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An Electrochemical Impedance-Based Thermal
Flow Sensor for Physiological Fluids
Alex Baldwin, Lawrence Yu, and Ellis Meng, Senior Member, IEEE

Abstract— A novel electrochemical-thermal flow sensor wasdeveloped for use in physiological liquids. The sensor was constructed out of a platinum resistive heater and platinum sensing
electrodes on a Parylene C substrate, rendering it flexible andfully biocompatible. During heating, changes in electrochemical
impedance across the sensing electrodes were used to detectchanges in temperature, and highly sensitive flow measurements
were achieved with overheat temperatures of only 1 °C. The
sensor’s biocompatibility and low overheat temperature make it
an ideal candidate for chronic in vivo applications. [2016-0170]
Index
Terms— Bioimpedance,
impedance,
biomedicaltransducers, microelectronic implants, in vivo, fluid flow,
microfluidics, fluid flow measurement, electrothermal effects,
temperature dependence, Parylene C, thermal flow sensing,electrochemical-microelectromechanical systems (EC-MEMS).

I. I NTRODUCTIONPRECISE and accurate knowledge of flow rate is a ubiquitous concern of both medicine and biology. Accurate flow
measurement is critical in applications such as drug delivery,
drug synthesis, and diagnostic tests. For in vivo applications,
implantable microscale flow sensors have been proposed tomonitor the flow of blood, cerebrospinal fluid, or other bodily
fluids, provide tracking of a disease state, or deliver diagnostic
information to clinicians. However, few flow sensors possess
all the capabilities necessary for medical applications, which
include compatibility with physiological fluid, low detection limit, low power consumption, and biocompatibility forin vivo implantation. Mechanical designs, which often mimic
macroscale turbine or cantilever flow sensors [6], possess
moving parts at risk of biofouling when exposed to biologicalfluids. Thus, designs having no moving parts are preferred forphysiological environments.
Microfluidic thermal flow sensors have been successfully
used for gas and liquid sensing in several applications andoperate by measuring convective heat transfer in a fluid asa function of flow rate [1]. Although many microfabricatedManuscript received July 22, 2016; revised September 7, 2016; accepted
September 16, 2016. Date of publication October 13, 2016; date of current
version November 29, 2016. This work was supported in part by the Universityof Southern California Coulter Translation Research Partnership Award and inpart by NSF under Award ECCS-1231994 and Award EFRI-1332394. Subject
Editor A. Zhang.
A. Baldwin and L. Yu are with the Biomedical Engineering Department,
University of Southern California, Los Angeles, CA 90089 USA (e-mail:abbaldwi@usc.edu; lawrence.yu@usc.edu).
E. Meng is with the Biomedical Engineering and the 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.2016.2614664

thermal flow sensors have been successfully developed andcommercialized for gas flow, there are additional challenges
involved in designing sensors for in vivo measurement ofphysiological liquids. Due to the high thermal conductanceof water and physiological fluids, sensors must be smaller
and operate at a higher overheat temperature to minimize heatdissipation losses. Sensors designed for implantation must also
minimize heating to avoid damaging the body, and must be
constructed from biocompatible materials that will not corrode
or degrade in vivo. Most prior reports of microfabricatedthermal flow sensors employ traditional semiconductor-basedmaterials, typically as silicon-based sensors employing a thinfilm microheater (often relying on silicon nitride membranes
for thermal isolation) with joule heaters and resistance temperature detectors (RTDs) fabricated from polysilicon [7]–[9],germanium [11]–[13], or metals [15]. These sensors haveachieved high accuracies and sensitivities over a wide rangeof flow velocities thanks to the high temperature sensitivityof doped polysilicon and germanium, but are not suitablefor chronic implantation due to their corrosion under chronic
exposure to physiological fluid [17]. Some sensors, designed
for use in water or physiological fluids, have been reportedwith metal heaters and temperature sensors fabricated onpolyimide [14], [18], Parylene C [10], [19], [20], or otherbiocompatible polymer substrates [16], [21]. Unfortunately,
metals have an order of magnitude lower sensitivity to temperature changes than semiconductor-based resistors, resulting
in low sensitivity and requiring high overheat temperatures.
The low sensitivity of current polymer-based thermal flow
sensors and the lack of biocompatibility in semiconductorbased thermal flow sensors has left a need for a flow sensor
which is both highly sensitive and biocompatible in design andmaterial. We have developed a thermal flow sensor designed
not only to function in physiological fluids, but to exploit theionic conductivity of several biological liquids which enables
improved sensitivity. Temperature-mediated changes in the
conductivity of the solution are monitored via high-frequency
electrochemical impedance to deduce flow rate, and the corresponding temperature sensitivity is an order of magnitudelarger than the temperature coefficient of resistance (TCR) in
metallic RTDs (Table 1). The new sensor is constructed ofa platinum resistive heater and sensing electrodes on a thinfilm Parylene C substrate. Parylene’s high thermal resistivity
reduces thermal conduction across the substrate. A time-offlight method for flow transduction was developed whichmeasures the rate of heat transfer due to both diffusion andconvection, enabling accurate measurements of low flow rates

