A DUAL MODE MICROBUBBLE PRESSURE AND FLOW SENSOR
Lawrence Yu and Ellis Meng
University of Southern California, Los Angeles, USAABSTRACT

electrochemical measurements. At the same time, an open
connection must be maintained to transduce pressure fromthe surrounding liquid environment to the confined μB.

We present the first microbubble (μB) based dual
mode sensor for static pressure and liquid flow. The
transducer harnesses the inverse pressure-volume relation
andhigh
precision
electrochemical
impedancemeasurement (2% RSD) to monitor size of electrolyticallygenerated μBs. By altering sensor orientation, staticpressure measurement (0-20 mmHg) or flow rate wastransduced (0-1000 µL/min). The design targets chronic invivo monitoring applications and therefore incorporates thefollowing features: non-hermetic packaging, absence ofmoving parts, low power consumption (<5 µW), andbiocompatible construction (Parylene and Pt).

INTRODUCTION
Implantable transducers enable around the clockmonitoring of localized in vivo pressure and fluid flow,which are useful indicators of patient health or diseaseprogression. Sensors may be implanted alongside coronary
stents or hydrocephalus shunts and provide accurate,noninvasive
monitoring
throughtelemetry.
A
microfabricated pressure sensor would typically comprise avacuum cavity sealed with a thin, flexible membrane that
deflects upon application of external pressure. Two (ormore) pressure sensors may be used in conjunction to infer
flow information on the basis of Bernoulli’s principle.
However, in vivo differential pressure flow metering ischallenged by the body’s corrosive environment and theconcomitant compromises in performance from hermetic
sealing of the sensing elements [1]. To address these issues,pressure transduction using microbubbles was investigatedwith an open package device in direct contact with theelectrolyte, but its use was limited to static conditions
without flow [2]. This work improves pressure
measurement resolution (5× increase), further elucidates
μB nucleation parameters necessary for pressuretransduction within a flowing liquid, and introduces a twoport structure for measuring pressure differential (Fig. 1,2). This marks the first demonstration of μBs as pressuretransducers for measurement of liquid flow and for dual
mode sensing of both pressure and flow.

Figure 1: 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 device, illustrating
confinement of microbubble.
Impedance based transduction of pressure using a fluid
filled sensor was first described by Ateya et al. through theuse of electrolytic bubbles formed in silicon microchannels
encapsulated in polydimethylsiloxane (PDMS) [5].
However, the device was limited to a rigid substrate withbulky fluidic interconnects and not suitable for in vivo
implantation. The open-sensor design and all Parylene
construction of this approach simplifies implementationand obviates requirements of hermeticity, greatly reducing
the footprint. To our knowledge, this approach is the first
application of such a sensing modality to microfabricatedpolymer-based transducers.

DESIGNThe microbubble pressure transducer (µBPT) utilizes
a pair of platinum electrodes in an open Parylene C
microchamber to first electrolytically generate amicrobubble and then subsequently monitor its size anddissolution rate through the use of electrochemical
impedance measurement (Fig. 1). Parylene (USP class VI
polymer and inert) is highly biocompatible [3], serves as
an excellent electrical insulator, and possesses greatmechanical strength [4].
A confinement chamber is necessary to create alocalized environment for the saturation of electrolyticallygenerated gas. The ability to isolate the μB and control theenvironmental conditions allows for consistent978-1-5090-1973-1/16/$31.00 ©2016 IEEE

Transduction Principle
In order to perform a pressure measurement, first a μB
is formed via electrolysis (Fig. 1). The electricallyinsulating bubble now resides in the measurement channel
region and is available for pressure transduction. A lowpower, high frequency alternating current applied acrossthe EI measurement electrodes permits monitoring of thevolumetric conductive path [6] and is measured as theelectrochemical impedance, Z:

325

MEMS 2016, Shanghai, CHINA, 24-28 January 2016     (1)where Rs is solution resistance, ρ is conductivity ofelectrolyte,  is the length of the fluid capillary formed bya μB residing in the microchannel, and Stot is the total crosssectional area occupied by the fluid capillaries [5]. Externalpressure of the liquid media is transferred through theinterface ports, which directly influences bubble size as
described by the Young-Laplace equation [7] and Boyle’slaw.
Figure 4: Photograph of microfabricated Parylene C
device attached to flat flexible cable (FFC) via zero
insertion force (ZIF) connector; connection scheme
provides robust and reversible electrical contact.

