1098

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 20, NO. 5, OCTOBER 2011

Impedance-Based Force Transduction Within
Fluid-Filled Parylene Microstructures
Christian A. Gutierrez, Member, IEEE, and Ellis Meng, Senior Member, IEEE

Abstract—We report on the use of electrochemical impedance(EI) as the basis for force transduction in Parylene-based microdevices. Electrolyte-filled microstructures were realized for extremely sensitive contact-force detection (10 mN range, ±0.023 mN
resolution) enabled by EI-based transduction and are a promising platform for next-generation biomedical sensing technology.
The design, fabrication, and characterization of Parylene-based
electrochemical-MEMS (EC-MEMS) devices capable of microNewton contact-mode force measurement are presented anddiscussed.
[2011-0040]
Index Terms—Biomimetic, contact sensor, electrochemicalMEMS (EC-MEMS), electrochemical sensing, electrochemicaltransduction, force sensor, impedance-based sensor, liquid encapsulation, liquid impedance, mechanotransduction, Parylene C.

Fig. 1. Model of single sensor unit. A compliant electrolyte-filled Parylene
microchamber enables EI-based force measurement. The sensor is fluidicallycoupled and open to the external environment through the fluidic access ports
and channels.

I. I NTRODUCTION

is a common approach; however, these technologies often lackthe force sensitivity required for the most sensitive applications(0.01–1 mN), and are geared toward general surgical grippingand grasping requiring force sensitivity in the range of 0.1–
11 N [25]. Traditional silicon-based techniques (cantilevers,
diaphragms, combs) often require additional encapsulation techniques to protect the sensors from wet environments, and therequisite packaging can often prohibit practical implementation[26]. Furthermore, protective elastic coatings reduce sensitivity
and function as low-pass spatial and temporal filters [27].
To address these challenges, we previously demonstrated theuse of Parylene as a substrate to fabricate flexible, electrolytefilled microstructures for sensing and actuating applications[28], [29]. The use of Parylene brings improved mechanicalcoupling over traditional materials for applications involving
intimate contact with biological tissues. The combination of
microfluidics, electrochemistry, and mechanics form the basisof electrochemical-MEMS (EC-MEMS) technologies that
present a promising platform for next-generation biomedical
sensors. Here, we report on the development of a Parylenebased EC-MEMS sensor capable of microNewton contactmode force transduction.

P

ARYLENE has gained popularity as a polymer material
across various fields of research due to its versatility, conformal room temperature deposition, favorable mechanical andelectrical properties, and biocompatibility (USP Class VI). The
ability to micromachine Parylene has led to its use in sensors[1]–[4], actuators [5], [6], and various microfluidic components
and systems [7]–[10]. This combination of attributes makes
Parylene attractive for many biomedical applications. In thispaper, we demonstrate the use of fluid-filled Parylene microstructures for highly sensitive force measurement intended
for biomedical applications.
A variety of biomedical force sensing technologies exist[11]–[13]; however, oftentimes the underlying operating mechanisms (capacitive, electrostatic, piezoresistive, optical), developed for nonbiological applications, remain the same [14], [15].
As a result, these techniques can be difficult to implementin practical applications where they could be most beneficial.
Biomedical force measurement is of particular interest to areas
such as tactile feedback for minimally invasive surgery [12], tissue health diagnostics [16]–[18], surgical tool instrumentation[19], and real-time monitoring of acute applied pressures during
delicate surgical procedures in the eye [20]–[22] or brain [23],[24] (e.g., retinal or cortical implant procedures). The use ofelastomeric materials functionalized with conductive particles
Manuscript received February 14, 2011; revised May 27, 2011; accepted
June 19, 2011. Date of publication July 29, 2011; date of current version
September 30, 2011. This work was supported in part by the Biomimetic
Microelectronic Systems Engineering Research Center of the NSF, an NSF
CAREER Award, and by the Bill and Melinda Gates Foundation (CG). Subject
Editor C. Liu.
The authors are with the University of Southern California, Los Angeles,
CA 90089 USA (e-mail: cagutier@usc.edu; 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.2011.2160935

II. D EVICE D ESIGN AND O PERATIONThe core of the device consists of a circular microchamber
structure filled with a conductive aqueous electrolyte and apair of microfabricated thin-film electrodes in contact with theelectrolyte solution (Fig. 1). Additional microfluidic channels at
the base of the chamber provide a pathway for fluidic coupling
between the microchamber interior and the external environment. A pair of electrodes is located opposite one another on
the interior perimeter of the chamber to maximize the path
length over which electrochemical impedance (EI) measurements are taken. A circular diaphragm design was selected here

1057-7157/$26.00 © 2011 IEEEGUTIERREZ AND MENG: IMPEDANCE-BASED FORCE TRANSDUCTION WITHIN FLUID-FILLED PARYLENE MICROSTRUCTURES

Fig. 2. (a) Electrochemical impedance transduction principle. (b) The electrical equivalent circuit model of the electrode–electrolyte interface.

for simplicity, however, additional sensor geometries are easily
realized given that the impedance sensing technique can beapplied between any electrode-pair arrangement provided thata fluid path exists. The use of Parylene C and thin-film platinum
in the device construction enables the fabrication of flexible
films with excellent biocompatibility. The sensors are ready touse after fabrication and, except for electrical connections, noadditional materials or packaging processes are required at thesensor’s mechanical interface.
The device is fluidically coupled to the external environment
and therefore requires immersion in an aqueous medium for
operation. This requirement, is however, compatible with thewet conditions usually encountered when working with soft
tissues. Filling is accomplished by immersion in the operating
environment, allowing the surrounding fluid to fill the chamberand function as the working electrolyte. The application of
external forces on the chamber surface deforms the compliant fluid-filled structure and redistributes the enclosed fluid[Fig. 2(a)]. Variations in the volumetric conductive path ofcurrent-carrying ions in the fluid register as a change in themeasured EI. In this manner, volumetric variations of an encapsulated electrolyte can be correlated to mechanical contact
forces exerted by external sources on the sensor. The resulting
impedance changes contain information about the magnitude ofthe externally applied force.
A. Electrochemical Considerations
The electrical equivalent circuit of the two-electrode system,
depicted in Fig. 2(b), is based on the simplified Randles model
where the parallel combination of the double layer capacitance(Cdl ) and charge transfer resistance (Rct ) describes phenomenon at each electrode interface. Many electrochemical sensing
applications employ potentiometric and amperometric detection techniques where changes to the electrode surface properties (due to binding and adsorption of species in the solution)
are used as the basis for detection. Here, however, we apply aconductometric technique where characteristics of the solution

