U.S. patent application number 11/592804 was filed with the patent office on 2007-08-23 for method and apparatus for visual neural stimulation.
Invention is credited to E.J. Chichilnisky, Robert J. Greenberg, Matthew J. McMahon, Chris Sekirnjak.
Application Number | 20070198066 11/592804 |
Document ID | / |
Family ID | 38429331 |
Filed Date | 2007-08-23 |
United States Patent
Application |
20070198066 |
Kind Code |
A1 |
Greenberg; Robert J. ; et
al. |
August 23, 2007 |
Method and apparatus for visual neural stimulation
Abstract
Existing epiretinal implants for the blind are designed to
electrically stimulate large groups of surviving retinal neurons
using a small number of electrodes with diameters of several
hundred .mu.m. To increase the spatial resolution of artificial
sight, electrodes much smaller than those currently in use are
desirable. In this study we stimulated and recorded ganglion cells
in isolated pieces of rat, guinea pig, and monkey retina. We
utilized micro-fabricated hexagonal arrays of 61 platinum disk
electrodes with diameters between 6 and 25 .mu.m, spaced 60 .mu.m
apart. Charge-balanced current pulses evoked one or two spikes at
latencies as short as 0.2 ms, and typically only one or a few
recorded ganglion cells were stimulated. Application of several
synaptic blockers did not abolish the evoked responses, implying
direct activation of ganglion cells. Threshold charge densities
were typically below 0.1 mC/cm2 for a pulse duration of 100 .mu.s,
corresponding to charge thresholds of less than 100 pC. Stimulation
remained effective after several hours and at high frequencies. To
demonstrate that closely spaced electrodes can elicit independent
ganglion cell responses, we utilized the multi-electrode array to
stimulate several nearby ganglion cells simultaneously. From these
data we conclude that electrical stimulation of mammalian retina
with small-diameter electrode arrays is achievable and can provide
high temporal and spatial precision at low charge densities. We
review previous epiretinal stimulation studies and discuss our
results in the context of 32 other publications, comparing
threshold parameters and safety limits.
Inventors: |
Greenberg; Robert J.; (Los
Angeles, CA) ; McMahon; Matthew J.; (Los Angeles,
CA) ; Sekirnjak; Chris; (Denver, CO) ;
Chichilnisky; E.J.; (Del Mar, CA) |
Correspondence
Address: |
SECOND SIGHT MEDICAL PRODUCTS, INC.
12744 SAN FERNANDO ROAD
BUILDING 3
SYLMAR
CA
91342
US
|
Family ID: |
38429331 |
Appl. No.: |
11/592804 |
Filed: |
November 3, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60733701 |
Nov 3, 2005 |
|
|
|
Current U.S.
Class: |
607/53 ;
607/141 |
Current CPC
Class: |
Y10T 29/49117 20150115;
A61N 1/36046 20130101 |
Class at
Publication: |
607/053 ;
607/141 |
International
Class: |
A61N 1/18 20060101
A61N001/18 |
Goverment Interests
GOVERNMENT RIGHTS NOTICE
[0002] This invention was made with government support under grant
No. R24EY12893-01, awarded by the National Institutes of Health.
The government has certain rights in the invention.
Claims
1. An electrode array for stimulating visual neurons comprising: A
non-conductive body; A plurality of conductive electrodes wherein
said electrode are less than 20 .mu.m in size and less than 60
.mu.m apart; and means for supporting said body in close proximity
to visual neurons.
2. The electrode array according to claim 1, wherein said
electrodes are arranged hexagonally.
3. The electrode array according to claim 1, wherein said
non-conductive body is a fluoro-polymer
4. The electrode array according to claim 1, wherein said
electrodes are arranged in a pattern longer in one dimension than
the other dimension; and Wherein said one dimension corresponds to
horizontal in a visual scene.
5. A flexible circuit electrode array adapted for neural
stimulation comprising: A polymer base layer; Metal traces
deposited on said polymer base layer, including electrodes suitable
to stimulate neural tissue; and A polymer top layer deposited on
said polymer base layer and said metal traces; Wherein said polymer
top layer defines openings smaller than said electrodes to overlap
said electrodes.
6. A flexible circuit electrode array adapted for neural
stimulation comprising: A polymer base layer; Metal traces
deposited on said polymer base layer, including electrodes suitable
to stimulate neural tissue; and A polymer top layer deposited on
said polymer base layer and said metal traces; wherein said
electrodes are less than 20. .mu.m in size and less than 60 .mu.m
apart
7. The flexible circuit electrode array according to claim 6,
wherein said polymer base layer, said metal traces and said polymer
top layer are curved to approximately the curvature of an eye.
8. The flexible circuit electrode array according to claim 6,
further comprising at least one bumper bonded to a peripheral edge
of said flexible circuit electrode array.
9. The flexible circuit electrode array according to claim 6,
further comprising a narrowed portion in a flexible circuit cable
portion of said flexible circuit electrode array.
10. The flexible circuit electrode array according to claim 6,
further comprising a stress relief membrane suitable for attachment
of said flexible circuit electrode array, wherein said stress
relief membrane is a more compliant material than said polymer base
layer.
11. The flexible circuit electrode array according to claim 9,
wherein said narrowed portion is suitable to pierce a sclera.
12. The flexible circuit electrode array according to claim 11,
wherein said narrowed portion is a diagonal fold in a flexible
circuit cable portion of said flexible circuit electrode array.
13. The flexible circuit electrode array according to claim 12,
where said diagonal fold is across a dogleg in said flexible
circuit electrode array.
14. The flexible circuit electrode array according to claim 13,
further comprising bond pads coupled to said metal traces on an end
of said flexible circuit electrode array opposite to said
electrodes and openings in said polymer top layer for said
electrodes and said bond pads.
15. The flexible circuit electrode array according to claim 8,
where said bumper is a continuous skirt covering at least of
portion of said flexible circuit electrode array.
16. The flexible circuit electrode array according to claim 8,
where said bumper is a continuous skirt covering at least of
portion of a cable portion of said flexible circuit electrode
array.
17. The flexible circuit electrode array according to claim 16,
further comprising a sleeve at least partially covering a flexible
circuit cable portion of said flexible circuit electrode array.
18. The flexible circuit electrode array according to claim 17,
wherein said sleeve and said bumper are a continuous body.
19. The flexible circuit electrode array according to claim 6,
wherein said polymer base layer, said metal traces and said polymer
top layer for a continuous electrode array and flexible circuit
cable where said flexible circuit cable forms a partial loop to
resist transmission of forces through said flexible circuit
cable.
20. The flexible circuit electrode array according to claim 6,
wherein said electrodes are arranged in a hexagonal pattern.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is related to and claims benefit of US
provisional application 60/733,2005, for Electrical Stimulation of
Mammalian Retinal Ganglion Cells with Multi-Electrode Arrays filed
Nov. 3, 2005. This application is related to and incorporates by
reference, U.S. patent application Ser. No. 11/207,644 Flexible
Circuit Electrode Array filed Aug. 19, 2005.
FIELD OF THE INVENTION
[0003] The present invention is generally directed to neural
stimulation and more specifically to an improved method of
providing artificial vision through electrical stimulation of
visual neurons.
BACKGROUND OF THE INVENTION
[0004] In 1755 LeRoy passed the discharge of a Leyden jar through
the orbit of a man who was blind from cataract and the patient saw
"flames passing rapidly downwards." Ever since, there has been a
fascination with electrically elicited visual perception. The
general concept of electrical stimulation of retinal cells to
produce these flashes of light or phosphenes has been known for
quite some time. Based on these general principles, some early
attempts at devising a prosthesis for aiding the visually impaired
have included attaching electrodes to the head or eyelids of
patients. While some of these early attempts met with some limited
success, these early prosthetic devices were large, bulky and could
not produce adequate simulated vision to truly aid the visually
impaired.
[0005] In the early 1930's, Foerster investigated the effect of
electrically stimulating the exposed occipital pole of one cerebral
hemisphere. He found that, when a point at the extreme occipital
pole was stimulated, the patient perceived a small spot of light
directly in front and motionless (a phosphene). Subsequently,
Brindley and Lewin (1968) thoroughly studied electrical stimulation
of the human occipital (visual) cortex. By varying the stimulation
parameters, these investigators described in detail the location of
the phosphenes produced relative to the specific region of the
occipital cortex stimulated. These experiments demonstrated: (1)
the consistent shape and position of phosphenes; (2) that increased
stimulation pulse duration made phosphenes brighter; and (3) that
there was no detectable interaction between neighboring electrodes
which were as close as 2.4 mm apart.
[0006] As intraocular surgical techniques have advanced, it has
become possible to apply stimulation on small groups and even on
individual retinal cells to generate focused phosphenes through
devices implanted within the eye itself. This has sparked renewed
interest in developing methods and apparatuses to aid the visually
impaired. Specifically, great effort has been expended in the area
of intraocular retinal prosthesis devices in an effort to restore
vision in cases where blindness is caused by photoreceptor
degenerative retinal diseases such as retinitis pigmentosa and age
related macular degeneration which affect millions of people
worldwide.
[0007] Neural tissue can be artificially stimulated and activated
by prosthetic devices that pass pulses of electrical current
through electrodes on such a device. The passage of current causes
changes in electrical potentials across visual neuronal membranes,
which can initiate visual neuron action potentials, which are the
means of information transfer in the nervous system.
