Over the past decade, electrocorticography (ECoG) has been used
for a wide set of clinical and experimental applications. Recently, there
have been efforts in the clinic to adapt traditional ECoG arrays to
include smaller recording contacts and spacing. These devices, which
may be collectively called “micro-ECoG” arrays, are loosely defined as
intercranial devices that record brain electrical activity on the sub-
millimeter scale. An extensible 3D-platform of thin film flexible micro-
scale ECoG arrays appropriate for Brain-Computer Interface (BCI)
application, as well as monitoring epileptic activity, is presented. The
designs utilize flexible film electrodes to keep the array in place without
applying significant pressure to the brain and to enable radial
subcranial deployment of multiple electrodes from a single craniotomy.
Deployment techniques were tested in non-human primates, and
stimulus-evoked activity and spontaneous epileptic activity were
recorded. Further tests in BCI and epilepsy applications will make the
electrode platform ready for initial human testing.
Electrocorticography (ECoG) has become an increasingly standard
clinical tool that has also spurred experimental exploration. Clinically
approved ECoG arrays are relatively large and new smaller scale
devices (micro-ECoG) are being developed using a variety of fabrication
methods, materials, and designs.1,2Micro-ECoG arrays are poised to
improve on both the degree of invasiveness and signal quality by utilizing
flexible electrode deployment through smaller craniotomies and by
increasing the spatial and spectral resolution with higher density and
smaller electrodes.3To date, most clinical micro-scale ECoG devices
have been built upon existing macro-ECoG platforms, and have not
leveraged contemporary microfabrication processes, which afford the
advantages of mass production, reproducibility and flexibility. While there
have been a number of previous studies that have described several
micro-fabrication techniques for producing micro-ECoG electrodes for
high resolution mapping experiments,3there remain a number of hurdles
towards the adoption of this technology in a clinical setting.
Standard ECoG fabrication, involving discrete electrodes and wires
embedded in silicone, has been pushed to its practical limits to approach
the micro-ECoG spatial scale.4-7Further advances may involve bio-
microelectromechanical systems (bio-MEMS).1-3,8Because the develop -
ment of BioMEMS devices utilizes photolithographic techniques, many
electrode variations can be produced in one batch, greatly reducing the
time required for the development and testing cycle. The interface or
connection between the electrode array and the amplifier is another
important consideration. Both awkwardly large connectors1and
meticulous hand soldering2,8have been used in previous systems. An
easily assembled platform with high-density solderless connections that
leverage the flexibility, and other material properties of thin-film micro-
ECoG arrays, would be a crucial component of a BioMEMS-based
electrode array. In this paper we describe the clinical application-specific
design, fabrication, flexible deployment, and in vivo testing of a custom
micro-ECoG platform in an animal model. Multiple electrode designs
derived from the same platform were explored to demonstrate the
creative synthesis of technology and clinical insight.
ELECTRODE PLATFORM OVERVIEW
With multiple clinical and experimental applications in mind, a
general-purpose electrode platform was designed to address the unique
properties and constraints of brain tissue and the subcranial space. The
surface of the brain is not static. It shifts and deforms continuously,
particularly following surgery,9so flexible circuit technology that can
move with the brain is an excellent match. Built on a layer of polymer that
is tens of micrometers thick, thin-film electrodes can be designed to
conform to a shifting brain. There are a wide range of available polymers
(parylene C,10Polyimide,11Liquid-crystal polymer12), metals (platinum,2
gold13) and possible electrode geometries, and we created a central plat -
form that is extendable to a wide range of materials and enables rapid
geometrical design iteration.
