International Functional Electrical Stimulation Society (IFESS): the Development of Controlled Neural Prostheses for Functional Restoration


and reprint requests to Ross Davis, MD, Neural Engineering Clinic. Clinical Neuroscience Center, 76 Eastern Avenue, Augusta, ME 04330. Email:

In 1998, the first combined INS-IFESS meeting occurred in Lucerne, Switzerland. Clinicians met with biomedical engineers, physiologists, allied health therapists and manufacturers who all became aware of each other's expertise. Members of one Society then realized that the other Society's members could be most important complementary partners in their future endeavors to create and use new controlled neural prostheses for the restoration of function. Let us explore some of the evolving fields now undertaken by members of IFESS. It is our aim to publish reviews and original papers by IFESS experts in this and following journal issues of Neuromodulation.

Functional electrical stimulation (FES) in broad terms, is the external control of stimulation to nerves to restore the original functioning of the target muscle(s), bladder, hearing, vision, breathing, etc. This history started in 1875 when the Medical Times ( 1) reported that G. Duchenne applied faradizing current to the lower extremities of a paraplegic subject. “Who does not recollect the astonishment exhibited in the clinic at that experiment of Duchenne of drawing from his bed a patient regarded as absolutely paraplegic, and loading him with the weight of a man of ordinary size, without him ever flinching under it?”

FES as we know it today, was initiated in 1961 by W. Liberson( 2), who attached surface electrode pads to the skin overlying the peroneal nerve of the affected leg in a hemiplegic patient. Synchronized electrical stimulation through the pad caused dorsiflexion of the ankle, so improving the swing phase of the gait. Peroneal nerve stimulators have since been applied to 5000 “foot-drop” patients.

In this issue of Neuromodulation, T. Sinkjaer has written a review( 3) on the use of chronic neural recording cuffs for sensory feedback in closed-loop controlled FES systems in humans. The controls of FES stimulators are categorized as open-loop, where there is no electronic feedback to the stimulator. Closed-loop control involves electronic feedback to the stimulator from sensors such as heel switches, position and pressure sensors, nerve cuff recordings, and electromicrograms (EMG). Hierarchical controls can be superimposed on the close-loop mechanism. Adaptive controls and learned controls by artificial neural networks are adding more precision and stability.

Following cardiac pacemaker technology, the first implanted electrical stimulators for the nervous system began in 1963 when Glenn et al.( 4) implanted a radio-frequency (RF) linked stimulator to rhythmically activate the phrenic nerve for long-term artificial respiration, particularly in high cervical spinal cord injured individuals. Several commercial respiratory assist systems are currently available and were reviewed by Creasey et al.( 5) in 1996. Research is currently investigating stimulation of the intercostal and/or abdominal muscles to augment respiration in individuals with a weak diaphragm as a method for providing cough. Others are investigating the use of intramuscular electrodes in the diaphragm as a less invasive and safe alternative to phrenic nerve cuffs.

There are considerable results in using FES for the restoration of function for those with spinal cord injury (SCI), particularly in the upper extremity of tetraplegia. Since 1977, FES has been used in tetraplegic individuals to provide functional hand grasp and release and was reviewed in 1996 by Triola et al.( 6). In 1987, an 8-channel system was first implanted and eliminated the need for percutaneous leads. This multiplex stimulator-receiver is surgically implanted with leads tunneled subcutaneously to the muscles used to control finger and thumb motion. Seven epimysial electrodes are placed on the muscles where their nerves enter (motor points). The eighth electrode is placed in the supraclavicular region and is used for sensory feedback. An RF inductive link provides the communication and power to the implant receiver-stimulator. The external portable control unit is battery-powered and programmed with the appropriate control and stimulus parameters which relate the opposite shoulder position to the stimulus output applied to each muscle. These systems have been successfully implemented in over 130 patients and have provided them with the ability to perform many activities of daily living. This implantable “Free Hand” system (NeuroControl Corporation, Cleveland, OH) was approved by the US Food & Drug Administration (FDA) for use in 1998. Further development of this FES system is planned with an increase in the number of stimulating channels to 10–12 as well as an implantable wrist angle sensor.

A review by Davis( 7) in 1997 described the use of surface-applied and implanted FES systems for functional restoration in the lower extremities for SCI individuals. Typically, the FES systems use stimulation of the conditioned quadriceps muscles to achieve standing. This is done in combination with voluntary arm action; and often for prolonged standing, an ankle-foot stabilizing orthosis is used (± knee brace). Locomotion has been produced by stimulating the peroneal nerve on one leg to produce a flexor withdrawl response and by quadriceps stimulation on the other leg. This combination is reversed with the next step. The forward propulsion is achieved with a combination of lower extremity movements and voluntary weight transfers through the upper extremities using either crutches or a walker. Hand switches allow individuals to adjust their own ambulation speed. The FDA has approved the Parastep FES surface-applied electrode system (Sigmedics, Inc., Northfield, IL) for standing and walking in SCI individuals.

