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Fluid Temperature Measurement in Aqueous
Solution via Electrochemical ImpedanceAlex Baldwin , Eugene Yoon, Trevor Hudson, and Ellis Meng , Fellow, IEEE

Abstract— A novel temperature transduction method using
the high-frequency electrochemical impedance between a pairof microelectrodes exposed to aqueous solution is reported. The
solution resistance of an aqueous ionic solution is highly temperature dependent, and the electrochemical impedance between
two microelectrodes is dominated by solution resistance whenmeasured at the appropriate frequency. Therefore, precise measurements of electrochemical impedance at the proper frequency
in a two-electrode system can be used to transduce solution
temperature. To demonstrate this method, a temperature sensor
composed of two thin-film platinum electrodes on a freestanding
Parylene C substrate was designed and fabricated. A platinum
resistance temperature detector was co-fabricated to provide a
simple means of benchmarking against an existing standard.
Transduction via electrochemical impedance was achieved bymeasuring the real part of impedance at the frequency wherephase was minimized. Fluid temperature was transduced between
15◦ C and 50◦ C with high sensitivity (−1.21%/◦ C) and resolution(0.02◦ C). A 4× improvement in sensitivity and resolution overthe conventional platinum resistance temperature detector wasachieved. The sensor design described here features flexible
construction with inert materials which facilitates future use inbiomedical or microfluidic applications.
[2019-0190]
Index Terms— Temperature measurement, electrochemical
impedance, biomedical sensor, thin film sensor, flexible sensor.

I. I NTRODUCTIONEMPERATURE measurement of aqueous solutions isvital in a number of different microfluidic, medical,
and biological applications. To measure flow rate, precise
temperature measurement can be utilized to create a thermal
flow sensor which tracks heat flow due to the moving liquid [1]. Applications of thermal liquid flow sensors include
flow cytometry [2], drug delivery dosing [3], and lab-on-chip
devices [4], [5]. In the medical realm, precise temperature
measurement in the body can be used to monitor blood

T

Manuscript received August 26, 2019; accepted September 3, 2019. Date of
publication September 18, 2019; date of current version December 4, 2019.
This work was supported in part by the U.S. National Science Foundation
under Award EFRI-1332394 and in part by the Alfred E. Mann Institute
Fellowship (AB). Subject Editor Y. Zohar. (Corresponding author: Alex
Baldwin.)
A. Baldwin is with Niche Biomedical Inc., Los Angeles, CA 90024 USA(e-mail: axbaldwin@gmail.com).
E. Yoon and T. Hudson are with the Department of Biomedical Engineering,
University of Southern California, Los Angeles, CA 90089 USA (e-mail:eugenejy@usc.edu; tqhudson@usc.edu).
E. Meng is with the Department of Biomedical Engineering, Universityof Southern California, Los Angeles, CA 90089 USA, and also with theMing Hsieh Department of Electrical and Computer 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 article are available
online at http://ieeexplore.ieee.org.
Digital Object Identifier 10.1109/JMEMS.2019.2939811

perfusion to an organ [6] or as part of a control system
to avoid these pathological states [7], [8]. Many of theseapplications involve prolonged exposure to saline or otheraqueous ionic liquids; therefore, a temperature sensing methodthat can withstand a corrosive saline environment is desirable.
State-of-the-art methods for fluid temperature measurement
include thermocouples and resistance temperature detectors(RTDs) [9], [10]. Thermocouples utilize the Seebeck effect,in which a temperature-dependent voltage occurs at the junction of two dissimilar conducting elements [11]. They are
generally constructed from nickel alloys and have temperature sensitivities up to 60 μV/◦ C, though their precision
is low (±2◦ C). Semiconductor junctions can also be usedas thermocouples, with sensitivities up to 110 mV/◦ C andprecisions approaching ±0.8◦C [9]. However, thermocouples
require encapsulation to protect conducting elements fromwater intrusion or shield the body from non-biocompatible
materials, further reducing sensitivity.
RTDs are conductive or semiconductive elements whose
resistance changes with temperature. Most metal RTDs use
platinum due to its highly linear response, but other metals
such as copper and nickel can be used as well. The standard
PT100 RTD is made of bulk platinum wire and definedby a resistance of 100  at 0◦ C [10]. Platinum RTDs canalso be microfabricated on thin-film substrates, though thetemperature sensitivity of thin-film platinum (0.198%/◦Cfor ∼100 nm films) is lower than in bulk platinum(0.385%/◦C) [12]. Laboratory-grade RTDs can operate
between −200 and 1000◦C with precision approaching±0.001◦C whereas commercial RTDs can have precisions
on the order of ±0.03◦C. Drawbacks of platinum RTDs
include low sensitivity and the requirement for encapsulation
in aqueous solutions, to prevent unwanted charge transfer
through the solution and biofouling. Semiconductor RTDs
are known as thermistors, and can possess temperature
coefficients of resistance approaching −2%/◦ C [9], [13].
However, semiconductors corrode and disintegrate during
chronic soaking in aqueous solutions [14]–[16].
This work proposes an alternate method of fluid temperaturetransduction via direct measurement of a fluid’s solution resistance. The solution resistance between electrodes immersed
in an aqueous ionic solution is highly sensitive to temperature changes, mainly due to decreases in water’s viscositywith increased temperature. Decreasing viscosity increases
the mobility of ions dissolved in solution, which lowers
a solution’s electrical resistivity. The temperature sensitivityof an aqueous solution’s resistivity matches or exceeds