The transcranial platform design centralizes all external
connections, and can be assembled quickly with flexible electrodes
using press-fit sockets. The platform has three parts: a printed circuit
board (PCB) for connections to an amplifier and other circuit
components; a plastic adapter ring for anchoring the platform to the
skull; and a flexible electrode array to interface mechanically and
electrically with the brain (Figure 1). The electrode is made with gold
and platinum on polyimide, a bio-compatible polymer that has several
advantageous properties. Polyimide is hydrophilic which allows it to
adhere strongly to moist brain tissue, flexible to accommodate the
gyrated cortex, and durable for long-term implantation durations. The
CLINICAL EEG and NEUROSCIENCE©2011 VOL. 42 NO. 4
A Micro-Electrocorticography Platform and
Deployment Strategies for Chronic BCI Applications
Sanitta Thongpang, Thomas J. Richner, Sarah K. Brodnick, Amelia Schendel, Jiwan Kim, J. Adam Wilson,
Joseph Hippensteel, Lisa Krugner-Higby, Dan Moran, Azam S. Ahmed, David Neimann, Karl Sillay and Justin C. Williams
From the Department of Engineering (S. Thongpang, T.J Richner, S.K. Brodnick, A.
Schendel, L. Krugner-Higby, J.C. Williams), and the Enzyme Institute (L. Krugner-Higby),
University of Wisconsin, Madison, Wisconsin; the Korea Electronic Technology Institute (J.
Kim), Gyeonggi-do, Korea; the University of Cincinnati (J.A. Wilson), Cincinnati, Ohio; the
Department of Neurological Surgery (J. Hippinsteel, A.S. Ahmed, D. Neimann, K. Sillay),
University of Wisconsin Hospital and Clinics, Madison, Wisconsin; and the Department of
Biomedical Engineering (D. Moran), Washington University, St. Louis, Missouri.
Address correspondence or requests for reprints to J.C. Williams, University of
Wisconsin-Madison, 1550 Engineering Drive, Room 3142, Madison, WI 53706.
electrode array can be placed either epidurally or subdurally. A PCB
mates the electrode to small high-density connectors (Mill-max and
Omnetics in this work), and a plastic adapter ring holds the PCB and
electrode array in place over the craniotomy. The housing ring is milled
out of an autoclavable X-ray transparent plastic, Ultem (McMaster-
Carr), to fit snugly in a 19 mm burr hole, similar to that used in many
standard clinical neurosurgical procedures.
The micro-ECoG electrode array was fabricated in a class 100
cleanroom. The electrode fabrication process must yield a biocom -
patible and non-toxic electrode array, and the entire process was
designed to avoid toxic chemicals. A photodefinable polyimide substrate
was chosen over other processes that use harmful Hydrofluoric acid
(HF) as an etchant for patterning. Additionally, photodefinable polymide
reduces the number of overall steps in the lithographic process.
Polyimide serves as the insulating substrate for the electrode and gold
traces which connect to each electrode site. Platinum was deposited on
the traces and electrode sites to improve electrical characteristics,
including decreasing the electrode impedance. Each electrode array
consisted of 16 or 32 electrode sites with the expectation that in some
instances multiple electrode arrays could be implanted in a single
craniotomy. The photolithographic transparency mask sets were
designed in Adobe Illustrator, and were printed on high-resolution
transparency masks (Fineline-Imaging, Colorado Springs, CO). The
shapes and sizes of electrodes were systematically varied, yielding a
wide array of electrode designs (Figure 2C).
