http://en.wikipedia.org/wiki/Functional_electrical_stimulation

http://www.christopherreeve.org/site/c.mtKZKgMWKwG/b.4453425/k.27A5/Functional_Electrical_Stimulation.htm

Functional Electrical Stimulation

Functional Electrical Stimulation (FES) applies small electrical pulses to paralyzed muscles to restore or improve their function. FES is commonly used for exercise, but also to assist with breathing, grasping, transferring, standing and walking. FES can help some to improve bladder and bowel function. There's evidence that FES helps reduce the frequency of pressure sores.
FES made a splash in 1983 when Nan Davis, a paraplegic student at Wright State University, got out of her wheelchair and "walked" to get her diploma. She was powered by an FES system and inspired a TV movie called "First Steps."
The Wright State technology soon emerged commercially in the form of a stationary bicycle (ergometer) called the Regys; users pedaled the bike using FES-stimulated leg muscles. Researchers soon noted that this form of FES provides real aerobic exercise in people who otherwise can't move on their own; it boosts heart and lung function, improves strength and circulation, builds muscle mass, even in people with high quadriplegia

Two companies make FES bikes in the U.S. Therapeutic Alliances, Inc., which originated the Regys 25 years ago, makes the Ergys 2 (www.musclepower.com). Restorative Therapies, Inc. offers the RT300-S which is operated from the wheelchair without the need for a transfer (www.restorative-therapies.com). RTI was started by Dr. John McDonald, the physician who got Christopher Reeve on an FES bike and who has claimed that FES helped Reeve get significant function back seven years after his C1 injury. According to McDonald, the FES bike can be more useful than just exercise. "We propose to use them for a totally different reason -- to promote regeneration and recovery of function." There is to date no support in the medical literature that FES affects recovery.
A doctor's prescription is needed for FES biking; each individual is given a program customized for run times, resistance, etc. The bikes cost in the range of $15,000. The manufacturers have yet to convince Medicare to pay for the devices. Some private insurance companies have reimbursed for them but many people access FES exercise in community settings, at health clubs and rehab clinics.
There are some risks associated with FES. Fracture of leg bones is possible due to loss of bone mineral density. Also, FES can trigger autonomic dysreflexia in upper-level injuries. People with severe spasticity, contractures, or osteoporosis are not good candidates.
Bladder or bowel FES:
Sacral stimulators are surgically implanted FES systems for on-demand control of the paralyzed bladder and bowel; these have been implanted in more than 1,500 paralyzed people, mostly in Europe. The stimulator, called the Finetech-Brindley device, has a strong track record for improving bladder and bowel control in the vast majority of users.
In 1999 a company called NeuroControl licensed the Brindley system and got FDA approval as the Vocare system. A company called NDI Medical more recently obtained the marketing rights to Vocare in the United States. See www.ndimedical.com.
Upper extremity:
About 15 years ago the FDA also approved an FES implant system to restores some hand and arm function to quads. The FreeHand system was well liked by the quads who used it; they gained significant function in grip, writing, eating, computer work, etc. Alas, NeuroControl dropped this from the market.
Walking:
There is a commercially available device called Parastep that is FDA approved for some paraplegics (T4 to T12 ) for "ambulation." Parastep, which has been approved by Medicare for reimbursement, facilitates gait by firing leg muscles; a front-wheeled walker fitted with a control pad is used. Contact www.sigmedics.com.
The future:
Brain-wave communication, it's the next big leap in neuroprosthetics and it's nearly here: in clinical trials, people area already controlling computer cursors and opening email with just their thoughts. Monkeys can precisely move robotic arms using only brain waves.
BrainGate is an investigational brain implant system from a biotech company called Cyberkinetics that places a computer chip into the brain; this monitors brain activity and converts the intention of the user into computer commands. The company is currently recruiting people with spinal cord injury, stroke or muscular dystrophy conditions for pilot clinical trials in Boston, Chicago and Rhode Island. Call the company for more: 508-549-9981.

http://www.wingsforlife.com/projects_detail.php?id=1&year=1

ELECTRICAL STIMULATION (EPIDURAL) OF THE SPINAL CORD TO ACHIEVE STANDING AND WALKING

Dr. Helmut Kern, MD / PhD, (Assoc. Prof.) (2006)
Wilhelminenspital, Vienna, Austria
Goal:	Rehabilitation
Method:	Epidural stimulation, electrophysiology

Networks of nerve cells located in the spinal cord control the musculature. The muscles are activated in a particular sequence, for example during walking. The control only works if it continuously receives electrical signals from the brain. In patients with paraplegia, the nerve fibres which transmit these signals to the spinal cord are severed. Previous research has shown that the musclecontrolling networks can be reactivated by external electrical stimulation of the injured spinal cord. This allows basic motor functions to be reestablished. Electrodes are attached subcutaneously near the spinal cord. The spinal cord is then electrically stimulated. Simple movement tests and measurements of muscle activity determine whether the motor function has been successfully activated. The project is supposed to enable paraplegic patients with residual control over the leg musculature to walk a few steps or to stand up from their wheelchair by means of electrical stimulation.
Status of the project: The project was ended on June 2007.
Results: This project was rated as successful. Stimulation of the posterior spinal cord from the surface of the body can recruit the appropriate neural structures to 1) generate muscle force and co-ordination effective to induce standing-up movement and 2) standing in complete (ASIA A,B) as well as incomplete spinal cord injured subjects (ASIA C) and to 3) improve functional muscle activation patterns for stepping in incomplete spinal cord injured
subjects (ASIA D).
Publications: Work resulting from this project was published in the journal ¨Artificial Organs〃 (Hofstoetter US, Minassian K, Hofer C, Mayr W, Rattay F, Dimitrijevic MR. Modification of reflex responses to lumbar posterior root stimulation by motor tasks in healthy subjects. Artif Organs. 2008 Aug;32(8):644-8.).

http://edition.cnn.com/2011/HEALTH/05/19/paralysis.implant.treatment/index.html

Doctors: Electrical implants a big step in paralysis treatment

By Val Willingham, CNN

May 20, 2011 -- Updated 0527 GMT (1327 HKT)

(CNN) -- Electrical stimulation from a spinal cord implant, mimicking the signals the brain would normally transmit to move the body, has allowed a paralyzed patient to stand on his own and walk on a treadmill with assistance, researchers said Thursday.
At a news conference in New York City, doctors introduced Rob Summers, a 25-year-old, paralyzed spinal cord injury patient from Oregon. He received continual epidermal stimulation to the lower spinal cord, researchers said, enabling the muscle and joint movements that are required to stand and, with assistance, to step.
"This stimulation causes changes in the brain and changes in the spinal cord," said Dr. V. Reggie Edgerton, a professor in the Departments of Integrative Biology and Physiology and of Neurobiology at the University of California, Los Angeles. "Now we need to know how this has occurred."
A therapy already in use in a number of U.S. hospitals is known as functional electronic stimulation (FES), in which stimulants or electrodes are placed on the skin above the muscles to help chronic spinal cord injury patients move their limbs and, in some cases, walk with assistance.
In the newly unveiled procedure, however, the electrodes are implanted in the patient's body. But even beyond that, there are distinct differences in the two procedures, the researchers said.
"In our case it's a constant signal," said Dr. Susan Harkema, a professor at the Department of Neurological Surgery at the University of Louisville. "FES tries to bypass the circuitry by stimulating the muscles externally with electrodes. We allow the spinal cord to control the muscles on its own."
The researchers said in the initial phase of their experiment, Summers was able to reach a standing position, supplying the muscular push himself while his spinal cord was being stimulated electrically. He could stand independently, bearing full weight, for up to four minutes at a time He also began to move his toes, ankles, knees and hips while being stimulated.
"This does not represent a cure for spinal cord injury," Edgerton said, "but it's something to build on."
At today's press conference, Summers, who became paralyzed in his lower extremities following a hit-and-run accident in July 2006, talked about his experience.
"I was able to stand independently, the third day we turned it on," said Summers. "I was amazed. I use it two hours a day now, and I hope to someday go back to playing baseball."