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BALDWIN et al.: FLUID TEMPERATURE MEASUREMENT IN AQUEOUS SOLUTION

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where C O (0, t) or C R (0, t) is the concentration of the oxidizedor reduced species at 0 displacement from the electrode at∗ or C ∗ is the bulk concentration of the oxidized
time = t, C O
Ror reduced species, αc is a transfer coefficient which can range
from zero to unity, n is number of electrons transferred, i 0 is
the exchange current density, F is Faraday’s constant, R is thegas constant, and T is temperature in Kelvin [23], [24]. Using
the linear approximation of an exponential (e x ≈ 1 + x), thisequation can be simplified to give an expression for chargetransfer resistance in response to small overpotentials:−i o n F
v0
RT
dv 0
RT
Rct =
=
di netn Fi oi net =
Fig. 1.
Temperature sensors are composed of 2000 Å thick platinum
electrodes (exposed area 150 × 250 μm2 ) sandwiched between 10 μm
thick layers of Parylene C and exposed to the solution. The electrochemical
impedance between two identical electrodes consists of the solution resistanceR S , as well as a charge transfer resistance Rct and double layer capacitance Cdl at each electrode interface. At high measurement frequencies,
the double layer capacitance acts as a short circuit and impedance can beapproximated as R S .

the sensitivities of semiconductor materials; for example,
the reported temperature coefficient of resistivity for humancerebrospinal fluid at body temperature was previously measured as −1.98%/◦C [17]. It follows that measuring solution
resistance using high-frequency electrochemical impedance
enables highly sensitive fluid temperature transduction.
Low-frequency electrochemical impedance was measured
across the terminals of a lithium-ion battery to transduce its
internal temperature [18], [19]. The temperature transduction
method was based on charge transfer resistance andnot investigated for other applications. In contrast,
this work utilizes the temperature sensitivity of highfrequency electrochemical impedance measured between two
microfabricated electrodes exposed to saline and demonstrates
precise temperature transduction. Previously, qualitative
measurement of high frequency electrochemical impedance
using microfabricated platinum electrodes was used to createa thermal flow sensor [20], [21]. Impedance measurements
were used to track the time of flight of a heat pulse, allowing
for measurement of low flow rates with minimal overheattemperature. The present temperature detector builds off
this early qualitative demonstration for flow measurement.
In comparison to earlier methods relying on charge transfer
resistance, potential toxic byproducts of Faradaic reactions
are avoided, and rapid measurement is possible.
II. T HEORY
The electrochemical impedance between two identical electrodes immersed in solution can be approximated as a simplified Randles circuit [22] which includes a charge transfer
resistance Rct and double layer capacitance Cdl at eachelectrode-electrolyte interface, and the solution resistance Rs
between electrodes (Fig. 1).
Rct describes charge transfer between the metal electrodeand the aqueous solution due to Faradaic redox reactions. For
a given overpotential v 0 , the net current i net is given by:
C R (0, t) v 0 (1−αc )n F/RT
Co (0, t) −v 0 αc n F/RT
i net = i o
e
−
e
Co∗
C ∗R
(1)

(2)
(3)