Briefly, the electrode fabrication process (Figure 2A) follows. After
cleaning a 4-inch silicon wafer with acetone, a 200 nm thick layer of
aluminum was evaporated (CHA-600 E-beam Evaporator) onto the
surface as a sacrificial layer. Polyimide (HD 4110, HD Microsystems)
was spin coated and patterned using photolithography (Mask aligner:
MA6 Suss MicroTec) following the manufacturer’s recommendation to
form a 12 µm thick base layer of polyimide (Figure 2A(a)). The poly -
imide was subsequently cured at 375 °C under a nitrogen atmosphere
for 2 hours (Cooke Oven). A metallization step (10 nm Cr/200 nm Au/20
nm Pt) defining the interconnection pads, connecting traces, and
electrode sites was performed using electron beam evaporation (CHA-
600 E-beam Evaporator) and a lift-off process. The second layer of
polyimide was then spin coated, patterned and cured under the same
conditions as the first base layer. Before releasing the device, oxygen
plasma (RIE Unaxis 790) was used to clean the surface for 2 minutes
(250 W, 20 Torr) ensuring that the remaining polyimide and solvent
residue on the electrode sites and contact pads were completely
removed (Figure 2A(d)). The thin film devices were released using an
anodic metal dissolution technique. In brief, the wafer was immersed in
a sodium chloride solution (2 M) with Pt counter electrodes under an
applied voltage of 0.8 V that allowed the dissolution of the aluminum
layer and release of the polyimide film. To complete the process, the
devices were soaked in deionized water, cleaned and dried.
Thin-film electrodes naturally roll into coils, form into springs, and
fold into 3-dimensional geometries (Figure 2B). These deformable
structures are well suited to interface with the nervous system at many
levels, including peripherally, via a catheter delivery system,
intracortically, or on the cortical surface with or without an intact dura
mater. Thin-film electrodes were first applied to peripheral nerves rolled
into a conduit or perforated to allow nerve fibers to regrow through the
electrode.14Other neural targets include deep brain structures, to which
the devices may be delivered via catheter, bundled with electrodes and
other sensors,15or intracortical targets using flexible substrates,16
similar to the traditional silicon based intracortical electrode.17The
micro-ECoG array depicted in Figure 2D(i) uses spring action to
interface epi- or subdurally while attached to the platform that fits in a
19 mm burr hole.
The 2-D array depicted in Figures 2D(i), 3A and 3B has a
rectangular grid of 32 electrode sites (300 µm diameter) spaced 1 mm
apart. The impedances of these electrodes in this array are nominally
5-10 kOhms at 1 kHz. The traces run up 4 flexible legs to connect to
the PCB. At the PCB, each electrode site has a corresponding multi-
leafed contact that is sandwiched between a pair of electrical sockets
(Mill-Max) (inset of Figure 3E). The leaves bend under pressure from
the conical shoulder of the socket pins, avoiding high stress regions
that would arise if the contact were simply annular, and that would
result in cracking of the contact pad metal layer. In principle only a
single leaf is needed, so the 6-leaf design adds redundancy to each
contact, increasing the yield. Mechanical connectors like these are
especially necessary when assembling the electrode into a 3-
dimensional platform, since soldering would be a challenging
alternative due to the difficulties inherent in aligning and maintaining
the individual arms while under bending stress.
ANIMAL AND SURGERY
Terminal surgeries were performed for initial electrode deployment
tests and to obtain acute recordings in rhesus or cynomolgus macaques
(n = 5). All terminal and survival animal procedures were approved by
the Institutional Animal Care and Use Committee (IACUC) at the
University of Wisconsin-Madison. In this experiment, the electrode was
deployed through one craniotomy, while a video was recorded with an
endoscope through a second craniotomy (Figure 4A). Due to the
flexible nature of the electrode, the array was inserted using a small
guide wire, which attaches to the front of the electrode (see small hole
at the distal end of the electrode in Figure 4D) and subsequently
retracted after the electrode is in place. This electrode design has the
potential of being used as a presurgical epilepsy diagnostic tool to find
a focal area before initiating the resection treatment.18
Chronic implant surgery
Survival surgery procedures approved by the University of
Wisconsin-Madison IACUC were done in rhesus macaques (n = 3)
using the most promising designs and deployment methods. Rhesus
macaques were chronically implanted with a miniaturized electrode
CLINICAL EEG and NEUROSCIENCE ©2011 VOL. 42 NO. 4
A schematic representation of the 3D-electrode platform on a primate brain
with the alternative ground and reference pads shown surrounding the
perimeter of the electrode array.
array over the right sensorimotor cortex region, centered 14 mm
anterior of the intra-aural line and 14 mm laterally off the sagittal suture
(Figure 5A). The monkeys were first anesthetized with Ketamine (10
mg/kg) for surgical preparation including vital monitor positioning and
intubation. Isoflurane (0.5-5%) was used for anesthetic induction and
maintenance during the non-recording parts of the procedure. Following
anesthetic induction, they were placed into a stereotaxic frame, and the
scalp was shaved and prepped with alternating povidone iodine and
alcohol. After the initial incision, exposed skull was cleaned and dried.