1099(namely solution resistance) are leveraged for transduction, andinterfacial phenomena occurring at the electrode surface areless important. Details regarding the electrochemical interface
and associated surface phenomena have been well studied
and can be found throughout the electrochemistry literature[30]–[34]; only a brief treatment relevant to the application
presented here is provided.
EI is measured through application of a small alternating current (or voltage) across the thin-film electrodes and tracking thecorresponding voltage (or current). These measurements can be
performed with very little power (< 100 μW) and produce only
reversible electrochemical species. Consequently, hydrolysis of
the electrolyte does not occur so long as the magnitude of theapplied AC signal is maintained within the “water-window” ofthe electrode material (∼1.4 Vpp for platinum) where current isgenerated though reversible faradaic reactions. At sufficiently
high frequencies, typically > 1 kHz, the contribution of theinterfacial components (Cdl and Rct ) to the overall impedanceis minimized as the capacitive reactance of Cdl becomes small(1). A limitation to this assumption often arises in the useof microelectrodes where electrode surface areas can be verysmall, thus increasing the reactance of Cdl appreciably. This is
typically a concern for sharp point-tip electrodes where noise
and attenuation can significantly impact measurement of small
signals

Z(w) = Rs +2Rct.
1 + jwRct Cdl(1)

The value of Cdl per unit area for bright platinum at 1 kHz
has been measured to be about 20 pF/cm2 [35]. The exposedarea of each platinum electrode used here is 55 × 25 μm(1375 μm2 ) yielding a capacitance of 275 pF at 1 kHz. The
values of both Rct and Cdl are in fact not constant as suggestedby the model, but have been empirically found to vary withfrequency scaling as 1/f 0.5 [36]. Thus, by appropriate scaling,
the capacitance value becomes 123 pF at 5 kHz and yields a
capacitive reactance of 260 kΩ which is in line with values
reported in the literature. Electrode impedances up to 5 MΩare not uncommon and have been utilized extensively in singleunit neuronal signal recording. For instances in which the electrode impedance becomes significantly larger, for example withvery small tip electrodes, platinization of the electrode surface
through electrodeposition has proved effective in dramaticallyincreasing surface area. An effective increase of 30–40 times
over smooth platinum has been reported [20], [37], [38]. The
charge transfer resistance, a parameter often difficult to measure experimentally, captures the resistance to changes in theequilibrium current within the double layer when an externalpotential is applied. It is generally considered to be the difference between the electrode impedance at very low frequency(< 0.1 Hz) and the solution resistance and is typically very high
(10–100 MΩ). This assumption was verified by a parametric fitof the EI spectrum of several thin-film electrodes using standard electrochemical analysis software (Gamry Instruments,
Echem Analyst) yielding Rct values in excess of 30 MΩ.
At the frequency utilized here (5 kHz), Rct does not make

1100

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 20, NO. 5, OCTOBER 2011

Fig. 3. Multilayered Parylene-based fabrication process performed at low temperature (90 ◦ C).

an important contribution to the overall electrode impedance(relative to the capacitive reactance of Cdl ).
It should also be noted that the exact values of Cdl and
Rct are of secondary importance given the use of a single
measurement frequency at which these parameters can beconsidered constant. Conductometric variations (eg. solution
resistance) are therefore the main source of changes to themeasured impedance. The use of a low conductivity electrolyte,such as deionized (DI) water (< 10 μS/cm), further ensuresthat the solution resistance is the dominant parameter in themeasured impedance value. Consequently, the impedance response can be approximately modeled as a simple classicalresistance RS , where ρ is the resistivity of the electrolyte, l isthe distance between the electrodes, and A is the cross-sectional
area of the fluid between the electrodes (2). The impedanceis, therefore, inversely proportional to the cross-sectional area
between the electrodes and is directly analogous to a variable“fluidic” resistor. This phenomenon enables the transduction
and measurement of physical stimuli acting on the electrolytefilled chamber
Δ|Z|f >1 khz ≈ ΔRS =

ρl
.
ΔA

(2)

The length scale over which such an approximation is valid isdetermined by the characteristic Debye length which accounts
for surface interactions occurring at the electrode surface. These
include adsorbed ions on the metal surface which form theHelmholtz double layer and nonspecifically adsorbed ions in
the adjacent diffuse layer. This length is also dependent on theconcentration of solvated ions and decreases with increasingionic strength. For a typical salt solution (eg. saline), thisdistance is < 10 nm but can range up to ∼950 nm in the caseof pure water [34], [39]. Beyond this region exists the bulk
solution (i.e., homogenous charge distribution) and extends
until a second electrode is encountered. For a two-electrodesystem in which the bulk solution resistance serves as the coretransduction mechanism, the minimum electrode separation
distance over which (2) is valid is estimated to be on the order∼2 μm for low conductivity electrolytes and ∼20 nm for a highconductivity electrolyte. An electrode separation of ∼175 μm,
utilized here, provides sufficient bulk solution to satisfy thisassumption under a variety of working electrolytes.

