of 11
Song
et al
.,
Sci. Adv.
2020;
6
: eaay9842 30 September 2020
SCIENCE ADVANCES
|
RESEARCH ARTICLE
1 of 10
APPLIED SCIENCES AND ENGINEERING
Wireless battery-free wearable sweat sensor powered
by human motion
Yu
Song
1,2
*, Jihong
Min
1
*, You
Yu
1
, Haobin
Wang
2
, Yiran
Yang
1
, Haixia
Zhang
2
, Wei
Gao
1†
Wireless wearable sweat biosensors have gained huge traction due to their potential for noninvasive health
monitoring. As high energy consumption is a crucial challenge in this field, efficient energy harvesting from
human motion represents an attractive approach to sustainably power future wearables. Despite intensive re-
search activities, most wearable energy harvesters suffer from complex fabrication procedures, poor robustness,
and low power density, making them unsuitable for continuous biosensing. Here, we propose a highly robust,
mass-producible, and battery-free wearable platform that efficiently extracts power from body motion through a
flexible printed circuit board (FPCB)–based freestanding triboelectric nanogenerator (FTENG). The judiciously
engineered FTENG displays a high power output of ~416 mW m
−2
. Through seamless system integration and effi-
cient power management, we demonstrate a battery-free triboelectrically driven system that is able to power
multiplexed sweat biosensors and wirelessly transmit data to the user interfaces through Bluetooth during
on-body human trials.
INTRODUCTION
Prolific research toward the development of wearable bioelectronic
technologies is greatly expanding the horizons of personalized health
monitoring (
1
4
). Wireless wearable devices offer a noninvasive
means of extracting real-time physiological parameters indicative of
health status and transmitting continuous data to a user device.
Wearable devices capable of detecting various vital signs such as pulse,
respiration rate, and temperature are being widely commercialized
and integrated into daily lifestyles (
5
9
). Sweat is another attractive
medium containing a multitude of molecular biomarkers including
electrolytes, metabolites, amino acids, hormones, and drugs that
wearable sensors can analyze (
10
15
). Continuous monitoring of
these biomarkers has the potential to supplement laboratory-based
blood tests in enabling real-time monitoring of daily health status as
well as early disease detection and management (
15
20
).
Over the past years, extensive interest and efforts have focused
on developing novel sensors and improving the wearability of these
platforms (
14
,
21
). Until now, most wearable sensor prototypes
relied on bulky and rigid battery packs to power the electronic cir-
cuitry for data acquisition, processing, and transmission. Flexible
batteries have been proposed to enable conformal contact on skin
(
22
,
23
), and incorporation of low-power electronics has substan-
tially reduced the power requirements of wearables and enabled the
use of small coin cell batteries. Despite these efforts, batteries still
face limitations in that they need to be charged and replaced fre-
quently. In addition, while unlikely, lithium ion batteries are sus-
ceptible to explosion, posing safety concerns. Battery-free systems
powered by near-field communication (NFC) are reported (
24
,
25
)
but could suffer from short operation distance. As an alternative,
energy can be harvested from renewable, portable, and sustainable
sources such as solar light, biofluids, and human motion to power
future wireless wearable electronics (
26
29
).
Triboelectric nanogenerators (TENGs), which convert the mechan
-
ical energy created by human motion into electrical energy via cou-
pling of inductive and triboelectric effects (
30
35
), offer a highly
attractive energy harvesting strategy to power wearable sweat sen-
sors during intensive physical activities as their operation is inde-
pendent from uncontrollable external sources such as sunlight or
wireless power transmitters. Despite the advantage, most of the
existing TENG-based devices suffer from low power intensity, inef-
ficient power management, and a lack of power continuity and lon-
gevity; thereby, the use of a TENG to continuously power a fully
integrated wireless wearable molecular sensor system has not been
reported (
28
,
36
39
).
Here, we propose a battery-free, fully self-powered wearable
system that consists of a highly efficient wearable freestanding-
mode TENG (FTENG), low-power wireless sensor circuitry, and a
microfluidic sweat sensor patch on a single
flexible printed circuit
board (FPCB) platform that can dynamically monitor key sweat
biomarkers (e.g., pH and Na
+
) (Fig. 1A). Such an FTENG-powered
wearable sweat sensor system (FWS
3
) is designed and prepared to
be compatible with traditional FPCB manufacturing processes, which
enables mass productivity and high reliability. Our freestanding
FPCB-based design, coupled with effective power management,
allows efficient energy harvesting from human skin, particularly
suitable for powering skin-interfaced wearables. With the aid of
waterproof medical tape, the FWS
3
can be conformally laminated
on the side torso to maximize the potential for energy harvesting
(Fig. 1, B and C). The integrated Bluetooth Low Energy (BLE) mod-
ule allows sensor data to be conveniently transmitted to a mobile
interface for health status tracking during exercise. This represents
the first demonstration of a fully integrated battery-free triboelec-
trically driven wearable system for multiplexed sweat sensing.
