1352

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 19, NO. 6, DECEMBER 2010

Parylene-Based Electrochemical-MEMS Transducers
Christian A. Gutierrez, Member, IEEE, and Ellis Meng, Senior Member, IEEE

Abstract—We report the design, fabrication, and characterization of electrochemical microelectromechanical systems(EC-MEMS) devices featuring encapsulated fluid as the basis
for transduction. Parylene microstructures, including discretechambers (square or circular geometry), are utilized as physical transducers for electrochemically mediated liquid impedance
transduction of physical phenomenon such as contact and force.
Parylene-based EC-MEMS technologies uniquely leverage advantages in size (< 500 μm diameter), packaging (no hermetic packaging necessary), power (nanowatts to microwatts), and flexibility
to address the physical sensing requirements of in vivo applications. Robust EC impedance (EI) sensor responses (up to 20%
from base-line) and discrimination of 200-nm chamber deflections
were possible using the EI transduction technique. Additionaltransducer configurations enabling electrolysis-based out-of-planeactuation and biomimetic mechanotransduction in microfluidic
channels are also presented.
[2010-0157]
Index Terms—Biomimetic, contact sensor, electrochemical
microelectromechanical systems (EC-MEMS), EC sensing, EC
transduction, electrolysis actuation, force sensor, impedance-based
sensor, liquid encapsulation, liquid impedance, mechanotransduction, Parylene C.

I. I NTRODUCTIONE

LECTROCHEMICAL (EC) transducer technology is wellestablished and used in a variety of applications. Electrochemically based sensing, for example, is broadly applied tofields ranging from analytical chemistry to macromolecular andcellular detection and analysis [1]. It was extensively appliedin capillary electrophoresis-based systems for macromolecular characterization of deoxyribonucleic acid and neurotransmitters [2]–[4] and gained popularity as a microfluidic-based
sensing and actuation scheme for lab-on-a-chip and micro total
analysis systems technologies [1], [5]–[7]. EC sensing was
perhaps most famously utilized in the invention of the Coultercounter, by Wallace H. Coulter, in the late 1940s [8], which

Manuscript received May 28, 2010; revised July 14, 2010; accepted
August 5, 2010. Date of publication October 11, 2010; date of current version
November 30, 2010. This work was supported in part by the Engineering
Research Centers Program of the National Science Foundation (NSF) underAward EEC-0310723, in part by an NSF CAREER Award (ECS-0547544),
and in part by the Bill and Melinda Gates Foundation (CG). Subject Editor
C.-J. Kim.
C. A. Gutierrez is with the Department of Biomedical Engineering and
the Biomimetic Microelectronic Systems Engineering Research Center, University of Southern California, Los Angeles, CA 90089 USA (e-mail:cagutier@usc.edu).
E. Meng is with the Department of Biomedical Engineering, University ofSouthern California, Los Angeles, CA 90089 USA, and also with the Ming
Hsieh Department of Electrical Engineering and also with the Biomimetic
Microelectronic Systems Engineering Research Center, University of Southern
California, Los Angeles, CA 90089 USA (e-mail: ellis.meng@usc.edu).
Color versions of one or more of the figures in this paper are available online
at http://ieeexplore.ieee.org.
Digital Object Identifier 10.1109/JMEMS.2010.2076791

utilizes the EC Coulter principle for particle detection andanalysis.
Despite the widespread use in such biological and chemical analysis systems, EC sensing technology has seen little
application in other modes of transduction particularly thatof physical quantities such as contact detection, force, andpressure. The inherent simplicity and sensitivity of EC transduction lend itself well to physical transduction applications.
Monitoring fluid response to an external physical stimulus via
EC impedance (EI) requires only the interfacing of electrodes
with the electrolyte fluid and the application of low-power electrical measurement signals; both are easily realized in standard
microelectromechanical systems (MEMS) devices. However,to date, EI-based sensing has been exploited only in limited
applications. Large-format EI-based tactile sensors for robotics
have demonstrated detection of multiple parameters, including
point contact, force, direction, object shape, and texture [9]–[11], and bubble-based hydrostatic pressure transduction wasshown within microfluidic channels using EI-based transduction [12]. EC-based physical transduction is an emerging field
ideally suited to the unique tools and capabilities of MEMS
fabrication techniques. EC-MEMS technology leverages theability to precisely pattern and define features (both electricaland mechanical) at the micrometer scale and the proven capability of microfluidic technology to realize novel transducer
modalities.
Polymer-based EC-MEMS is of particular interest for its potential application to in vivo settings where sensing of physicalparameters such as contact, force, pressure, and strain mustoccur in a compliant low-modulus wet environment. These
conditions are particularly difficult for standard transduction
techniques, such as piezoresistive or capacitive, which generally cannot operate in wet conditions and often require hermetic packaging. Quantification of parameters such as contact
force and pressure is of particular importance for neuromodulation technologies such as epiretinal electrode arrays forretinal prostheses that intimately contact nervous tissues atscales below 500 × 500 μm2 [13], [14]. Excess mechanicalloading of sensitive tissues, such as retina, can lead to adverse
effects, including mechanical insult and apoptotic cell death[15]–[17]. Thus, understanding the extent of this acute andchronic contact is critical to maintaining tissue viability, device
efficacy, and overall functionality. However, the magnitude ofthis mechanical interaction has not been quantitatively studied
and is currently unknown. There is evidence that suggests thatmicronewton-to-millinewton forces can be expected [18]–[21];however, measurement of these interfacial forces in situ requires novel sensor technology. Polymer-MEMS sensors basedon EI transduction offer a novel approach with many advantages, including improved mechanical matching to tissue (in1057-7157/$26.00 © 2010 IEEEGUTIERREZ AND MENG: PARYLENE-BASED ELECTROCHEMICAL-MEMS TRANSDUCERS

