REVIEW OF PERIPHERAL NERVE ELECTRODE
TECHNOLOGY
Dhonam Pemba
Introduction
The electrode serves as the critical component of Functional Electrical Stimulation, Brain
Machine Interfaces and Neural Prosthesis. The choice of electrode use in any stimulation system
would depend on the application. Electrode selectivity varies from measuring surface EMG
signals to signal neuron spikes, while invasiveness is inversely proportional to selectivity.
Surface electrodes are the least invasive while penetrating intraneural electrodes are the most
invasive, accordingly their respective resolution increases with proximity to the neurons.
Electrode types can be grouped according to their location; on the surface, epimysial,
intramusclar, extraneural, epineural, interfasciular, intraneural.
I.
Surface Electrodes
Surface Electrodes usually consist of stainless steel disks placed on the skin external to
the desired muscle to stimulate. The electrode are able to record motor unit action potentials and
EMG signals[1–3] and have shown promise in rehabilitation applications[4–7]. Although the
electrodes demonstrate capability of detecting muscle activity, their use in for finer motor control
would be impractical. The electrodes would require significant current to stimulate muscle for
functional motor movement which would cause muscle fatigue and tissue damage. In addition,
the large area of the electrodes would produce very low selective stimulation, and accurate
muscle recoding for finer control would require significant post signal processing and
computation to accurately differentiate motor activity readings[8], [9]. Magnetoneuraphy is an
alternative non-invasive method to record magnetic fields in the action in the peripheral
nerves[10] but this system would again only be practical for stimulation and depend significantly
on post processing[11].
II.
Muscle Electrodes
To overcome the disabilities of surface electrodes, researchers and clinician have place
electrodes directly in contact with the muscle fiber. Methods of muscle stimulation include
epimysial electrodes surgically sutured to the muscle[12–14][15], and intramuscular electrodes
that are placed inside the muscle[11], [16–19]. The most widespread design of intramsuclar
electrodes consist of fine wire inserted into targeted muscle. A novel intramuscular electrode
system called the Bionic Neurons(BIONs) was developed to be a self-contained unit. The BIONs
are implanted into the muscle with a needled, and are individually addressable, receive power
externally and stimulate motor units.[20], [21]. Although epimysial and intramuscular electrodes
are able to stimulate and record motor unit action potential, there are several issues preventing
success in neuroprosthetic applications. Similar, to surface electrode, the muscle electrodes
possess low selectivity due to their distance from the nerve. Even though the electrodes are
placed on the muscle, the stimulation actually evokes action potential in the nerve rather thus a
higher current would be required stimulate the nerve muscle impedance as opposed to an
electrode placed on the nerve directly. In addition, due to the graded recruitment of muscle[22],
[23], muscle stimulation would result in muscle fatigue and discomfort[24], [25]. Finally, to
achieve complex and fine motor control, electrodes would need to be placed on each individual
muscle making the system inconvenient and high maintenance.
III.
Nerve Electrodes
Stimulating the nerve directly has significant advantages of surface, epimysial and
intramuscular electrodes. By directly stimulating the nerve, less current can be used while
achieve higher selectivity, the electrode can be placed away from the contracting muscle thereby
increasing durability, and avert graded recruitment issues to reduce muscle fatigue. Electrodes
are either placed outside the epineurium, between epineurium and perineurium, insde the
perineurium or through the fascicule.
Extraneural Electrodes
Extraneural electrodes contact the nerve but are placed outside the epineurium, and including
epnineural, helical, book, and cuff electrodes.
Epineural electrodes are electrodes that are sutured outside of the nerve. The electrodes are
wires, discs or buttons and had success with less complex stimulation needs such as pain therapy
or respiration.[26–29] The disadvantages with them are their lower relative selective to other
neural electrode and more complicated surgery to ensure stable suture.
G.S Brindley pioneered the book electrodes, which have demonstrated success in bladder
control by contacting the sacral spinal roots. The device is constructed from implanting
electrodes inside channels that are cut out of block biocompatible insulation. The electrode
design resembles book, where the gaps between the pages resemble the channels[30–34].
