is high compared to the number of functioning electrodes based on recording criteria (see Recording section).

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FIG. 7. Comparison of recordings acquired in a 4-min time interval. (A) The mechanical stimulation was evoked by brushing over parts of the paw. Channels 1, 5, and 10 seem to record relevant nerve activity. (B) Rotation of the leg around the knee joint was the stimulation paradigm (channels 2, 4, and 12).

Recording experiments. Responses were recorded from several electrodes simultaneously. Recording experiments were conducted over 2–3 days with recordings made with the electrode array about every 3 h. Multiunit activity was evoked and localized by brushing of the animal’s paws and digits or by limb rotation around a joint. Action potentials could be recorded from muscle spindle receptors that encode information about leg or toe rotation in a certain direction, or from hair follicles. Figs. 7A and 7B show multiunit signals typically recorded with the UEA. The two sets of simultaneous recordings are from the same set of 13 electrodes and were recorded sequentially over a few-minute interval. Periodic brushing of a few hairs on the top of one digit evoked the responses in Fig. 7A, whereas those in Fig. 7B

were evoked by leg rotation around the knee. It is clear from these results that the UEA recorded selectively from the fibers in the nerve; electrodes 1, 6, and 10 recorded units that were well-driven by brushing the digit, while electrodes 2, 4, and 12 recorded units that were well-driven by rotation around the joint. The firing rate of channel 12 was tonic and proportional to the angle of displacement. The continuous firing of channel 12 in Fig. 7A shows that the leg was slightly rotated at the time of recording. Channel 12 seemed to record from a spindle in a muscle that flexes the leg, whereas channels 2 and 4 recorded from a spindle in one of its antagonist muscles as it encoded rotation in the opposite direction. Leg rotation during these experiments was as large as 50°, which resulted in a relatively large displacement of the nerve and the array (1–2 mm). However, this seemed to neither influence the

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FIG. 8. Single-unit identification on an individual channel. Each waveform shows averaged unit responses on a single channel. All data was taken within a 36-h period. Five different single units can easily be identified.

quality nor the stability of the recordings, nor the stimulation thresholds. Not all implanted electrodes were able to record multiunit activity (either spontaneous or evoked). While 29% of all the electrodes we implanted over the course of these experiments had recording capabilities, values ranged from 10 to 83% in individual experiments and clearly depended upon the quality of the surgery and implantation. In the most recent experiments more than 50% of the electrodes had recording capabilities due to continued improvements of our surgical techniques. In some experiments, single units were extracted from the multiunit recordings using custom-designed signal processing software. The amplitudes of the evoked responses in different experiments ranged from 15–200 ␮V. An example of five single units recorded from a single microelectrode at various times over a 48-h period is shown in Fig. 8. The waveforms shown in this figure were obtained by averaging over many action potentials. All waveforms of a single unit on a given channel were extracted

and aligned in time. The clustered data is comprised of the averaged waveforms of up to 210 of these single units at any given time point. While there were small changes in the amplitudes of the action potentials over this period, we seldom saw monotonic changes in the data. As in this example, analysis of many recordings showed that our electrodes often recorded from three, but occasionally up to five, different single units. The quality of single- and multiunit recording is defined by their signal-to-noise (S/N) ratios. The calculation of the S/N ratio is described by Nordhausen et al. [29]. In most cases, only electrodes with S/N ratios larger than three were recorded so that single units could be extracted and, in some cases, only the best signals of a set were recorded. Fig. 9 shows a histogram of the number of electrodes that recorded responses in a given S/N ratio range. S/N ratios ranged from 2 to 36, with the mean S/N ratio of 10 (⫾7). Perhaps the best index of the chronic stability of an implanted electrode is its ability to record neural activity from the

FIG. 9. Histogram of S/N ratios of electrodes. The data from three experiments have been included. The S/N ratio was obtained by calculating the average amplitude of every spike over time and dividing it by the average (RMS-Noise ⴱ sqrt(2)) over time.