Fig. 1. Model of Parylene-based EC-MEMS device. An electrolyte-filled
microchamber forms the core of the transducer, and an in-line stiction valve
provides liquid encapsulation capability. The stiction valve, as depicted, is open.
(Device layout adapted from [23].)

contrast to rigid materials such as silicon and glass), small
footprint, low power consumption, simple construction, andexcellent sensitivity [22].
In this paper,we present the development of a Parylene-based
EC-MEMS device utilizing encapsulated fluid to mediate transduction designed specifically for the measurement of interfacial
contact forces between implants and tissue. The key elementsof the device design are an electrolyte-filled Parylene chamber
which provides a soft compliant contact surface and a minimumof two microelectrodes exposed to the fluidic contents of thestructure (Fig. 1). The exposed electrodes serve as the electricalinterface for EI measurements. To ensure liquid entrapment
within the chamber structure, an in-line annular-plate stictionbased valve is utilized. The stiction valve design and concept
were previously described by Zhang et al. [23]. Here, theirapplication to EI-based contact sensing at scales below 500 ×
500 μm2 using Parylene microchamber structures is presented
and discussed. Additional EC-MEMS applications related tomicroactuation and fluid-enabled mechanotransduction are alsopresented.

1353

At sufficiently high frequencies (> 1 kHz), Cdl essentiallyacts as a short, and the solution resistance RS dominates the
impedance response. It is generally desirable for the fluid tohave a high resistivity to ensure that the measured impedanceof the series circuit (electrode leads plus fluid) is dominatedby RS . Consequently, the impedance response can be approximately modeled as a simple classical resistance, where ρ isthe resistivity of the electrolyte, l is the distance between theelectrodes, and A is the cross-sectional area of the fluid between
the electrodes [see (2)]. The impedance is therefore inversely
proportional to the cross-sectional area between the electrodes
and is directly analogous to a variable resistor. This phenomenon enables the transduction and measurement of physicalquantities such as contact force and contact pressure acting on
the electrolyte-filled chamber
Z=

1

+ RS
1
Rct + jwCdl|Z|f >1 khz ≈ RS =

ρl
.
A

(1)
(2)

It is desirable to apply an excitation signal to the electrodesystem in such a way that the voltages developed across themetal–electrolyte interfaces are sufficiently low to avoid nonreversible Faradaic reactions, which could potentially corrode
the metal contacts and cause electrolysis of the electrolytic
fluid [26]. This is accomplished by applying a small alternating
current (typically < 1 μA) and measuring the resulting voltage
across the electrodes. As a result, this measurement is low
power and typically requires no more than a few microwatts.
At these low powers, no significant increase in temperature
is observed as most of the energy is expended by reversible
isothermal Faradaic reactions at the electrode surface, resulting
in negligible ohmic losses.
III. D EVICE D ESIGN

II. BACKGROUND AND T HEORETICAL C ONSIDERATIONSWhen a voltage is applied to an electrode in contact with anelectrolyte, the electrode interacts capacitively with the electrolyte, generating ionic current flow by non-Faradaic mechanisms [24]. John E. B. Randles, one of the most important
theorists in the field of electrochemistry, devised a circuit
model in 1947 to describe the electrode–electrolyte junction[25]. He described the interface as a capacitor, the Helmholtz
double-layer capacitance Cdl , in parallel with a resistance tocharge transfer Rct . This parallel combination describes theelectrode–electrolyte interface and is in series with a solution
resistance RS [see (1)].
By applying an alternating current, the impedance of theelectrolyte volume around the electrodes can be measured. In
our device structure, the electrolyte volume is confined near
the electrodes by a polymer microchamber. Thus, externallyapplied forces to the microchamber result in volumetric deformation of the electrolyte that induces corresponding volumetricvariations to the fluid path of current-carrying ions between theelectrodes. The resulting impedance changes contain information about the magnitude of the externally applied force (Fig. 2).

Device architecture consists of the following: 1) anelectrolyte-filled Parylene chamber that provides a soft compliant contact surface and 2) interdigitated electrodes exposedto the fluidic contents trapped in the chamber that serve as theelectrical interface for EC measurements (Fig. 3). Microchamber structures of square and circular geometries were fabricated. The primary design goal was to enable the controlled
entrapment of small liquid volumes on the order of ∼1 nL
for physical EC transducers. A polymer structural material
was favored because of the material compliance enabling the
fabrication of an easily deformable structure at scales below
500 × 500 μm2 . Furthermore, the development of such technologies for implantable or in vivo applications warrants theuse of flexible biocompatible materials that are amenable tostandard microfabrication techniques. These design goals wereaddressed by using Parylene C (United States Pharmacopoeia
Class VI biocompatibility) as the structural material in theconstruction of our devices [27], [28]. Additionally, the ability
to fabricate these devices completely out of Parylene C enables
the release from rigid substrates and provides the ability toplace these sensors on nonplanar surfaces (such as tissue) and1354

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 19, NO. 6, DECEMBER 2010

Fig. 2. Depiction of EI sensing principle using an encapsulated liquid volume. (a) Under applied forces, the cross-sectional area between the electrodes decreases,
resulting in an increase in RS and, thus, impedance magnitude. (b) Randles equivalent circuit model, where Cdl is the electrode double-layer capacitance, RS isthe solution resistance, and Rct is the charge transfer resistance.
TABLE ID EVICE D IMENSIONS

Fig. 3. Key dimensional design parameters for transducer and annular platestiction valve. Dimension a represents the side length for square chambers anddiameter for circular transducer chambers.

