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An Electrochemical Microbubble-Based
MEMS Pressure Sensor
Lawrence Yu, Christian A. Gutierrez, and Ellis Meng, Senior Member, IEEE

Abstract— A novel pressure transducer concept is introduced
in which external pressure variations induce changes in thevolume of a trapped microbubble (µB) in an electrolyte solution.
These volumetric changes are monitored by electrochemical
impedance-based measurements and may be used for pressure
tracking applications. Microbubbles are nucleated on-demand byelectrolysis within a confinement chamber with high precision(measured size RSD ∼1%). This pressure transducer concept
was developed specifically for operation in liquid environments and features biocompatible construction, small footprint(<0.1 mm2 ), and low power consumption (<1 nW). This combination of features is ideal for in vivo pressure monitoring
applications.
[2015-0203]
Index
Terms— Pressure measurement,
electrochemical
device, impedance-based sensing, Parylene C, electrochemicalmicroelectromechanical systems (EC-MEMS).

I. I NTRODUCTIONCHRONICALLY
implantabletransducers
enablearound the clock monitoring of localized in vivo
pressures [2]–[4], a useful indicator of patient health ordisease progression. One region where hydrostatic pressure
is highly regulated is within the skull [5], [6]. For thosewith hydrocephalus, this ability to regulate intracranialpressure (ICP) is impaired and potentially life threatening [8].
Treatment involves the implantation of a catheter to shunt
the excess cerebrospinal fluid (CSF) into another region of
the body where it can be safely processed [10], but failuresoccur at a high rate [11], [12] and there is currently no way
to reliably and accurately monitor shunt status noninvasively.
Current clinical practice to interrogate ICP involves lumbar
infusion and tapping into the shunt reservoir, which isaccurate but invasive [13]–[15]. Meanwhile, noninvasive
imaging techniques (ultrasound, MRI) do not provide theaccuracy necessary for clinical use [16]–[19]. Sensors canManuscript received July 17, 2015; revised October 1, 2015; accepted
November 4, 2015. Date of publication November 22, 2015; date of current
version February 1, 2016. This work was supported by the National Science
Foundation through the Division of Electrical, Communications, and Cyber
Systems under Grant ECCS-1231994. Subject Editor D. Devoe.
L. Yu is with the Biomedical Engineering Department, University ofSouthern California, Los Angeles, CA 90089 USA (e-mail: lawrence.
yu@usc.edu).
C. A. Gutierrez was with the University of Southern California,
Los Angeles, CA 90089 USA. He is now with Google Inc., Mountain View,
CA 94035 USA (e-mail: chris.a.gutierrez@gmail.com).
E. Meng is with the Biomedical Engineering and Electrical Engineering
Department, 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.2015.2499167

Fig. 1.
Cartoon of commonly utilized deflecting diaphragm basedtransduction mechanism and several possible failure modes.

be implanted during the shunting surgery and provide
the promise of accurate, noninvasive monitoring throughtelemetry. These microfabricated sensors typically comprise
an air cavity sealed with a thin, flexible membrane that
deflects upon application of external pressure [20]. Examplesof this transduction principle were implemented in early
MEMS (microelectromechanical systems) sensors [21] andcan be found today in a FDA approved implanted pressuretransducer (CardioMEMS, St. Jude Medical, MN, USA).
Despite advances in sensor development, the inhospitablein vivo environment remains a formidable challenge,
where the corrosive aqueous environment and foreign body
response [22] can cause eventual degradation of hermetic
packaging or obscure transducing elements and induce sensor
drift (Figure 1).
To address these issues, alternative approaches have been
investigated involving the use of trapped gases in liquid forpressure sensing. Instead of measuring membrane deflection,
the geometry of the liquid-gas boundary is observed as
it changes with pressure [23]–[26]. This sensing principle
harnesses the liquid environment, does not require hermetic
packaging, and is immune to biofouling. This work utilizes
precisely generated single microbubbles that are localized
within a microfabricated confinement structure to track pressure by measuring their pressure-responsive size changes using
electrochemical impedance. The exclusive use of polymer
materials has the additional benefits of reduced complexity,
mechanical flexibility, and streamlined integration for othermedical applications.
II. D ESIGN
The microbubble pressure transducer (μBPT) utilizes apair of platinum electrodes embedded within a Parylene C
microstructure to first electrolytically generate a microbubble
and then subsequently measure its size and dissolution rate through the use of electrochemical impedance

1057-7157 © 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.