[0008] Based on this mechanism, it is possible to input information
into the nervous system by coding the information as a sequence of
electrical pulses which are relayed to the nervous system via the
prosthetic device. In this way, it is possible to provide
artificial sensations including vision.
[0009] One typical application of neural tissue stimulation is in
the rehabilitation of the blind. Some forms of blindness involve
selective loss of the light sensitive transducers of the retina.
Other retinal neurons remain viable, however, and may be activated
in the manner described above by placement of a prosthetic
electrode device on the inner (toward the vitreous) retinal surface
(epiretial). This placement must be mechanically stable, minimize
the distance between the device electrodes and the visual neurons,
and avoid undue compression of the visual neurons.
[0010] In 1986, Bullara (U.S. Pat. No. 4,573,481) patented an
electrode assembly for surgical implantation on a nerve. The matrix
was silicone with embedded iridium electrodes. The assembly fit
around a nerve to stimulate it.
[0011] Dawson and Radtke stimulated cat's retina by direct
electrical stimulation of the retinal ganglion cell layer. These
experimenters placed nine and then fourteen electrodes upon the
inner retinal layer (i.e., primarily the ganglion cell layer) of
two cats. Their experiments suggested that electrical stimulation
of the retina with 30 to 100 uA current resulted in visual cortical
responses. These experiments were carried out with needle-shaped
electrodes that penetrated the surface of the retina (see also U.S.
Pat. No. 4,628,933 to Michelson).
[0012] The Michelson '933 apparatus includes an array of
photosensitive devices on its surface that are connected to a
plurality of electrodes positioned on the opposite surface of the
device to stimulate the retina. These electrodes are disposed to
form an array similar to a "bed of nails" having conductors which
impinge directly on the retina to stimulate the retinal cells. U.S.
Pat. No. 4,837,049 to Byers describes spike electrodes for neural
stimulation. Each spike electrode pierces neural tissue for better
electrical contact. U.S. Pat. No. 5,215,088 to Norman describes an
array of spike electrodes for cortical stimulation. Each spike
pierces cortical tissue for better electrical contact.
[0013] The art of implanting an intraocular prosthetic device to
electrically stimulate the retina was advanced with the
introduction of retinal tacks in retinal surgery. De Juan, et al.
at Duke University Eye Center inserted retinal tacks into retinas
in an effort to reattach retinas that had detached from the
underlying choroid, which is the source of blood supply for the
outer retina and thus the photoreceptors. See, e.g., E. de Juan, et
al., 99 Am. J. Ophthalmol. 272 (1985). These retinal tacks have
proved to be biocompatible and remain embedded in the retina, and
choroid/sclera, effectively pinning the retina against the choroid
and the posterior aspects of the globe. Retinal tacks are one way
to attach a retinal array to the retina. U.S. Pat. No. 5,109,844 to
de Juan describes a flat electrode array placed against the retina
for visual stimulation. U.S. Pat. No. 5,935,155 to Humayun
describes a retinal prosthesis for use with the flat retinal array
described in de Juan.
[0014] Recent attempts to restore vision in the blind have met with
extraordinary success. Electrical stimulation of retinas in people
with neurodegenerative diseases has demonstrated the potential for
direct excitation of neurons as a means of re-establishing sight.
Long-term retinal implants in several profoundly blind people were
shown to produce perceptions of light and allowed for the detection
of motion and discrimination of very simple shapes (Humayun 2003;
Humayun et al. 2003). Such achievement brings hope to the millions
of people worldwide who suffer from photoreceptor loss due to
advanced retinitis pigmentosa or age-related macular degeneration
(Heckenlively et al. 1988; Klein et al. 1997). It is expected that
ten years from now, macular degeneration will become the single
leading cause of legal blindness with an incidence as high as 5.5%
in people over 65 (Klein et al. 1997). While degenerative diseases
result in severe damage to photoreceptors, inner retinal neurons
survive at fairly high rates (Stone et al. 1992; Santos et al.
1997; Kim et al. 2002) and may be electrically excitable. The
fundamental concept underlying retinal neuroprosthetic devices is
to electrically activate those residual neurons by bypassing the
damaged photoreceptors, thus achieving artificial vision in
otherwise blind patients. Of several prosthetics designs,
epiretinal implants specifically target ganglion cells by
positioning electrodes in close proximity to the inner surface of
the retina.
[0015] In spite of recent successes, the current implants are but a
first step toward restoring sight. To create useful vision,
stimulating electrodes must be arranged in two-dimensional arrays
that generate a visual image made up of a matrix of discrete
perceptions of light. Psychophysical studies suggest that foveal
implants may provide the user with an acceptable level of mobility
if they contain a minimum of about 600 electrodes (Cha et al.
1992a; Cha et al. 1992b). To achieve this number or greater,
electrodes must be tightly packed, necessitating small stimulation
sites. At present a typical epiretinal implant contains tens of
electrodes with diameters of a few hundred pm, spaced several
hundred pm apart (Humayun 2003). Considering that such electrodes
are much larger than the cells they stimulate, the need for
implants with hundreds or thousands of much smaller electrodes is
apparent.
[0016] The success of the next generation of implantable devices
will be tied to our understanding of how to activate neurons with
extracellular electric stimuli applied to the retinal surface
through electrodes that approach cellular dimensions. Little is
known about the parameters which would permit reliable retinal
stimulation with small electrodes. When the electrode surface area
is reduced, current density and charge density increase rapidly,
and high charge densities are known to cause tissue damage by
electrochemical reactions (Pollen 1977; Brummer et al. 1983;
Tehovnik 1996). A detailed in vitro analysis of small electrode
stimulation is thus a prerequisite for developing such implants for
use in human patients.
[0017] A comprehensive literature review reveals that the
feasibility of stimulation with arrays of small electrodes in
mammalian tissue has not been adequately tested. The majority of
studies involving retinal stimulation have used needle-shaped
probes with one or two conductors at the end of an insulated rod,
such as platinum wires or concentric microelectrodes. In its
simplest form, such stimulating probes are made of metal wires
several hundred .mu.m in diameter, exposed at the tip and insulated
elsewhere (Doty and Grimm 1962; Humayun et al. 1994; Nadig 1999;
Weiland et al. 1999; Suzuki et al. 2004).
[0018] Others have attempted to utilize stimulating microprobes
with tip diameters of 25 .mu.m or smaller (Dawson and Radtke 1977;
Wyatt et al. 1994; Rizzo et al. 1997; Jensen et al. 2003). However,
the geometry of such probes differs greatly from the planar disk
electrode design developed for current epiretinal implants.
Stimulation, furthermore, is always limited to a single stimulation
site, precluding the study of stimulation using multiple electrodes
and their interaction effects. The use of multi-electrode arrays
for retinal stimulation has been mainly limited to large electrodes
with diameters between 100 and 1500 .mu.m (Greenberg 1998; Humayun
et al. 1999; Hesse et al. 2000; Walter and Heimann 2000; Humayun et
al. 2003; Rizzo et al. 2003b). Multi-electrode arrays with smaller
electrodes (around 10 .mu.m diameter) have been utilized to
stimulate the retina in the subretinal space (Zrenner et al. 1999;
Stett et al. 2000). Grumet has used an array to selectively
stimulate the axons of retinal ganglion cells, using a separate
distant array to record somatic spikes (Grumet 1999; Grumet et al.
2000). No study has targeted mammalian ganglion cell bodies for
direct epiretinal stimulation using planar electrodes with surface
areas below 200 .mu.m2. In this study we establish thresholds for
stimulation of ganglion cells in rat, guinea pig, and primate
retina using electrodes with surface areas of 30-500 .mu.m2
(diameters of 6-25 .mu.m). We then used these parameters to further
investigate frequency dependence, pharmacology, and spatial
interaction effects of stimulation. Our arrays use planar disk
microelectrodes very similar to those utilized in present
epiretinal prosthetics, but smaller by an order or two of
magnitude. We conclude our analysis by discussing the results in
the context of the pertinent literature. Early and preliminary
portions of this work have been presented elsewhere (Sekirnjak et
al. 2005).
SUMMARY OF THE INVENTION
[0019] Existing epiretinal implants for the blind are designed to
electrically stimulate large groups of surviving retinal neurons
using a small number of electrodes with diameters of several
hundred .mu.m. To increase the spatial resolution of artificial
sight, electrodes much smaller than those currently in use are
desirable. In this study we stimulated and recorded ganglion cells
in isolated pieces of rat, guinea pig, and monkey retina. We
utilized micro-fabricated hexagonal arrays of 61 platinum disk
electrodes with diameters between 6 and 25 .mu.m, spaced 60 .mu.m
apart. Charge-balanced current pulses evoked one or two spikes at
latencies as short as 0.2 ms, and typically only one or a few
recorded ganglion cells were stimulated. Application of several
synaptic blockers did not abolish the evoked responses, implying
direct activation of ganglion cells. Threshold charge densities
were typically below 0.1 mC/cm2 for a pulse duration of 100 .mu.s,
corresponding to charge thresholds of less than 100 pC. Stimulation
remained effective after several hours and at high frequencies. To
demonstrate that closely spaced electrodes can elicit independent
ganglion cell responses, we utilized the multi-electrode array to
stimulate several nearby ganglion cells simultaneously. From these
data we conclude that electrical stimulation of mammalian retina
with small-diameter electrode arrays is achievable and can provide
high temporal and spatial precision at low charge densities. We
review previous epiretinal stimulation studies and discuss our
results in the context of 32 other publications, comparing
threshold parameters and safety limits.