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TABLE IT EMPERATURE C OEFFICIENT OF R ESISTANCE (TCR) OF M ETALSAND S EMICONDUCTORS W HICH A RE C OMMONLY U SED IN
R ESISTIVE T EMPERATURE S ENSORS , AS W ELL AS
S EVERAL I ONIC S OLUTIONS I NCLUDING
H UMAN C EREBROSPINAL F LUID [1]–[3]Fig. 1. Temperature sensitivity of conductivity, which is comparable to TCR,
of an aqueous solution of sodium ions at infinite dilution.

and bidirectional flow detection using only a single temperature sensor. Here we present the operating theory, fabrication,and characterization of the sensor and sensing mechanism.
II. T HEORY
A. Electrolyte Conductivity
The conductance, and thus the electrochemical impedance,of an aqueous electrolyte solution is highly dependent on temperature. Electrical conduction through an electrolyte solution
arises from the movement of charged ions, and the solution’s
ionic conductance describes the movement of these ions in
an electric field. This movement is determined by the sumof forces between ions and the surrounding solution. Bulk
conductance is directly proportional to ionic concentration but
a molar conductivity can be defined as  = G/c, where G is
the bulk conductance, and c is the molarity of ions [22]. The
limit of conductivity as concentration goes to zero is known
as the state of ‘infinite dilution’ and is useful to consider sincein this state, conductivity depends on interactions between asingle ion and the solution. If the mobility of an ion is definedas the limiting velocity per applied force, the conductivity at
infinite dilution is
λ0 = uq F(1)where u is ionic mobility, q is the charge on the ion, andF is Faraday’s constant [23]. Stoke’s law gives the first
approximation for ionic mobility as a product of the dragbetween a spherical ion and a continuous solution, and thislaw can be extended by including dielectric losses producedby moving charges to give3 ε0 − ε∞u −1 = 4πηr + q 2 2 3 τ0
8
ε0 r

(2)

where η is water’s viscosity, r is ionic radius, ε0 and ε∞
are the high-frequency and low-frequency limits of the solution permittivity, and τ0 is the Debye relaxation time [24].

All temperature-dependent terms in this expression are parameters of water, and viscosity is the term which has the largest
effect, indicating that the effect of changing temperatures onionic mobility is largely independent of the ionic species
present and is closely related to changes in water’s viscosity.
Viscosities, permittivities, and Debye relaxation times ofwater at temperatures from 10-60°C were used to calculate conductivity at infinite dilution of a common ion(sodium) [25]–[27]. Figure 1 shows the percent change inthis conductivity per °C, which is equivalent the temperaturesensitivity of the solution analogous to TCR of a solid.
The temperature coefficient of conductivity at infinite dilution is a good indicator of the response of an electrolyte
solution to changes in temperature, but forces exist between
ions that must be taken into account for real solutions.
Debye, Huckel, and Onsager summarize these forces in theexpression⎛⎞
2 eF 2
3 eF
√
z
qz