RESULTS
Design of
the
FWS
3
The FTENG consists of an interdigital stator and a grating-patterned
slider (Fig. 1D). To obtain a strong electrification effect, polytetra-
fluoroethylene (PTFE) and copper are used as tribo-pairs in the
1
Andrew and Peggy Cherng Department of Medical Engineering, California Insti-
tute of Technology, Pasadena, CA 91125, USA.
2
National Key Lab of Micro/Nano
Fabrication Technology, Peking University, Beijing 100871, China.
*These authors contributed equally to this work.
†Corresponding author. Email: weigao@caltech.edu
Copyright © 2020
The
Authors, some
rights reserved;
exclusive licensee
American Association
for the Advancement
of Science. No claim to
original U.S.
Government
Works. Distributed
under a Creative
Commons Attribution
NonCommercial
License 4.0 (CC BY-NC).
on October 1, 2020
http://advances.sciencemag.org/
Downloaded from
Song
et al
.,
Sci. Adv.
2020;
6
: eaay9842 30 September 2020
SCIENCE ADVANCES
|
RESEARCH ARTICLE
2 of 10
flexible FTENG.
The FTENG is fabricated through commercial
FPCB technology (as illustrated in fig. S1), and the detailed dimen-
sional parameters are shown in fig. S2. The inter-electrode distance
was optimized through transferred charge density studies of the
FTENGs (fig. S3). The stator and slider are patterned through pho-
tolithography as the periodically complimentary interdigital struc-
ture and the grating structure, respectively. With an electroless
nickel/immersion gold (ENIG) surface finish on the electrode area,
the stator is further laminated by PTFE.
The reusable flexible cir
-
cuitry in conjunction with a disposable microfluidic sweat sensor
patch can continuously perform electrochemical measurements of
key biomarkers in sweat (Fig. 1E). During exercise, the FTENG gener
-
ated power is stored and released from a capacitor as controlled by
the power management integrated circuit (PMIC) (Fig. 1F). When
fully charged, the storage capacitor releases its stored energy, which
is regulated to a stable voltage to power the BLE programmed system
on a chip (PSoC) module and instrumentation amplifiers for acqui-
sition and transmission of potentiometric measurements over BLE.
Characterization of
the
FTENG
The working mechanism of the FTENG can be explained as the
coupling effect of contact electrification and in-plane sliding-induced
charge transfer, as illustrated in Fig. 2A. As copper is more tribo
-
electrically positive than PTFE, electrons accumulate on PTFE
during the sliding process. In the initial state, the grating slider fully
overlaps with one stator electrode and no charge flow occurs be-
tween the interdigital stator electrodes due to the electrostatic equi-
librium. The unidirectional sliding process results in a charging
flow between stator electrodes until the grating slider fully overlaps
with the second stator electrode with reversed polarity. This work-
ing process was further validated by numerical simulation using
COMSOL Multiphysics (fig. S4). The detailed model of the FTENG
under open-circuit and short-circuit conditions is explained in fig.
S5 and notes S1 and S2. The optical microscopic image of our pro-
posed FPCB-based FTENG and the typical short-circuit current
(
I
SC
) profiles of an FTENG at different working frequencies are
shown in Fig. 2 (B and C). The FTENG operates consistently at
varying frequencies of 0.5, 1.25, and 3.3 Hz, resulting in the maxi-
mum
I
SC
of 8.39, 19.11, and 42.25
A, respectively. The open-circuit
voltage (
V
OC
) attained at the frequency of 0.5 Hz is presented in fig.
S6A, where the signal polarity of the envelope waveform oscillates
rapidly along the sliding process. To evaluate the use of our FPCB-
based FTENG as a power source, the voltages and powers are mea
-
sured under a series of different load resistances (Fig. 2D), with a
working frequency of 1.5 Hz for actuation. An increase of resistance
beyond 1 megohm leads to a rapid increase of voltage. With a load
resistance of 4.7 megohms, the FTENG reaches a maximum power
output of 0.94 mW (corresponding to 416 mW m
−2
).
