Functional Electric Stimulation
for sensory and motor functions: Progress and Problems

Dimiter Prodanov1,2, Enrico Marani1,2, Jan
Holsheimer1

1 Biomedical Signals
& Systems Group, Faculty of Electrical Engineering, Twente University,
Enschede, The Netherlands

2Neuroregulation Group,
Department of Neurosurgery, Leiden University Medical Center, Leiden, The
Netherlands

Address for correspondence:
E. Marani,
Neuroregulation Group, PO Box 9604, 2300 RC Leiden, The Netherlands, e-mail: [email-masked]

Published in Biomedical Reviews, 2003, 14:23-50.

Copyright - pre-acceptance version © 2003 by the authors. However,
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To avoid copyright issues, here follows the pre-acceptance version
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Abstract

Functional Electrical Stimulation (FES)
tries to restore the lost functions of the nervous system by means of electrical
stimulation. FES is recapitulated in this review.  An overview is given of
the relevant moments in history. The principal notions behind FES are outlined,
in which special attention is dedicated to the electrodes for selective
stimulation. FES applications encounter specific research problems, which are
extensively covered. The major clinical applications are overviewed: pacing of
the heart, respiratory pacing of the diaphragm, the rehabilitation and
restoration of locomotion and hand functions, and the restoration of bladder and
bowel functions. An analysis of the current applications based on the
evidence-based clinical approach is performed.

FES for restoration of the lost
locomotor functions is a rapidly developing area that is on the verge of
broadening its applications. Evidence for this is the increasing number of
various neuroprosthetic devices that have been developed recently. Some of them
are already beyond the prototype stage. The present FES treatments combined with
conventional occupational and physical therapy still remain the most promising
approach in rehabilitating Spinal Cord Injury patients and stroke patients. FES
in combination with neuromodulation becomes a viable option in the treatment of
various urological disorders, however it is still in a rather empirical state of
operation.

The
presented overview of the FES applications shows that despite the substantial
amount of research performed in the field which led to successful development of
cardiac and phrenic pacemakers and cochlear prostheses, a number of fundamental
scientific problem still remains to be solved before one will see a comparable
degree of effectiveness and penetration in common medical practice for the
locomotion neuroprostheses and urological FES appliances. The major up-to-date
directions of research in clinical and theoretical aspects are summarized.

History
of electrical stimulation
<Figure
1 Galvani>
<Figure
2 Friction machine>

Principal
notions in FES
Electrochemical
processes at the tissue-electrode interface
Voltage-controlled
and current-controlled stimulation
Monopolar,
bipolar and tripolar stimulation
Cable
properties of the nerve fibers
Active
properties of the nerve fibers
<Table
1>
<Figure
3 McNeal Model>

Stimulation
Electrodes
Surface
electrodes
Intramuscular
and epimysial electrodes
Cuff
electrodes
<Figure
4 cuff electrode >
<Figure
5 3D MEA>
Intraneural
electrodes
Intraneural
wire electrodes
Intraneural
multi-electrode arrays (MEAs)

Research
Problems
Biocompatibility
of the implantable devices: general considerations
Chemical
biocompatibility
Mechanical
biocompatibility
Implantation
and the Nerve damage
Tissue
reaction to implantation
Nervous
Tissue reaction to electrical stimulation
Selectivity
concepts
Surface
stimulation vs. Implantation

Contemporary
Clinical applications of FES
The
cochlear implant 20 Drop-foot
stimulator
<Figure
6 Peroneal Drop-foot stimulator>
Restoration
of Lower limb Function
Hand
disabilities – reaching and grasping
Phrenic
stimulation

Clinical
trials / Evidence based medicine for FES
Efficacy
of FES after stroke
Efficacy
of FES in urological conditions
Detrusor
sphincter dysynergia
Overactive
bladder
Experimental
and theoretical research
Experimental
research
Visual
prosthesis: blind eye vs. normal brain
Spinal
Cord Stimulation in motor disorders

Theoretical
Modeling studies
Research
on spatially-selective nerve stimulation
<Figure
7>
Research
on fiber size selective stimulation

Conclusions

References

History of
electrical stimulation

Functional Electrical Stimulation
(FES) is a part of the broader field of electrotherapy. The term was introduced
relatively late in the history of FES by Moe and Post (1)
. In general, history of electrical stimulation follows the theoretical and
technological breakthroughs in electricity and magnetism and it is based on the
developments in Neuroanatomy and Neurophysiology, since electrical stimulation
is mainly an application of electro-magnetic theory in the nervous system.
Nevertheless, the history of electrical stimulation has its own particular
subjects of which the major ones, related to the nervous system, are given in
this review.

In the field of Neuroanatomy the
application of ethanol to fix brain tissue (Felix Vicq d’Azyr, 1748-1794) opened
the microscopic approach for the study of the brain, ending up with the
detection of the sub cellular protrusions of the neuron and its supporting
cells. One should notice that the detection of the micro-architecture of
peripheral nerves (Van Leeuwenhoek, 1719) was remarkably accurate but forgotten
for nearly a century. The descriptions of Dutrochet in 1824 (2) and
Remak in 1838 (3) once
again depicted the tubular structure of the entities in the nerve, together with
the detection of myelinated and non-myelinated fibers.

Albrecht Haller (1708-1777)
described, the “irritation phenomenon of muscles”: in fact the contraction and
relaxation of a muscle. In his notion, it was a ”force” that passed along the
nerves that could induce this excitation of the muscles. In addition, Haller
made the first concept on sensibility. He found that tissues by themselves lack
sensations, but they were noticed and relayed by nerves and their
endings.

Luigi Galvani (1737-1798), who
developed the so-called “galvanic” element, by means of which nerves could be
stimulated, studied the muscle contractions. Later on in 1792, Allesandro Volta
(1745-1827) could demonstrate that the galvanic element used electricity. So,
the nerves were since considered to be the electric cabling of the body.

Figure 1. Galvanic experiments on human muscles. From G.
Aldini, Essai théorique en experimental sur le Galvanisme, avec série
d’experiences… Paris, Fournier fils, 1804 Courtesy: Prof. H. Beukers, Leiden
University

In the mean
time, friction electricity was in use and the first capacitor, “de Leidse fles”
(1746), was developed by Muschenbroeck. From 1784 till 1791 the biggest
friction-electrical machine was build by John Cuthbertson. Two men had to move
two enormous glass discs. Each disc was frictioned by four cushions. The
electricity on the glass discs had to be induced into the conductors, a series
of Leiden jars. The potential difference reached nearly 330 kV. Van Marum
(1750-1837), biologist, physician and botanist used this machine to do
electrophysiological research for ten years. This shows that around the
beginning of the XIX century systematic application of electricity in research
and medicine had already started.

Figure 2. The enormous friction electrical machine
together with the Leiden jars built by J. Cuthbertson. Courtesy: Teylers
Museum

By the middle of
the XIX century with the discovery of static electricity, capacitance and
electromagnetic inductance, physiologists could study muscle contractions, nerve
conduction and cell excitation.

In the
nineteenth century J. Müller (1801-1858) introduced the law of Specific Nerve
Energies: “the kind of sensation, following the stimulation of a sensory nerve,
depends not on the mode of stimulation but on the nature of the sense organ with
which the nerve is linked”. His own disciple Helmholtz (1821-1894) measured the
velocity of the nervous impulse and Emil DuBois-Reymond (1818-1896), also a
disciple of Müller, proved the existence of the resting membrane potential in
the nerve fibers and also proved that each impulse is directly related to a
change in potentials.

Sir Charles Bell
(1774-1842) and Francois Magendie (1783-1855) established the neuronal basis for
the reflex. The dorsal root contains afferent, sensory information, while the
ventral root relays the motor function.

The final idea
that the central nervous system is the controller of the organism was brought
forward by T. H. Huxley (1825-1895) in his Manual of the Anatomy of the
Invertebrate Animals (1817) and became generally accepted. This concept
became the fundament for the effective application of electrical stimulation. In
this period medical applications were numerous but short lived and poorly
understood. Electricity was reported to cure rheumatism, neuralgia, insomnia and
even cold feet (4)
.

In XX
century, the major technological advances in electronics and computing
(discovery of the diode, triode, transistor and integrated circuits) lead to the
development of contemporary FES appliances. The main push in the applications of
FES was the successful development and clinical applications of the heart
pacemakers.

Literature shows controversy over
who was the first inventor of the artificial heart pacemaker (5) . In
1932 Dr. Albert Hyman described an “artificial pacemaker” that he had used with
experimental animals to resuscitate the “arrested” heart. About that time, the
Australian physician Mark C. Lidwill, along with the physicist Major Edgar
Booth, built a portable pacemaking unit. It was demonstrated in 1931, while
Hyman developed his device in 1930-31. Lidwill's apparatus had one pole applied
to the skin and another into the appropriate cardiac chamber (6) .
Later on, the Hyman device was successfully used in patients

The application of cardiac
pacemakers to stimulate the human heart dates as far back as 1954. However, it
was not until the 60s when pacemakers were adopted on a great scale in clinical
practice. Their potential for different applications in the beginning was not
recognized both by the medical profession and by the industry. The other reason
was the immaturity of the technology and the implantation techniques.
Stimulation was carried out by skin electrodes, which left uncomfortable burns
when used for a couple of days or more. Later, electrode wires leading through
the skin were tried, but infections along the wires were an unsolvable problem
(6) .
Its solution was the implantation of the entire pacemaker. Dr. Rune Elmqvist
designed the world's first implantable pacemaker. The first 10 patients for the
implantable pacemaker were operated in 1960 (7) . In
the 1970’s, heart pacemakers demonstrated the reliability of implanted
electronic devices and showed the benefits in terms of improved quality of
life.

In parallel went the attempts to
apply stimulation of the diaphragm for respiratory pacing. For some 200
years, electricity has been applied to the phrenic nerves to activate the
diaphragm. However, the actual therapeutic opportunity arose in the
1950s, at the same time the heart pacemakers were developed. The first attempts
to pace the diaphragm with implanted electrodes were carried out between 1948
and 1950 by Sarnoff and co-workers (8) . One
of the most important prerequisites of the routine clinical applications was the
introduction of the long-term electrical stimulation by the radio frequency
inductive method introduced around the end of the 1950s (9)
. In 1966, Dr William Glenn at Yale University performed the first clinical
application to patients with hypoventilation of central origin. The ventilation
support was made by radio frequency pacing of the diaphragm. Pacing for total
ventilatory support in patients with respiratory paralysis was applied in 1971.
In the early 1980’s the first commercial phrenic nerve pacers were
introduced.

As a result of the successful
application of the pacing of the heart, the technological advances achieved
during the development of the pacemakers were also successfully applied in other
neural prostheses. Restoration of hearing to the profoundly deaf was an early
target.

A report by Gersuni and Volokhov
in 1936 (10)
indicated that hearing could be produced over a normal frequency range, and that
it persisted following the surgical removal of the tympanic membrane and
ossicles. One of the first recorded attempts to stimulate the auditory nerve was
made by Lundberg in 1950, who employed sinusoidal currents during a
neurosurgical operation. The patient could only hear noise.

However, a milestone for the
development of the auditory prosthesis was the electrical stimulation of
the acoustic nerve of a deaf man by an implanted electrode achieved by Djourno
and Eyries (11) .
They used an implanted induction coil connected with one end to the inner ear
electrode and with the other to an indifferent electrode at the temporalis
muscle. Another coil, outside the body was held over the implanted coil to
inductively transmit the signal from a microphone. The usefulness of the device
was very limited since the patient could recognize only few words from the
transmitted signal ('papa', 'maman', and 'allo'). This operation, however,
opened the doors for further attempts to interface the sense organs.

The early investigators were
subjected to severe criticism by both neuroscientists and clinical professionals
for the insufficient safety studies and unsatisfactory design of the implants
used. The safety of the cochlear implantation improved in the late 1950s with
the evolution and developments of microsurgical techniques and stapes surgery.
The indications and contraindications for this implantation were elaborated in a
broad debate between the clinicians and the promoters of the cochlear
prostheses. Questions were also raised about which device should be used and
about the safety of the electrical stimulation. The seventies could be
considered as a decade of controversy concerning the implant itself and its use.
Finally after more than twenty years of on-going experimentation, in 1984, the
cochlear implant received FDA approval for use in adults in USA. Nowadays the
cochlear implant is a historical example of how a neural prosthesis can
revolutionize treatment.

In 1929, Foerster described the
effects of electrical stimulation of the human visual cortex (12) .
The subjects reported for “perception” of a small point of light during
stimulation, which was later called a phosphene This result was reproduced many
times in both sighted subjects and blind subjects. The idea that concurrent
stimulation of many sites in the brain could produce a single coherent image was
postulated by Krieg (13) ,
based on the retinotopy in the visual cortex.

As a second milestone for the
development of neural prostheses can be accounted the implantation of prosthesis
in the visual cortex of a blind woman in 1968 by Lewin and Brindley (14) .
The prosthesis did not benefit the patient much (she reported only about
occasional phosphenes in her visual field), but demonstrated that it is possible
to transfer large amount of information in the visual system. Giles Brindley,
one of the greatest protagonists in the FES field, attained recognition and
acceptance of the applications of neural prostheses in the clinical field.

The results obtained by the
Brindley implant and the early cochlear stimulator showed the need to stimulate
selectively small groups of neurons or fibers in order to transmit meaningful
information into the central nervous system.

Attempts to
restore the lost functions of the paralyzed leg muscles were performed for the
first time in 1961 by Liberson and colleagues (15) .
The system was developed to compensate for the “drop foot” problem in hemiplegic
stroke patients. By stimulating the peroneal nerve, the prosthesis triggered
ankle dorsiflexion, eversion and inversion. Since 1961, a number of
neuroprostheses for restoration of walking, hand and arm functions have been
designed and tested with various patients and with different levels of success.

Apart from these well-endorsed
success stories, other applications of FES still suffer from major difficulties
regarding the clinical acceptance and patient compliance. Two examples are given
here:

Long and
Masciarelli introduced prosthesis for the hand that could not meet the
expectations of the patients and it was later on abandoned.

In the 70s,
attempts to treat refractory epilepsy were also carried out by means of
stimulation of the cerebellar cortex. Later on, the performed double blind
clinical trials did not confirm the claimed effects (16)
.

In
conclusion: History of FES shows that pacing of the heart emerged at
different places and its success stimulated pacing of the diaphragm.

The criticism of
the early cochlear prostheses could be surmounted because of tenacious research.
However, one might consider these first successful applications as “lucky
shots”, based on the simple reactions of the involved organs:

For the cardiac
muscle and the diaphragm – simple contraction, and since the tonotopically
organized inner ear can easily decode the transmitted signal, the success of the
cochlear implant was understandable.

Neuroprostheses
for locomotion as well as those for reaching and grasping, on the other hand,
require selective stimulation of the fibers in the nerves, which innervate the
muscles making the movements. Therefore, in the case of selective nerve
stimulation, the research should be directed towards localization of the types
of (myelinated) axons and their spatial organization.

FES tries to restore the lost
functions of the nervous system by means of electrical stimulation. This
can be achieved by either rehabilitation thus improving recovery or by
prosthesis (artificial substitution of a dysfunctional part) of the lost
function or orthosis (device around organs in the case of impairment) of
the diminished function. Example of the latter is the peroneal stimulator
for foot-drop and of the former the stimulators of the phrenic nerve. FES is
also sometimes referred to as Functional Neuromuscular Stimulation or FNS. This
approach is limited only to the sensory and the motor functions. The initiation
of reflex activity is, though, not the aim of FES but may be a secondary and an
useful collateral effect.