0 ⎠ c (3) = 0 − ⎝ 2  23 +2
24πε RT ε RT
3πη ε RT
where c is the electrolyte concentration in moles per liter,
0 is the infinite dilution conductivity, z is the valency numberof each ionic species, R is the gas constant, e is the electron
charge, and q describes the symmetry of an ionic species [28].
Calculating this conductivity for various ionic species reveals
that the temperature coefficient of conductivity decreases asconcentration increases. However, the difference in temperature coefficient between infinite dilution and most physiological fluids is minimal. For example, cerebrospinal fluid andblood have approximate ionic molarities of 295 mM [29]and 345 mM [30], which have a difference in temperature
coefficient from a solution at infinite dilution of 0.19% and0.21% respectively.
Changes in conductivity of an electrolyte solution canbe detected by measuring the electrochemical impedance
between two immersed electrodes. Typically, three electrodes
are used for electrochemical measurements, with one electrodeproviding a stable reference voltage in the electrolyte. This
allows for sensitive measurements at the electrode-electrolyte
interface and has been used for a number of amperometricbased sensors. However, to measure changes in the bulk

BALDWIN et al.: ELECTROCHEMICAL IMPEDANCE-BASED THERMAL FLOW SENSOR FOR PHYSIOLOGICAL FLUIDS

Fig. 2. Our sensor transduces flow through the transfer of heat from a
resistive heater to flowing fluid and from the fluid to a pair of electrodes.
The electrodes sense changes in temperature via changes in electrochemical
impedance (Zsense ). Orange represents the temperature distribution being
distorted by flow, blue represents the flowing liquid and light blue represents
the polymer substrate.

The maximum rate of change of temperature at electrodes
can be used to transduce flow velocity in both convection and diffusion-dominated states, since the heating profile around a heater is affected by both convection anddiffusion. Measuring the maximum rate of change allows
flow transduction down to arbitrarily low flow velocities andfor both positive and negative flow directions. To validate
this approach, a 1-dimensional finite difference simulation
was used. Assuming perfect insulation and an infinitely
long fluid cylinder, the explicit finite difference equation forheat transfer via the convection-diffusion equation in one
dimension is
T jn − T jn−1
t

=α

n−1
T jn−1+ T jn−1+1 − 2T j−1−V
conductivity of the electrolyte a high frequency signal ispassed between electrodes, and at the proper frequency theelectrode-electrolyte interface is completely bypassed and areference electrode is no longer necessary.
B. Flow Measurement
To enable bidirectional flow measurement at low flow rates,
a time of flight method was developed requiring a heater anda single pair of impedance electrodes, where the rate at which
the heated solution reaches the electrodes is used to transduce
flow velocity. Figure 2 shows the basic setup of our sensor.
The transfer of heat away from a heater can be describedusing the convection-diffusion equation:∂T= ∇∗ (α∇T ) −∇ ∗ (VT) + H (x, t)(4)∂twhere T is temperature, α is the fluid’s thermal diffusivity,
V is the fluid’s velocity vector at any point, and H (x, t)
is any change in temperature forced upon the system [31].
The impulse response to the convection-diffusion equation isa constantly widening and shifting Gaussian curve, and whenconvection is the dominant method of heat transfer, the peakof this Gaussian can be tracked to derive the flow velocity.
However, for low flow rates and over short distances, diffusion
is the dominant method of heat transfer, and the Gaussian
will tend to dissipate before the peak moves a significant
distance. The relative dominance of convection over diffusion
is described by the Peclet numberLv(5)αwhere L is length and v is the scalar flow velocity [32]. Whenthe Peclet number is greater than 1, the system is dominatedby convection. The highest flow velocity we tested was around
800 μm/s, which with a temperature sensor spaced 1 mm
away from the heater and using the thermal diffusivity of purewater (α = 143 × 10−9 m2 /s) gives a Peclet number of 5.59.
This indicates that the system is slightly convection-dominatedbut that reducing flow velocity would lead to a diffusiondominated state. Therefore, a method of flow measurement
is needed which operates in both diffusion-dominated andconvection-dominated states.
Pe =

