Th1E.005

A DUAL FUNCTION PARYLENE-BASED BIOMIMETIC TACTILE SENSOR AND ACTUATORFOR NEXT GENERATION MECHANICALLY RESPONSIVE MICROELECTRODE ARRAYS
C.A. Gutierrez* and E. Meng
University of Southern California, Los Angeles, California, USA
ABSTRACT
We present the first multimodal Parylene-based
biomimetic platform with the ability to measure tactile
forceswhile
simultaneously
enabling
activestimulation/recording of neural tissue via movablemicroelectrodes.
Aliquid-filled
microchamber
encapsulating interdigitated microelectrodes performs thedual functions of actuation and tactile sensing. Theelectrodes are impedance-based tactile sensing elements
and sites for electrolytic gas generation for pneumatic
control of microelectrode position. Impedance sensitivityof up to 1.7 %/μm and out-of-plane positioning up to 8 μm
were demonstrated. We also report on themechanotransduction of biomimetic channel structures.

KEYWORDS
Parylene C, Tactile sensing, Movable electrodes.

INTRODUCTION
Most neural prostheses function by recording and/orstimulating tissue via microelectrodes. Some notable
examples include cortical prosthesis [1] and retinal
prosthesis [2, 3] efforts which employ microelectode
arrays (MEAs) to restore function by selective electricalstimulation of the surrounding tissues.
Traditionally, MEAs have been electrode-only andresearch efforts primarily focus on flexibility for
conformal application, high density electrodes, andimplantability. These characteristics can be addressed byusing polymers such as Parylene as the structural material.
Parylene is used for its United States Pharmacopoeia(USP) Class VI biocompatibility, mechanical strength(~4GPa), MEMS compatibility and flexibility whichenables released MEAs to conform to three-dimensional
tissue surfaces unlike MEAs bound to rigid substrates.
Although these considerations offer great improvements,
the efficacy of electronic neural prostheses, such as theintraocular retinal prosthesis, is ultimately dependent on
the electromechanical coupling between the device andtissue [3]. Factors such as proximity, temperature,pressure, and post-implantation micromotion all impact theoverall function of a neural prosthesis. These factors
cannot be adequately addressed using current electrodeonly technologies.
Here we present an enabling technology for nextgeneration multimodal MEAs. This dual function device iscapable of providing key tactile information at thetissue/device interface and has the ability to respond

978-1-4244-4193-8/09/$25.00 ©2009 IEEEmechanically by adjusting the out-of-plane electrodeposition. We also show biomimetic operation via axon-liketactile sensing structures. The quantification of
mechanically induced strains at the tissue-prosthesis
interface will provide additional understanding of thebiological and electrophysiological effects of neural
prosthetic implants.

DEVICE DESIGNThe key elements of this platform design are a liquidfilled chamber encapsulating interdigitated electrodes anda self-sealing stiction valve. The device has two possible
configurations: (1) a tactile sensor and (2) a movable
electrode actuator (Figure 1a, b). Liquid encapsulation isaccomplished by immersing the device in a liquid whichthen fills a chamber through the stiction valve port. The
device is then exposed to air where the evaporative drying
process will cause the valve to collapse due to stiction,
thereby sealing the remaining liquid in the chamber.

Figure 1: (a) Structure of basic functional unit for tactilesensing. Stiction valve enables liquid encapsulation in Parylene
chamber. (b) Movable electrode configuration; optional top
electrode is pneumatically actuated as chamber membrane
deflects upward. (c)-(d) Single fabricated devices.

Operating principles
The dual modes of operation are depicted in Figure 2.
Tactile sensing utilizes the electrochemical impedance ofthe encapsulated liquid and is dependent on liquid volume
and ionic conduction path between the two electrodes(Figure 2a). Under tactile deformation, the volumetricconduction path within the chamber is altered, analogous
to a variable resistance. This phenomenon enables
measurement of the tactile force acting on the chamber

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Transducers 2009, Denver, CO, USA, June 21-25, 2009

membrane.
Out-of-plane actuation of the top electrode isaccomplished by applying a constant current (~5 μA for afew seconds) across the chamber electrodes to initiate
electrolysis. The evolution of hydrogen and oxygen gas
within the chamber generates an internal pressure causing
the electrode to deflect upward (Figure 2b).

Figure 2: (a) Tactile sensing modality. External forces deformchamber altering liquid conductive path between bottom
electrodes. (b) Movable electrode modality. Increased chamberpressure from electrolytic gas generation pneumatically actuates
top electrode.

ri/ro and p is a numerical function of φ and υ.

DEVICE FABRICATION
A standard soda-lime wafer was optionally treatedwith silane A-174 Parylene adhesion promoter. Released
devices did not require this step. Parylene C was deposited(~10 μm) on the entire wafer. Photoresist was then spunon and pattered for metal lift-off forming chamberelectrodes, leads, and contact pads. A sacrificial
photoresist layer (10-22 μm) of AZ 4620 was patterned todefine the valve and chamber structures. Parylene (2-8μm) was then deposited followed by an optional dual-layer
photoresist (AZ1518/AZ4620) lift-off process to pattern
the stimulation/recording electrodes atop the chamber
structures. The dual layer lift-off process with a globally
exposed bottom layer was necessary to achieve continuity
between substrate and the top of the chamber. TheParylene covering the contact pads and stiction valve port
was etched using oxygen plasma (100 W, 100 mT).
Devices were diced and soaked in acetone to remove thesacrificial photoresist. After rinsing in isopropyl alcoholand DI water, dies were immersed in an electrolyte (DIwater) (Figures 3, 4).

