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Parylene-Based Electrochemical-MEMS Force
Sensor for Studies of Intracortical Probe
Insertion Mechanics
Brian J. Kim, Christian A. Gutierrez, and Ellis Meng, Senior Member, IEEE
Abstract— To investigate the mechanical interactions between
the implanted cortical multielectrode probes and brain tissue,
a Parylene C-based electrochemical-microelectromechanical systems force sensor array was developed. The array consists ofseven linearly distributed sensor units arranged along the lengthof a flexible Parylene C microchannel. The seven sensor units areformed by eight fluidically coupled and adjacent platinum electrode pairs enclosed within the microchannel. Deformation of the
top surface of the mechanically compliant microchannel changes
the volumetric conduction path between the pairs of sensing electrodes, and therefore, the measured electrochemical impedance,
which is proportional to the magnitude of the contacting force.
Each sensor unit demonstrated a linear response from 0 to 60 mN
with a sensitivity of 0.13 ± 0.01 percentage change inimpedance/µN (%/µN; mean ± SE, n = 6). The sensor arrays
were mounted onto a ceramic intracortical probe and inserted
into the tissue phantoms to verify in situ functionality and assessinterfacial probe mechanics. Probe surface force distribution wasmeasured under various insertion speeds and the results indicatedthat interfacial forces are distributed nonuniformly along theprobe shaft length, concentrating within the first 1 mm of theadvancing probe tip. Faster insertion speeds were also found to
decrease the magnitude of the interfacial forces, suggesting thatthe tissue strain during cortical implantation may be minimized
through appropriate selection of the insertion speed.[2014-0320]
Index Terms— Force sensor, electrochemical-microelectromechanical systems (EC-MEMS), electrochemical sensing,electrochemical transduction, impedance-based sensor, liquid
impedance, Parylene C, mechanotransduction.

I. I NTRODUCTIONCHRONICALLY implanted intracortical probes provide
a means for recording neural signals or applying
electrical stimuli within the sensory and/or motor areasof the brain, fueling their development for clinicalbrain-machine interfaces (BMIs) and neuromodulation.
However, the potential of these implantable neural prosthetics

Manuscript received October 15, 2014; revised January 8, 2015; accepted
April 1, 2015. Date of publication April 16, 2015; date of current version
September 29, 2015. This work was supported in part by the Engineering
Research Centers Program through the National Science Foundation underAward EEC-0310723, in part by the University of Southern California Provost
Ph.D. Fellowship, and in part by the Bill and Melinda Gates Foundation.
Subject Editor A. J. Ricco.
B. J. Kim is with the Department of Biomedical Engineering, University ofSouthern California, Los Angeles, CA 90089 USA (e-mail: brianjk@usc.edu).
C. A. Gutierrez is with Fluid Synchrony, LLC, Los Angeles, CA 90089 USA(e-mail: christian@fluidsynchrony.com).
E. Meng is with the Department of Biomedical Engineering, Ming Hsieh
Department of Electrical Engineering, 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.2015.2420043

as enabling technologies is challenged by the delicate
interfacial mechanics between the probe and the soft corticaltissue it penetrates [1]. The initial stab wound injury from
implantation of rigid metal or silicon probes and the resultant
disruptions of the blood-brain-barrier (BBB), combined withthe resulting foreign body response have been found todiminish their efficacy for chronic recordings [2]. In addition,
cortical micromotion due to vascular pulsing induces aconstant mechanical agitation between the soft tissue andrigid probes [3]. The corresponding effects on chronic probe
performance have yet to be fully understood, largely due toinadequate measurement technology.
Previously, external load cells were used to measuremacroscale normal insertion forces [4]–[6] as well as tocompare differing probe tip designs [4], [7], but out-of-planestresses acting on the probe surface could not be detected,
providing limited insight into interfacial phenomena. Also,
studies comparing normal insertion force measurements alone
have found poor correlation between normal force magnitudes
and chronic probe performance [7]. Modeling [8]–[10]and ex vivo probe implantation experiments utilizing
cross-sectional viewing planes [11] have provided great
insight into the theoretical distribution of strains and forces
during implantation, at the probe-tissue interface, and during
micromotion events. However, quantitative data obtained fromsurgical and in vivo conditions to capture these mechanicalinteractions during insertion and micromotion events in realtime is lacking.
Technologies to integrate sensors on probes have alreadybeen developed, but many of these sensors are difficult toapply or integrate and provide limited access to interfacialmechanics. For example, CMOS sensors integrated on a siliconprobe were used to measure effects of insertion-derived probe
deflection stresses during implantation [12]. Though resultant
probe deflection due to insertion into tissue does capturesome interfacial forces, the method cannot detect interfacial
forces normal to the probe from the surrounding neuraltissue. Also, these sensors are currently limited to siliconprobes. Furthermore, commercial probes readily available toneuroscience researchers do not offer the option of sensor
integration. A more versatile approach is necessary to enable
comparative studies between different electrode designs andmaterials. Sensors must be able to: accommodate small shankareas, integrate with CMOS-incompatible electrode shankmaterials, and operate in aqueous and saline environments.
By utilizing sensor technology in addition to histological1057-7157 © 2015 IEEE. Personal use is permitted, but republication/redistribution requires IEEE permission.
See http://www.ieee.org/publications_standards/publications/rights/index.html for more information.