B. Device Mechanics
Due to the circular chamber geometry, the force–deflection
response most closely follows that of a pressure-loaded circular diaphragm with fixed edges. Although these boundary
conditions are not strictly met due to slight deformation of thechamber sidewalls, the analytical results of this model sufficiently describe the sensor response. By adjusting parameterssuch as membrane thickness t, diaphragm radius a, and heighth, the force–deflection performance can be easily tuned for aparticular range (3) [40]:
P =

16
Et3
7 − u Et 3
y0 +
y
2
4
3(1 − u ) a3(1 − u) a4 0

(3)

where y0 is center deflection, u is Poisson’s ratio, E is modulus
of elasticity, and P is pressure (assumed to be uniformly applied force distributed over surface area). Adjusting the sensor
height will set the maximum deflection range of the sensors
and impose an upper limit to the applied forces before themembrane touches down onto the substrate below. As will be
discussed later, the effect of the integrated fluidic channels andports only becomes significant if they are occluded or blocked.
The most significant factor to the mechanical behavior is theviscous damping imposed by the surrounding solution which
limits operation to low frequency (< 200 Hz).
III. D EVICE FABRICATIONThe entire fabrication process was realized at relatively low
temperatures (90 ◦ C) which enables its use on a variety ofpolymer substrates and facilitates adoption in temperaturesensitive processes (Fig. 3). Standard surface micromachining
techniques were employed on a standard soda lime substrate(500 μm thick). First, Parylene was deposited (10 μm) on awafer optionally treated with A-174 adhesion promoter (Specialty Coating Systems, Inc., Indianapolis, ID). The adhesion
promoter facilitated handling of test devices which were diced
into 1 cm2 dies for characterization purposes and ensured
adhesion to the substrate. Platinum electrodes (2000 Å) were
deposited in an electron beam evaporator and patterned bylift-off using an image reversal photoresist with a negative
sidewall profile (1.5 μm, AZ 5214-E, AZ Electronic Materials,
Branchburg, NJ). Next, a Parylene insulation layer (1 μm)

GUTIERREZ AND MENG: IMPEDANCE-BASED FORCE TRANSDUCTION WITHIN FLUID-FILLED PARYLENE MICROSTRUCTURES

Fig. 4.

1101(a) Optical micrograph of single device [41] and (b) scanning electron micrograph of device structure ([10] © IEEE 2010).

was deposited. Insulation over the electrodes was removed
by reactive ion etching in oxygen plasma (100 W, 100 mT)
utilizing a photoresist mask (4 μm, AZ 4400, AZ Electronic
Materials, Branchburg, NJ). Photoresist (2–4 μm, AZ 4400,
AZ Electronic Materials, Branchburg, NJ) was then spun-on
and patterned forming a sacrificial mold for the fluidic accessports. This was followed by Parylene deposition (2 μm) andselective etching of Parylene over the fluidic port openings andelectrode sites.
Removal of the photoresist etch mask proved to be difficultat this point due to the exposed fluidic port openings andresulted in erosion of the underlying sacrificial structures.
Directly spinning an additional layer of photoresist on topof the existing etch mask at a calibrated speed enabled
successful patterning of a second sacrificial photoresist layer(AZ 4620, AZ Electronic Materials, Branchburg, NJ) to form
the microchamber mold (18–21 μm height, 200 μm diameter).
A thorough softbake of this layer was essential to avoid solvent
degassing and to minimize rounding of chamber featuresduring the following steps. A final layer of Parylene (4.2 μm)
was then deposited comprising the membrane and sidewallsof the chamber structure. The Parylene was etched using a
10–12 μm thick etch mask (AZ 4620, AZ Electronic Materials,
Branchburg, NJ) to re-open the fluidic access ports on theexterior of the chamber exposing the underlying photoresist
allowing for access and removal by immersion in baths ofacetone and isopropyl alcohol. Due to the nonplanar structuresduring this final etch step, a thicker mask was necessary to
prevent mask erosion and unwanted etching of the top edgesof the chamber. Following a 24 h soak in the solvent baths, thedevices were transferred to a DI water bath for rinsing (∼6 h)
and finally placed in a bath containing the desired electrolyte
solution. Removal of sacrificial materials, rinsing, and filling
occurred simultaneously across multiple devices in a waferlevel batch process. The Parylene film was optionally released
from an untreated soda lime substrate by cutting with a razor
blade, a process which was facilitated by immersion in water.
Fig. 4(a) is an optical micrograph of a device following
removal of the sacrificial photoresist and filling with DI water.
Fig. 4(b) is a scanning electron micrograph of a device prior
to removal of any sacrificial materials showing the chamber
protruding from the surface of the Parylene film. A thin layer ofgold has been evaporated over the samples for scanning electron
microscopy to avoid excess charging of the Parylene film.

IV. E XPERIMENTAL M ETHODS
Test dies containing several devices were diced from a soda
lime wafer substrate and mounted on a custom printed circuitboard (PCB) with a cyanomethacrylate adhesive. Electricalconnectivity to PCB-mounted dies was established by wirebonding prior to removal of the sacrificial photoresist contained
within the devices. A water-proof marine epoxy was thenapplied underwater and served two purposes: 1) insulationof the wirebond connections and 2) formation of a well tocontain electrolyte such that devices were submerged during
testing. Insulating the wirebonds was critical to avoid shortingof leads when the well was filled with electrolyte. This step
was performed underwater to prevent devices from drying andcollapsing due to stiction when exposed to air. Outside thewell, electrical connections were soldered to contact pads andwere kept dry during testing. The PCB-mounted devices weremaintained in DI water for 24 h to allow full curing of theepoxy well. Although the ionic content of DI water is low, it
is still capable of functioning as an electrolyte with a solution
resistance typically in the 0.5–2 MΩ range. Various othersolutions could function as an electrolyte, for example, humanblood, body fluids, and cell culture media are all compatible
with EI measurements [42].
Experiments were conducted on a floating optical table andcontrolled via a custom LabVIEW interface (National Instruments, Austin, TX) providing both control and measurement
capability. PCBs were mounted onto an X-Y positioning stage
secured to the optical table. Force measurements were acquiredwith a 50 g load cell (LCFA-50, Omega Engineering, Stamford, CT) mounted to a computer-controlled micropositioning
stage providing accurate z-axis positioning capability (Z812,
Thorlabs, Newton, NJ) (Fig. 5). This enabled real-time contactmode force measurement during application of user-defined
deflection profiles. EI measurements were acquired (5 kHz,
1 Vpp ) concurrently by a precision LCR meter (E4980A, Agilent Inc, Santa Clara, CA) and logged by LabVIEW software.
A flat, polished probe tip measuring 500 μm in diameter
was attached vertically in line with the load cell. The probe
tip provided a defined contact surface large enough to deflect
the entire sensor surface Fig. 6(a). Alignment of the tip was
performed visually through an optical microscope attached to agoose-neck CCD capture device for flexibility. Contact between
the probe tip and sensor was verified by concurrent increases
in the measured force and impedance prior to application of