The project, which is outlined in the latest issue of the medical journal Lancet, was funded by the Christopher and Dana Reeve Foundation, which provides money for research on spinal cord injuries.

See images of functional electrical stimulation

http://www.youtube.com/watch?v=Lxl6xT3L3Qc

http://www.mrisafety.com/safety_article.asp?subject=182

RF Bion Microstimulator

Surgically implanted neurostimulators and electrodes may be utilized to provide functional electrical stimulation of the affected site. However, these devices may be associated with considerable surgical morbidity and expense. As such, there has been an on-going effort to develop technology that combines the reliability of using an implanted device with a low morbidity and low cost procedure. This effort has yielded a miniaturized, implantable device designed for functional electrical stimulation.

In 1988, Heetderks first demonstrated the feasibility of using a millimeter-sized, neural prosthetic implant. Over the years, this so-called "microstimulator" evolved to its present form. The microstimulator now exists as a relatively small, wireless, digitally controlled stimulator that is implanted using a minimally invasive procedure to provide electrical pulses to a muscle or nerve. This device receives power and command signals by inductive coupling from an externally worn coil that generates a radiofrequency magnetic field. The microstimulator is currently undergoing clinical trials to assess its therapeutic effect on a variety of neurological disorders including urinary incontinence, shoulder subluxation, drop foot, ventilator-dependant respiratory deficiencies, and sleep apnea.

RF BION Microstimulator
The implantable RF BION Microstimulator (Alfred E. Mann Foundation for Scientific Research, Valencia, CA and Advanced Bionics Corporation, Valencia, CA) is a wireless device designed for functional electrical stimulation of the peripheral nervous system. This hermetically-sealed implant is a small, lightweight, cylindrical-shaped device (length, 16.6-mm; diameter-2.4 mm; mass, 0.265-g) made of a ceramic tube closed on each end by titanium caps and contains components made from titanium, gold, copper, ferrite, platinum, iridium, silicon, zirconium, and tantalum. The active electrodes are welded on each end cap: an iridium disk on the cathodal side and a platinum-iridium eyelet on the anodal side.

The microstimulator receives power and digital commands via a 2-MHz radiofrequency magnetic field link generated from an external coil that is worn by the patient. The RF Bion Microstimulator produces asymmetric, biphasic, capacitively coupled constant-current pulses. Stimulation circuitry and a receiving multi-turn loop antenna are contained within the microstimulator. The antenna is wound around two pieces of ferrite, cut from a cylinder of radius 0.74-mm. Because of the small size of this microstimulator, it may be implanted through a specially designed, trocar-based 12- or 14-gauge implant tool or via a small surgical opening for placement near a nerve or at the motor unit of a muscle. A suture passed through one of the end caps allows the microstimulator to be maintained acutely in a properly implanted position.

MRI and the RF BION Microstimulator
Safety information for the use of a magnetic resonance imaging (MRI) procedure (i.e., imaging, angiography, functional imaging, spectroscopy, etc.) in a patient with the RF BION Microstimulator is highly specific to the type of MR system and conditions used to determine safety criteria for this implant. Safety information for MR procedures described herein pertains to the use of MR systems operating with static magnetic fields of 1.5-Tesla, gradient magnetic fields of 20-Tesla/second or less, and a whole body averaged specific absorption rate (SAR) of 2.0 W/kg or less for 15-minutes of imaging (i.e., per pulse sequence).

Warning
The effects on the RF BION Microstimulator of performing MRI procedures using other MR systems and other conditions have not been determined.

The following information describes testing that was performed on the effect of magnetic resonance imaging (MRI) procedures on an implanted RF BION Microstimulator.

MR procedures MUST ONLY be performed according to the following information:

Magnetic Field Interactions.
Magnetic field interactions (translational attraction and torque) were assessed for the RF BION Microstimulator in association with exposure to a 1.5-Tesla MR system. The RF BION Microstimulator displays relatively moderate magnetic field interactions. [The deflection angle measured for the RF BION Microstimulator was 58 degrees. Thus, the magnetic force is 1.6 times the gravitational force (0.42 g). The maximal magnetic torque was calculated to be equal to 3.9 times the "gravity torque", the product of the implant's length and weight. Therefore, torque can be viewed as a force equivalent to 0.98 g acting on one end of the RF BION Microstimulator with the other end fixed.] Because of the internal ferrite material (mass of 32 mg), the intended in vivo use of this implant must be considered.

Notably, tissue encapsulation around the RF BION Microstimulator is a major factor that will retain this implant in situ. In vivo research demonstrated that counter-forces provided by encapsulation (i.e., fibrous tissue) that occurs 2 to 3 weeks after implantation will prevent the RF BION Microstimulator from being moved or dislodged in association with the exposure to a 1.5-Tesla MR system. Therefore, considering that the mass of the RF BION Microstimulator is relatively small (0.25 g), and this implant does not contain any sharp ends, magnetic field interactions associated with exposure to a 1.5-Tesla MR system will not move or dislodge this implant after a conservative post-op waiting period of 6 to 8 weeks.

Interactions with Time-Varying Magnetic Fields.
The potential interactions of the RF BION Microstimulator with magnetic resonance imaging (MRI)-related time-varying magnetic fields (gradient and RF) were theoretically and experimentally investigated with respect to a 1.5-Tesla/64 MHz MR system.

Effects of Radiofrequency Fields.
Based on theoretical analysis and in vitro experiments, the presence of the RF BION Microstimulator in a patient undergoing an MRI procedure at a whole-body-averaged specific absorption rate (SAR) of 2.0 W/kg for 15-min. (per pulse sequence) will not result in a substantial increase in temperature. Therefore, a patient with the RF BION Microstimulator is not at increased risk with respect to MRI-related heating of this implant under the conditions used for this assessment.

Effects of Gradient Fields.
The electric field in the body induced by the MRI pulsed gradient magnetic fields may exceed 5 V/m. Given the size and shape of the RF BION Microstimulator, this field intensity induces minimal voltage (95 mV) along the length of the implant. Importantly, concentration of the gradient currents will be no greater than that resulting from the concentration rising from bones surrounded by tissue. Therefore, no physiological impact is expected by the interaction between the RF BION Microstimulator with the gradient currents in the tissue of a patient with this implant undergoing an MR procedure at 1.5-Tesla/64 MHz. Accordingly, a patient with the RF BION Microstimulator may safely be exposed to the gradient magnetic fields (dB/dt) at levels that are currently allowed for 1.5-Tesla MR systems operating with conventional pulse sequences and standard accessories (e.g., surface RF coils).