BRIEF DESCRIPTION OF THE DRAWINGS
[0020] FIG. 1A planar view of the distribution of electrodes in the
referred retinal array over retinal tissue.
[0021] FIG. 1B is side view of the preferred retinal array.
[0022] FIG. 1C is a waveform showing a stimulation pattern.
[0023] FIG. 2A-C are waveforms of evoked responses.
[0024] FIG. 3A-D are waveforms showing the relationship between
short and long latency spikes.
[0025] FIG. 4A-C are waveforms showing the results of
pharmacological manipulations.
[0026] FIG. 5A-B are diagrams showing the spatial spread of
stimulation.
[0027] FIG. 6 is a strength duration curve.
[0028] FIG. 7 is a graph of threshold current and charge.
[0029] FIG. 8A-C are waveforms showing frequency dependence of
stimulation.
[0030] FIG. 9 is a waveform illustrating continuous low-frequency
stimulation.
[0031] FIG. 10A-B are waveforms illustrating multiple site
stimulation.
[0032] FIG. 11A-D are graphs showing dependence of threshold on
electrode size.
[0033] FIG. 12-A-D are graphs showing a summary of threshold
information from literature analysis.
[0034] FIG. 13 is a graph showing threshold charge density vs.
charge.
[0035] FIG. 14 is a perspective view of the implanted portion of
the preferred retinal prosthesis.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0036] The following description is of the best mode presently
contemplated for carrying out the invention. This description is
not to be taken in a limiting sense, but is made merely for the
purpose of describing the general principles of the invention. The
scope of the invention should be determined with reference to the
claims.
[0037] Recent attempts to restore vision in the blind have met with
extraordinary success. Electrical stimulation of retinas in people
with neurodegenerative diseases has demonstrated the potential for
direct excitation of neurons as a means of re-establishing sight.
Long-term retinal implants in several profoundly blind people were
shown to produce perceptions of light and allowed for the detection
of motion and discrimination of very simple shapes (Humayun 2003;
Humayun et al. 2003). Such achievement brings hope to the millions
of people worldwide who suffer from photoreceptor loss due to
advanced retinitis pigmentosa or age-related macular degeneration
(Heckenlively et al. 1988; Klein et al. 1997). It is expected that
ten years from now, macular degeneration will become the single
leading cause of legal blindness with an incidence as high as 5.5%
in people over 65 (Klein et al. 1997). While degenerative diseases
result in severe damage to photoreceptors, inner retinal neurons
survive at fairly high rates (Stone et al. 1992; Santos et al.
1997; Kim et al. 2002) and may be electrically excitable. The
fundamental concept underlying retinal neuroprosthetic devices is
to electrically activate those residual neurons by bypassing the
damaged photoreceptors, thus achieving artificial vision in
otherwise blind patients. Of several prosthetics designs,
epiretinal implants specifically target ganglion cells by
positioning electrodes in close proximity to the inner surface of
the retina. In spite of recent successes, the current implants are
but a first step toward restoring sight. To create useful vision,
stimulating electrodes must be arranged in two-dimensional arrays
that generate a visual image made up of a matrix of discrete
perceptions of light. Psychophysical studies suggest that foveal
implants may provide the user with an acceptable level of mobility
if they contain a minimum of about 600 electrodes (Cha et al.
1992a; Cha et al. 1992b). To achieve this number or greater,
electrodes must be tightly packed, necessitating small stimulation
sites. At present a typical epiretinal implant contains tens of
electrodes with diameters of a few hundred pm, spaced several
hundred .mu.m apart (Humayun 2003). Considering that such
electrodes are much larger than the cells they stimulate, the need
for implants with hundreds or thousands of much smaller electrodes
is apparent. To match the intrinsic resolution of the visual
system, an advanced implant would devote one electrode to every
ganglion cell. This requires that each electrode be similar in size
to a ganglion cell (about 5-20 .mu.m). Instead of 4 affecting
hundreds or thousands of cells, each electrode would evoke a few
spikes in a few retinal ganglion cells.
[0038] The success of the next generation of implantable devices
will be tied to our understanding of how to activate neurons with
extracellular electric stimuli applied to the retinal surface
through electrodes that approach cellular dimensions. Little is
known about the parameters which would permit reliable retinal
stimulation with small electrodes. When the electrode surface area
is reduced, current density and charge density increase rapidly,
and high charge densities are known to cause tissue damage by
electrochemical reactions (Pollen 1977; Brummer et al. 1983;
Tehovnik 1996). A detailed in vitro analysis of small electrode
stimulation is thus a prerequisite for developing such implants for
use in human patients.
[0039] A comprehensive literature review reveals that the
feasibility of stimulation with arrays of small electrodes in
mammalian tissue has not been adequately tested. The majority of
studies involving retinal stimulation have used needle-shaped
probes with one or two conductors at the end of an insulated rod,
such as platinum wires or concentric microelectrodes. In its
simplest form, such stimulating probes are made of metal wires
several hundred .mu.m in diameter, exposed at the tip and insulated
elsewhere (Doty and Grimm 1962; Humayun et al. 1994; Nadig 1999;
Weiland et al. 1999; Suzuki et al. 2004). Others have attempted to
utilize stimulating microprobes with tip diameters of 25 .mu.m or
smaller (Dawson and Radtke 1977; Wyatt et al. 1994; Rizzo et al.
1997; Jensen et al. 2003). However, the geometry of such probes
differs greatly from the planar disk electrode design developed for
current epiretinal implants. Stimulation, furthermore, is always
limited to a single stimulation site, precluding the study of
stimulation using multiple electrodes and their interaction
effects.
[0040] The use of multi-electrode arrays for retinal stimulation
has been mainly limited to large electrodes with diameters between
100 and 1500 .mu.m (Greenberg 1998; Humayun et al. 1999; Hesse et
al. 2000; Walter and Heimann 2000; Humayun et al. 2003; Rizzo et
al. 2003b). Multi-electrode arrays with smaller electrodes (around
10 .mu.m diameter) have been utilized to stimulate the retina in
the subretinal space (Zrenner et al. 1999; Stett et al. 2000).
Grumet has used an array to selectively stimulate the axons of
retinal ganglion cells, using a separate distant array to record
somatic spikes (Grumet 1999; Grumet et al. 2000). No study has
targeted mammalian ganglion cell bodies for direct epiretinal
stimulation using planar electrodes with surface areas below 200
.mu.m2.5 In this study we establish thresholds for stimulation of
ganglion cells in rat, guinea pig, and primate retina using
electrodes with surface areas of 30-500 .mu.m2 (diameters of 6-25
.mu.m). We then used these parameters to further investigate
frequency dependence, pharmacology, and spatial interaction effects
of stimulation. Our arrays use planar disk microelectrodes very
similar to those utilized in present epiretinal prosthetics, but
smaller by an order or two of magnitude. We conclude our analysis
by discussing the results in the context of the pertinent
literature. Early and preliminary portions of this work have been
presented elsewhere (Sekirnjak et al. 2005).
Methods
Retinal Preparation
[0041] This study used retinal tissue from 55 adult rats, 6 guinea
pigs, and one macaque monkey. The average body weight was 289.+-.5
g for rats (Long-Evans), 420.+-.55 g for guinea pigs, and 4 kg for
the macaque monkey (Macaca radiata). Rodent eyes were enucleated
after decapitation of animals deeply anesthetized with 10 mg/kg
Xylazine and 50 mg/kg Ketamine HCI. Primate eyes were obtained from
terminally anesthetized macaque monkeys used by other
experimenters, in accordance with institutional guidelines for the
care and use of animals. Immediately after enucleation, the
anterior portion of the eye and vitreous were removed in room light
and the eye cup placed in bicarbonate-buffered Ames' solution.
Vitreous removal in rats was aided by a homemade extractor which
allowed for rapid but gentle separation of retina and vitreous gel.
The success rate for vitrectomies performed in this manner was 92%.
Pieces of retina 1-2 mm in diameter (FIG. 1A) were separated from
the retinal pigment epithelium and placed flat on the electrode
array, with the ganglion cell layer facing the array (FIG. 1B). The
tissue was held in place by weighted nylon netting positioned over
the array. The preparation was then mounted on a circuit board
attached to an inverted microscope and continuously superfused at
room temperature with Ames' solution bubbled with 95% oxygen and 5%
carbon dioxide at a flow rate of 2-4 ml/min. Pharmacological agents
(TTX, kynurenic acid, CNQX, AP-5, cadmium chloride) were added
directly to the perfusion solution.
Multi-Electrode Array
[0042] The array consisted of a planar hexagonal arrangement of 61
extracellular electrodes, approximately 0.5.times.0.5 mm2 in total
size (FIG. 1A). These electrodes were used both to record action
potentials extracellularly from ganglion cells (Meister et al.
1994; Chichilnisky and Baylor 1999), and to apply current to the
tissue for stimulation. In some experiments, different neighboring
electrodes were used for stimulating and recording.
