Available online at www.sciencedirect.com

Sensors and Actuators B 132 (2008) 449–460

Flexible parylene-based multielectrode array technology
for high-density neural stimulation and recording
Damien C. Rodger a,b,∗ , Andy J. Fong a , Wen Li a , Hossein Ameri b , Ashish K. Ahuja b ,Christian Gutierrez b , Igor Lavrov c , Hui Zhong c , Parvathy R. Menon a , Ellis Meng b ,Joel W. Burdick a , Roland R. Roy c , V. Reggie Edgerton c , James D. Weiland b ,Mark S. Humayun b , Yu-Chong Tai aa California Institute of Technology, Pasadena, CA, USAb University of Southern California, Los Angeles, CA, USAc University of California, Los Angeles, CA, USA

Available online 12 November 2007

Abstract
Novel flexible parylene-based high-density electrode arrays have been developed for functional electrical stimulation in retinal and spinal cord
prosthetics. These arrays are microfabricated according to a single-metal-layer process and a revolutionary dual-metal-layer process that promises
to meet the needs of extremely high-density stimulation applications. While in many cases thin-film platinum electrodes in parylene C would be
sufficient, high surface-area platinum electroplating has been shown to extend the lifetime of stimulated electrodes to more than 430 million pulses
without failing. Iridium electrode arrays with higher charge delivery capacity have also been fabricated using a new high-temperature stabilized
parylene variant, parylene HT. In addition, a new heat molding process has been implemented to conform electrode arrays to approximate thecurvature of canine retinas, and a chronic implantation study of the mechanical effects of parylene-based electrode arrays on the retina overa 6-month follow-up period has provided excellent results. Retinal stimulation from these parylene-based electrode arrays in an isolated tiger
salamander preparation was shown to be comparable to light stimulation in terms of generation of action potentials in the inner retina. Finally,
electrode arrays have also been implanted and tested on the spinal cords of murine models, with the ultimate goal of facilitation of locomotion
after spinal cord injury; these arrays provide a higher density and better spatial control of stimulation and recording than is typically possible using
traditional fine-wire electrodes. Spinal cord stimulation typically elicited three muscle responses, an early (direct), a middle (monosynaptic), and alate (polysynaptic) response, classified based on latency after stimulation. Stimulation at different rostrocaudal levels of the cord yielded markedly
different muscle responses, highlighting the need for such high-density arrays.
© 2007 Elsevier B.V. All rights reserved.
Keywords: Artificial vision; BioMEMS; Biomimetic; MEA; Neural prosthesis; Parylene C; Parylene HT; Retinal prosthesis

1. Introduction
Low-electrode-density neural prostheses have shown incredible promise, enabling those with severe hearing impairments
to recognize speech [1] and those blind from such devastating outer retinal diseases as retinitis pigmentosa (RP) (where
the photoreceptors are damaged but the remaining inner retinal circuitry remains largely intact [2]) to perceive visual data[3]. Subjects with prototype sixteen-electrode retinal prostheses∗ Corresponding author at: California Institute of Technology, Department of
Bioengineering, 1200 E. California Building, M/C 136-93, Pasadena, CA 91125,
USA. Tel.: +1 626 395 8477; fax: +1 626 584 9104.
E-mail address: dcrodger@mems.caltech.edu (D.C. Rodger).

0925-4005/$ – see front matter © 2007 Elsevier B.V. All rights reserved.
doi:10.1016/j.snb.2007.10.069

can even distinguish between objects such as plates, cups, andknives, and perceive directions of movement in high-contrast
environments far better than would be possible by chance alone[4]. However, there exists a need for a multielectrode array(MEA) technology that is capable of increasing the densityof neural stimulation beyond its current limits, while ensuringbiocompatibility and device longevity. In fact, for retinal stimulation in particular, it has been shown that room navigation
could be significantly improved with an electrode array comprising one thousand or more electrodes, and furthermore such anarray would likely enable facial recognition and large type reading [5]. The next-generation retinal prosthesis for patients withdiseases like RP and age-related macular degeneration (AMD)(Fig. 1), then, requires a high-density flexible electrode array

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A new high-temperature stable [10] and ISO 10993 biocompatible [11–13] fluorinated variant of parylene, parylene HT, thatis similar in many respects to parylene C, has also been used tofabricate iridium electrode arrays. We show that while evaporation and patterning of iridium was unsuccessful on parylene
C due to the high melting temperature of iridium, parylene HTlends itself to such a process, and we present this as anotherpossible technology for ensuring good charge delivery to neuraltissue.
2. Fabrication methods
2.1. Single-layer process
Fig. 1. Placement and components of envisioned next-generation intraocularretinal prosthesis.