Apart from the clinically
recognized “success stories” of electrical stimulation FES for restoration of
the lost locomotor functions is a rapidly developing area that is on the verge
of broadening its applications. Evidence for this is the increasing number of
various neuroprosthetic devices that have been developed recently. Some of them
are already beyond the prototype stage. That is why this paper will deal further
on primarily with the electrical stimulation of the motor system.

Principal
notions in FES

Electrochemical processes at the tissue-electrode
interface

Electrical stimulation of
biological tissues with metal electrodes requires the flow of ionic charge in
the biological tissue. This flow is induced by both capacitive and faradic
mechanisms. The capacitive mechanism involves periodic charge and discharge of
the electrode double layer. There is no charge transfer across the
electrode/electrolyte interface. This is an ideal mechanism for charge
injection, but it is limited by the maximal amount of charge that can be
transferred before the dielectric breaks down. This is approximately 20
mC/cm2. Because the charge required for physiological stimulation
exceeds this limit, all electrical stimulation is performed by faradic charge
injection. The faradic mechanism involves charge transfer across the
electrode/tissue interface and, therefore, electrochemical reduction/oxidation
processes. These reactions can be either reversible or irreversible. All
irreversible electrochemical reactions are undesirable since they alter the
chemical composition of the extracellular fluid, producing cytotoxic products or
bringing about large changes in pH. For every material used in the production of
electrodes, there is a charge limit for the reversibility of the electrochemical
process. This charge limit depends on the particular properties of the material,
the shape of the electrode, and its size, and the stimulation waveform. The
temporal shape of the stimulation wave is the most important according to
Robelle and Rose (1990) (17)
.

In order to diminish the oxidation
and dissolution of the electrodes generally a noble metal like
Pt, Ir or Au is used as
material for the anode. Of the non noble metals 316L stainless steel is also
used for intramuscular electrodes (18) .
At the cathode, however, O2 molecules are reduced to
O2-, which in turn react to produce free radicals. The
free radicals are known to damage the cell membranes and DNA molecules. During a
typical neurostimulation pulse, a substantial amount of the O2
molecules is reduced to anions (19) .

Part of the free radicals
generated near the cathode can be neutralized by oxidation when an anodic pulse
follows the cathodic pulse. Morton et al. (1994) (19)
concluded that the reduction of oxygen, and thus the generation of free radicals
are restricted most when the cathodic pulse has a small width and when an anodic
pulse with the same charge immediately follows the cathodic pulse. This
principle is known as biphasic charge-balanced stimulation (20) .
The continuous stimulation with biphasic, charge-balanced pulses without delay
between the pulses are more beneficial with regard to nerve degeneration than
those with delays (21) .

The continuous stimulation with
biphasic, charge-balanced pulses was shown not to have harmful effects to
sciatic nerves (22)
.

Voltage-controlled and current-controlled
stimulation

Historically the implantable FES
systems in clinical applications have generally been voltage-controlled, as for
example are cardiac pacemakers. In contrast, the current-controlled devices
dominate in experimental set-ups. In the voltage-controlled devices, the output
voltage V of the pulse generator is regulated. Generally V is kept
constant, thus creating a rectangular voltage pulse, the related current obeying
Ohm’s Law. In contrast, in the current-controlled devices the current I
is kept constant, whereas V is automatically adapted according to the
value of the impedance Z, which may vary over time due to the reaction of
the surrounding tissue to the implant. The current needed for stimulation does
not change chronically and thus no corrections are needed. In the
voltage-controlled systems, however, the voltage needed for stimulation is
influenced by the changing value of Z over time and thus the amplitude of
stimulation has to be adapted at least in the first two months after
implantation when the fibrous encapsulation layer develops.

Since the current consumption
determines battery life of an implanted pulse generator, it is important that
the stimulation current is determined in the first place. This is, however, not
possible when a voltage- controlled pulse generator is used, since I depends on
the unknown value of Z. Therefore, the battery life cannot be accurately
predicted from V. Nevertheless, it is often erroneously assumed that a
reduction of V reflects a reduction of I and vise-versa.

Monopolar,
bipolar and tripolar stimulation

Since the cathodic threshold for
nerve fibers is 3-7 times lower than the anodic threshold current, cathodic
stimulation is by far the most efficient way (23) .
In monopolar stimulation, the active electrode (in or near the neuronal target)
is therefore the cathode and the distant, indifferent electrode, is an anode. In
that situation, the current injected by the cathode is distributed more or less
evenly in all directions.

Activation of nerve fibers always
happens near a cathode in bi-, tri- and multipolar stimulation. For the optimal
stimulation, in bi-, tri- and multipolar stimulation, the position of the
cathode is more important then the position of the anode. In the clinical
applications of FES phenomena such as cathodic block, anodic excitation and
anodic block are not likely to occur (24) ,
(25)
. The threshold current for stimulation of a nerve fiber increases with the
distance between the cathode and the fiber and is inversely related to the fiber
diameter. The preferential stimulation of large fibers is favored most when
short pulses (~ 60 ms) are applied. Smaller fibers can be activated more easily
when pulses are wider. The current consumption is minimized when stimulation is
given with a “guarded” cathode parallel to the fiber bundle as in peripheral
nerve stimulation and Spinal Cord Stimulation (SCS).

Cable
properties of the nerve fibers

The theory of the current flow in
electric cables was developed for submarine cables by lord Kelvin (1855). It was
first applied in studies on excitable cells towards the end of XIX
century.

Cable theory opens the possibility
to derive equations for current flow in cylindrical nerve and muscle fibers
subjected to voltage changes which are small enough for the membrane properties
to be linear; or equations that describe finite cables, since in most
experiments the length of the preparation is not large enough, compared to the
space constant l for the cable to be regarded as infinite.

This means that the behavior of
the impulse in an axon can be predicted.

Active
properties of the nerve fibers

The nerve fibers conduct
information by means of action potentials. The action potentials propagate at a
velocity proportional to the fiber diameter.

Table
1. Morphology and electrophysiology of the fibers in Peripheral nervous
system

Fiber type

class

MCV1
(m/s)

Diameter (μm)

Function

Aα

myelinated

70-120

12-20

motor somatic muscle fibre;
proprioception; muscle spindle annulospiral; proprioception; Golgi tendon
organ

Aβ

myelinated

70-120

5-12

proprioception; muscle
spindle flower spray; exteroception; touch & pressure

Aγ

myelinated

30-70

3-6

motor somatic muscle
spindle

Aδ

myelinated

12-30

2-5

exteroception; pain
temperature (some ) touch

B

thin myelinated

3-14

<3

motor autonomic
preganglionic

C

unmyelinated

.5-2.5

0.4-1.2

exteroception; pain reflex
responses

C

unmyelinated

.5-2.5

0.3-1.3

motor autonomic
postganglionic sympathetic

1Mean Conduction
Velocity

Moreover, the threshold of
excitations also depends of the fiber type.

In the myelinated fibers, the
conduction is saltatory, with the active involvement of the nodes of Ranvier.
For the first time the action potential of the non-myelinated fibers were
mathematically described by Hodgkin and Huxley in 1952 (26) .

The electrical behavior of the
myelinated nerve fiber can be represented by a simple cable network. McNeal was
the first to introduce this model (27) in
order to calculate how a stimulation-induced extracellular field affects nodal
transmembrane voltages. According to this model, the node of Ranvier that is
closest to the cathode will be excited first when the stimulus current is
sufficiently high. The results of this theoretical approach are in accordance
with the general observation that the threshold stimulus of nerve fiber
excitation is smallest near the cathode and rises with the increase of the
distance (23) .
As a result, cathodic stimulation gives rise to a depolarization of several
nodes, whereas a larger number on both sides of the cathode are hyperpolarized
to a lesser extent, thus creating a virtual anode.

Figure
3. The McNeal cable model for the axon. Each node of Ranivier is represented
by an equivalent electric scheme. Vi – intramembrane potential, Ve –
extamembrane potential, Ra –intraaxonal resistance, R – variable membrane
resistance, C- membrane capacity, V – resting membrane
potential

In anodic stimulation the opposite
effects occur. The virtual anode and virtual cathode effects
underlay phenomena such as cathodic block and anodic excitation. Anodic
excitation is possible when the anodic current is so large that the virtual
cathodic depolarization on either side of the hyperpolarized membrane generates
action potential. In cathodic block, the opposite is true – if the cathodic
current is large enough the induced virtual anodic hyperpolarizations can block
the propagating action potential.

The McNeal model is the basis of
all the theoretical research performed in the area of FES.

Stimulation
Electrodes

Spatially selective stimulation of
the nerve fibers within a given nerve bundle is important, because different
muscles are innervated by fibers possessing a specific localization within a
nerve bundle, and moreover often a mixture of efferent and afferent fibers is
present in the nerve. Such mixed populations may lead to undesirable sensations
during stimulation. In the case of several nerve fibers innervating different
muscle fibers of one muscle, cyclic stimulation can be performed which
guarantees an overall stimulation frequency, while stimulation frequency and
thus fatigue of a single muscle fiber can be kept low (28) .

Many different designs of
electrodes exist. They employ different materials and geometries of the
contacts. Most of them are used only in animal experiments, and only few are
employed in clinical FES applications as well.

Several types of stimulation
electrodes for FES have been developed. A simple subdivision is

Surface (skin)
electrodes

Intramuscular
and epimysial electrodes

Extraneural
electrodes (cuff or helical electrode around a peripheral nerve)

Intraneural
wire electrodes

Intraneural linear, 2D and 3D
multi-electrode arrays

An extensive
account on the cuff, intrafascicular and sieve electrodes designs and
fabrication technology can be found in (29) and
(30)
.

Surface
electrodes

Transcutaneous stimulation is performed
with self-adhesive or non-adhesive electrodes that are placed on the subject’s
skin, above the peripheral nerve (respectively the motor fibers). The metal
electrode is connected to the skin with a saline bridge.

Criteria
for surface electrodes are low impedance with even current spread, flexibility,
ease of application and removal and lack of skin irritation (31) .
The application is simple. Damage of the skin only can occur by the diffusion of
soluble electrochemical products. The large distances to the stimulation targets
and the insulation of the skin and the fat tissue demand high stimulation
strengths, which in turn results in low selectivity. Complications are rare and
include burns, skin irritation, erythema and local pain.

Intramuscular and epimysial electrodes

Muscle electrodes have their
stimulating surfaces onto or inside a striated muscle. In the former case, they
are sutured to the epimysium of the muscle and, therefore, are called epimysial.
If terminal motor branches are present, the electrodes exert their action by
stimulating the nerve endings. Therefore, it is correct to consider them also as
nerve electrodes. In comparison with the transcutaneal electrodes they produce
contractions with lower currents and with greater selectively. For these
reasons, muscle-based electrodes are preferable for situations that require
independent control of several isolated muscles.

Depending on their intended
application, intramuscular electrodes can be introduced either percutaneously
(see for example (32) )
or in an open surgical procedure (33) .
However, intramuscular electrodes can also activate neural structures near their
stimulating tips other than those intended, thus recruiting muscles in addition
to the ones targeted (34) .
Nevertheless, intramuscular electrodes are considered safe and effective means
to produce strong and isolated contractions of single muscles.

Complications include infections
and wire breakdowns. In an effort to avoid the use of percuataneous leads for
muscle electrodes, Loeb (2001) (35) and
Cameron et al. (1997) (36)
developed implantable muscle stimulators that can be controlled telemetrically,
and are inductively powered from external magnetic devices in order to overcome
infections. According to the promoters of such an approach, this would
ultimately alleviate the burdens involved with maintenance of the awkward
external electronic equipment.

Cuff
electrodes

Cuff electrodes circumscribe the
nerve. They were introduced in the 1970s. The first cuff electrodes had a very
simple cylindrical rigid design. Modern cuff electrodes use flexible materials
and adaptable geometries.

They are
implanted by a simple surgical procedure, and their size can be adapted to the
diameter of the nerve of interest. If needed, they can be surgically fixed to
adjacent tissues.

Various designs have been
developed and applied such as the helix-shaped electrode (37) , the spiral-cuff
(38)
, and the so-called "half-cuff" electrode of (39) .
Another particular design is the arrangement of flexible interdigitating
sub-units with microelectrodes along a backbone-like carrier (40)
.

The main push for their continuing
development was the inadequate spatial and fiber selectivity they provided at
first. The first cuff electrodes could stimulate only the most superficial
fibers of the nerves (29)
.

With a sufficient number of
electrodes within the cuff, high spatial selectivity of stimulation can be
achieved by combinations of longitudinal and transversal currents produced by
appropriately switched electrodes (41) .
Longitudinally aligned tripolar dot electrodes on the surface of a nerve trunk
restrict excitation to superficial nerve trunk regions more successfully than
will monopolar dot electrodes do (42) .
Selectivity was dependent on the relative location of the electrode contacts and
the nerve fascicles, as well as the size and relative spacing of neighboring
fascicles of a 12 contact spiral nerve cuff electrode (43) .
On the basis of computer modeling results Deurloo et al. (2003) (44)
concluded that transverse steering provides more selective stimulation than
longitudinal steering. However, the problem of different fiber diameters, which
leads to different activation thresholds of the fibers, is not solved by these
methods. Goodall et al. (1996) found that large fibers were activated before
smaller with a cuff electrode containing 12 electrodes arranged in four
longitudinal tripoles, irrespective of the fiber position (45) .
Transverse current from an anode positioned opposite to the stimulating cathode
was found to improve spatial selectivity, and position selectivity was enhanced
when the ratio of transverse current to longitudinal current was increased.

The use of silicone and helical
designs in cuff electrodes has improved surgical access and reduced nerve damage
(46)
, (47) and
(41)
.

Figure 4. A schematic view of a cuff electrode. Courtesy:
K Deurloo, from (96) p.
66

Figure 5. (a) Scheme of the University of Twente
128-electrode 3-D glass-silicon array (UT-128 array), mounted on a CMOS , mixed
mode, processing chip with dimensions 4x4 mm. Needle length is 600, 425 or 250
μm, width at tip is 15 μm, needle spacing is 120 μm. (b) Details of the
dimensions and materials used for the UT-128 array. (c) A ‘sea’ of sawn and
etched silicon needles of three different lengths, embedded in a glass matrix.
Courtesy: W. Rutten, from (30).

Intraneural
electrodes

In an attempt to overcome the
selectivity problems with stimulation cuff electrodes intraneural electrodes
were developed. They are another type of electrodes directly contacting nerve
fibers within a given nerve bundle.

Intraneural
wire electrodes

The first penetrating electrodes
were simply thin metal wires or needles, which were inserted into the nervous
tissue. They are used for both recording and stimulation. An interesting,
non-silicon approach for intrafascicular stimulation is the use of tethered
platinum microwires. Nannini and Horch (1991) (48)
tested the performance of Pt-Ir intrafasicular electrodes (25
mm diameter) implanted in nerves innervating the gastrocnemius and soleus
muscles. The electrodes were originally developed for recording purposes. They
concluded that intrafascicular recording electrodes are also a potentially
suitable technology for FES. Yoshida and Horch (1993) (49)
used dual intrafascicular electrodes to study activation of nerve fibers by
pairs of Pt-Ir wire electrodes implanted within single
fascicles of the nerve innervating the gastrocnemius muscle in cats, in order to
determine whether intrafascicular electrodes can activate nerve fibers in
different fascicles independently of each other and if they can be used to
activate separate subsets of axonal populations within a single fascicle.