To obtain interoperative recordings, a continuous rate infusion of
ketamine: fentanyl at a dosage of 10-20 mg/kg/hr for ketamine and
0.0001-0.001 mg/kg/hr for the fentanyl was initiated just before the skull
was opened and isoflurane concentration maintained at 0.5% or less. A
small hole was drilled through the skull until dura was exposed, and a
stainless steel screw (size 4-40) was used for attachment of a ground
wire. Then a small craniotomy (19 mm in diameter) was made using a
spherical burr (stryker) designed for use with a high-speed surgical drill.
After exposing the dura, the micro-ECoG array was placed epidurally
onto the sensorimotor area of the cortex. GelFoam (Pharmacia and
Upjohn Co, New York, NY) was placed on top of the array.
Coupling device and brain
To harness the spring-action of the polymer, and allow the brain to
move freely without damage, a multi-leg electrode design was used.
Very small forces cause the legs to bend, and these same forces help
keep the device in contact with the dura or pia mater without breaking
CLINICAL EEG and NEUROSCIENCE©2011 VOL. 42 NO. 4
A) Electrode fabrication process using
lithography to pattern Polyimide, metal
evaporation to deposit the traces and
electrode sites, and anodic metal
dissolution to release the electrodes. B)
Various electrode designs that show the
flexibility of the polyimide polymer
substrate, making it suitable for a wide
range of applications. C) The layout of a
variety of electrode designs with different
shapes, electrode sizes and inter-
electrode spacing. D) The realization of
the design layout in (C) with
corresponding platform and connectors.
A) Four legs allow the 2-D array to translate along 3 axes and rotate about an additional 2 axes. The
flexible legs maintain only a minimal amount of pressure to ensure close contact between the electrode
and tissue. B) The hydrophilicity of the polyimide helps the electrode adhere to the dura or pia mater. C)
and D) An x- axis translation of the brain tissue phantom (0.6% agarose gel) up to 1 cm, representing the
constant contact of the electrode array with dura due to small brain movement in all x-, y-, z- and 3
rotational axes in vivo. E) A cross-section of the electrode platform shows an assembled electrode and
PCB sitting atop a plastic adapter ring inside a craniotomy. The adapter ring can be temporarily fixed to the
cranium. The inset shows the multi-leafed electrode contact to the electrical sockets on the PCB. F)
Diagram illustrating the flexible legs (longer leg cable than those shown in A-E), which flare out radially
under the skull to cover a larger area of cortex. G) and H) Electrode sites come in contact with the dura as
the electrode is deployed. The dura can be left intact (pictured) or removed.