Although the book electrode has had success with spinal root stimulation, the device is too bulky
and not designed peripheral nerve stimulation. Recently, a group has attempted to miniaturize the
book electrodes for placing individual nerve fascicles between the pages rather than spinal
roots[35]. However the miniature book electrode would require the dissection of the nerve into
individual fascicles and manual placement in the grooves.
Cuff electrodes are arguably the most widely used nerve configuration[36–56], they were
designed to improve upon the epineural electrodes by improving the signal to noise and ease of
implantation. In general the cuff design composes of an electrode in contact with the epineurium
that is encircles the never with biocompatible insulation. Original designs did not utilize
micromachining technology and were silicon rubber tubes with a slot[37], blocks of insulating
material with a channel[36][57], or flexible insulating material wrapped around a nerve[54],
[58].Figure from Loeb et al summarizes in detail the fabrication of several other nerve cuff
deigns[59]. With the advancement of thin film micromachining technology, electrode designs
have included polyimide as the substrate material, and improved electrode materials and
designs[40], [60–65]. Most of the polyimide cuff electrodes are fabricated by depositing and
patterning metal between layers of polyimide, and then shaped into a cuff. A critical issue with
cuff electrodes is to seal the slot in which the nerve was inserted which was originally done with
a silicon elastomer[49], [50], [66],suture[54], [58] or tension[37]. Recent designs have attempted
to close the slot by using hinges[45], or advanced polymers with self-coiling abilities[51], [52].
In addition to sheets of biocompatible insulation encircle the nerve, spiral and helical cuff
electrodes have been designed[67–73]. Helical electrodes improve fitting while decreasing
compression due to self-sizing nature of the spiral. Although cuff electrodes have widely been
used, stimulation superficially[74] thus the greatest disadvantages are in the difficulty in
selective action of particular fascicles.
To improve of selectivity of nerve cuff electrodes, the cuff electrode was redesigned to
reshapes the nerve to bring central axon population closer to the electrodes[75], [76]. The flat
interface nerve electrode(FINE) were silicon elastomer shaped with a rectangular center opening
1.4 times the cross-sectional area of the nerve(figure [76]). The electrodes have shown
effectiveness in controlling gross movements such as knee extension[77], [78] but not complex
movements such as ankle control.[79] More recently, Poly (DL-lactic-co-glycolic) acid (PLGA)
was added to the a FINE electrode to improve control of closure[80]. The PLGA was attached to
a stretched FINE electrode, and upon degradation the FINE electrode deformed to its original
equilibrium state.
Figure 2.1: FINE electrode (Left), SPINE electrode (Right).
Longitudinal electrodes
The inventors of the FLAT electrode had previously developed an interfasicular electrode
which makes contacts which penetrates the epinerium. The slowly penetrating interfascicular
nerve electrode (SPINE) is similar to a traditional cuff electrode but includes blunt electrodes
that penetrate epineurm. The SPINE electrode is an interfascicular electrode that is in-between
extraneural and intraneural electrodes. Electrodes that penetrate the perineurium are considered
intraneural while electrodes outside the epineurium are identified as extraneueral. Intraneural
electrodes improve selectivity and signal to noise ratio and reduce cross talk by placing the
electrodes inside the nerve and in contact with individual fascicules[81]. Intraneural electrodes
possess superior selectivity to extranenural electrodes[74], [82]and have shown the most promise
for use in complex motor control neuroprosthesis[83–89]. Restoring hand function is one of the
most complex motor movements for a neuroprothesis because high selectivity would be needed
to stimulate specific muscles without recruiting others. Success in of stimulating independent
specific muscles along with natural hand movements has beech achieved by implanting an
intraneural electrode[87]. Intraneural electrodes have shown the ability to detect motor activity in
amputee and have helped provide tactile and proprioceptive feedback to discriminate objects
without visual cues[90]. By implanting intraneural electrodes, stimulation was able to evoke
discrete tactile and proprioceptive sensation in a missing limb.[84], [85]
Longitudinally Implanted Intrafascicular electrodes(LIFE) were originally thin insulated
wires securely inserted under the perenerium which stimulate and record from individual
fascicules. Electrode areas were exposed by heating the insulating layer and the penetrating end
was electrochemically sharpened with potassium nitrite[91]. To reduce the mechanical mismatch
between tissue and electrode, and decrease cellular encapsulation, flexible substrates were used
over the original wire[83], [92]. Polymer fiber was first as a flexible alternative to metal
wires, polymer-based longitudinal intrafascicular electrode (polyLIFE) were fabricated by
sputter deposition of metals over Kevlar fiber, and dip coating of silicone insulation[83], [93].