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FIG. 10. Histogram of duration versus percentage of single units active. The results of two experiments are shown. The first experiment lasted 36 h. The mean duration for which a unit was present was 15.5 ⫾ 8.5 h. The second experiment was terminated after 61 h. Here, on average, a unit was present for 34.1 ⫾ 24.9 h. The duration of a single unit was determined as the period between when it was seen first and last.

same identified single unit for prolonged periods of time. We used this index to validate the recording stability of the UEA as a peripheral nerve interface. Our stability experiments were conducted over a 2–3-day time course where we monitored the amplitudes and kinetics of single units. Sixteen channels were selected for recording by the S/N ratio of the recorded neural activity. In some cases, single units were lost or became less well-defined, and in other cases, new single units appeared after some time. Experiments were terminated after different time periods, depending on the number of single units left and the condition of the animal. Fig. 10 is a plot of the percentage of units vs. the duration they are active. Two different experiments were summarized here. The first experiment had a higher num-

ber of single units, but less than 50% of the single units lasted longer than 18 h. There were less single units recorded in the second experiment, but 50% of the units were present longer than 50 h. The changes over time of a single unit in one experiment can be seen in Fig. 11. The amplitude changed little over the 36-h period of this experiment. The level of anesthesia and displacements of the electrodes probably caused the minor changes seen in this figure. Histology Our recordings and stimulation thresholds suggested that our electrodes were located within individual fascicles. To verify if

FIG. 11. Single unit changes over time. An identified single unit was tracked over a 36-h time period. There were only slight changes in the kinetics of the waveform.

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FIG. 12. Light microscopic pictures of cross sections of a cat sciatic nerve. The tissue thickness in both cases was 1 ␮m. (A) Note the position of the electrode tip in the fascicle; bar 175 ␮m. (B) This picture represents a closeup of (A). A small compression wave can be seen in the tissue around the tip of the electrode; bar 100 ␮m.

this was the case, and to examine the nerves for evidence of trauma associated with the implantation procedures, we performed a histological analysis on some of the nerves from these acute experiments. Sections of 1 ␮m thickness were cut orthogonal to the nerve, stained with Richardson’s stain, and inspected using light microscopy. The figures shown here are from sciatic nerves in cats fixed 5 h after implantation of the electrode array. It can easily be seen that the electrode tips were indeed located within the fascicles (Fig. 12A). The electrode array was implanted directly into the nerve,

and fibers along the upper part of the electrode do not seem affected by the implantation. The electrodes were inserted approximately two-thirds into the nerve. The technique used mainly stained myelin sheaths as visible on the slide. A blowup of the nerve where the electrode tip was located showed evidence of a small region of compression of the fibers under the electrode tip (Fig. 12B). This compression zone extended below the electrode tip in a half circle with a radius of 150 ␮m. The density of the axons in the compression region was about 2.5 times higher than in unaffected areas. It cannot be

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BRANNER AND NORMANN TABLE 2 COMPARISON OF EXTRANEURAL AND INTRANEURAL PERIPHERAL NERVE ELECTRODES Investigator

Extraneural Electrodes Grill et al. [10]

Koller et al. [19]

Type of Electrode

Cuff electrode; 4 tripolar Pt electrode configuration Epineurial electrode; stainless steel

Site of Implantation

Type of Experiment

Sciatic nerve, cat

Acute stimulation

Sciatic nerve, rat

Chronic stimulation

Loeb et al. [23]

Cuff electrode; 3 Pt electrodes

Sciatic, sural and superficial peroneal nerve, cat

Chronic stimulation and recording

McNeal et al. [25]

Cuff electrode; 3 bipolar Pt electrodes

Tibial nerve, cat

Acute/chronic stimulation

Rozman et al. [36]

Cuff electrode; 7 bipolar Pt electrodes

Acute stimulation

Sweeney et al. [42]