YU et al.: ELECTROCHEMICAL MICROBUBBLE-BASED MEMS PRESSURE SENSOR

Fig. 2. Cartoon of device operation, illustrating top down view of Parylene
microstructure filled in electrolyte. (a) Nucleation of microbubble via
electrolysis. (b) Impedance measurement of microbubble during dissolution.
(c) 3D rendering of μBPT, illustrating confinement of microbubble.

measurement (Figure 2). Parylene (USP class VI polymer andinert) and platinum construction is highly biocompatible [22],serves as an excellent electrical insulator, and possesses greatmechanical strength [27], [28].
Bubble nucleation is commonly implemented through aphase change with a heating element (boiling) [29] orthrough a Faradaic charge transfer reaction (electrolysis) [30].
Considering restrictions on power and local heat generation
for biological implants, electrolysis was selected for thisapplication. Detection of microbubble size has been implemented through a variety of methods (i.e. ultrasonic, optical);the electrochemical impedance sensing modality was selected
because it can be performed using the same electrode pair thatgenerates the microbubble.
Microbubbles in contact with a hydrophobic surface (suchas that of Parylene C) are in an energetically favorable
state and resist detachment [31]. Device orientation becomes
an important consideration if buoyancy forces acting upon
microbubbles are sufficiently large to carry them away fromsensing electrodes. Postural changes occur frequently andhave a noted effect on ICP, so sensor orientation dependence
must be minimized. A chamber is also necessary to create
a localized environment for the saturation of electrolyticallygenerated gas. The ability to isolate and control the conditions in this environment allows for consistent electrochemicalmeasurements. Thus, a confinement chamber was designed to
localize microbubbles while maintaining an open connection
to the surrounding liquid environment.
Despite the widespread use of electrochemical detection
techniques in biological analysis systems, examples ofelectrochemical (EC) based sensing technologies used formeasurement of physical quantities such as force andpressure has been limited. The potential of impedance based145transduction of pressure using a fluid filled sensor was firstdescribed by Ateya et al. through the use of bubbles formedin silicon microchannels encapsulated in polydimethylsiloxane (PDMS) [32]. However, the device was confined to a rigid
substrate with bulky fluidic interconnects and not suitable forin vivo implantation. The open-sensor design and all Parylene
construction of this approach simplifies implementation andobviates requirements of hermeticity. To our knowledge, thisapproach is the first application of such a sensing modality to
microfabricated polymer-based transducers.
The performance metrics of existing in vivo pressuretransducer technologies can be used to guide sensor design.
Physiologically relevant pressures in hydrocephalus patientsfalls between 6.6 to 13.3 kPa (commonly reported units:−50 to 100 mmHg) and are taken as a gauge measurementwith respect to atmospheric pressure with a supine patient. The
neuroanatomy of the central nervous system (brain and spine)
is bathed in CSF, an incompressible liquid. The addition of a
microbubble (10 orders of magnitude smaller in volume compared to cranial compartment) can be used to intermittently
track the pressure within the cranial compartment.
III. T HEORETICAL C ONSIDERATIONS
A. Electrolytic Microbubble Nucleation
Faraday’s laws of electrolysis described the relationshipbetween the mass of a substance altered at an electrodeduring electrolysis and the quantity of electricity delivered tothe electrode. With the application of a constant current forelectrolysis, the relationship can be described as [33]:
I te
(1)
n=
Fz
where n represents the amount of substance altered(in moles), I is the magnitude of the applied current, te is theduration of the current application, F is the Faraday constant,and z the valence number of the substance.
The liquid medium contained in the confinement structure
must be an electrolyte in order to transduce pressure asa function of the size of the bubble which affects thevolumetric conduction path. Common examples foundin vivo are CSF, blood, and tears. Within phosphate bufferedsaline (PBS), a commonly used electrolyte to mimic theconditions found in vivo, the predominant electrolysis
reactions are as follows [34]:
2H2 O + 2e− → H2 (g) + 2O H − , E° = −0.83V
2Cl − → Cl2 (g) + 2e− , E° = +1.36V(2)Bubble nucleation results in ion concentration andpH gradients; typical electrolysis events generate
subnanomolar (∼10−11) concentration gradients that canbe readily neutralized by buffering.
B. Microbubble Dynamics
The fluid dynamics of a spherical bubble suspended in
an infinitely large body of liquid are described by theRayleigh-Plesset equation [24]:


4μ
1
2σ
3 2
−
ṙ
(3)
r r̈ + ṙ =p g − p∞ −
2
ρd
r
r

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where r represents the bubble radius as a function of time
(dots denote time derivatives); pg and p∞ are the pressure of
the gas within the bubble boundary and the external pressure
taken at a distance infinitely far away from the bubble, respectively; σ denotes surface tension; μ represents fluid viscosity;and ρd is fluid density. For pressure oscillation frequencies
significantly lower than the bubble resonant frequency, thepressure-radius relationship can be described as [24], [35]: r b 2σ0−p∞ = pi(4)rrWith the polytropic exponent b ≈ 1, the equation resembles
the Young-Laplace equation and describes the balance offorces at the gas-liquid interface due to surface tension andexternal pressure. For CSF, the surface tension is relatively
constant and homogeneous, and thus the bubble size can bedirectly correlated to the external pressure. During dissolution,hydrostatic pressure exerts forces normal to the liquid-gas
boundary, where the geometry is dictated by the shape ofthe microbubble confinement chamber. Within the confines ofa hydrophobic Parylene [22] microstructure, the gas exhibits
nonlinear dissolution characteristics due to effects of thesurface chemistry on surface tension [36].
At larger time scales, the gas within the bubble will dissolve
into the surrounding under-saturated liquid, as governed by
Henry’s law:c∞ = apg

TABLE ID IFFUSION C OEFFICIENTS OF G AS -L IQUID C OMBINATIONS

Fig. 3.

Simplified Randles circuit model of electrodes in electrolyte.

Fig. 4.

Electrochemical measurement of bubble size in electrolyte.

(5)

where c∞ is the concentration of the dissolved gas in thebulk liquid, a is a temperature dependent constant (also known
as Henry’s law constant) characteristic of the particular gasliquid combination, and pg is the partial pressure of the gas
acting on the liquid surface (i.e. the gas within the bubble).
For a fixed external pressure (e.g. the regulated ICP of anormal patient), the gas concentration c must satisfy (5) for
the gas within the bubble to remain in equilibrium with thesurrounding liquid, else the bubble will grow or shrink in size.
The effects of gas dissolution on bubble size was describedby Epstein and Plesset [37]:k(cs − c∞ ) 1
1
dr
=
+
(6)
dt
ρ∞ + 2τ/3r r(πkt)1/2
where k is the diffusion coefficient (see TABLE 1 for examples) of the specific gas-liquid combination, cs is the saturation
concentration of gas in the liquid at the bubble surface,
ρ∞ is the density of the gas within the bubble, and τ is themodified surface tension. The primary product in the electrolysis of PBS as described in (2) is H2 , thus we expect the
diffusion coefficient to be near that of the H2 -H2 O combination (TABLE 1). This diffusion process is limited by masstransfer and is relatively slow; the typical dissolution rate of
bubbles with radius 100 μm are estimated to be on the orderof ∼0.1 μm/s.
C. Electrochemical Sensing
The electrochemical impedance of an electrolyte as measured by immersed microelectrodes can be represented by
the well-known Randles circuit model (Figure 3) [38].