Warning: Unconventional or non-standard MR techniques have not been assessed for the RF BION Microstimulator and, therefore, must be avoided.

MRI Procedures and Function of the RF BION Microstimulator.
The effects of MRI procedures on the functional aspects of the RF BION System were assessed. In vitro testing was performed on a phantom with nine RF BION devices placed in various orientations relative to the MR system.

MR imaging was conducted using 1.5-Tesla/64 MHz MR system (transmit/receive RF body coil) to perform 15 different MR imaging pulse sequences, as follows: (1) T1-weighted, spin echo pulse sequence, (2) T1-weighted, fast spin echo pulse sequence, (3) T2-weighted, spin echo pulse sequence, (4) T2-weighted, fast spin echo pulse sequence, (5) Gradient echo, two dimensional pulse sequence, (6) Gradient echo, three dimensional pulse sequence, (7) Fast gradient echo, two dimensional pulse sequence, (8) Fast gradient echo, three dimensional pulse sequence, (9) Fast spoiled gradient echo, two dimensional pulse, sequence, (10) Fast spoiled gradient echo, three dimensional pulse sequence, (11) T1-weighted, spin echo pulse sequence with high whole-body averaged specific absorption rate (1.1 W/kg), (12) T1-weighted, fast spin echo pulse sequence with high whole-body averaged specific absorption rate (1.1 W/kg), (13) T2-weighted, spin echo pulse sequence with high whole body averaged specific absorption rate (1.1 W/kg), (14) T2-weighted, fast spin echo pulse sequence with high whole-body averaged specific absorption rate (1.1 W/kg), and (15) Gradient echo, three dimensional pulse sequence with magnetization transfer contrast with high whole-body averaged specific absorption rate (1.1 W/kg)

The findings indicated that there was no apparent damage or alteration in the functional aspects of the RF BION Microstimulators. Thus, the RF BION Microstimulator was demonstrated to maintain full functionality after exposure to the MR imaging conditions indicated above.

Warning
The effect of using any other MR imaging pulse sequence or procedure on the RF BION Microstimulator is unknown.

Additional MRI Safety Guidelines for the RF BION Microstimulator
-DO NOT use MR systems other than 1.5 Tesla MR systems.
-Continuously monitor the patient using visual and audio means (i.e., intercom system) throughout the MR procedure.
-Instruct the patient to alert the MR system operator of any unusual sensations or problems so the MR system operator can terminate the MRI procedure, if needed.
-Provide the patient with a means to alert the MR system operator of any unusual
sensations or problems that may be experienced during the MR procedure.

Post-MRI Procedure Recommendations for the RF BION Microstimulator
After undergoing an MRI procedure, the RF BION Microstimulator should be checked to ensure that it is working properly. To do so, stimulation threshold measurements should be obtained and compared to pre-MRI procedure threshold levels.

http://www.medscape.com/viewarticle/542356_2

The BION Devices: Injectable Interfaces With Peripheral Nerves and Muscles.: Clinical Need

Many neurological and orthopedic disorders reduce or eliminate voluntary recruitment of muscles. Such loss degrades the ability to perform motor tasks or to maintain muscles, connective tissues, and metabolic systems that depend on muscle activity for their function and integrity. Clinicians have long understood the usefulness of electrical stimulation to activate paralyzed and paretic muscles,4 but electrical muscle stimulation has not gained widespread acceptance as a standard clinical therapy. The poor acceptance seems linked to limitations in the previously available technologies, which are described as follows. 1) Transcutaneous stimulators using surface electrodes on the skin are difficult to position. It is also difficult to adjust these devices to achieve selective and reproducible effects on muscles. At stimulation intensities sufficient to activate muscles, they typically stimulate skin afferents as well. Recruitment of skin afferents may give rise to unpleasant sensations. Furthermore, surface electrodes may produce cutaneous irritation and may be considered cosmetically unappealing. 2) Percutaneous wire electrodes are mechanically vulnerable, may fail mechanically from repeated flexion, and may provide a conduit and nidus for chronic infection. 3) Surgically implanted multichannel stimulators are expensive and invasive. Thus, they are impractical for many applications, particularly in prevention and rehabilitation. Implanted stimulators tend to be designed for very specific sites and uses. Consequently, a particular system can be difficult to customize to the special needs of the patient or to extend or modify function after implantation.

Technology

In this report, we describe a new "platform" technology that was designed to have broad applicability, as described later. It consists of wireless micromodules, each of which receives power and command signals by inductive coupling from an external antenna.^[8,34,36] One or more of these BIONs can be injected in a simple outpatient procedure similar to a Botox injection. Four generations of this technology are now in clinical trials or are being developed. All are designed to stimulate myelinated sensory or motor axons, typically in peripheral nerves or muscles. They are not used to stimulate muscle fibers directly because muscle fibers are much less excitable and generally require pulse strengths that are not achievable with this technology.^[30]

The BION1: Radiofrequency-Powered, Continuously Controllable Stimulator

As originally proposed by Loeb, et al.,^[36] BION1 single-channel implants receive power and digital command signals from an inductive coil that is worn over or positioned near the implants. An implant generates a stimulation pulse when it detects a match to its address. The pulse duration can be set from 2 to 512 µsec in 2-µsec steps, and the pulse current is regulated from 0.2 to 30 mA (17-V compliance) in two ranges of 16 steps each.^[8] Up to 3000 commands per second can be transmitted, permitting fine control of activation in many muscles concurrently. Such implants are intended primarily for therapeutic muscle stimulation, in which patterns of electrical stimulation are used to build strength in hypotrophic muscles. Extensive preclinical testing has demonstrated that these implants form a mechanically and functionally stable and biocompatible interface when used to stimulate muscles continually.^[7,19,35] From results obtained in accelerated in vivo testing, we infer that the implants will function indefinitely.^[35] The clinical experience as of this writing is described in our article.
One version of this device (BION1 AMI; Fig. 1) is produced by the Alfred Mann Institute for Biomedical Engineering at USC in Los Angeles, California, and has been approved for and used in clinical investigational studies in the US, Canada, and Italy (see later discussion). Its electronic components are housed in a hermetically sealed glass capsule, which is 2 mm in diameter ⊙ 16 mm in length. Each device delivers monophasic stimulation pulses through a tantalum capacitor electrode, which is immediately recharged to the compliance voltage to maintain the charge balance required for safe long-term use in the body. The counterelectrode is made from pure iridium, which forms a conductive, nonpolarizing oxide that can be used safely to deliver the maximal stimulus pulses.^[44] Another version of this technology with the same electronic functionality is produced by the Alfred Mann Foundation for Scientific Research in Valencia, California, for additional clinical trials not described in this article. It is housed in a ceramic package (2.5-mm diameter ⊙ 16.5-mm length) and incorporates an internal electrolytic capacitor to achieve charge balance through platinumiridium electrodes (BION1 AMF; Fig. 1).

(Enlarge Image)

Figure 1. Photograph depicting the BIONs now in clinical use. ABC = Advanced Bionics Corp.; AMF = Alfred Mann Foundation; AMI = Alfred Mann Institute; dia = diameter.

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Description: http://img.medscape.com/fullsize/migrated/542/356/nf542356.fig1.gif

Figure 1.