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Fig. 2. Schematic of electrochemical impedance-based transduction methodof contact forces utilizing a flexible, mechanically compliant microchamber.
Fig. 1. Equivalent circuit model of measuring the electrochemical impedance
between a pair of electrodes in an electrolyte. Note the use of the simplified
Randles circuit model at the two electrode-electrolyte interfaces.

methods, benchtop or in vivo studies (such as comparing
probe designs, materials, and insertion parameters) can be
conducted to correlate interfacial mechanical interaction
data with chronic neural probe integration and performance.
In addition, these integrated devices can be used to effectively
characterize micromotion-related forces and their effects on
long-term performance. Understanding these design factors
and corresponding mechanical interactions at the prosthesistissue interface in vivo is one large step towards the realizationof reliable probe technologies that minimize tissue damage andthe related immunological responses that hinder the progressof chronically implanted neural prosthetic technologies [13].
To address this unmet need, we adapted ourParylene C-based electrochemical-MEMS (EC-MEMS)
sensing technology [14]–[16] to transduce force and pressure
distributions directly on the probe surface. Here, we present
a sensor layout capable of integrating directly onto existing
implantable ceramic electrode shanks, with its subsequent
device packaging and testing, including thermal annealing.
We also demonstrate measurement, for the first time, of thenormal force distribution produced along the shank surface inagarose tissue phantoms.
II. O PERATING P RINCIPLE AND D ESIGN
EC-MEMS sensors use electrochemical impedance as asensing modality [14], [17], [18]. These sensors usually consistof a pair of electrodes exposed to an electrolyte (a conductive fluid) enclosed within a Parylene C (hereon referred toas Parylene) microstructure. Integrated port openings allow
fluid exchange with the local environment. To understandthe impedance-based transduction mechanism, it is useful toconsider the well-known simplified Randles circuit model ofthe electrode-electrolyte interface [19].
A more detailed discussion of the simplified Randles circuit
model for EC-MEMS sensors can be found in previous
work [14], but is briefly summarized here. In the model,
the double-layer capacitance (Cdl ) and the charge transfer
resistance (Rct ) approximate the complex charge transfer
phenomena from the electrolyte by the electrodes (Fig. 1).

At this electrode-electrolyte interface, there exist two layersof charge known as the double-layer, which is formed bycorresponding ions and molecules in the electrolyte responding
to excess charges on the surface of the electrode [20]. The
reversible reactions at this interface are responsible for thecapacitive and resistive characteristics observed as charge istransferred between the electrode and electrolyte [21], [22].
An alternating current (AC) signal applied to the
EC-MEMS sensors is used to measure the electrochemical
impedance between a pair of electrodes in an electrolyte.
If the frequency of this AC signal is sufficiently high ( f meas ),the double-layer effects are diminished as the solution
resistance (Rs ) dominates the impedance response (Fig. 1).
Thus, in this sensing configuration, the sensor is effectively
a variable resistor, and a number of variables (conductive
media (ρ), distance between the electrodes (l), and crosssectional area ( A)) can affect a change in the measured
solution resistance, and thus electrochemical impedance,
accordingly. The complex impedance for a two-electrodesystem is expressed in (1) with the high-frequency
approximation provided in (2).
Z=

1
1
2Rct
jωCdl RctjωCdl Rct
+
R
+
=
+ Rs
s
1
1
1 + Rct jωCdl
jωCdl + Rct
jωCdl + Rct

(1)
ρl
(2)
Z ≈ Rs = ; for sufficiently high ωAThe principle of sensor operation relies on external contact
forces deforming a compliant electrolyte-filled structure in thenormal direction and redistributing the fluid contained within
the chamber. This alters the volumetric ionic conduction
path of current-carrying ions in the fluid, in effect modifying
the cross sectional area between the two electrodes, andregistering as a change in the magnitude of the solution
impedance (Fig. 2). In this manner, impedance variations canbe correlated to mechanical interfacial contact forces exerted
by the tissue in contact with the top of the sensor array.
However, because of the use of an electrochemical sensing
scheme, the sensor is particularly sensitive to changes inelectrolyte compared to other membrane-based force sensors.
There exists a trade-off between electrolyte sensitivity andmechanical flexibility; current membrane-based force sensors

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Fig. 3. (a) Optical micrograph of full sensor array and integrated Parylene cable. Microchannel structure is highlighted in pink. (b) Scanning electron
micrograph (SEM) image showing top view of the sensor array with a fluidic port at the end of the microchannel. (c) Top-down image of
Parylene microchannel (pink) indicating the electrodes and fluidic ports located at the ends. (d) SEM showing the cross section of the 20 μm
Parylene microchannel sensing structure. (e) Microelectrode layout and sensor numbering based on adjacent electrode pairs.