Intraneural
multi-electrode arrays (MEAs)

A simple classification consist
of

a) Linear
multielectrodes:

Wedge-shaped microprobes carry a
line of electrode sites for recording and stimulation, the so-called
"one-dimensional" (1D-) arrays (50)
.

b) Two- and three-dimensional
multielectrodes

The 2D-array consists of
electrodes with tips ending in the same plane. The tips are configured either as
a bundle of wires or galvanically grown needles. The substrate carrying the
electrodes is either needle- or wedge-shaped to allow penetration of the nervous
tissue which makes recording from and stimulation of axons possible not only on
the surface but also in a well-defined depth within the tissue, e.g. within the
fiber bundle. When the array is slanted or the needles have otherwise
variable lengths the electrode is named a 3D-array. Implantation of any such a
device is always associated with some damage of the nervous tissue, cell death
of neurons and disruption of axons. Moreover, stiffness of many models may also
lead to mechanical damage of nervous tissue. Thus, the efforts are directed to
miniaturize the penetrating parts of the implant and to use materials that are
more flexible.

Rutten et al. produced a 3D needle
array with 128 recording sites on one electrode placed on the tip of a needle
(51) . The needles are made of silicon and are embedded into a glass substrate.
They vary in height from 250 to 600 um and have a distance of 120 um, with a tip
size of 15×15 um. The array is intended to serve as an interface between a
stimulator and a peripheral nerve. The different length of the needles allows
stimulation in the volume of the nervous tissue. The dimensions of the array
were based on calculations for optimal stimulation, based on random uniform
distribution of the nodes of Ranvier in the peripheral myelinated
axons.

For signal recording and
stimulation over a larger area, e.g. areas of the cortex, two-dimensional MEAs
are necessary. An example of such a 2D array is the Utah Electrode Array
initially developed by Nordhausen and colleagues (52) .
It is designed for implantation in the human visual cortex and is intended as
prototype in development of visual prostheses (53) .
It provides a multichannel interface to the visual cortex using a large number
of 1.5mm long electrodes (typically 100 in a 10 x 10 square grid) projecting out
from a very thin (200 um) substrate and are separated from each other by 400 um.

With the technological advance in
the field of constructing flexible polymeric substrates, so-called "flexible
nerve plates" came into development (54) .
They consist of a flexible substrate, which carries a Micro Electrode Array and
it is possible to add other elements of a microelectronic
circuitry.

In general the silicon
microtechnology based fabrication leads to improvement of recording structures,
which made it possible to create not only planar two-dimensional Micro Electrode
Arrays but also intraneural electrode structures for in vivo
measurements.  Alike the cuff electrodes, geometry and arrangement of the
intraneural electrode arrays are currently optimised to fit with geometrical
nerve fiber characteristics (fiber position in a bundle, diameter of the fibres
and their orientation). While Smit (1996) argues for intrafascicular
stimulation, the proposed approach could be employed only in acute experiments,
thus limiting the use of the obtained knowledge for clinical applications (55)
.

In theory intraneural electrodes
should give better spatial selectivity than the cuff electrodes (56, 49),
but the distribution of the nodes of Ranvier in the peripheral nerves makes the
stimulation a probabilistic process. The implanted intraneural electrode may
simply miss the target fiber population when the topography of the nerve fibers
is unknown.

Also since during implantation
these intraneural electrodes must penetrate the relatively tough epineurium and
the perineurium of the nerve, they can cause short- and long-term nerve damage
during the electrode insertion procedure (57, 58, 59, 60 ).
Nevertheless, histological results suggest that intraneural electrodes can be
used chronically in peripheral nerves (61, 62, 63).

The main prerequisite for
successful application of the modern sophisticated electrode designs and
stimulation protocols is detailed knowledge of the topography of the nerve
fibers within the nerve bundle and their function. Moreover, a really reliable
fixation of the cuff must be achieved because otherwise the cuff may rotate
around the nerve or shift along the nerve, both leading to a serious loss of
selectivity.

In conclusion: The implantable
electrodes are surgically set in: around, within or adjacent to a nerve trunk or
root. By means of transcutaneous electrodes it is difficult or impossible to
selectively activate individual muscles, especially deep muscles such as hip
flexors (64, 65).
Large currents may be required to drive sufficient charge through the skin and
intervening tissues between the electrode and the peripheral nerve. In many
cases, cutaneous pain receptors are excited, and patients with preserved or
heightened sensation may find it difficult to tolerate surface stimulation at
the levels required to produce a functional motor response.

Advantages of the implantable
electrodes over the superficial, epimysial and intramuscular include:

The current needed to excite an
action potential in the target fibers is smaller in at least an order of
magnitude over the current required for the latter designs.

Muscle length and limb position
are expected to have substantially less effect on the recruitment
characteristics in the case of an implantable electrode.

Cuff-electrodes are easy to
implant and are generally less invasive to the nerve than intraneural electrodes
are.

Despite all new developments in
electrodes, the robustness of stimulation is hampered by the structure of the
nerve:

The random distribution of the
nodes of Raniver

The typical
fasciculation

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.

Research
Problems

The major issues
in the development of a neural prosthesis from the point of view of the
neuroscientist are:

Biocompatibility of the materials, which
determines the long-term use of the device.

Detailed
neuroanatomical and neurophysiological knowledge of the system, which in the
end determines the effectiveness of the device.

Biocompatibility of the implantable devices: general
considerations

Implantable devices should be able
to survive in the rather aggressive internal milieu of the body without causing
significant tissue reaction. The major prerequisite for the application of
implants is that the organism accepts the implant, i.e. that the implant is
biocompatible. This also holds for objects that come in close contact with the
skin or mucose. An accepted definition of biocompatibility is "the ability of a
material to perform with an appropriate host response in a specific application"
(66)
. The selection and evaluation of materials and devices intended for use in
humans requires a structured program of assessment to establish biocompatibility
and safety.

Current regulations, whether in
accordance with the U.S. Food and Drug Administration (FDA) (ISO 10993-1/EN
30993 standard, since 1995), the International Organization for Standardization
(ISO), or E.U. regulation bodies (The EU council directive - 93/42/EEC, since
1993), as part of the regulatory clearance process require conduction of
adequate safety testing of the finished devices through pre-clinical and
clinical phases (67)
. An extensive account on the biocompatibility may be found in the standard
ISO 10993-1/EN 30993.

In brief, the notion of
biocompatibility encloses non-toxicity, non-immunogenity, non- carcinogenicity,
non-mutagenity and haemocompatibility. An implant can be considered
biocompatible if it gives negative results on the following
tests:

Cytotoxicity

Sensitisation

Genotoxicity

Implantation

Chronic
Toxicity

Carcinogenicity

Irritation
Intracutaneous

Acute Systemic
Toxicity

Subchronic
Toxicity

In other words, it includes the
whole behaviour of the implant in its biological environment.

According to Heiduschka and Thanos
(1998), the "biosafety" and the "biofunctionality" also have to be considered
(29) . Biosafety means that the implant does not harm its host in any way, and
biofunctionality means that the implant acts in the body as it was intended. In
addition, "biostability" is important which means that the implant must not be
susceptible to attack of biological fluids, proteases, macrophages or any
metabolic substances. For example, implants may be subject to continuous attack
by hydrolytic enzymes (68) or
free radicals produced by monocytes and/or cell lysis. Stability of implanted
material is important not only for a stable function, but also because
degradation products may be harmful to the host organism. Overview about
biological reactions to implanted materials can be found in Ratner et al. (1996)
(69)
.

Chemical
biocompatibility

Chemically appropriate implant
materials are designed to be as inert as possible. If chemical reactions are to
be expected they should be minimal and all resulting products should be inert.
Candidate materials for neuroprothesis pass especially rigorous testing since
they must remain inert not only passively but also when subjected to electrical
stimulation. Typical materials for nerve electrode arrays include
platinum-iridium alloys or stainless steel for the conducting parts and epoxy
resins, polytetrafluoroethylene (PTFE, Teflon®), silicone rubbers and
polyimide for insulators (70, 29) .
These polymers are biocompatible, electrically insulating and stable.

The amount of platinum ions
released into the surrounding tissue may be neglected even after long-term
stimulation. During the last years, iridium has been of increasing importance
because a stable oxide film can be formed on the surface of iridium electrodes.
This oxide film has a big charge delivery capacity and is, for this reason, well
suited for stimulating electrodes. Platinum and iridium are established
materials in microelectronics, and carbon can be deposited onto microelectronic
structures. Glassy carbon or carbon fibers are also used as electrode materials,
and they are biocompatible and stable, though they have a higher roughness than
metals.

The bulk properties of polymers as
backbones for electrodes can be modified to a certain degree and also surface
modification procedures are performed in order to improve biocompatibility.
However, certain surface modifications like autologous protein coatings also
enhance the adhesion of pathogenic bacteria to the implant surface (71, 72).
For a review on the chemical modifications employed in the production for neural
prostheses, see (29)
.

Mechanical
biocompatibility

The mechanical biocompatibility of
an implant depends on the basic mechanical properties of the tissue and the
implant. Ideally, the implantable device should have mechanical properties
similar to the tissue in which it is implanted. This includes flexibility,
strength and durability. It is also important to consider the tissue motion
within the implantation site. For example, flexibility is important in the
peroneal foot-drop stimulator and not in the devices for spinal cord stimulation
and anterior lumbo-sacral root stimulator.

For any implant, the allowable
size must be defined relative to its function. It is not necessarily true that
increased function is associated with larger size. The geometrical
characteristics of an implant are of great importance too. For example, sharp
edges and blunt corners should be avoided.

The primary goal for a cuff
electrode design is to minimize the size, however any constriction injury in the
implanted site should be avoided. Such designs have been associated with
incidence of morphological changes in neural tissue like axonal degeneration,
demyelinization and fibrosis (37) .
Thus, the recommendation had been to implant cuffs with an internal diameter
equal to 150% of the nerve diameter (73) .
This recommendation was criticized as to be actually more likely to create
neuronal trauma (37) .
It also limits the degree of selectivity that is obtained during extraneural
stimulation (41, 31) .

Implantation
and the Nerve damage

The key factor determining the
damage of electrode implantation is the disruption of the microcirculation of
the nerve, which results in edema. This holds especially for intrafascicular
electrodes. Another factor is the damage to the endoneurium and the disruption
of the homeostasis, maintained by the blood-nerve barrier. The mechanisms of the
damage during implantation are still debatable. According to Agnew and McCreery
(1990) on the first place stands direct mechanical interaction between the
electrode and the nerve (70) .
Other mechanisms that may play a role during implantation are:

Surgical trauma to either the
neural microvasculature or the nerve itself.

Pressure caused by
post-surgical edema, seroma formation, or excessive fibrous encapsulation of
the implant.

Reduced mobilization of the
nerve caused by excessive scar tissue formation, which could fuse the nerve to
the surrounding tissues.

Undue tension in the electrode
leads.

The transmission of forces from
muscles to the electrode array and the nerve.

Obstruction of the
microstimulation

For epineural and intraneural
designs, though, the most important factors are the implant procedure itself and
transmission of tension via the leads.

Tissue
reaction to implantation

Implantation usually causes
foreign body reaction involving the natural inflammatory defensive response of
the body. The acute response starts with protein adsorption onto the
surface of the implant. Therefore, the surface properties like roughness and
surface chemical composition, which depend on the production process, are
important for the grade of the subsequent response (74) .
The late response produces the typical cellular and humoral reaction
including tissue edema, cytokines secretion, neutrophil, lymphocyte and
macrophage migration and adhesion. This is a complex process mediated by the
secreted cytokines.

Chronically a fibrotic
capsule around the foreign body is formed, consisting of collagen fibers and
fibroblasts. The thickness and the structure of the capsule is generally a
measure of the severity of the response. The severity of the reactions can be
determined by the amount of the giant Langerhans cells and
macrophages.

The electrical properties of this
fibrotic capsule were studied by Grill and Mortimer (1994) (75).
They found that resistivity of the encapsulating tissue had a frequency
dependency between 10 Hz and 1kHz and decreased from 454 ± 123 to 193 ± 98
Ohm.cm and was frequency -independent between 1 kHz and 100 kHz with a mean
value of 195 ± 88 Ohm.cm. Moreover, the fibrotic capsule may lead to
displacement of the electrode positions and changes in the tissue impedance (75)
.

The results from early experiments
(37,
76)
, indicated that nerves implanted with snuggly fitting spiral cuffs showed signs
of sustained trauma. The typical pattern was a crescent-shaped region
containing thinly myelinated axons, proliferation of subperineurial connective
tissue, and a reduced axon density. In a more extensive study of the tissue
reaction to the spiral cuff, 4 of 44 nerves exhibited peripheral crescent-shaped
areas which contained thinly myelinated fibers with an apparent reduction in
axon density, but there was no correlation between the cuff-to-nerve diameter
ratio and the presence of morphological abnormalities. Similar results have been
found in cuff electrode studies conducted by researchers at the Huntington
Medical Research Institutes (77)
.

Recently Grill and Mortimer (2000)
(78)
reported focal areas of abnormal neural morphology including perineurial
thickening, endoneurial fibrosis, thinly myelinated axons, and focal reduction
in the density of myelinated axons in chronic implantation experiments with cuff
electrodes.

Intrafascicular electrodes can
also produce morphological changes in neural tissue (63) .
Chronic implantations of intraneural coiled-wire electrodes showed endoneurial
fibrosis and edema, loss of nerve fibers with the large myelinated fibers being
the most susceptible, and variable shifts in the excitation threshold. However,
there were no changes in conduction velocities indicating that no significant
damage occurred due to implantation (61) .
The bulbous enlargement formed at the point where the electrode penetrated the
perineurium was associated with focal nerve fiber compression, demyelination,
edema, and fibrosis (61, 79) .
Electrodes sutured to the epineurium have also led to neural damage including
endoneurial edema, endoneurial fibrosis, loss of axons, and reduction of myelin
(80,
81)
. In general, intraneural electrodes are associated with greater risk of trauma
than the cuff-electrodes (82)
.

Nervous
tissue reaction to electrical stimulation

Brief stimulation of the
peripheral (83) or
the cranial nerves results in increased expression of Immediate Early Genes
(IEG) in the related neuronal cell bodies (for example c-Fos (84) ).
The family of the Immediate Early Genes consists of approx. 30 genes, with
related proteins acting as transcription factors (85) .
Members of the IEG family in general are activated shortly after cell
stimulation and without the requirement for de novo protein
synthesis.

Brief unilateral electrical
stimulation of the cochlear nerve (120–250 mA, 5 Hz, 30 min) in anaesthetized
rats with a biphasic current resulted in increased expression of c-Fos in the
ipsilateral ventral and bilateral dorsal cochlear nucleus (84) .
Intracochlear electrical stimulation with a cochlear implant in rats lead to
changes in the phosphorylation state of the cAMP response element binding
protein (CREB) and the expression of the IEG family members c-fos and
egr-1 in a tonotopically precise pattern in the central auditory neurons
(86)
. These neurons resided nearly in all auditory brainstem nuclei. Moreover,
effects of electrical stimulation were identified in the medial vestibular
nucleus and the lateral parabrachial nucleus. Regionally, CREB was
dephosphorylated, wherever immediate-early gene expression went up. These
massive stimulation-dependent modulations of transcription factors in the
ascending auditory system indicate ongoing plastic changes as a consequence of
the stimulation of the inner ear.

In the spinal cord, Molander et
al. (1992) (87)
showed that in chronically axotomized nerves, c-fos is expressed after
electrical stimulation of the fibers. Stimulation of the normal sciatic nerve at
C-fiber intensity resulted in c-fos protein-positive cells within the sciatic
projection territory in the ipsilateral dorsal horn (lamina I and outer lamina
II) and stimulation of A/B fibers had little effect. The expression was
delimited to the projection areas of the sensory A/B-fibers (ipsilateral laminae
II, III and IV and in the gracile nucleus) in young animals. This suggests that
the excitability of these neurons is increased by nerve injury.