1017

( x)2
n−1
T jn−1+1 − T j −1
2 x

+H

(6)

where T jn is the temperature at time n t and position
j x [33]. The stability criteria for this method is x < 2αV2and t < 2αx . Choosing 50 μm for x and 5 ms fort satisfies these criteria. The initial temperature was set to
25 °C, and after 1 second, the heat at x = 0 increased by
2 °C to simulate constant current delivered to the heater. Heat
conduction through the polymer substrate was considered to be
negligible, since the polymer used (Parylene C) has a thermal
conductivity an order of magnitude lower than the surrounding
fluid (0.084 W/m·K, versus 0.596 W/m·K for water) [34].
Figure 3 shows the simulation results, which confirm thatthe heating profile 1 mm away from the heater is alteredby changes in the flow velocity and that the maximum rateof change of temperature can be used to transduce flow.
This method of flow transduction overcomes some limitationsof traditional time of flight measurements and allows thetemperature measurement electrodes to be placed close to
the heater, allowing for smaller devices with lower overheattemperatures and enabling the accurate measurement of low
flow velocities.
III. S ENSOR D ESIGN
Sensors which used the above principles to transduce flow
velocity were designed and fabricated, each sensor consistingof a resistive heater and a pair of exposed platinum electrodes
deposited on a thin-film Parylene C substrate, as show in
Figure 4. The heater is 1 mm long and 250 μm wide andconsists of a serpentine platinum trace with a trace widthand spacing of 25 μm. Each heater has a DC resistanceof 750-800 , depending on process variations. The measurement electrodes are spaced 750 μm apart perpendicular tothe direction of flow, with each electrode having a geometricsurface area of 20000 μm2 . Two pairs of electrodes were
fabricated on each sensor to enable accurate measurement
over a large range of flow rates. Two different sensor variations were designed, one with electrodes spaced 0.5 mm
and 2 mm away from the center of the heater and one withelectrodes spaced at 1 mm and 3 mm. Each die also contained a previously reported pressure sensor [35] and patency
sensor [36].

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Fig. 5. A sensor die (A) just after being released from its silicon carrier
wafer and (B) packaged in a luer lock connector for fluidic testing.

Fig. 3. (A) Simulations show that the temperature profile 1 mm away fromthe heater is dependent on flow and (B) that the maximum rate of change oftemperature is related to flow velocity.

Fig. 4. Fabricated sensor die showing the resistive heater and two sets ofimpedance electrodes. The die also contains previously reported pressure andpatency sensors.

pads were exposed using a switched chemistry deep reactive
ion etching process in oxygen plasma and a second cutout
etch was used to separate devices. The free-film devices
were released from the carrier wafer by gently peeling while
immersed in deionized water, and any residual photoresist wasremoved by washing in acetone.
Electrical connections were made through a zero insertion
force (ZIF) connector soldered onto a flat flexible cable. The
contact pads of the released device were first attached to a
250 μm thick PEEK (polyetheretherketone) backing using
cyanoacrylate glue to achieve the thickness necessary for
insertion into the ZIF connector. The ZIF connector was thenencapsulated with biocompatible EpoTek 353-NDT epoxy. For
flow testing, the sensors were packaged into a luer lock, shownin Figure 5. Using a luer lock connector enables devices toeasily attach to shunts and catheters currently used in hospitals,
including external ventricular drains for cerebrospinal fluid,which is our primary target application. A slot was milled
into the top of a luer lock connector with an inner diameter of
3.25 mm and the sensor was inserted and sealed with EpoTek
353-NDT epoxy such that one end of the sensor fixed to theconnector and one free-standing. Care was taken to place thesensor films as close to the center of the luer lock connector
as possible so that the resistive heater and electrodes are in
the region of maximum flow velocity. An acrylic jig was used
to hold the sensors in place while the epoxy dried.
V. F LOW T ESTING AND C HARACTERIZATION

IV. FABRICATION
Fabrication followed previously reported surface micromachining processes for Parylene C MEMS devices described indetail elsewhere and briefly summarized here [37]. A 12 μm
thick Parylene substrate was deposited on a silicon carrier
wafer, and the heater, electrodes and contacts were definedin a single layer of 2000 Å thick sputtered platinum. Components were patterned with UV lithography followed by metal
deposition and lift-off before being insulated in a second
12 μm thick Parylene layer. Electrode surfaces and contact