Fig. 4.

Operation principle of the Parylene microchannel-based interfacial force sensor (a) pre and (b) post insertion into tissue.

require bulk, hermetic packaging to resist the high saline
environment of the body. By leveraging the electrolytic
environment for sensing, the EC-MEMS sensor design
eliminates the bulky packaging required by current sensors
and results in a low-profile and flexible device ideal for achronic implant.
The intracortical Parylene-based EC-MEMS sensor
array technology is based on distributed sensors
arranged along a Parylene microchannel-sensing element(100 μm × 4.5 mm × 24 μm) to match the footprint of aceramic-based cortical probe [23], notable for exceptional
chronic recording and electrochemical sensing performancewithin monkey cortical work, for initial testing. The
microchannel encloses eight fluidically-coupled platinum (Pt)electrodes (100 μm ×130 μm) arranged at a pitch of 500 μm,
with adjacent pairs forming a sensor unit (Fig. 3). Each arrayconsists of seven sensor “pixels” along the length of theprobe; sensor 1 begins at the tip of the probe and sensor 7 is
at its base (Fig. 3e). This sensor configuration was selected tomaximize sensor density over the area of interest defined by
the cortical probe layout and to maintain transducer symmetry.

Etched fluidic ports at the channel ends allow the surrounding
electrolyte to fill the channel. Fluid can flow freely between
the external environment and the channel via these accessports. A Parylene flat flexible cable (FFC) with eight leads
was also integrated into the device design in order to establish
electrical contact via a zero-insertion-force (ZIF) connector.
Because of the flexible and thin Parylene microchannel
structure (4 μm thickness), contact forces (i.e. elastic andspring forces) from surrounding cortical tissue will deform the
top surface along the length of the probe (Fig. 4). It is important to note here that the sensors are insensitive to hydrostaticforces because the fluidic ports allow for an open fluidic
connection between the inside and outside of the channel; thesensing element deforms primarily due to mechanical contact
forces on the top channel surface. The seven sensor pixels
along this length quantify the distribution of interfacial forcesthat the probe imparts on the surrounding cortical tissue during
insertion, extending a similar concept from previous sensor
technology developed to quantify interfacial forces produced
during implantation of an epiretinal implant onto the retinal
surface [24]. The flexible substrate (total thickness = 10 μm)

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Fig. 5. Overview of fabrication process for the Parylene-based EC-MEMS sensor array. (a) Platinum electrodes and contact pads were patterned by liftoff
onto a Parylene substrate; an insulation layer was also deposited and etched to reveal electrode sites. (b) Sacrificial photoresist was spun on and patterned toform the microchannel structure. (c) Another Parylene layer was deposited to form the deformable top membrane structure and fluidic ports were etched into
the top using oxygen plasma; the device perimeter was then etched using O2 plasma. (d) Devices were released off the carrier wafer and soaked in acetone
to remove the sacrificial photoresist. Then the devices were filled with 1× PBS prior to testing.

also allows for simple, low-profile instrumentation of theceramic cortical probes. The microchannel and electrodelayout can be modified to fit probe footprints of a variety ofdesigns, demonstrating its robustness and versatility.
III. D EVICE FABRICATION AND PACKAGING
Standard surface micromachining processes were used forsensor array fabrication (Fig. 5). All fabrication processes were
performed at low temperatures (90 °C) to prevent excessive
thermal cycling of the Parylene structural material. Parylene
was chosen as the sensor material for its biocompatibility,
mechanical strength, electrical insulation properties, andcompatibility with micromachining processes.
Platinum electrodes and contact pads (2000 Å) were
patterned by lift-off on a Parylene substrate (5 μm) coatedonto a silicon carrier wafer. Following deposition of aParylene insulation layer (1.5 μm), openings for electrodes
and contact pads were etched using oxygen plasma.
A sacrificial photoresist (PR) process was used to form themicrochannel, but because of the relatively tall height (20 μm),
two PR layers (AZ 4620; AZ Electronic Materials,
Branchburg, NJ) were spun on (each 10 μm thick) toaccurately and uniformly establish the desired microchannelheight. Following this, a 4 μm thick layer of Parylene wasdeposited to form the final device structure. This thickness
was important as it set the sensor mechanics; thicker films
would displace less than thinner microchannels. A thicknessof 4 μm was chosen based on previous sensor work on
measuring contact forces between epiretinal implants andretinal tissue [14]. Fluidic access ports at the channel ends
were opened by oxygen plasma. Finally, to facilitate precise
shaping of the array for attachment to a fine ceramic shanktip, arrays were singulated using oxygen plasma.
Devices were released from the wafer by stripping
the protective photoresist mask for the cut-out etch withacetone, and then submerging the wafer in DI water. The
hydrophobicity of the Parylene polymer allowed the arrays
to easily separate from the native oxide of the silicon carrier
wafer and lift-off the surface. Sacrificial photoresist within
the channel was then removed by immersion in acetone
and isopropyl alcohol following rinse in deionized water.
The individual arrays were then let to air dry overnight.
The mechanical integrity of the Parylene microstructure wassufficient to avoid stiction during the drying process.