METHODS
Fabrication
The devices (Fig. 3) were fabricated using surface
micromachining techniques on a polymer substrate withthin film electrodes, a process based on previously reported
techniques [2]. Platinum was e-beam deposited (2000 Åthick) and patterned on a Parylene coated (12 μm thick)silicon wafer. Following deposition of a 10 μm Parylene
insulation layer, sacrificial photoresist was patterned toestablish the microchannel. A final 4 μm thick layer ofParylene was used to enclose the microchannel, and accessports were etched using switched-chemistry oxygen
plasma etching process in a deep reactive ion etching(DRIE) tool [8]. The entire fabrication process wasperformed at relatively low temperatures (<90°C) toprevent thermal degradation of Parylene.
Devices were released in an acetone bath, and anelectrical connection was established with a ZIF connector(Fig. 4) using previously reported methods [9]. The
microchannel was first soaked in isopropyl alcohol tofacilitate filling with 1× PBS (phosphate buffered saline),
an electrolyte that simulates in vivo conditions.

Figure 2: Cross section profile of device containing a
trapped microbubble, with fluidic interfaces oriented (a) inline with and (b) normal to direction of liquid flow.
Pressure differential due to flow is captured as differencein pressure applied through ports P1 and P2.
The Venturi effect harnesses Bernoulli’s principle,
which describes the pressure differential that accompanies
a change in fluid flow speed. By placing the fluidic
interface ports parallel to the direction of flow (Fig. 2a), theconfined bubble will experience this pressure differential
which is a direct measure of flow rate. When the device isoriented such that the ports are perpendicular (Fig. 2b) to
the flow, the bubble experiences no pressure differential
due to flow but instead transduces the static pressure. Thus,
the sensor may be operated in two different modes bychanging the orientation of the device with respect to thedirection of liquid flow.

Figure 3: Micrograph of sensors soaked and filled withelectrolyte. A variety of microchannel dimensions wereutilized to enable testing of various electrolyte and flow
parameters.
Figure 5: Schematic overview of test setup, illustrating
electrical and fluidic connections.
Experimental Setup
The μB transducer was housed within a custom acrylic
test fixture, oriented parallel and normal to applied flow. A
peristaltic pump (Watson-Marlow, Wilmington, MA) was326

connected to the flow chamber (3 mm x 2 mm x 10 mm)
and electrical connections were routed to custom
multiplexing circuitry controlled via LabVIEW (Fig. 5).

Figure 7: Tracking of dissolution time (τdissolution) for
repeated bubble nucleation events. Once stabilized, thelocal electrolyte environment was saturated with dissolved
gas and repeatable impedance measurements wererealized.
Figure 6: Representative impedance measurement ofelectrolytically generated microbubble, with indicators forparameters of interest. (a) Impedance of microbubble
immediately following electrolysis. (b) Microbubble begins
detachment from microchannel. (c) Microbubble rapidly
recombines to fluid in presence of platinum electrode.

For a given constant pressure environment andelectrolysis parameters, impedance tracking ofmicrobubble size during dissolution was performed. The
ΔZ metric (Fig. 6) was observed to have an inversely
proportional relationship (ΔZ = 60496 + (1.23 ×
106/(pressure (mmHg) + 148.85)), R2 = 0.999) to arandomly applied pressure (Fig. 8), suggesting that theimpedance measurement is a direct measurement of thebubble volume relationship as described by the YoungLaplace equation and Boyle’s law. The possibility ofabsolute pressure monitoring is promising with further
calibration of saturation conditions.

Electrochemical impedance spectroscopy (EIS) results
indiciated that 10 kHz was the optimum measurement
frequency (minimum system phase) for maximizing thesolution resistance component of the impedance response.
Microbubbles were electrolytically generated in 1× PBS
under conditions of 1-4 μA for 2-5 s. Prior to and following
current injection, impedance measurements were taken(Fig. 6) with a precision LCR meter (Agilent E4980A).

RESULTS AND DISCUSSION
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 repeated andimpedance measurements were taken until τdissolution wasconsistent to within 10% (Fig. 7). This priming process
takes ~ 50 minutes and lasts for ~ 3 hours. The ΔZ andτdissolution (as illustrated in Fig. 6) from a primed environment
impedance response curve (i.e. the 5th to 10th current
injections in Fig. 7) were then analyzed for pressuretransduction. Notably, the size of the nucleated bubble
remained consistent (~1% RSD) while attempting tosaturate the local dissolved gas environment, as seen in thelow variance of ΔZ (mean = 37.9 kΩ, standard deviation
574.6 Ω, n = 10).