capable of stimulating the inner retina and a high-lead-countcable to allow for high-resolution vision. We present the firstflexible parylene-based MEAs designed for functional electricalstimulation in retinal prostheses, and the extension of this technology to enable stimulation of central nervous system structures
after spinal cord injury. In addition to presenting a high surfacearea electroplating technology that extends stimulation electrodelongevity to at least 430 million pulses, which is ample for many
applications, we address the importance of a novel paryleneenabled dual-metal-layer fabrication methodology that permits
complex electrode arrangements while alleviating the traditional
problems of electrode crowding and electrode size restrictions
caused by wire routing. Promising chronic biomechanical stability results in canine eyes and acute neural recording andstimulation results in an in vitro retinal preparation and in vivo inmurine spinal cords are presented, demonstrating the ability ofthese parylene-based arrays to both record from and stimulate theneuronal targets of interest. All animal procedures conformed
to the ARVO Statement on the Use of Animals in Ophthalmic
and Vision Research.
The advantages of using parylene C as the structural material for such neuroprostheses, when compared with technologies
based on the use of other materials such as PDMS, polyimide[6] and silicon [7], include parylene’s pinhole-free conformality due partly to its unique room-temperature chemical
vapor deposition process, its low water permeability, its chronic
implantability as an ISO 10993, United States Pharmacopeia(USP) Class VI material (highest biocompatibility class for plastics in the United States), and its high flexibility and mechanicalstrength (Young’s modulus ∼4 GPa). Since parylene is depositedat room temperature (we have verified this using Temp-Plate
irreversible temperature recorders traceable to NIST (Wahl
Instruments, Inc., Asheville, NC, USA)), the coating process ispost-integrated-circuit (IC) compatible. Parylene C is also optically transparent, enabling the anatomy to be seen through thecable and the array during ophthalmic surgery, post-implantation
examination, and follow-up. While many groups use parylene C
as a coating of their arrays for many of these reasons, we havechosen to use it as the main substrate for our devices [8,9], aparadigm that leverages these advantages to the greatest extent.

Single-metal-layer parylene C-based electrode arrays are fabricated as shown in Fig. 2. A photoresist sacrificial layer is
optionally spun on a standard silicon wafer. Approximately
8 ␮m of parylene C is then vapor-deposited in a PDS2010
system (Specialty Coating Systems, Indianapolis, IN, USA)on the entire wafer. An LOR3B photoresist layer (Microchem
Corporation, Newton, MA, USA) and an AZ1518 layer (AZElectronic Materials, Branchburg, NJ, USA) are spun on topof the parylene, exposed in a 10× reduction GCA Mann 4800
DSW wafer stepper (General Signal Corporation, Stamford, CT,
USA) or a Kasper 2001 contact aligner (Kasper Instruments,
Inc., Sunnyvale, CA, USA) depending on the required resolution of the electrode array, and developed to achieve a liftoff
pattern consisting of contacts, conductive traces, and electrodes.
After hard-bake, approximately 2000–5000 Å of platinum, withor without a 200 Å titanium layer, is then e-beam evaporated(SE600 RAP, CHA Industries, Fremont, CA, USA) on thewafer. The subsequent photoresist strip generates the desired
single-layer metallization pattern. An approximately 7-␮mthick coating of parylene C is then deposited, followed by a spin
coating of photoresist. This photoresist etch mask is exposedover the areas of the electrodes and contact pads and to pattern theoverall array geometry, and the entire wafer is then subjected to areactive-ion etch in oxygen plasma, removing the parylene insulation over the electrodes and the parylene surrounding the array.
The photoresist mask is then removed with solvent. Finally, if asacrificial photoresist layer was used, the array is released fromthe substrate in an acetone bath. If no sacrificial layer was used,
it is peeled from the silicon in a water bath. Ultimately, for mostcases, the sacrificial photoresist layer is unnecessary.
Due to the higher thermal stability of parylene HT (long-term
stability at 350 ◦ C, intermittent exposures up to 450 ◦ C [12]), it
was surmised that the material would be better suited to iridiumarray fabrication than would parylene C (the melting temperature of iridium is 2447 ◦ C whereas that of platinum, for which aparylene C substrate works well, is 1772 ◦ C [14]). Consequently,
a minor modification of this process is made to fabricate iridium
electrode arrays. A thin parylene C layer (∼2.4 ␮m) is depositedon the silicon wafer followed by a thicker layer of parylene HT(∼5.7 ␮m) in a PDS2035 system (Specialty Coating Systems).
The parylene C layer facilitates fabrication and subsequent
release, whereas the HT layer provides the necessary thermal
stability. The dual photoresist layer is spun and patterned, and theD.C. Rodger et al. / Sensors and Actuators B 132 (2008) 449–460

451

Fig. 2. Fabrication process for parylene-based single-metal-layer MEAs. A simple modification of this process is required for iridium electrode array fabrication.

iridium (∼800 Å) is then e-beam evaporated on the wafer. After
liftoff, a final parylene HT layer (∼5.4 ␮m) is deposited and patterned as with the parylene C arrays, and the arrays are released ina water bath. This process was compared to an identical process
performed using only parylene C structural layers.