Fig. 6.
Image showing integrated Parylene flex cable inserted into a
ZIF-ZIF connector for electrical connection to the sensor.

Following fabrication, sensors underwent post-processing
steps for electrical packaging to the measurement system.
Electrical connections to the integrated Parylene cable wasmade via an 8 contact ZIF connector (8 channel, 0.5 mm pitch;
Hirose Electric Co., Simi Valley, CA), a hinge-based connector
that allows for a zero-insertion force method when the cable
is inserted into the bay of the connector. As the connector
requires a cable thickness of 300 μm, a polyetheretherketone (PEEK) polymer backing was applied as a stiffener on
the back of the contact pad region. This connection scheme
enabled rapid, epoxy-less connections to multiple contact pads
supported on our flexible Parylene substrate [25] (Fig. 6).
The ZIF connector used in this packaging method wasa ZIF-ZIF adaptor that connected the sensor array to acommercial polymer flexible cable (Molex Inc., Lisle, IL)
that was thicker and more robust for handling; theZIF-ZIF adaptor was made by soldering the legs of
two separate ZIF connectors together. The addition of thepolymer flexible cable allowed for more robust, repeatableconnections into the custom measurement PCB via another
ZIF connector compared to the thin Parylene cable.
IV. E XPERIMENTAL M ETHODS
To measure the electrochemical impedance from all sevensensors of the array, a custom multiplexer was designedthat allowed for user-controlled switching between various
pairs of electrodes for impedance measurement using aprecision LCR meter (Agilent E4980A; Agilent Technologies,
Santa Clara, CA). The multiplexer was controlled digitally

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Fig. 7. (a) Load-displacement setup for mechanical and calibration testing of Parylene force sensor array. (b) Optical micrograph taken using a camera
underneath testing setup illustrating deflection probe (pink) displacing into sensor element.

through a LabVIEW graphical user interface (GUI) which
cycled through adjacent electrode pairs. Impedance wasmeasured at 1 kHz, with an amplitude of 1 Vpp .
Characterization of the sensors was carried out on thebenchtop in a custom acrylic jig designed to keep the sensor
array immersed in an electrolyte (1× PBS; Fig. 7a). A flat,
circular, probe tip (diameter = 630 μm), attached to a
motorized z–axis stage, was used to indent the microchannel
in 2 μm increments until reaching the maximum deflectionof 20 μm (Fig. 7b). The probe diameter was chosen to
cover the entire surface of a sensor unit (60,000 μm2 ).
An in-line load cell (0-50 g, LCFA-50; Omega Engineering,Stamford, CT) was used to measure the applied forceof the indenting probe during displacement. During thiscalibration procedure, real-time impedance was measured
across all sensor units using LabVIEW software. A 10-minuterelaxation step was used between calibrations of sensors
to ensure that the microchannel structure returned to thebaseline reading (impedance ≤ 0.1% of initial value). Also,
to further decouple the sensor units of the array, sensors were
calibrated in an order chosen to prevent the sequential testingof two adjacent sensors.
For the case in which these sensors are to be used inchronic implant conditions, additional steps are necessary
to counter observed soaking-induced delamination failure ofmulti-layered Parylene C devices [26]. Thermal annealingof Parylene devices has been reported to improve interlayer
adhesion and device lifetime during prolonged soak
conditions [27]–[31]. The annealing process is thought
to increase polymer chain entanglement between the
Parylene layers and provide additional mechanical anchoring
between the chains to resist delamination and improvemoisture barrier performance [32]. Annealing is achievedat temperatures above the glass transition temperature of
Parylene (Tg = 60-90 °C) in order to provide enough energy
for movement of the polymer chains. However, the heat
treatment of Parylene is also known to produce changes incrystallinity of the bulk polymer (hence the importance ofParylene’s thermal budget during device fabrication) [33],which may influence device performance. Because of this,device performance pre- and post-annealing was evaluated toassess any mechanical changes due to the annealing process.
Sensor arrays were placed within a vacuum oven
(10 mTorr), which was ramped at 1.6 °C/min to 200 °C.