The role of the IEGs is implied in
the structural plasticity of the nervous systems (88, 89).
One may speculate that the continuous improvement of the responses of the
individuals after early rehabilitation, in contrast with late rehabilitation, is
due to FES being able to guide the plastic phenomena occurring after injury in a
direction towards restoration of function. This line of thought is supported
also in the review on animal experimentation (90)
showing accumulation of physiological and behavioral data that adaptive
(plastic) processes also occur within spinal circuits. The potential ability of
the spinal cord to “learn” has obvious implications for altering and improving
locomotor function after injury.

Recent studies in man indicate
that a significant population of stroke and Spinal Cord Injury (SCI) patients
could also benefit from FES rehabilitation (91, 92). In
particular, it was found that stroke patients as well as incomplete SCI patients
subjected to intensive FES treatment were able to recover grasping or walking
function faster and better compared to patients who did not participate in the
FES treatment post injury.

In the peripheral nervous system,
prolonged high-frequency (f> 20-50 Hz) electrical stimulation of a
peripheral nerve induces a typical type of neural injury, called early axonal
degeneration (EAD), with a characteristic “salt-and-pepper” like mixture of
“normal” and damaged fibers. The primary damage is observed as collapse of the
myelin into the axoplasm and is restricted primarily to the large caliber axons
(Aa, Ab type). It is followed by degradation and phagocytosis of the axon. (77, 93, 82) .
The neuronal injury originating from electrical stimulation has two mechanisms
of occurrence:

a)
Direct injury resulting from the electrochemical process near the
electrodes. Brummmer and Turner (1972) (94)
have shown that the rate of production of compounds by electrochemical
reactions and the type of compounds produced are directly related to the
charge density (charge transferred per unit area of the electrode surface).
The charge density near the surface of an electrode determines the extent of
depolarization and hyperpolarization induced in the neurons close to the
electrodes. If such depolarization is abundant and going on for a long time
the second mechanism of injury comes into play.

b)      The
excitotoxic neuronal injury, which is caused by the excitatory
neurotransmitter glutamate through its NMDA receptors. This view is
supported also by the finding that MK-801, a potent NMDA receptor antagonist,
is a neuroprotective factor during prolonged electrical stimulation (95).
The mechanism of the induced neuronal cell death is necrotic. Such damage
occurs when the stimulus charge density or the charge per phase is
large.

Selectivity
concepts

In the paradigm of the implantable
devices, selective stimulation of individual components of multifascicular
nerves would allow control of several muscles with less hardware (43) .
Different aspects of selectivity may be distinguished depending on the
structures being stimulated. Here we propose modification of the classification
of Smit (55)
.

Muscle selectivity is
the possibility to activate any specific muscle by peripheral nerve
stimulation. This implies stimulation of only a particular part of a
(peripheral) nerve i.e. spatial selectivity. In other words, it implies
control at the level of the individual muscle

Spatial selectivity is
the possibility to stimulate a particular region inside a nervous
structure.

The peripheral and the spinal
nerves generally consist of several morphologically distinct sub-organs,
called fascicles. Therefore, a suitable definition of fascicle
selectivity is the stimulation of only one fascicle in a multifascicular
peripheral nerve without spread of the activation to other fascicles. It is a
special case of spatial selectivity. Fiber selectivity is another
special case of the spatial selectivity. It can be defined as activation of
single nerve fibers.

The nerves consist of axons, which in general have different sizes.
Usually the fiber size distribution is multimodal. Size selectivity can
be defined as activation of fibers of a particular size group.

Surface stimulation vs. Implantation

An important advantage of the
surface FES systems is that they do not require surgical intervention, with its
inherent risks. They also can be removed at any time if contraindications arise.
Another advantage is that the transcutaneous electrode can be placed directly
above the target muscle, and thus muscle-selective stimulation can be achieved
(96)
. A typical example where surface stimulation can be used is restoration of hand
function. Surface FES can be applied at a very early stage of the
rehabilitation, during the recovery and reorganization period of the central and
peripheral nervous systems (plasticity), allowing early benefit for the patient.
FES training during recovery may help a subject to restore a function to the
point that he/she no longer needs a neuroprosthesis.

Popovic et al. advocate an early
start of FES as part of the rehabilitation in stroke patients. This benefits the
use of surface FES systems since the objective of the treatment is to help
patients to relearn the grasping task rather than to provide them with a
permanent assisting system (91) .
In SCI subjects there are two possibilities: implantation and surface
stimulation. For the implantation, the patient should have reached stable
neurological status.Surface FES is also preferred in treatment of spasticity in
order to strengthen the antagonistic muscles and in recovery of simple
movements, which requires stimulation by few electrodes with easy positioning
(4)
.

A disadvantage is that only a
small part of a muscle can be stimulated, causing a rapid development of muscle
fatigue. Furthermore, the reproducibility of muscle force is insufficient due to
changes in the muscle geometry during contraction. Another issue is the
displacement of the electrodes, which may occur during use of the FES device
(96) .

The advantages of the implantable
FES systems are that the stimulation does not depend on the geometry of the
muscle and a single electrode can stimulate several muscles with low energy
consumption. Therefore, implantable stimulation is preferred in the following
cases (4)
:

When deep
muscles that are difficult to stimulate by surface electrodes have to be
recruited

For complex
movements which require a large number of electrodes in a limited
space

To avoid pain
and burns, caused by stimulation of the pain receptors and the
skin.

The disadvantages are that the
existing implantable FES systems still do not provide enough selectivity of
stimulation, the surgical trauma associated with the implantation and some
unresolved issues in the geometry of the electrodes (97)
.

In conclusion: Implantation of a
FES device induces acute, late and chronic inflammatory response. The least
effect is caused by the chemical reaction due to the inert device materials
used. The biggest damage comes from mechanical stress or is due to
surgery.

Contemporary
Clinical applications of FES

The cochlear
implant

The cochlear implant relies on the
assumption that there are enough auditory nerve fibers left for stimulation near
the electrodes. Although the first fully implantable cochlear prostheses
preceded the fully implantable heart pacemakers, in the 1970’s, heart pacemakers
achieved high reliability and improved quality of life of the patients. In
the USA, FDA nevertheless approved the first cochlear implant device for
clinical use not earlier than 1984. Over the past 20 years of clinical
experience more than 20.000 people worldwide have received cochlear implants.

Cochlear implantation has a
profound impact on hearing and speech perception in postlingually deafened
adults. Most individuals demonstrate significantly enhanced speech reading
capabilities, attaining scores of 90 -100% correct on everyday sentence
materials (98) .
Moreover, according to the NIH Consensus Statement the cochlear implant is the
first, and still the only, neural prosthesis that is aiding a significant
portion of a disabled population. In recapitulation - while the cochlear implant
has reached wide acceptance and maturity as device, most of the other
neuroprosthetic applications are still in childhood.

Drop-foot
stimulator

The first neuroprosthesis for
walking was developed in 1961 by Liberson and colleagues (15) . This system was
developed to compensate for the “drop-foot” problem. The drop-foot is a
pathological condition, caused by diminished ability to use the muscles that
lift the foot. It may be caused by stroke, cerebral palsy, multiple sclerosis,
or neurological trauma. By stimulating the peroneal nerve, the prosthesis
elicited ankle dorsiflexion, eversion and/or inversion thus allowing the subject
to make a step with the disabled leg. Since then a multitude of devices has been
developed, mostly in former Yugoslavia. The Fepa system was proposed by Vodovnik
and colleagues (1978) (99) . Currently there
are several commercial systems all based on surface stimulation - MikroFES
(Jozef Stefan Institute of Science, Slovenia), Odstock 2 (Salisbury District
Hospital, UK) (100)
and WalkAide (Neuromotion, Canada) (101) .
The drop-foot stimulators are commonly controlled by a foot switch. A recent
overview on the different FES systems can be found in Popovic et al. (2001a) (97)
.

So far, most
systems are external with a surface electrode over the peroneal nerve just below
the head of the fibula.

More recently,
radio-frequency transmitter enabled implants have been developed: for example,
the new implantable two-channel system (Finetech, UK) (102)
with subepineural electrodes in the n. peroneus profundus and n. peroneus
superficalis. It has been tested in 10 patients. This dual channel implant will
soon be registered with CE-mark. It has the ability to stimulate independently
the deep and superficial branch of the peroneal nerve, thus allowing the
correction for excessive eversion and inversion.

Odstock 2 is a surface
stimulation system that has been used mostly in the clinical environment. It is
based on the stimulator proposed by Liberson. The device can be used as an
assisting aid or as a training device to strengthen muscles and improve
voluntary control. Additionally, the device has a role in physiotherapy in gait
re-education allowing isolated components of the gait cycle to be practiced
under therapist control. (103,
104)
The Odstock 2 was perceived by the users to be of considerable benefit. A
comprehensive clinical follow-up service is essential to achieve the maximum
continuing benefit from FES-based orthoses. A recent review on the subject can
be found in Lyons et al. (2002) (105)
and (97).
Odstock 2 and the MikroFES, have been fitted to more than 500 subjects. Thus
far, only the WalkAid has been FDA approved.

Figure 6. The peroneal stimulator (Finetech, UK) (see
text) Courtesy: J. Holsheimer, G. Bulstra and H.E. van der Aa

Restoration
of Lower limb Function

In 1963, Kantrowitz reported the
first application of FES to a T-3 paraplegic patient who achieved standing by
surface stimulation of the gluteus and the quadriceps for brief periods. The
first systems were external devices enabling walking and were developed by Bajd
and Kralj in the 1970s through 1980s in Lublijana, Slovenia (106) .
The initial devices only used two surface stimulation channels for each leg to
produce standing and walking. The stimulation was delivered to both knee
extensors. Devices with 4 channels have been proposed later (107) .
Systems used by patients included stationary bikes for exercise, transfer
systems to enable the patient to move between a wheelchair and a bed, transfer
systems between standing and sitting, and walking aids (106) .
The targeted patient group was paraplegics after SCI. Patients used a rolling
walker frame or crutches for support. About 50 patients successfully used the
system (108) .
One of the first implantable multichannel stimulators was developed by Brindley
with electrodes applied on the femoral and the gluteal nerves (109) .
An overview of the contemporary devices can be found in (97) .
In brief they are Parastep, LARSI (110) ;
FESmate, HAS, RGO, Praxis24 and the implanted FES system proposed in (111)
.

The only commercialized system,
developed initially by Graupe and colleagues (112,
113) ,
is Parastep (Sigmedics Inc., USA). It is based on the earlier work of Kralj (106) .
The target patient group consists of the complete or almost complete T4-T12 SCI
patients. The system uses surface electrodes and no orthotic aids. The Parastep
system was applied to more than 400 subjects and was the first FDA-approved FES
system.

Although FES of locomotion gains
clinical acceptance, it is still in its experimental phase of development.
At this stage, FES of locomotion is so far incapable to deliver a complete
treatment for the target groups of patients. Only in several cases, it may
provide pathogenetic treatment. The main impediment of its development is the
insufficient selectivity of stimulation and the small number of applications in
patients, which hinders assessment of any beneficial clinical
outcome.

Hand
disabilities – reaching and grasping

In tetraplegic and stroke
patients, the most important target for achieving a high level of independence
in active daily life is the restoration of the hand function. Thus, the main
objective in applying FES in such patients is to improve the hand function by
creating a reliable and long lasting power grasp, or a smooth pulp-pinch grasp
that is needed to manipulate small objects. Neuroprostheses for grasping are
used to restore or improve grasping function in tetraplegic and stroke subjects.
The patients that would benefit most from such a neuroprosthesis would be C5-7
quadriplegics. The available neuroprostheses for grasping enable the restoration
of the two most frequently used grasping styles, the palmar and the lateral
grasp (114) .
The best known grasping neuroprostheses are the Freehand system (115) ,
Handmaster NMS (116) ,
Bionic Glove (117) ,
NEC FESMate system (118) ,
ETHZ-ParaCare neuroprosthesis (114) ,
the systems developed by Rebersek and Vodovnik  (119)
and the Belgrade Grasping system (120) .
A recent overview can be found in (91)
.

In 1997, FDA approved for clinical
use the Freehand system (NeuroControl Corp., Cleveland) for restoration
of the hand grasp in quadriplegics. The research and development process took 25
years, mostly carried out by the researchers at the Case Western Reserve
University. The device consists of a stimulator and electrodes implanted in
wrist and forearm muscles, a "joystick" controller implanted in the opposite
shoulder, and an external processing unit (115) .
The joystick-like sensor placed on the chest relaxes and tightens the hand as
the shoulder moves back and forth. A detailed description of the prosthesis can
be found in (121) .
This prosthesis has been experimentally implanted in more than 130 people. With
training, most patients with this device can open and close their hand in two
different grasping movements and lock the grasp in place by moving their
shoulder in different ways. A recent small clinical study with 6 patients (122)
is a sign of increasing acceptance and application of the Freehand system
outside USA. Researchers reported that all subjects were able to grasp, move and
release more objects within the 30-s test period with the neuroprosthesis than
without it. In 85% of the occasions, the six subjects expressed a preference for
using the neuroprosthesis to perform these activities in daily living. Twelve
months after rehabilitation, five of the six subjects still used the
neuroprosthesis regularly. One of the main advantages of the Freehand system is
that the time needed to put on (donning) and to take of (doffing) the system is
significantly shorter compared to most surface stimulation FES systems. On the
other hand, the Freehand system can be applied only 18-24 months after the
injury and is only suitable for SCI subjects and not individuals suffering from
stroke (114)
.

The Bionic Glove is a
hybrid system that utilizes a glove with FES electrodes. The detection of wrist
extension stimulates finger flexion; therefore, the control of the device is
directly related to the normal grasping sequence. Only patients who have
sufficient wrist extension strength can use this type of device. The controls of
the Freehand system and the Handmaster are not related to the normal movements
included in the motor function, whereas the Bionic Glove uses mechanical sensors
to detect wrist extension and the control is therefore directly coordinated by
the grasping sequence. Bionic Glove can significantly improve independence in
patients with C5-C7 spinal cord injury if their initial Functional Independence
Measure and Quadriplegia Index of Function scores are 20% to 50% of the maximum
values. The study is part of a multicenter clinical trial (123)
.

The Handmaster is an
orthosis for grasping with three pairs of surface stimulation electrodes. This
system can be used to generate a grasping function in tetraplegic and stroke
patients. Originally, this system was envisioned as an exercise and
rehabilitation tool, but it is also used as a permanent prosthetic system.

One of the
advantages of the Handmaster is that it is easy to put on and to take off. The
Handmaster is predominately used as an exercise tool for stroke subjects and is
commercially available in a limited number of countries.

An interesting attempt is the
MeCFES system (myoelectrical controlled functional electrical stimulator)
where the residual myoelectric signals from the paretic wrist extensor (m.
extensor carpi radialis) are used to control stimulation of either the wrist
extension (i.e., the same muscle) or thumb flexion. Initial results of six
spinal cord lesion patients and one stroke patient show improvement of the wrist
extension (124)
.

With the exception of the Freehand
and NEC-FES systems, all other neuroprostheses for grasping are FES systems with
surface stimulation technology. Only the Freehand and Handmaster systems are
currently available on the market while the other neuroprostheses are produced
in laboratory environments.

All of the discussed
neuroprostheses for grasping have demonstrated, in clinical trials in stroke
and/or SCI subjects, improvement of the grasping function. These systems
confirmed that the FES technology in principle could facilitate comfortable and
secure grasp. However, the grasp strategies that can be provided with the
existing neuroprostheses for grasping are very limited and can only be used for
a restricted set of grasping and holding tasks.

In conclusion, it is too early to
consider any of the existing systems, including Freehand and Handmaster, as
success because of the small number of patients using the systems (about 150 for
Freehand and a comparable amount for Handmaster).