1102

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 20, NO. 5, OCTOBER 2011

Fig. 5. Experimental setup. A load cell with attached probe tip was mounted to a motorized micropositioning stage with sub micrometer Z-axis positioning
capability. Deflection experiments were performed with sensors completely submerged.

A. Finite-Element ModelingFig. 6. (a) Flat polished probe tip mounted to load cell used for uniform
deflection across entire sensor footprint. (b) High resolution force probe (3 mmlong, 50 × 50 μm) with microNewton resolution (Femtotools). Smaller probe
provides high resolution force measurements not possible with the standard
load cell and approximates deflection characteristics under a central point force.

user-defined deflection profiles. For high resolution force measurements, a silicon force probe (50 × 50 μm tip) with microNewton force resolution was utilized (ST-150, Femtotools
HmGb, Zurich, Switzerland). The smaller probe head allowed
for precise force measurements applied centrally on the sensor
surface Fig. 6(b). Assuming that a centrally applied point force
F can be translated to an evenly distributed pressure P over amembrane of radius r through the relation F = P πr2 , it canbe shown that a centrally point-force-loaded diaphragm will
deflect four times that of uniformly loaded diaphragm [40]. This
relation provides a useful and approximate basis for comparison
between high resolution measurements and those performed
with the load cell. High resolution force measurements wereattained up to a maximum of 3000 μN (the probe overload
limit).
V. R ESULTS AND D ISCUSSION
The typical deflection profile consisted of a constant velocity(2 μm/s) downward deflection followed by release, with a dwell
time of approximately 3 s at the maximum deflection. This
profile was repeated twice for each depth between 6 and 18 μm
in increments of 2 μm [Fig. 7(a)]. During sensor deflection,
temporal responses of both force and impedance were recorded
at a sampling rate of 4 Hz. Over a period of several minutes,
these responses were robust and repeatable with deflection
steps clearly discernable through elevated impedance and force
measurements [Fig. 7(b)].

A nonlinear finite-element model (FEM) of the Parylene microchamber was developed and tested to model its
force–deflection performance (COSMOSWorks 2009, Dassault
Systèmes SolidWorks Corporation, Concord, MA). A nonlinear
model provided greater accuracy in capturing sensor dynamics
outside of the small deflection range. Typical values reported in
the literature for thin-film Parylene were utilized for this model[43]. Forces in the range of 10 mN (applied normally over thefull sensor contact surface) indicated deflections of 10–12 μm
or approximately 50% of device span and agree quite wellwith experimental results (Fig. 8). Compared to the analyticalsolution for circular diaphragms, the FEM simulation produced
larger deflections for a given force which is reasonable given
the fixed, rigidly clamped boundary conditions assumed by the
analytical solution.
Experimental data was attained using a load cell which
provided a large force measurement range but low precisionparticularly at small loads which may account for some ofthe discrepancy between the experimental data and the FEM
particularly in the range below 5 mN. Another source mayarise from the model assumption of a strictly normal application of force on the diaphragm throughout the deflection
range. This is somewhat different experimentally as slightdeviations from normal are expected as diaphragm curvature
is introduced. Yet, another possible source is the introductionof a slight convexity to the diaphragm which is not captured
in the FEM model. It is reasonable to expect some slightdeviation from an ideally flat diaphragm when utilizing verythin flexible films filled with solution. At low convexity ratios,in the region of small deflection (< 2 μm), the response ischaracterized by a very close (or slightly lower) deflection
than predicted followed by a region of greater deflection than
predicted by the models [40]. The net result would be a noticeably steeper gradient than predicted in the region below
5 mN. The response curve at large deflection would remain
largely unaffected. Nevertheless, these models are useful inproviding an estimate of the range of loads expected over thefull range of sensor deflection. Although the simulations andanalytical models do not capture any coupling between the fluid
and sensor microstructures, the measured responses still agree

GUTIERREZ AND MENG: IMPEDANCE-BASED FORCE TRANSDUCTION WITHIN FLUID-FILLED PARYLENE MICROSTRUCTURES

1103

Fig. 7. (a) Typical deflection profile. Deflections were performed twice at each depth and incremented by 2 μm after each cycle. (b) Typical force and impedance
response measured concurrently during deflections cycles. Load cell force response is shown with typical region of interest for high resolution measurements
indicated.

Fig. 8. Nonlinear finite-element simulation of the force–deflection loading response. Comparison of theoretical, simulated and measured force–deflection
responses (mean ± SE, n ≥ 10).

Fig. 9. (a) Normalized deflection as a function of radial distance from the center of the sensor. Shaded area indicates cross-sectional area calculated for one
deflection curve (membrane touch-down). The dependence of normalized cross-sectional area with maximum center deflection is shown in the inset. (b) Theoreticalcurve and representative data demonstrating the relationship between impedance and deflection.

quite well with both simulated and analytical solutions. Thissuggests that for the aqueous electrolytes, surface tension andviscosity have a minimal effect on the mechanical performanceof the sensor under static conditions. At higher loading frequencies, these factors will likely play a larger role in the sensor
dynamics but their effects under these conditions were notinvestigated here.