Equation 3 implies a direct correlation between Rct and temperature. However, the byproducts of Faradaic reactions can betoxic [25], so measurement of Rct in biological or microfluidic
systems is undesirable. Furthermore, Rct only factors in to
the impedance of microelectrodes at frequencies <1 Hz [26],preventing rapid measurement which limits the utility ofthe technique. Therefore, practical electrochemical impedance
measurements between microelectrodes directly exposed tosolution will be composed of Cdl and/or R S .
The capacitance Cdl is created by the double layer ofelectrostatically-bound ions that builds up at the electrodeelectrolyte interface [27]. Helmholtz, Gouy-Chapman, Stern,
Graham, and others have modeled this phenomenon, but the
precise relationships between Cdl and temperature remains anarea of ongoing research [28], [29].
Solution resistance, R S , is highly sensitive to changes intemperature. The solution resistance between two electrodes
depends on the orientation of and distance between electrodes,
but generalized expressions can be developed by considering
an electrolyte solution’s resistivity or its reciprocal value, conductivity. Bulk conductivity is proportional to ionic concentration, but a molar conductivity  = G/c can be defined, where
G is bulk conductance and c is concentration [30]. As ionicconcentration goes to zero, interionic forces disappear. Thisis known as the state of infinite dilution and can be modeledusing Stoke’s law as:
λ0 =

qF
4πηr

(4)

where λ0 is conductivity at infinite dilution, q is ionic charge,
η is water’s viscosity, and r is ionic radius [31]. Only viscosity
varies significantly with temperature; therefore, the temperature coefficient of solution resistance at infinite dilution is
proportional to water’s viscosity [20].
At concentrations above infinite dilution, solution conductivity will decrease. Interionic forces at concentrations
above infinite dilution can be modeled using Debye-HuckelOnsager theory [32]; the Kohlrausch equation can then be
employed to generate the following expression for solution
conductivity:⎛⎞√z 2 eF 2
x z 3 eF
0
0


+
λ ⎠ c
(5)
=λ −⎝
22
3πη ε0 RT
24πε0 RT ε RT

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Fig. 2. The temperature coefficient of solution resistance for a solutionof sodium ions at infinite dilution and at the osmolarity of cerebrospinal
fluid (CSF). At infinite dilution, the temperature coefficient ranges from
−3.39%/◦ C at 0◦ C to −1.04%/◦ C at 100◦ C, with a value of −1.91%/◦ C
at body temperature (37◦ C). Increasing ionic concentration decreases temperature sensitivity; at the osmolarity of CSF (295 mOsm), the temperature
coefficient of solution resistance decreases to −1.75%/◦ C at body temperature.
Figure generated via MATLAB.

where c is electrolyte concentration in moles per liter, z is thevalency number of each ionic species, e is the electron charge,
and x describes ionic symmetry [33].
This expression for solution conductivity can be used to
calculate the sensitivity of solution resistance to temperature.
Empirical measurements for the viscosity [34], permittivity [22], and Debye relaxation time [23] of pure water and thehydrated radius of a sodium ion [35] were used to calculate
temperature coefficients of solution resistance of a solutionof sodium ions (Fig. 2). At infinite dilution, the temperature
coefficient of resistance calculated using Equation 5 ranged
from −3.39%/◦C at 0◦ C to −1.04 %/◦ C at 100◦C. Equation 6 can be used to calculate the temperature coefficientof solution resistance at concentrations above infinite dilution; temperature sensitivity decreases at higher concentrations
compared to infinite dilution, but there is minimal change
in temperature coefficient between 100 mOsm and 1 Osm.
At body temperature (37◦C), the temperature coefficient ofsolution resistance is −1.75%/◦C at a typical osmolarity of
cerebrospinal fluid (295 mOsm [36]), compared to −1.91%/◦C
at infinite dilution.
The predictable and highly sensitive relationship between
R S and temperature suggests a possible method for temperature transduction. By measuring electrochemical impedance athigh frequencies where phase is minimized, Cdl and Rct are
rendered negligible and impedance consists only of R S . Therefore, fluid temperature can be transduced using high-frequency
electrochemical impedance.
III. D ESIGN AND FABRICATION
A temperature sensor was designed and fabricated totransduce fluidic temperature via high-frequency impedance.
The sensor consisted of a pair of platinum electrodes on a

Fig. 3. Impedimetric temperature sensing was tested using a pair of platinum
electrodes on a Parylene C substrate. A microfabricated platinum RTD on thesame substrate was used for benchmarking.