Phrenic
stimulation

Of those people who sustain a high
cervical spinal cord injury, a substantial number initially require some form of
mechanical ventilation of their lungs. Most of them will be able to breathe
spontaneously with recovery, but 25% will remain dependent upon some form of
ventilation support for the remainder of their lives. Most of these injuries
occur in young, otherwise healthy people with a life expectancy of 20 years or
more, assuming they have access to appropriate medical care. In patients with
cervical lesions of the spinal cord, direct stimulation of the phrenic nerve can
be applied for respiratory pacing. This is one of the earliest clinical
applications of FES.

Diaphragm pacing is useful for
conditions in which the brain stem respiratory centers provide little or no
activation of the respiratory muscles, i.e. central hypoventilation syndrome,
Arnold-Chiari malformation/brain stem dysfunction and high quadriplegia.
Suitable patients are those with an intact phrenic nerve motor neuron pool in
their cervical spinal cord on both sides. This can be checked by measuring the
phrenic nerve conduction velocity and the diaphragm electromyogram (EMG). To do
this phrenic nerves are stimulated by needle or surface electrodes in the neck
and the diaphragm muscle EMG is recorded from electrodes placed low on the chest
at the front. The patient wears an external radio-frequency (RF) transmitter
over an implanted receiver, and a stimulating current is induced without the
need for any percutaneous wires. The introduction of phrenic nerve pacing more
than two decades ago by Dr William Glenn and associates at Yale
University has provided many ventilator-dependent tetraplegic
patients with freedom from mechanical ventilation (9, 125,
126)
.

More than 1.200 phrenic nerve
stimulator implantations have been performed throughout the world since 1968.
Patients from several months of age to over 80 years of age have been
successfully implanted and paced for long periods. Many patients have been
successfully helped for more than 10 years; the longest period of pacing in a
patient has been more than 20 years. In patients with ventilator-dependent
quadriplegia, phrenic nerve pacing provides significant clinical advantages
compared with mechanical ventilation. This technique, however, generally
requires a thoracotomy with its associated risks; in-patient hospital stay and
it carries some risk of phrenic nerve injury.

During the past decade, diaphragm
pacing has also been attempted in small infants. The appropriate stimulation
parameters are low stimulus frequency, short inspiration time and moderate
respiratory rate. In a clinical trial in 33 pediatric patients, the mean time to
failure was 56 months, which is acceptable for limited application of the
pacemakers (127) .
Among the benefits of the phrenic nerve pacing can be outlined:

Increased
mobility of the patients

Improved
speech

Improved sense
of well-being and reduced anxiety due to elimination of fear of ventilator
disconnection, elimination of ventilator tubing, elimination of ventilator
noise, more physiological breathing and tracheostomy closure in some
patients

Reduced
incidence of respiratory tract infections

To recapitulate: one may
consider the phrenic stimulator as a major improvement of the quality of life of
the patient, which also reduces the overall costs for nursing of the
patient.

Clinical
trials / Evidence based medicine for FES

Efficacy of
FES after stroke

Meta analysis showed that in
post-stroke hemiparetic patients FES promotes recovery of muscle strength if
included as part of the rehabilitation (based on the paretic muscle force
measurements). The study was based on the results of clinical trials dating
between 1978 and 1992 (128) .
There is, however, more evidence for improvement of the motor control of the
upper extremity after stroke (129) . The meta
analysis was based on six randomized controlled trials on the therapeutic
electrical stimulation on motor control and functional abilities. After stroke,
up to 81% of the individuals develop shoulder subluxation, a condition
frequently associated with poor upper limb function. FES on the shoulder muscles
has been used in treatment of this condition. The results of seven (four early
and three late) trials indicate when added to conventional therapy FES is
beneficial early after stroke for prevention of shoulder subluxation (130).
Burridge et al. in a randomized controlled trial in subacute single
stroke patients measured the effect of the Odstock Dropped Foot Stimulator as a
supplement to physiotherapy (100). The results of
32 subjects showed increase in walking speed and beneficial Physiological Cost
Index, if the stimulator had been used.

Efficacy of
FES in urological conditions

Stimulation of the sacral spinal
nerves or the sacral spinal roots or the pelvic nerves can be used to restore
bladder function in patients with voiding disorders. Several therapeutic
techniques have become established in clinical urology as part of therapy at an
increasing number of specialized centers.

As part of the treatment of
different kinds of lower urinary tract dysfunctions, refractory to conservative
treatment, Tanagho and Schmidt introduced sacral nerve neuromodulation in
the 80s (131) .
Neuromodulation is now carried out by stimulation of the S3 sacral nerve. In
brief, an electrode is placed in S3 foramen. The electrode has to be brought
into the ventral side of the opening in front of the S3 ventral ramus. This
procedure is minimally invasive, as compared to sacral neurostimulation. The
patient is sent home with an external pulse generator for a few days as part of
an evaluation test. Responders are then implanted with a permanent sacral
foramen implant and an implantable pulse generator (132)
.

Since its early applications,
neuromodulation has grown in popularity and the indications for this procedure
are multiplying. Among them are now urge incontinence and sensory urgency,
idiopathic chronic urinary retention, pelvic pain and interstitial cystitis
(133).
Complications in general are minimal and included electrode migration, electrode
failure and pain at the implantable pulse generator site (134).

The sacral neurostimulation
approach proposed in the 1980s by Brindley (135)
consists of intradural implantation of a so-called book electrode to stimulate
S2-S3 ventral roots, together with posterior S2-S4 rhizotomy. The primary
purpose of the Brindley bladder stimulator is to improve bladder emptying,
thereby eliminating urinary infection and preserving kidney function in
suprasacral SCI patients. It also assists in defecation and
enables male patients to have a sustained full erection. Nowadays the Brindley
bladder stimulator is being developed by the company Finetech in UK.

Detrusor
sphincter dysynergia

The suprasacral SCI patients
develop detrusor sphincter dyssinergia, which imposes serious health risk. The
results from the first 500 patients implanted with the Brindley stimulator show
usage in 424 after a mean follow up of 4 years. (136) .
Another study reported 38 patients with a complete spinal cord lesion (137) .
During the follow-up period, ranging up to 12 years all patients had increased
bladder capacity and reduced residual urine volumes; 31 patients were continent;
29 males could achieve a sustained full erection; 27 patients used the implant
for bowel function. The long long-term favorable effects of the stimulator were
also confirmed in follow-up studies on urodynamics (138)
and cost-effectiveness (139)
in 52 patients.

Overactive
bladder

The overactive bladder comprises a
spectrum of conditions ranging from urgency-frequency syndrome to urge
incontinence. The incidence of the overactive bladder increases with age, with a
prevalence exceeding 4% to 5.5% of the population. The efficacy of the
neuromodulation therapy was evaluated in a prospective 12-center study in
Europe, Canada, and the United States conducted form 1993 to 1999 with the
InterStim system (Medtronic, Inc., Minneapolis, Minnesota). The study enrolled
184 patients for urinary urge incontinence, 220 for urgency-frequency, and 177
for retention for a total of 581 patients (140) .

Urge Incontinence and Sensory
Instability

Neuromodulation is used in
patients with urge incontinence and with therapy-resistant idiopathic detrusor
instability, where in a follow-up trial results showed clinically significant
improvement of the quality of life (133) .
In patients with detrusor hyperactivity implantable neuroprosthetic devices also
lead to improvement in the urodynamic parameters (141) .
Out of 20 patients with urge urinary incontinence, presented by Thon and
colleagues (2003) (142) ,
17 showed an improvement of more than 50% compared to the baseline, which
persisted for more than a year of follow-up. Elabbady and colleagues (1994) (143)
presented their results in 9 patients with urgency frequency and/or urge
incontinence: frequency improved by 73%, urgency by 42% and incontinence by 50%.
However, the number of patients is small to derive statistically valid
conclusions for the success rate of the operation.

On the other hand, the outcome of
the multi-center trial is promising. Results demonstrate that after 3 years, 59%
of 41 urinary urge incontinent patients showed greater than 50% reduction in
leaking episodes per day with 46% of patients being completely dry. After 2
years, 56% of the urgency-frequency patients showed greater than 50% reduction
in voids per day (140)
.

Idiopathic Non-Obstructive
Chronic Urinary Retention

The results for the idiopathic
urinary retention of the multi-center trial were also separately reported in (144) .
A total of 177 patients with urinary retention refractory to standard therapy
were enrolled in the study. Compared to the control group, patients implanted
with the InterStim system had statistically and clinically significant
reductions in the catheter volume per catheterization. Successful results were
achieved in 83% of the implant group with retention compared to 9% of the
control group at 6 months. Temporary inactivation of sacral nerve stimulation
therapy resulted in a significant increase in residual volumes but effectiveness
of sacral nerve stimulation was sustained through 18 months after
implant. In another study, Thon and colleagues (2003) (142)
reported that from 33 patients with chronic urinary retention implanted
permanently with neuroprosthesis 23 showed a long-lasting significant
improvement, but in the remaining 10 the improvement did not reach 50% compared
to baseline. In 7 patients with chronic retention, Vapnek and Schmidt (2003)
reported for success in 5 cases (145)
and Elabbady and colleagues (1994) (143)
presented success in 8 of 8 cases. Results of all the presented studies
demonstrate that sacral nerve stimulation is effective for restoring voiding in
patients with retention who are refractory to other forms of
treatment.

Pelvic pain and
discomfort

Pelvic pain or discomfort is a
very common symptom associated with other storage or voiding dysfunctions. In
the available literature on sacral root neuromodulation, associated pelvic pain
has improved from 85% to 90% when postimplant status was compared to baseline
(134,
142).

In conclusion: The correlation
between the clinical outcome and the urodynamic test results is poor. The
results of the presented studies on neuromodulation show that 40% of the
selected patients do not qualify for the procedure and the effect of treatment
is not enduring: voiding dysfunctions return as soon as the neuroprosthesis is
switched off. It is thought that the effect is based on antidromic
stimulation of the inhibitory neurons in the spinal cord. Due to the fact that
the stimulators are voltage-controlled so far, the amount of injected current
into the tissues is unknown and the fibers which are stimulated are also not
known (may be both myelinated or non-myelinated) and also the direction of
stimulation (ortho- antidromic). More basic and clinical research needs to be
performed before this treatment can be introduced as a routine procedure in
patients with serious voiding dysfunction refractory to conservative measures
(146)
.

Experimental and theoretical
research

Experimental
research

Visual
prosthesis: blind eye vs. normal brain

Blindness can
result when any step of the optical pathway—the optics, the retina, the optic
nerve, visual cortex, or other cortical areas involved in the processing of
vision — sustains damage. In Germany, 17.000 patients become blind every year
for whom there is no effective treatment or cure (1).

Public discussion of electricity's
effect on vision dates to 1751, when it was addressed by Benjamin Franklin
following his celebrated kite-and-key experiment. Despite some advocates, the
idea of treating blindness through electrical stimulation did not catch on. In
the last 30 years extensive experimental research for development of visual
prostheses has been performed. In 1967-68 the experiments of Brindley and Lewin
(14)
showed the feasibility of the long-term interface with the visual system. The
approach of the cortical prostheses has its foundations in the observations of
the visuotopic organization of electrically evoked phosphenes in the occipital
cortex. This have led a number of investigators to propose that electrical
stimulation of visual cortex via arrays of electrodes might provide the
profoundly blind with a limited form of functional vision. Several groups
investigated the stimulation with cortical surface electrodes – the so called
“cortical prostheses” of Brindley, Dobelle (1974) (147) ;
cortical penetrating electrodes of Bartlett and Doty (1980) (148)
and of Schmidt and Hambrecht (1996) (149) .
However influential it is, so far this idea has very little practical progress.
The neurons in the cortex represent textures, depth (displacements), angles and
brightness /colors of the object. The perception of images is not based on
pixelation; therefore, all the bitmap-based stimulation approaches are doomed to
fail. On the other hand, the cortical stimulation approach may provide the only
therapeutic approach for individuals with non-functional retinas and/or optic
nerves.

Recently Dobelle (2000) (150)
reported for the development of a visual prosthesis providing a sort of
"artificial vision" to a blind volunteer by connecting a digital video camera,
computer and associated electronics to the visual cortex of his brain. As an
alternative to the cortical stimulation, two other approaches have been also
investigated.

The idea to realize a visual
prosthesis by stimulating the optic nerve was conceived by Mortimer and
Veraart (1998) (151)
at the beginning of the 1990's. The first implantation was in 1998 and the
patient was able to localize single bright spots of light, but high spatial
resolution cannot be expected with such a stimulation arrangement. This
prosthesis may show promise if an ultra-high electrode count array of
intraneural electrodes are implanted in the optic nerve for providing better
selectivity of stimulation. Although ingenious, the approach does not take into
account so far the information processing steps taking places in the retinal
ganglionic cells.

At the end of the 1980's, an
entirely new approach was undertaken by the teams of M. Humayun, of John Hopkins
University (149)
and that of J. Rizzo of Harvard University (156) ,
in association with the Massachusetts Institute of Technology. They developed an
implantable electrode array for the retina itself, further referred to as
retinal implant, in order to stimulate the retinal ganglionic cells (152) ,
whose extensions form the optic nerve. This prosthesis is referred to as
"epiretinal implant". Another kind of retinal implant is the subretinal implant
developed by Chow and colleagues (1997) (153)
in Chicago and Zrenner in Tubingen (1997) (154)
.Both teams (155,
156),
have stimulated the retina of blind patients with epiretinal electrodes that
were transiently inserted into the eye through a scleral opening. Both groups
reported a sensation of light patterns by the patients, but perception of
geometric patterns was reported in only a few instances. Recently a permanent
implantation has been made in a blind volunteer (157)
.

All results in the field in
the past 30 years are still far from true object recognition. However, they do
demonstrate the feasibility of generating perception of light patterns in blind
people.

Spinal Cord
Stimulation in motor disorders

Apart from the pain suppression of
chronic intractable pain, peripheral vascular disease and angina pectoris,
spinal cord stimulation has been employed in motor disorders for control of
spasticity in SCI patients. Attempts in this direction have been carried out in
the 1980s for the first time by Richardson et al. (1979) (161) .
Recently Pinter et al. (2000) (162)
stimulated the L2-L3 dorsal roots of the spinal cord in SCI patients. Results
demonstrated reduction of the muscle hypertonia of the lower
extremities.

A possible future application of
SCS is to contribute to the restoration of the motor functions of the lower
extremity (158,
159) .
The reasoning behind this idea is that if we are able to control the Central
Pattern Generator (CPG) in the motor spinal cord we circumvent many of the
selectivity problems pointed out so far. There is accumulating evidence for the
existence of CPG of the locomotion in primates and men (160).
This requires a Multi Electrode Array to be placed on the spinal cord such that
it is able to stimulate the musculotopically-organized motor pools of the lower
limbs.

Theoretical
Modeling studies

As part of the design of different
types of electrodes for stimulation a substantial amount of theoretical and
experimental research on selectivity has been performed. The electrical behavior
of the myelinated nerve fiber can be represented by a simple cable network. The
first to introduce this model was McNeal (1976) (27) in
order to calculate how a stimulation-induced extracellular field affects nodal
transmembrane voltages. According to this model, the node of Ranvier, which is
closest to the cathode, will be excited first when the stimulus current is
sufficiently high. The results of this theoretical approach are in accordance
with the general observation that the threshold stimulus of a nerve fiber
excitation is smallest near the cathode and rises with the increase of the
distance (23) .

Research on
spatially-selective nerve stimulation

The feasibility of selective
activation of peripheral nerve fascicles was demonstrated by McNeal and Bowman
(1985) (31) ,
who found that with proper fit and positioning a single circumneural sleeve with
multiple electrode contacts could control selectively the activation of two
antagonist muscle groups innervated by a common nerve trunk. In their study,
Sweeny et al. (1990) (42)
performed numerical modeling and experimental testing of a nerve cuff technique
for selective stimulation of superficial peripheral nerve trunk regions. Two
basic electrode configurations ("snug" cuff monopolar and tripolar
longitudinally aligned dots) have been considered. Both modeling and
experimentation suggested that longitudinally aligned tripolar dot electrodes on
the surface of a nerve trunk would restrict excitation to superficial nerve
trunk regions more successfully than monopolar dot electrodes would do.
Transverse anodal "steering" improved the spatial selectivity of both monopolar
and tripolar electrode configurations (42)
.