Impedance is measured using an alternating current signal atsufficiently high frequencies such that the solution resistancedominates the impedance response. The complex impedanceis reduced to its real component and can be represented using
the DC model, where resistance is inversely proportional to thecross-sectional area between electrodes, as described by thefollowing:|Z | f ≥10k H z ≈ Rs =

ρe l
ST ot

(7)

where Z is the complex impedance measured by theelectrodes, Rs is the solution resistance, ρε is the resistivityof the electrolyte, l the distance between electrodes, andST ot the effective cross sectional area of the ionic conduction
path. A suspended microbubble behaves as a nonconductive
void within the conductive electrolyte (Figure 4), altering thecross-sectional area and thus the overall solution resistance.
This phenomenon enables the transduction and measurementof hydrostatic pressure acting upon microbubble size. The sensitivity of the device is dependent on electrode arrangement,
sensor dimensions, and the volume of electrolyte subject toionic current flow.
The geometry of the gas-liquid interface and ionic
conduction path is largely dictated by the microbubble confinement structure. Rectangular and cylindrical confinement

YU et al.: ELECTROCHEMICAL MICROBUBBLE-BASED MEMS PRESSURE SENSOR

Fig. 5.

147

Cross section of bubble in rectangular fluid channel.
Fig. 7.
Micrograph of fabricated devices, with various microchannel
confinement structures to enable testing of various electrolyte parameters.

Fig. 6. Cartoon of simplified fabrication process; steps depict cross section
view. (a) Deposition of Parylene and Pt lift-off. (b) Patterning of sacrificial
photoresist, deposition of Parylene, and fluidic interface port etch. (c) Releasedevice from silicon substrate and soak in electrolyte to fill confinement
structure.

chambers were fabricated and evaluated [39]. The geometryof a microbubble confined within a cylindrical chamber isradially symmetric and allows for orientation independent
operation. When a microbubble is trapped within a chamberwith rectangular cross section, the gas tends to fill the middleof the cross sectional area and leave pockets of liquid at thecorners (Figure 5). The cross section profile does not change
significantly as the bubble changes in size, while the length ofthe conduction path changes. As described by Ateya et al. [32],the area of this cross section is described as follows:


cos θ
r2
(8)
cos (π/4 + θ ) − (π/4 − θ )ST ot = 4 A 2
R sin π/4
where A is the total cross sectional area of the channel,
R is the inscribed inner radius of the channel, r the radiusof curvature of the corner meniscus, and θ the contact angleof the microbubble with the chamber wall. This type ofionic conduction path improves sensitivity by leveraging largechanges in impedance observed in the submicrometer sizedliquid capillaries found at the corners.
IV. FABRICATION
The device was fabricated using surface micromachining
techniques on a polymer substrate with thin film electrodes, a process based on previously reported techniques(Figure 6) [40]. The entire fabrication process was performedat relatively low temperatures (<90°C) to prevent thermal
degradation of Parylene.
A layer of Parylene (12 μm, Specialty Coating Systems,
Indianapolis, IN) was deposited onto a silicon wafer, which
serves as a temporary support substrate. Platinum wasdeposited (2000 Å thick) via an electron beam evaporator
and features were defined using a lift-off process to form
electrodes and traces. An additional 12 μm of Parylene wasdeposited to serve as an insulation layer. To expose theelectrode active areas, the insulation area was removed by theuse of switched-chemistry deep reactive ion etching (DRIE)
in oxygen plasma [41] with a dual layer photoresist mask
(15 μm, AZ4620; AZ Electronic Materials, Branchburg, NJ).
Next, photoresist was patterned (10 μm, AZ4620) to serve asa sacrificial mold for the microbubble confinement structure.
A final 4 μm thick layer of Parylene was deposited to
enclose the sacrificial photoresist structure, and access ports
were patterned and opened with oxygen plasma reactive ion
etching (RIE). Another switched-chemistry DRIE process was
performed to etch down to the underlying silicon wafer,forming the die outlines for device release. The remainderof the etch mask and sacrificial structure PR were strippedby soaking the wafer in an acetone bath. By submerging thewafer in water, free film devices were released and separated
from the underlying silicon substrate (Figure 7). It is worthy tonote that the confinement structures did not experience stiction
upon drying.
A robust, epoxyless, reversible electrical connection scheme
was established [42], minimizing handling of the thin Parylene
devices. A hinge-based zero insertion force (ZIF) connector
for flat flexible ribbon cables was utilized (Hirose ElectricCo., Simi Valley, CA), enabling a secure connection to beestablished with the exposed contact pads (8 or 12 channel
with 0.5 mm pitch). Because the connector mandates a cable
thickness of 300 μm, the free film Parylene devices were
affixed to a 0.01 inch (∼250 μm) thick polyetheretherketone (PEEK) sheet with a cyanoacrylate based adhesive(All Purpose Krazy Glue, Westerville, OH) and cut down
to fit into the connector. The ZIF connector was directly
soldered onto a commercially available polymer flexible cable(Molex Inc., Lisle, IL). The use of this thicker, more robust
polymer cable results in a secure, repeatable connection
to a measurement printed circuit board (PCB) via another
ZIF connector.
V. E XPERIMENTAL M ETHODS
Multiplexing circuitry (Analog Devices ADG1206;
Norwood, MA) was utilized to control device connections to
instrumentation and facilitate usage of multiple devices on
the same die. Because current injection during electrolysis
disrupts impedance measurement, device connections weretime division multiplexed between the current source(Keithley Model 2400 Sourcemeter; Keithley Instruments
Inc., Cleveland, OH) and LCR meter (Keysight E4980A,
Keysight [formerly Agilent] Technologies, Santa Rosa, CA).