electrode contact with the brain surface. This is important for chronic
ECoG monitoring, as in clinical applications the brain can move relative
to the skull. In addition to the forces imparted by the electrodes legs, the
hydrophilic nature of the polyimide substrate helps the electrode
maintain its position on the brain. Polyimide is unique in that it becomes
increasingly hydrophilic with repeated contact to water.19Thus, the initial
hydrophilicity of polyimide allows the electrode to be placed easily
during deployment, and the electrode then stays in place as the
polyimide substrate’s hydrophilic structure nature binds reversibly to
moist tissues such as the dura or pia mater (Figure 3A and 3B). The
multi-leg design leverages the flexible and hydrophilic properties of
polyimide to maintain contact even as the brain translates relative to the
cranium in 6 axes (x-, y-, z-, and 3 rotational axes). Figures 3C and 3D
illustrate the translation in the x-axis for approximately 1 cm while the
electrode array remains in contact with the brain tissue phantom model
(0.6% agarose gel20). The y- and z- and 3 rotational axes showed similar
results to those shown in Figures 3C and 3D by moving the brain model
in different axes while keeping the electrode array stationary. Sudden
translations of 5 mm did not affect electrical properties of the electrode
nor did rotations (data not shown). The device has greatest translational
flexibility in the axis normal to the electrode surface, and somewhat
reduced rotational flexibility along the same axis, but this matches the
expected movements of a patient’s brain. The legs of the electrode
exert a tiny amount of pressure to keep the electrode surface in contact
with the brain. Only an estimated 0.52 Pa is applied when the legs are
in the compressed position (Figure 3B), which is equivalent to 52 mg of
force per square centimeter. As can be observed from Figure 3B, this
force is insufficient to even break the surface tension of water, and is
unlikely to cause any depression of the brain surface. It does however
produce enough of a downward force to maintain the electrode in
contact with the cortical surface during relative brain motion.
Using an unfurling technique (Figure 3F), the electrode array could
be deployed several centimeters distal from the craniotomy if the cable
arms are designed to be longer than that shown in Figure 3A-3E. This
technique would allow the electrodes to be deployed from a craniotomy
away from the site of injury, minimizing further trauma and scarring at
the injury site. During unfurling, the electrode never came close to the
critical radius of curvature that would damage its electrical integrity.
Bench tests showed that more than 95% of electrode sites maintained
electrical continuity at a 1 mm radius of curvature (data not shown).
Images of the unfurling electrode deployment method depicted in
Figure 3E are shown in Figures 3G and H. These images, taken during
a terminal procedure in a cynomolgus macaque, show that the
electrode sites come into contact with the brain surface as the
electrode strip unfurls. With the same principle in mind, a multi-leg
flared electrode was also prototyped (image not shown). Multi-leg
electrodes can flare out radially from a small 5 mm burr hole to cover
an area of cortex 3 times as large as the unfurling electrode, and
enable a minimally invasive surgical procedure for mapping or
diagnosing any abnormal brain function in a specific area or provide for
a large area of coverage in a chronic neural recording application.
Alternative electrode designs and
deployment methods for clinical applications
Weighed against a more extensive craniectomy, deployment
through a moderately sized craniotomy is a less invasive option. For
example, to keep the craniotomy relatively small, but extend the spatial
reach of the electrode, a 16-site strip electrode was designed (Figure
2C(iii) and 2D(iii)) that may be deployed epidurally. This electrode
might be appropriate for monitoring epileptic activity (Figure 4D) in a
large area if its radial position were changed. With some practice, it can
be deployed leaving the dura intact. After removing any dural
adhesions that may be present, the electrode can be inserted with the
aid of an elevator. These deployment techniques were tested in a non-
human primate spontaneous epilepsy model.
During the non-human primate epilepsy model procedure, record -
ings from the micro-ECoG were able to detect spontaneous epileptic
activity in an anesthetized rhesus macaque that had a previous open-
chambered cranial implant for several years. The implant was removed
and the animal was used as a breeder for three years prior to terminal
recording. The monkey had no history of grand mal seizures, but less
severe epileptic episodes may have been present. Spontaneous seizure
activity under ketamine/fentanyl continuous rate infusion (ketamine, 20
mg/kg/hr and fentanyl, 1.5 µg/kg/hr), the anesthetic combination used in
this study, was characterized by large depolarizations (maximum of 2
mV) seen across multiple sites. Interestingly, the signal amplitude de -
creased over the span of four neighboring electrode sites spaced 1 mm
CLINICAL EEG and NEUROSCIENCE©2011 VOL. 42 NO. 4
Long strip electrode suitable for acute recording. A) A representation of a strip
electrode inserted through a very small (10 mm) burr hole simultaneously with
an endoscope for real time imaging of electrode position. B) and C) Real time
images of electrodes in contact with the dura, taken from the scope which was
situated on the opposite side of the location of electrode entry. D) Epileptic
activity observed from an acute recording using the same long strip electrode
array. Spontaneous epileptic activity recorded under anethesia from 4
neighboring electrode sites shows the signal amplitude of up to 2 mV, droping
over a spatial scale resolvable only by micro-ECoG arrays.