With a sufficient number of
electrodes within a so-called “cuff electrode”, high selectivity of stimulation
can be achieved by either longitudinal or transversal currents produced by
appropriately switched electrodes (41) .

Goodall et al. (1996) (45)
found that transverse current from an anode positioned opposite the stimulating
cathode improved spatial selectivity, and position selectivity was enhanced when
the ratio of transverse current to longitudinal current was increased. Based on
computer modeling Deurloo et al. (2003) concluded that transverse steering
provides more selective stimulation than longitudinal steering (44)
.

Figure
7. Transverse section of the 3D volume conductor model of
a monofascicular nerve surrounded by an insulating cuff with an electrode
contact on its inner surface.

B Recruitment contours showing
excitation regions in the nerve trunk. c- cathode, a –anode; a: central
cathode b: longitudinal tripole + additional anode c: transverse
tripole.

Courtesy: K Deurloo, from (96) A
p. 15 and B p.30

Selectivity was dependent on the
relative location of the electrode contacts and the nerve fascicles, as well as
the size and relative spacing of neighboring fascicles (43) .

Deurloo et al. (1998) (163)
performed theoretical research on how to improve the selectivity performance of
multi-contact cuff electrodes for stimulation of peripheral nerves. As
combination of controlled fiber size and spatial selectivity seems hard to
achieve, the focus was primarily on spatial selectivity. They found that the
transverse tripole is the only configuration that maximizes activation
selectivity for a small (cylindrical) bundle of fibers in the periphery of a
monofascicular nerve trunk (163)
(see also Fig 7). Inverse recruitment was less pronounced than for the other
configurations. Therefore, they recommended transverse tripolar stimulation
since it did not change the shape of the recruitment contours, despite the
lowering of the excitation threshold, which might occur in chronic implantation
fibrosis on the implantation site.

However, in acute animal
experiments, where the recruitment characteristics of muscle selective nerve
stimulation by a multi-contact nerve cuff electrode was studied, the results
showed that only in a few cases transverse bi- and tripolar stimulation provided
a better selectivity than monopolar stimulation (164) .
In accordance with the results of the modeling studies, bi- and tripolar
stimulation required higher stimulus currents than monopolar stimulation,
whereas maximum recruitment and slopes of recruitment curves were lower. Due to
the variability in the number and size of the fascicles and their position in
this nerve, sufficient reproducibility for the selectivity could not be
obtained.

The theoretical research performed
by Rutten et al. (1991) (56) and
Yoshida, and Horch (1993) (49)
suggested that intraneural electrodes should give better spatial selectivity
than cuff electrodes. In this line of reasoning Rutten et al. (1999) (165)
argue that the best way to control individual fibers is to stimulate in close
proximity of the nodes of Ranvier, which implies use of intrafascicular
electrodes. However, the distribution of the nodes of Ranvier in the peripheral
nerves makes the stimulation a probabilistic process. The experimental studies
show that, despite the theoretical expectations, when the topography of the
motor nerve fibers is unknown, the implanted intraneural electrode may simply
miss the target fiber population. Several attempts for such stimulation have
been made so far with 1D arrays. The obtained information gave insight on how to
improve the design (shape, distance between the electrodes) of the later on
produced 3D arrays.

In acute nerve implantation
experiments eight 5- to 24-wire-MEAs were used in order to investigate whether
the electrodes could selectively stimulate single motor units. The results
revealed partial blocking of neural conduction, similar to that reported with
microneurographic insertion with single needles (166) .
The eight arrays were capable of evoking threshold forces selectively with an
average efficiency of 0.81 (or 81%)

Frieswijk et al. (1998) (167)
performed animal experiments and model simulations of monopolar, intrafascicular
nerve stimulation in order to study force-current relationships (recruitment
curves). They found that the conductivity of the extraneural medium is of prime
importance to the resulting recruitment curves: an insulating extraneural medium
generally leads to steeper curves with lower threshold currents than a
well-conducting extraneural medium. The statistical comparison of experimental
and model results suggested clustering of the motor fibers, originating from a
single muscle, within the same fascicle. This clustering manifested itself
mainly by an increased spread in threshold currents, as opposed to the situation
where the fibers are distributed uniformly throughout the entire
fascicle.

In conclusion, the insufficient
neuroanatomical knowledge of the spatial organization and topography of fibers
in the peripheral nerve hampers the definitive choice for the design of the
electrode.

Research on
fiber size selective stimulation

In myelinated axons, there is a
positive correlation between the internodal distance and the axonal diameter.
Thus, the large fibers (A, B) have nodes further apart than the small fibers. As
a result, in an electric field the larger fibers have a larger potential
difference between adjacent nodes. This leads to a lower threshold of the large
fibers in comparison with the small ones. Therefore, during stimulation, the
large myelinated fibers fire first. Unfortunately, this is the reverse of the
natural order of recruitment, and it results in fast development of muscle
fatigue. By activating/blocking stimuli delivered by multiple cuff electrodes
the order of recruitment can be reversed, but the fibers still can not be chosen
(168) .
Better fiber size selectivity can be achieved by intraneural stimulation, as
outlined by Rutten et al. (1999) (165)
.

The order in which the nerve
fibers in a peripheral nerve or a spinal root are excited by a stimulus pulse is
predominantly related to both the fiber diameter and the distance between the
fiber and the cathode (169) .
In the conducting media around the electrode, the current density is inversely
proportional to 2nd- 3rd power of the distance.
Accordingly, the threshold current is increased at the same rate as shown
empirically (23) .
In clinical applications, due to the generally limited amplitude range, only the
large fibers will be recruited. For example in spinal cord stimulation, the
maximal therapeutic amplitude should not exceed 170% of the paresthesia
threshold perception. Calculations show that the nerve fibers smaller than 9um
in diameter will not be recruited. Another relevant aspect of the recruitable
large fibers is their density within the fascicle (169) .

Fang and Mortimer (1991) (170,
171)
studied selective activation of small fibers without activating larger fibers in
the same nerve trunk. In the proposed nerve stimulation system,
quasitrapezoidal-shaped current pulses were delivered through a tripolar cuff
electrode to effect differential block by membrane hyperpolarization. The
quasitrapezoidal-shaped pulses with a square leading edge, a 350 microsecond(s)
plateau, and an exponential trailing phase ensured the block of propagating
action potentials and prevented the occurrence of anodal break excitation. The
tripolar cuff electrode restricted current flow inside the cuff and thus
eliminated the undesired nerve stimulation due to a "virtual cathode." The
subsequent animal experiments confirmed that larger alpha motor axons could be
blocked at lower current levels than smaller alpha motor axons, and that all
alpha fibers could be blocked at lower current levels than gamma fibers, and the
blocking threshold correlated with the fiber diameter Fang and Mortimer (1991 b)
(171).
For the same purpose, van Bolhuis et al. (2001) (172)
suggested the use of two pulse generators independently supplying short
supramaximal cathodal stimulating pulses (0.5 ms) and long subthreshold cathodal
inactivating pulses (1.5 s) to the sciatic nerve. Results showed that
propagation of action potentials was selectively blocked in nerve fibers of
different diameters by adjusting the strength of the inactivating current (172)
.

It has been shown that cathodal
pre-pulses inverse the recruitment order

(173,
174) .
Grill and Mortimer (1997) (173)
demonstrated that subthreshold membrane depolarization generated a transient
decrease in neural excitability and thus an increase in the threshold for
stimulation by a subsequent stimulus pulse. When a depolarizing stimulus pulse
was applied immediately after the subthreshold depolarization, nerve fibers far
from the electrode could be stimulated without stimulating fibers close to the
electrode. Thus subthreshold depolarizing pre-pulses allowed selective
stimulation of nerve fibers far from the electrode

In a realistic model of a nerve
fiber surrounded by a cuff electrode Deurloo et al. (2001) (174)
showed that it is also possible to stimulate small fibers without exciting large
ones. The applied model requires that, in the case of monopolar stimulation with
a cuff electrode, the cuff length should not exceed twice the internodal length
of the fibers to be blocked. Similarly, the distance between cathode and anodes
should not exceed the internodal length of these fibers when tripolar
stimulation is used (174) .

Conclusions

The principal requirements to any
substituting structure are features mimicking some of the biological functions
of nerves and replacing these functions depending on the scope of implantation.
Profound knowledge about the development, functional and structural organization
of the nervous system is a prerequisite for any attempt to establish meaningful
recording or stimulation in order to substitute for a given function or
modality.

A number of proposed ideas for
electronic implants are based, however, on a more "engineer-like" way of
thinking rather than considering the complexity of the biological systems. The
necessary for the engineering development reductionist approach often
substitutes basic problems for secondary ones. For example, the concept of
spatial selectivity in stimulation is often substituted for fascicle
selectivity, which equalizes the fiber representation of a particular muscle
contraction (dynamic view) to a coincidentally delimited anatomical region along
the course of a nerve (static view).

The presented overview of the FES
applications shows that despite the substantial amount of research performed in
the field which led to successful development of cardiac and phrenic
pacemakers and cochlear prostheses, a number of fundamental
scientific problem still remains to be solved before we see a comparable degree
of effectiveness and penetration in common medical practice for the locomotion
neuroprostheses and urological FES appliances.

Although FES of locomotion
gains clinical acceptance, it is still in its experimental phase of
development.  There are several interrelated, but different,
issues

-
Effectiveness and reliability of the devices

-
Technical support

-
The place of the devices in the complex treatments of the CVA, SCI, multiple
sclerosis etc.

Despite the significant technical
progress achieved in the last 10 to 15 years in the FES field, there is a
consensus that these systems are not sufficiently advanced and that they need
further development. Complexity of human motion greatly diminishes the area of
the animal experimentation in this field and puts the stress on computer
modeling. This complexity, however, makes the leap from a model to a patient
even more difficult than from an animal model to human, which one could see in
the performance on many implantable FES systems. This has only been done in
limited number of cases, which in turn diminishes the validity of the obtained
results and leads to case descriptions rather than generalizations. This in turn
increases the interval between trial and error (the most prominent approach in
experimentation) and as a result slows down the research and development process
for new applications.

Nevertheless, the present FES
treatments combined with conventional occupational and physical therapy still
remain the most promising approach in rehabilitating SCI patients and stroke
patients. The need of training, which is seen by some as a shortcoming of FES
(97)
, in our opinion may be turned into advantage if it is integrated in the overall
treatment and rehabilitation process.

The mechanism of action of
electrical neuromodulation is still poorly understood, which leaves it in a
rather empirical state of operation. The Brindley neurostimulation method is a
real FES application. Until now hundreds of patients have been treated
successfully with the Brindley approach in bladder voiding disorders. However,
the posterior rhizotomy, which is required, is a rather crude approach towards
blocking the spinal reflexes. Nevertheless, FES is effective in improving the
bladder function in the overactive bladder states and in SCI patients it gives
long-term favorable results, which makes it a viable therapeutical option for
the indicated groups of patients.

The research performed in the
field of the visual prostheses showed the complexity of the information
processing by the (human) visual system. Even the modest goals of the projects
so far are ambitious compared with the information content of the reported
visual phenomena. We would like to finish with a quotation from Dobelle (150) ,
which applies in the end to any prosthetic treatment: “Development of
implanted medical devices such as this artificial vision system progresses in
three stages. First there is speculation, then there is hope, and finally there
is promise.”

References

1 .  Moe JH,
Post HW. Functional electrical stimulation for ambulation in hemiplegia. J
Lancet 1962; 82: 285-288.

2.  Dutrochet
RJH. Recherches anatomiques et physiologiques sur la structure intime des
animaux et vegétax et sur leur mobilité. Paris: Bailliére, 1824.

3.  Remak
R. Observationes anatomicae et microscopiscae de systematis nervosi
structura. Berolini: Reimerans, 2003.

4.  Rattay
F. Electrical nerve stimulation: theory, experiments and applications.
Wien: Springer-Verlag , 1990.

5.  Gold GR.
Cardiac pacing-from then to now. Medical Intsrumentation 1984; 18:
15-21.

6.  Greatbatch
W. Origins of the implantable cardiac pacemaker. J Cardiovasc.Nurs. 1991;
5: 80-85.

7.  Chardack
WM, Gage AA, Greatbatch W. Experimental observations and clinical experiences
with the correction of complete heart block by an implantable, self-contained
pacemaker. Trans.Am.Soc.Artif.Intern.Organs 1961; 7: 286-295.

8.  Sarnoff SJ,
Gaensler EA, Maloney JV. Electrophrenic respiration: the effectiveness of
contralateral ventialtion during activity of one phrenic nerve.
J.Thoracic.Surg. 1950; 19: 929.

9.  Glenn WW,
Hageman JH, Mauro A, Eisenberg L, Flanagan S, Harvard BM. Electrical stimulation
of excitable tissue by radiofrequency transmission. Ann Surg 1964; 160:
338-350.

10.
Clark GM, Tong YC. Cochlear Prostheses. Melbourne: Churchill Livingstone,
1990.

11.
Djourno A, Eyries C. Prothèse auditive par excitation électrique a distance du
nerf sensoriel a l'aide d'un bodinage inclus a demeure. Presse Med 1957;
35: 14-17.

12.
Foerster O. Beitrage zur Pathophysiologie der Sehbahn und der Sehsphare.
J.Psychol.Neurol., Lpz 1929; 39 : 463-485.

13.
Krieg W. Functional Neroanatomy. New York: Blakiston, 1953: 207-8.

14.
Brindley GS, Lewin WS. The visual sensations produced by electrical stimulation
of the medial occipital cortex. J Physiol 1968; 194: 54-5P.

15.
Liberson WT, Holmquest HJ, Scot D, DOW M. Functional electrotherapy: stimulation
of the peroneal nerve synchronized with the swing phase of the gait of
hemiplegic patients. Arch.Phys.Med.Rehabil. 1961; 42: 101-105.

Hambrecht FT. The History of Neural Stimulation and Its Relevance to Future
Neural Prostheses. In: Agnew WF, McCreery DB, editors. Neural Prostheses:
Fundamental Studies. Englewood Cliffs: Prentice Hall, 1990: 1-25.

17.
Robblee LS, Rose TL. Electrochemical guidelines for selection of protocols and
electrode materials for neural stimulation. In: Agnew WF, McCreery DB, editors.
Neural Prostheses: Fundamental studies. Englewood Cliffs: Prentice Hall,
1990: 26-66.

18.
Bhadra N, Peckham PH. Peripheral nerve stimulation for restoration of motor
function. J Clin.Neurophysiol. 1997; 14: 378-393.

.
Morton SL, Daroux ML, Mortimer JT. The role of oxygen reduction in electrical
stimulation of neural tissue . J.Electrochem.Soc. 1994; 141 :
122-130.

20.
Lilly JC, Hughes JR. Brief, noninjurious electric waveform for stimulation of
the brain. Scince 1955; 121: 468-469.

21.
McCreery DB, Agnew WF, Yuen TG, Bullara LA. Damage in peripheral nerve from
continuous electrical stimulation: comparison of two stimulus waveforms. Med
Biol.Eng Comput. 1992; 30: 109-114.

22.
McCreery DB, Agnew WF, Yuen TG, Bullara LA. Relationship between stimulus
amplitude, stimulus frequency and neural damage during electrical stimulation of
sciatic nerve of cat. Med Biol.Eng Comput. 1995; 33: 426-429.

