Polymer-Based Microelectrode Arrays

Scott Corbett

1,2

, Joe Ketterl

, and Tim Johnson

MicroConnex, 34935 SE Douglas St., Suite 110, Snoqualmie, WA, 98065

Advanced Cochlear Systems, 34935 SE Douglas St., Suite 110, Snoqualmie, WA, 98065

ABSTRACT

We have developed flexible, polymer-based electrodes for potential medical applications

including neural recording and stimulation. Using various combinations of liquid crystal

polymer (LCP) substrates, implantable grade silicone and polyimide, we have developed and

tested several prototype multi-layer, polymer electrodes. We report here on two specific

electrodes. In the first case, a multilayer electrode consisting of high-melt temperature liquid

crystal polymer (LCP) material with patterned electrodes of sputter deposited and plated gold,

laminated together with a lower-melt temperature LCP, was produced. Iridium oxide was

deposited on the exposed electrode sites to facilitate effective charge transfer for neural

stimulation. The electrode was designed for acute implantation in a cat cochlea and contained 12

contacts, with a pitch of 200 microns. The small contact spacing allowed testing of electric field

focusing techniques both in vitro and in vivo. We subjected the electrodes to electrical and

mechanical tests to assess its likely suitability as a long-term biomedical implant. Chronic

electrical leakage testing indicated ionic permeability of the low and high temperature LCP

interface that was higher than that desired. In a second case, we produced a mock circuit using

high-melt LCP and medical grade low durometer silicone in place of the low-melt LCP as the

interlayer adhesive. Mechanical and electrical testing of the hybrid design indicated the potential

to fabricate cochlear electrodes containing up to 72 contacts with a footprint and mechanical

performance similar or better than current commercially available cochlear implant arrays

(containing up to 24 elements). Multi-layer polymer electrode technology offers the opportunity

to create new electrodes with higher numbers of channels, offering improved performance in

neural stimulation applications including cochlear implants, retinal arrays, deep brain stimulators

and paraplegic remobilization devices.

INTRODUCTION

Neurostimulation systems are an emerging class of medical devices promising

miraculous treatments for otherwise intractable medical conditions [1]. The most successful

example to date is the cochlear implant, which directly stimulates auditory neurons to partially

restore hearing in the deaf [2]. Other emerging treatments include retinal arrays to restore sight,

deep brain stimulators to treat Parkinson’s disease and paraplegic remobilization systems [3]. All

of these devices rely on electrode arrays to transfer the electrical current that elicits neural

responses. To date, most electrodes have been patterned after cardiac pacemaker electrode

technology, relying on insulated wires, platinum foil or metal electrodes and medical grade

silicone encapsulation. These mechanical assemblies are very expensive and limited in

complexity and therefore, utility. In the case of neuroprosthetic devices, particularly cochlear

implants and retinal arrays, it is highly desirable to have as high a density and as great a number

of active electrode channels or stimulation sites as possible. In this case, higher numbers of

channels implies greater information exchange to the perceptual sense. Unfortunately, there is a

limit to the number of channels that can be formed using conventional wire-based technology

and mechanical assembly.

Logically, this area is ripe for the implementation of advanced materials, fabrication and

manufacturing techniques developed in the integrated circuit and flexible printed circuit board

industries. Examples include micromachining methods, advanced materials and

photolithographic based etching techniques developed for micro-electrical-mechanical systems

(MEMS). Indeed, silicon-based microelectrodes, developed by researchers at the University of

Michigan, have been used in animal studies [4]. The literature also contains references to planar

electrodes manufactured using polyimide, a flexible substrate media commonly used for flexible

circuit fabrication. However, planar manufacturing methods introduce problems of their own.

Materials compatible with existing manufacturing methods are either stiff (i.e. silicon) or

perform poorly when exposed to body fluids. The dielectric properties of polyimide have been

demonstrated to degrade during in vitro testing in saline solution [5]. Another challenge is

converting the planar structure to a three-dimensional implant.