Although polyLIFEs improved on flexibility over traditional LIFEs, the active sites precision had
not improved. Microfabrication technology was implemented in the newest generation of LIFEs
that allowed control over electrode geometry. Fabrication of thin film LIFE (tfLIFE) was similar
to polyimide cuff electrodes, where lithography based patterning of metal was sandwiched
between two polymer layers[92], [94], [95]. LIFEs offered precise stimulation and record from
individual fascicules, but to control complex motor movements various fascicules would need to
be selectively stimulated. Thus to selectively stimulate specific fascicules within a nerve fiber
bundle would require several implanted LIFEs.
Transverse Penetrating
Selective stimulation of various fascicules within a nerve fiber bundle would be
impractical with LIFEs and unachievable with extraneural electrodes, therefore the best option
for electrodes for complex motor function controlled would be transversely penetrating probes.
a. Wires
Hodgkin and Huxley first demonstrated that action potentials in a squid axon could be
detected with a simple wire inserted into a micro glass pipette[96], [97].The basic microwires are
easily fabricated and has demonstrated success in neural activity from cortex and peripheral
nerves [98–103]. Wire based microelectrodes are constructed from small diameter wires that are
then electrochemical sharpened, and finally insulated[99], [104]. Although, these
microelectrodes have had success recording neural activity, the fact that selective and
independent activation of different fasciculus would require several electrodes limits its
application in use with neuroprothesis.
b. MEMS Penetrating Electrodes
The advancement of MEMS has allowed for microfabrication precision and repeatability
to transfer to neural electrodes construction. By incorporating microfabrication technology such
as ebeam deposition, sputtering, metal etching and lift off, MEMS based penetrating electrodes
now have the capabilities to precisely define electrode geometry as well as a whole larger choice
of materials. Bulk and surface micromaching capabilities allow for precision probe shape design
and thin film technology allow a larger choice of dielectrics and insulation. MEMS based neural
probes were pioneered by University of Utah and University of Michigan. Integrated Circuit
technology was first implemented at Stanford by Dr. Wise[105], but Wise moved to University
of Michigan and developed with Najafi a planer probe, known famously as the Michigan
probe[106], [107]. Along the same time as the Michigan probe was developed, Dr. Norman at
the University of Utah produced a three dimensional probe, known as the Utah Electrode
Array(UEA)[108], [109].
The Utah Electrode
Dr. Richard Norman’s group at the University of Utah fabricated an electrode array
referred to as the Utah Electrode Array or UEA[108–116]. The group fabricated a ten by ten
electrode array with overall dimensions of 4.2 mm by 4.2 mm Each electrode is needle shaped
and can range from 1 to 1.5 mm in length long, with a 80µm diameter at the base. The needles
project to a .2 mm thick glass-silicon composite base, which also serves as insulation, see figure
6d[9]. Glass insulation is poured between grooves that are made with a dicing saw in the
backside of a silicon square block. The glass/silicon base is grinded down, and patterned with
aluminum. The device is flipped over and columns are diced on the top side, and finally the
columns are sharpened and insulated. Insulation on the tips is removed by poking the arrays
through aluminum foil and etching the exposed tips. This method has been improved recently
with parylene C deposition, and a custom holder for photoresist spinning[113]. Other groups
have improved charge capabilities of the UEA by coating with new materials[117] or roughening
the tips[118].