23.
Ranck JB, Jr. Which elements are excited in electrical stimulation of mammalian
central nervous system: a review. Brain Res. 1975; 98:
417-440.

24.   Wee
AS, Leis AA, Kuhn AR, Gilbert RW. Anodal block: can this occur during
routine nerve conduction studies? Electromyogr.Clin.Neurophysiol. 2000;
40: 387-391.

25.   Wee
AS. Anodal excitation of intact peripheral nerves in humans.
Electromyogr.Clin.Neurophysiol. 2001; 41: 71-77.

26.
Hodgkin AL, Huxley AF. A quantitative description of membrane current and its
application to conduction and excitation in nerve. J Physiol 1952; 117:
500-544.

27.
McNeal DR. Analysis of a model for excitation of myelinated nerve. IEEE
Trans.Biomed.Eng 1976; 23: 329-337.

28.
Talonen PP, Baer GA, Hakkinen V, Ojala JK. Neurophysiological and technical
considerations for the design of an implantable phrenic nerve stimulator. Med
Biol.Eng Comput. 1990; 28: 31-37.

29.
Heiduschka P, Thanos S. Implantable bioelectric interfaces for lost nerve
functions. Prog.Neurobiol. 1998; 55: 433-461.

30.
Rutten WL. Selective electrical interfaces with the nervous system.
Annu.Rev.Biomed.Eng 2002; 4: 407-452.

31.
McNeal DR, Bowman BR. Selective activation of muscles using peripheral nerve
electrodes. Med Biol.Eng Comput. 1985; 23: 249-253.

32.
Marsolais EB, Kobetic R. Functional walking in paralyzed patients by means of
electrical stimulation. Clin.Orthop. 1983; 30-36.

33.
Wuolle KS, Van Doren CL, Thrope GB, Keith MW, Peckham PH. Development of a
quantitative hand grasp and release test for patients with tetraplegia using a
hand neuroprosthesis. J Hand Surg.[Am.] 1994; 19: 209-218.

34.
Marsolais EB, Scheiner A, Miller PC, Kobetic R, Daly JJ. Augmentation of
transfers for a quadriplegic patient using an implanted FNS system. Case report.
Paraplegia 1994; 32: 573-579.

35.
Loeb GE, Peck RA, Moore WH, Hood K. BION system for distributed neural
prosthetic interfaces. Med Eng Phys. 2001; 23: 9-18.

36.
Cameron T, Loeb GE, Peck RA, Schulman JH, Strojnik P, Troyk PR. Micromodular
implants to provide electrical stimulation of paralyzed muscles and limbs.
IEEE Trans.Biomed.Eng 1997; 44: 781-790.

37.
Naples GG, Mortimer JT, Yuen TGH. Overview of peripheral nerve electrode design
and implantation. In: Agnew WF, McCreery DB, editors. Neural Prostheses:
Fundamental studies. Englewood Cliffs: Prentice Hall, 1990:
107-45.

38.
Naples GG, Mortimer JT, Scheiner A, Sweeney JD. A spiral nerve cuff electrode
for peripheral nerve stimulation. IEEE Trans.Biomed.Eng 1988; 35:
905-916.

39.   Kim
JH, Manuelidis EE, Glenn WW, Fukuda Y, Cole DS, Hogan JF . Light and
electron microscopic studies of phrenic nerves after long-term electrical
stimulation. J Neurosurg. 1983; 58: 84-91.

40.
Klepinski, J. R. Implantable neural electrode. 1994. USA.

41.
Veraart C, Grill WM, Mortimer JT. Selective control of muscle activation with a
multipolar nerve cuff electrode. IEEE Trans.Biomed.Eng 1993; 40:
640-653.

42.
Sweeney JD, Ksienski DA, Mortimer JT. A nerve cuff technique for selective
excitation of peripheral nerve trunk regions. IEEE Trans.Biomed.Eng
1990; 37: 706-715.

43.
Grill WM, Jr., Mortimer JT. Quantification of recruitment properties of multiple
contact cuff electrodes. IEEE Trans.Rehabil.Eng 1996; 4:
49-62.

44.
Deurloo K, Holsheimer J. Fascicular Selectivity in Transverse Stimulation with a
Nerve Cuff Electrode: A Theoretical Approach. Neuromodulation 2003; 6:
258-269.

45.
Goodall EV, de Breij JF, Holsheimer J. Position-selective activation of
peripheral nerve fibers with a cuff electrode. IEEE Trans.Biomed.Eng
1996; 43: 851-856.

46.
Loeb GE, Peck RA. Cuff electrodes for chronic stimulation and recording of
peripheral nerve activity. J Neurosci.Methods 1996; 64:
95-103.

47.
Tarver WB, George RE, Maschino SE, Holder LK, Wernicke JF. Clinical experience
with a helical bipolar stimulating lead. Pacing And Clinical
Electrophysiology: PACE 1992; 15: 1545-1556.

48.
Nannini N, Horch K. Muscle recruitment with intrafascicular electrodes. IEEE
Trans.Biomed.Eng 1991; 38: 769-776.

49.
Yoshida K, Horch K. Selective stimulation of peripheral nerve fibers using dual
intrafascicular electrodes. IEEE Trans.Biomed.Eng 1993; 40:
492-494.

50.
BeMent SL, Wise KD, Anderson DJ, Najafi K, Drake KL. Solid-state electrodes for
multichannel multiplexed intracortical neuronal recording. IEEE
Trans.Biomed.Eng 1986; 33: 230-241.

51.
Rutten WL, Frieswijk TA, Smit JP, Rozijn TH, Meier JH. 3D neuro-electronic
interface devices for neuromuscular control: design studies and realisation
steps. Biosens.Bioelectron. 1995; 10: 141-153.

52.
Nordhausen CT, Maynard EM, Normann RA. Single unit recording capabilities of a
100 microelectrode array. Brain Res. 1996; 726: 129-140.

53.
Branner A, Normann RA. A multielectrode array for intrafascicular recording and
stimulation in sciatic nerve of cats. Brain Res.Bull. 2000; 51:
293-306.

54.
Meyer, J-U., Blau, C., and Stieglitz, Th. Multikontaktierung von Nerven mit
Mikrostrukturen. Spektrum D.Wissensch. 98-102. 1995.

55.
Smit, J. P. A. Selective Motor Stimulation using Endoneural
Prostheses. 1-24. 1996. Twente University.

56.
Rutten WL, van Wier HJ, Put JH. Sensitivity and selectivity of intraneural
stimulation using a silicon electrode array. IEEE Trans.Biomed.Eng
1991; 38: 192-198.

57.
Fried K, Frisen J, Mozart M. De- and regeneration of axons after minor lesions
in the rat sciatic nerve. Effects of microneurography electrode penetrations.
Pain 1989; 36: 93-102.

58.
Rice ASC, Andreev NY, McMahon SB. The consequences of microneurography
electrode-induced injury of peripheral nerves observed in the rat and man.
Pain 1994; 59: 385-393.

59.
Rice AS, McMahon SB. Peripheral nerve injury caused by injection needles used in
regional anaesthesia: influence of bevel configuration, studied in a rat model.
British Journal Of Anaesthesia 1992; 69: 433-438.

60.
Rice ASC, McMahon SB, Wall PD. The electrophysiological consequences of
electrode impalement of peripheral nerves in the rat. Brain Research
1993; 631: 221-226.

61.
Bowman BR, Erickson RC. Acute and chronic implantation of coiled wire
intraneural electrodes during cyclical electrical stimulation. Annals Of
Biomedical Engineering 1985; 13: 75-93.

62.
Hallin RG, Wiesenfeld-Hallin Z, Duranti R. Percutaneous microneurography in man
does not cause pressure block of almost all axons in the impaled nerve fascicle.
Neuroscience Letters 1986; 68: 356-361.

63.
Lefurge T, Goodall E, Horch K, Stensaas L, Schoenberg A. Chronically implanted
intrafascicular recording electrodes. Annals Of Biomedical Engineering
1991; 19: 197-207.

64.
Stanic U, Bajd T, Valencic V, Kljajic M, Acimovic R. Standardization of
kinematic gait measurements and automatic pathological gait pattern diagnostics.
Scand.J Rehabil.Med 1977; 9: 95-105.

65.
Petrofsky JS, Phillips CA. Discharge characteristics of motor units and the
surface EMG during fatiguing isometric contractions at submaximal tensions.
Aviat.Space Environ.Med 1985; 56: 581-586.

66.
Williams DF. Definitions in Biomaterials. Progress in Biomedical
Engineering. Amsterdam: Elsevier, 1987: 54.

67.
Bollen, L. S. and Svendsen, O. Regulatory Guidelines For Biocompatibility Safety
Testing. Medical Plastics and Biomaterials , 16. 1-5-1997.

68.
Salthouse TN. Cellular enzyme activity at the polymer-tissue interface: a
review. J Biomed.Mater.Res. 1976; 10: 197-229.

69.
Ratner BD, Hoffman A, Lemons JE, Schoen FJ. Biomaterials Science - An
Introduction to Materials in Medicine. New York : Academic Press,
1996.

70.
Agnew WF, McCreery DB. Neural Prostheses: Fundamental studies. Englewood Cliffs:
Prentice Hall, 1990.

71.
Vaudaux PE, Francois P, Proctor RA, McDevitt D, Foster TJ, Albrecht RM et
al. Use of adhesion-defective mutants of Staphylococcus aureus to define the
role of specific plasma proteins in promoting bacterial adhesion to canine
arteriovenous shunts. Infect.Immun. 1995; 63: 585-590.

72.
Wassall MA, Santin M, Peluso G, Denyer SP. Possible role of
alpha-1-microglobulin in mediating bacterial attachment to model surfaces. J
Biomed.Mater.Res. 1998; 40: 365-370.

73.
AAMI. Association for the Advancement of Medical Instrumentation American
National Standard for Implantable Peripheral Nerve Stimulators. 1984.
ANSI/AAMI NS15-1984.

74.
Romero, E. Feasibility of Long-term Nerve Stimulation with the Self-sizing
Spiral Cuff Electrode: Histological and Physiological Study. 1-207.
2000.  Université catholoque de Louvain.

75.
Grill WM, Mortimer JT. Electrical properties of implant encapsulation tissue.
Ann Biomed.Eng 1994; 22: 23-33.

76.
Naples GG, Mortimer JT, Scheiner A, Sweeney JD. A spiral nerve cuff electrode
for peripheral nerve stimulation. IEEE Trans.Biomed.Eng 1988; 35:
905-916.

77.
Agnew WF, McCreery DB, Yuen TG, Bullara LA. Histologic and physiologic
evaluation of electrically stimulated peripheral nerve: considerations for the
selection of parameters. Ann Biomed.Eng 1989; 17: 39-60.

78.
Grill WM, Mortimer JT. Neural and connective tissue response to long-term
implantation of multiple contact nerve cuff electrodes. J
Biomed.Mater.Res. 2000; 50: 215-226.

79.
McNeal DR, Waters R, Reswick J. Experience with implanted electrodes.
Neurosurgery 1977; 1: 228-229.

80.
Koller R, Girsch W, Liegl C, Gruber H, Holle J, Losert U et al. Long-term
results of nervous tissue alterations caused by epineurial electrode
application: an experimental study in rat sciatic nerve. Pacing
Clin.Electrophysiol. 1992; 15: 108-115.

81.
Girsch W, Koller R, Gruber H, Holle J, Liegl C, Losert U et al.
Histological assessment of nerve lesions caused by epineurial electrode
application in rat sciatic nerve. J Neurosurg. 1991; 74:
636-642.

82.
Agnew WF, McCreery DB, Bullara L, Yuen TG. Effects of Prolonged Electrical
Stimulation of Peripheral nerve. In: Agnew WF, McCreery DB, editors. Neural
Prostheses: Fundamental Studies. Englewood Cliffs: Prentice Hall, 1990:
147-68.

83.
Liang F, Jones EG. Peripheral nerve stimulation increases Fos immunoreactivity
without affecting type II Ca2+/calmodulin-dependent protein kinase, glutamic
acid decarboxylase, or GABAA receptor gene expression in cat spinal cord.
Exp.Brain Res. 1996; 111: 326-336.

84.
Nakamura M, Rosahl SK, Alkahlout E, Gharabaghi A, Walter GF, Samii M. C-Fos
immunoreactivity mapping of the auditory system after electrical stimulation of
the cochlear nerve in rats. Hearing Research 2003; 184: 75-81.

85.
Lanahan A, Worley P. Immediate-Early Genes and Synaptic Function.
Neurobiology of Learning and Memory 1998; 70: 37-43.

86.
Illing RB, Michler SA, Kraus KS, Laszig R. Transcription Factor Modulation and
Expression in the Rat Auditory Brainstem Following Electrical Intracochlear
Stimulation. Experimental Neurology 2002; 175: 226-244.

87.
Molander C, Hongpaisan J, Grant G. Changing pattern of c-FOS expression in
spinal cord neurons after electrical stimulation of the chronically injured
sciatic nerve in the rat. Neuroscience 1992; 50: 223-236.

88.
Illing RB. Activity-dependent plasticity in the adult auditory brainstem.
Audiol.Neurootol. 2001; 6: 319-345.

89.
Platenik J, Kuramoto N, Yoneda Y. Molecular mechanisms associated with long-term
consolidation of the NMDA signals. Life Sci. 2000; 67:
335-364.

90.
Muir GD, Steeves JD. Sensorimotor stimulation to improve locomotor recovery
after spinal cord injury. Trends in Neurosciences 1997; 20:
72-77.

91.
Popovic MR, Popovic DB, Keller T. Neuroprostheses for grasping.
Neurol.Res. 2002; 24: 443-452.

92.
Wieler M, Stein RB, Ladouceur M, Whittaker M, Smith AW, Naaman S et al.
Multicenter evaluation of electrical stimulation systems for walking.
Arch.Phys.Med Rehabil. 1999; 80: 495-500.

93.
Agnew WF, McCreery DB, Yuen TG, Bullara LA. Evolution and resolution of
stimulation-induced axonal injury in peripheral nerve. Muscle Nerve 1999;
22: 1393-1402.

94.
Brummer, S. B. and Turner, M. Electrochemical aspects of neuromuscular
stimulation. Functional Neuromuscular Stimulation: Report of a
Workshop. 1972. Washington, National Academy of Sciences.

95.
Agnew WF, McCreery DB, Yuen TG, Bullara LA. MK-801 protects against neuronal
injury induced by electrical stimulation. Neuroscience 1993; 52:
45-53.

96.
Deurloo, K. Selectivity in extraneural stimulation of peripheral
nerves. 8-21. 1999. Twente University.

97.
Popovic MR, Curt A, Keller T, Dietz V. Functional electrical stimulation for
grasping and walking: indications and limitations. Spinal Cord. 2001; 39:
403-412.

98.
Cochlear implants in adults and children. NIH Consens.Statement 1995; 13:
1-30.

99.
Vodovnik L, Kralj A, Stanic U, Acimovic R, Gros N. Recent applications of
functional electrical stimulation to stroke patients in Ljubljana .
Clin.Orthop. 1978; 64-70.

100.    Burridge JH, Taylor PN, Hagan SA, Wood DE,
Swain ID. The effects of common peroneal stimulation on the effort and speed of
walking: a randomized controlled trial with chronic hemiplegic patients.
Clin.Rehabil. 1997; 11: 201-210.

101.    Wieler, M., Naaman, S., and Stein, R. B.
WalkAid: An improved functional electrical stimulator for correcting
foot-drop. 101-104. 1996. Cleveland, Ohio. Proc 1st Ann Conf Int Functional
Electrical Stimulation Society.

102.    Kenney L, Bultstra G, Buschman R, Taylor P,
Mann G, Hermens H et al. An implantable two channel drop foot stimulator:
initial clinical results. Artif.Organs 2002; 26: 267-270.