Fig. 4. Fabrication on a Parylene C substrate give sensors flexibility, enabling
simple packaging and integration into catheters and shunts used in the clinic.

Parylene C substrate (Fig. 3). Electrodes had exposed areas
150×250 μm2 with centers 750 μm apart. A platinum resistor,
placed 1 mm away from the electrodes on the same Parylene
C substrate, was used as an RTD for benchmarking. This RTD
consisted of a serpentine trace 25 μm wide with nominal
DC resistance of ∼450 . Construction out of platinum on
a Parylene C substrate leaves sensors flexible, [37] (Fig. 4)
for simpler integration into microfluidic packaging. In addition, both Parylene C and platinum are non-toxic and havea long history of use in chronically implantable medical
devices [38], [39].
Devices were fabricated using previously-reported polymer
micromachining techniques [40], [41]. First, a 10 μm thicklayer of Parylene C was deposited onto a 4” silicon carrier
wafer using a SCS LabCoter 2 chemical vapor deposition
system (Specialty Coating Systems, Indianapolis, IN, USA).
Next, a mask for metal deposition composed of 2 μm thick
AZ 5214 photoresist (Integrated Micro Materials, Argyle, TX,
USA), was spin-coated and patterned using UV lithography.
A 2000 Å thick layer of platinum was electron-beam deposited

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BALDWIN et al.: FLUID TEMPERATURE MEASUREMENT IN AQUEOUS SOLUTION

and patterned via liftoff. A second 10 μm thick layer ofParylene C was deposited as insulation, then electrodes andcontact pads were defined by a switched-chemistry deep
reactive ion etch in oxygen plasma [41] using an Oxford
Plasmalab 100 inductively coupled plasma etcher (Oxford
Instruments, Abington, UK). For oxygen plasma etching,an etch mask was defined using a 15 μm thick layer of AZ
P4620 photoresist (Integrated Micro Materials, Argyle, TX,
USA). Devices were released from the carrier wafer by gently
peeling while immersed in deionized water, then cleaned bysoaking in acetone, isopropyl alcohol, and deionized water
for ten minutes each. After release, devices were thermallyannealed at 200◦C for 48 hours in a vacuum environment;
thermal annealing increases crystallinity, decreasing flexibility
slightly but improving adhesion between adjacent Parylene C
layers [40], [42]–[45].
Electrical connection to devices was achieved by attachinga 250 μm thick PEEK (polyetheretherketone) backing tocontact pads using cyanoacrylate glue and inserting into a zero
insertion force connector [46] (ZIF; Hirose Electric Co. Ltd.,Tokyo, Japan). Electrical connections were protected during
soaking using EpoTek 353ND-T epoxy (Epoxy Technology
Inc., Billerica, MA, USA).
IV. T ESTING M ETHODS
Sensors were primarily tested in phosphate-buffered saline(1× PBS), a common physiological fluid analog. 1× PBS is
similar in composition to cerebrospinal fluid and possesses
a nominal resistivity of 50 ·cm. For comparison, the conductivity of human cerebrospinal fluid has been measured as
68.7±1.1 ·cm at room temperature and 55.5±0.7 ·cm atbody temperature [17]. Transduction was also evaluated inconcentrated PBS (10×PBS, resistivity 5 ·cm) and deionized(DI) water (resistivity 18 M·cm).
For electrochemical impedance spectroscopy (EIS), sensors
were sealed in glass vials filled with 1× PBS and placedin a water bath, which was heated to temperatures between
30 and 50◦C. A Gamry Reference 600 potentiostat wasused for EIS measurement. For single-frequency impedance
tests, solution was heated from room temperature (∼20◦C)
to 50◦C using a hot plate or cooled to 15◦ C with ice packs.
A thermocouple/multimeter (Newport TrueRMS Supermeter,±2◦ C accuracy) was used to monitor fluid temperature andcalibrate the co-fabricated RTD; RTD calibration was achievedby comparing resistance measurements to thermocouple dataover multiple cycles and generating a linear regression. Data
was collected from impedance electrodes using an Agilent
E4980A precision LCR meter with a measurement amplitudeof 0.1 VPP , to stay within the low overpotential regime. Data
was acquired from the RTD using a Keithley SourceMeter
with 10 μA bias current. The impedance between electrodes
and the resistance of the platinum RTD were simultaneously
collected and analyzed using LabVIEW.
To evaluate drift, a sensor die was sealed into a silicone
tube using EpoTEK 353-NDT epoxy and exposed to room
temperature 1× PBS for 24 hours. Fluid was flowed past
the sensor at 100 μL/min using a Watson-Marlow 120U