The heater resistance and electrode impedance werecharacterized in phosphate-buffered saline (PBS), an
isotonic solution commonly used to mimic solutions(pH 7.4, ionic molarity 280 mM), at temperatures between
25 and 50 °C. Tests were also performed in artificial
cerebrospinal fluid (aCSF), which has a slightly higher ionicmolarity and includes calcium and magnesium to better mimic
the ionic environment in in vivo. aCSF was prepared by
combining 8.66 g NaCl, 0.224 g KCl, 0.206 g CaCl2 · 2H2 O,
0.163 g MgCl2 · 6H2 O, 0.214 g Na2 HPO4 · 7H2 O, and 0.027 g

BALDWIN et al.: ELECTROCHEMICAL IMPEDANCE-BASED THERMAL FLOW SENSOR FOR PHYSIOLOGICAL FLUIDS

1019

Fig. 6. To transduce flow rate, the impedance during heating was normalized
to a baseline value and the minimum rate of change in impedance, which
occurred 1-2 s after heater activation, was recorded.

NaH2 PO4 · H2 O in 1 L of double-distilled water [38]. The
heater’s DC resistance was measured using a Keithley 2400
SourceMeter and electrochemical impedance spectra was
acquired between 1 Hz and 1 MHz using a Gamry R600
potentiostat. Fluidic testing was performed by flowing PBS
through packaged sensors using a Watson Marlow 120U
peristaltic pump with a flow rate range of 0 to 400 μL/min.
Constant current pulses were delivered to the heater using a
Keithley 2400 SourceMeter and electrochemical impedance
was measured at 100 kHz and 1 VPP using an Agilent E4980A
LCR meter. Sensor performance was first characterized under
static conditions using heater currents between 0.5 and 6 mA,
corresponding to overheat temperatures between 0 and 9 °C,
and was then characterized at flow rates between -400 and
400 μL/min using a 2 mA heater current. Sensors were alsocharacterized in 0.25, 0.5, 1, and 2× dilutions of PBS, whichhave ionic concentrations of 0.07, 0.14, 0.28, and 0.56 mol/L
respectively, as well as tap water. Additional characterization
was performed in PBS at temperatures between 19 and 30 °C,
with the exact temperature at the heater confirmed throughmeasurement of the DC resistance of the heater.
To transduce flow velocity, the electrochemical impedanceacross a pair of sensing electrodes was recorded by a LabVIEW program at a rate of 5 Hz. The baseline impedance
without heating was measured, and any deviations fromthis baseline due to heating were normalized as a percent
change. The instantaneous rate of change was then calculated while the heater was active, and the minimum rate ofchange, corresponding to the fastest increase in temperature
at the electrodes, was taken as a measurement of the flow
velocity (Figure 6).
VI. R ESULTS
Characterization of the electrodes at different temperatures
showed that the largest temperature-based changes in electrochemical impedance occurred at frequencies greater than
∼60 kHz, corresponding to the frequency range at which
phase approaches zero. To ensure that all measurements werein this range, all impedance measurements were performed
at 100 kHz. Heater characterization showed that the TCR of

Fig. 7. (A) The overheat temperatures at the heater and (B) the response ofan electrode pair 1 mm away for various current levels at no flow.

the resistive heater was approximately 0.16 %/°C, which iswithin the expected range for thin-film platinum [39]. From
the TCR, the overheat temperature was calculated during flow
testing by recording the heater’s resistance during current
injections. Figure 7 shows the heater overheat temperatureand the impedance of electrodes 1 mm away normalized tothe pre-heating value. A heater current of 2 mA provided anadequate impedance response with an overheat temperatureof only 1.04 °C. To minimize power consumption and reduce
the risk of damaging tissue or other structures due to excessive
heating, all subsequent tests used a heater current of 2 mA.
A volumetric flow rate of 1 μL/min will result in a flow
velocity of 2 μm/s when sent through a tube with a diameterof 3.25 mm. Figure 8A shows the impedance response ofelectrodes 1 mm away from the heater operating at an overheat temperature of 1°C under different flow velocities. The
response profile and the minimum rate of change of impedance
was found to vary with flow velocity and to be consistentacross multiple trials (Figure 8B).
Figure 9 shows the minimum rate of change of impedanceat flow velocities between −800 and 800 μm/s. The rateof change versus flow velocity takes the form of a steadily
increasing curve at small negative and positive flow velocities,
but begins to level off at large negative flow velocities. The

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Fig. 10. The sensor response of electrodes spaced 0.5 mm, 1 mm, and 2 mm
away from the heater. The response of electrodes at 4 mm was negligible.