Photograph depicting the BIONs now in clinical use. ABC = Advanced Bionics Corp.; AMF = Alfred Mann Foundation; AMI = Alfred Mann Institute; dia = diameter.

The BION2: Radiofrequency-Powered, Continuously Controllable Stimulator and Sensor

Work is underway to add sensing and back-telemetry systems to the BION1 platform. A new low-frequency inductive link can transmit a mix of inward command data (128 kb/second) and outward sensor data (40 kb/second) over the same 480-kHz carrier that provides power for the implants.^[55] The sensory modalities being developed include electromyographic envelope detection (for myoelectric control), a two-axis direct-current accelerometer based on microelectromechanical systems technology,^[61] and a form of "BIONic muscle spindle" based on range-finding readings between implants that move with the muscles in which they are placed.^[35] Implants with both sensing and stimulation capabilities are intended for progressively more ambitious research on FES to reanimate paralyzed limbs (Fig. 2). The more advanced clinical applications such as reach and grasp function for quadriplegic patients depend on parallel development of control algorithms and fitting software that are outside the scope of this report.^[12-14]

(Enlarge Image)

Figure 2. Artist's conception of radiofrequency-powered BION2 implants being used to stimulate contractions in paralyzed muscles and to sense and transmit outward information about the posture and movement of the limb to an external control unit.

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Description: http://img.medscape.com/fullsize/migrated/542/356/nf542356.fig2.gif

Figure 2.

Artist's conception of radiofrequency-powered BION2 implants being used to stimulate contractions in paralyzed muscles and to sense and transmit outward information about the posture and movement of the limb to an external control unit.

The BION3: Rechargeable, Battery-Powered, Programmable Stimulator

Advanced Bionics Corp. in Valencia, California, produces a commercial implant, "bion" (labeled BION3 ABC in Fig. 1), that is approved for sale in Europe and is undergoing premarket approval in the US. It incorporates a rechargeable lithium cell in a somewhat larger ceramic package (3.3-mm diameter ⊙ 27-mm length). It can be programmed to produce various preset patterns of stimulation for a few days (depending on stimulus parameters) before being recharged by an inductive link similar to that used in the radiofrequency-powered BIONs (BION1 and BION2). It is designed primarily for neuromodulation applications, in which electrical stimulation is used to modify neural activity in a dysfunctional pathway (for example, urinary urge incontinence; see later section). The battery life is projected to be 5 to 10 years depending on discharge and recharge cycles.

The BION4: Rechargeable, Battery-Powered, Intercommunicating Stimulator and Sensor

As the sophistication of FES applications grows, the numbers and physical distribution of channels required for stimulation and sensing will grow beyond feasibility for low-frequency, radiofrequency-powered devices such as the BION2. Researchers at the Alfred Mann Foundation are working on a new class of BION implants with rechargeable battery power similar to the BION3 but with a high data-rate communication protocol that allows large numbers of implants to exchange information freely among themselves and with an external controller

http://www.rehab.research.va.gov/jour/03/40/6/kilgore.html

Durability of implanted electrodes and leads in an upper-limb neuroprosthesis

Kevin L. Kilgore, PhD; P. Hunter Peckham, PhD; Michael W. Keith, MD; Fred W. Montague, MS;
Ronald L. Hart, MS; Martha M. Gazdik, BS; Anne M. Bryden, OTR/L; Scott A. Snyder, PhD;
Thomas G. Stage, BS

Department of Orthopaedics, MetroHealth Medical Center, Cleveland, OH; Department of Biomedical Engineering, Case Western Reserve University, Cleveland, OH; Department of Veterans Affairs, Rehabilitation Research and Development Service, Cleveland, OH

Abstract  Implanted neuroprosthetic systems have been successfully used to provide upper-limb function for over 16 years. A critical aspect of these implanted systems is the safety, stability, and reliability of the stimulating electrodes and leads. These components are (1) the stimulating electrode itself, (2) the electrode lead, and (3) the lead-to-device connector. A failure in any of these components causes the direct loss of the capability to activate a muscle consistently, usually resulting in a decrement in the function provided by the neuroprosthesis. Our results indicate that the electrode, lead, and connector system are extremely durable. We analyzed 238 electrodes that have been implanted as part of an upper-limb neuroprosthesis. Each electrode had been implanted at least 3 years, with a maximum implantation time of over 16 years. Only three electrode-lead failures and one electrode infection occurred, for a survival rate of almost 99 percent. Electrode threshold measurements indicate that the electrode response is stable over time, with no evidence of electrode migration or continual encapsulation in any of the electrodes studied. These results have an impact on the design of implantable neuroprosthetic systems. The electrode-lead component of these systems should no longer be considered a weak technological link.

Key words: electrodes, functional electrical stimulation, implants, neuroprosthesis, spinal cord injury, survival analysis.

Abbreviations: AbPB = abductor pollicis brevis, EDC = extensor digitorum communis, EIP = extensor indicus proprious, EPL = extensor pollicis longus, FDP = flexor digitorum profundus.

This material was based on work supported by the Department of Veterans Affairs, Rehabilitation Research and Development Service; the National Institutes of Health, Neural Prosthesis Program, contract NO1-NS-2344; the National Institutes of Health, Clinical Research Center, Case Western Reserve University, M01 RR00080-31; and the Food and Drug Administration, Orphan Products, grant FD000832. Support for this study was also provided by the Department of Orthopaedics, MetroHealth Medical Center, Cleveland, Ohio.

Address all correspondence and requests for reprints to Kevin L. Kilgore, PhD; MetroHealth Medical Center, Hamann 601, 2500 MetroHealth Drive, Cleveland, OH 44109; 216-778-3480; klk4@cwru.edu.

INTRODUCTION

Implanted neuroprosthetic systems have been successfully used to provide motor function in spinal cord injury (SCI). To date, the majority of these systems have used multiple muscle-based electrodes and a central implanted pulse generator located in the torso. The design of these systems, therefore, puts stringent requirements on the electrode, lead, and connector assembly [1].

Although electrodes and leads are common components of implanted systems, such as pacemakers, cochlear implants, respiratory assist devices and bladder/bowel stimulators, these systems do not place the same requirements on the electrode and lead as motor neuroprotheses. First, multiple leads are typically required, because multiple muscles are activated and electrodes are distributed over a relatively large area of the body. Second, the leads are required to cross many joints so that electrodes reach some of the more distal muscles. Third, fine control over the degree of contraction is required to produce highly coordinated movements.

In this paper, we report retrospectively on the reliability and stability of electrode, lead, and connector systems used in a neuroprosthesis providing upper-limb function [1-4]. These systems used 8 or 10 functioning electrodes and required some leads to cross the shoulder, elbow, and wrist joints to stimulate muscles that must be activated in a graded fashion. The clinical performance and acceptance of first- and second-generation upper-limb neuroprostheses have been reported previously [3,4]. These systems provide a stringent test for implanted electrode-lead systems.

METHODS

Neuroprosthetic System Description

The implanted neuroprosthetic system used in this study consisted of an implanted stimulator, electrodes, leads, and connectors, as well as external components, as shown in Figure 1. A radio frequency inductive link was used to communicate with, and power, the implanted device. The implanted stimulator had 8 or 10 leads connected to stimulating electrodes. An in-line connection was used to connect the implant device to each lead intraoperatively. Each electrode was implanted in a separate location, typically in the muscles of the hand, forearm, and upper arm. In some cases, an electrode was placed in the supraclavicular region to provide sensory feedback.