in other locations and environments which are very difficult toaccess with traditional materials.
To minimize footprint and provide automatic liquid encapsulation capability to each device, an in-line Parylene-based stiction valve was integrated next to the sensing chamber. Due tothe liquid-mediated device functionality and the desire to avoid
tedious encapsulation techniques such as epoxy or sealants, aself-sealing approach was adopted through the use of a stiction
valve. Parylene-based stiction valves were successfully demonstrated for long-term liquid encapsulation and employ simple
design guidelines and straightforward fabrication techniques[23], [29]. Integration of the automatic fluid sealing mechanism
in line with the devices also enables batch sealing of multiple
devices on a wafer, providing significant packaging advantages
over traditional serial fluid sealing techniques.
A. Stiction-Mediated Liquid Encapsulation
Stiction is usually considered a notorious and unwanted
failure mechanism in MEMS and can render devices inoperable
unless accounted for appropriately [30], [31]. This phenomenon, however, is beneficial when utilized as a valve sealingmechanism. The analysis and design guidelines for annularplate stiction valves were previously reported in [23] and [32]and adopted here. Briefly, stiction of a suspended Parylene
annular plate occurs for critical number Nc > 1, where
Nc =

σro4
p(φ, υ)Dh2(3)

with surface tension term σ = 2γLA cos θC and flexural rigidity
D = (Et3 /12(1 − υ 2 )). Here, γLA is the liquid-air surface tension, θC is the Parylene contact angle, E is Young’s Modulus,
t is the membrane thickness, h is the membrane height, υis Poisson’s ratio, φ is the ratio ri /ro , and p is a numericalfunction of φ and υ. To ensure liquid entrapment within thechamber, stiction valve dimensions were selected to ensure thatthe condition Nc  1 was satisfied (Table I).
IV. D EVICE FABRICATION
The devices were fabricated utilizing standard surface micromachining techniques on a 3-in soda-lime wafer (see Figs. 4
and 5). The entire process was achieved at relatively low
temperatures (below 100 ◦ C), which is compatible with avariety of polymers. The wafer was optionally treated with
Silane A-174 Parylene adhesion promoter (Specialty Coating
Systems, Inc., Indianapolis, ID). Inclusion of the adhesion
promoter ensures that the underlying Parylene substrate, uponwhich the devices are surface micromachined, adheres to thesoda-lime glass supporting substrate during wafer dicing. A
10-μm Parylene C film was then vapor deposited on the entire
wafer. A 2000-Å layer of platinum was deposited by e-beam
evaporation and patterned by a liftoff process briefly described
here. A dual-layer photoresist process was used to create
an undercut sidewall profile to facilitate metal liftoff. First,
AZ1518 photoresist (AZ Electronic Materials, Branchburg, NJ)
was spun at 4000 r/min followed by global exposure. Then,
AZ4400 photoresist (AZ Electronic Materials, Branchburg, NJ)
was applied at 4000 r/min and patterned. Following a short
descum in oxygen plasma, the Pt film was e-beam evaporated
and patterned, forming the chamber electrodes, leads, andcontact pads. A sacrificial photoresist layer (h = 10−22 μm)
of AZ4620 (AZ Electronic Materials, Branchburg, NJ) was
spun and patterned to define the chamber and valve structures(ro = 165 μm). A second layer of Parylene (t = 2−8 μm) wasGUTIERREZ AND MENG: PARYLENE-BASED ELECTROCHEMICAL-MEMS TRANSDUCERS

Fig. 4. Fabrication process for EC-MEMS Parylene-based contact sensor.
(a) Parylene deposition. (b) Deposit and pattern electrodes. (c) Pattern sacrificial photoresist. (d) Deposit Parylene. (e) Etch Parylene over valve. (f) Dissolve
sacrificial photoresist and replace with electrolyte. (g) Initiate evaporative
stiction valve closure. (h) Release from substrate.

1355

To prepare the devices for use, it was necessary to remove thesacrificial photoresist contained within the Parylene transducer
chamber and fill the device with the desired electrolyte. This
was accomplished by first immersing the entire die in an
acetone bath for approximately 30 min. The acetone dissolved
the protective photoresist layer left on the surface from theprior dicing step as well as the sacrificial photoresist within
the Parylene chamber which escaped through the valve orifice
opened in the final Parylene etching step. Following immersion
in acetone, the devices were immediately transferred to an isopropyl alcohol bath to rinse away the acetone and dissolve thephotoresist (∼15 min). The devices were then transferred to adeionized (DI) water bath to rinse away any remaining solvents.
Following the rinse, the devices were placed in a bath of thedesired electrolyte for about 30 min to ensure complete filling.
The entire sacrificial photoresist removal and filling process isdiffusion driven, so adequate time is required for each complete
solution replacement step within the Parylene microchamber.
Devices were then removed from the final bath and exposed tothe ambient laboratory environment (∼ 25 ◦ C and 35% relative
humidity) for drying. Within minutes, evaporation through thevalve orifice created the necessary conditions for the stictioninduced collapse of the annular valve plate to the substrate,
thereby encapsulating the remaining liquid in the chamber.
V. E XPERIMENTAL M ETHODS
A. Impedance Measurement Interface

Fig. 5. Fabricated Parylene microchamber structure with in-line stiction
valve. Interdigitated platinum electrodes are integrated and exposed to thechamber liquid environment.

Fig. 6. (a) Released Parylene film with integrated EC-MEMS devices.
(b) Released device film adhered to curved surface.

then deposited. The Parylene over the contact pads and valve
orifice (ri = 50 μm) was then etched in oxygen plasma (100 W
and 100 mT) using a photoresist etch mask (AZ 4620). Due tothe comparable etch rates (approximately 1:1) of Parylene andphotoresist, the thickness of the etch mask was at least 1.5 times
the thickness of the Parylene film being etched. After etching,
the remainder of the photoresist etch mask was left on the surface to serve as a protective layer during dicing. Square test dies
(1 cm2 ), each containing several test devices, were diced forcharacterization and testing. For wafers untreated with A-174,
devices were released at this stage by cutting the Parylene filmwith a razor blade and peeling from the substrate (Fig. 6).