Cuff electrode; 12 tripolar electrodes

Common peroneal nerve, dog Sciatic nerve, rabbit

Intraneural helix; Wire helix, stainless steel

Tibial nerve, rabbit and cat

Chronic stimulation

Branner et al. [3]

Intrafascicular electrodes; 25–100 silicon needles

Sciatic nerve, cat

Acute stimulation and recording

Liang et al. [22]

Intrafascicular electrodes; 16 Au electrode array

Sciatic nerve, frog

Acute stimulation

Nannini et al. [28]

Intrafascicular electrodes; longitudinal Pt/Ir wires

Sciatic nerve, cat

Chronic stimulation and recording

O’Donovan et al. [30]

Ventral Root electrodes; up to 12 Pt/Ir microwires Intrafascicular electrodes; 12 electrode silicon array Interfascicular electrodes; Pt contacts on penetrating silicone elements Intrafascicular electrodes; stainless steel wires

L5 Ventral Root, cat

Chronic stimulation and recording

Common peroneal nerve, rat Sciatic nerve, cat

Acute stimulation

Intraneural Electrodes Bowman et al. [2]

Rutten et al. [38]

Tyler et al. [46]

Veltink et al. [49]

Common peroneal nerve, rat

inferred from this figure whether these compressed axons will recover or degenerate. We will address this issue in a future study with histology from chronically implanted animals. DISCUSSION The UEA has been used successfully in stimulating cortical tissues, and in recording single- and multiunit neural activity in sensory and motor cortex of rats, cats, and monkeys on a chronic basis. This suggests that the unique architecture of the UEA may

Acute stimulation

Acute stimulation

Acute stimulation

Stimulation Paradigm

Charge Per Phase

4–5 mA, 10 ␮s duration, 0.5 Hz, monophasic; tetanic contraction 400 ␮A, 600 ␮s duration, 28 Hz, biphasic; stimulation threshold 20–50 ␮A, 100 ␮s duration, 1–3 Hz, biphasic; stimulation threshold

40–50 nC/phase

100–150 ␮A, 200 ␮s duration, single monophasic pulse; stimulation threshold 1–10 mA, 50 ␮s duration, 20 Hz, biphasic; tetanic contraction 120–250 ␮A, 100 ␮s duration, single monophasic pulse; stimulation threshold

20–30 nC/phase

70–200 ␮A, 200 ␮s duration, 50 Hz, monophasic; stimulation threshold 10 ␮A, 200 ␮s duration, 1–2 Hz, biphasic pulse; stimulatin threshold 5–15 ␮A, 100 ␮s duration, single biphasic pulse; stimulation threshold 75 ␮A, 50 ␮s duration, single biphasic pulse; half maximal contraction 5–10 ␮A, 50 ␮s duration, single mono- or biphasic; stimulation threshold 10 ␮A, 100 ␮s duration, single biphasic pulse; stimulation threshold 2.5 mA, 10 ␮s duration, 0.5 Hz, monophasic; stimulation threshold (6.5 mA; 90% Activation) 15 ␮A, 60 ␮s duration, 70 Hz, monophasic pulse; stimulation threshold

14–40 nC/phase

240 nC/phase

2–5 nC/phase

50–500 nC/phase

12–25 nC/phase

2 nC/phase

.5–15 nC/phase

3.75 nC/phase

.25–5 nC/phase

1 nC/phase

25 nC/phase

.9 nC/phase

make it a suitable interface to the peripheral nervous system. Two of the unique aspects of the UEA architecture that have supported its chronic application in the central nervous system (CNS) is its very large surface area and its very thin extraneural silicon substrate from which the electrodes project out into the nervous tissue. When implanted, the 100 electrodes of the UEA provide a mechanical interface, which causes the UEA to completely associate with the cortical tissues into which it is implanted. This provides a mechanical stability that has allowed measurements of identified