2.2. Dual-layer processDual-metal-layer electrode arrays are fabricated as shown in
Fig. 3. Approximately 8 ␮m of parylene C is first deposited ona silicon wafer with the optional photoresist sacrificial layer,

Fig. 3. Fabrication process for parylene-based dual-metal-layer MEAs.

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forming the underside of the electrode array. A platinum ortitanium–platinum metal liftoff process is used to define traces
with 16 ␮m pitch and 2000–3000 Å thickness. A second parylene deposition (∼1 ␮m) forms the insulation between the two
metal layers. At this point, 6 ␮m × 6 ␮m vias are patterned in
the insulation layer over the ends of the traces using an O2
plasma reactive-ion etch. A second step-coverage optimized
liftoff process is used to define a second metal layer comprising electrodes and traces, while at the same time achieving
electrical continuity between the underlying traces and the overlying electrodes. A final parylene coating approximately 7 ␮m
thick forms the top insulation. The electrodes are exposed andthe overall geometry of the implant is defined in a final setof O2 reactive-ion etches using a thick photoresist etch mask.
Finally, the arrays are peeled from the wafer in a water bathor released through removal of the sacrificial photoresist in
acetone.
2.3. Annealing and heat molding
The implants are typically annealed for 2 days at
200 ◦ C in a vacuum oven with nitrogen backfill to optimize
parylene–parylene adhesion. Under accelerated-lifetime passive
soak test conditions, such an annealing process has been shownto increase the extrapolated mean time to failure of devices
to 20 years or more [9]. In addition, an important aspect ofthese parylene arrays is that they are heat moldable. If thearrays are contoured and confined to a geometry of interest for
the target application during this annealing process, we havediscovered that this conformation will be maintained duringsterilization, implantation, and follow-up. In order to maintainthe planarity of the spinal cord arrays during annealing, they
are clamped between two flat pieces of Teflon or glass slides
coated with aluminum foil. The retinal arrays, on the other hand,are shaped using a custom 6061 aluminum mold comprising arecessed concave region and a mating stainless steel sphere thatapproximates the curvature of the retina. During annealing, thearray region is sandwiched between the sphere and the mating surface, while the cable is pressed flat against an aluminum
plateau.
2.4. Electroplating
As a possible mechanism for extending the longevity ofchronically pulsed electrodes, we are investigating electroplated films of high surface-area platinum. Specially designed
thin-film platinum electrode arrays, consisting of sixteen 150␮m-diameter electrodes of 3000 ␮m center-to-center spacing,were fabricated according to the single-layer process. They were
immersed in an aqueous ammonium hexachloroplatinate solution in a specialized jig and six were plated according to Whalen
et al. [15] at a plating potential of −0.6 V (vs. an Ag/AgCl
reference electrode) for 1.5 h. The others remained unplated.
Electrochemical tests were performed to evaluate the efficacy ofthis plating step in extending electrode longevity under chronic
pulsing.

Fig. 4. Isolated larval tiger salamander retina (darker region) overlying parylenebased platinum electrode array. Arrow indicates 10-␮m-diameter electrode used
for recording traces in Figs. 5 and 6. Asterisk identifies 200-␮m-diameter stimulating electrode used to generate action potentials seen in Fig. 6.

3. Results and discussion
3.1. In vitro retinal recording and stimulation
Parylene C-based arrays of thin-film platinum electrodes,
comprising four 200-␮m-diameter stimulating electrodes and 56
recording electrodes of 10 ␮m diameter were fabricated according to the single-metal-layer process on a glass substrate. These
were placed in a bicarbonate perfusate under a microscope andconnected to a stimulus generator and preamplification board(Multi Channel Systems MCS GmbH, Reutlingen, Germany)[16]. As shown in Fig. 4, a retina isolated from larval tiger
salamander (Ambystoma tigrinum) was placed retinal ganglion
cell side down on the array (to simulate epiretinal stimulation),and a remote platinum ground electrode was introduced to thebath. In order to assess retinal health, and for comparison withelectrical stimulation results, a white light pulse 40 ms in duration was applied to the tissue. As shown in Fig. 5, a voltage
trace of the activity of the cells overlying one of the recording electrodes (electrode indicated with an arrow in Fig. 4),a robust ON response was detected after the phototransduction delay, followed by the expected OFF response ∼50 ms
later. Subsequently, with the lights off, a 20 ␮A, 400 ␮s/phase,
cathodic-first biphasic electrical pulse was applied between thestimulating electrode indicated with an asterisk in Fig. 4 andthe ground electrode. The voltage trace from the same recording electrode as before is shown in Fig. 6. This stimulation wasconsistently repeatable over a 50 pulse train with a 400 ms interpulse interval, and other stimulating electrodes were also capableof “epiretinally” stimulating other cells in the retinal slice. As isclear from these results, the parylene-based platinum electrode
was able to stimulate the tissue and elicit a response similar to
the response generated from a light pulse in this intact retina.
Given these results and the knowledge garnered from clinical

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453

Fig. 5. Recording of ON response followed by OFF response of cells overlying
electrode denoted with an arrow in Fig. 4 to a full-field white light stimulusof 40 ms duration. Pulse began at 0 ms on the abscissa, with delay until ON
response due to phototransduction.