Description: Figure 1. Implanted upper-limb neuroprosthesis used in this study.

Muscle activation was accomplished through electrical stimulation. Stimulus pulses consisted of a constant current balanced charge waveform. The stimulating phase was a square cathodic pulse of 0 to 200 µs in duration and 20 mA in amplitude. In infrequent cases, a stimulation amplitude of less than 20 mA was used. Stimulus frequency was 12 to 16 Hz for all muscle-based electrodes. For sensory feedback, frequency varied from 4 to 60 Hz.

Electrode-Lead Design

Stimulus output was delivered through lead wires that traveled subcutaneously to the distal electrode termination. Each lead had an in-line connector. Two types of electrodes have been used. These components are shown in Figure 2 and are described in the following paragaphs.

Description: Figure 2. Electrode, lead, and connector components analyzed in this study.

Lead

The lead cable consisted of seven Type-316LVM stainless steel wires (each was 0.034 mm in diameter, Fort Wayne Metals, Fort Wayne, Indiana), organized into a single seven-filament strand and insulated with PFA-Teflon^TM (Dupont, Wilmington, Delware) by Temp-Flex (S. Grafton, Massachusetts). The lead was fabricated by winding two lengths of cable in tandem, forming a double helix of two conductors electrically insulated along their length but shorted to each other at the electrode and connector end. The coiled lead was placed inside medical-grade Silastic^TM (Dow Corning, Midland, Michigan) tubing. The lead outer diameter was approximately 1.3 mm.

Connector

A single-conductor in-line connector was used to independently connect each electrode to the implanted stimulator. The connector consisted of two male plugs that mate with a center spring and were enclosed by an external insulating cuff. The male plug and the associated strain relief spring were integral parts of the terminal end of both the implant lead and the electrode lead. The connector measured 30 mm long from end to end of the strain relief springs. The outside diameter measured 3.5 mm at the suture cuffs (widest point) [5].

Electrodes

Implanted electrodes were used to direct the stimulus current so that it activated the desired neural structures. Two types of electrodes were used in upper-limb implanted systems during this study. They were epimysial and intramuscular electrodes [6,7].

Epimysial Electrode.

This electrode was placed on the muscle surface. The electrode was a Pt-Ir disk embedded in a silastic backing. The lead wire was welded to the back of the disk with a resistance welder in an inert gas environment. The silastic backing was reinforced with Dacron. The electrode was sewn onto the muscle epimysium with sutures through the Dacron backing. With 4-0 Dacron suture, five sutures were placed around the perimeter of the electrode backing, each tied with four knots. The diameter of the stimulating disk was 5 mm.

Intramuscular Electrode.

This electrode was inserted into the muscle belly. The electrode consisted of 316LVM stainless steel wire coiled around the outside of the lead tubing. The stimulating surface was 2 mm long and had an approximate surface area of 14.5 mm². A 2 mm-long polypropylene barbed anchor on the tip of the electrode maintained the position of the electrode in soft tissue. The electrode was inserted into the muscle with a probe and cannula system, as described by Memberg et al. [7].

Y-Branch

The total number of leads coming from the implant can be reduced by taking advantage of the two separate conductors in each lead wire [8]. A Y-branch termination was constructed that terminated in two separate male plugs to obtain separate access to each conductor. The individual cables were split during fabrication to form a Y-junction or Y-branch. Each single conductor cable was formed into a separate lead wire with the use of the same fabrication methods used for the two conductor leads. This Y-branch had no splice or connection between conductors of the tandem and single helix leads. The Y-branch was reinforced with a molded silastic strain relief and was only used in devices with 10 electrodes.

Subjects

Data were collected from a single clinical series conducted at two sites: the Louis Stokes Department of Veterans Affairs (VA) Medical Center and MetroHealth Medical Center, both in Cleveland, Ohio. All subjects had sustained a traumatic SCI and were tetraplegic at the C5 or C6 level. All subjects were implanted by one surgeon (MWK), all procedures and technology received appropriate Institutional Review Board (IRB) and Food and Drug Administration (FDA) approvals prior to usage, and all subjects gave appropriate written consent.

Preoperative Preparation

Before surgery, each subject underwent a period of muscle conditioning using surface stimulation for at least 1 month [2]. This produced stronger, more fatigue-resistant muscles in preparation for intraoperative placement of electrodes.

Implantation Procedure

The implant device, leads, and electrodes were placed during an operative procedure with the patient under general anesthesia, as described in detail elsewhere [1]. Epimysial electrodes were sutured on the muscles in an open surgical procedure. Intramuscular electrodes were inserted into the muscle belly with the use of a probe and cannula. Leads were tunneled from the electrodes up the arm to a connector in the mid-humeral area. Tunneling was accomplished with a blunt plastic probe that was 6 mm in diameter and approximately 30 cm long (Scanlan tunneler, Scanlan International, St. Paul, Minnesota). Suction was placed on one end of the tunneler, and the leads were fed into the opposite end. Sterile saline was used to cause the leads to be drawn into the tunneler. When the leads were completely drawn into the tunneler, it was withdrawn, leaving the leads in place. Up to 10 leads can be passed through a single tunneler with this method.

The stimulator unit was implanted in the pectoral region with the leads tunneled subcutaneously to the humeral connector site, where the spring connector was used to connect the lead from the stimulator unit to the lead from the electrode. Where leads crossed joints, the surgeon routed the lead near the neutral axis whenever possible to try to minimize the stress on the leads. Postoperatively, the patient was placed in a long arm cast for 3 weeks for electrode stabilization. Elective modifications to the postoperative care were made if the patient received additional surgical alterations to the upper limb, such as tendon transfers, with the result that some patients remained in a cast for as long as 4 weeks. After cast removal, muscle conditioning with the use of the neuroprosthesis was initiated at a low level for the muscle strength to be rebuilt. This condition typically continued for at least a month before the patient was trained to use the system for functional activities. Patients underwent 1 to 3 weeks of rehabilitation training, after which they were released home to use the device for functional activities.

Monitoring of System Integrity

We monitored the implanted neuroprosthetic system to verify the biological safety and integrity of the implanted components. The mechanical and electrical integrity of the implanted components was also monitored. These tests included intraoperative verification, X rays, electrode thresholds, surface potential measurements, and patient and staff reports of technical and medical incidents.

Intraoperative Verification

Prior to final wound closure, the entire system was tested with a sterile radio frequency (RF) link to the implanted stimulator. The surgeon tested each electrode individually to ensure that the proper response was obtained by visual observation of joint movement.

X rays

X rays of the entire upper limb and shoulder were taken at the time of surgery. Subsequently, if an unexplained change in electrode response was identified by the subject or investigator, the images were repeated. Leads, electrodes, and connectors were easily identified on a standard X ray, but only gross separation of wires could be identified with this method (i.e., the fractured ends need to be separated by at least 1 mm).