A new class of dielectric laminates have been developed, based on counter-rotating die

extrusion of liquid crystal polymer (typically Vectra A950, a thermotropic Type I all-aromatic

polyester (Ticona, Florence, KY) [6]. LCP has excellent dimensional stability, low moisture

absorption and is potentially available as a Class VI FDA implantable grade material. The

coefficient of hygroscopic expansion of LCP is 1.5 ppm/%, and its water absorption is 0.1 %,

which is ten times less than that of polyimide. The superior electrical and mechanical properties

of LCP, plus the compatibility with planar circuit fabrication techniques, make LCPs compelling

candidates for high-density neural implants.

We explored LCP materials and planar fabrication methods intended for two neural

stimulation applications: cochlear implants and retinal stimulation arrays. We report here on the

prototype cochlear implant arrays that we developed.

EXPERIMENT

We developed a multilayer polymer circuit architecture and fabrication process (Figures

1 and 2). [7,8].

We used commercially available LCP laminates in various combinations supplied by

several vendors including: Foster-Miller (Waltham, MA), Gore Japan (no longer in the US

market), Nippon Steel (NY, NY), Rogers (Rogers, CT), and Yamaichi Electronics (Japan).

Individual 50 µm thick sheets of high-temp LCP laminate (T

=335°C) were plasma etched using

an Advanced Plasma Systems Model B-Series-24-2 (250 mTorr pressure: 19% Ar, 81% O

,), and

sputter deposited with 200 angstroms of chromium followed by 2500 angstroms of gold (using a

modified MRC sputter system (KDF, Rockleigh, NJ) with computer-controlled 6 kW DC

Advanced Energy power supply). Traces were patterned using a proprietary UV laser direct-

write process utilizing an Electro Scientific (ESI, Portland OR) Model 5200 micro-machining

system employing a tripled YAG (355 nm) laser.

1a 1b

Figure 1a. Multilayer electrode schematic showing multiple layers of high melt-temp LCP

(darker substrates) patterned with sputter deposited gold trace elements, laminated together using

lower melt-temp LCP substrate material (lighter color substrate). Traces are routed to the side of

the structure connecting with UV laser-drilled openings, which are subsequently metalized with

gold followed by iridium oxide. The structure is patterned with a UV laser, cutting through the

plated openings and dielectric layers to create electrode sites (200µm spacing). Alternative

versions include placing the electrode sites on the top or bottom surface, connected through the

openings, which are insulated through subsequent application of dielectric. Figure 1b.

Completed electrode.

Figure 2a. Schematic of cross-section of multi-layer laminate with one trace layer, showing high

melt temp LCP layers (one layer of which has traces patterned shown in red) laminated together

with lower melt-temperature LCP. Figure 2b. Actual cross-section.

50 µm thick, low-temperature LCP laminate (T

=280°C) was placed between the trace

layers as an adhesive, forming a multilayer stack. The layers were aligned by placing laser-cut

alignment pinholes referenced to the trace pattern on each layer and placing the laminates

together in a custom-built pin alignment fixture [9]. The entire structure was then placed into a

Carver Model C laboratory press (Wabash, IN) with added custom heating platens and Watlow

temperature controllers (St. Louis, MO), after which a temperature/pressure/time profile was

applied to complete the lamination. The lamination process relies on differential melt profile of

the low and high temp LCP materials. The low melt LCP slightly melts to form an adhesive

interface between the high temperature LCP layers, which remain stable during the lamination

process.

Hi-temp LCP

Low-temp LCP

Gold trace

200

50 µ m

High melt LCP

Low melt LCP

Gold electrode

Gold traces

After lamination, openings were laser drilled through the laminate, followed by sputter

deposition of gold to fill the openings, after which electrode sites were patterned on the top

layers using direct-write laser patterning, followed by electrolytic plating of additional gold. A

space filling circuit overmold was incorporated on the array to better position the contacts near

the modiolus of the cat cochlea. NuSil MED-4011 (Carpinteria, CA), a medical grade silicone

currently used in commercial implants, was chosen as an overmold material, using an EDM-

processed aluminum mold as the forming fixture. The resultant electrode is shown in figure 3.

DISCUSSION

We performed several tests to assess the suitability of the arrays for short- and long-term

implantation. A series of five cat implants were performed by researchers at the University of

California, San Francisco, which demonstrated the suitability of the arrays for short-term use in

animal subjects. Although beyond the scope of this paper, the researchers were successful in

using the LCP high-density arrays to demonstrate field-focusing methods to reduce channel

interaction in a mammalian cochlea [10].