Impedance measurements were made with a high-precision
impedance converter board (AD5933 Analog Devices, Norwood, MA). The AD5933 combines an onboard frequency
generator with a 12-bit 106 sample-per-second analog-to-digitalconverter (ADC). The frequency generator allows an external
complex impedance to be excited at a known frequency. The
impedance response signal was sampled by the onboard ADC,
and a discrete Fourier transform was obtained using onboard
signal processing. The converter board was interfaced via
USB to a PC where a custom LabVIEW interface (National
Instruments, Austin, TX) was developed for real-time dataacquisition. An ac drive signal of 1 Vpp at 5 kHz was selected
to avoid any irreversible Faradaic reactions at the electrodesurface. At this frequency, the double-layer capacitance was
effectively shorted, bypassing the charge transfer resistance,
which allowed the solution resistance to dominate the measured
impedance.
B. Experimental Setup
Characterization experiments were conducted on a 1-cm2
test die mounted on a glass slide. A drop of mineral oil was used
to minimize evaporative water loss through the Parylene. DI
water was selected as the electrolyte because of the inherently
high solution resistance. Due to trace impurities and dissolved
gases, such as carbon dioxide, from the ambient environment,
DI water exhibits a finite and measurable impedance. Impedance measurements were generally in the megaohms range
at the measurement frequency.

1356

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 19, NO. 6, DECEMBER 2010

Fig. 8. (a) Open stiction valve immediately following removal from bath.
(b) Closed valve. Interference rings (shown in contrast-enhanced inset image)indicate areas where the membrane transitions from collapse (center) to freestanding (periphery).

Fig. 7. Experimental setup.

The experimental setup was built on a floating optical table
and consisted of a motorized micropositioning stage with submicrometer accuracy (Thorlabs Inc., Newton, NJ) to precisely
control the indentation depth of an attached microprobe stylus
tip (Fig. 7). The stylus tip had a diameter of approximately
100 μm. The motorized stage and impedance converter board
were interfaced through a custom LabVIEW interface for realtime data acquisition.
VI. R ESULTS AND D ISCUSSION
A. Liquid Encapsulation
Stiction valve sealing and liquid encapsulation capability
were evaluated by using a sample die containing several identical devices under optical observation. The devices were filled,
as described previously, with DI water as the electrolyte. The
die was then removed from the bath and exposed to ambientlaboratory conditions to initiate fluid encapsulation. Successful
stiction valve sealing was clearly evidenced by the appearanceof a dark central region indicating conformal contact between
the annular plate and substrate. Emergence of interference rings
between the central region and the periphery (as the valve platetransitioned from being pinned to the substrate to freestanding
near the anchored edges) indicated complete circumferential
sealing (see Figs. 8 and 9). Encapsulation of water within
Parylene membranes was previously demonstrated [23], [29],[33]. Although Parylene has found ubiquitous application asa protective encapsulant, it is not hermetic. The water vapor
transmission rate (WVTR) through the Parylene material itself
is the limiting factor to encapsulation lifetime. Furthermore, forParylene films less than about 8 μm in thickness, pin holes,
pores, and defects can exist that contribute to gas transport
permeation through the film, thereby limiting the moisturebarrier performance [34].
The WVTR (5.6 × 10−12 g · μm/μm2 · h) of Parylene
films was previously reported in literature [35]. Although DI
water was utilized as the electrolyte here, alternative solutions,such as ethylene glycol-based conductive fluids, may be used

Fig. 9. (Top) SEM of collapsed stiction valve. (Bottom) SEM cross section
taken through the transducer chamber.

for increased resistance to evaporation. For devices with
2-μm-thick Parylene membranes, the water vapor permeation
through the Parylene film imposed a significant limitation totesting in the open-air laboratory environment. The volume
of encapsulated liquid varied significantly over time, thereby
introducing an unacceptable level of drift in addition to limiting
total encapsulation time. Several methods exist to address thisissue. Increasing the Parylene membrane thickness increases
the encapsulation time proportionally, but a thicker membraneis also less compliant and results in a lower critical number Nc
required for stiction valve operation. Annealing of Parylene
films (200 ◦ C) was shown to reduce the WVTR by nearly 50%
in addition to improving Parylene–Parylene adhesion [14],[35]. Additional coatings such as thin film metals have also
improved barrier performance [23]. To mitigate evaporative

GUTIERREZ AND MENG: PARYLENE-BASED ELECTROCHEMICAL-MEMS TRANSDUCERS

1357

Fig. 10. (a) Square chamber with interdigitated electrodes. (b) Circular chamber with circumferential and central electrodes.

Fig. 12. Time-averaged impedance responses to 1-μm deflection and release
steps. Responses are shown for one full deflection and release cycle.

Fig. 11. (Top) Real-time deflection profile of the sensor membrane. Data show
stepwise stage movement with attached stylus tip contacting sensor surface.
(Bottom) Temporal response of sensor impedance to deflection. A 250-μmdiameter sensor response is shown and is representative of typical sensor
response.

effects, experiments were performed with a drop of mineral oil
placed over the devices.
B. Contact Sensing
Devices were tested as physical contact sensors. Both squareand circular chambers were fabricated. It was found that morerobust responses were obtained with the circular chamber design. This may be in part due to the circularly symmetricelectrode spacing in contrast to the serpentine electrode gap
with the interdigitated design (Fig. 10). Applied deflections insuch a design caused cross-sectional area changes along onlya small region of the available electrode gap, resulting in small
impedance changes (< 0.25% from baseline) that were difficult
to detect using our measurement circuitry. Circular chambers of
250-, 300-, and 500-μm diameters were tested. Each device hada chamber height of 22–25 μm and a membrane thickness of
8 μm. The sensors were deflected (and released) in 1-μm steps.
Temporal sensor response was robust and clearly tracked
membrane deflection (Fig. 11). A clear sensor response wasattained with deflections as small as 200 nm (0.8% resolution of
full-scale deflection). In all experiments, deflection was limited
to 22 μm so as to not fully deform the chamber or achieve
contact with the substrate which may scratch or otherwise
damage the Parylene membrane. Impedance variations up to
20% from base line were achieved at maximum deflection, with
250-μm-diameter chambers demonstrating the large-signal response available with EI-based sensing. Sensor performancewas consistent and reproducible for individual devices, butsome variation was observed in the nominal resting (zerodeflection) impedance between devices. This variation was
typically less than 10% and was attributed to the presence ofa trace amount of residual photoresist particulates visible in thechamber electrolyte.
The impedance response to the first 10 μm of deflection wasless pronounced than the response near full deflection. Timeaveraged impedance responses clearly demonstrated this nonlinear behavior (Fig. 12). To further elucidate this relationship,
cross-sectional area was calculated using theoretical deflection
equations for rigidly clamped circular diaphragms [36]y=3(1 − μ2 )P 2
(a − r2 )2
16Et3