The PTFE of the FPCB-based FTENG displays superior durabil-
ity performance over conventional micro-pyramid polydimethylsi-
loxane (M-PDMS) and wrinkled PDMS (W-PDMS) after 20,000
working cycles (Fig. 2E), where the
V
OC
experiences minimal decay
Na
+
H
+
A
B
We
arable sw
eat
sensor
system
DE
F
C
Signal
processi
ng
Sw
eat
biosens
ing
Wi
reless
trans
missi
on
Heal
th status
alar
m
Energy harvestin
g
V
out
BLE PSo
Cm
odule
AD
C
Instrum
ent
am
plifiers
Voltage
regulato
r
User
interfac
e
GP
IO
V
S
PM
IC
FT
EN
G
~
V
DD
Biosensor
s
FP
CB slider
FP
CB stator
FT
EN
G
PT
FE
Sensor
patch
Flexible
circuitr
y
Fig. 1. Battery-free FWS
3
for wireless and noninvasive molecular monitoring.
(
A
) Schematic illustrating the FWS
3
that integrates human motion energy harvesting,
signal processing, microfluidic-based sweat biosensing, and Bluetooth-based wireless data transmission to a mobile user interface for real-time health status tracking.
(
B
and
C
) Optical images of an FPCB-based FWS
3
, which can be worn on a human side torso. Scale bars, 4
cm. (
D
) Schematic diagram of the FPCB-based FTENG with a
grating slider and an interdigital stator. (
E
) Schematic diagram of the FWS
3
showing a microfluidic-based sweat sensor patch interfacing with the flexible circuitry.
(
F
) System-level block diagram showing the power management, signal transduction, processing, and wireless transmission of the FWS
3
from the FTENG to the bio-
sensors, then to the user interface. Photo credit: Yu Song, California Institute of Technology.
on October 1, 2020
http://advances.sciencemag.org/
Downloaded from
Song
et al
.,
Sci. Adv.
2020;
6
: eaay9842 30 September 2020
SCIENCE ADVANCES
|
RESEARCH ARTICLE
3 of 10
as displayed in fig. S6B.
The scanning electron microscopy (SEM)
images in Fig. 2E and fig. S7 reveal the morphologies of different
triboelectric materials before and after the durability test: The PTFE
shows great mechanical robustness without scratches, while both
M-PDMS and W-PDMS suffer notable surface damage. The in-
fluence of normal force and shear force (reflected by sliding fre-
quency) on the performance of the FTENG during regular use is
demonstrated in fig. S8: The peak output voltage increases and then
saturates with the increase of normal forces, while the output volt-
age remains stable under varied sliding frequencies for a given nor-
mal force. The FTENG is mechanically robust and shows similar
electrical output even under a high normal force of 100 N.
The
FTENG’s response is stable after 1000 bending cycles (radius of cur-
vature, 5 cm) (fig. S9) and under varied physiological temperatures
(fig. S10). Moreover, the FTENG could maintain high performance
after 100 washing cycles, indicating superior wearable capabilities
over conventional TENGs (fig. S11 and table S1). In designing
future TENG-powered devices, it is important to consider factors
such as cost, materials, mechanical properties, and power density.
TENGs commonly prepared through fabric weaving and polymer
coating processes are very low cost but limited by low fabrication
resolution and reproducibility. In contrast, the FPCB-based FTENG
offers a high-resolution, cost-effective, and mechanically robust en-
ergy harvesting solution.
To meet the high energy demands of the wearable sensors, one,
three, and six panels of the FTENGs, designed by taking consider-
ation of the size of human side torso, were further evaluated by ca-
pacitor charging (Fig. 2F). The output of the FTENG was rectified
with a full-wave rectifier. These different FTENG layouts were each
actuated for 30 working cycles to charge capacitors ranging from
10 to 1000
F (Fig. 2G). For the capacitor with 1000
F, 0.03, 0.12,
and 0.19 V could be obtained for one, three, and six panels, respectively,
80
90
100
110
Material
V
/
V
0
(%
)
Initial state
After 20,000 cycles
M-PDMS
W-PDMS
PTFE
0.01
0.11
10
100
0
70
140
210
0.001
0.01
0.1
1
Voltage (V
)
Power
(mW)
04
08
01
20
0.0
0.5
1.0
1.5
2.0
2.5
Time (min)
Voltage (V
)
10
1001
000
0.01
0.1
1
10
100
1-panel
3-panel
6-panel
Voltage (V
)
0.0
0.6
1.2
1.82
.4
–5
0
–2
5
0
25
50
0.5 Hz
1.25 Hz
3.3 Hz
Time (s)
F
Stator
D
E
C
PT
FE
ENIG
Po
lyimid
e
Copper
A
(iii) Final
stat
e
++
+
++
+
++
+++
+
(i)
Initial
stat
e
Slider
Stator
++
+
++
+
++
+++
+
Load
Slidin
g
(ii) Intermediate
state
++
+
++
+
++
++
+
+
Current
B
GH
3-pane
l FTE
NG
& 47-
capac
itor
Slider
ENIG
PoIy
imid
e
ENIG
FP
CB-based
interd
igit
al stator
PTFE
M-PDMS
W-
PDMS
1-panel
6-panel
3-panel
3 / 6 -
panel
1-pane
l
Fig. 2. Characterization of the FTENG for energy harvesting from human motion.