Fig. 8. (A) The percent change in impedance at electrodes 1 mm away fromheater during constant 1°C heating and (B) the instantaneous rate of change.
The minimum (peak) of the rate of change is used to transduce flow velocity.

Fig. 11. A comparison of sensor response in 1x PBS and aCSF, which showsno significant differences.

Fig. 9. The minimum rate of change of impedance 1 mm away at flow
velocities from −800 to 800 μm/s.

maximum standard deviation between −200 and 200 μm/s
is 43.3 μm/s.
Figure 10 shows the response of sensors with electrodes
spaced 0.5, 1, and 2 mm away from the heater. Changing
the heater-electrode spacing changes the dynamic range of thesensor, and electrodes spaced 0.5 mm away from the heater
showed a higher dynamic range than the electrodes at 1 mm.
However, measurements from electrodes at a 0.5 mm spacing

were observed to be noisier than measurements at 1 mm.
Sensors with a 4 mm electrode-heater spacing showed no
response.
To ensure that results using PBS can be generalized tophysiological fluids, a sensor was tested in both PBS andartificial cerebrospinal fluid (aCSF), which contains 150 mM
Na+ , 3.0 mM K+ , 1.4 mM Ca2+ , 0.8 mM Mg2+ , 1.0 mM
PO43− , and 155 mM Cl− . Figure 11 shows the results, which
reveal no significant differences in the response.
The baseline impedance of electrodes in different dilutions of PBS ranged from 3.67±0.04 k in 2× PBS to
20.7±0.18 k in 0.25× PBS, and electrodes in tap water
showed a baseline impedance of 82.5±8.9 k (n = 15,
mean ± SD). Figure 12 shows the sensor’s response in thesesolutions. As concentration decreased, the magnitude of thesensor’s response increased slightly, though this change is only
greater than standard deviation between tap water and the mostconcentrated PBS solution.
Figure 13 shows the sensor’s response in 1× PBS at ambient
fluid temperatures between 19 and 30°C. The baseline rate ofchange of the sensor decreases slightly over this range.

BALDWIN et al.: ELECTROCHEMICAL IMPEDANCE-BASED THERMAL FLOW SENSOR FOR PHYSIOLOGICAL FLUIDS

Fig. 12. (A) The minimum rate of change versus flow velocity at various
dilutions of PBS. 0.25, 0.5, 1, and 2× PBS have ionic concentrations of
70, 140, 280, and 560 mM respectively. (B) The sensor response versus thebaseline impedance of the temperature sensing electrodes.

1021

Fig. 14. Comparison of sensor response between (A) 1°C and 2°C and(B) 1°C and 10°C. Higher overheat temperatures are somewhat noisier but
possess a significantly higher dynamic range, resulting in a net improvement
in measurement resolution.

A sensor was characterized using a 1, 2, and 10 °C overheattemperatures, which correspond to 2, 2.9, and 6.2 mA currents
delivered to the resistive heater. Figure 14 shows the sensor
response results at these different operating regimes. When theoverheat temperature is increased from 1 to 2 °C, the resolution
improves significantly, with the standard deviation between−200 and 200 μm/s decreasing from 28.6 to 17.3 μm/s.
Increasing the overheat temperature to 10 °C further decreases
the resolution to 13.1 μm/s, though this improvement is notproportional to the change in temperature due to higher measurement variations seen during higher temperature operation.
VII. D ISCUSSION
Fig. 13. The sensor response using 1°C overheat temperature at ambient
temperatures between 19°C and 30°C. The minimum rate of change becomesslightly less sensitive at higher temperatures.