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Fig. 5.
Electrochemical impedance spectroscopy between two platinum
electrodes at temperatures between 30 and 50◦ C. As temperature increased,
the real impedance at high frequencies decreased.

peristaltic pump to prevent local variations in the bulk ionicconcentration. Impedance and RTD resistance were recorded
once per minute, and a Sensirion SLI-1000 operating in its
temperature transduction mode (±0.1◦C precision) was used
to confirm measurements.
V. E XPERIMENTAL R ESULTS
EIS from 1 Hz to 100 kHz was performed at 5◦ C temperature intervals from 30 to 50◦ C (Fig. 5). Only the real part ofimpedance was recorded in order to remove any interference
from parasitic capacitances. Results showed a clear temperature dependency of impedance at high frequencies.
The platinum RTD was calibrated by comparing its resistance to thermocouple measurements in 1×PBS, revealing atemperature coefficient of resistance of 1.21 /◦ C. Temperature was then cycled between 20 and 50◦C while simultaneously measuring the impedance across the electrodes at
100 kHz, the DC resistance of the RTD, and fluid temperaturefrom the thermocouple. The real part of impedance correlatedwith temperature with a sensitivity of -58.29 /◦ C (Fig. 6).
The resolution, calculated as 3× the average RMS noise
across the tested temperature range divided by the linearized
temperature coefficient of resistance between 20 and 50◦ C,
was 0.02◦C for impedance-based temperature sensing and0.08◦C for the platinum RTD. Direct comparison between
temperature measurements using the calibrated RTD and real
impedance values exhibited an inverse relationship (Fig. 7).
Minimal hysteresis was observed over 8 cycles. Both electrode
impedance and RTD resistance responded to ±30◦ C changes
in temperature in less than 200 ms.
Temperature sensing was then evaluated in DI water and
10× PBS. First, EIS between the microelectrodes was performed in DI water and 10× PBS at room temperature (Fig. 8),
revealing ideal measurement frequencies of 1 kHz for DI waterand 1 MHz for 10× PBS. Baseline impedance magnitudes
were 1.77 M at 1 kHz for DI water and 803  at 1 MHz

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1064

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 28, NO. 6, DECEMBER 2019the next 19 hours. Over this time, the electrode impedance
dropped approximately 5.3% from its baseline value while
the RTD resistance increased by 0.29%. Solution conductivity
was measured before and after the experiment (A212, ThermoFisher, Waltham, MA, USA) to monitor evaporation, which
detected an increase from 17.01 to 17.68 mS/cm. Correcting
for this conductivity change results in only a 0.2◦C error
between temperature measurements using the electrodes versus
using the platinum RTD (Fig. 10). This drift was not observed
during tests at room temperature.
VI. D ISCUSSION

Fig. 6.
Real impedance at 100 kHz and RTD resistance compared tothermocouple measurements. Impedance was roughly linear with temperature,with a temperature coefficient of resistance −58.29 /◦ C and resolution of
0.02◦ C. This compared favorably to the RTD, which showed a temperature
coefficient of resistance of 1.21 /◦ C and a resolution of 0.08◦ C.

Fig. 7. Real impedance showed an inverse relationship with temperature,
as measured using the platinum RTD on the same substrate. Minimal hysteresis was observed during repeated temperature cycles from 15 to 45◦ C.