(
A
) Schematic illustration of the working mechanism and charge distribution of the
FTENG. (
B
) Microscopic and optical images of an FPCB-based interdigital stator with an ENIG surface finish on the patterned electrode area. Scale bars, 200
m and 5
mm.
(
C
) Current output of an FTENG under different working frequencies. (
D
) Peak voltage and the corresponding average power of an FTENG at different external load resist
-
ances (
n
=
5). Working frequency, 1.5
Hz. (
E
) Durability of M-PDMS–, W-PDMS–, and PTFE-based stators after 20,000 test cycles.
V
and
V
0
represent the peak open-circuit
voltage after and before the durability test. Inset images, SEM images of M-PDMS, W-PDMS, and PTFE after 20,000 test cycles. Scale bars, 5, 50, and 50
m. (
F
) Schematic
of the flexible FTENG with different stator layouts (one, three, or six panels in parallel) and corresponding slider layouts. (
G
) Comparison of voltages of the capacitors
ranging from 10 to 1000
F charged with one-, three-, and six-panel FTENGs for 30 cycles. (
H
) Long-term stability of a three-panel FTENG in charging a 47-
F capacitor
under a working frequency of 1.5
Hz for 2 hours.
on October 1, 2020
http://advances.sciencemag.org/
Downloaded from
Song
et al
.,
Sci. Adv.
2020;
6
: eaay9842 30 September 2020
SCIENCE ADVANCES
|
RESEARCH ARTICLE
4 of 10
showing great charging capability. At a working frequency of
3.3 Hz, the six-panel FTENG showed a maximum transferred
charges (
SC
) of 15.73
C during one working cycle (fig. S12A).
Meanwhile, fig. S12B portrays the charging-discharging curves of
different capacitors charged to 2 V with a three-panel FTENG at a
working frequency of 2 Hz. The three-panel FTENG, actuated at a
working frequency of 1.5 Hz, was able to repetitively charge a 47-
F
capacitor over a 2-hour duration from 0 to 2 V (Fig. 2H), indicating
high long-term cycling stability. The FTENGs can also be used to
charge capacitors of various capacitances at different cycle lengths
(fig. S13). Depending on the application, connecting several FTENGs
in parallel can be a practical and attractive strategy to greatly in-
crease power output.
Design and
characterization of
microfluidic sweat
sensor patch
Figure 3A depicts the schematic of the dual biosensor array used for
sweat analyte analysis based on ion-selective electrodes (ISEs).
Laser-engraved microfluidic channels are assembled onto the sens-
ing patch. Detailed fabrication procedures are listed in Materials
and Methods and fig. S14. The Ag/AgCl reference electrode is coated
with polyvinyl butyral (PVB) to sustain a steady potential for the
potentiometric measurements of various electrolytes in sweat re-
gardless of solution ionic strength. For pH analysis, the deprotona-
tion of H
+
atoms on the surface of the electrodeposited polyaniline
(PANI) layer is measured as an indicator of bulk H
+
concentration.
Na
+
concentration measurements are facilitated by a thin ion-selective
membrane containing a Na
+
ionophore X and a poly(3,4-ethylene-
dioxythiophene (PEDOT):poly(sodium 4-styrenesulfonate) (PSS)
layer in between the gold electrode layer and sodium ion-selective
membrane as an ion-electron transducer that can minimize the
potential drift of the biosensor. As depicted in Fig. 3 (B and C), the
pH and Na
+
sensors display near-Nernstian sensitivities of 56.28
and 58.63 mV per decade concentration, respectively, in physiolog-
ically relevant pH levels (4 to 8) and Na
+
concentrations (12.5 to 200 mM).