Following a thermal soak time of 48 hours, the devices
were cooled overnight (∼15 hours) under vacuum, and thenremoved for post-process characterization. Untreated sensors
were released off the wafer and also tested for comparison.
As the mechanical and electrochemical properties of thesensor drive its performance, the effects of annealing onthese two aspects were analyzed via scanning electron
microscopy (SEM), profilometry, and load-deflection testsof the sensing structures immersed in an electrolyte
solution (1× PBS). Two point electrochemical impedance
spectroscopy (EIS; 1× PBS, 20 Hz–1 MHz), a commonlyused electrochemical technique to assess electrode surface
properties, was conducted between adjacent electrodes usinga Gamry Reference 600 potentiostat (Gamry Instruments,
Warminster, PA) to determine changes to the electrochemical
properties of the sensing electrodes.
V. R ESULTS AND D ISCUSSION
The sensor mechanical and force response was characterized
on unannealed devices after the optimal electrochemical
impedance measurement frequency was evaluated. The effectsof thermal annealing on sensor performance were thenevaluated and compared to un-annealed samples. Finally,
the sensor array was mounted onto a ceramic cortical probe
for benchtop insertion tests to demonstrate proof-of-concept
sensor operation.
A. Obtaining fmeas
Prior to calibration, EIS was performed across all sensor
electrode pairs to determine the optimal fmeas at which
the solution resistance dominates the complex impedance
response. An excitation frequency of ∼2 kHz was found tocorrespond to the maximum resistive response at a phasevalue near 0° (Fig. 8). However, to ensure signal fidelity
from parasitic capacitances of the measurement system,
all subsequent impedance measurements were performed
at 1 kHz, which was sufficient to isolate the solution resistance.
B. Sensor Characterization
The Parylene microchannel was found to deform uniformly
along the length of the channel, as the displacement per
applied force was similar across runs for each sensor (Fig. 9).

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Fig. 8. Impedance (a) magnitude and (b) phase plots of 18 individual sensors. An f meas of 1 kHz (dashed line) was selected and corresponds to maximum
resistive response at a phase value near 0°.

Fig. 9.
Load-displacement plot for the Parylene microchannel sensing
element of the EC-MEMS sensor array indicating that operational uniformity
was maintained despite their differing locations along the fluidically coupledchannel.

Plotting cyclic loading/unloading force-displacement data
generated curves that capture the sensor’s dynamic mechanicalbehavior. The channel structures were deformed at a rateof 1 μm/s, held at a maximum deflection of 20 μm for 20 s
and then returned to the initial position at 1 μm/s; the next
loading cycle was initiated 30 s following the return to resting
position. Five consecutive loading/unloading cycles wereapplied. Results indicated negligible hysteresis regardlessof the sensor position (Fig. 10). Although higher frequency
dynamic analysis was not performed, the frequency responseof the Parylene microchannel structure in air was calculated to
be 570 kHz using FEM software (COMSOL, Burlington, MA).
It is expected that this frequency would be further reduced
to account for fluidic damping to a conservative estimateof 10-100 Hz, as observed in other resonant frequency studies
of Parylene thin film diaphragms [14], [34]. Fortunately, as
the interfacial force dynamics within the cortex are very slow

(maximum ∼3–5 Hz of micromotion due to vascularpulsing [5]), the current sensor design is more than sufficientfor this application. Additional dynamic loading tests werealso performed to assess the effects of varying step-sizes(4, 6, 12, 20 μm) on possible deflection-linked dynamic
mechanical responses. Results indicated reproducible andrepeatable displacement-load force curves for varying
displacements and small hysteresis over five cycles (0.01 Hz)
at each depth. In some instances however, considerable
hysteresis was observed, which suggests yield and process
variations that can create non-uniform microchannel structures
and thus varying mechanical responses (data not shown).
A sensitivity of 0.13 ± 0.01 percentage change inimpedance/µN (%/µN; mean ± SE, n = 6) was found foreach sensor unit (Fig. 11) from the calibration curve, withina sensor working range of 0-60 mN. Without utilizing any
signal conditioning techniques, the resolution of the sensor isdependent on the accuracy of the impedance measuring device.
Using the high precision LCR meter (“short” measurement
function, ± ∼0.2% impedance magnitude), the resolution of
the measurement systems was calculated to be ∼±1 mN.
Mechanical crosstalk was also assessed by indenting asingle sensor and monitoring the responses of the othersensors in the array concurrently (Table I). Crosstalk was
calculated as the ratio of the percentage change in impedanceof the site recorded to that of the site indented. The coloredtable facilitates visualization of relative crosstalk magnitudes.
As indicated in the table below, mechanical crosstalk between
sensors was generally ∼2-10%, allowing for independent
discrimination of forces based on location along the lengthof the sensor array.
C. Annealing Effects on Sensor Performance
Detailed results and analysis of the effects of thermal
annealing on sensor performance can be found in previous
work [35], but key points are briefly mentioned here forcompleteness. Load-deflection tests revealed a ∼1.6× increasein the stiffness of the structure following annealing,