Fig. 6. Typical recording of response of cells overlying same recording electrodeas in Fig. 5 to a 20 ␮A, 400 ␮s/phase, cathodic-first biphasic electrical pulse
from “epiretinal” stimulating parylene-based platinum electrode denoted withan asterisk in Fig. 4. Spike at 0 ms on abscissa is stimulation artifact.

trials with prototype arrays fabricated of other materials, it is
not unreasonable to presume that our arrays will most likely beable to stimulate retinal tissue in other species, including human.
To this end, the chronic biostability of implanted parylene-basedarrays was also evaluated.

3.2. Chronic retinal implantation
Chronically implantable retinal electrode arrays comprising 1024, 75-␮m-diameter electrodes arranged in a complex
biomimetic pattern that closely mimics the density of ganglion

Fig. 7. (a) Fabricated biomimetic electrode array with 60 of 1024 75-␮m-diameter electrodes connected through dual-layer process with U.S. dime for size comparison;(b) SEM of electrode highlighting vias to underlying traces; (c) heat molded and annealed retinal electrode array with retained spherical curvature (arrow denotes
retinal tack hole).

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Fig. 8. Fundus photographs (left) showing parylene MEAs tacked to the right retina of both animals and FAs (right) showing normal vessel perfusion under thearrays. Arrows point to retinal tack.

cells in the human retina [17] were fabricated according to thedual-layer process (Fig. 7a), with 60 of the electrodes connected
via two traces each to facilitate electrical conductivity verification. The strength of metal adhesion was verified using a Scotch
tape test, which demonstrated that direct platinum evaporation is
feasible without the need for a titanium adhesion layer. Electrical testing demonstrated a typical impedance of approximately
5 k, which includes two 8-␮m-wide traces of 20 mm length,
as well as two via step junctions connecting underlying traces
to the overlying electrode (vias are shown in Fig. 7b, a scanningelectron micrograph (SEM) of a single electrode). Each via had
an impedance of less than 12.5 . These arrays were successfully molded to the approximate curvature of the canine retina(Fig. 7c) using the custom mold, and sterilized using ethyleneoxide gas.
Our biomimetic arrays were implanted in the right eye of
two canines through a 5-mm pars plana incision after vitrectomy, and were affixed to the retina using a retinal tack modified
by the addition of a PDMS washer (to account for the thinnature of the parylene-based arrays). Follow-up in both animals
was conducted for 6 months using fundus photography, fluorescein angiography (FA), in which blood is fluorescently stained
to assess vessel perfusion in the retina, and optical coherence

tomography (OCT), an interferometric technique that enables
cross-sectional imaging of the retina. Fundus photography andFAs of both animals, examples of which are shown in Fig. 8, consistently demonstrated that vessel filling underneath the array
was normal. Obstruction and vessel leakage would have beenvisualized if the array were placing excessive pressure on theretina. In addition, OCT demonstrated that the electrodes wereconsistently less than 50 ␮m away from the ganglion cell layer
in both animals (typical OCTs of both animals are shown in
Fig. 9), an outcome that theoretically affords excellent electricalcoupling between the electrodes and the electrically excitable
cells of the retina. It is important to note that in the OCT of thesecond canine, the scan was taken along a segment furthest fromthe tack site, where one might expect the least proximity. Evenat this location, this array remained in very close apposition
throughout the 6-month implantation. Post-enucleation histology has since confirmed the excellent biostability seen duringfollow-up.
The dual-metal-layer process is wholly enabled by the useof parylene as an insulating layer. The low dielectric constantof parylene (∼3.1 at 1 kHz [13]) enables this layer to be verythin while still minimizing capacitive crosstalk between overlapping and adjacent metal lines. The vapor-deposition process

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Fig. 10. Parylene MEA for murine spinal cord stimulation and recording.

stimulation with implanted parylene-based arrays in an animalmodel.
3.3. In vivo spinal cord recording and stimulation

Fig. 9. OCTs of both animals showing very close apposition (<50 ␮m) of thearrays to the retinal ganglion cell layer.