apart—a spatial scale unresolvable by traditional ECoG or EEG
technology (Figure 4D).
Multiple strip electrodes, combined with a 2-D array, could be
deployed from a single craniotomy, resulting in electrode densities
superior to the current standard of care. As a thin and flexible polymer
device, it was easily removed when recordings were completed. This
electrode design utilizes the flexibility of the electrode to reduce the
required craniotomy size. The flexibility of the electrode also helps
reduce the risk of dural bleeding and reduce other surgical
complications. The high resolution of the microECoG spacing could
allow precise localization of pathological neural activity and could help
inform a neurosurgical team in clinical applications. For clinical
applications, especially epilepsy monitoring, a device that has higher
spatial representation, such as micro-ECoG where the focal change of
cortical physiology may span as few as 1-2 electrode sites, could allow
the diagnosis to be more accurate. Recent studies have suggested
that macro-epileptic seizure activity might be the result of the
coordinated synchrony of numerous spatially distinct “micro-seizure”
domains, whose activity may proceed the large scale activity
traditionally used as clinical evidence during seizure monitoring.21
Chronic in vivo tests
Initial signal recording procedures were obtained using standard
methods to elicit stimulus-evoked potentials, both interoperatively (as
described above) and chronically.22For chronic recordings, light
anesthesia (Ketamine, 10-15 mg/kg) and analgensia (Buprenorphine,
0.005-0.03 mg/kg) were administered intramuscularly or subcutane -
ously 30-60 minutes before the start of an intramuscular stimulation
recording session and electrical stimulation evoked potentials were
recorded. Multiple (n = 60) short cathodic electrical stimulation (2 ms
pulse) of current 2-4 mA was applied with 25 gauge needle electrodes.
Neural activity was sampled at 3051 Hz using an RZ2 amplifier (Tucker-
Davis Technologies, Alachua, FL). A band-pass filter was set at 0.1 to
200 Hz. Longitudinal baseline recording for power spectrum analysis of
the micro-ECoG over 8 weeks was performed. The power spectra of
channels 1-16, within the area of interest for the foot electrical
stimulation, are illustrated in Figure 5B. Baseline recordings, evoked
potentials and electrode impedance were obtained for 5 minutes weekly
to ensure the long-term stability of the micro-ECoG device. The
baseline power spectrum was calculated using the Matlab toolbox
Chronux 2.023from 30 seconds of data from each recording session. A
CLINICAL EEG and NEUROSCIENCE©2011 VOL. 42 NO. 4
Long-term micro-ECoG recording comparison. A) An electrode array and platform design that fits a 19 mm craniotomy in comparison with a quarter coin. B) Position
of the electrode array on a non-human primate brain (Rhesus macaque), situated over sensorimotor cortex. C) Power spectrum (with 95% confidence intervals) of an
example electrode site (highlighted in blue in (Figure 5B)) showing stability of the recorded baseline signal over an 8 week period. D) Power spectra of channels 1-16
over 8 weeks show long-term stability in long-term implant (60 Hz line noise omitted). E) Evoked potential response due to contralateral hind limb electrical stimulation
(2.5 mA, 2 ms) compared to control (no stimulus) (week 2, 6 and 8).