Electrode Thresholds

The electrode threshold was defined as the lowest stimulus level at which a visible response was obtained [9]. In all cases, the stimulus pulse amplitude was 20 mA and the frequency was 12 Hz. We recorded thresholds by slowly increasing the stimulus pulse duration in 1 µs increments to each electrode individually while observing each upper-limb joint for movement. Thresholds were recorded at 1 to 2 months postimplant and at 6 and 12 months postimplant. Thresholds were repeated any time the subject reported a possible change in grasp response.

Surface Potential Measurements

The presence of an electrical stimulus from an implanted system can be accurately recorded with the use of surface electrodes on the skin [10]. A reference electrode was placed over the implant capsule where the anode was located. The sensing electrode was placed at variable distances along the arm (typically 10 cm intervals). A voltage potential was recorded during the delivery of the stimulus. The voltage increased as a function of the distance between the two electrodes and could be as high as 6 V. A broken lead was confirmed by changes in the expected surface potential map and by changes in the stimulus waveform recorded on the skin surface [10].

Medical and System Incident Reports

Physiological and technical incidents were recorded and evaluated as reported by each subject. Subjects were regularly contacted and queried regarding incidents during their first 2 years postoperative, and followed at approximately yearly intervals thereafter. Subjects were instructed to contact the research staff before having surgery to allow antibiotics to be prescribed.

Statistical Methods

Threshold Stability

Electrode threshold measurements were recorded at irregular intervals for subjects once they were beyond 1-year postimplant. To determine whether some electrodes demonstrated a long-term increase in threshold as a function of time, we examined each electrode that met the following criteria:

1. Electrode threshold recorded at least 60 days postimplant.

2. At least four threshold values, each recorded on different days.

3. Threshold data must span at least a 2-year interval.

A line was matched to the threshold data points with the use of a linear least squares estimation. We evaluated the slope of the estimated line to determine if it was statistically significantly higher than zero at p = 0.05. Where the slope was statistically higher than zero, the projected increase in threshold over 50 years was estimated. If the threshold increase over that period was projected to be less than 50 µs (i.e., slope <1 µs/yr) , the slope was determined to be insignificant clinically. A threshold increase of this rate would not be functionally noticeable to the user and, in the worst case, might require reprogramming of the neuroprosthesis approximately every 20 years.

Electrode data during the first 60 days postimplant were not considered in this analysis because electrode thresholds undergo a settling period during that time. Typically thresholds are slightly higher immediate postoperative and then settle to a consistent value at 2 to 3 months postimplant [1].

Correlation of Intraoperative and Postoperative
Electrode-Muscle Output Characteristics

It is desirable to determine if the electrode-muscle output characteristics (i.e., the muscle forces as a function of stimulus level) that are observed intraoperatively correlate to the postoperative output characteristics [9,11,12]. A poor correlation would result in many electrodes needing to be repositioned surgically. However, obtaining direct measurements of the electrode-muscle output characteristics intraoperatively is difficult because quantitative measurements of muscle force and joint position are time-consuming and instrumentation-intensive. Therefore, we have established four easily recorded criteria to estimate this correlation. These are-

1. Accurate recruitment of targeted muscle. The postoperative electrode response should demonstrate that the first muscle recruited by the electrode is the muscle that was targeted intraoperatively.

2. Isolated response. Stimulus delivered to the electrode should only recruit the muscle or muscle groups that were targeted intraoperatively. If the electrode recruited a second muscle within a pulse duration change of 2 µs from the target muscle, then it was defined as nonisolated [11].

3. Low threshold. The threshold should be below 50 µs

4. Electrode moved or replaced. An electrode that was moved (surgically) or replaced indicates a possible failure of identifying the desired response intraoperatively.

Survival Analysis

A Kaplan-Meier analysis was used [13], which enables data from surviving and failed electrodes to statistically predict the longevity of a population. The proportion surviving is equal to the geometric sum of one minus the ratio of the number of failures to the number of electrodes at risk for failure. The predicted surviving proportion is constant between points of failure. Therefore, a meaningful confidence interval for survival can only be calculated at points of failure. The hazard rate was also calculated at each failure point, with the use of the formula of number at risk divided by the interval between failures.

RESULTS

Between August 1986 and December 1999, 28 arms in 27 patients received the upper-limb neuroprostheses. Twelve arms had C5 motor function and the remaining sixteen arms had C6 motor function. Patients were at a median age of 32 years at the time of implantation (range 21 to 47 years). The active use of the neuroprosthesis varied considerably among the patients, from regular daily use to occasional use [3,14]; therefore, the total stimulation time experienced by each electrode varied from approximately 1,000 hours to over 50,000 hours. Overall, 238 electrodes were implanted, with an average follow-up time of 7.1 years (range: 3.2 to 16.4 yr). There were 204 epimysial electrodes and 34 intramuscular electrodes. Intramuscular electrodes were not introduced until the tenth subject in the series (1995) and were not used extensively until 1997. One subject has nine intramuscular electrodes and one epimysial electrode.

Electrodes were grouped into various "regions" of the body, depending upon the location of the motor point or skin area (for sensory electrode) of the muscle to be excited. The electrode locations are shown in Table 1. Total lead length (from stimulator package to electrode termination) varied from 28 to 83 cm and depended on electrode placement, lead routing, and subject size.

Table 1. Electrode locations and lead length.
Location	Distribution (%)	Average Total Lead Length (cm)
Chest	6	32
Upper Arm	4	38
Forearm	64	55
Hand	26	75

No cases have been reported where failure of a component of the neuroprosthesis resulted in the inability of the subject to use the neuroprosthesis for functional activities. Of the 238 electrodes in the series, 234 (98.3%) remained intact throughout the study. Three (1.3%) were broken and one (0.4%) was infected. These electrodes are discussed in detail subsequently. Three subjects died during the study at 3.3, 6.0, and 9.4 years postimplantation (representing 25 electrodes). Therefore, 209 electrodes continue to be used in functioning neuroprosthetic systems.

Survival analysis using Kaplan-Meier showed that there was a 98.7 percent probability for an electrode to be intact at 16 years, as shown in Figure 3. At the latest failure point (1.9 years), the 95 percent confidence interval for the survival probability is 97.3 to 100.0 percent. The hazard rate at each of the three failure points ranged from 0.6 percent a year to 0.8 percent a year. At present, 31 electrodes are older than 10.0 years, 102 electrodes older than 7.5 years, and 191 electrodes older than 5.0 years.

Description: Figure 3. Kaplan-Meier survival analysis for all electrodes and leads.

Correlation of Intraoperative and Postoperative
Electrode-Muscle Output Characteristics

Less than 1.5 percent of the electrodes exhibited a lower threshold for a muscle that was not the targeted muscle intraoperatively. An additional 2 percent of the electrodes did not have an isolated response (second muscle recruited within 2 µs of the first muscle) and another 2 percent had thresholds higher than 50 s. Therefore, 94.5 percent of the electrodes met our criteria for correlation between the targeted response in surgery and the response achieved after surgery. Many of the electrodes with a less than a desirable response were those placed in weak, partially denervated muscles.

Three epimysial electrodes have been moved as part of a subsequent surgical procedure. In two cases, the surgeon repositioned the electrode on the muscle in an attempt to obtain better recruitment properties. One electrode was placed on the extensor digitorum communis (EDC) and was moved after 2 years. The difference in response was minimal. The second electrode was placed on the flexor digitorum profundus (FDP) and was moved after 10 months. The surgeon moved a third electrode, which was originally implanted on the pronator quadratus muscle, to the ulnar nerve near the flexor carpi ulnaris tendon to obtain a new function. In all cases, after the electrode was moved, the patient was recasted for 3 weeks. No other incidents have occurred with these electrodes (at least 6 years of follow-up).