Stiction valve
The design guidelines for annular stiction valves werepreviously investigated in [4, 5]. Stiction occurs for criticalnumber N c > 1 wherep (φ ,υ )(1)
Dh 2
with σ = 2γ LA cos θ C , D = ( Et 3 / 12(1 − υ 2 )) and where γLA
is the liquid-air surface tension, θC is the Parylene contactangle, E is Young’s Modulus, t is the membrane thickness,
h is the membrane height, ν is Poisson’s ratio, φ is the ratio
Nc =Figure 3: Released flexible array.

σro 4

RESULTS AND DISCUSSION
Stiction valve
Valves (OD: 330 μm, ID: 100 μm) were fabricated andsealed within a few minutes of evaporation drying atambient conditions (Figure 5).

Figure 4: Fabrication process for tactile sensor/actuator platform. Steps 1-8 show tactile sensor fabrication. Steps 1-8 plus A-C(added between steps 4 and 5) show movable electrode fabrication.

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Figure 5: (a) Stiction valve open. (b) Stiction valve closed. (c)Nomarski image highlighting regions of collapse. Border fringes
indicate areas where membrane has not touched down.

membrane surface due to electrolysis (Figure 9). On
average the maximum deflection increases with larger
chambers. This is due to the greater compliance of thelarger membranes which allows for greater deflections.
Maximum deflections of ~8 μm were achieved with
500 μm square chambers. Corrugated membranes or
bellows may extend the actuation range but electrodetethering will need to be addressed.

Experimental setup
Experiments validated the two modes of operation:
impedance-based tactile sensing and actuation ofelectrodes. Tactile sensing was evaluated by indenting thechamber membrane with a micromanipulator stylus. Stylus
displacement was measured under a 100x microscope
objective with 1μm vertical resolution. The same
technique was used to measure electrode displacement.
Impedance was measured using a voltage divider withactive rectification to attain the DC level of the AC
response (Figure 6). A LabVIEW-controlled Keithley
2400 precision multimeter was used to record the real-time
impedance response. The circuit was calibrated using an
Agilent E4980A precision LCR meter.

Figure 7: Percent change in solution impedance vs. membrane
indentation depth for circular chamber devices. Error bars
correspond to measurement resolution of microscope. Table lists
the maximum impedance sensitivities obtained.

Figure 6: Impedance measurement circuit.

Tactile sensing
Impedance-indentation results for circular chamberswith diameters ranging from 300 to 500 μm were obtained.
Percent change in impedance per micron of indentation
was greatest for 500 μm chambers with a sensitivity of
1.69 %/μm (Figure 7). Indentation depth was limited by
the as-fabricated chamber height and is easily increased.
This result indicates a trade-off between sensor response
and size. The greater displacement of liquid volumes inlarger chambers creates larger impedance responses at thecost of overall footprint.
Figure 8 shows real-time impedance based tactile
detection capability. A stylus was used to indent themembrane down to the substrate several times over the fullvertical travel of the suspended membrane height.
Mechanically responsive electrodes
Electrode actuation was demonstrated by applying aconstant current (~5 μA for a few seconds) across thechamber electrodes and measuring deflection of theFigure 8: Real-time tactile sensor operation capturing several
quick indentations followed by a sustained indentation throughthe full dynamic range of devices. Intermediate levels ofdeformation were also readily detectable.

Figure 9: Electrolysis actuation. Maximum chamber deflection
for square chambers. 5 μA is typically applied for about 5seconds to achieve deflection.

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Modeling
Nonlinear finite-element modeling (FEM) was used tosimulate membrane tactile deformation under the stylus tip(Figure 10a). Forces of 6 mN are required for maximal
deflection of 8 μm thick membranes. The force sensitivityof the device depends on chamber parameters, such as
membrane thickness and chamber size [6], and can be
designed to achieve a desired range. Forces acting locallyon tissues at a neural prosthetic interface may range fromseveral hundred micronewtons to a few millinewtons [7-9].

and is a modality of interest for new bio-inspired devices.

Figure 12: Axon-mimetic configuration for tactile detection.

CONCLUSION
A dual-function device for next generation MEAs wasdesigned and fabricated. The integration of tactile
information with electrode actuation enables smart MEA
interfaces that can dynamically adjust to specific patientneeds and provide an improvement over electrode-only
designs. Work is under way to integrate these devices withcurrent neural prostheses for in-vivo experiments.

ACKNOWLEDGEMENTS
Figure 10: (a) Nonlinear FEM simulations of membrane
deflection under a normally applied force to center of chamber.
6 mN required for ~18 μm membrane indentation.
(b) Resultant stress distribution with 0.1 psi internal chamberpressure and (c) membrane deflection of 10 μm.

Linear modeling of membrane actuation was used tooptimize routing of the top electrode leads (Figure 10b).
The low stress in the connecting channel region during
deflection ensures minimum membrane tethering by leads
and reduces the possibility of metal fracture during
deflection (Figure 11).

Figure 11: Arrayed devices for movable MEA. Electrode leads
are optimally routed over the connecting channel.

Biomimetic tactile sensing
Fabrication of channels enables axon-like structures tomimic mechanotransduction phenomenon seen in neuronalsignal propagation. Multiple electrodes can populate achannel segment allowing for downstream tactiledetection. An indentation of the channel was detectable up
to 1 mm away from the measurement site (Figure 12). The
ability to detect tactile information along the length of achannel extends the physical sensing range of the sensor

This work was funded in part by Engineering ResearchCenters Program of the NSF under Award Number EEC0310723 and the Bill and Melinda Gates Foundation.

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CONTACT* C.A. Gutierrez, tel: 213-821-3949; cagutier@usc.edu

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