1.57 second window, a time-bandwidth product of 3, and 5 tapers were
used to calculate the power spectrum. The 95% confidence intervals
were found with the jackknife resampling method to compare weeks 1
and 8 (highlighted in blue in Figure 5B). The power spectrum appeared
to be stable in most of the electrode sites (Figure 5D). In sensorimotor
rhythm, a mu band (8-12 Hz) and gamma band (30-40 Hz) were
present and recorded by the ECoG device. Note that the frequencies
of 58 to 62 Hz were omitted in Figure 5D as they represent main
electrical frequency noise. The power spectra of channels 17-32 are
not shown for space considerations, but also appear to be stable.
With 1 mm spacing, our non-human primate electrode arrays
showed a distinctive evoked potential amplitude difference between
adjacent electrode sites when the contralateral foot was electrically
stimulated (Figure 5C). Higher evoked potentials were observed over
the stimulated part of the cortex and then faded out slowly towards
the anterior (channels 17-32). Repeated electrical pulse (2.5 mA, 2
ms) stimulation evoked potentials up to 100 µV in amplitude. This
was large compared to control baseline recordings in which no
stimulus was applied. Over 8 weeks, the evoked potential responses
were similar in both amplitude and waveform (Figure 5E). This
verifies that our epidural micro-ECoG has an acceptable signal
quality and a high spatial resolution, and does not have signal
properties that are drastically different than previously reported from
both micro- and macro-electrode recordings.
Thin-film neural interfaces are becoming a significant research
technology with many potential applications in clinical settings including
advanced neuroprosthetics. With their high degree of flexibility, thin-film
electrodes can flex with the brain, assemble into 3-D structures, and
deploy through reduced surgical openings. Precise fabrication tools
utilizing photolithographic techniques enable high electrode densities
for greater spatial resolution than standard ECoG arrays.
Micro-ECoG arrays may some day replace existing clinical ECoG
technologies for many applications as the current generation of ECoG
arrays with discretely wired electrode sites in silicone have reached
practical size limits. Electrodes of many sizes and shapes are easily
fabricated in parallel using thin-film processes using rapid prototyping
process, and so the set of possible electrode designs, the “design
space,” is quite large. Many electrode designs can be developed with the
same fabrication process to fit any application including rodent and non-
human primate studies, and eventual human clinical trials. Based on
signal considerations, the electrode site size and spacing can be
matched to the spectral and spatial characteristics of the signal,
respectively, or to the anatomical layout of the gyri. This may someday
make it possible to customize electrodes for individual patient needs in
clinical applications. On the mechanical side, the thickness of the
polyimide substrate can be increased for added strength during
deployment, or reduced for deployment strategies requiring additional
electrode flexibility. Theoretical and computational modeling can help
guide the design process, but due to the rapid prototyping nature of the
manufacture process, there are also opportunities to directly test new
devices. 3-D electrode designs have not yet been utilized in the micro-
ECoG community, but 3-D electrodes can make superior contact with a
moving brain. To keep the electrode array in constant contact with the
brain surface (i.e., to self seat), the highly hydrophilic nature of the
material is crucial for accommodating the brain movement inside the
skull. In addition, electrodes can be easily repositioned, as the adhesion
mechanism is reversible. This platform also allows the surgeon to deploy
the device precisely and safely with minimal difficulty, because the whole
platform can be placed without having to directly handle the delicate
electrode array. The electrode itself does not need to be directly
manipulated as is the case with most penetrating MEMS electrodes.
High-yield solderless connections can also reduce the man-hours
required for assembly of multiple devices, providing a viable pathway
towards the scale-up necessary for clinical manufacturing. It would be
very difficult to assemble the device shown in Figure 3, if the connections
at the end of each of the arms had to be individually soldered or epoxied.
The combination of the polyimide surface properties and its
flexibility allows for tight coupling of the electrode with the convoluted
cortical surface. As this technology moves towards clinical
applications, the degree of flexibility is important to accommodate the
gyro-cephalic brain.24Other polymers (e.g., parylene C10and silk
fibroin24) also have desirable material properties. The optical transpar -
ency of parylene is desirable in a research setting where imaging or
perhaps optogenetics25may be used.