Skin surface stimulation has a lack of specificity plus the inability to penetrate to deeper muscles in paraplegic subjects. It can also produce multisynaptic and multisegmental flexor responses resulting in multilevel spastic reflexes. These problems have been overcome by implanting stimulating electrodes directly onto peripheral motor nerves and nerve roots before (epineural) or at the entry point on the target muscle (epimyseal), and using deeply placed wire electrodes at or near muscle motor points (intramuscular). From 1970 on, implanted electrodes have been used to achieve standing, stepping, and bicycling. Presently, there are 3 implantable FES systems with 12–22 stimulus outputs reported with ± orthoses for use in the clinic setting.

The majority of SCI patients develop urological problems mainly due to bladder dysfunction with the whole of the urinary tract permanently in danger. In 1998, the FDA approved the extradural anterior sacral root FES stimulator (Vocare, NeuroControl Inc., Cleveland, OH). This device evolved from Brindley's work in 1976. With his first stimulator, the electrodes placed on intradural sacral roots; some 1500 have since been implanted( 8–10). At implant surgery, posterior sacral root rhizotomies at the conus level are done completely in SCI patients to allow an increase in bladder volume without uninhibited contractions and with a reduction/abolition of autonomic dysreflexia. A major drawback of anterior sacral root stimulation is the simultaneous contraction of striated urethral sphincter muscle and the smooth detrusor muscle. This results because the ventral sacral root consists of large-diameter myelinated somatic fibers innervating the urethral sphincter (via pudendal nerves), and small-diameter myelinated preganlionic parasympathetic fibers innervating the detrusor muscle (via pelvic nerves). Hence, anterior sacral root stimulation leads to detrusor sphincter dyssynergia with the risk of elevated bladder pressure and incomplete bladder emptying. With the use of timed bursts of stimulation, the bladder detrusor continues to contract as the sphincter intermittently relaxes. Benefits include improvements in continence, with absolute continence often achieved, and increased bladder capacity and compliance. Abolition or diminution in upper urinary tract dilatation with improvement in renal function was observed with the abolition of high-pressure ureteric reflux. Lower residual urinary volume occurred with fewer urinary tract infections.

FES is used to control heterotopically transferred skeletal muscles( 7), as in latissimus dorsi muscle transfer for cardiac assistance (cardiomyoplasty), gracilis muscle transfer for dynamic anal myoplasty for fecal incontinence, and dynamic gracilis myoplasty for bladder incontinence.

For obstructive sleep apnea (OSA), hypoglossal (HG) nerve stimulation has produced tongue protrusion by genioglossal recruitment in 15 cases as reported in 1997( 11). Direct HG nerve stimulation below the arousal threshold can improve airflow in patients with OSA. The future development of a closed-loop system appears essential for the success of the HG stimulator for OSA. A sensor would detect the inspiratory phase or the lack of it (with possibly an added respiratory gas sensor) and feed this information back to the implanted stimulator.

In the hearing impaired, cochlear prostheses( 12) have been a major FES success with over 30,000 implanted. These are used for sensory transduction losses from damage to, loss of, or absence of the inner hair cells of the organ of Corti. The cochlear implant restores hearing by means of direct electrical stimulation of the neurons that convey auditory information from the ear to higher processing centers. To identify most human speech sounds, it is necessary for the listener to receive information simultaneously about the relative intensities of sound in several different frequency bands. The acoustic information is detected by a microphone and filtered electronically to determine the relative energy in each of 4–20 frequency bands. This information is used by a microprocessor to compute the appropriate electrical stimulation applied to each of 4–20 sites of stimulation in the cochlear (or in the brain stem when the areas bilateral loss of the auditory nerves: auditory brain stem stimulation). In 1981, the first commercial multichannel prosthesis to obtain premarket approval from the FDA was built by Cochlear Ltd. (Lane Cove, New South Wales, Australia) which has 22 ring-shaped electrode contacts threaded into the scala tympani. There is at least three other multichannel cochlear prostheses.

There have been searches for vision prostheses starting from 1968. In 1997, Girvin and Fodstad( 13) and Davis( 7) reviewed the investigations that placed stimulating electrodes on the occipital lobe surface (visual cortex) to develop a prosthesis for the blind that might be of value in reading and mobility. With this electrical stimulation, human subjects were able to see circumscribed and often punctate sensations of light, called phosphenes. The problems with surface stimulation include high currents for eliciting phosphenes, interactions between phosphenes generated by simultaneously stimulated electrodes, and occasional persistence of phosphenes following cessation of stimulation.