Figure 2. 2: Utah Electrode Array- Utah Electrode Device including circuit received in May 2008(Top),
Varying Length Utah Electrode(Bottom Left) [119], 10 by 10 Utah Electrode(Bottom Right) [111]
An improved design with varying length electrodes has also been design to allow stimulation of
different regions of a nerve bundle, see Figure 2.2 [110]The authors have also designed a
pneumatically actuated insertion method to minimize tissue damage their system was able to an
array of 100 needles electrodes into the cerebral cortex with minimal by using a minimum
insertion speed for safe entry was 8.3m/s for a depth of 1.5 mm into a cat cortex[109].
There are certain draw backs with the UEA, first of all the fabrication is tedious due to the
numerous manual dicing steps required to create one array. Preferably, a more lithography based
approach would provide The UEA lacks on chip circuitry thus requiring lead wires that bulk up
the size of the device. As can be seen in Figure 3, the electrode arrays are tiny in comparison to
the whole device. The circuit board, itself is over 8cm long; the size of this device makes long
term implantation impractical. In addition, its structural integrity is in question because it is a
composite material and thus introducing an interfacial region. One additional issue is that glass
has a lower service temperature than the silicon, which might cause processing obstructions.
The Michigan Probe
At Michigan Center for Neural Communication Technology planer iridium stimulating
arrays were fabricated using surface micromachining and thin-film technology[119–122]. The
probes are fabricated by selectively diffusing boron to define the shape of the probes. The
thickness of the devices is defined by the thickness of the boron layer. A three layer stack of
silicon dioxide, silicon nitride and another layer of silicon dioxide, serve as dielectrics and
insulation of the interconnects. Phosophorous doped polysilicon is patterned and serves as
interconnects, which is then covered with the three layer stack of dielectrics once more. Next
bond pads are made with gold and chromium for adhesion, while electrode sites are made with
iridium with titanium for adhesion. The device is finally released by using anisotropic silicon
etch composed of ethylenediamine, pyrocatechol, and water.(EDP) Integrated ribbon cables as
thin as 2 µm and as thick as 20, as well as a sharper tip angles of less than 10 degrees can also be
fabricated using shallow boron diffusion[122]. Later improvement of the original cables was to
use polyimide which provides more flexibility and improves durability of the failure prone
silicon cables[123]. These probes were able to be integrated with CMOS[119] on chip circuitry.
Although planer they can be assembled into three dimensional arrays using orthogonal lead
transfers between the planer probes and a platform; Gold electrode plated leads are bent and then
the probes are manually inserted into a platform[120]. Despite the great improvements and
advancements on the Michigan probe, there are still problems limiting their use in neural
implants. Although, the planer geometry is defined by microfabrication capabilities, the depth is
dependent on boron diffusion, and therefore limited. In addition, the mechanical mismatch
between nerve implantation and these electrodes causes problems of chronic implantation. For
long term implantation, a wireless system must be integrated to remove the output leads. There is
also the consideration of the tissue-electrode interface. Improvements of the electrodes have been
made with bioactive coatings. These techniques have lowered the impedance caused by
unfavorable cell activity by coating the electrodes with the conducting polymer polypyrrole
(PPy) doped with biomolecules having cell adhesion[124], [125]. Although these systems have
improved the interface between electrodes and tissue, there is still the issue of chronic immune
response. Preferable drug eluting system must be incorporated to prevent this immune response.