Functional Stability

Stability of physiological responses obtained from the electrodes was also excellent. Of all 238 electrodes, the average pulse width for threshold was 15 µs (range of 2 µs to 82 µs). The threshold depended on the implanted muscle, as shown in Table 2. Larger muscles, such as the FDP and triceps, tended to have higher thresholds.

Table 2. Electrode thresholds for upper-limb muscles.
Target Muscle or Nerve	n	Average Threshold (µs)	Standard Deviation (µs)
Ulnar Nerve	7	4	2
Third Dorsal Interosseous	5	10	4
Adductor Pollicis	24	10	10
Extensor Pollicis Longus	27	11	6
Abductor Pollicis Brevis	28	12	9
Extensor Carpi Ulnaris	9	14	8
Extensor Digitorum Communis	24	14	11
Second Dorsal Interosseous	6	16	12
Flexor Pollicis Longus	20	17	15
Flexor Digitorum Profundus	22	21	21
Triceps	9	21	22
Flexor Digitorum Superficialis	29	25	18
n = number of electrodes implanted

A more extensive analysis was performed on 81 electrodes from 11 subjects that had at least four recorded threshold values spanning at least 2 years, with the first data point being at least 60 days postimplantation. We used a linear fit to identify whether the thresholds showed a trend that was significantly different than zero. Seventy-six of these eighty-one electrodes (93.8%) had slopes that were not significantly different from zero. Of the remaining five electrodes, all had an estimated slope of less than 0.8 µs a year, which was well below the rate that was considered clinically significant. Therefore, 100 percent of the electrodes demonstrated long-term stability in their threshold responses.

Thresholds for the first group of electrodes implanted have been recorded regularly over the course of 16 years [1]. These data are shown in Figure 4. There is session-to-session variability in the observed threshold, but the estimated slope for each electrode is less than 0.74 µs a year.

Description: Figure 4. Epimysial electrode threshold measurements for subjects.

Surface potential mapping was recorded regularly in the first three subjects, but the response was determined to be quite predictable and therefore only needed to be measured if a suspected lead fractured. An example of a typical surface potential map is shown in Figure 5.

Description: Figure 5. Typical surface potential map.

Mechanical Durability

Across all 238 electrodes, only three electrode-lead mechanical failures occurred. Each of these incidents occurred within 2 years postimplant. We have identified that repeated use is the most likely cause of failure in only one of the three breakages. No failures of the lead occurred more than a few centimeters proximal to the electrode tip, and no failures or separations of the connectors occurred. None of the mechanical failures appear to have been the result of the lead being flexed across the joint, because no failures occurred near the joints.

Failed Electrode-Case 1

In subject 2K, the epimysial electrode used for sensory feedback broke secondary to the implant device rotating within the body. The rotation of the device caused the sensory lead to become wound around the remaining leads. The lead eventually pulled apart in tension and broke near the electrode termination. At the time of repair, we determined that the implant had made 16 complete revolutions within the body. The implant was unwound and we determined that all the leads from the implant device were intact and had not been damaged. The distal lead of the sensory electrode was disconnected at the connector site and a new sensory electrode implanted in the same location. This device has now been operational for 10 years without further incident. Because of this early incident, the surgeon now sutures the implant device in place to prevent rotation or migration within the body, and no additional incidents have occurred.

Failed Electrode-Case 2

In subject 1W, an epimysial electrode placed on the abductor pollicis brevis (AbPB) muscle broke after 2 years of regular use. The break was identified through routine electrode threshold measurements and verified with surface potential measurements. The surface potentials for the AbPB electrode, as compared to the nearby AdP electrode, are shown in Figure 6. The surface potentials indicated that the breakage did not include the lead insulation and/or that the break was near the electrode tip. The electrode was surgically removed and examined. The break was in the lead wire just proximal to the electrode disk. We suspect that this fracture occurred from repeated pressure on the thenar eminence. This patient pushed a manual wheelchair and used the palm of his hand against the rims of the wheelchair to accomplish propulsion, undoubtedly placing repeated high stresses directly on the electrode. Because of this incident, we now place the AbPB electrode in a more medial location so that it is more protected from pressure to the palm.

Description: Figure 6. Surface potential map for failed abductor pollicis brevis

Failed Electrode-Case 3

In subject 2C, an intramuscular electrode placed in the second dorsal interosseous muscle developed a high threshold and an altered response. This electrode had functioned normally for 442 days. The electrode exhibited an elevated threshold, and the surface potential measurements indicated that the resistance of the lead/electrode or electrode/tissue interface had increased. We determined the latter by examining the stimulus pulse as recorded from the skin surface, as shown in Figure 7. Under normal conditions, the constant-current stimulus pulse delivered by the implant produces a falling edge that gradually reaches a plateau after a few microseconds. When the electrode impedance exceeds the capacity of the stimulator to maintain a constant current pulse, the recorded waveform changes. The falling edge exhibits an overshoot before reaching a plateau. The increased threshold and impedance indicated a broken electrode. The electrode was removed during an unrelated surgical procedure and replaced with a new electrode in the same muscle. Examination of the lead uncovered a region of fracture that was approximately 10 cm proximal to the tip of the electrode. The silastic insulation was intact.

Description: Figure 7. Recording surface potential for a single stimulus pulse

We could not determine the cause of the fracture. The subject underwent a surgical procedure near the location of the lead fracture just before the electrode exhibited the altered response. In addition, the subject had accidentally hit his hand on a metal door just before the altered response, causing bruising. The fracture mechanism is consistent with either cause. Based on our analysis, it is unlikely that the fracture occurred secondary to flexion fatigue.

Tissue Response

In a few cases, when an electrode has been replaced or moved, there has been the opportunity to observe the encapsulation response of the body to the electrodes. The encapsulation is similar to what we have observed in the dog model [10], consisting of a 0.5 to 1.0 mm-thick transculescent encapsulation that securely holds the electrodes in place. The sutures are still visible and tied.

Infection/Rejection

A single incident of a localized infection at an electrode occurred. In subject 1N, a localized infection developed at the site of the sensory electrode termination near the mid-clavicle. The infection appears to have been initiated by a suture in the incision site. To ensure that the infection would not track up the electrode lead, we cut the distal lead proximal to the infection site and removed the distal end of the electrode. The infection resolved without further incident. The sensory electrode was not replaced in this subject.

DISCUSSION

This study presents data demonstrating the long-term viability of implanted electrodes, leads, and connectors for use in upper-limb neuroprostheses. Electrode and lead failures or infections have been anticipated to be significant causes of failure in motor neuroprostheses for limb control [15-21], but our results demonstrate that the total failure incidence for any reason was less than 2 percent. This incidence can be reduced by careful electrode placement. The results further indicate that the electrodes remain in the location in which they are placed and that the tissue response consists of a thin encapsulation that is formed within the first 1 to 3 months postimplant and is entirely stable thereafter.