Using a 1°C overheat temperature is desirable to reduce
power consumption and avoid tissue damage in in vivo
applications, but higher overheat temperatures would provide
greater dynamic range and potentially increase sensor resolution, which would be useful for non-implanted applications.

Results from sensor testing matched simulation results inthat the maximum rate of change of temperature at downstream electrodes, which corresponded to the minimum rateof change of impedance, can be used to transduce flow
velocity (Fig. 15). However, experimental results showed asignificantly lower dynamic range compared to the simulation,
most likely due to heat loss through the walls of the luer
lock connector, which assumed to be zero in the simulation.
At negative and low positive velocities, the minimum rateof change was related to flow velocity as predicted by the

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Fig. 15. A comparison between simulated results and experimental results,
both with approx. 2°C heating. The results show a similar pattern over the
majority of flow rates, though due to heat loss through the walls of the flow
channel the experimental results have a lower dynamic range.

simulation, but at high positive velocities the rate of changebegan to deviate slightly from what the simulation predicted.
This may have been due to the use of constant current toheat the fluid, which would result in a small decrease inheater temperature as the flow velocity increases. The dynamic
range at higher flow velocities could therefore be improved byusing a constant temperature at the resistive heater, though thiswould necessitate more complex feedback circuitry.
The decrease in sensor response with both increasing
ambient temperature and increasing ionic concentration also
followed theoretical predictions, although the changes in sensitivity with ionic concentration were found to be insignificant
within an order of magnitude difference in ionic concentration.
However, when used in an environment where large changes
in ionic concentration are expected, the baseline impedanceof the sensing electrodes without heater activation could be
used to calibrate sensor response, though this would requirea priori calibration over the expected ionic concentration
range. Changes in sensitivity with ambient temperature were
small but significant compared to the sensor’s dynamic range
for future applications, the heating resistor may be used as anambient temperature sensor for real-time calibration, since theresistance of platinum is highly stable and increases linearlywith temperature. Similar devices were fabricated which used
a second platinum resistor as a temperature sensor, but after
initial testing these were discarded due to the low TCR of theresistor and inductive coupling between the heater and resistor,
which overwhelmed any flow-based effects.
The dynamic range of the sensor was relatively small
when using a 1 °C overheat temperature, but was improved
during operation with 2 and 10 °C overheat temperature.
Slightly increasing the overheat temperature result in noticeable improvements in sensor resolution, and for ex vivo
applications, overheat temperatures even higher than 10 °C
may be useful, as long as those temperatures do not affect thesurrounding environment. Association for the Advancement ofMedical Instrumentation standards limit temperature increases
in implants to between 1 and 2 °C [40], so any implantabledevice will need to stay within this range, but even with aminimal overheat temperature, the sensor can transduce flow
velocity ranges relevant to cerebrospinal fluid flow throughhydrocephalus shunts [41] or external ventricular drains [42]as well as capillary blood flow [43]. Arterial blood flow has
a much higher average velocity, but even without increasing
overheat temperature, the sensor can be tailored to measure
higher flows by increasing the distance between electrode andheater. Multiple pairs of electrodes could also be used totransduce a wider range of flow velocities. Outside of the bodythere are many applications from cell culturing to microfluidic
diagnostics that could benefit from low flow rate sensing ofelectrolytic solutions. Furthermore, benchtop devices could
have higher operating temperature limits compared to in vivo
applications and could therefore be tuned to a much wider
range of flow velocities.
The small size of our sensor and its construction on a
flexible substrate allow it to be easily integrated into shunts,
catheters, and other equipment, and this design could befurther miniaturized to access additional applications. Future
work will investigate changing the size and orientation ofelectrodes in order to improve sensitivity and the use ofcalorimetric sensing to obtain a larger measurement range
and decrease the time per measurement. Future work will
focus on reducing measurement variation in order to improve
resolution. This variation was observed to be much smallerwhen the sensor was tested without flow compared to when
fluid was flowing, indicating that the pulsatile operation of thepump may be reflected in the sensor’s output. The sensor may
also move when subjected to flow, as only one end of thesensor is affixed to the wall of the luer lock connector; thismovement may further contribute to measurement variation.
Variation between devices was also present, which may havebeen a result of small deviations in the placement of the
Parylene sensor die from the center of the flow channel. Noncentered placement of the sensor would result in a lower flow
velocity at the sensor die, which may have caused a baseline
shift. Before packaging an assumption was made that, due
to water’s high thermal conductivity, heat would spread tothe region of highest flow velocity quickly enough that small
deviations from center would not have a noticeable effect on
sensor performance. Future work will test this assumption
and evaluate whether an alternative packaging scheme could
decrease inter-device variability.
Table 2 shows select parameters of several commercial
and laboratory flow sensors. T is the overheat temperatureused and range is either the total range over which effective
flow measurements can be made or the largest range tested,whichever is smaller. Resolution for commercial sensors is
usually defined as 3× the standard deviation, but in literature
the value usually reported is the smallest flow rate difference
at which a difference in measurements was detected. There
are very few commercial thermal flow sensors which areable to sense liquids traveling at low flow velocities, and ofthese the Sensirion sensor is the most effective, reporting aminimum detectable flow velocity of ∼14 μm/s with a very
wide detectable flow range. However, all available commercial
liquid flow sensors use silicon-based sensing elements andBALDWIN et al.: ELECTROCHEMICAL IMPEDANCE-BASED THERMAL FLOW SENSOR FOR PHYSIOLOGICAL FLUIDS