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Fig. 10. Loading and unloading cycle results of (a) four different sensors at maximum deflection and (b) at various deflection depths indicating negligible
hysteresis.
TABLE IC ROSSTALK WAS C ALCULATED AS THE R ATIO OF THE P ERCENT I MPEDANCE C HANGE OF THE S ITE R ECORDEDOVER T HAT OF THE S ITE I NDENTED . M EAN ± SE, n = 3. ∗ F OR ROWS 1 AND 4, n = 2

thin-film metal and underlying Parylene [37]. These changes
decreased the sensitivity of the sensor by 24%. FEM utilizing
COMSOL to model the observed changes following annealing(i.e. increased Parylene stiffness and decrease in electroactive
surface area) suggested that the increase in stiffness of the
Parylene film contributed the most to the differences in sensor
performance for these devices. As the sensing mechanism ischiefly derived from the mechanics of the sensing structure,
any changes to the stiffness of the polymer have a significant
impact on sensor response.
D. Instrumentation of Cortical Probes

Fig. 11.
Obtained calibration curve for the EC-MEMS sensor showing
linearity within the force range of 0-60 mN.

likely due to a combination of microchannel height shrinkage
and an increase in the elastic modulus of the annealed
Parylene attributable to increased polymer crystallinity [36].
EIS results indicated a decrease in the surface roughness ofthe electrodes following annealing, due to smoothing of theInstrumentation of a representative neural probe, aninert ceramic cortical probe [38] notable for exceptional
chronic recording and electrochemical sensing performancewithin monkey cortical work, was achieved (Fig. 12).
To ensure that the recording and sensing capabilities of
the ceramic probe were not lost, the sensor arrays wereattached to the back-side (non-electrode-side) of the probes
using a thin layer of biocompatible super glue (Adhesive
Systems MG 100 USP Class VI; Adhesive Systems, Inc.,
Frankfort, IL). The sensor array was manually aligned andadhered to the ceramic probe. Handling tabs designed into

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Fig. 12. (a) Image of the ceramic cortical shank instrumented with fully
packaged sensor array. (b) Optical micrograph of probe surface illustrating
the electrodes of the sensor array. Microchannel sensing element has beenhighlighted yellow.

the array allowed for precise manual placement. For benchtop
testing, instrumented ceramic probes were mounted toa custom acrylic jig which allowed integration with anautomated probe insertion setup.
E. Benchtop Insertion Experiments
The cortical sensor array was demonstrated on benchtop
by replicating an implantation via probe insertion into
0.5% agarose, a commonly used model for corticaltissue [39], [40], immersed in a 1× phosphate bufferedsaline (PBS) solution to mimic in vivo conditions. Sensor
arrays were filled via immersion with 1× PBS electrolyte
following a short immersion in isopropyl alcohol (a Parylene
wetting step to facilitate diffusion of the polar PBS into
the microchannel). Instrumented ceramic probes werethen inserted to a depth of 3.5 mm into agarose at threedifferent speeds: 0.01 mm/s (slow), 0.03 mm/s (medium),and 0.1 mm/s (fast), held at the maximum displacement forfive minutes, and extracted at the same speed while sensor
array impedances were measured in real time. Impedance wasmeasured by a LabVIEW-interfaced precision LCR meter(1 Vpp sinusoid, 1 kHz) via a multiplexing PCB for multichannel impedance measurement across the seven sensors of
the array.
An external 50 g load cell was used for measurementof normal forces generated during insertion. The probe andload cell were affixed to a motorized micropositioning stage
that controlled insertion speed and depth into the tissue
phantom (Fig. 13a). The motorized stage, LCR meter, and
PCB digital switching all interfaced with LabVIEW to givethe user control of insertion speed and depth as well asa real-time image (graphical) of the impedance changes(interfacial forces) across the ceramic probe (Fig. 13b).
Two insertion experiments were performed. In both insertion
experiments, some of the observed normalized impedancevalues (and thus forces) greatly exceeded the calibrated linear
range of the sensor. Possible sources for the observed are:(1) observing force ranges much higher than expected and(2) differing displacement profiles (i.e. non-normal to
top surface) of the sensor from what was characterized. Slightchanges in electrolyte conductivity from agarose gel filling
into the channel through the fluidic ports during insertion
may occur, but may be unlikely as diffusion from a gel into
solution would occur at a longer time scale than mechanical1541

