PARYLENE-BASED ELECTROCHEMICAL-MEMS FORCE SENSORARRAY FOR ASSESSING NEURAL PROBE INSERTION MECHANICS
Brian J. Kim1, Christian A. Gutierrez1, Greg A. Gerhardt2, and Ellis Meng1
1
University of Southern California, Los Angeles, CA, USA
2
University of Kentucky Chandler School of Medicine, Lexington, KY, USAABSTRACTDESIGN AND OPERATIONWe present the first use of a Parylene-based
electrochemical-MEMS (EC-MEMS) sensor array for
instrumentation of ceramic-based neural electrode probes.
The sensor array consists of a liquid-filled Parylene-based
microchannel and an array of enclosed electrodes thatmonitor local variations in impedance during mechanicaldeformation of the channel. The array provides real timemeasurement of out-of-plane interfacial forces produced
directly on the electrode shank surface (<5 mm2) during
insertion of the probe. We demonstrate the ability toexamine the relative force distribution of interfacial forces
produced on the shank surface during insertion, thereby
providing a clearer understanding of probe insertion
mechanics. Our approach enables, for the first time,robust mechanical instrumentation of electrode shanks
providing a means for assessing the poorly understoodinterfacial mechanics between neural probes and tissue.

The linear array of seven sensors is constructed on aflexible Parylene C substrate and within a Parylene C
microchannel (100 μm width, 4.5 mm length) that serves
as a contact surface. Platinum microelectrodes (2000 Åthick) exposed to the fluidic contents of the channelinterior provide an electrical interface for electrochemical
impedance measurements; pairs of adjacent electrodes
form sensor units (Figure 1). Etched fluidic ports atchannel ends allow the surrounding electrolyte to fill thechannel. Fluid can flow freely between the externalenvironment and the channel via the access ports.

INTRODUCTION
Current designs of implantable neural probes forneuronal recording and stimulation largely ignore
mechanical considerations, which may contribute tolimited performance and lifetime in vivo. Understanding
the link between insertion-induced mechanical damageand electrode shank design is necessary to guide thedesign of next generation probes, such that tissue damage
and the related immunological responses are minimized[1]. Instrumentation of the electrode surface with tiny
physical sensors captures important mechanicalinteractions between the probe and tissue that impact
probe shape, material, and insertion parameters.
Previously, external load cells were used to measuremacroscale insertion forces [2], but limited insight into
interfacial phenomena was gained and out-of-planestresses acting on the shank surface could not be detected.
Recently, CMOS sensors integrated on a silicon shankwere used to measure interfacial stresses [3]. This design,
however, was limited to measurement of stresses in thebulk material and only compatible with silicon-basedneural probes. A more versatile approach is necessary toenable comparative studies between different electrodedesigns and materials.
Sensors must be able to:accommodate small shank areas, integrate with CMOSincompatible electrode shank materials, and operate inaqueous and high salt environments. To address thisunmet need, we adapted our Parylene-based EC-MEMS
sensing technology to transduce force and pressure
distributions directly on the electrode shank surface [4].
We present a new sensor layout capable of integration
directly onto existing implantable ceramic electrodeshanks [5] and demonstrate, for the first time, the relative
force distribution produced along the shank surface inagarose tissue phantoms.

978-1-4673-0325-5/12/$31.00 ©2012 IEEE

Figure 1: Optical micrograph of Parylene force sensor
array (100 μm width, 4.5 mm length, and 20 μm height).
Magnified view illustrates the layout of the thin film
platinum electrodes and the position of the fluidic ports.

Figure 2: Illustration depicting transduction method of
the sensor array. Prior to insertion, the ionic currentremains undisturbed between electrodes. Insertion intoagarose (tissue phantom) results in interfacial forces that
deform the channel and changes sensor impedance.

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MEMS 2012, Paris, FRANCE, 29 January - 2 February 2012

During insertion, external contact forces deform thecompliant fluid-filled structure in the normal direction andredistribute the fluid contained within the chamber.
Changes to the ionic conduction path of current-carrying
ions in the fluid register as a change in the magnitude of
the solution impedance (Figure 2). Thus, impedance
variations can be correlated to mechanical interfacial
contact forces exerted by the tissue in contact with the topof the sensor array.

end to allow for proper connection with a custom ZIF-ZIF
adaptor connector. The adaptor facilitated connection to acommercially available flat flexible cable that enabled
integration with the rest of the test system.