Figure 2.3:Michigan probe arrays-Single probes are planer(Top Left) [127], 3 dimensional integration of
planer electrodes(Top Right)[128], Cmos Integration(Bottom Left)[120], polyimide cable used instead of
silicon ribbon(Bottom Right)[124]
Silicon Penetrating Electrodes
The UEA and Michigan probe pioneered the field of MEMS based microelectrodes,
however the advancements of microfabrication processes and technology brought a whole new
era of MEMS based probes. In addition, to the UEA and Michigan probe, researchers have
developed various other probes using different materials. In our laboratory, Dr. Jian Wu
developed a three dimensional silicon intraneural electrode arrays that utilizes DRIE instead of
the dicing saw used in the UEA fabrication[126]. Further improvements by Dr. Wu included the
incorporation of a form fitting SU-8 layer to fit the electrode array with peripheral nerves[127].
In addition, Dr. Wu also fabricated a novel curvature-controlled electrode for use in cochlear
implants; the probe was made out of polypyrrole which deforms by application of electric
field[128]. Researchers in California Institute of Technology utilized emerging parylene
technology to develop a more durable alternative to the silicon ribbon and polyimide cables of
the Michigan probe[129]. To improve on the uniformity and circuit integration, groups have
moved toward dry etching of probe shape. The first probes developed with dry etching were
limited to how deep they could etch. Therefore they used a combination of Reactive Ion Etching
to define the probe shape, and used backside wet etching to release the probes[130][131]. With
the development of deep reactive ion etching (DRIE), thicker probes could be etched. To
provide a thicker support structure groups have utilized deep reactive ion etching(DRIE) of SOI
wafers[132–134]. DRIE is a dry etching technique and has more control than the wet etching
used in the Michigan probe. In addition by using a SOI wafer, the probe thickness is defined by
the topside silicon while the support structures depth is limited to the thickness of the SOI wafer.
DRIE has also been used by our group to fabricate MEMS based alternative to UEA[126], [135]
which requires less manual fabrication with features defined by lithography instead of a dicing
saw. Chapter 2 and 3 of this will further optimize our previously fabricated three dimensional
electrodes.
Figure 2. 4:Silicon probe with parylene wire(Left)[129], Probe made by combination of dry and wet
etching (Right) [131].
Polymer Based Probes
Silicon based probes provide excellent recording and stimulating capabilities in acute but
have not had as parallel success in chronic use. A key issue which degrades electrode charge is
the cellular response from the mechanical mismatch between tissue and electrode. Fabrication of
flexible neural probes has been a key research issue and has resulted in shift towards polymer
based probes. Polymer based probes all have a similar fabrication procedure that involves
depositing a sacrificial layer on silicon, then depositing the first polymer, next patterning metal
on the first polymer layer, then depositing another polymer layer, exposes the electrodes and
finally dissolving the sacrificial layer to release the polymer probe[65], [136–145]. Although,
there are a plethora of choices of polymers, polyimide[65], [136], [146], [147],
Benzocyclubuten[148–150] and Parylene[143–145], [151], [152] have been used most
extensively in neural probes. Along with a more cost effective and easier fabrication procedure,
polymer probes can be assembled easier than their silicon counter parts. Takeuchi et al
demonstrated three dimensional batch assembly of their polymer probes by using magnetic
force[153]. The main disadvantage of flexible probes is that they are not strong enough to
penetrate the nerve. But this can be overcome by using a microneedle to first puncture the
epineurm and perineum[140] integrate a silicon backbone with the polymer[137] or use a
biodegradable polymer to increase mechanical stiff for insertion and later dissolve to increase
flexibility[147].
Cellular Response Probes
Although flexible probes provide on improvement on long time viability over their
silicon counterparts, there still are more promising methods to further improve the lifetime of a
chronic implant. The main culprit of electrode failure is due to the host’s systems cellular
response. However, the chronic electrode has already by developed over 20 years ago. The
neurotrophic electrode consists of gold wires and neurotrophic factors inside of a glass tip[154–
156]. By implanting autologous growth factors into the tip, the electrode integrates with the host
tissue by inducing neurite growth, and have achieved viability of 4 years in humans[155].