for parylene ensures that this layer can remain thin while still
being pinhole free, where a spin-on coating, on the other hand,
would prove problematic. The process depends on optimal stepcoverage of the parylene sidewall during evaporation, which is
aided, in part, by the slightly isotropic nature of the O2 plasmaetch of parylene [18]. The dual-layer process proffers considerable advantages over the more traditional single-layer approach.
Design of single-layer electrode arrays is usually hindered by
the need to route traces amongst the electrodes. This tends tocause crowding of traces and electrodes into groups, an organization that may not be optimal for stimulating the tissue ofinterest. In addition, this has a propensity to constrain the geometric area of the electrodes in the MEAs to smaller sizes, andthus reduces the number of electrodes possible in a given area.
The dual-layer process obviates these problems by enabling
traces to pass under overlying electrodes without making contact to them, having the effect of both relaxing the constraints
on electrode size and number and enabling more complex electrode organization (such as the biomimetic one presented in thiswork). Although the arrays fabricated here had just 60 electrodes of connectivity with 120 traces total, this was without
making full use of both layers for wire routing and connectionof electrodes. In order to not make traces unnecessarily narrow
and of too high impedance, we believe an extension of this process to three or more metal layers will be necessary to achieve1024 electrodes of total connectivity. Indeed, this fabrication
process is easily extendable to such structures through the addition of extra layers of parylene and metal. Given the encouraging
biostability results presented here and the ability of these arrays
to stimulate retinal tissue, future studies will include chronic

Spinal cord arrays, consisting of five or ten electrodes of
250 ␮m diameter (Fig. 10), were fabricated using the singlelayer process. Interelectrode spacing was controlled so thateach array covered four to five segments of the lumbosacral
spinal cord. These were connected via Clincher connectors (FCI,
Versailles Cedex, France) to the stimulation and recording electronics.
Under isoflurane anesthesia, the spinal cord electrode arrays
were implanted epidurally on spinal cord segments L2-S1 in
nontransected mice. The electrodes were oriented linearly along
the rostrocaudal extent of the cord. Recording capability wasassessed by using the electrode array to record spinal cord potentials evoked by tibial nerve stimulation. Following stimulation of
the tibial nerve, somatosensory evoked potentials were recorded
from the cord dorsum at three lumbosacral levels (P1–P3, rostral
to caudal). The recorded waveform consisted of three response
peaks, two of which are clearly depicted in Fig. 11 (N1 and N3).
These findings closely mirror results reported previously in a

Fig. 11. Peak amplitudes of somatosensory evoked potentials (N1 and N3)
recorded from three levels of the rostrocaudal spinal cord (P1–P3). Example
waveform at top shows approximate response times.

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the somatotopic distribution of locomotor circuits in the spinal
cord.
3.4. Thin-ﬁlm vs. electroplated platinum
The ability of the electroplating process to extend electrodelongevity was evaluated by chronically pulsing single-layer
thin-film and platinum-plated electrodes in PBS. Cyclic voltammograms (CVs) were performed in 250 mM aqueous sulphuricacid prior to plating and periodically during pulsing to estimate
electrode surface area and overall electrode health. In addition, voltage responses to the current pulses were recorded andelectrochemical impedance spectroscopy (EIS) was performed
regularly. The plated and unplated electrodes were pulsed continuously for 50 days, or until failure occurred, at 100 Hz with a
Fig. 12. Effect of interelectrode spacing on muscle recruitment. Smaller spacings yielded graded muscle activation while larger spacings yielded maximal
amplitude responses at low currents.

study using conventional spinal cord recording electrodes [19]demonstrating that the recording capability of the array electrodes matches that of conventional electrodes. By measuring
the difference in the response latencies obtained at each electrode position (corresponding to different levels of the spinalcord), and by utilizing the known, fixed interelectrode spacing, accurate measurements of the conduction velocities wereobtained. The properties of these responses can potentially be
used to diagnose the progressive recovery of the spinal cord asa result of treatments provided after a spinal cord injury.
To test the capability of the electrode array to act as amulti-channel stimulating device for generating hindlimb movements, constant-current monophasic stimulus pulses (amplitude:
50–850 ␮A, frequency: 0.3–10 Hz, pulse duration: 0.5 ms) wereapplied to the spinal cord between each of the array electrodes and a ground electrode located near the shoulder, while
muscle activity was monitored using electromyogram (EMG)
recordings of the tibialis anterior and medial gastrocnemius
muscles. Stimulation generated a typical three-component EMG
action potential consisting of an early (direct motor), a middle(monosynaptic), and a late (polysynaptic) response, classified
by post-stimulus latency. The appearance and magnitude of eachof these responses was correlated with the choice of electrodeposition and stimulation parameters. Varying the interelectrodespacing between the stimulating electrode and a ground electrode now on the array affected the sensitivity with which thetarget muscles were activated. The recruitment profile of themedial gastrocnemius middle response is shown in Fig. 12 using
interelectrode spacings of 1500 ␮m (filled bars) and 4500 ␮m(unfilled bars). Using the smaller spacing, graded muscle activation was achieved. With the larger spacing, approaching amonopolar configuration, the muscle quickly attained maximal
activation at low currents. Thus, the specific goal and sensitivity requirements of a particular motor task may dictate optimal
interelectrode spacing and whether a monopolar or bipolar
configuration is chosen. Currently, we are developing higherdensity MEAs capable of bilateral stimulation for assessing

Fig. 13. Voltage responses to a current pulse for (a) an unplated electrode, documenting the process of electrode failure, and (b) a plated electrode, showing
steady responses throughout the 50-day test.