[0043] The array was microfabricated on a glass substrate, with
indium tin oxide leads and silicon nitride insulation (Litke 1998;
Litke et al. 2003). Each electrode was formed by microwells (holes
in the silicon nitride layer) which were electroplated with
platinum prior to an experiment (FIG. 1A, B). This was accomplished
by submersing the array in a 0.0025N HCI solution containing 1%
chloroplatinic acid and 0.01% lead acetate and applying voltages of
1-5 V through 10 M.OMEGA. resistors for 10-120 sec. Electrode size
was determined by well diameter (5, 6, 8, 10, 12, or 14 .mu.m) as
well as the amount of platinum deposited in each well. Final
electrode diameter varied between approximately 6 and 25 .mu.m,
with a fixed inter-electrode spacing of 60 .mu.m. The geometric
electrode area (#r2) was used to calculate current and charge
densities; however, platinum tends to deposit in a granular
fashion, rendering the effective electrode area significantly
larger (Mathieson et al. 2004). A circular chamber glued on the
glass plate allowed for perfusion of saline solution. A 4 cm-long
platinum wire loop integrated into the chamber served as distant
ground. All stimulations were performed using a monopolar
configuration (electrode to distant ground).
Electrical Stimulation and Recording
[0044] Unless otherwise noted, experiments were performed on a
setup allowing for simultaneous recording of all 61 electrodes and
stimulation on multiple electrodes. The array was connected to a
circuit board containing two custom-made readout ASICs (Application
Specific Integrated Circuit) which amplified, filtered, and
multiplexed signals from the 61 electrodes and sent them to ADC
cards installed in a PC. The board also contained two computer
controlled ASICs capable of sending current pulses to any
configuration of electrodes (Dabrowski et al. 2005). A dim level of
illumination was maintained during the entire experiment (room
lights or microscope illuminator). Recording and stimulation were
controlled by interface software (Labview). Extracellular
potentials were recorded from all 61 electrodes, digitized at 20
kHz (Litke 1999), and stored for off-line analysis.
[0045] The available discrete stimulation pulse current amplitudes
were: 0.6, 0.8, 1.0, 1.2, 1.5, 1.7, 2.0, 2.3, 2.7, 3.0, 3.3, 4.0,
4.7, 5.3, 6.0, 6.7, 8.0, 9.3, 10.7, 12.0, 13.3, and 16.7 .mu.A
(several threshold curves reported in FIG. 6 were determined using
an earlier stimulus generator capable of delivering current
amplitudes as low as 0.1 .mu.A. This device was not used in
subsequent experiments). The pulse consisted of a cathodic
(negative) current pulse of amplitude A and duration d, followed
immediately 8 by an anodic (positive) pulse of amplitude A/2 and
duration 2d (FIG. 1C). Reported current values always refer to the
negative phase amplitude A. Pulse duration was 0.05, 0.1, 0.2, 0.5,
or 1 ms and always refers to the duration d of the cathodic phase.
All pulses were individually calibrated to produce biphasic stimuli
with zero net charge. The pulse shape could be inverted in time to
yield an anodic-first stimulus. Stimulation frequency was varied
between 0.25 and 300 Hz. Experimental protocol Many ganglion cells
show maintained activity under diffuse retinal illumination (Troy
and Robson 1992) and fire spontaneous spikes in isolated pieces of
retina. Stimulation on a particular electrode was typically
attempted if spontaneous extracellular spikes could be recorded
from that electrode. This approach guaranteed that the electrode
was properly platinized and confirmed that ganglion cells in the
vicinity of the electrode were alive. Typically, at least half of
the platinized electrodes on an array showed spontaneous activity
from at least one cell. Spontaneous spikes were readily
distinguished from evoked spikes since they bore no temporal
relationship to the stimulus pulse, while evoked spikes were locked
to the stimulus onset. Stimulation was typically attempted by using
the lowest current settings and was then increased systematically
if no response was seen. Threshold was defined as the current
setting which produced a spike with nearly every stimulus pulse
(.gtoreq.90% of trials) while stimulating at 1-2 Hz. Latency was
defined as the time between stimulus pulse onset and the first
deflection of the evoked spike. Unless otherwise stated, threshold
current, threshold charge, and threshold charge density always
refer to the negative phase of the biphasic, charge-balanced
stimulus pulse. For pharmacological manipulations, a minimum drug
perfusion time of 5-10 minutes was allowed before responses were
recorded.
Data Analysis
[0046] Multi-electrode data was analyzed offline using Labview,
Matlab, and Igor Pro. Means and group data were calculated in
Microsoft Excel. Images were processed in Adobe Photoshop. 9
Chronaxies were calculated by fitting power functions y=a/x+b and
y=a/xp+b (Lapicque 1907; Ranck 1975; Holsheimer et al. 2000) or
exponentials y=b/(1-e-x/a) (Lapicque 1907; Plonsey and Barr 1988;
Greenberg 1998) to the strength-duration data. The asymptote
(coefficient b) was defined as the rheobase; chronaxie was
calculated as (a/b)1/p, a/b, or aln2 for power and exponential fit
functions, respectively. Given the small number of data points
available for some cells, fit quality and resulting parameters
differed for the individual functions and thus values from all
three are reported in the Results section. Autocorrelations of
evoked and spontaneous spiking were obtained by generating
histograms of spike times and interspike intervals, respectively.
On average, about 37 spikes were used per histogram. Spontaneous
histograms were aligned so that time=0 coincided with the
occurrence of the peak of the first evoked spike.
[0047] Power function fit lines to literature data and R2 values
were calculated in Igor Pro by fitting linear functions to the
logarithmic plots of threshold parameters. Statistical comparisons
were done by performing a Student t-test (two-tailed, equal
variance) with a significance limit of p<0.05. Errors and error
bars reported in this study are standard errors of the mean (SEM),
unless otherwise stated.
[0048] Threshold artifact subtraction To reveal spikes with
latencies of less than 2 ms, a novel digital subtraction technique
was used. Spikes obscured by the stimulation artifact (which
typically lasted for several milliseconds) were made visible by
increasing the stimulation current until a possible spike threshold
was reached. Just below threshold, the recorded traces changed
shape noticeably on about half of the stimulus trials (for example,
a change in curvature or peak height), indicating that a possible
spike hidden inside the artifact was elicited on some trials (see
FIG. 2B). Subsequently, the digital difference between two such
traces was calculated. Since the artifact itself was identical in
both traces, the subtraction cleanly revealed the spike inside the
stimulus artifact. Typically, several traces with and without a
suspected spike were averaged before subtraction to increase the
signal over the noise. This method necessitated recording the
spikes on an electrode immediately adjacent to the stimulation
electrode, since the signal recorded at the stimulation site
usually saturated the amplifier and was not suitable for
subtraction. The results obtained were comparable to or better than
those reported for artifact suppression by local curve fitting
(Wagenaar and Potter 2002). We further verified this technique in 6
cells by applying tetrodotoxin (TTX, 1 .mu.M) to the bath solution.
The stimulus artifact recorded in TTX was then subtracted from the
traces containing obscured spikes. In these cells, the resulting
subtracted spikes were identical to the spikes obtained using the
above method. An example of this is shown in FIG. 4A.
Literature Analysis
[0049] Data from epiretinal stimulation studies were compiled as
follows. Threshold current, current density, charge, and charge
density necessary to elicit a ganglion cell response were median or
mean values as reported in each study. When a list of thresholds
was reported, an average value was calculated; when a range of
thresholds was reported, the minimum or the median value was used;
in some cases, a "typical" value was taken from a representative
example or figure. Whenever possible, a missing parameter was
calculated from reported parameters, for instance threshold current
from reported threshold charge (Humayun et al. 1999; Suzuki et al.
2004), surface area from charge density (Dawson and Radtke 1977),
or current from reported charge density (Nadig 1999). In a few
cases it was not possible to calculate a parameter and it was then
estimated from other publications by the same author or the same
group (asterisks in Table 1). When such substitution was not
possible, the study was not included (Crapper and Noell 1963;
Benjamin et al. 1994; Narayanan et al. 1994; Kuras and Gutmaniene
1997). Several studies were represented by multiple entries when
different values of parameters were reported (electrode size,
duration, pulse shape) or when several drastically different
results were reported for the same parameters (such as for two or
more human subjects). The geometric surface area was calculated
from the reported electrode geometry (circular or rectangular for
planar electrodes: #r2 or I2, cylindrical for exposed wires: #r2h,
conical for cone tips: #r(r2+h2)1/2, spherical for ball electrodes:
4#r2). When two or more electrodes were reported to be coupled
electrically and used simultaneously, the surface area was
multiplied accordingly. When a stimulus consisted of high frequency
pulse trains, the effective pulse duration was taken as the number
of pulses per train times the single-pulse duration (Walter and
Heimann 2000; Laube et al. 2003). Whenever possible, data from
normal animals, not those with degenerated retinas were used.
[0050] For plotting the neural injury limit, cat cortical tissue
data from McCreery et al. (1990) was fit to the equation
log(Q/A)=k-log(Q), where Q is the charge in nC and Q/A is the
charge density in mC/cm2 (Shannon 1992; Merrill et al. 2005). The
data can be fit with a coefficient k varying between 1.7 and 2.0;
both values were used for the injury limit plots in FIG. 13.
Results
[0051] We electrically stimulated pieces of isolated mammalian
retina while simultaneously recording spiking activity in ganglion
cells. The properties of evoked spikes are presented first,
followed by strength-duration relationships, temporal properties,
and the results from multi-electrode stimulation. Stimulation at
individual array electrodes resulted in all-or-none spikes recorded
at latencies between a few hundred .mu.s and tens of ms. Of the 208
successfully stimulated ganglion cells, 189 were from rat, 11 from
guinea pig, and 8 from monkey. Most responses consisted of one or
two spikes, although in some cells later spikes were recorded.