deformation during insertion at these speeds. Despite thishowever, the percentage change in impedance data suggests
a trend in interfacial pressures experienced as a function of
insertion speed.
Noting that sensors 1-7 begin at the shank tip and move toits base, data indicated that a majority of the interfacial forces
during insertion are within the first millimeter of the electrodeshank (the first sensor). This is also evidenced by the results
across all three insertion speeds; the maximum percentagechange in impedance of the first sensor is considerably
higher than that of the remaining sensors higher up on theshank (Table II). We attribute the interfacial forces experienced
at the tip during insertion to the tissue displacement andpropagation of the electrode shank track generated by theprobe. Additional force is imposed by tissue as it is being
displaced during insertion.
By varying insertion speed, we observed that interfacial
forces at the shank tip (first millimeter) decrease as the
speed increases. Slower insertion speeds allow for a longer
travel time through tissue, increasing shank-tissue adhesion
and thus the frictional forces between the shank and tissue.
Results confirm that faster speeds decrease the interfacial
forces experienced at the tip (by tissue displacement andfrictional forces) [1] as the maximum impedance change islower by nearly 60%.
This effect of speed on interfacial forces is consistentthroughout the length of the shank. At the slowest speed,
increased frictional forces resulting in additional arraydeformation during insertion give rise to higher average
impedance change in other sensors; Sensors 2-7 recorded
an average maximum impedance change of 12% compared
to 2% and 1% at the medium and fast speeds, respectively.
Therefore, faster speeds also reduce adhesion forces between
the shank and surrounding tissue along the length of the tip.
VI. I NSERTION M ECHANICS C ONSIDERATIONSBenchtop insertion data indicated that interfacial forceswere concentrated at the tip. As these interfacial forces area result of tissue strain during insertion, we would expect tofind a similar trend in vivo with high tissue strains localized at
the tip. This is supported by modeling [8], [10] andin vitro [13] efforts that assess tissue strains during theinsertion of stiff probes. The probe track is often compared
to a mode 1, or opening mode, of a crack, which producesa triangular profile that lends to higher interfacial forces at
the tip. These results are further confirmed by an ex vivo study
by Bjornsson et al [11], where probe insertions establishedtwice the tissue strain at the tip compared to the base of theprobe. Even within in vivo experiments [41], [42], increased
glial response and reduced neuronal presence was observedat the tip of the probe, most likely due to the relatively
high forces and tissues strains that can cause additional
damage. Our work and those in literature suggest that probe
designs should have electrodes concentrated away from thetip to minimize recording zones within these highly damaged
regions.
In addition to probe design, another important factor that
influences tissue damage is the speed of insertion. The precise

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JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 24, NO. 5, OCTOBER 2015

Fig. 13. (a) Schematic of benchtop insertion testing setup. (b) Screenshot image of user GUI during benchtop testing: (Left) a video stream of the insertion
process using a USB microscope. (Right) heat map of GUI to show the distribution of force magnitudes along the length of the probe. Images illustrate theprogression of the instrumented probe from (i) before insertion, (ii) only tip inserted, (iii) half of probe inserted.
TABLE IITABLE OF O BSERVED D ISTRIBUTIONS OF I MPEDANCES A LONG THE L ENGTH OF THE P ROBE FOR T WO D IFFERENT I NSERTION E XPERIMENTS

effect of speed on tissue damage is still debated, but in general,slower speeds can drag and compress the surrounding
tissue and create a path of damage that can extend thekill zone of the implant beyond the probe track as far
as 300 μm away [11], [42]. Faster insertion speeds can
reduce tissue damage by swiftly transecting the tissue andvasculature during implantation, which cause significantly lesstissue compression [42], [43]. In comparing insertion speeds,
our results indicated that faster speeds produced smaller
interfacial forces, which suggests smaller tissue strains (andthus damage) during insertion. This result is further supported
by ex vivo insertion mechanics studies where faster insertion
rates (∼2 mm/s) were found to decrease microvasculaturedamage and tissue strain compared to slower speeds(125-500 μm/s) [11]. Unfortunately, tissue responses
to insertion speeds of 125 and 500 μm/s [11] and10 and 100 μm/s [42] were found to be similar, suggestingthat insertion mechanics are comparable within the rangeof 10-500 μm/s for probes with cross sectional dimensionsof ∼100×100 μm. Probes with larger dimensions (e.g. ∼1 mm
in diameter for deep-brain stimulation electrodes) mightobserve greater differences between speeds.
Insertion speeds greater than 2 mm/s have been known
to compensate for blunt probe design, as observed ina speed of 8.3 m/s used to pneumatically implant thethree-dimensional, 100 site Utah electrode array to avoid
elastic compression of the cortical surface [44]. On theother side of the spectrum, some efforts target very slow
speeds, on the range of neuronal or microglial movement (0.42 μm/s), to displace cells and vasculature without