Figure 8: Normalized impedance change (ΔZ) is directly
correlated to bubble size and confirms the Boyle’s law
pressure-volume relation.
Dissolution rate was also shown to have a correlation
with pressure (Fig. 9). The observed relation wasapproximately linear (dissolution rate = 100.70 + 0.67 ×pressure (mmHg), R2 = 0.998), as confirmed by Henry’s
law and the Epstein-Plesset equation [10]. Provided that theelectrochemical measurement is taken when theenvironment properly primed, this metric provides amethod to check the ΔZ measurement on the same
nucleated microbubble.

327

ACKNOWLEDGEMENT
This work was funded by the NSF under award
number ECCS-1231994 and EFRI-1332394. The authors
would like to thank Mr. J. Nishida, Mr. A. Jibodu, and themembers of the USC Biomedical Microsystems
Laboratory for their assistance.

REFERENCES[1] L. Bowman and J. D. Meindl, "The packaging of
implantable
integrated
sensors,"
Biomedical
Engineering, IEEE Transactions on, pp. 248-255,
1986.
[2] C. A. Gutierrez and E. Meng, "A subnanowatt
microbubble pressure sensor based on electrochemical
impedance transduction in a flexible all-parylene
package," in 24th IEEE International Conference on
Micro Electro Mechanical Systems, MEMS 2011,January 23, 2011 - January 27, 2011, Cancun,
Mexico, 2011, pp. 549-552.
[3] T. Y. Chang, V. G. Yadav, S. De Leo, A. Mohedas, B.
Rajalingam, C.-L. Chen, et al., "Cell and protein
compatibility of parylene-C surfaces," Langmuir, vol.
23, pp. 11718-11725, 2007.
[4] C. Shih, T. A. Harder, and Y.-C. Tai, "Yield strengthof thin-film parylene-C," Microsystem Technologies,
vol. 10, pp. 407-411, 2004.
[5] D. A. Ateya, A. A. Shah, and S. Z. Hua, "Impedancebased response of an electrolytic gas bubble topressure in microfluidic channels," Sensors andActuators, A: Physical, vol. 122, pp. 235-241, 2005.
[6] J. E. B. Randles, "Kinetics of Rapid Electrode
Reactions," Discussions of the Faraday Society, vol. 1,
pp. 11-19, 1947.
[7] B. Ran and J. Katz, "The response of microscopic
bubbles to sudden changes in the ambient pressure,"
Journal of Fluid Mechanics, vol. 224, pp. 91-115,
1991.
[8] E. Meng, P.-Y. Li, and Y.-C. Tai, "Plasma removal ofParylene C," Journal of Micromechanics andMicroengineering, vol. 18, p. 045004, 2008.
[9] B. J. Kim, C. A. Gutierrez, G. A. Gerhardt, and E.
Meng, "Parylene-based electrochemical-MEMS force
sensor array for assessing neural probe insertion
mechanics," in Micro Electro Mechanical Systems(MEMS), 2012 IEEE 25th International Conference
on, 2012, pp. 124-127.
[10] P. S. Epstein and M. S. Plesset, "On the Stability ofGas Bubbles in Liquid-Gas Solutions," Journal ofChemical Physics, vol. 18, pp. 1505-1509, 1950.
CONTACT*E. Meng, tel: +1-213-8213949; ellis.meng@usc.edu

Figure 9: Dissolution rate (Rdissolution in Fig. 6) was foundto be linearly related to hydrostatic pressure.
Measurements under liquid flow for the device
oriented parallel to flow are presented in Fig. 10 andpressure was found to increase with increased flow, as
described by Bernoulli’s principle (R2 = 0.92). Whenoriented normal to direction of flow, flow sensing wasnegligible and only static pressure was measured.
Switching between pressure and flow sensing can beachieved by changing sensor orientation with respect todirection of liquid flow.

Figure 10: Differential pressure-flow relationship. Withsensor oriented normal to flow, measured pressure isnegligible. When sensor ports are oriented parallel to flow,
measured pressure differential confirms square root
relationship as described by Bernoulli’s principle.

CONCLUSION
A unique sensing mechanism for hydrostatic pressure
measurement and liquid flow utilizing impedance based
measurements of electrolytically generated microbubbles
was developed and demonstrated. With the exceptional
biocompatibility characteristics of the open-to-liquid
packaging, this robust sensing mechanism can be usedreliably for chronic in vivo measurements. Additional
sensor characterization is planned and performance will beassessed in a clinical setting to achieve first-in-humandemonstration.

328