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Fig. 9.
Electrochemical impedance spectroscopy of sensor response.
Arrow indicates measurement frequency, where system is most resistive(minimum phase).
Fig. 8. Schematic overview of test setup, illustrating electrical and fluidic
connections.

An acrylic jig with an airtight chamber and fluidic connections was custom fabricated for benchtop characterizationof the sensors. Once the devices were mounted, the chamber
was purged of air and filled with an electrolyte (e.g. 1× PBS,
artificial cerebrospinal fluid). The jig was connected to acustom built, regulated (ControlAir T900X; ControlAir Inc.,
Amherst, NH) nitrogen pressure source using PEEK tubing(1/16 in. O.D., 0.055 in. I.D.) prefilled with electrolyte.
A commercially available sensor (Omega PX480A) was placedin line to measure the applied pressure during testing.
Multiplexing, pressure regulation, and data acquisition
were computer controlled through a LabVIEW based graphical user interface (Figure 8). The devices were placed
under a stereomicroscope to provide visual confirmation ofdevice wetting and microbubble generation during benchtop
testing. Operation of the sensor can be isolated into separate aspects (electrochemical impedance measurement andelectrolytic bubble generation), and each was characterizedindividually. Afterwards, mechanical response to pressure was
evaluated.
VI. R ESULTS
A. Electrochemical Impedance Measurement
Electrochemical impedance spectroscopy (EIS, 20 Hz1 MHz) was used to verify and assess electrode surface
properties in Fig. 9. Measurement was conducted between
adjacent electrodes with a Gamry Reference 600 potentiostat(Gamry Instruments, Warminster, PA). The optimal frequencyof measurement ( f measurement ) where the complex impedance
response is dominated by solution resistance, was found tobe where the phase response was nearest 0°, corresponding
to an excitation frequency of 10 kHz for 1× PBS (phosphatebuffered saline). For this electrochemical impedance measurement, power draw was ∼1 nW (∼1 nA, 1 Vp-p ). Electrochemical impedance measurements are sensitive to parameterssuch as temperature and concentration of ionic species inthe solution. However, both are relatively constant within the
cerebrospinal fluid [43], [44].

Fig. 10.
Representative measurement of electrolytically generated
microbubble, with indicators for parameters of interest. (a) Impedance of
microbubble immediately following electrolysis. (b) Bubble begins detachment from microchannel. (c) Bubble shrinks rapidly in presence of platinum
electrode. Points (a), (b), and (c) correspond to the micrographs in Figure 2.