B. Force Response
The ability to accurately track deflections using EI ultimately
determines the resolution and sensitivity of this approach.
Membrane deflection as function of radial distance r is shownin Fig. 9(a) for several levels of increasing force. The dependence of cross-sectional area on central membrane deflection (y0 , r/a = 0,) is highlighted in the inset. With all other

1104

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 20, NO. 5, OCTOBER 2011

Fig. 10. Impedance–force responses as measured with load cell and large probe tip. (a) Impedance–force curves of a single sensor. Two deflection cycles are
performed at each deflection depth. (b) Average responses of five unique devices in 0–10 mN force range. Averaged points were interpolated from data. Error bars
include loading/unloading hysteresis (mean ± SE, n ≥ 8).

parameters constant, the cross-sectional area is directly proportional to deflection, (A ∝ ky0 ) where k is a negative constantthat depends on a and h. The impedance however, is an inverse
function of the cross-sectional area between the electrodes,
as per (2), (|Z| ∝ 1/A), and therefore also varies inverselywith sensor deflection (|Z| ∝ 1/y0 ). The impedance responseis therefore expected to vary nonlinearly in similar fashion
to capacitance-based sensors that utilize electrode separation
as the measurement variable. EI-based tracking of deflection
proved to be robust and repeatable Fig. 9(b). Without utilizing
any additional signal conditioning techniques, it was possible
to resolve membrane deflections at submicron levels (typically
250 nm) yielding a calculated minimum force resolution of±22.5 μN. Although this relationship is inherently nonlinear,
it is clear that impedance provides an accurate and reliable
method for tracking sensor deflection.
Plotting real-time impedance–force data generates characteristic response curves capturing sensor behavior during loading/unloading cycles. The data were repeatable and robustwith maximum forces measured up to 79 mN at full sensor
deflection. Impedance–force curves at various levels of deflection for a single sensor are provided in Fig. 10(a). The
impedance–force response was reproducible and repeatable
over fourteen deflection cycles. The most sensitive range ofmeasurement was found below 8–10 μm of deflection coinciding favorably with a region of linearity and minimal hysteresis.
Beyond this range, impedance–force response demonstrated
built-in hysteresis likely due to viscoelastic deformation of
the Parylene. FEM modeling of Von Mises stresses at theseelevated deflection values indicated stress magnitudes near theyield strength of Parylene (59 MPa) [43] and provided anindication of the practical limits for completely elastic sensor
operation. It is worth noting that following an initial rapid
increase where response is quite linear, the impedance response
plateaus beyond about 10 μm of deflection coinciding with thenonlinear region of the force–deflection curve.
A similar characteristic response was observed across fiveunique sensors tested with a minimum of eight deflections
cycles each [Fig. 10(b)]. Maximum changes in the nominal
impedance ranged 35–60 kΩ (measured in the plateau region> 3000 μN) over steady-state impedance values. An average

TABLE IP ERFORMANCE OF F IVE U NIQUE S ENSORS W ITH DI WATER
AS W ORKING E LECTROLYTE

sensitivity of 10.6 ± 3.7 Ω/μN (mean ± SD, n = 5) was
measured in the linear region of operation (< 3000 μN).
A summary of key performance parameters is provided inTable I. Although individual sensor responses were quite consistent, some variation was observed between individual sensors. The average nominal value of the resting impedance was
1.11 MΩ(n = 5) but ranged from 0.5–1.8 MΩ. Variation in the
maximum normalized sensor response was also observed andranged between 4% and 10%. These differences are thought
to arise from photoresist residue remaining in the chambers
which can impact the conductivity of the medium. The removal
process is diffusion limited, and although sufficient time wasallotted for near complete removal across a majority of devices,
some residue was still visible in a few devices. Further variability may also be introduced during fabrication, specifically incontrolling membrane thickness and sensor height from batch
to batch.

C. Diaphragm Convexity
Device performance is a function of several factors including
membrane thickness, device diameter, and convexity [40]. The
effect of diaphragm convexity was clearly noted by the responseof devices that underwent hard-baking following the patterningof the chamber sacrificial layer. It is well known that additional
baking results in photoresist reflow and consequently, morerounded features. The additional baking step therefore results in
a convex photoresist sacrificial layer for the chamber as verified
by surface profilometry. This shape is replicated in the Parylene
chamber diaphragm and results in a response with very different

GUTIERREZ AND MENG: IMPEDANCE-BASED FORCE TRANSDUCTION WITHIN FLUID-FILLED PARYLENE MICROSTRUCTURES

1105

Fig. 11. Force–deflection data attained with high resolution force probe. (a) Convexity was observed following the addition of a baking step during fabrication.
Representative data for one loading cycle with a high resolution force probe is shown. Error bars indicate time-averaged force values for each 2 μm deflection step
applied (mean ± SD, n ≤ 20). The generalized characteristic response of a convex diaphragm is provided in the inset. (b) Response of flat diaphragm provided
for comparison.

Fig. 12. High resolution force and impedance responses of two individual sensors during a loading/unloading cycle (2 μm step increments). Data were acquired
at 4 Hz. (a) Sensor response with unobstructed fluidic access. (b) Sensor response with obstructed fluidic access, provided for comparison.

characteristics. Convex diaphragms exhibit significant hysteresis including a dead-band region over which “snap action” occurs. Snap action is characterized by a sudden step-like increasein deflection at a threshold force level. However, this “snap”
point occurs at different points during the loading/unloading
cycle resulting in the observed hysteresis. Measurements withthe high resolution force probe were particularly useful incapturing the convexity effect which was evident only at low
force levels and indicated nearly 5 μm of travel prior to thesnap point for the sensor tested [Fig. 11(a)]. It should be
noted that the initial distance required to reach the snap point
must be accounted for when comparing to the force–deflection
curves of a flat diaphragm [Fig. 11(b)]. Due to the smaller
probe tip used for these measurements, the magnitude of theforces generated is expected to be approximately 3–4 times
smaller than those generated under the large tip condition. This
characteristic response was observed over four separate runs
on one device. For force sensing applications, this behavior
is generally undesirable; however, it may be useful in otherapplications such as push-buttons and threshold detection. The
response can be tailored by controlling the bake time and thus
convexity ratio of the diaphragm.