We have demonstrated that efficient recordings of abnormal elec -
trical activity may be obtained from a very small area of the brain with a
long strip electrode (Figure 4). Traditional epilepsy monitoring requires
temporarily implanting an FDA approved standard clinical ECoG array
that involves a potentially risky surgical procedure. The strip electrode
can be applied not only for epilepsy monitoring, but also for cortical map -
ping. A large cortical area can be mapped from a small burr hole by
changing the long strip electrodes radial position. None of the previous
studies has exploited this promising feature and integrated it into a
particularly versatile micro-ECoG design. Other studies1-2proposed lar -
ger micro-ECoG arrays that cover up to a half of the hemisphere. These
designs may be practical for cortical mapping, but for neuroprosthetics,
a small electrode array that covers the sensorimotor cortex may be suffi -
cient to decode limb movement activities.26-27In addition, the 3-D platform
electrode array also demonstrated long-term signal recording stability in
a chronic rhesus macaque model. The power spectra of the baseline re -
cording and the evoked potentials of all sites were stable over 8 weeks.
Before moving forward to clinical trials, an acceptable chronic implant
electrode has to address issues related to robustness, signal quality and
stability, and biocompatibility. Further work will be necessary to fully char -
acterize the relationships between signal quality, electrochemical
measures and histological reactions to the presence of the device.
Micro-ECoG arrays may someday rival the utility of silicon and
micro wire based neuroprosthetics interfaces that are currently
employed. Arm movements in 3-D have been decoded from spiking
data28but similar 3-D movement decoding has not yet been achieved
with micro-ECoG. A better understanding of the micro-ECoG signal
and its relationship to other brain signals may improve its efficacy in
BCI applications. Its overall utility is still promising though, since
penetrating electrodes often lose the ability to isolate spikes over a
period of weeks or months.16,29
EEG and ECoG have traditionally had broader clinical application
than penetrating electrodes, because of the relatively high risk encoun -
tered by penetrating electrodes. Relatively less invasive micro-ECoG
devices may someday replace existing macro-ECoG technologies. The
flexibility, customizability, and quick assembly of the platform presented
above leads to many clinical and experimental paths. BCIs rely on high
quality information-dense signals extracted from the cortex. Our non-
human primate data show similar utility for supporting results for evoked
activity applications, such as P300 or similar motor cortical BCI and
CLINICAL EEG and NEUROSCIENCE©2011 VOL. 42 NO. 4
epileptic monitoring. Further validation in animal models of epilepsy and Download full-text
BCI will help to prepare the electrode platform for initial human testing.
Advances in the thin-film electronics industry have created new
opportunities for micro-ECoG devices. The next puzzle-piece borrowed
from industry could be integrated amplifier circuits housed directly on
the electrode substrate. A complete micro-ECoG platform will someday
include the electrode sites, amplifiers,3analog to digital converters,
wireless transmission and power all on a single flexible thin-film device
which can be inserted subcranially through a very small craniotomy. In
principle, all of the components are ready, but skilled technology
integration and collaboration will be required to continue to push this
promising technology forward.
This work was supported by grants from the National Institutes of
Health (NIH NIBIB 1R01EB009103-01 and 2R01EB000856-06) and in
parts by the Wallace H Coulter Foundation Institutional Translational
Partnership and a refinement grant from the Center for Alternatives to
Animal Testing (terminal experiments) to LKH.
DISCLOSURE AND CONFLICT OF INTEREST
S. Thongpang, T.J. Richner, S.K. Brodnick, A. Schendel, J. Kim, J.A.
Wilson, J. Hippensteel, L. Krugner-Higby, D. Moran, A.S. Ahmed, D.
Neimann, K. Sillay, J.C. Williams have no conflicts of interest in relation
to this article.
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