Although these electrodes have achieved great success in brain computer interface
applications[157–160], using them in peripheral nerve neural interfaces will be less successful
due to the difficulty in miniaturizing the electrodes and defining active sites. Researchers at the
University of Michigan developed a probe to reduce tissue encapsulation around the electrode
sites by placing them in a finer lateral area region[161–163]. They were able to increase
penetrating strength while placing the electrodes on a smaller area by incorporating an open
architecture design. Most recently, the same group utilized parylene-N covered by PEGMA to
prevent protein adoption and cellular response[164]. Unlike the neurotrophic electrode, the openarchitecture probes did not contain neurotrophic factors to induce neurite growth, but current
research has shown that neurite growth can be inducing by electrical stimulation or neural
adhesion molecular coating [165]. Groups have developed electrically stimulated nerve guidance
channels to guide nerve growth as an alternative to nerve grafts[166–168]. Although, the
guidance channels were not developed for prosthetics, the idea of regenerative electrodes for
neuroprostheses were one of the first conceived[169–172].
Figure 2.5: Electrodes design on small surface area to reduce encapsulation[174]
Regenerative Electrodes
Regenerative Electrodes are another group of penetrating electrodes. They are also known as
sieve electrodes and implanted on severed nerves. The electrical activity and/or neurotrophic
factors induce never axon growth through the holes in the electrode[40], [171], [173–177].
Although, these electrodes provide the highest selectivity they are limited to use only severed
nerves or stumps of amputees.
Figure 2.6:Sieve Electrode(Left)[175], Implantation Setup(Right)[177]
Fluidic Delivery Probe
Figure 2. 7: Various Fluidic Probe Fabrication Techniques-Bonding(Left), Sacrificial layer(Middle),
sealing buried channel with dielectrics(Right).
The ideal neural probe would have the highest selectivity while evoking the minimal
immune response. An electrode with fluidic delivery capabilities could combine the selectivity of
penetrating electrodes, the axon regenerative abilities of sieve electrodes and the tissue
encapsulation reduction capabilities of neurotopic electrodes. A MEMS based approach would
allow for high repeatability, scalability, and precise control over design, while a fluidic channel
could delivery either drugs and growth factors to prevent or enhance cellular responses. Figure
2.7 provides an overview of the various approaches to fabricate fluidic channels. University of
Michigan researchers have pioneered neural electrode progress, Kensall Wise and David
Anderson developed the commercially viable Michigan probe, and their student’s Daryl Kipke
and Jamie Hetke continued the legacy and success with NeuroNexus technologies. That legacy
still continues today with David Martin leading the field in electrode materials with his startup
Biotectix. It would seem fitting that one of the first neural electrodes simultaneous fluidic and
electrical capabilities would originate at the University of Michigan[178].Chen et al, fabricated
in fluidic channel in the traditional Michigan probes by using shallow boron diffusion to define
the channel mask, RIE to etch the channel, EDP to form a continues flow channel by
undercutting the mask, and growing and depositing dielectrics to seal the channel[178].
Although, the probes were capable of delivering fluids, they are more practical in research on the
cellular response of chemicals due to the extremely shallow depth of the channels. The channels
were limited to 15 um due to the channel fabrication and sealing process. In addition, the final
release of the probes use EDP as in the Michigan probe fabrication which could potentially
destroy CMOS circuitry. To improve CMOS compatibility, UC Berkeley researchers used Deep
Reactive Ion Etching of a SOI wafers to both define and release the probes. Their fluidic channel
was defined in similar fashion but with a nitride mask, and KOH was to undercut the mask, and
low pressure chemical vapor deposition of oxide was used to seal the channels[132].
Figure 2. 8: Silicon Fluidic Probes- Figure Michigan fluidic probe(Left)[180], Berkeley probe
(Right)[132].
To improve flexibility, Michigan researchers recently fabricated a parylene C flexible fluidic
probe by filling an etched a channel in parylene with sacrificial photoresist and then applying
another layer on top to seal the channel[152], [179]. Limitations from this method arise from
non-planar deposition of sacrificial photoresist and parylene thickness deposition. The current
fluidic probes available at NeuroNexus are made from their Michigan probe mounted to a fluidic
tube.