D. Temporal and Frequency Response
The effective spring constant of the diaphragm was estimatedfrom the small deflection response and measured to be approximately 266 N/m. The resonant frequency fn of the Parylene
diaphragm with a clamped edge boundary condition in air was
calculated to be fn = 811 Hz as determined by the relation
1/2
where σ is the density of the
fn = 9.22 t/a2 [E/σ(1 − μ2 )]diaphragm material [40]. A more accurate value accounts forthe viscous damping of the fluid surrounding the diaphragm.
Scaling by (1 + β)−1/2 , where β = 0.699(σf a/σh) and σf /σis the density ratio of the fluid to diaphragm material, yields
a resonant frequency of 222 Hz. It is expected that further
accounting for the hydraulic resistance across the two accessports will reduce this value even further. However, given that
there was no observable time constant in the data that wereacquired at a rate of 4 Hz [Fig. 12(a)], a conservative estimate
would place the nominal value of the resonance frequency on
the order of 10–50 Hz, which in most cases is sufficient for
tracking forces generated in surgical settings over time scalesof minutes to hours. The flow resistance through channels/port
structures was not found to be a significant factor to proper
device operation under the conditions of interest; clear step-like

1106

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 20, NO. 5, OCTOBER 2011

responses were observed during both deflection and release.
Further characterization would however be necessary to fully
assess the frequency response under rapid loading/unloading
cycles.
In some cases, however, the force response exhibited aclear time constant during both deflection and release. The
value of the time constant was estimated from the data to be
10–12 s [Fig. 12(b)]. This was assessed by applying discrete
deflection/release steps in 2 μm increments. The slow-adapting
behavior was most notable in the force response; however, it
was also observed in the impedance response during the release
cycle. The observed behavior was attributed to flow constriction
through the fluidic access ports and possible port collapse. It
is important that fluid flow into and out of the structure is
not impeded so as to not inhibit the mechanical behavior of
the chamber. Occlusion or insufficient opening of the fluid
access ports during fabrication significantly altered the desired
mechanical behavior. Additional channels or larger ports mayhelp in reducing the chance of failure due to blockage as well
as reducing the effects of hydraulic resistance.
E. Applications and Additional Considerations
To our knowledge, this is the first demonstration of force
measurement utilizing EI as the basis for transduction withina microdevice. This technique was previously demonstrated
for robotic tactile sensing applications at larger scales withbiomimetic models of the human digit [44] but required advanced signal conditioning to extract force and feature information. Compared to other polymer-based approaches, EI-basedtransduction provides a new alternative to resistive and capacitive techniques [45], [46]. A significant advantage over conventional elastomeric approaches is the potential for reducing
crosstalk between sensor elements. Because the sensors are
independent and not embedded within a bulk material, thelateral crosstalk encountered by embedded approaches is significantly reduced. Silicon-based approaches also predominantly
use capacitive or piezoresistive transduction and have produced
a variety of sensor technologies. Many of these technologies
have been introduced for tactile feedback as part of minimallyinvasive surgery applications with some success [12]; however, packaging and flexibility issues remain. The combinationof sensitivity and simple packaging afforded by fluid-filled
impedance-based sensors provides an attractive alternative tothese technologies.
For neural prosthetic applications involving the implantationof devices into the brain, spinal cord, or retina, very few sensor
solutions exist. The integration of sensor technology for suchapplications could assist in improving implant design, material
selection, and the understanding of biomechanical interaction
at these interfaces. Many neural prosthetic applications employ
small probes [47], flat electrode arrays, or flexible substrates[20], [22] making integration of existing sensor technologies
into these systems challenging. The Parylene-based technology
presented here addresses the need for biocompatibility, waterproofing, and a packaging approach tailored to the working
environment. Force measurement capability below 1 mN is also
desirable because the nervous tissues of the body are known

to be some of the softest [48], [49] with Young’s modulus
values typically in the 1–20 kPa range. Impedance-based force
sensors offer a unique sensing approach to these operating
conditions. A potential limitation to the current approach liesin the Parylene-Parylene interfaces which can be susceptible todelamination if soaked for an extended period of time. Annealing of the Parylene films above the glass transition temperature
has proved useful in addressing this issue but was not performed
here; however, no such delamination was observed during thecourse of our experiments. Although Parylene was selected as
the structural material in this design, other materials such as
silicones or polyimides could potentially be used in conjunction
with the impedance-based sensing technique to realize novel
sensor designs.
VI. C ONCLUSION
Fluid-filled Parylene-based microstructures were designed,
fabricated, and demonstrated for high sensitivity force measurements. A low temperature fabrication process was developed
utilizing multilayered Parylene structures and stacked sacrificial photoresist. This technique requires minimal packaging
and utilizes the ambient operating solution as the basis forelectrochemical measurements. Impedance-based transductionof externally applied forces was shown to be a highly sensitive measurement technique as is compatible with liquid
environments. The small form factor, flexible substrate, easeof integration, and highly sensitive response provides a usefulcombination of features making these sensors attractive for use
with biomedical force sensing applications.
ACKNOWLEDGMENT
The authors would like to acknowledge Mr. B. Kim forinvaluable assistance with the experimental setup and device
characterization. We would also like to thank Dr. D. Zhu andthe members of the USC Biomedical Microsystems Laboratory
for their assistance.
R EFERENCES[1] Y. Suzuki and Y.-C. Tai, “Micromachined high-aspect-ratio Parylene
beam and its application to low-frequency seismometer,” in Proc. IEEE
MEMS, Kyoto, Japan, 2003, pp. 486–489.
[2] P.-J. Chen, D. C. Rodger, M. S. Humayun, and Y.-C. Tai, “Unpowered
spiral-tube Parylene pressure sensor for intraocular pressure sensing,”
Sens. Actuators A, Phys., vol. 127, no. 2, pp. 276–282, Mar. 2006.
[3] C. A. Gutierrez and E. Meng, “A subnanowatt microbubble pressure
sensor based on electrochemical impedance transduction in a flexible
all-Parylene package,” in Proc. IEEE MEMS, Cancun, Mexico, 2011,
pp. 549–552.
[4] Z. Fan, J. M. Engel, J. Chen, and C. Liu, “Parylene surfacemicromachined membranes for sensor applications,” J. Microelectromech. Syst., vol. 13, no. 3, pp. 484–490, Jun. 2004.
[5] P.-Y. Li, R. Sheybani, C. A. Gutierrez, J. T. W. Kuo, and E. Meng, “AParylene bellows electrochemical actuator,” J. Microelectromech. Syst.,
vol. 19, no. 1, pp. 215–228, Feb. 2010.
[6] T. N. Pornsin-Sirirak, Y. C. Tai, H. Nassef, and C. M. Ho, “Flexible
Parylene actuator for micro adaptive flow control,” in Proc. IEEE MEMS,
Interlaken, Switzerland, 2001, pp. 511–514.
[7] H.-S. Noh, Y. Huang, and P. J. Hesketh, “Parylene micromolding, a rapid
and low-cost fabrication method for Parylene microchannel,” Sens. Actuators B, Chem., vol. 102, no. 1, pp. 78–85, Sep. 2004.
[8] X.-Q. Wang, Q. Lin, and Y.-C. Tai, “A Parylene micro check valve,” inProc. IEEE MEMS, Orlando, FL, 1999, pp. 177–182.