(4)

where y is the deflection, P is the pressure, t is the membranethickness, a is the diaphragm radius, E is the modulus ofelasticity, r is the radial distance, and μ is Poisson’s ratio.
The areas under the deflection curves were calculated andplotted against maximum deflection (Fig. 13). Cross-sectional
area through the central axis of the diaphragm is at a minimum
under the imposed deflection and is assumed to account for amajority of the resistive variations. Area clearly varies linearlywith deflection, but the reciprocal dependence of impedanceon the cross-sectional area dictates the observed nonlinear
behavior [see (2)].
Furthermore, sensor sensitivity to deflection was found tovary with chamber size with smaller chambers exhibiting
greater impedance–deflection sensitivity. This is in part due toproportionally larger fraction of the smaller chamber’s surface
area that is deflected by the stylus (fixed diameter), but generally smaller volumes (created by decreasing lateral footprint)

1358

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 19, NO. 6, DECEMBER 2010

Fig. 13. Variation of theoretical cross-sectional area with deflection.

can be expected to be more sensitive to deflections for agiven sensor height. The nonlinear sensor response is analogous
to capacitive sensors, where plate separation is modulated
and capacitance is inversely proportional to plate separation.
Linearization techniques are available to transform calibrated
nonlinear responses into a linear output. Once calibrated, linearization transformations can be easily implemented and programmed off-chip [37], [38].
The mechanical loads required to produce deflection of a
(250-μm-diameter and 8-μm-thick) chamber membrane were
evaluated analytically and with finite-element modeling (FEM,
COSMOSWorks 2009, Dassault Systèmes SolidWorks Corp,
Concord, MA). Under a linear small-deflection approximation(deflection-to-thickness ratio < 0.3) [36] for a rigidly clamped
circular diaphragm, a pressure of 1.27 MPa (62.3-mN total
force) distributed over the sensor surface was required to produce a 20-μm deflection. This value agreed well with the FEM
result of 1.36 MPa. If a thin membrane under nonlinear large
deflection is assumed (deflection-to-thickness ratio > 5) [36],a uniform 4.5-MPa (270-mN) pressure is required. The actual
sensor response in the range of 20-μm deflection (maximum
deflection-to-thickness ratio ≈ 2.5) is expected to be boundedby these two conditions. This response is sensitive to membranethickness and radius, scaling with the thickness cubed andinversely with the radius to fourth power [see (4)]. Therefore,
the membrane thickness and lateral chamber dimensions can beused to effectively tune the sensor response over a wide rangeof loading forces (micronewton to millinewton).
Impedance-based detection of surface deformations using
a fluid-filled large-format (1-cm2 ) sensor was previously described by Helsel et al. using a planar grid of gold–chromium
electrodes patterned on a rigid Al2 O3 substrate, ethylene
glycol-based conductive fluid, and latex skin [39]. Additional
physical parameters such as force and frequency were measured
by Wettels et al. using an array of integrated wires within abiomimetic model of the human digit [9]. Our MEMS-based
approach, featuring an integrated stiction valve, reduces thesensor footprint drastically (250–500 μm), utilizes only twoelectrodes, and is batch fabricated on a flexible substrate. The
reduction in chamber size enables precisely patterned structures for single-point pixelated contact sensing over a definedFig. 14. Impedance response at various locations along the length of the channel (mean ± SE, n ≥ 3). Inset shows the schematic representation of channel
layout with peripheral stiction valves and corresponding optical micrographwith measurement and indentation sites indicated.

and controlled surface. Furthermore, the problem is reduced
to the measurement of a single resistive signal correlated ina straightforward fashion to the deflection of the membrane.
Although application dependent, chamber sizes on this scale
can eliminate the problem of multisite contact discrimination
by simply being smaller than the contacting body.
The basic requirements of an electrolyte in contact withelectrodes enable applications at various size scales. EI-based
sensing at submicrometer dimensions appears feasible provided
that fluid handling at this scale can be realized. This could
lead to applications, for example, in subcellular mechanicalcharacterization. Unique geometric sensor configurations arealso possible, extending beyond simple diaphragms to interconnected microfluidic structures.

VII. A DDITIONAL A PPLICATIONS
A. Mechanotransduction Along Channels
Additional applications of EC-MEMS can be realized byleveraging the unique liquid-based transduction mechanism.
Fluidically coupled structures of varying geometries can be
fabricated, such as interconnected channels, to form uniquesensor configurations not possible with traditional transduction
approaches. A long microfluidic channel (150 μm × 2.5 mm ×
10 μm) was fabricated using the same fabrication processdescribed previously. Stiction valves were placed peripherally
along the edges of the channel for liquid encapsulation purposes. A single reference electrode spanned the entire lengthof the channel, while additional electrodes were placed at junctions in close proximity to the reference at specific locations
within the channel. Each electrode junction acted as a single
sensor element.
A mechanical stimulus (maximum deflection) was applied
at various locations along the length of the channel, while
the impedance response was monitored at a single junction.
Deflections were detectable up to 1.2 mm away from themeasurement site (Fig. 14). The amplitude of the measured impedance change decreased with distance. The ability to detect