Figure 2. 9: Parylene fluidic probe, channel height 5 micrometers[153]
Other groups have similarly used sacrificial layers and sealants to construct fluidic neural
probes[180–182]. Similarly to the fluidic Michigan probes, sacrificial methods to construct
fluidic probes using polyimide or parylene limit the channel depth. Most recently, Kuo et al,
inserted a wire into a parylene fluidic probe fabricated from sacrificial method, and
thermoformed the electrode into a three dimensional shape with fluidic openings up to 300
um[183]. Boer et al developed the buried channel technology to increase the dimensions of
silicon etched microfluid channels. The process involves DRIE of trench, sidewall coverage of
trench with either silicon dioxide or LPCVD silicon nitride, removing the coverage on the
bottom of trench with RIE, creating buried channel with Isotroptic dry etching, KOH, or HNA,
and sealing the channel seal trench with LPCVD of polysilicon dioxide or, silicon nitride[184].
Figure 2. 10:Silicon fluidic probe-DRIE used to make trench (Top), isotropic etching to make channel,
and sealed with dielectrics(Bottom).[186]
An alternative to sealing the channel with dielectrics or using sacrificial materials is
bonding the channel roof to walls of microchannel. Metz et al defined channel walls with
polyimide and the channel roof was constructed by bonded and releasing an another polymide
layer that was mounted on mylar foil[138].Parker et al used thermocompression bonding Ti foils
to fabricate embedded channels[185]. The European Union funded NeuroProbes project utilized
wafer bonding to fabricate silicon probes with fluidic channels. Wafer bonding allows for
increased channel depth that is now limited to silicon thickness rather than polymer deposition or
undercutting mechanics. The NeuroProbes project used DRIE to define channels and the probe
shape in one wafer, then mounted wafer to serve as the roof of the channel, the mounted wafer
was grinded down to the desired thickness, and DRIE was used to etch out the probe shape on
the mounted wafer[186]. The probes then could be assembled into three dimensional
configuration with a PDMS based microfluidic connection[187]. Polyimide and parylene probes
have the advantage in flexibility, tissue stiffness matching, simplicity of fabrication, while
silicon based probes allow for greater control of channel depth and stiff enough for neural
penetration.
SU8 is a low cost negative resist developed by IBM with single coat spin thickness of 750nm
to 450 um[188] and has excellent biocompatibility[189]. Interestingly one of the first uses of
SU8 in neural application in 2004[190] and 2005[191] were used as a molds. Park et al
fabricated constructed double molds for biodegradable polymer needles by sharpening SU-8
cylindrical posts with RIE[191]. The SU8 molds were used to create PDMS molds which were
further used for biodegradable polymers. Kim et al created hollow metallic needles using
sacrificial SU8. Tapered SU8 posts were created by backside exposure through glass, which
subsequently was covered with metal, then polished to create a hole in posts, and plasma was
used to remove the SU8[190]. Lu et al reported one of the first SU8 neural electrodes were in
2006[192] [193], where they fabricated planar SU8 electrodes with grooves for nerves to be
inserted. In 2011, Ribeiro et al developed a low cost, simple solution to silicon penetrating
electrodes using SU8;They deposited gold in-between two layers of SU8 with omnicoat used to
separate the probes from the silicon substrate[194]. In 2012, a Spanish research group under the
ERANET-Neuron project fabricated planar SU8 probes under a similar method to Ribeiro et al,
with the main difference being the deposition of sacrificial aluminum instead of omnicoat for
device separation[195]. In 2009, the Spanish group first demonstrated the viability of SU-8
based neural electrodes with fluidic channels[196], and most recently the same group reported in
a different approach to create SU8 embedded microchannels [197]. Their newest involves
bonding two patterned SU-8 layers, but crosslinking will increase thermal stability, and glass
transition temperature and make reflow and bonding difficult. An alternative to bonding would
allow for a simpler process, more accurate alignment, and stronger interface sealing the channels.