1023

TABLE II
A COMPARISON OF S EVERAL T HERMAL F LOW S ENSORS , B OTH C OMMERCIAL AND R EPORTED IN L ITERATURE

high overheat temperatures, rendering them incompatible withchronic in vivo use. Many other thermal flow sensors designed
for water or other liquids have been reported in literature, butthese have high overheat temperatures, low sensitivities, or
non-biocompatible materials which prevent their use in vivo.
VIII. C ONCLUSION
A thermal flow sensor which uses electrochemical
impedance to measure changes in temperature was designed
and tested. The use of impedance as well as a novel timeof flight method for transducing flow allows for bidirectional
measurement of low flow velocities with an overheat temperature of only 1°C. Accurate flow transduction is possible despite
variations in ionic concentration, and in-use temperature
calibration is possible thanks to the use of a platinum resistive
heater. Due to its flexible, robust, and biocompatible design
this sensor is useful in a wide range of applications for whichcurrent thermal flow sensors are not suitable, including chronic
implantation in the human body
ACKNOWLEDGMENT
The authors would like to thank Madelina Pratt andWilla Jin for their help with characterization and testing and
Dr. Brian Kim and the members of the Biomedical
Microsystems Laboratory of USC for all of their help andsupport.
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Alex Baldwin received the B.S. degree inelectrical engineering from the University of
Arkansas, Fayetteville, in 2013, and the M.S. degree
in biomedical engineering from the University ofSouthern California in 2015, where he is currently
pursuing the Ph.D. degree in biomedical engineering with the Biomedical Microsystems Laboratory.
He is a recipient of the USC Graduate School
Fellowship. He is currently working on implantablephysiological sensors constructed using polymer
micromachining.

Lawrence Yu received the B.S. degree in biomedical
engineering, M.S. degree in electrical engineering,and Ph.D. degree in biomedical engineering fromthe University of Southern California in 2009, 2010,
and 2016, respectively. His work in the Biomedical
Microsystems Laboratory focused on the development of a sensor platform for in vivo applications.
He was a recipient of the Walt Disney Foundation
Scholarship and the Rose Hills Foundation 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 from the
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, Universityof Southern California, Los Angeles, where she has
been since 2004. She also holds a joint appointment
with the Ming Hsieh Department of ElectricalEngineering. Her current research interests
include bioMEMS, implantable biomedical microdevices, microfluidics,
multimodality integrated microsystems, and packaging. She was the Viterbi
Early Career Chair in the Viterbi School of Engineering. She is a memberof Tau Beta Pi, the Biomedical Engineering Society, the Society of WomenEngineers, and the American Society for Engineering Education. She wasa 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.