MULTIELECTRODE ARRAY FOR SCIATIC NERVE RECORDING/STIMULATION single unit responses from cat visual cortex for up to 4 months [24]. While this architectural feature is an asset once it is implanted, the large surface area of the UEA presents a complication to its use as a multichannel neural interface: how to implant devices of this complexity in nervous tissues. By using a high velocity insertion technique, Rousche and Normann [35] have shown that the UEA can be safely implanted in sensory cortex. These findings have been subsequently validated in scores of implantations of UEAs into sensory and motor cortex. The peripheral nervous system presents the UEA with a number of significant challenges not presented by the CNS. One main concern involves high velocity insertion of the UEA through the epineurium and the perineurium. While the epineurium can be surgically breached and the individual fascicles exposed, such an approach is not only laborious, but traumatizes the nerve and can result in the formation of neuromas. One of the main goals of this study was to see if the insertion of the UEA through the epineurium is possible. If so, this procedure would greatly facilitate the use of the UEA in basic and applied studies of the peripheral nervous system (PNS). The single-unit recordings shown in Figs. 7, 8, and 11 provide compelling evidence that the UEA can be inserted directly through the epineurium and perineurium and that electrodes were located very close to viable neurons. The low stimulation thresholds and the histological analysis additionally indicate that the UEA electrode tips penetrated into individual fascicles and provided a good electrical interface to individual axons in the fascicles. One motivation for the development of the UEA was its potential as a general purpose interface for neuroprosthetic applications. When implanted in the PNS, such a neuroprosthetic system could be used to record sensory information from remnant mechanoreceptors in the extremities, and/or to stimulate motor neurons that would activate various muscles in the skeletal and the urogenital systems [34,54]. Of course, these applications must involve chronically implanted systems, and the work described here has only focused on acute applications. However, we would like to place the UEA in the context of other peripheral nerve interfaces. We chose to limit the comparison to the stimulation capabilities of other devices. Table 2 summarizes these devices. Depending on the literature references, Table 2 contains the stimulus used to obtain either minimal or maximal contraction of the target muscle. It is clear from the table that penetrating electrode arrays have great potential in both basic neuroscience as well as in neuroprosthetic applications in the PNS. With the exception of Loeb and Peck [23], threshold values determined with the UEA are in the range obtained with intrafascicular/intraneural electrodes, whereas interfascicular and extraneural electrodes show higher values. Cuff electrodes generally have threshold values that are one to two orders of magnitude larger. Using the data given, the histogram of the UEA (Fig. 5) suggests that about 20% of the electrodes were either implanted interfascicularly or not close to viable motor neurons. Further, the relative ease by which large numbers of electrodes can be implanted in the peripheral nerve reduces the time and surgical difficulties associated with the implantation of other electrode systems. However, before the UEA can be considered for use in neuroprosthetic applications in the PNS, many problems must be addressed. Perhaps the most difficult problem involves implanting an array in a nerve that experiences large motions relative to surrounding fascia, musculature, and bone. The large surface area of the UEA may help mitigate this problem, but will not eliminate it. An equally perplexing problem has been and remains the tethering forces produced on the implant by the lead wires that must eventually leave the body through a transcutaneous interconnect system. This problem could be circumvented with integrated circuit

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telemetry systems built into the implant, but such systems are experimental [4,45,55] and not yet commercially available. Finally, the long-term biocompatibility of the array in a peripheral nerve must be demonstrated. If these problems can be overcome, implant systems like the UEA that use penetrating electrodes to contact individual nerve fibers could form the foundation of neuroprosthetic systems for a variety of clinical applications: phrenic nerve pacing, control for bowel and bladder voiding, and controlled movement of the musculoskeletal systems. ACKNOWLEDGEMENTS

The authors thank E. M. Maynard, Ph.D., D. J. Warren, and K. S. Guillory for their assistance, and V. Ngo and Y. Zang for manufacturing the electrode arrays. Thanks are also due to R. B. Stein, Ph.D., of the University of Alberta for reviewing this manuscript. This work was supported by the German Academic Exchange Service (DAAD), a State of Utah Center of Excellence Grant, and a NSF grant IBN94-24509.

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