for 10× PBS, compared to 5.55 k at 100 kHz for 1× PBS.
Temperature coefficients were found to be remarkably constant, even over a >103 change in solution resistance.
Electrode impedance and RTD resistance were monitored
in flowing PBS at room temperature for 24 hours (Fig. 9).
Both the RTD and the electrodes recorded ambient temperature fluctuations throughout the day. During the first threehours of measurement, the temperature as measured by the
RTD increased approximately 2◦ C, a change which was notreflected in the real impedance measurement or the Sensirion
temperature data. Between 3 and 24 hours, there was nodetectable drift between the impedance measurement, the RTD
resistance, and the Sensirion temperature data.
To examine drift, a sensor was immersed in 1× PBS andplaced in an environmental chamber (EC0-A, Sun Electronic
Systems Inc., Titusville, FL, USA) set to 37◦ C. After twohours of warmup time, over which time both the impedancebased temperature sensor and the RTD detected an increase
from 21 to 37◦ C, the chamber maintained its temperature
within the resolution of the chamber’s thermocouple for

The experimental results demonstrate that electrochemical
impedance is highly sensitive to changes in temperature. Thiseffect was used to measure sub-degree temperature fluctuations
at a higher resolution than platinum RTDs, which are considered the gold standard for fluid temperature transduction.
High-speed and high precision measurement of changes influid temperature is useful in thermal anemometry; previous
work has utilized the qualitative application of impedancebased temperature transduction to create highly sensitive liquid
thermal flow sensors [20], [21].
Construction out of thin-film platinum on a Parylene C
substrate renders this sensor flexible and resistant to corrosion.
A flexible sensor is advantageous for packaging and willallow easy integration into the walls of shunts and catheterscurrently used in the clinic. However, flexing of the substrate
may change the distance between electrodes, creating drift
in impedance magnitude and preventing accurate temperature
measurement. For accurate temperature transduction, electrodeposition and spacing should be fixed prior to calibration
or baseline measurements. In addition, the use of Parylene C as a substrate material limits operation to temperatures
below its glass transition temperature, which lies between
60 and 90◦ C [47], [48]. The current version of the sensor
is designed for use between room temperature and 37◦ C,
which do not exceed this temperature limit; future versions
for industrial applications could be fabricated on borosilicate
glass to avoid thermal degradation issues.
During 24 hours of continuous measurement at room
temperature there was no significant drift in the electrode
impedance compared to the RTD. However, when held at 37◦ C
significant drift was observed in impedance measurements.
This drift is most likely caused by increasing ionic concentration in the testing solution due to evaporation, as suggested
by conductivity meter readings before and after testing. Aftercorrecting for changes in solution composition, drift remained
below 0.2◦C over a 21-hour period.
Future target applications for the temperature sensor
described in this work will focus on temperature transduction
in vivo. Variations in physiological fluid temperature can affectthe calibration of biochemical sensors, so sensitive temperature measurement is useful to reduce noise and increase
accuracy. Measuring the temperature of blood can also be
useful for hemodynamic assessment and for calibration ofblood pressure or flow sensors, and chronic measurement of
CSF temperature can be used to monitor the efficacy of hydrocephalus treatment. Electrochemical impedance measurements

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BALDWIN et al.: FLUID TEMPERATURE MEASUREMENT IN AQUEOUS SOLUTION

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Fig. 8. Left: EIS showing the impedance phase across temperature sensing electrodes in DI water, 1× PBS, and 10×PBS. Higher ionic concentrations resulted
in a shift of the ideal measurement frequency (where phase is minimized) towards higher frequencies. Right: Relative impedance compared to thermocouple
measurements in DI water, 1× PBS, and 10× PBS. Temperature coefficients of real impedance in the resistive range stayed relatively constant despite a >103
change in solution resistance. Real impedance was measured at 1 kHz in DI water, 100 kHz in 1× PBS, and 1 MHz in 10× PBS.

Fig. 9.
Electrode impedance at 100 kHz in 1× PBS was measured
once per minute for one day and compared to temperature as measured
by the co-fabricated RTD. Real impedance and RTD temperature between
6 and 12 hours of measurement shows sub-degree changes in room temperature due to cycling of the air conditioning system of the laboratory.

are independent of the substrate and electrode material whenmeasured at the appropriate frequency, since any surface
effects are bypassed. Therefore, temperature measurement via
electrochemical impedance can be extended to other rigid or
flexible substrates and can be implemented in pre-existing
medical devices such as pacemakers or continuous glucose
monitors that already contain electrodes exposed to aqueous
solution. Future studies will characterize the effects of biofouling on the measurement signal, and self-cleaning techniquessuch as cyclic voltammetry or electrolysis will be evaluated.
Outside of the body, impedance-based temperature measurement has applications in microfluidics, diagnostic testing, andchemical processing, since these fields all require precise control of fluid temperature to manipulate reaction kinetics. The
chemical inertness, biocompatibility, and corrosion resistanceof both Parylene C and platinum make the sensor described