GUTIERREZ AND MENG: PARYLENE-BASED ELECTROCHEMICAL-MEMS TRANSDUCERS

1359

TABLE II
M AXIMUM ACTUATED D EFLECTION FOR S QUARE AND C IRCULARPARYLENE C HAMBERS (h = 10 μm AND t = 2 μm) (n ≥ 5)

Fig. 15. Same transducer structure with an additional surface microelectrodepatterned on the top exposed surface of the chamber membrane.

tactile information along the length of a channel extends thephysical sensing range from that of a single point using a single
chamber-based sensor to a linear range within the fluidicallycoupled structure. Multiple electrodes can populate a channel
segment, enabling further downstream tactile detection. The
mechanical transduction phenomenon along the length of thesingle channel could be considered biomimetic and analogous
to that of neuronal signal propagation along the length of anaxon. This kind of biomimetic sensing modality opens thepossibilities for new bio-inspired mechanotransduction devices
and applications. For example, these devices could be arranged
in concentric circular patterns and applied to grasp and tactile
feedback systems for minimally invasive robotic surgery, providing the ability to dynamically adjust tactile receptive fields
and resolution. Alternatively, the addition of surface features
such as bumps or “whiskers” could enable detection of physicalparameters such as flow detection and roughness.

B. Electrolysis-Based Microactuation
In addition to sensing capabilities, EC-MEMS devices arecapable of providing actuation through electrolysis of water.
For demonstrative purposes, a surface microelectrode was integrated on the surface of the chamber structure. The inclusion ofan electrode on the surface of the chamber enables a new device
application: out-of-plane electrode actuation (Fig. 15). Thisconfiguration could be applied to neural stimulation or recording where out-of-plane microactuation of electrodes could
provide improved electromechanical coupling through minuteinterfacial mechanical adjustments. Fabrication consists of anadditional dual-layer liftoff process with a globally exposed
bottom layer. This was essential to achieve continuity between
metal on the substrate level and on top of the chamber. FEM
of chamber deformation was conducted to establish areas ofleast stress for purposes of metal trace routing (data not shown).
Minimum stress occurred in the connecting region between thechamber and valve. Routing metal traces through this region
minimized tethering effects on the chamber by the metal during
deformation and also minimized strain on the fragile traces.
Platinum metallization and patterning were carried out as described previously.
The application of a constant current (∼ 5 μA for a few
seconds) across the internal chamber electrodes initiated elec-

trolysis and the evolution of hydrogen and oxygen gas within
the chamber. The accumulation of this gas generated high internal pressure, causing the chamber membrane and electrode to
deflect upward. Maximum deflection measurements of squareand circular devices are summarized in Table II. The results
indicate that larger maximum deflections were achievable aschamber dimensions increased due to increased membranecompliance with size. In general, circular chambers exhibitedlarger more uniform deflections than the square chambers.
Electrolytic gas generation was concentrated in the center ofthe circular chambers, while gas generation was widespreadand erratic in the square chambers. This difference is likely
due to the interdigitated electrode geometry of the squarechambers as well as the propensity for gases to accumulate near
the edges and corners of the square chamber. The maximum
achievable actuation of this device configuration was limitedby the cracking pressure of the in-line stiction valve. Above
this pressure threshold, the valve plate popped open, thereby
limiting the amount of pressure that can be generated for actuation purposes. This pressure was estimated analytically to beabout 6895 Pa (∼1 psi) by equating the pressure for maximum
measured center deflection with valve cracking pressure. This
limitation could potentially be resolved by using a multilayered
configuration, thereby integrating the valve inside the chamber in such a manner that the generation of internal pressure
serves to push down the valve plate, further sealing it. This
would allow the buildup of high pressures and enable greater
membrane deflection. Pressure generation of several hundred
psi is possible by electrolysis but unlikely to be sustained within
the transducer structure due to the known weak interfacialadhesion of the Parylene–Parylene interface. Strategies such as
annealing, anchoring, or molten Parylene adhesion [40] may
be employed to improve adhesion performance; however, asystematic comparison of the improved interfacial adhesion
afforded by these techniques is still required.

VIII. C ONCLUSION
EC-MEMS devices utilizing encapsulated fluid to mediatetransduction and Parylene as the structural material have been
designed, fabricated, and tested. Parylene microstructures, including square/circular chambers and long channels, have been
evaluated and validated as highly sensitive physical transducers
with applications to contact sensing, biomimetic mechanotransduction, and out-of-plane microelectrode actuation.

1360

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 19, NO. 6, DECEMBER 2010

These transducers are ideally suited for applications requiring
highly sensitive contact force measurements of soft nonplanar
surfaces in wet environments. Integration of these sensors
with state-of-the-art Parylene-based retinal prosthesis willallow interrogation of important physical phenomena,
in situ, for the first time. Validation experiments utilizing exvivo soft tissues are currently under way.