GUTIERREZ AND MENG: IMPEDANCE-BASED FORCE TRANSDUCTION WITHIN FLUID-FILLED PARYLENE MICROSTRUCTURES[9] P.-J. Chen, D. C. Rodger, E. M. Meng, M. S. Humayun, and Y.-C. Tai,
“Surface-micromachined Parylene dual valves for on-chip unpowered microflow regulation,” J. Microelectromech. Syst., vol. 16, no. 2, pp. 223–
231, Apr. 2007.
[10] 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.
[11] S. Yang and T. Saif, “Micromachined force sensors for the study of cellmechanics,” Rev. Sci. Instrum., vol. 76, no. 4, pp. 044 301-1–044 301-8,
Apr. 2005.
[12] M. E. H. Eltaib and J. R. Hewit, “Tactile sensing technology for minimal
access surgery—A review,” Mechatronics, vol. 13, no. 10, pp. 1163–1177,
Dec. 2003.
[13] J. Tegin and J. Wikander, “Tactile sensing in intelligent robotic
manipulation—A review,” Ind. Robot, Int. J., vol. 32, no. 1, pp. 64–70,
2005.
[14] A. Menciassi, A. Eisinberg, M. C. Carrozza, and P. Dario, “Force sensing microinstrument for measuring tissue properties and pulse in microsurgery,” IEEE/ASME Trans. Mechatronics, vol. 8, no. 1, pp. 10–17,
Mar. 2003.
[15] K. Noda, K. Hoshino, K. Matsumoto, and I. Shimoyama, “A shear stresssensor for tactile sensing with the piezoresistive cantilever standing inelastic material,” Sens. Actuators A, Phys., vol. 127, no. 2, pp. 295–301,
Mar. 2006.
[16] L. Han, J. A. Noble, and M. Burcher, “A novel ultrasound indentation
system for measuring biomechanical properties of in vivo soft tissue,”
Ultrasound Med. Biol., vol. 29, no. 6, pp. 813–823, Jun. 2003.
[17] D. B. Plewes, J. Bishop, A. Samani, and J. Sciarretta, “Visualization andquantification of breast cancer biomechanical properties with magnetic
resonance elastography,” Phys. Med. Biol., vol. 45, no. 6, pp. 1591–1610,
Jun. 2000.
[18] A. Samani, J. Bishop, C. Luginbuhl, and D. B. Plewes, “Measuring theelastic modulus of ex vivo small tissue samples,” Phys. Med. Biol., vol. 48,
no. 14, pp. 2183–2198, Jul. 2003.
[19] M. Tokida, T. Obara, M. Takahashi, M. Yamato, and S. Konishi, “Integration of cell sheet sucking and tactile sensing functions to retinal pigment
epithelium transplantation tool,” in Proc. IEEE MEMS, Hong Kong, 2010,
pp. 316–319.
[20] D. C. Rodger, A. J. Fong, W. Li, H. Ameri, A. K. Ahuja, C. A. 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.
[21] D. C. Rodger, W. Li, H. Ameri, A. Ray, J. D. Weiland, M. S. Humayun,
and Y.-C. Tai, “Flexible Parylene-based microelectrode technology for
intraocular retinal prostheses,” in Proc. IEEE Nano/Micro Eng. Mol. Syst.,
2006, pp. 743–746.
[22] J. D. Weiland, W. Liu, and M. S. Humayun, “Retinal prosthesis,” Annu.
Rev. Biomed. Eng., vol. 7, pp. 361–401, 2005.
[23] J. Rosenhorn and N. H. Diemer, “Reduction of regional cerebral bloodflow during brain retraction pressure in rat,” J. Neurosurg., vol. 56, no. 6,
pp. 826–829, Jun. 1982.
[24] T. W. Berger and D. L. Glanzman, Toward Replacement Parts for theBrain. Cambridge, MA: MIT Press, 2005.
[25] J. Dargahi, S. Najarian, and B. Liu, “Sensitivity analysis of a novel tactile
probe for measurement of tissue softness with applications in biomedical
robotics,” J. Mater. Process. Technol., vol. 183, no. 2/3, pp. 176–182,
Mar. 2007.
[26] H. Roham, S. Najarian, S. M. Hosseini, and J. Dargahi, “Design andfabrication of a new tactile probe for measuring the modulus of elasticity
of soft tissues,” Sens. Rev., vol. 27, no. 4, pp. 317–323, 2007.
[27] G. Vásárhelyi, M. Ádám, É. Vázsonyi, I. Bársony, and C. Dücso, “Effectsof the elastic cover on tactile sensor arrays,” Sens. Actuators A, Phys.,vol. 132, no. 1, pp. 245–251, Nov. 2006.
[28] C. A. Gutierrez and E. Meng, “Parylene-based electrochemical-MEMS
transducers,” J. Microelectromech. Syst., vol. 19, no. 6, pp. 1352–1361,
Dec. 2010.
[29] 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. IEEE TRANSDUCERS,
Solid-State Sens., Actuators, Microsyst. Conf., Denver, CO, 2009,
pp. 2194–2197.
[30] M. E. Orazem and B. Tribollet, Electrochemical Impedance Spectroscopy.
Hoboken, NJ: Wiley, 2008.
[31] M. A. Schwarz and P. C. Hauser, “Recent developments in
detection methods for microfabricated analytical devices,” Lab Chip,
vol. 1, pp. 1–6, 2001.