B. Electrolytic Bubble Generation
Microbubbles were electrolytically generated by applying
direct current pulses (1–5 μA for 2–5 seconds). Impedancemonitoring during μB nucleation enabled precise control of
bubble formation [45] and was used to determine the overallcharge injection during electrolysis. Optical microscopy was
also used to confirm nucleation and containment of the bubble
within the microchannel (Figure 11). Electrolysis current wasmaintained under 10 μA to avoid delamination of the metalParylene interface and for longevity of device [46]. Observed
power draw was <60 μW for electrolytic bubble generation.
Following application of current, electrochemical impedance
measurement was used to verify and quantify the growth ofthe microbubble.
Electrochemical impedance was recorded prior to (∼10 s)
and following electrolysis (Figure 10). The recorded
impedance prior to nucleation served as a baseline. Immediately following electrolysis, the sharp increase in impedance
observed is the result of the nonconductive gas displacing
the electrolyte solution. The gases within the microbubble

YU et al.: ELECTROCHEMICAL MICROBUBBLE-BASED MEMS PRESSURE SENSOR

149

Fig. 11.
Micrograph of microbubble dissolution within microchannel.
(a) Microbubble immediately following nucleation. (b) Liquid-gas boundary
adheres to sidewall of microchannel. (c) Following detachment, microbubble
volume diminishes quickly.

gradually diffuse back into solution, thus lowering themeasured impedance until its return to baseline. Data was
acquired at 5 Hz and was terminated once impedance returned
to within 1% of the initial baseline impedance value. Several
features of interest (as graphically depicted in Figure 10)
were extracted from the temporal impedance response usinga MATLAB script. Baseline was defined as the average
impedance prior to current injection [], Z is the initial
impedance response referenced to baseline [], Rdissolution was
defined as the slope within between the (a) and (b) regionsin the impedance response as depicted in Figure 10. The
dissolution time τdissolution was defined as the time elapsed
from the current injection to the point when the impedance
returns to 5% of Z.
The presence of local maxima (e.g. around 200 s
in Figure 10) can be explained by interactions of the bubble
with the Parylene surface. During dissolution, the liquid-air
interface adheres and appears to “stick” to the sidewalls of the confinement chamber due to interactions withthe surface, leading to increased measured electrochemical impedance [32]. The impedance decreases more rapidly
once the microbubble detaches from the channel sidewall(Figure 11b). Several different channel geometries were tested,each of which featured characteristic local minima and maxima in the impedance response due to the differences inchannel width which impacts microbubble shape. Dissolution
was defined in the region where the microbubble was incontact with the channel sidewalls, the configuration in which
the impedance measurement is most sensitive.
C. Priming of Electrochemical Environment
To maintain consistency between bubble nucleation events,
the local electrolyte environment was saturated with dissolved
gas by repeatedly injecting current to induce electrolysis.
Current injections were performed (Figure 12) and theimpedance measurements were taken until τdissolution was consistent to within 10% (Figure 13). This priming process lasted
approximately 50 minutes and only needs to be performed
once. The Z and τdissolution (as illustrated by Figure 10)
from a primed environment impedance response curve

Fig. 12. Successive nucleation events and tracking of microbubble dissolution
to establish gas saturation state.

Fig. 13. Tracking stabilization and dissolution time with successive bubble
nucleation events.

(i.e. the 5th to 10th current injections in Figure 12) werethen analyzed for pressure transduction. Notably, the size ofthe nucleated bubble remained consistent (∼1% RSD) while
attempting to saturate the local dissolved gas environment, asseen in the low variance of Z (mean = 37.9 k, standard
deviation 574.6 , n = 10).
Note that depressurization of the local environment isanother commonly used method to saturate the liquid electrolyte, as utilized by previous studies [47], [48]. However, thismethod is incompatible with the eventual in vivo applicationand thus was not used for device characterization.
D. Pressure Transduction
For a given constant pressure environment and set electrolysis parameters, impedance tracking of microbubble size during
dissolution was performed (Figure 14). The Z metric wasobserved to have an inversely proportional relationship ( Z =60496+(1.234E6/(pressure (mmHg)+148.85)), R2 = 0.999)
to a randomly applied pressure (Figure 15), suggesting thatthe impedance measurement is a direct measurement of the
bubble volume relationship as described by the Young-Laplace