tearing them, and have also demonstrated improved tissue
immune response [45]. Overall these results suggest thatspeeds (whether extremely fast or slow) are chosen based on
the probe structure and design to minimize tissue damage,and should be considered on an individual probe to probe
basis. It is the hope that additional experiments with thistechnology, in combination with histological data and in vivo
neurophysiological measurements, can help to clarify or
confirm some of these debated topics to realize the promiseof neural prosthetics.
VII. C ONCLUSION
EC-MEMS sensors are straightforward to fabricate
and implement. Here, we have demonstrated the design,
fabrication, characterization, and benchtop implementation of
a Parylene-based force sensor array for studying mechanicalinteractions in neural probes. We demonstrated that annealing
introduces both mechanical and electrochemical changes that
effect sensor performance. We have also presented a method
to attach these thin, low-profile sensors to ceramic corticalprobes and quantitative data on interfacial forces in modelexperiments involving insertion speed. Previous work using
these sensors also validated their capability on measuring
micromotion events in gel phantom models [23]. Force ranges
and sensitivity for a particular EC-MEMS force sensor canalso be tuned by changes in device geometry and thickness
to accommodate interfacial force ranges expected during
insertion. The ease of fabrication, simple wet transduction
mechanism, and low-profile integration opens possibilities foruse of this technology in a variety of in vivo applications.

KIM et al.: PARYLENE-BASED ELECTROCHEMICAL-MEMS FORCE SENSOR

Specifically in looking at insertion mechanics, an analysis
limited to normal insertion forces might suggest that slower
insertion speeds benefit insertion by decreasing the forcesencountered during probe implantation. However, it is also
important to consider the contribution of interfacial forces
prior to reaching a conclusion on insertion parameters. Faster
insertion speeds not only reduce the forces at the shank tip
created during insertion, but also the frictional forces andthus tissue strains observed. Measurement of interfacial forces
elucidates the locations and magnitudes of tissue displacement
and frictional forces during insertion. Instrumentation ofdifferent neural interfaces with sensors such as these can aid
in the further development of more reliable chronic neural
prosthetics as enabling devices.
ACKNOWLEDGMENT
The
authors
would
like
to
acknowledge
Mr. E. Peter Washabaugh VI,
Mr.
Louis
Jug,
and
Ms. Alejandra Gonzales Calle, for invaluable assistancewith device characterization. We would also like to thankDr. Greg Gerhardt of the University of Kentucky for providing
the ceramic cortical probes, and Dr. Donghai Zhu and themembers of the USC Biomedical Microsystems Laboratory
for their assistance.
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Brian J. Kim received the B.S. degree in biomedical
engineering and mathematics from Duke University,
Durham, and the M.S. degree in biomedical engineering from the University of Southern California,
Los Angeles, in 2010 and 2013, respectively,
where he is currently pursuing the Ph.D. degree inbiomedical engineering (Provost’s Ph.D. Fellow)
with the Biomedical Microsystems Laboratory.
His research interests are in the development ofParylene-based microelectromechanical systems
devices for intracortical applications, specificallyon implantable Parylene-based sensor systems. He was a recipient of the
Howard G. Clark Award for Independent Research from Duke University,
the Viterbi School of Engineering Early Ph.D. Fellowship, and the Universityof Southern California Provost’s Ph.D. Fellowship.

Christian A. Gutierrez received the B.S. degrees
in electrical engineering, and business economics
and management, in 2005; the M.S. degree inelectrical engineering from the California Instituteof Technology (Caltech), Pasadena, in 2006; andthe Ph.D. degree in biomedical engineering fromthe University of Southern California (USC),
Los Angeles, in 2011. He conducted research with
the National Science Foundation’s Biomimetic
Microelectronic Systems Engineering ResearchCenter and the USC’s Biomedical Microsystems
Laboratory toward the development of novel impedance-based micro andnanotechnologies for biomedical applications, and focused specifically on
the development of retinal prosthesis technologies to restore sight to theblind. While with the Micromachining Laboratory, Caltech, he worked on
identifying, analyzing, and implementing power harvesting and storage
circuitry for microelectromechanical systems-based power harvesting devices.
He is the Co-Founder and Chief Technology Officer of Fluid Synchrony, LLC,
where he develops high-performance implantable micropump technology for
drug delivery applications. He was a recipient of the Caltech Presidential
Scholarship, the Bill and Melinda Gates Millennium Fellowship, and the
USC Viterbi School of Engineering Kunzel Fellowship.

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 engineeringfrom the California Institute of Technology(Caltech), Pasadena, in 1997, 1998, and 2003,
respectively. She held the Viterbi Early Career Chairwith the Viterbi School of Engineering. She has
been a Professor with the Department of Biomedical
Engineering, University of Southern California,
Los Angeles, since 2004. She currently holds a
joint appointment with the Ming Hsieh Department
of Electrical Engineering. Her research interests include bioMEMS,
implantable biomedical microdevices, microfluidics, multimodality integrated
microsystems, and packaging. She 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 Womenin Science and Engineering Scholarship, the Caltech Alumni Association
Donald S. Clark Award, and the Caltech Special Institute Fellowship. She hasreceived the NSF CAREER and the 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.