Muscle-based electrodes provide an excellent method of delivering electrical stimulation to paralyzed muscles. These electrodes provide selective stimulation, deliver minimal trauma to the underlying tissue, are easily implanted, and are easily repositioned. There are several situations in which the neuromuscular anatomy dictates the need for muscle-based electrodes. For example, the extensor pollicis longus (EPL), which is the key muscle for providing thumb extension, must be activated strongly and in isolation from the nearby finger extensors. Isolating activation of the EPL from the more superficial EDC and extensor indicus proprious (EIP) is virtually impossible with the use of surface electrodes [22]. The multiple branching of the distal portion of the radial nerve precludes the distal placement of a nerve-cuff style of electrode isolated to the EPL motor branch. Therefore, the only existing method to reliably activate the EPL in isolation is to use a muscle-based electrode. This illustrates why we expect that muscle-based electrodes will be a necessary part of motor neuroprostheses for the near future.

Surgical installation of implanted electrodes does not present any unusual technical difficulties, especially in the upper limbs where all muscles are relatively easy to access and expose. The use of an epimysial-style mapping probe allows the surgeon to move the electrode over the entire surface of the muscle while observing the stimulated response in real-time. Once the optimum location has been identified, the electrode can be temporarily sutured in place and the response tested again. Evaluation of the electrode response should be performed after the overlying muscles and skin are returned to their natural position. Then by varying the stimulus levels, the physician should perform an intra-operative evaluation of the electrode-muscle output characteristics. Important electrode-muscle output characteristics include the change in muscle force as a function of the muscle length and the change in muscle force output as a function of the stimulus intensity [11]. When these principles are followed, our results show that the postoperative response corresponds to the intraoperative response in 95 percent of the electrodes. In addition, the average threshold for the electrodes in this series was 15 µs which is considerably lower than the average threshold of 33 µs obtained for percutaneous electrodes in the same group of muscles [12], indicating that the fully implanted electrodes can be placed more accurately.

Three electrodes were moved during revision procedures to gain improved performance of the electrode. All of these cases occurred during the first 13 subjects in this study. Our experience has been that we have not been able to find a significantly better electrode response in the attempted revision procedures, indicating that the initial placement was probably already close to optimal.

The long-term physiological response of the body to the presence of the electrode materials and to the stimulus current was an important concern in early studies [1,6,10]. The results of this study further confirm that the presence of the electrode materials (Pt-Ir, 316 SS, silicone) are well tolerated by the body and do not pose a source of long-term irritation, even when placed on (or in) contracting skeletal muscles. We found no evidence that daily delivery of electrical current through these electrodes at the stimulus levels used in this study (20 m, 200 µs) electrode surface area: 10 mm²) caused any adverse response. No evidence was found of any progressive muscle weakness as a result of the chronic daily stimulation.

Although histology has not been performed on humans because of the risk of losing viable muscle tissue in patients who are already weak, we have had opportunity to grossly observe the tissue response to the implanted components during revision surgery. In three cases, the implant device itself was exposed after implantation periods of 6 months, 2 years, and 9 years. In all cases, we found that a well-formed capsule with a thin wall surrounded the implant components. No cases indicated any infection, inflammation, or ongoing encapsulation.

The absence of long-term drift in the thresholds for the stimulating electrodes indicates that substantial encapsulation does not continue beyond the first few weeks after implant. Electrode thresholds have been consistent regardless of patient activity or stimulation time. This result suggests that the stimulation itself has no effect on the electrode-tissue interface.

There is a clinical concern that the presence of any foreign material within the body might be a potential site for infection. The presence of long leads running the length of the arm to the chest might also present a path for migration of infection. In this series of patients, only a single localized infection of one electrode was identified, and this infection was probably due to the suture in the wound above the electrode. The infection did not track along the electrode lead, but the lead was removed as a precaution. Other systemic infections have not tracked to the implantable components in this series of patients. Patients have reported unrelated infections, such as pneumonia, urinary tract infections, and even cellulitis, without any adverse affect from the implanted components. Although patients were instructed to take antibiotics as a precaution with any infection, compliance was difficult to gage.

The incidence of mechanical failures with the electrodes, leads, and connectors was extremely small in this study. Only three failures were encountered, and most likely, only one of these three failures was an actual failure caused by repeated lead flexure. Including all leads in our analysis, the 98.7 percent survival for 16 years is excellent. This compares favorably to the reported survival rates of pacemaker, defibrillator, and spinal cord stimulator leads [23-28]. The connectors had no failures.

In the single subject where the lead appears to have undergone fatigue failure (subject 1W), extenuating circumstances were found. The electrode was placed on the surface of the AbPB so that it was quite superficial. This electrode was easily palpated and was located in a region that was subject to high stress. People with tetraplegia frequently place substantial body weight on their hands for weight shifts and transfers and stress on their arms and shoulder for wheelchair propulsion. This particular subject also pushed his own wheelchair and was otherwise quite active. All of these activities are likely to put stress on the thenar eminence located directly under the electrode. Because of this experience, we have now modified the placement of the electrode so that it is more medial, deeper in the tissue and is therefore more protected from external stresses. We have had no further incidents of this type in the 7 years since this issue was first identified.

All three of the lead failures and the single electrode infection occurred within 2 years of implantation. Currently, 191 electrodes have been functioning longer than 5 years, with no failures or infections within this group. The trend of earlier failures is similar to that reported for other implantable leads [24,27,28]. Within the time frame of the present study (15 years), failure is not positively correlated to total time implanted. In addition, we have not had a lead failure occur in this study for the past 3 years, which may indicate that we have identified and rectified some of the possible failure mechanisms.

Even though electrode failure is an infrequent occurrence, identifying these failures quickly and confidently is important so that they can be corrected. We have developed a complete battery of methods that can be used to identify and analyze a possible failed electrode prior to surgical replacement, as described in this paper. In all three incidents where an electrode failed, the subject reported a different response or sensation from the neuroprosthesis. In one case, after the subject described changes in the location of the stimulus sensation, an X ray was used to reveal the problem. However, our experience is that only gross failures can be identified by X ray. In the other two cases, electrode threshold measurements, when compared to previously recorded values, indicated an unusually elevated threshold. The use of surface potential recordings allowed further confirmation and localization of the broken lead. To determine if the electrode impedance has increased above the expected range, one can use the shape of the recorded waveform. Theoretically, surface potential measurements can be used to identify a proximal current leakage through the lead insulation, but we have yet to confirm this in a human because the silastic insulation has produced a sufficient seal to prevent measurable current leakage at the break site. In our experience, yearly electrode thresholds provide the most straightforward means of identifying a failed or failing electrode.

CONCLUSION

In 27 patients who received an upper-limb neuroprosthesis, 238 electrodes have been studied. There is at least 3 years follow-up on all electrodes, with a maximum follow-up time of 16 years. Three lead failures occurred and one localized infection occurred during the study, resulting in an overall failure rate of less than 2 percent. The expected lifetime of the electrodes is over 98 percent at 16 years. The device-tissue interface consists of minimal encapsulation that is stable over time. The results indicate that complications of a neuroprosthesis caused by device failure, electrode breakage, lead breakage, infection, or rejection are extremely low and are not a source of major concern. These results have an impact on the design of implantable neuroprosthetic systems. The electrode and lead component of these systems should no longer be considered a weak technological link.

R a n d a l l M o r t o n BBus BComp BPsych