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linear (dissolution rate = 100.70 + 0.669∗pressure (mmHg),
R2 = 0.998), as confirmed by Henry’s law (5) and the
Epstein-Plesset (6) equations. Provided that the electrochemical measurement is properly primed, this metric provides amethod to check the Z measurement on the same nucleatedmicrobubble. Using this metric for pressure transduction, theestimated resolution is approximately ±20 mmHg.
The range of this sensing modality is dependent on anumber of factors and is currently still under investigation. As mentioned in the theoretical considerations section,
the amount of injected charge and the dimensions of theconfinement chamber directly influence the volume and lengthof microbubbles. Preliminary pressure measurements at aphysiologically relevant range (0-20 mmHg) were achievedby utilizing smaller microbubbles and smaller confinement
structures.
Fig. 14. Measurement of electrolytically generated microbubble and dissolution at various pressures.

VII. C ONCLUSION
Successful development of a unique sensing mechanism forhydrostatic pressure measurement utilizing impedance based
measurements of electrolytically generated microbubbles has
been demonstrated. Electrochemical impedance measurements
within Parylene-based microfluidic structures were possible
and demonstrate the feasibility of intracranial pressure measurement. Combined with the exceptional biocompatibility
characteristics of the open-to-liquid packaging, this robust
sensing mechanism can be used reliably for chronic measurements. Implementation within a completely flexible Parylene
substrate has the potential to address the unmet need forreliable, chronic transducers for the monitoring of hydrocephalus treatment. Further calibration of electrolyte saturation
characteristics enables usage in a multitude of wet sensing
applications.

Fig. 15. Initial impedance response
size, in relation to pressure.

Z, a direct correlation of microbubble

ACKNOWLEDGMENT
The authors would like to thank Mr. Jeff Nishida,
Mr. Aduragbemi Jibodu, and the members of the USC
Biomedical Microsystems Laboratory for their assistance.
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Fig. 16.

Dissolution rate in relation to hydrostatic pressure.

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Lawrence Yu received the B.S. degree in biomedical
engineering and the M.S. degree in electricalengineering from the University of Southern
California (USC), Los Angeles, in 2009 and 2010,
respectively, where he is currently pursuing the
Ph.D. degree in biomedical engineering with the
USC Biomedical Microsystems Laboratory. He iscurrently working on the development of a sensor
platform for in vivo applications. He was a recipientof the Walt Disney Foundation Scholarship and theRose Hills Foundation Fellowship.

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JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 25, NO. 1, FEBRUARY 2016

Christian A. Gutierrez received the B.S. degrees inelectrical engineering and business economics andmanagement, in 2005; the M.S. degree in electricalengineering from the California Institute of Technology (Caltech), Pasadena, in 2006; and the Ph.D.
degree in biomedical engineering from the University of Southern California, Los Angeles, in 2011.
He conducted research with the National Science
Foundation’s Biomimetic Microelectronic Systems
Engineering Research Center and the USC Biomedical Microsystems Laboratory toward the development of novel impedance-based micro- and nanotechnologies for biomedical
applications and focused specifically on the development of retinal prosthesis
technologies to restore sight to the blind. While at the Caltech Micromachining
Laboratory, 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
with Fluid Synchrony LLC, which develops high performance implantablemicropump 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 engineering fromthe California Institute of Technology (Caltech),
Pasadena, in 1997, 1998 and 2003, respectively. She
is currently a Professor and the Chair of the Department of Biomedical Engineering with the Universityof Southern California, Los Angeles, where she has
been since 2004. She also 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 held the Viterbi Early Career Chair with
the Viterbi School of Engineering. She is a Member of Tau Beta Pi, theBiomedical Engineering Society, the Society of Women Engineers, and theAmerican 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, and the 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.