From 1990 on, investigators from and those funded by the Neural Prosthesis Program at the National Institutes of Health began the systematic design, development, and evaluation of safe and effective means of microstimulating cortical tissue. The expectation was that by implanting microelectrodes with exposed tip sizes the same order of magnitude as the excited neurons, much more selective stimulation could be achieved, resulting in more precise control of neuroneal function. For a visual prosthesis, it was hoped that this would result in reduced phosphene interaction and higher information transfer rates into the visual system. They did validate these findings in three awake and sighted humans who had volunteered for visual cortex stimulation while undergoing excision of epileptic foci under local anesthesia. They found that electrical current thresholds for producing phosphenes by intracortical microstimulation were 10–100 times lower than those produced by stimulation with nonpenetrating cortical surface electrodes, that phosphenes could be resolved with simultaneously stimulated electrodes as close as 700 μm, and that the phosphenes had simple forms that were stable and predictable. These results were encouraging for the possibility of a visual prosthesis based on intracortical microstimulation.

In 1992, the NIH group reported on a fourth subject, totally blind for 22 years, who was similarly implanted and stimulated with intracortical microelectrodes. The microelectrode had exposed tip sizes approximating the size of cortical neurons that were used to activate small populations of neurons. The subject could recognize simple patterns with the limited number of electrodes available. Phosphenes were produced at currents as much as 1000 times lower than required with electrodes placed on the cortical surface. Within rather narrow limits the brightness in size and color of the visual sensations could be varied with changes in the stimulus level. Phosphenes produced by simultaneous stimulation through two electrodes 500 μm apart could be resolved but not when electrodes were spaced at 250 μm. These results are compatible with the possibility of a visual cortex prosthesis for the blind but complex image recognition experiments must await future implants with larger arrays of electrodes. Such a prosthesis would utilize a miniature television camera and electronic processor to convert the camera output to a signal suitable for controlling a multichannel (up to 1024) stimulator with its array of visual intracortical microelectrodes.

The most common form of blindness in the western world is age-related macular degeneration. This condition affects 10 million Americans, at least 10% of whom are legally blind. The most common form of inherited blindness in the world is retinitis pigmentosa that affects 1.2 million people. Both of these conditions share the same problem of damaged rods and cones, while other cells, particularly the ganglion cells that connect the eye to the brain remain healthy. It is the aim of several groups( 7) to restore vision to these patients by making use of this healthy connection. Their goal has started with the development of implantable, intraocular, microchip stimulators. These chips are now being used in animals, and either rest on or below the retinal surface of the ganglion cell bodies so that controlled electrical stimulation will hopefully create visual perception. Prosthetic stimulation at the cortical level has several drawbacks, first the prosthesis is placed underneath the skull and second, only a limited portion of the striate cortex is accessible to stimulation electrodes. In 1998, an optic nerve prosthesis was developed and intracranially implanted by Veraart et al.( 14) in a blind volunteer with retinitis pigmentosa. The self-sizing spiral cuff nerve electrode around the right optic nerve had four contacts. Flickering was studied using repeated identical stimulation at various repetition frequencies and maintained up to 3.6 s. Flicker was usual below some 4–5 Hz. Around 5–10 Hz, flicker fusion occurred. When frequency increased, around 16–64 Hz, the mean perceived intensity usually tended to increase. Furthermore, at flicker fusion, the perception threshold diminished. When fusion occurred, perception usually disappeared after some 1–3 s, which corresponds to the described perception durations for stabilized images.

A further development in microelectronics has lead to the MicroStimulator, which was developed at the Alfred E. Mann Foundation (Santa Clarita, CA). This was in conjunction with the Queens University Bio-Medical Engineering Unit (Kingston, ON, Canada) and the Illinois Institute of Technology/Pritzker Institute (Chicago, IL), plus funding by the NIH Neural Prosthesis Program( 15). The implanted device is approximately 2 mm in diameter and 10 mm in length, and can be injected into muscle or adjacent to a nerve through a 12-gauge “intracath” needle. Inside this glass cylindrical capsule is a microcoil and a custom integrated chip. An external control system, connected to an overlying coil antenna, delivers both RF power and data for the implanted MicroStimulator(s) that deliver electrical stimulation as current-regulated pulses.

A microprocessor-based control system suitable for miniaturization as a wearable controller has been built and used successfully in animal experiments at Queens University. This device is undergoing its first clinical use in preventing or reducing shoulder subluxation, pain, and dysfunction in hemiparesis due to stroke. Multiple MicroStimulators have been implanted for use as a multichannel FES system with a single external coil and transmitter. This research group is also developing telemetry capabilities in this module to monitor bioelectric and mechanical events so enabling the controller programmer to have a closed-loop system.


The past 25–30 years has seen the evolution of chronic electrical stimulation of the nervous system that evolved from cardiac pacemaker technology. The quest for new targets has expanded, especially in the efforts to restore function in those with spinal cord injuries, stroke, blindness, and deafness. It has required teamwork of a very high order among the clinicians, biomedical engineers, physiologists, and manufacturers to achieve the various implantable neural prostheses with safer electrodes/arrays and efficient stimulation parameters. The development of implantable sensors and/or the ability to record information from sensory nerves linked with these new multichannel stimulators is and will produce closed-loop computer controlled interactive systems for safer and more efficient restoration of function.