Response Latencies
[0052] We classified spikes with latency .gtoreq.2 ms as
long-latency spikes, and earlier responses as shortlatency.
Latencies above 10 ms were infrequently observed and virtually no
spikes occurred more than 20 ms after stimulation onset. Typically,
only long-latency spikes could be readily discerned since the
stimulus artifact obscured the first few milliseconds of the
recording. FIG. 2A shows two spontaneously firing ganglion cells
and their responses to single stimulus pulses. While the primate
cell (top) responded with a distinct spike at latency 5.6 ms, the
guinea pig response (bottom) was obscured by the stimulus artifact.
To isolate the evoked short-latency spike, a threshold artifact
subtraction method was employed (see Methods). Briefly, the
artifact was selectively eliminated by recording several traces
near threshold and subtracting those traces which did not contain
evoked spikes (FIG. 2B). This method was typically employed when a
neighboring electrode was used for stimulation in lieu of the
recording electrode, since this configuration reduced the artifact
below amplifier saturation levels and allowed the artifact to be
subtracted. The result for the guinea pig cell is shown at the
bottom of FIG. 2B: a spike was revealed at 0.25 ms latency. For 86
spikes in rat, visible without artifact subtraction, the average
latency was 7.6.+-.0.3 ms, while 48 artifact-subtracted spikes had
a latency of 0.73.+-.0.05 ms. Nearly all short-latency spikes
occurred at <1 ms; the shortest latencies recorded in this study
were around 0.2 ms. Latency histograms for both short- and long
latency spikes are shown in FIG. 2C.
[0053] Evoked spikes usually resembled the recorded spontaneous
spikes, but occasionally spikes from a different cell were
elicited. Short-latency spikes in particular tended to be of
identical shape as the spontaneous spikes. This is shown in the
inset of FIG. 2B: the evoked spike resembled the spontaneous spike.
Two further examples are shown in FIG. 4A and in the inset to FIG.
8B. Of 48 subtracted short latency spikes, 42 unambiguously matched
the spontaneous spike.
[0054] To elucidate the origin of long-latency spikes, the method
of digital artifact subtraction was applied to recordings which
contained both short- and long-latency spikes. It seemed possible
that each long latency spike was in fact the second spike of a pair
response and not a solitary spike. Indeed, analysis of 20 cells
revealed that the occurrence of long-latency spikes (6.4.+-.0.3 ms)
was always associated with short-latency responses (0.7.+-.0.1 ms).
An example from guinea pig retina is shown in FIG. 3: while the raw
data traces (A) showed only three long-latency spikes (asterisks),
the artifact-substracted traces (B) revealed that every
long-latency spike was preceded by a short-latency spike at 0.35
ms. Furthermore, an analysis of spontaneous spiking activity showed
that spike doublets spontaneously occurred in this cell. This is
shown at the bottom of FIG. 3B: the autocorrelation histogram of
spontaneous spikes showed a peak at a latency similar to that of
the evoked long-latency spikes. Thus, the evoked spikes occurred
with timing expected from the spontaneous activity. A second
example from rat retina is shown in FIG. 3C for a cell with
long-latency spikes at 7 ms. Spike timing analyses were performed
in a total of 8 cells, with similar results: the spontaneous
interspike intervals matched the typical intervals between short-
and long-latency spikes. These results indicate that some cells
responded to a single stimulus pulse with a spike pair, with the
first spike obscured by the artifact, and that this tendency toward
paired spiking was evident in the spontaneous activity of the cell.
The method of analyzing spike timing was further utilized to
calculate the approximate latency of obscured short-latency spikes
when only long-latency spikes were available. FIG. 3D shows an
example of a cell in which a large stimulus artifact precluded the
use of the artifact subtraction method; only long-latency spikes
were discernible. By aligning the peaks of the two histograms, a
short-latency spike (dashed box) was inferred at times <1 ms.
Similar results were found in 3 cells and suggest that
short-latency responses can be deduced from the observance of
long-latency spikes. Lastly, we compared the spike latencies of
long-latency responses evoked with stimulation electrodes of
different diameters, which ranged from 6 to 25 .mu.m in this study.
No systematic difference was observed when large rather than small
electrodes were used and average latencies for the smallest
electrodes (6-9 .mu.m) were similar to the largest (20-25 .mu.m):
8.2.+-.0.7 ms and 7.1.+-.0.4 ms, respectively (p>0.2; n=32
cells).
Pharmacological Manipulations
[0055] Several ion channel antagonists were applied to the
perfusion solution to further investigate the evoked responses. To
ascertain that the observed spikes were of neuronal origin, the
sodium channel blocker tetrodotoxin (TTX, 1 .mu.M) was added to the
perfusion solution. In 3 guinea pig and 17 rat cells, all spikes
(both spontaneous and evoked) disappeared within seconds of drug
application, confirming their identity as neuronal action
potentials. An example is shown in FIG. 4A: application of TTX
eliminated evoked short-latency spikes, leaving only the pulse
artifact. When this artifact was subtracted from the control
responses, an evoked spike was revealed (FIG. 4A). The waveform of
this spike did not differ from that derived by threshold artifact
subtraction or the spontaneous spike recorded at this electrode
(FIG. 4A, bottom). Similar results were found in 5 cells
[0056] To investigate whether the applied current pulses acted
directly on ganglion cells or involved more distant cells with
synaptic connections to the recorded cell, blockers of synaptic
transmission were added to the perfusion solution. A combination of
the following agents was used: the broad spectrum glutamate
antagonist kynurenic acid (1 mM), the NMDA-receptor blocker APV
(400 .mu.M), and the AMPA-receptor blocker CNQX (75 .mu.M). FIGS.
4B and C show examples of responses from two cells, recorded before
and after addition of the blockers. Spike shapes, latencies, and
response rates were unchanged, even in the cell with spikes at
latency 15 ms (FIG. 4C). No systematic differences between spikes
elicited in control and drug conditions were observed in any of 9
cells. These findings suggest that ganglion cells were activated
directly, not trans-synaptically, and further corroborate the
notion that apparent long-latency spikes (such as in FIG. 4B, C)
are not solitary spikes, but part of a two-spike response.
[0057] In separate experiments, the calcium channel blocker cadmium
chloride (100-250 .mu.M) was applied to the perfusion solution to
abolish synaptic transmission (not shown). In 10 cells, evoked
spikes were still observed after drug application, indicating that
the observed spikes were not produced by mechanisms involving
calcium-dependent synaptic transmission. Minimal thresholds and
spatial spread Spikes were evoked in ganglion cells using currents
between 0.6 and 5 .mu.A. When stimulated with 0.1 ms pulses, the
average threshold current for 78 rat cells stimulated under similar
conditions was 0.81.+-.0.03 .mu.A, corresponding to a charge of
81.+-.3 pC and a charge density of 0.073.+-.0.005 mC/cm2. In many
cases, the lowest current setting of our stimulator (0.6 .mu.A)
yielded a superthreshold response, indicating that the reported
average thresholds may be overestimated.
[0058] Thresholds were lowest when the recording electrode, rather
than a neighboring electrode, was also used for stimulation. To
examine whether spikes could be elicited by stimulating at a
distance from the recording site, electrodes immediately adjacent
to the recording electrode were used to stimulate. FIG. 5A shows
average results for 8 cells, stimulated with 1, 3, or 6 adjacent
electrodes. The goal was to elicit the same long-latency spike
using the different configurations of stimulation sites shown.
Thresholds for spike initiation increased several-fold, depending
on the number of active electrodes. In particular, when a single
neighboring electrode was used for stimulation, about 3 times more
current was needed compared to stimulation at the recording
electrode. This indicates that a resolution of the order of the
electrode spacing or finer (.ltoreq.60 .mu.m) can be achieved with
minimal threshold stimulation.
[0059] The preceding results suggest that stimulation using low
stimulus amplitudes (<0.1 mC/cm2) usually affected only cells in
the vicinity of the stimulation electrode. To further verify this,
in 35 low amplitude stimulation experiments (average charge density
0.071.+-.0.004 mC/cm2), all electrodes surrounding the stimulation
electrode were inspected for evidence of evoked spikes which
differed in latency, shape, or reliability from the ones recorded
on the center electrode. Such additional spikes would indicate
recruitment of neurons at nearby locations. Of 186 neighboring
electrodes analyzed for long-latency spikes, only one showed an
additional evoked spike. However, additional evoked spikes were
frequently seen on surrounding electrodes when the current was
increased several-fold, suggesting recruitment of cells tens of
.mu.m distant, consistent with the results shown in FIG. 5A. Still
higher currents sometimes elicited spikes on non-neighboring
electrodes, more than 150 .mu.m away from the stimulation
electrode.
[0060] To further investigate spatial spread of activation, a more
detailed analysis was performed to detect short-latency spikes
around the stimulation electrode. In 4 experiments, we applied TTX
and subtracted the averaged stimulus artifact on each electrode
individually to reveal additional short latency spikes, as in FIG.