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457

Fig. 14. Photomicrograph of an iridium electrode on a parylene C surface after liftoff (left) compared with an iridium electrode on a parylene HT surface after liftoff(right). Parylene HT enables fabrication of flexible iridium electrode arrays through thermal evaporation.

60 ␮A (∼0.34 mC/cm2 of geometric electrode area, close to theupper limit of charge density for platinum, typically considered
to be 0.35 mC/cm2 [20]) biphasic cathodic-first current pulse
with 1 ms per phase and a 100 ␮s inter-phase delay.
The estimated surface area of the electrodes (determined bycomparing the areas under the hydrogen desorption peaks in theCVs) increased approximately 40–50 times between before andafter plating. Under pulsing, the voltage responses of both plated
and unplated electrodes remained stable for approximately 29
days, at which point the unplated electrodes showed signs offailure. Voltage responses for one such electrode on day 26,
29, and 31 are overlaid in Fig. 13a, which documents the progression of failure. The plated electrodes, on the other hand,remained intact for much longer, most surviving more than 50days, or 430 million pulses, at which point the testing goal was
met and the test was stopped (overlaid voltage responses for one
such electrode, showing the voltage responses at day 26, 31,
and 50, are shown in Fig. 13b). In addition, the electrochemical
impedance magnitudes at 1 kHz of a typical plated and unplated
electrode averaged approximately 1.2 and 4.4 k, respectively
for the 3 weeks of testing. A dramatic jump in impedance toapproximately 0.87 M was observed for the unplated electrode at the time of failure (approximately day 29), while theplated electrode demonstrated only minor variability in its low
impedance throughout the 430-million-pulse trial. These preliminary data corroborate the evidence that plating of the electrodes
is beneficial to longevity, and suggest that high surface-area
platinization of electrodes can have a dramatic effect on extending electrode life, while lowering electrochemical impedance tocharge delivery. Future work will include replication of thesetests and chronic pulsing at high temperatures for longer times
to further accelerate and assess the possible modes of failure.

investigation into the possibility of fabricating parylene-based
flexible iridium electrode arrays spawned from such considerations. Usually, iridium or iridium oxide are deposited throughsputtering [22,24] or electroplating [25,26], especially when
they are applied to flexible substrates, a likely reason being thatthe high melting-temperature of the material makes evaporation
on polymers quite foreboding (thermal evaporation on thermallyconductive substrates such as silicon is possible, however [27]).
Our comparison of iridium array fabrication on parylene C
and fabrication on parylene HT yielded remarkable results. As
can be seen in Fig. 14, when e-beam evaporated and patterned onparylene C, the iridium film cracked and the liftoff was incomplete. Electrically, the traces from the contacts to the electrodes
were non-conductive. However, when the arrays were fabricatedusing the hybrid parylene C/HT process, the evaporation andliftoff on the HT surface occurred routinely (Fig. 14). In addition, electrical testing of contact-electrode-contact circuits using
sets of 45-␮m-wide traces ranging 56–70 mm in length revealed
impedances ranging from approximately 12.8–16.7 k, depending on the lengths of the traces in the circuit (the high resistanceis due to the thin nature of the iridium wires). An SEM

3.5. Iridium electrode arrays
While the charge delivery capacity of platinum is sufficientfor many applications, that of iridium is significantly higher [21],as it has four oxidation states as opposed to two [22]. In addition,
iridium can be cycled to form an activated iridium oxide film
(AIROF) [23] with even higher charge delivery capacity. Our

Fig. 15. SEM of a fabricated parylene HT-based iridium MEA. Compositional
analysis through EDS verified exposed material as iridium.