Electrical leakage testing was performed to assess the ionic permeability of the LCP

structure, particularly the LCP-LCP interface. Measurements were made on bulk LCP samples,

bulk silicone (Nusil MED-6640) and multilayer laminate LCP structures with embedded traces.

Samples were soaked in buffered saline solution with a voltage bias (ranging from 1-10V DC)

applied to the traces within the sample, and leakage to a reference platinum electrode measured

using an electrometer (Keithley Instruments, 6517A). Table A summarizes the bulk leakage

measurements of the LCP and silicone materials converted to resistivity values.

Table I. Measured electrical resistivity of bulk and fabricated samples.

Sample Resistivity

(ohm-cm)

Immersion Time (weeks) at

37°C

Bulk LCP 5.0 x 10

Bulk Silicone 6.6 x 10

Embedded traces in LCP

multilayer circuit (edge of array)

2.9 x 10

Embedded traces in LCP

multilayer circuit (center of array)

1.9 x 10

Figure 3. LCP array potted into silicone

mold to form space-filling cochlear implant

for cat. White substrate is LCP material

with gold stimulation contact shown coated

with iridium oxide. Photo courtesy of

Steven Rebscher and Russell Snyder of

University of California, San Francisco.

200 µm

As seen in table I, there was a significant decrease in resistivity from the bulk material

values in the multilayer electrodes we tested, indicating ionic penetration through the LCP-LCP

interface, even though these structures showed good mechanical adhesion as demonstrated

through pull testing (typical adhesive strength 1 lb/in). Although we were successful in using the

arrays in a short term acute implant setting, the decrease in resistivity indicates the arrays would

likely fail in a long-term implantable context. There were other significant problems encountered

in producing the arrays. The process window for successful mechanical adhesion of the laminate

substrates is extremely narrow with too much heat and pressure leading to dimensional

instability of the dielectric layers, and too little pressure leading to overt delamination.

In an attempt to understand the increased leakage, we imaged the LCP surfaces using

scanning electron microscopy (SEM). SEM photographs of the low-temp LCP substrate after

plasma cleaning and before lamination are shown in figure 4. The SEMs indicate filler materials

(copolymers and silicon dioxide) have been added to the bulk polymer, which may increase the

ionic permeability of the interface.

Figure 4. SEM image of LCP material before (left) and after (right) plasma etching. Plasma

etching conditions were 20 minutes at 150 W and 300 mTorr in oxygen plasma. Residue

remaining at the surface is likely ceramic powder or molecular “grains” which have a lower etch

rate than the bulk LCP polymer. The vendor reported the LCP has ceramic filler, up to 5% by

weight. EDAX spectra revealed a significant Si content in the LCP, indictating the material also

contains a small amount of silica-based filler. Magnification is 4,000X.

We were motivated to develop alternative designs, not only on account of the loss of

electrical insulation shown in the LCP arrays, but also because the stiffness of the finished array

was higher than desired for the intended cochlear implant application. In an alternative design we

have replaced the low-temp LCP adhesive layer with medical grade, low durometer silicone

adhesive (NuSil 6640). Initial testing indicates significantly improved mechanical performance

with electrical testing underway. Multilayer structures were fabricated in a form factor designed

to eventually incorporate 72 channels in a human sized cochlear implant. The flexibility of the

arrays compared favorably with existing commercially available implants, indicating that if the

electrical leakage of the multilayer structure is stable, the devices may be suitable for ultimate

long-term implantation.

CONCLUSIONS

We have developed and tested a multilayer LCP electrode containing 12 electrode sites

spaced at 200 µm. The arrays were successfully used in an acute animal implant, but showed a

reduction in electrical isolation after soaking in buffered saline for several months. An alternative

design incorporating medical grade low durometer silicone as a replacement adhesive between

LCP layers is currently under evaluation, with early results indicating improved performance for

the intended cochlear implant application

ACKNOWLEDGMENTS

This work was funded by NIH Phase II SBIR Grant, 5 R44 DC004614-03 (Advanced

Cochlear Systems), and NIH SBIR Grant 2 R44 NS40647-02 (MicroConnex).

REFERENCES

At the date this paper was written, URLs or links referenced below were active. Neither

the authors nor the Materials Research Society warrant or assume liability for the content or

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