Fig. 10. Impedance and RTD resistance were measured during 21 hours ofsoaking in an environmental chamber set to 37◦ C. Over this time, temperature
measurements by the impedance sensor drifted by approximately 0.2◦ C
compared to temperature measurements by the platinum RTD.

in this work widely applicable throughout the biomedical andindustrial spaces.
VII. C ONCLUSION
A novel method of temperature transduction in aqueous
solutions was explored using a polymer MEMS sensor. The
sensor consisted of two exposed platinum electrodes embedded
in a thin-film Parylene C substrate; temperature was transduced by measuring the electrochemical impedance between
the electrodes at the frequency where phase was minimized.
According to the Randles model, at this frequency, impedance
will be dominated by solution resistance, which is highly
temperature sensitive due to decreases in water’s viscosity withincreasing temperature. Experimental results showed that thesensor transduced fluid temperature with higher sensitivity andlower noise than a platinum RTD co-fabricated on the same

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substrate; sensitivity was consistent over a wide range of ionicconcentrations. This microfabricated polymer temperature sensor is flexible and corrosion resistant, lending itself to a wide
range of applications in the chemical and biomedical fields.
Future development will focus on practicality for chronic
in vivo temperature transduction.
ACKNOWLEDGMENT
The authors would like to thank Dr. Brian Kim of
Verily Life Sciences for his ideas and encouragement,
Dr. Donghai Zhu of the Keck Photonics Lab for assistancewith fabrication, and all the members of the Biomedical
Microsystems Lab of USC for their help and support.
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Alex Baldwin received the B.S. degree in electricalengineering from the University of Arkansas, Fayetteville, AR, USA, in 2013, and the M.S. and Ph.D.
degrees in biomedical engineering from the University of Southern California (USC), in 2015 and2018, respectively. His graduate and post-doctoral
work focused on improving the treatment of hydrocephalus using chronically implantable sensors andremote patient monitoring. He is currently developing novel neurostimulation-based therapies at Niche
Biomedical Inc. His research interests include bioelectronic medicine, chronically implantable medical device, and biomedical
sensors. He was a recipient of the USC Viterbi Fellowship and the Alfred E.
Mann Institute Fellowship.
Eugene Yoon received the B.S. degree inchemical and biomolecular engineering from theJohns Hopkins University, Baltimore, in 2015, andthe M.S. degree in biomedical engineering fromthe University of Southern California (USC),
Los Angeles, in 2017. He is currently pursuing the
Ph.D. degree in biomedical engineering with USC,
where his research is focused on bioMEMS sensors
and implantable microdevices.

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Trevor Hudson received the B.S. degree in biomedical engineering from Rice University, in 2014, andthe M.S. degree in biomedical engineering from the
University of Southern California (USC), in 2016.
He is currently pursuing the Ph.D. degree with
USC, where he researches implantable impedimetricsensors and polymer micromachining.

Ellis Meng (M’02–SM’09–F’16) received the B.S.
degree in engineering and applied science andthe M.S. and Ph.D. degrees in electrical engineering from the California Institute of Technology (Caltech), Pasadena, in 1997, 1998, and 2003,
respectively. She is currently a Professor of biomedical engineering and the Vice Dean of Technology Innovation and Entrepreneurship of theViterbi School of Engineering, University of Southern California, Los Angeles, where she has been,
since 2004. She also holds a joint appointment in
the Ming Hsieh Department of Electrical and Computer Engineering. Her
research interests include bioMEMS, implantable biomedical microdevices,microfluidics, multimodality integrated microsystems, and packaging. She isalso a Fellow of ASME, BMES, NAI, and AIMBE. She held the Viterbi Early
Career Chair at the Viterbi School of Engineering and was the Department
Chair, from 2015 to 2018. She is also an Inaugural Holder of a Gabilan
Distinguished Professorship in Science and Engineering. Her honors include
the NSF Career award, Wallace H. Coulter Foundation Early Career Award,
2009 TR35 Young Innovator Under 35, Viterbi Early Career Chair, ASEE
Curtis W. McGraw Research Award, and 2018 IEEE Engineering in Medicine
and Biology Society Technical Achievement Award.

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