103.    Taylor PN, Burridge JH, Dunkerley
AL, Wood DE, Norton JA, Singleton C et al. Clinical use of the
Odstock dropped foot stimulator: its effect on the speed and effort of walking.
Arch.Phys.Med Rehabil. 1999; 80: 1577-1583.

104.    Taylor PN, Burridge JH, Dunkerley
AL, Lamb A, Wood DE, Norton JA et al. Patients' perceptions of the
Odstock Dropped Foot Stimulator (ODFS). Clin.Rehabil. 1999; 13:
439-446.

105.    Lyons GM, Sinkjaer T, Burridge JH, Wilcox
DJ. A review of portable FES-based neural orthoses for the correction of drop
foot. IEEE Trans.Neural Syst.Rehabil.Eng 2002; 10: 260-279.

106.    Kralj A, Bajd T, Turk R. Electrical
stimulation providing functional use of paraplegic patient muscles. Med
Prog.Technol. 1980; 7: 3-9.

107.    Bajd T, Kralj A, Turk R, Benko H, Sega J.
The use of a four-channel electrical stimulator as an ambulatory aid for
paraplegic patients. Phys.Ther. 1983; 63: 1116-1120.

108.    Kralj A, Bajd T, Turk R. Enhancement of gait
restoration in spinal injured patients by functional electrical stimulation.
Clin.Orthop. 1988; 34-43.

109.    Brindley GS, Polkey CE, Rushton
DN. Electrical splinting of the knee in paraplegia. Paraplegia 1979;
16: 428-437.

110.    Donaldson ND, Perkins TA, Worley
AC. Lumbar root stimulation for restoring leg function: stimulator and
measurement of muscle actions. Artif.Organs 1997; 21: 247-249.

111.    Kobetic R, Triolo RJ, Marsolais EB. Muscle
selection and walking performance of multichannel FES systems for ambulation in
paraplegia. IEEE Trans.Rehabil.Eng 1997; 5: 23-29.

112.    Graupe D, Davis R, Kordylewski H, Kohn KH.
Ambulation by traumatic T4-12 paraplegics using functional neuromuscular
stimulation. Critical Reviews in Neurosurgery 1998; 8:
221-231.

113.    Graupe D, Kohn KH. Functional neuromuscular
stimulator for short-distance ambulation by certain thoracic-level
spinal-cord-injured paraplegics. Surg Neurol. 1998; 50:
202-207.

114.    Popovic MR, Keller T, Pappas IP, Dietz V,
Morari M. Surface-stimulation technology for grasping and walking
neuroprosthesis.  IEEE Eng Med Biol.Mag. 2001; 20: 82-93.

115.    Scott TR, Peckham PH, Keith MW. Upper
extremity neuroprostheses using functional electrical stimulation. Baillieres
Clin.Neurol. 1995; 4: 57-75.

116.    Snoek GJ, IJzerman MJ, in 't Groen FA,
Stoffers TS, Zilvold G. Use of the NESS handmaster to restore handfunction in
tetraplegia: clinical experiences in ten patients. Spinal Cord. 2000; 38:
244-249.

117.    Prochazka A, Gauthier M, Wieler M, Kenwell
Z. The bionic glove: an electrical stimulator garment that provides controlled
grasp and hand opening in quadriplegia . Arch.Phys.Med Rehabil. 1997; 78:
608-614.

118.    Handa Y, Handa T, Ichie M, Murakami H,
Hoshimiya N, Ishikawa S et al. Functional electrical stimulation (FES)
systems for restoration of motor function of paralyzed muscles--versatile
systems and a portable system. Front Med Biol.Eng 1992; 4:
241-255.

119.    Rebersek S, Vodovnik L. Proportionally
controlled functional electrical stimulation of hand. Arch.Phys.Med
Rehabil. 1973; 54: 378-382.

120.    Popovic, D., Popovic, M., Stojanovic, A.,
Pjanovic, A., Radosavljevic, S., and Vulovic, D. Clinical evaluation of the
Belgrade Grasping System. 247-250. 1998. Vienna, Austria. Proc 6th Vienna
Internat Workshop on Functional Electrical Stimulation.

121.    Keith MW. Neuroprostheses for the upper
extremity. Microsurgery 2001; 21: 256-263.

122.    Carroll S, Cooper C, Brown D, Sormann G,
Flood S, Denison M. Australian experience with the Freehand System for restoring
grasp in quadriplegia. Aust.N.Z.J Surg 2000; 70: 563-568.

123.    Popovic D, Stojanovic A, Pjanovic A,
Radosavljevic S, Popovic M, Jovic S et al. Clinical evaluation of the
bionic glove. Arch.Phys.Med Rehabil. 1999; 80: 299-304.

124.    Thorsen R, Spadone R, Ferrarin M. A pilot
study of myoelectrically controlled FES of upper extremity. IEEE Trans.Neural
Syst.Rehabil.Eng 2001; 9: 161-168.

125.    Glenn WW, Hogan JF, Loke JS, Ciesielski TE,
Phelps ML, Rowedder R. Ventilatory support by pacing of the conditioned
diaphragm in quadriplegia. N.Engl.J Med. 1984; 310: 1150-1155.

126.    Glenn WW, Phelps ML, Elefteriades JA, Dentz
B, Hogan JF. Twenty years of experience in phrenic nerve stimulation to pace the
diaphragm. Pacing Clin.Electrophysiol. 1986; 9: 780-784.

127.    Brouillette RT, Marzocchi M. Diaphragm
pacing: clinical and experimental results. Biol.Neonate 1994; 65:
265-271.

128.    Glanz M, Klawansky S, Stason W, Berkey C,
Chalmers TC. Functional electrostimulation in poststroke rehabilitation: a
meta-analysis of the randomized controlled trials. Arch.Phys.Med
Rehabil. 1996; 77: 549-553.

129.    de Kroon JR, van der Lee JH, IJzerman MJ,
Lankhorst GJ. Therapeutic electrical stimulation to improve motor control and
functional abilities of the upper extremity after stroke: a systematic review.
Clin.Rehabil. 2002; 16: 350-360.

130.    Ada L, Foongchomcheay A. Efficacy of
electrical stimulation in preventing or reducing subluxation of the shoulder
after stroke: a meta-analysis. Aust.J Physiother. 2002; 48:
257-267.

131.    Schmidt RA, Tanagho EA. Feasibility of
controlled micturition through electric stimulation. Urol Int. 1979; 34:
199-230.

132.    Groen J, Bosch JL. Neuromodulation
techniques in the treatment of the overactive bladder. BJU.Int. 2001; 87:
723-731.

133.    Hassouna MM, Siegel SW, Nyeholt
AA, Elhilali MM, van Kerrebroeck PE, Das AK et al. Sacral
neuromodulation in the treatment of urgency-frequency symptoms: a multicenter
study on efficacy and safety. J Urol. 2000; 163: 1849-1854.

134.    Bosch JL, Groen J. Sacral (S3) segmental
nerve stimulation as a treatment for urge incontinence in patients with detrusor
instability: results of chronic electrical stimulation using an implantable
neural prosthesis. J Urol. 1995; 154: 504-507.

135.    Brindley GS, Polkey CE, Rushton DN. Sacral
anterior root stimulators for bladder control in paraplegia. Paraplegia
1982; 20: 365-381.

136.    Brindley GS. The first 500 patients with
sacral anterior root stimulator implants: general description. Paraplegia
1994; 32: 795-805.

137.    van der Aa HE, Alleman E, Nene A, Snoek G.
Sacral anterior root stimulation for bladder control: clinical results.
Arch.Physiol Biochem. 1999; 107: 248-256.

138.    Van Kerrebroeck EV, van der Aa HE, Bosch JL,
Koldewijn EL, Vorsteveld JH, Debruyne FM. Sacral rhizotomies and electrical
bladder stimulation in spinal cord injury. Part I: Clinical and urodynamic
analysis. Dutch Study Group on Sacral Anterior Root Stimulation. Eur.Urol
1997; 31: 263-271.

139.    Wielink G, Essink-Bot ML, van Kerrebroeck
PE, Rutten FF. Sacral rhizotomies and electrical bladder stimulation in spinal
cord injury. 2. Cost-effectiveness and quality of life analysis. Dutch Study
Group on Sacral Anterior Root Stimulation. Eur.Urol 1997; 31:
441-446.

140.    Siegel SW, Catanzaro F, Dijkema HE, Elhilali
MM, Fowler CJ, Gajewski JB et al. Long-term results of a multicenter
study on sacral nerve stimulation for treatment of urinary urge incontinence,
urgency-frequency, and retention. Urology 2000; 56: 87-91.

141.    Bosch JL, Groen J. Neuromodulation:
urodynamic effects of sacral (S3) spinal nerve stimulation in patients with
detrusor instability or detrusor hyperflexia. Behav.Brain Res. 1998; 92:
141-150.

142.    Thon WF, Baskin LS, Jonas U, Tanagho EA,
Schmidt RA. Neuromodulation of voiding dysfunction and pelvic pain. World J
Urol 2003; 9: 138.

143.    Elabbady AA, Hassouna MM, Elhilali MM.
Neural stimulation for chronic voiding dysfunctions. J Urol. 1994; 152:
2076-2080.

144.    Jonas U, Fowler CJ, Chancellor MB, Elhilali
MM, Fall M, Gajewski JB et al. Efficacy of sacral nerve stimulation for
urinary retention: results 18 months after implantation. J Urol. 2001;
165: 15-19.

145.    Vapnek JM, Schmidt RA. Restoration of
voiding in chronic urinary retention using neuroprothesis. World J Urol
2003; 9: 142.

146.    Marani E. Innervation of the Mature Human
Pelvis. The Pelvis: Another View. Enschede: Twente University Press,
2002: 95-112.

147.    Dobelle WH, Mladejovsky MG. Phosphenes
produced by electrical stimulation of human occipital cortex, and their
application to the development of a prosthesis for the blind. J Physiol
1974; 243: 553-576.

148.    Bartlett JR, Doty RW. An exploration of the
ability of macaques to detect microstimulation of striate cortex. Acta
Neurobiol.Exp.(Wars.) 1980; 40: 713-727.

149.    Schmidt EM, Bak MJ, Hambrecht FT, Kufta CV,
O'Rourke DK, Vallabhanath P. Feasibility of a visual prosthesis for the blind
based on intracortical microstimulation of the visual cortex. Brain 1996;
119 ( Pt 2): 507-522.

150.    Dobelle WH. Artificial vision for the blind
by connecting a television camera to the visual cortex. ASAIO J 2000; 46:
3-9.

151.    Veraart C, Raftopoulos C, Mortimer JT,
Delbeke J, Pins D, Michaux G et al. Visual sensations produced by optic
nerve stimulation using an implanted self-sizing spiral cuff electrode. Brain
Res. 1998; 813: 181-186.

152.    Grumet AE, Wyatt JL, Jr., Rizzo JF, III.
Multi-electrode stimulation and recording in the isolated retina. J
Neurosci.Methods 2000; 101: 31-42.

153.    Chow AY, Chow VY. Subretinal electrical
stimulation of the rabbit retina. Neurosci.Lett. 1997; 225:
13-16.

154.    Zrenner E, Miliczek KD, Gabel VP, Graf HG,
Guenther E, Haemmerle H et al. The development of subretinal
microphotodiodes for replacement of degenerated photoreceptors. Ophthalmic
Res. 1997; 29: 269-280.

155.    Weiland JD, Humayun MS, Dagnelie G, de Juan
E Jr, Greenberg RJ, Iliff NT. Understanding the origin of visual percepts
elicited by electrical stimulation of the human retina. Graefes
Arch.Clin.Exp.Ophthalmol. 1999; 237: 1007-1013.

156.    Rizzo JF, III, Wyatt J, Loewenstein J, Kelly
S, Shire D. Methods and perceptual thresholds for short-term electrical
stimulation of human retina with microelectrode arrays. Invest
Ophthalmol.Vis.Sci. 2003; 44: 5355-5361.

157.    Humayun MS, Weiland JD, Fujii GY, Greenberg
R, Williamson R, Little J et al. Visual perception in a blind subject
with a chronic microelectronic retinal prosthesis. Vision Res. 2003; 43:
2573-2581.

158.    Barbeau H, McCrea DA, O'Donovan MJ,
Rossignol S, Grill WM, Lemay MA. Tapping into spinal circuits to restore motor
function. Brain Res.Brain Res.Rev. 1999; 30: 27-51.

159.    Kargo WJ, Giszter SF. Rapid correction of
aimed movements by summation of force-field primitives. J Neurosci. 2000;
20: 409-426.

160.    Dimitrijevic MR, Gerasimenko Y, Pinter MM.
Evidence for a spinal central pattern generator in humans. Ann
N.Y.Acad.Sci. 1998; 860: 360-376.

161.    Richardson RR, Cerullo LJ, McLone
DG, Gutierrez FA, Lewis V. Percutaneous epidural neurostimulation in
modulation of paraplegic spasticity. Six case reports. Acta
Neurochir.(Wien.) 1979; 49: 235-243.

162.    Pinter MM, Gerstenbrand F, Dimitrijevic MR.
Epidural electrical stimulation of posterior structures of the human lumbosacral
cord: 3. Control Of spasticity. Spinal Cord. 2000; 38:
524-531.

163.    Deurloo KE, Holsheimer J, Boom HB.
Transverse tripolar stimulation of peripheral nerve: a modelling study of
spatial selectivity. Med Biol.Eng Comput. 1998; 36: 66-74.

164.    Deurloo KE, Holsheimer J, Bergveld P. Nerve
stimulation with a multi-contact cuff electrode: validation of model
predictions. Arch.Physiol Biochem. 2000; 108: 349-359.

165.    Rutten WL, Smit JP, Frieswijk TA, Bielen JA,
Brouwer AL, Buitenweg JR et al. Neuro-electronic interfacing with
multielectrode arrays. IEEE Eng Med Biol.Mag. 1999; 18: 47-55.

166.    Smit JP, Rutten WL, Boom HB. Endoneural
selective stimulating using wire-microelectrode arrays. IEEE
Trans.Rehabil.Eng 1999; 7: 399-412.

167.    Frieswijk TA, Smit JP, Rutten WL, Boom HB.
Force-current relationships in intraneural stimulation: role of extraneural
medium and motor fibre clustering. Med Biol.Eng Comput. 1998; 36:
422-430.

168.    Zhou BH, Baratta R, Solomonow M.
Manipulation of muscle force with various firing rate and recruitment control
strategies. IEEE Trans.Biomed.Eng 1987; 34: 128-139.

169.    Feirabend HK, Choufoer H, Ploeger
S, Holsheimer J, van Gool JD. Morphometry of human superficial dorsal and
dorsolateral column fibres: significance to spinal cord stimulation.
Brain 2002; 125: 1137-1149.

170.    Fang ZP, Mortimer JT. A method to effect
physiological recruitment order in electrically activated muscle. IEEE
Trans.Biomed.Eng 1991; 38: 175-179.

171.    Fang ZP, Mortimer JT. Selective activation
of small motor axons by quasi-trapezoidal current pulses. IEEE
Trans.Biomed.Eng 1991; 38: 168-174.

172.    van Bolhuis AI, Holsheimer J, Savelberg HH.
A nerve stimulation method to selectively recruit smaller motor-units in rat
skeletal muscle. J Neurosci.Methods 2001; 107: 87-92.

173.    Grill WM, Mortimer JT. Inversion of the
current-distance relationship by transient depolarization. IEEE
Trans.Biomed.Eng 1997; 44: 1-9.

174.    Deurloo KE, Holsheimer J, Bergveld P. The
effect of subthreshold prepulses on the recruitment order in a nerve trunk
analyzed in a simple and a realistic volume conductor model. Biol.Cybern.
2001; 85: 281-291.