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of a fabricated parylene HT-based iridium electrode array isshown in Fig. 15. The composition of the electrode surface
was analyzed via energy dispersive spectroscopy (EDS) andverified to be iridium. This work constitutes, to our knowledge, the first time that iridium electrode arrays have beensuccessfully fabricated on a parylene substrate through thermal evaporation. Future work includes electrochemical analysis
and chronic pulsing of these and similarly fabricated iridiumarrays.
4. Conclusion
Single and dual-metal-layer fabrication processes forparylene-based electrode arrays have been outlined and demonstrated as robust techniques for building flexible MEAs. These
revolutionary MEAs have the ability to stimulate and record
from neural tissue in the retina and spinal cord, and demonstrate
excellent biostability when chronically implanted in contact withcanine retinas. The parylene-enabled dual-metal-layer process
allows increased electrode density while obviating many of theissues typically associated with single-layer arrays, such as constraints on electrode size and electrode crowding due to wire
routing. It is a simple matter to extend this process to increase thenumber of metal and parylene insulation layers, as will likely benecessary to fabricate a fully connected 1024-electrode deviceof approximately the same geometry. High surface-area platinum electroplating technology has shown encouraging resultsin terms of extending electrode life while also decreasing electrochemical impedance. In addition, evaporated iridium electrodearrays have been fabricated using a novel high-temperaturestable variant of parylene, parylene HT, further adding to therepertoire of parylene-based technologies for retinal, spinal, andother neural interfaces.
Acknowledgments
This work was supported in part by the Engineering ResearchCenters Program of the National Science Foundation under NSF
Award Number EEC-0310723, and by a fellowship from theWhitaker Foundation (D.R.). The authors especially wish to
thank Dr. Jack Whalen and Ms. Aditi Ray for their help withthe electroplating and electrochemical testing experiments, Dr.
Saloomeh Saati for her help with the in vivo canine studies, Mr.
Matthew Behrend for his work on the isolated retina setup, Ms.
Ronalee Lo for her help with the fabrication of the PDMS washers used with the retinal tacks, Mr. Andrew Pullin for his help
with fabrication of the heat-forming mold, and Mr. Trevor Roperfor his always valuable assistance.
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Biographies
Damien C. Rodger received his BS degree in electrical engineering (magna
cum laude with honors) from Cornell University in 2000. He is currently an
MD/PhD candidate at the Keck School of Medicine of the University of Southern
California and the California Institute of Technology conducting research inbioengineering on bioMEMS for ophthalmic and neural use. He held a Whitaker
Foundation Graduate Fellowship from 2003 to 2006 and is a member of theIEEE EMBS and the Association for Research in Vision and Ophthalmology(ARVO).
Andy J. Fong received his MS degree in biomedical engineering at the University of California, Los Angeles in 2002. Currently he is pursuing his PhD
in bioengineering at the California Institute of Technology. His dissertation
research focuses on developing new technologies for treating spinal cord injury
using a combination of robotically assisted physical rehabilitation, pharmacological treatments, and multi-locus electrical stimulation.
Wen Li received the BS degree in material science and engineering from
Tsinghua University, Beijing, China, in 2001, the MS degree in microelectronics
from Tsinghua University, Beijing, China, in 2003, and the MS degree in electrical engineering from California Institute of Technology in 2004. Currently,
she is working toward a PhD in electrical engineering at the California Institute of Technology. Her research interests include flexible MEMS devices andsystems for power and data transfer in retinal prostheses and other biomedical
applications.
Hossein Ameri graduated from Tehran University of Medical Sciences in
1993. He received his Ophthalmology and Vitreoretinal Surgery training
in Ireland and obtained his FRCSI from the Royal College of Surgeonsin Ireland, and MRCOphth from the Royal College of Ophthalmologists
in London in 2000. During his ophthalmology training at Cork University Hospital, he was also an official lecturer at University College Cork.
He moved to the Doheny Eye Institute in Los Angeles in 2004 to pursue
his research on retinal prostheses and retinal vascular disease. He is currently an assistant professor of ophthalmology at the University of Southern
California.
Ashish K. Ahuja received his BS and MS degrees from Columbia University in applied physics and electrical engineering in 1999 and 2001,
respectively. He worked in the Condensed Matter Physics Department at Bell
Laboratories for 2 years. He then received his PhD in electrical engineeringelectrophysics in 2007 from the University of Southern California. He iscurrently a Post-Doctoral Research Fellow at Second Sight Medical Products working towards the development of retinal prosthetic implants for theblind.
Christian Gutierrez received his BS degree in electrical engineering from the
California Institute of Technology in 2005. He continued at Caltech and received
his MS in electrical engineering in 2006. He is currently under fellowship pursuing a PhD in biomedical engineering at the University of Southern California.
His research interest is in the development of novel MEMS devices with applications to medicine and he is currently involved in research for an intraocularretinal prosthesis.
Igor Lavrov received his MD in 1999 from Military Medical Academy andhis PhD in 2002 from the Pavlov Institute of Physiology. In 2002 he joined
the University of Louisville School of Medicine as a postdoctoral fellow, and in2005 the laboratory of Dr. Edgerton at the University of California, Los Angeles.
His research interests have been focused on understanding spinal cord functionalorganization and mechanisms of spinal cord reorganization after injury, artificial
activation of spinal cord circuits responsible for locomotion, and developing
rehabilitation strategies for spinal cord injury.
Hui Zhong received her MD in 1983 from Hebei Medical College. She has
been at the University of California, Los Angeles since 1994. Currently, she459