ACKNOWLEDGMENT
The authors would like to thank B. Kim, M. Pahwa,
C. McCarty, and M. Nguyen for their invaluable assistancewith device characterization. The authors would also like tothank Dr. D. Zhu and the members of the Biomedical Microsystems Laboratory, University of Southern California, for theirassistance.
R EFERENCES[1] M. A. Schwarz and P. C. Hauser, “Recent developments in detection
methods for microfabricated analytical devices,” Lab Chip, vol. 1, no. 1,
pp. 1–6, Sep. 2001.
[2] D. J. Beebe, G. A. Mensing, and G. M. Walker, “Physics and applicationsof microfluidics in biology,” Annu. Rev. Biomed. Eng., vol. 4, no. 1,
pp. 261–286, 2002.
[3] R. P. Baldwin, T. J. Roussel, M. M. Crain, V. Bathlagunda, D. J. Jackson,
J. Gullapalli, J. A. Conklin, R. Pai, J. F. Naber, K. M. Walsh, and
R. S. Keynton, “Fully integrated on-chip electrochemical detection forcapillary electrophoresis in a microfabricated device,” Anal. Chem.,
vol. 74, no. 15, pp. 3690–3697, Aug. 2002.
[4] W. R. Vandaveer, IV, S. A. Pasas-Farmer, D. J. Fischer, C. N. Frankenfeld,
and S. M. Lunte, “Recent developments in electrochemical detection for
microchip capillary electrophoresis,” Electrophoresis, vol. 25, no. 21/22,
pp. 3528–3549, Nov. 2004.
[5] D. Erickson and D. Li, “Integrated microfluidic devices,” Anal. Chim.
Acta, vol. 507, no. 1, pp. 11–26, Apr. 2004.
[6] S. Z. Hua, F. Sachs, D. X. Yang, and H. D. Chopra, “Microfluidic actuation
using electrochemically generated bubbles,” Anal. Chem., vol. 74, no. 24,
pp. 6392–6396, Dec. 2002.
[7] A. P. Papavasiliou, D. Liepmann, and A. P. Pisano, “Electrolysis-bubble
actuated gate valve,” in Proc. Solid-State Sensor Actuator Workshop,
Hilton Head, SC, 2000, pp. 48–51.
[8] W. H. Coulter, “Means for counting particles suspended in a fluid,” U.S.
Patent 2 656 508, Oct. 20, 1953.
[9] N. Wettels, V. J. Santos, R. S. Johansson, and G. E. Loeb, “Biomimetic
tactile sensor array,” Adv. Robot., vol. 22, no. 8, pp. 829–849, 2008.
[10] A. R. Russel, “A tactile sensory skin for measuring surface contours,”
in Proc. IEEE Region 10 Int. Conf.—Technology Enabling Tomorrow:
Computers, Communications and Automation Towards the 21st Century,
Melbourne, Australia, 1992, pp. 262–266.
[11] A. R. Russel and S. Parkinson, “Sensing surface shape by touch,” in Proc.
IEEE Int. Conf. Robot. Autom., Atlanta, GA, 1993, pp. 423–428.
[12] D. A. Ateya, A. A. Shah, and S. Z. Hua, “Impedance-based responseof an electrolytic gas bubble to pressure in microfluidic channels,” Sens.
Actuators A, Phys., vol. 122, no. 2, pp. 235–241, Aug. 2005.
[13] J. D. Weiland, W. Liu, and M. S. Humayun, “Retinal prosthesis,” Annu.
Rev. Biomed. Eng., vol. 7, pp. 361–401, 2005.
[14] D. C. Rodger, A. J. Fong, W. Li, H. Ameri, A. K. Ahuja, C. Gutierrez,
I. Lavrov, H. Zhong, P. R. Menon, E. Meng, J. W. Burdick, R. R. Roy,V. R. Edgerton, J. D. Weiland, M. S. Humayun, and Y.-C. Tai, “Flexible
parylene-based multielectrode array technology for high-density neuralstimulation and recording,” Sens. Actuators B, Chem., vol. 132, no. 2,
pp. 449–460, Jun. 2008.
[15] L. Colodetti, J. D. Weiland, S. Colodetti, A. Ray, M. J. Seiler,
D. R. Hinton, and M. S. Humayun, “Pathology of damaging electricalstimulation in the retina,” Exp. Eye Res., vol. 85, no. 1, pp. 23–33,
Jul. 2007.
[16] D. Reigada, W. Lu, M. Zhang, and C. H. Mitchell, “Elevated pressure
triggers a physiological release of ATP from the retina: Possible rolefor pannexin hemichannels,” Neuroscience, vol. 157, no. 2, pp. 396–404,
Nov. 2008.
[17] A. B. Majji, M. S. Humayun, J. D. Weiland, S. Suzuki, S. A. D’Anna,
and E. de Juan, Jr., “Long-term histological and electrophysiological results of an inactive epiretinal electrode array implantation indogs,” Invest. Ophthalmol. Vis. Sci., vol. 40, no. 9, pp. 2073–2081,
1999.
[18] B. C. Basinger, A. P. Rowley, K. Chen, M. S. Humayun, and J. D. Weiland,“Finite element modeling of retinal prosthesis mechanics,” J. Neural Eng.,
vol. 6, no. 5, p. 055 006, Oct. 2009.
[19] Y.-B. Lu, K. Franze, G. Seifert, C. Steinha, F. Kirchhoff, H. Wolburg,
J. Guck, P. Janmey, E.-Q. Wei, J. Kas, and A. Reichenbach, “Viscoelastic properties of individual glial cells and neurons in the CNS,” PNAS,
vol. 103, no. 47, pp. 17 759–17 764, Nov. 2006.
[20] H. A. Quigley, “Neuronal death in glaucoma,” Prog. Retin. Eye Res.,
vol. 18, no. 1, pp. 39–57, Jan. 1999.
[21] I. Suzuma, T. Murakami, K. Suzuma, H. Kaneto, D. Watanabe, T. Ojima,Y. Honda, H. Takagi, and N. Yoshimura, “Cyclic stretch-induced reactive oxygen species generation enhances apoptosis in retinal pericytes
through c-Jun NH2 -terminal kinase activation,” Hypertension, vol. 49,
no. 2, pp. 347–354, Feb. 2007.
[22] C. A. Gutierrez and E. Meng, “A dual function parylene-based biomimetic tactile sensor and actuator for next generation mechanically responsive microelectrode arrays,” in Proc. Transducers, Denver, CO, 2009,
pp. 2194–2197.
[23] H. Zhang, S. Wang, and Y. Xu, “Study and applications of a parylene
self-sealing structure,” in Proc. MEMS, 2006, pp. 282–285.
[24] L. S. Robblee and T. L. Rose, “The electrochemistry of electrical stimulation,” in Proc. Annu. Int. Conf. IEEE Eng. Med. Biol. Soc., Philadelphia,
PA, 1990, pp. 1479–1480.
[25] J. E. B. Randles, “Kinetics of rapid electrode reactions,” in Proc. Discuss.
Faraday Soc., 1947, pp. 11–19.
[26] L. S. Robblee and T. L. Rose, “Electrochemical guidelines for selectionof protocols and electrode materials for neural stimulation,” in Neural
Prostheses Fundamental Studies, W. F. Agnew and D. B. Mcreery, Eds.
Englewood Cliffs, NJ: Prentice-Hall, 1990, pp. 26–59.
[27] L. Wolgemuth, “Assessing the performance and suitability of parylene coating,” Med. Device Diagn. Ind., vol. 22, no. 8, pp. 42–49,
2000.
[28] Cookson Specialty Coating Systems. [Online]. Available: http://www.
scscoatings.com/parylene_knowledge/specifications[29] C. A. Gutierrez and E. Meng, “Improved self-sealing liquid encapsulation in parylene structures by integrated stackable annular-plate stiction
valve,” in Proc. IEEE MEMS, Hong Kong, 2010, pp. 524–527.
[30] C. H. Mastrangelo and C. H. Hsu, “Mechanical stability and adhesion of
microstructures under capillary forces. I. Basic theory,” J. Microelectromech. Syst., vol. 2, no. 1, pp. 33–43, Mar. 1993.
[31] C. H. Mastrangelo and C. H. Hsu, “Mechanical stability and adhesion of
microstructures under capillary forces. II. Experiments,” J. Microelectromech. Syst., vol. 2, no. 1, pp. 44–55, Mar. 1993.
[32] Z. Wang and Y. Xu, “Theoretical and experimental study of annularplate self-sealing structures,” J. Microelectromech. Syst., vol. 17, no. 1,
pp. 185–192, Feb. 2008.
[33] S. Matsumoto and N. Ichikawa, “New methods for liquid encapsulationin polymer MEMS structures,” in Proc. MEMS, Tucson, AZ, 2008,
pp. 415–418.
[34] A. Tanioka, N. Fukushima, K. Hasegawa, K. Miyasaka, and N. Takahashi,“Permeation of gases across the poly(chloro-p-xylylene) membrane,” J.
Appl. Polym. Sci., vol. 54, no. 2, pp. 219–229, Oct. 1994.
[35] P. R. Menon, W. Li, A. Tooker, and Y. C. Tai, “Characterization of water
vapor permeation through thin film parylene C,” in Proc. Transducers,
Denver, CO, 2009, pp. 1892–1895.
[36] M. D. Giovanni, Flat and Corrugated Diaphragm Design Handbook,vol. 11. New York: Marcel Dekker, 1982.
[37] K. Kasten, N. Kordas, H. Kappert, and W. Mokwa, “Capacitive pressure
sensor with monolithically integrated CMOS readout circuit for hightemperature applications,” Sens. Actuators A, Phys., vol. 97/98, pp. 83–
87, Apr. 2002.
[38] P. Hille, R. Höhler, and H. Strack, “A linearisation and compensation
method for integrated sensors,” Sens. Actuators A, Phys., vol. 44, no. 2,
pp. 95–102, Aug. 1994.
[39] M. Helsel, J. N. Zemel, and V. Dominko, “An impedance tomographic tactile sensor,” Sens. Actuators, vol. 14, no. 1, pp. 93–98,
1988.
[40] R. Huang and Y. C. Tai, “Parylene to silicon adhesion enhancement,” inProc. Transducers, Denver, CO, 2009, pp. 1027–1030.