EXPERIMENTAL METHODS
Fabrication
Standard surface micromachining processes wereused for sensor array fabrication (Figure 3).
Allfabrication processes were performed at low temperatures(90 ºC) to prevent excessive thermal cycling of theParylene C structural material. Parylene C was chosen asthe sensor material for its biocompatibility, mechanicalstrength, electrical insulation properties, and ability toform pinhole free films via chemical vapor deposition atroom temperature.
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 forelectrodes and contact pads were etched using oxygen
plasma. A 20 μm high sacrificial photoresist layer was
spun-on to form the structure for the microchannel
sensing element, and a 4 μm thick layer of Parylene wasdeposited to form the final device structure. Fluidic
access ports at the channel ends were opened by oxygen
plasma. Finally, to facilitate precise shaping of the arrayfor attachment to a fine ceramic shank tip, arrays were cut
out using a switched chemistry (C4F8 and O2 plasma)etching process [6].

Figure 4: Released sensor array prior to attachment toshank. Integrated contact pads allow simple, solder-less,and repeatable connections via a ZIF connector.
Packaged arrays were adhered to the ceramicelectrode shank tip using a thin layer of biocompatible
superglue (Adhesive Systems MG 100 USP Class VI)(Figure 5b).

(a)(b)
Figure 5: a) Instrumented ceramic electrode shanks inacrylic jigs for insertion experiments. The white flat
flexible cables connect to a ZIF connector on a custom
PCB. b) Optical micrograph of the sensor array affixed
to two different ceramic electrode shank designs.
Biocompatible adhesive was chosen to minimize
immunological responses to the instrumented shank. The
superglue is soluble in acetone so that sensors can beremoved from the ceramic shanks, allowing reuse ofelectrodes. For in vitro testing, instrumented ceramicelectrode shanks were interfaced to an acrylic test jig,
which allowed attachment to the insertion setup (Figure5a).

Figure 3: Abbreviated process flow for array fabrication
andattachment
utilizingstandard
surfacemicromachining techniques. Cross-section taken laterally
through device microchannel.
Devices were released from the wafer by stripping theprotective photoresist mask for the cut-out etch, and thensubmerging the wafer in DI water. The hydrophobicity of
the Parylene polymer allowed the arrays to easily separate
from the silicon wafer and lift-off the surface. Sacrificial
photoresist was then removed by immersion in acetone
and isopropyl alcohol and arrays were dried (Figure 4).

Experimental Setup
Insertion experiments were conducted into agarose
phantom models of cortical tissue (0.5%) in a bath of 1×
phosphate buffered saline (PBS) to mimic wet, in vivo
conditions. The setup consisted of the acrylic jig attached
to a 50 g load cell, for normal insertion force
measurements, all affixed to a motorized micropositioning
stage that controlled insertion speed and depth into thetissue phantom.
Impedance was measured by aLabVIEW-interfaced precision LCR meter (1 Vp-psinusoid, 10 kHz) via a multiplexing PCB for multichannel impedance measurement across the seven sensorsof the array. Electrochemical impedance spectroscopy(EIS) yielded 10 kHz as the optimum frequency toArray Packaging
Prior to attachment to ceramic shank tips, arrays werefirst electrically packaged to a zero-insertion force (ZIF)
connector (Hirose 8 channel, 0.5 mm pitch) (Figure 5a).
Details of the epoxy-less, ZIF connection scheme forParylene cables are described in our prior work [7].
Briefly; a PEEK stiffener was attached to the contact pad125

measure solution impedance to minimize system phase(Figure 6). The motorized stage, LCR meter, and PCB
digital switching all interfaced with LabVIEW to give theuser control of insertion speed and depth as well as a realtime image (graphical) of the impedance changes(interfacial forces) across the ceramic electrode shank.

Figure 8: Representative plot of insertion experiments(sharp electrode shank design at the 0.01 mm/s insertionspeed). Normal insertion forces are in purple. Plotshows the insertion of instrumented ceramic electrodeshanks into agarose with a 5 minute hold at the maximum
displacement (3.5 mm).

Figure 6: EIS magnitude and phase plots for sensor
electrodes. 10 kHz was chosen to bypass capacitiveeffects at the electrode-electrolyte interface and maximize
solution impedance (phase of ~0°).

Table 1: Summary of agarose insertion experiments of thesharp shank design for three insertion speeds.
Speed(mm/s)

Prior to insertion experiments, sensor arrays werefilled via immersion with 1× PBS electrolyte following ashort immersion in isopropyl alcohol (a Parylene wetting
step to facilitate diffusion of the polar PBS into themicrochannel). Instrumented ceramic electrode shanks
were inserted 3.5 mm (Figure 7) 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
for five minutes, and extracted at the same speed.