1107[32] J. E. B. Randles, “Kinetics of rapid electrode reactions,” Discuss. Faraday
Soc., vol. 1, pp. 11–19, 1947.
[33] D. Erickson and D. Li, “Integrated microfluidic devices,” Anal. Chim.
Acta, vol. 507, pp. 11–26, 2004.
[34] A. J. Bard and L. R. Faulkner, Electrochemical Methods: Fundamentals
and Applications, 2nd ed. Hoboken, NJ: Wiley, 2001.
[35] D. A. Robinson, “The electrical properties of metal microelectrodes,”Proc. IEEE, vol. 56, no. 6, pp. 1065–1071, Jun. 1968.
[36] R. C. Gesteland, B. Howlandt, J. Y. Lettvin, and W. H. Pitts, “Comments on microelectrodes,” Proc. IRE, vol. 47, no. 11, pp. 1856–1862,
Nov. 1959.
[37] C. de Haro, R. Mas, G. Abadal, J. Munoz, F. Perez-Murano, and
C. Dominguez, “Electrochemical platinum coatings for improving performance of implantable microelectrode arrays,” Biomaterials, vol. 23,
no. 23, pp. 4515–4521, Dec. 2002.
[38] J. J. Whalen, J. D. Weiland, and P. C. Searson, “Electrochemical deposition of platinum from aqueous ammonium hexachloroplatinate solution,”
J. Electrochem. Soc., vol. 152, no. 11, pp. C738–C743, 2005.
[39] B. Bhushan, Handbook of Micro/Nanotribology, 2nd ed. Boca Raton,
FL: CRC Press, 1999.
[40] M. D. Giovanni, Flat and Corrugated Diaphragm Design Handbook,vol. 11. New York: Marcel Dekker, 1982.
[41] C. A. Gutierrez, C. McCarty, B. Kim, M. Pahwa, and E. Meng, “An
implantable all-Parylene liquid-impedance based MEMS force sensor,”in Proc. IEEE MEMS, Hong Kong, 2010, pp. 600–603.
[42] R. Bashir, “BioMEMS: State-of-the-art in detection, opportunities andprospects,” Adv. Drug Deliv. Rev., vol. 56, no. 11, pp. 1565–1586,
Sep. 2004.
[43] C. Y. Shih, T. A. Harder, and Y. C. Tai, “Yield strength of thin-filmParylene-C,” Microsyst. Technol., vol. 10, no. 5, pp. 407–411, Aug. 2004.
[44] 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.
[45] E.-S. Hwang, J.-H. Seo, and Y.-J. Kim, “A polymer-based flexible tactile
sensor for both normal and shear load detections and its application
for robotics,” J. Microelectromech. Syst., vol. 16, no. 3, pp. 556–563,
Jun. 2007.
[46] L. Beccai, S. Roccella, A. Arena, F. Valvo, P. Valdastri, A. Menciassi,
M. C. Carrozza, and P. Dario, “Design and fabrication of a hybrid silicon
three-axial force sensor for biomechanical applications,” Sens. Actuators
A, Phys., vol. 120, no. 2, pp. 370–382, May 2005.
[47] K. A. Moxon, S. C. Leiser, G. A. Gerhardt, K. A. Barbee, andJ. K. Chapin, “Ceramic-based multisite electrode arrays for chronic
single-neuron recording,” IEEE Trans. Biomed. Eng., vol. 51, no. 4,
pp. 647–656, Apr. 2004.
[48] L. Galli-Resta, P. Leone, D. Bottari, M. Ensini, E. Rigosi, and E. Novelli,“The genesis of retinal architecture: An emerging role for mechanical interactions?” Prog. Retinal Eye Res., vol. 27, no. 3, pp. 260–283,
May 2008.
[49] 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.

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 currently works at the National Science Foundation’s Biomimetic Microelectronic Systems Engineering Research Center, where he is a Co-President of the Student LeadershipCouncil. His past research was carried out in the Caltech Micromaching
Laboratory and was focused on identifying, analyzing, and implementing
power harvesting and storage circuitry for microelectromechanical-systemsbased power harvesting devices. He is currently conducting research toward thedevelopment of novel micro- and nanotechnologies for biomedical applications.
His specific research is focused on the development of retinal prosthesis
technologies to restore sight to the blind.
Mr. Gutierrez was a recipient of the Caltech Presidential Scholarship, Billand Melinda Gates Millennium Fellowship, and the USC Viterbi School ofEngineering Kunzel Fellowship.

1108

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 20, NO. 5, OCTOBER 2011

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.
She is an Associate Professor in the Departmentof Biomedical Engineering, University of Southern
California, Los Angeles, where she has been since2004. She currently holds a joint appointment in
the Ming Hsieh Department of Electrical Engineering. Her research interests include bioMEMS, implantable biomedical microdevices, microfluidics, multimodality integrated
microsystems, and packaging. She held the Viterbi Early Career Chair in theViterbi School of Engineering.
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. In 2009, she was recognized as one of the TR35 Technology
Review Young Innovators under 35.