4A. In 2 such experiments with a stimulus strength of around 0.1
mC/cm2, no short-latency spikes were found outside the 60 .mu.m
radius around the stimulation electrode. In 2 further experiments
stimulated at 0.21 and 0.35 mC/cm2, spikes were detected as far
away as 160 .mu.m. One example of strong-stimulus stimulation is
illustrated in FIG. 5B: while the majority of electrodes on the
array recorded no evoked spikes, four separate responses were
elicited in the vicinity of the stimulation site. The spikes from
these stimulated cells were each detected on 2 or more electrodes
and the electrode recording the largest spike amplitude can be used
to infer the approximate location of the soma. Most evoked spikes
(circles, squares, diamonds) were recorded within 60 .mu.m of the
stimulation electrode, but one cell was detected nearly 160 .mu.m
away (triangles). These results show that the radius of stimulated
ganglion cells can be controlled by adjusting the stimulus
strength.
[0061] The above results were obtained by applying cathodic-first
pulses (FIG. 1C). For most cells, thresholds were slightly higher
when the anodic phase was delivered first: in 18 cells stimulated
with 0.05 or 0.1 ms anodic-first pulses, spike thresholds were
115.+-.5% of the thresholds measured using cathodic-first
pulses.
Strength-Duration Relationship
[0062] The current required to elicit a spike depended strongly on
pulse duration. In all three species tested, higher currents were
required to evoke a spike when shorter pulses were applied.
Durations were varied from 50 .mu.s to 1 ms and several resulting
strength-duration curves are shown in FIG. 6. In the examples
plotted here, electrode diameter, stimulation configuration, and
spike latency differed considerably across cells, resulting in a
wide spread of threshold curves. Nevertheless, the slopes of these
curves were similar in monkey, guinea pig, and rat, indicating that
the threshold-duration relationship was independent of the species.
To characterize each strength-duration curve by a time-constant and
an asymptote, power functions or exponentials were fit to the data
(see Methods). Rheobase is defined as the asymptote of the fit
curve (Ranck 1975; Loeb et al. 1983) and chronaxie, the classical
measure of responsiveness of a neuron, as the duration at which the
threshold current is twice the rheobase (Lapicque 1907). The
average chronaxie of 34 cells such as those shown in FIG. 6 was
407.+-.45 .mu.s for 1/x fits, 338.+-.81 .mu.s for power fits, and
212.+-.28 .mu.s for exponential fits. The average rheobase was
0.51.+-.0.12 .mu.A, 0.60.+-.0.11 .mu.A, and 0.76.+-.0.16 .mu.A,
respectively. The fit quality was generally highest for power fits.
Seven cells were from monkey, 7 from guinea pig, and 20 from rat
and all responses were long-latency spikes. When the group of 20
rat cells was divided into those stimulated with the recording
electrode and those stimulated at a neighboring site, no difference
in chronaxie was found (p>0.5).
[0063] To facilitate comparison of thresholds in a single species
and to illustrate the influence of stimulation electrode position,
FIG. 7 shows averaged data from 25 ganglion cells in rat. The cells
were stimulated under identical conditions using electrodes with
similar diameters (average 10.4.+-.0.5 .mu.m) and only long-latency
responses were included. The solid line plots thresholds for
stimulation at the recording electrode (13 cells), while the dashed
line shows results from stimulation at a neighboring electrode (12
cells). As in FIG. 5A, eliciting a spike required several-fold
higher currents when the site of stimulation was at an adjacent
electrode. Since the charge delivered during the cathodic phase of
the pulse is often used as a measure for stimulation strength, the
inset plots charge thresholds for pulse durations up to 0.2 ms:
charges were consistently below 200 pC, corresponding to charge
densities below 0.25 mC/cm2.
[0064] To further corroborate the above notion that short- and
long-latency spikes constitute doublet responses, we measured
strength-duration curves of responses with latencies <2 ms. The
average strength-duration relationship in 14 cells with
short-latency spikes (latency 0.69.+-.0.08 ms) was similar to that
of long-latency responses: chronaxies determined from fit curves
(as above) were 571.+-.149 .mu.s, 299.+-.52 .mu.s, and 311.+-.113
.mu.s; none of these values was significantly different from
long-latency chronaxies (p>0.1; n=45 cells). These results
suggest that the same neuronal element was excited in both short-
and long-latency responses.
Frequency Dependence
[0065] To mimic natural spike trains, a retinal implant must be
capable of delivering pulses and evoke spikes at a wide range of
stimulation frequencies. Furthermore, continual stimulation at
higher frequencies may be a requirement for generating sustained
percepts of light. We tested stimulation at pulse frequencies of up
to several hundred Hz. To examine high-frequency responses, two
closely spaced pulses were applied, with the inter-pulse interval
corresponding to frequencies of up to 200-300 Hz. High-frequency
stimulation was deemed successful when spikes were evoked following
the second stimulus pulse. Pulse pairs were applied for 10-20
seconds at intervals of 0.5 seconds and at stimulus strengths of
about twice threshold. In 9 cells tested, spikes were evoked on the
second pulse on 99.+-.1% of trials at 100 Hz, and 94.+-.4% at 200
Hz. Three cells were stimulated with 300 Hz pulse pairs and all
responded at >90% of trials. All responses were short-latency
spikes (latency 0.8.+-.0.2 ms). Data from such an experiment is
shown in FIG. 8A: a super-threshold 300 Hz pulse pair reliably
produced two short-latency spikes, more clearly seen in the
artifact-subtracted traces shown below. A superposition of several
stimulus trials is shown to demonstrate repeatability. In 4 cells,
TTX was added to the bath solution to facilitate artifact
subtraction and spike detection (not shown).
[0066] To test responses to brief periods of sustained
high-frequency stimulation, 9 cells were continuously stimulated
for 5-20 seconds at frequencies of up to 100 Hz. FIG. 8B shows an
example of responses to over 70 stimulus pulses near spike
threshold, delivered at 32 Hz. Short-latency spikes were evoked on
roughly half of the trials (arrow) and were used to subtract the
artifact (inset). 32 Hz stimulation evoked spikes indistinguishable
from those produced by 2 Hz stimulation or spontaneous
activity.
[0067] Response rates, defined as the number of evoked spikes in a
stimulation period, were measured at sustained pulse frequencies of
up to 100 Hz at stimulus strengths of about twice threshold (FIG.
8C). Short-latency spikes (closed symbols) showed a slight
reduction of the response rate at 50 Hz (<20%) and a significant
drop at 100 Hz. We also observed a gradual reduction in spike
amplitude throughout the stimulation period at frequencies above
about 32 Hz (not shown). Strikingly, long-latency responses (open
symbols) were robust only up to 5 Hz and virtually no spikes were
observed above 10 Hz. This observation was corroborated in one cell
with a short-latency spike (0.7 ms) which was followed by a spike
at latency 5 ms: stimulation at low frequencies consistently evoked
both responses, while only the short-latency spike was observed at
frequencies above 8 Hz.
[0068] We conclude that short-latency spikes can be reliably evoked
in ganglion cells at pulse frequencies up to about 50 Hz and that
late spikes are suppressed at moderate frequencies.
Sustained Stimulation
[0069] Chronic retinal implants must be capable of delivering
effective stimulation pulses over a period of many hours each day.
To determine whether sustained low-frequency stimulation could
reliably evoke spikes, we extended our stimulation period to the
longest duration that was experimentally feasible.
[0070] Two cells were continuously stimulated for 30 minutes, and
two additional cells for 4.5 hours. The longest sustained
stimulations were performed using 0.8 .mu.A pulses with 0.1 ms
duration, delivered at frequencies of 1-2 Hz, and corresponding to
a charge density of about 0.04 mC/cm2 per pulse. FIG. 9 shows an
example of spikes evoked before and after a 4.5 hours stimulation
period: the cell showed robust responses after having been
stimulated with over 16,000 pulses. A slight increase in threshold
and spike latency (about 20%) was noted at the end of the
stimulation period.
Multi-Electrode Stimulation
[0071] To generate artificial vision, a functional retinal implant
requires independent activation of many closely-spaced electrodes.
To investigate responses to spatial stimulation patterns, the
multi-electrode array was utilized to stimulate at several
electrodes simultaneously. Our goal was to demonstrate that
simultaneous activation of two or more nearby electrodes did not
influence each other. If that were the case, spikes elicited during
multi-electrode stimulation should not differ in threshold, shape,
or number from individual stimulations.
[0072] We selected 7 sites which clearly showed evoked long-latency
spikes when stimulated individually. These evoked spikes differed
in spike shape and latency, but had similar thresholds. All 7
electrodes were subsequently activated simultaneously using 0.8
.mu.A pulses (0.1 ms duration) and the responses recorded. FIG. 10A
shows spikes evoked at these sites and their locations on the
array. Simultaneous stimulation evoked seven distinct responses on
seven spatially disparate electrodes.
[0073] To establish that the spikes evoked by simultaneous
stimulation did not differ from those evoked by individual
stimulation, traces recorded at each electrode under both
conditions were compared. Two examples are shown in FIG. 10B:
individually evoked spikes (single) were identical to
simultaneously evoked spikes (all) for both electrodes shown here.
Furthermore, FIG. 10B demonstrates that stimulation at neighboring
electrodes evoked independent responses. While this was expected
given the low currents utilized here (see FIG. 5A), these data
clearly establish that adjacent electrodes (1 and 2) did not
influence each other during simultaneous stimulation. Only a
small-amplitude deflection was recorded on electrode 1 at the
latency of the spikes seen on electrode 2 (arrowhead), indicating
that the cell stimulated by electrode 2 was probably located close
enough to electrode 1 to be recorded as small spikes. To further
ensure spatial precision, all 22 inactive electrodes surrounding
the 7 active stimulating electrodes were inspected for spikes.