0.01
0.030.1Maximum
normal force(mN)

7.99
8.94
15.60

Maximum
ΔZ% :Sensor 1(shank tip)

87 %
38 %
28 %

Averagemaximum ΔZ% change:
Sensors 2-7(n=6, mean ±SE)
12 ± 3%2 ± 1%
1 ± 0.6%

Shank Interfacial Forces
By correlating the magnitude of impedance changes(normalized) to a force magnitude, we are able to obtaincomparisons of the relative magnitudes of forces that areexperienced across the ceramic electrode tip. Initial finite
element modeling results indicate that the force range thatcan be measured with this device geometry is in the rangeof tens of milli-Newtons (up to 60 mN for a near
maximum 18 μm chamber deflection).
Noting that sensors 1-7 begin at the shank tip andmove to its base, data indicate that a majority of theinterfacial forces during insertion are within the firstmillimeter of the electrode shank (the first sensor; Figure8). This is also evidenced by the results across all three
insertion speeds; the maximum percentage impedancechange of the first sensor is considerably higher than thatof the remaining sensors higher up on the shank (Table 1).
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.
In varying insertion speed, we also observed that theinterfacial forces at the shank tip (first millimeter)
decrease as the speed increases. In analyzing thisphenomenon we must also consider the effects offrictional forces exhibited by the electrode shank during
insertion. Slower insertion speeds allow for a longertravel time through tissue, increasing shank-tissueadhesion and thus the frictional forces between the shankand tissue. Results confirm that faster speeds decrease theinterfacial forces experienced at the tip (by tissue

Figure 7: Spacing of sensor array on a sharp ceramicelectrode shank. A displacement of 3.5 mm into agarose
covers up to 6 sensors of the array.

RESULTS AND DISCUSSION
Normal Insertion Forces
Normal neural probe insertion forces were reported
by others, but are briefly mentioned here to provide acomplete picture of insertion mechanics. There are twomain forces that dominate mechanics during probe
insertion: tissue displacement forces enacted by the shankon tissue and frictional forces caused by the shank tip
moving through tissue [1]. Measurements of normal
insertion forces combine these two phenomena into asingle measurement to provide unclear information on
insertion mechanics. We observed that with increasing
insertion speeds, the normal insertion forces increase:with forces nearly doubling with a 10× increase in speed(Table 1). We hypothesize that this is due to the increased
resistance to insertion by the tissue that is encounteredwith increases in insertion speed.

126displacement and frictional forces) [1] as the max
percentage impedance change is lower by nearly 60%(Table 1).
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 percentage 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.
In summary, a mechanics analysis limited to normal
insertion forces might suggest that slower insertion speeds
benefit insertion by decreasing the forces encounteredduring probe implantation.
However, before thisconclusion can be made, it is also important to consider
interfacial forces and the effects of tissue displacement
and frictional forces; faster insertion speeds not only
reduce the forces at the shank tip created during insertion,
but also the frictional forces observed. Measurement ofinterfacial forces elucidates the locations and magnitudesof tissue displacement and frictional forces during
insertion.
Force ranges and sensitivity for a particular ECMEMS force sensor can be tuned by changes in device
geometry and thickness to accommodate interfacial force
ranges expected during insertion. Currently, further
characterization of the sensor array to obtain a calibration
curve between force and impedance change is underway.

CONCLUSION
We designed, fabricated, and tested a Parylene-based
EC-MEMS sensor array for instrumentation of a ceramicelectrode shank. We demonstrated the ability to track therelative distribution of interfacial forces produced along
the shank surface during insertion. We determined thatinterfacial forces primarily act on the first millimeter ofthe shank which imposes a significant portion of the tissue
displacement forces, and that faster insertion speeds
reduce interfacial forces experienced along the shank tip.
We also demonstrated tracking of micromotion in
implanted neural probes.
Following additional
characterization, in vivo studies with instrumented shanks
will be performed to gain a more complete understandingof insertion mechanics.

ACKNOWLEDGEMENTS
This work was funded in part by the Engineering
Research Centers Program of the NSF under Award
Number EEC-0310723, the University of Southern
California Provost Ph.D. Fellowship (BK), and the Billand Melinda Gates Foundation (CG). The authors would
like to thank Dr. D. Zhu and members of the USC
Biomedical Microsystems Laboratory for their assistance.

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Micromotion
Electrode shank instrumentation also provides thebenefit of additional sensing capabilities.
More
specifically, it has been reported that micromotion of thebrain occurs during craniectomies due to pulsed bloodflow and respiration [8]. Micromotion events can affectimplanted probes and result in disrupted recording due topositional changes. We also demonstrate the ability todetect micromotion events such that sensory feedback canbe used to distinguish neural signals from micromotion
events (Figure 9).

Figure 9: Results from simulated micromotion
experiments. Micromotion was produced by displacing
the agarose substrate ±1 mm normal to shank face.
Highlighted region contains 3 micromotion events.

CONTACT*B.J. Kim, tel: +1-213-8213949; brianjk@usc.edu

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