GUTIERREZ AND MENG: PARYLENE-BASED ELECTROCHEMICAL-MEMS TRANSDUCERS

Christian A. Gutierrez (M’09) received B.S. degrees in electrical engineering and business economics and management and the M.S. degree inelectrical engineering from the California Instituteof Technology (Caltech), Pasadena, in 2005 and2006, respectively. He is currently working toward
the Ph.D. degree in biomedical engineering (under
fellowship) at the University of Southern California(USC), Los Angeles.
He is also currently with the National Science
Foundation’s Biomimetic Microelectronic Systems
Engineering Research Center, USC, where he is a Copresident of the student leadership council. His past research was carried out in the Caltech
Micromachining Laboratory and was focused on identifying, analyzing, andimplementing power harvesting and storage circuitry for microelectromechanical-systems-based power harvesting devices. He is currently conducting research toward the development of novel micro- and nanotechnologies
for biomedical applications. His specific research is focused on the developmentof retinal prosthesis technologies to restore sight to the blind.
Mr. Gutierrez was a recipient of the Caltech Presidential Scholarship and iscurrently a Bill Gates Millennium Scholar. He was also a recipient of the USC
Viterbi School of Engineering Kunzel Fellowship.

1361

Ellis Meng (M’02–SM’09) received the B.S. degree
in engineering and applied science and the M.S.
and Ph.D. degrees in electrical engineering from the
California Institute of Technology (Caltech),
Pasadena, in 1997, 1998, and 2003, respectively.
Since 2004, she has been with the Departmentof Biomedical Engineering, University of Southern
California (USC), Los Angeles, where she is currently an Associate Professor. She currently holds a
joint appointment with the Ming Hsieh Department
of Electrical Engineering, USC. She is currently theThrust Leader for Interface Technology and the Associate Director of Educationand Student Diversity Researcher at the National Science Foundation (NSF)’s
Biomimetic Microelectronic Systems Engineering Research Center, USC. She
held the Viterbi Early Career Chair in the USC Viterbi School of Engineering. Her research interests include biological microelectromechanical systems,
implantable biomedical microdevices, microfluidics, multimodality integrated
microsystems, and packaging.
Dr. Meng is a member of Tau Beta Pi, the Biomedical Engineering Society,
the Society of Women Engineers, and the American Society for Engineering
Education. She was a recipient of the Intel Women in Science and Engineering
Scholarship, the Caltech Alumni Association Donald S. Clark Award, andthe Caltech Special Institute Fellowship. She has also received the NSF
CAREER and Wallace H. Coulter Foundation Early Career Translational
Research Awards.