is a project scientist in the Neuromuscular Plasticity Laboratory in the Brain
Research Institute. Her research interest is neuromuscular plasticity after spinal
cord injury.
Parvathy R. Menon is currently an undergraduate at the California Institute ofTechnology and will receive her BS in engineering and applied science in bioengineering in 2008. Prior to joining the Caltech Micromachining Group, she
worked as an Axline Scholar in the Microdevices Lab of the NASA Jet Propulsion Laboratory. Additionally, she has worked in the Network Architecture Lab(NAL) of Intel Corporation and at the Oregon Graduate Institute (OGI). In thefall of 2009, she will be pursuing graduate studies in bioengineering or medicine.
Ellis Meng received her BS degree in engineering and applied science from the
California Institute of Technology in 1997. She pursued her graduate studies
in electrical engineering and received her MS in 1998 and PhD in 2003 at thesame institution. She is now an assistant professor of biomedical engineering
at the University of Southern California. In the National Science Foundation
Biomimetic MicroElectronic Systems Engineering Research Center (BMES
ERC) she is a Thrust Leader for Interface Technology and the Associate Directorof Education and Student Diversity. She is a member of Tau Beta Pi, the American
Society of Mechanical Engineers (ASME), the Society of Women Engineers (SWE), and the Association for Research in Vision and Ophthalmology(ARVO).
Joel W. Burdick received his BS in mechanical engineering from Duke University and MS and PhD degrees in mechanical engineering from Stanford
University. He has been with the Department of Mechanical Engineering at the
California Institute of Technology since May 1988. His current research interests
include robotic locomotion, sensor based robot motion planning, multi-fingered
robotic grasping, medical applications of robotics, applied nonlinear control
theory, and neural prosthetics.
Roland R. Roy received the MSc and PhD degrees in exercise physiology from
Michigan State University, East Lansing, in 1973 and 1976, respectively. He
is currently a researcher in the Brain Research Institute, University of California, Los Angeles. His primary research interests concern the plasticity of
the neuromuscular apparatus when subjected to a chronic increase (exercise,functional overload, hypergravity) or decrease (spinalization, immobilization,weightlessness, hind limb unloading, spinal cord isolation) in neuromuscular
activity.
V. Reggie Edgerton received his PhD in exercise physiology from Michigan
State University. He has been at the University of California, Los Angeles, since
1968 and is currently a professor in the Departments of Physiological Scienceand Neurobiology and a member of the Brain Research Institute. Dr. Edgerton’slaboratory focuses on how, and to what extent, the nervous system controls
protein expression in skeletal muscle fibers as well as how the neural networks
in the lumbar spinal cord of mammals, including humans, control stepping andhow this stepping pattern becomes modified by chronically imposing specificmotor tasks on the limbs after complete spinal cord injury.
James D. Weiland received his BS from the University of Michigan in 1988.
After 4 years in industry with Pratt & Whitney Aircraft Engines, he returned toMichigan for graduate school, earning degrees in biomedical engineering (MS1993, PhD 1997) and electrical engineering (MS 1995). In 1999 he was appointed
an assistant professor of ophthalmology at Johns Hopkins. Dr. Weiland is now the
Director of the Intraocular Retinal Prosthesis Lab at the Doheny Retina Institute,and is an associate professor of ophthalmology and biomedical engineering at
the University of Southern California. He is a member of the IEEE EMBS, theBiomedical Engineering Society, and the Association for Research in Vision andOphthalmology.
Mark S. Humayun received his BS from Georgetown University in 1984, his
MD from Duke University in 1989, and his PhD from the University of North
Carolina, Chapel Hill in 1994. He finished his training by completing an Ophthalmology residency at Duke and a Fellowship in Vitreoretinal Diseases at
Johns Hopkins Hospital. Currently, Dr. Humayun is a professor of ophthalmology, biomedical engineering, and cell and neurobiology at the Universityof Southern California. Dr. Humayun is the Director of the National Science
Foundation Biomimetic MicroElectronic Systems Engineering Research Center (BMES ERC). He is also the Director for the Department of Energy Artificial

460

D.C. Rodger et al. / Sensors and Actuators B 132 (2008) 449–460

Retina Project that is a consortium of five Department of Energy labs and four
universities, as well as industry.
Yu-Chong Tai received his BS degree from National Taiwan University, and the
MS and PhD degrees in electrical engineering from the University of California
at Berkeley. He is currently a professor of bioengineering and executive officer

and professor of electrical engineering at the California Institute of Technology,
and director of the Caltech Micromachining Laboratory. His current researchinterests include flexible MEMS, bioMEMS, MEMS for retinal and neural
implants, parylene-based integrated microfluidics, neuroprobes/neurochips, andHPLC-based labs-on-a-chip.