Summary
We presented a review of electrodes for use in neuroprosthetic, brain machine and functional
electrical stimulation devices. Intraneural electrodes are necessary to achieve the high selectivity
needed to reach central axons of nerve fibers for complex motor control. The tradeoff with
increasing selectivity is the increase in invasiveness. Current research understands the need for a
chronic electrode to make neuroprosthetic systems a reality. Table 2.1 summarizes the progress
made in multi-electrode arrays. Initially, silicon was the gold standard used for neural electrodes,
but as researchers attributed the cellular response that degrades the electrode to the mechanical
mismatch between silicon and neuronal a trend was pushed to more flexible materials. At the
same time, the demand for a probe capable of simultaneous fluidic delivery, and electrical
abilities as created because the neurotropic electrode, and others studies demonstrated that
neurotrophic factors and drugs can extend the lifespan of probes[154][198], [199][200] while
enhance nerve regeneration[201], [202]. Research and progress suggest that the next generation
of probes could be made from SU8 as it is biocompatible with customizable flexibility, IC
compatibly, and photolithography defined.
Table 2.1: Progress in multielectrode array fabrication
University/Company
Year
Material
Description
Recording and stimulating
Recording and stimulating electrode made with EDP
and boron diffusion
Wire enclosed with neurtrophic factors enclosed in glass
cone
Stanford University
1970
Si
University of Michigan
1985
Si
Georgia Institute of Technology
1989
Si/Glass
University of Utah
1991
Si
University of Michigan
1994
Si
University of Michigan
1997
Si/dielectrics
University of Southampton
2000
SOI
Fraunhofer-Institute
2000
Polyimide
University of Michigan
2003
Si/Polyimide
University of Tokyo
2003
Arizona State University
2004
Parylene
Benzocyclobute
ne
Arizona State University
2004
Polyimide/SI
University of Tokyo
Georgia Institute of Technology
Emory Universtiy
2004
Parylene
2005
PGA
University of Michigan
California Institute of
Technology
2005
Polyimide
2006
Si/Parylene
Universtiy of Texas at Dalas
2006
SU8
Recording and stimulating
Silicon Ribbon cables made with shallow diffusion of
Boron and EDP etching on non doped Si
Fluidic channel made by undercuting mask with wet
etchant and seal channel by growing dielectrics
Recording and stimulating
Recording and stimulating electrodes sandwiched
between polyimide
Polyimide cables made with same method as Stieglitz et
al
Recording and stimulating
Recording and stimulating
Polyimide with silicon backbone tip for stiffness
improvement
Fluidic probe made with sacrificial layer photoresist
SU8 needles used to mold PDMS. PMDS mold used for
PGA needle
Fluidic probe made by same method as Takeuci et al
Parylene Cable made with same methods as Takeuchi et
al
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K. Lee et
al[136]
Takeuchi et
al[144]
Park et
al[191]
Pellinen et
al[152]
Pang et
al[129]
Recording and stimulating
Lu et al[192]
SU8 shank and parylene electrode sites for reduced
Seymour et
tissue encapsulation
al[163]
University of Michigan
2006 SU8/ Parylene
University of California at Santa
Fluidic channel made by thermocompression of two Ti
Parker et
2007 Titantium
Barbara
foil layers
al[185]
SU8 fluidic probe made by embedded chrome to prevent Fernández et
crosslinking
al[196]
Spain*
2009 SU8
University of Southern
Fluidic channel made by thermoforming parylene with
California
gold wire
2012 Parylene
Kuo et al[183]
Parylene, SU8 and polyimide coated with PEG for
Cheng et
stiffness improvement
al[147]
Polymer coated with PEG
2013 PEG
Fluidic channel made by bonding of two patterned SU8
Altuna et
Spain**
2013 SU8
layers
al[197]
*Collaboration between IKERLAN, S. Coop.(Company), University of Barcelona, Ciberned, and Spanish National Research
Council
** Collaboration between IKERLAN, S. Coop.(Company) , Ciberned, Spanish National Research Council and Cajal institute
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