While four adjacent electrodes showed small spikes that were
recorded on one of the 7 stimulation electrodes, none recorded new
spikes. Multi-electrode stimulation was performed 5 times using
various electrode arrays and spatial patterns, with results very
similar to the data presented above. We conclude that evoking
independent spikes on multiple electrodes spaced 60 .mu.m apart is
feasible with minimal cross-electrode interaction.
Electrode Size
[0074] This study employed arrays with electrodes that varied in
diameter between 6 and 25 .mu.m. To determine the influence of
electrode size on thresholds, stimulation results were compared in
a set of 86 cells for which the exact platinum disk diameters of
the stimulating electrode was measured. FIG. 11 shows thresholds as
a function of electrode diameter, both for cells stimulated using
pulse durations of 0.1 and 0.05 ms. Current, charge, current
density, and charge density are plotted. All spikes were long
latency spikes (average 7.6.+-.0.3 ms).
[0075] Threshold current and charge (FIG. 11A, B) increased by a
factor of 2-3 between the smallest and the largest diameters,
indicating that with smaller electrodes, less current and charge
injection was necessary to elicit spikes in ganglion cells.
Conversely, current density and charge density (FIG. 11C, D) was
drastically decreased for electrodes larger than 10-15 .mu.m. Note
that the lowest threshold values plotted here may have been
overestimated since the minimal available current setting (0.6
.mu.A) often evoked a super-threshold response.
Discussion
[0076] This study used dense arrays of small-diameter electrodes to
electrically stimulate rat, guinea pig, and primate retina. We
described the responses of individual ganglion cells to a wide
range of pulse configurations and spatial stimulation patterns and
showed that effective stimulation is feasible with high temporal
and spatial precision. Our findings imply that the electrode size
of future epiretinal prosthetics may safely approach the cellular
dimensions of retinal ganglion cells.
Evoked Spikes
[0077] Long-latency spikes (>2 ms) were readily observable,
while spikes with shorter latencies could only be observed with
digital artifact subtraction. Note that this classification into
"short" and "long" latencies differs somewhat from that of other
researchers: Jensen et al. (2005) defines short as 3-5 ms and long
as .ltoreq.9 ms; Stett et al. (2000) classifies spikes at latencies
1-10 ms as early and spikes at 10-20 ms as delayed. The definition
of Crapper & Noel (1963) is more similar to the one used in
this study: immediate spikes were defined as those around 0.5 ms
latency and later responses as 5-15 ms. Early spikes. The earliest
observed responses occurred within several hundred .mu.s of
stimulation onset and probably represent the immediate activation
of the ganglion cell spike generator. Actual latencies are in
effect shorter than the reported values by 50-100 .mu.s due to a
delay introduced by the amplifier circuitry. Further, if measured
from the termination of the cathodic phase, true minimal latencies
observed in this study amount to 50-150 .mu.s. While these values
are lower than the latencies in many studies, sub-millisecond
spikes have been reported by a few authors (Crapper and Noell 1963;
Grumet et al. 2000). Late spikes. We showed that evoked spikes at
longer latencies are preceded by short-latency spikes.
[0078] Long-latency spikes can thus be used to infer short-latency
spikes even when the magnitude of the pulse artifact confounds
direct observations of early spikes. However, it is possible that
thresholds for long-latency spikes are systematically
overestimated, since short-latency responses typically occur at
lower stimulation currents than doublet responses. Multi-spike
responses are consistent with known intrinsic firing properties of
ganglion cells, in which doublets or triplets in spike trains occur
with interspike intervals of several ms (e.g. Devries and Baylor
1997). In our study, interspike intervals ranged from 4 to 16 ms,
with the majority of long latency spikes occurring around 5-7 ms
after the short-latency spikes. We favor the interpretation of
long-latency spikes as part of a doublet response over other
suggestions such as conduction delays (Jensen et al. 2005b) or
intracellular charging mechanisms for the following reasons: the
small electrodes and currents used here make activation several
millimeters from the recording site exceedingly unlikely; in each
case tested, every long-latency spike was preceded by a
short-latency spike; and earlier studies have not employed artifact
subtraction methods, thus seeing only later spikes.
[0079] Furthermore, long-latency spikes are only observed at
stimulation frequencies below 10 Hz, suggesting that higher
repetition rates suppress multi-spike bursts. Stimulation
thresholds Safety of stimulation. An important prerequisite of
implantable stimulators is their capability to deliver current that
is safe, yet efficient. Unsafe stimulation can originate from two
sources: electrochemical destruction of the stimulating electrode
(such as corrosion), and neural tissue damage induced by toxic
products near the electrode or by neuronal hyperactivity. Several
electrochemical safety limits have been proposed, such as the
often-stated non-gassing limit of 0.3-0.4 mC/cm2 for platinum
electrodes (Brummer and Turner 1977). More recently, limits as low
as 0.1 mC/cm2 for cathodic stimulation with platinum electrodes
have been recommended (Rose and Robblee 1990).
[0080] Thresholds for tissue injury in cortex have been shown to
arise from the synergistic interaction between charge and charge
density: as the charge is increased, the charge density for safe
stimulation decreases (McCreery et al. 1986; McCreery et al. 1990;
Merrill et al. 2005). The McCreery data show that no histologically
detectable damage is produced with low-charge stimulation (<50
nC) even when the charge density is >1 mC/cm2, while for pulses
delivering a higher charge (1 .mu.C), the damage threshold is
<0.1 mC/cm2. In the absence of detailed threshold measurements,
concerns have been raised regarding the feasibility of using
small-diameter electrodes in human patients, since they have been
suggested to require much higher charge densities for threshold
stimulation than large electrodes (Brummer et al. 1983; Loeb et al.
1983; Greenberg 1998). However, we found in this study that
threshold stimulus pulses are characterized by low currents (around
1 .mu.A), low charge injection (around 100 pC) and low charge
densities (around 0.1 mC/cm2) despite the small electrode size.
Several cells had threshold charge densities of less than 0.03
mC/cm2, an order of magnitude lower than the platinum electrode
safety limit. Furthermore, while we have used the geometric
electrode area to calculate current and charge densities, the
effective electrode area likely was significantly larger.
Electroplated platinum tends to deposit in granular surface
structures which greatly increase the area of metal in contact with
the solution. It has been reported that the fractal-like platinum
deposits can increase the surface area by up to 100 times (Kim and
Oh 1996; Mathieson et al. 2004). Thus, all density values reported
here should be considered upper limits, further reducing the
likelihood of electrochemical electrode damage.
[0081] Our results complement data recently reported for
small-diameter needle electrodes, which have described threshold
charge densities between 0.15 and 0.3 mC/cm2 (Wyatt et al. 1994;
Rizzo et al. 1997; Jensen et al. 2003; Wilms et al. 2003). Distance
between electrode and cells. One factor contributing to the low
thresholds in this study is the tight contact between electrodes
and tissue. This was a requirement in our experiments since
extracellular spikes cannot be recorded without close juxtaposition
of the retina to the array. Novel techniques to minimize the gap
between retina and epiretinal implant are being developed (Schanze
et al. 2002; Johnson et al. 2004) and may ensure close contact in
future prosthetic devices. Optimal electrode size. We observed
lower threshold current and charge for the smaller electrodes in
this study than for the larger ones (see FIG. 11). However, the
resulting charge density is increased for smaller electrodes. As
electrode diameter drops below about 10 .mu.m, the decrease in
surface area outweighs the current decrease. It has been suggested
that for electrodes smaller than the cellular size (about 10 .mu.m)
the electric field is concentrated in too small an area for
effective stimulation (Palanker et al. 2004). Thus, electrode
diameters around 10-15 .mu.m may be the optimal size for selective
single cell stimulation and might be an ideal compromise between
excellent spatial resolution and high charge density. This size
range would also have less stringent requirements on the distance
between electrode and cells, since stimulation with <10 .mu.m
electrodes is disproportionately more sensitive to this distance
(Palanker et al. 2004). Clearly, this issue will need to be
re-addressed once technical advances in retinal prosthetics call
for even smaller electrodes as the ratio of electrodes to ganglion
cells approaches 1. Spatial resolution. One consequence of the low
required stimulation strengths was the exceedingly localized nature
of stimulation: excited cells were limited to a narrow radius
around the stimulating electrode and pharmacology experiments
further confirmed that ganglion cells were directly activated:
spikes were not suppressed in the presence of CNQX, APV, and
kynurenate, which block excitatory transmission in the retina
(Fujimoto and Toyoda 1991; Stett et al. 2000). This is a much more
local effect than can be achieved with larger electrodes: indirect
spikes sensitive to synaptic blockers have been reported for 125
.mu.m electrodes (Jensen et al. 2002; Ziv et al. 2002) and larger
electrodes (Greenberg 1998; Shimazu et al. 1999). Our results from
simultaneous stimulation using multiple electrodes further confirm
that the current spread in the plane of the electrode array is
small enough to allow for independent activation of cells using
neighboring stimulation electrodes.
[0082] Thresholds increase with the distance between stimulating
and recording electrode on the array (see FIG. 5A). The observed
increases are similar to those of the cathodal stimulation map
reported by Jensen et al. (2003): stimulating about 60 .mu.m away
from the center of the receptive field required 2-8 times more
current to elicit a spike. We conclude from these observations that
retinal implants with small electrodes can achieve a high spatial
resolution, since the low applied currents activate single (or at
most a few) ganglion cells.
Chronaxies and Site of Activation.
[0083] The use of pulses significantly longer than chronaxie
contributes little to the evoked response, stipulating pulse
durations smaller than chronaxie to insure that most of the applied
charge contributes to evoking a response (Tehovnik 1996). Thus,
from the chronaxies measured in this study (around 100-400 .mu.s)
we conclude that optimal pulse durations should not exceed this
range. The measured values are similar to those reported in other
studies (Crapper and Noell 1963; Greenberg 1998; Grumet et al.
2000; Jensen et al. 2005b) and can further be used to identify the
neuronal element most likely excited by electrical stimulation. Our
chronaxies match those reported for activation of axons (Nowak and
Bullier 1998; Grumet et al. 2000; Holsheimer et al. 2000), since
cell bodies and dendrites have chronaxies of 1-10 ms (Ranck 1975;
Holsheimer et al. 2000). Because the initial axon segment near
somas is more excitable than cell bodies (Porter 1963; Nowak and
Bullier 1998; Greenberg et al. 1999; Schiefer and Grill 2002), the
juxtasomal electrode used here likely activates this initial region
on the axon and action potentials subsequently back propagate a
short distance to elicit the recorded somatic spike. Computational
models suggest that excitation occurs near the junction of ganglion
cell soma and axon or slightly more distal on the axon (Fohlmeister
and Miller 1997; McIntyre and Grill 1999; Schiefer and Grill
2002).
[0084] It is difficult to experimentally rule out the activation of
passing axons, in particular since tests designed to identify
antidromic responses (Fuller and Schlag 1976) would not distinguish
between initial axon segment excitation and more distant axon
activation. Nevertheless, activation of passing axons is deemed
less likely by the fact that ganglion cell axons have high
thresholds away from their initial segment (Loeb et al. 1983).
Since in retinal ganglion cells of most mammals (including human
and rat) the axon remains unmyelinated within the retina, sodium
channels are found uniformly throughout the distant unmyelinated
region (Boiko et al. 2003). At the initial axon segment, however,
the density of sodium channels is exceptionally high (Wollner and
Catterall 1986), with clustering of the Nav1.6 subunit in
particular (Boiko et al. 2003). This difference in channel density
between the initial and distant region can amount to an order of
magnitude or more (Ritchie et al. 1976; S. R. Levinson, personal
communication). Like the nodes of Ranvier in myelinated fibers
(McIntyre and Grill 2000), the initial portion of unmyelinated
axons constitutes the most likely site of electrical excitation,
perhaps at the "thin segment" 10-40 .mu.m from the cell body
(Fohlmeister and Miller 1997; Boiko et al. 2003).
Comparative Literature Analysis
[0085] To discuss data from this study in the context of previous
work, a comprehensive review of the published literature was
composed. Table 1 summarizes 32 studies that have reported
epiretinal stimulation thresholds. These studies span several
orders of magnitude in electrode size and can thus be used to
elucidate threshold trends. Several key parameters have been
graphed in FIG. 12, along with best fit lines and correlation
coefficients. To facilitate comparison of different electrode
geometries across studies, threshold parameters were plotted
against the geometric electrode surface area (see Methods).
Variability.
[0086] Several factors contribute to the relatively wide scatter of
points in FIG. 12. Threshold was defined inconsistently from study
to study, spanning the range of 50 to 90% probability of eliciting
a spike, cortical recordings, and human percept reports.
Furthermore, while the majority of studies utilized charge-balanced
biphasic pulses, several reported monophasic stimulation (typically
cathodal), leading to lower thresholds in some cases. Moreover,
human studies (open symbols) usually involved degenerated retinas,
while animal studies were typically performed on normal tissue.
Finally, studies which measured retinal responses by monitoring
cortical activity (triangles) may have overestimated spike
thresholds in ganglion cells, since the concerted activity of many
cells is typically required for a cortical response.
Parameter Trends.
[0087] Both threshold current and charge (FIG. 12A, B) decrease
dramatically as electrode size is reduced. Correlation was highest
for threshold charge, since it takes into account both current and
pulse duration, which varied across studies. These trends confirm
that smaller electrodes require several orders of magnitude lower
currents to elicit responses. They also mirror the results found
within this study over a much more narrow range of electrode sizes:
current and charge thresholds were small for small stimulating
electrodes and large for large electrodes (FIG. 11A, B). Current
density also increases somewhat when electrode size is reduced
(FIG. 12C), but such a trend is not seen in the plot of charge
density thresholds (FIG. 12D). While the large variability does not
permit an accurate fit to the data, there is no definitive change
of charge density with electrode size, such as is seen for
threshold charge over 4-5 orders of magnitude. In fact, charge
density is virtually independent of electrode size for electrodes
smaller than 104 .mu.m2 (disk diameter about <100 .mu.m). This
trend is in contrast to the increased charge densities observed for
the smallest electrodes used in this study (see FIG. 11D), which is
probably due to an effect restricted to electrodes smaller than
.about.10 .mu.m diameter (see above).
[0088] Representative data from this study (monkey and rat) have
been included in FIG. 12 and fit well with the trends established
by the published literature. Our data substantiate the main
conclusion from this analysis: small electrodes require much less
charge injection for threshold stimulation than larger electrodes,
but the accompanying increase in charge density is almost
negligible.
Stimulation Safety.
[0089] Since both charge and charge density must be considered when
discussing stimulation safety (Merrill et al. 2005), FIG. 13 shows
a plot of both parameters for the same set of studies. Two types of
safety limits were included (dotted lines): the often used
electrochemical limits for platinum electrodes (0.35 and 0.1
mC/cm2) and the limits for neural injury from cortical stimulation
data (k=1.7 and 2.0; see Methods). To show the spread of thresholds
measured in this study, FIG. 13 includes data points for all cells
stimulated using 0.05 ms pulses and electrodes with diameters
between 6 and 25 .mu.m (crosses). Most human studies (and several
animal studies) fall near or outside of the safe region formed by
the limit curves, possibly because degenerated retina requires
higher currents to produce phosphenes in humans. This plot further
validates our claim that the small electrodes used in this study
can safely stimulate mammalian retina: except for the data
collected using the smallest electrodes, most thresholds are well
within all safety limits.
Outlook
[0090] The purpose of this study was to elucidate basic stimulation
parameters to test whether a future generation of implants could
incorporate a design using significantly smaller electrodes than
are presently available. We used planar microelectrode arrays that
closely resemble those currently in use for chronic human testing
(Humayun et al. 2003), but contain much smaller electrodes at a
much smaller electrode spacing. We suggest that future implants
could directly activate ganglion cells instead of affecting large
areas of retina by indirect stimulation, making possible a
reasonable spatial resolution of artificial sight. One can envision
high-resolution arrays containing thousands of stimulation sites
with diameters around 10-20 .mu.m and separation between electrodes
of 20-60 .mu.m. While years away, results from this study suggest
that there is no fundamental hindrance to the feasibility of such a
device. Once implanted, the stimulus parameters can be adjusted to
stimulate individual or small overlapping groups of ganglion cells,
depending on the desired phosphene size. By utilizing low currents,
activation of axon bundles can be avoided. As a next step toward
the development of such implants, further experiments using small
electrodes with degenerated retina are warranted.
[0091] FIG. 14 shows a perspective view of the implanted portion of
the preferred retinal prosthesis. While the invention has broad
applicability to neural stimulation, the preferred embodiment is a
retinal prosthesis. A flexible circuit 1 includes a flexible
circuit electrode array 10 which is mounted by a retinal tack (not
shown) or similar means to the epiretinal surface. The flexible
circuit electrode array 10 is electrically coupled by a flexible
circuit cable 12, which pierces the sclera and is electrically
coupled to an electronics package 14, external to the sclera.
[0092] The electronics package 14 is electrically coupled to a
secondary inductive coil 16. Preferably the secondary inductive
coil 16 is made from wound wire. Alternatively, the secondary
inductive coil 16 may be made from a flexible circuit polymer
sandwich with wire traces deposited between layers of flexible
circuit polymer. The electronics package 14 and secondary inductive
coil 16 are held together by a molded body 18. The molded body 18
may also include suture tabs 20. The molded body 18 narrows to form
a strap 22 which surrounds the sclera and holds the molded body 18,
the secondary inductive coil 16, and the electronics package 14 in
place. The molded body 18, suture tabs 20 and strap 22 are
preferably an integrated unit made of silicone elastomer. Silicone
elastomer can be formed in a pre-curved shape to match the
curvature of a typical sclera. However, silicone remains flexible
enough to accommodate implantation and to adapt to variations in
the curvature of an individual sclera. The secondary inductive coil
16 and molded body 18 are preferably oval shaped. A strap 22 can
better support an oval shaped coil.
[0093] The preferred prosthesis includes an external portion (not
shown) which includes a camera, video processing circuitry and an
external coil for sending power and stimulation data to the
implanted portion.
[0094] Accordingly, what has been shown is an improved method of
stimulating neural tissue for improved response to brightness.
While the invention has been described by means of specific
embodiments and applications thereof, it is understood that
numerous modifications and variations could be made thereto by
those skilled in the art without departing from the spirit and
scope of the invention. It is therefore to be understood that
within the scope of the claims, the invention may be practiced
otherwise than as specifically described herein.
* * * * *