Both sensors show excellent selectivity, reproducibility, and long-
term stability (figs. S15 to S17), and their responses remain stable
under different physiological temperatures (fig. S18), making them
suitable for wearable continuous monitoring.
Laser patterned microfluidic layers were attached onto a poly-
ethylene terephthalate (PET) sensor substrate in a sandwich struc-
ture (medical-tape/PDMS/medical-tape) for controlled and automated
on-body sweat sampling (Fig. 3D and fig. S19). To validate the per-
formance of the microfluidic system, dynamic biosensing was per-
formed during continuous flow injection of Na
+
solutions at varying
physiologically relevant sweat rates (1, 2, and 4
l min
−1
) (Fig. 3E).
When the Na
+
concentration was switched from 50 to 200 mM at a
flow rate of 2
l min
−1
, the Na
+
sensor took ~2 min to reach new
stable readings. The high temporal resolution is repeatable over
multiple concentration change cycles (Fig. 3F). The flexible micro-
fluidic sensor patch can conformally adhere to human skin (Fig. 3G)
and shows excellent mechanical stability through rigorous bending
tests (radius of curvature, 2 cm), indicating their potential for wear-
able applications in various physical activities (Fig. 3, H and I).
System-level integration for
energy management and
low-power biosensing
As aforementioned, the FWS
3
consists of an interdigital FENG
stator, a PMIC, a low-dropout voltage regulator, two low-power
instrumentation amplifiers, and a BLE PSoC module seamlessly
integrated onto a polyimide-based FPCB.
In addition, the full plat
-
form requires a grating-patterned FTENG slider and a microfluidic
sensor patch. For design compatibility and flexibility, the FTENG
and electronic circuitry were designed on a single PCB design soft-
ware. The detailed part list and circuit diagram for the flexible cir-
cuitry are shown in figs. S20 and S21, respectively. A block diagram
illustrating the electrical connections between the modules is shown
in Fig. 4A. For optimal power management, a commercial energy
harvesting PMIC was adopted to manage power generated by the
FTENG with minimal power waste. With the aid of a bridge rectifi-
er that converts the high voltage AC signal generated by the FTENG
into a DC signal, the PMIC stores the FTENG generated power in
two capacitors in parallel (220 and 22
F). The three SET_
V
OUT
resistors set the programmable threshold and hysteresis voltages
such that the stored power is released only when absolutely neces-
sary through built-in switch control logic. When the voltage of the
storage capacitors (
V
STORE
) reaches 3.5 V, the capacitors supply en-
ergy to the load/output (
V
OUT
) until
V
STORE
decreases to 2.2 V.
At
2.2 V, the PMIC’s control unit disconnects the storage capacitors
from the load/output until the storage capacitors are charged back
to 3.5 V.
When supplied by the storage capacitors, the load/output
voltage is regulated to 2.2 V by a voltage regulator to provide a stab
le
voltage for the precise measurement circuitry.
Efficient power management is matched with low-power mea-
surement via low-power instrumentation amplifiers with shutdown
modes, and low-power data transmission via connectionless BLE
advertisements to enable FTENG-powered wearable and wireless
sweat analysis. Every time the storage capacitor is charged to 3.5 V,
the BLE PSoC module initiates one ~510-ms operation cycle as por-
trayed in the flow diagram (Fig. 4B). After startup of the main pro-
cessor, the PSoC pulls a general-purpose input/output (GPIO) pin
high to wake up the two instrumentation amplifiers from shutdown.
After initialization of the instrumentation amplifiers, the PSoC’s
embedded 12-bit ADC (analog-to-digital converter) samples and
averages 32 potentiometric measurements acquired through the
instrumentation amplifiers. After ADC measurements, the instru-
mentation amplifiers are shut down to minimize power consump-
tion. The BLE submodule of the PSoC requires the 32-kHz watch
crystal oscillator (WCO) for accurate operation, which has a maxi-
mum startup time specification of 500 ms. Therefore, after the ADC
measurements, the main processor of the PSoC starts the WCO and
goes into deep sleep for 500 ms where it consumes ~2
A.
Then, the
BLE stack initializes and the ADC measurements are advertised to a
nearby
BLE observer user device. The detailed power consumption
breakdown of the circuit including the voltage regulator, BLE PSoC
module, and two instrumentation amplifiers is shown in Fig. 4C.
When supplied with 2.2 V, the circuit consumes an average of 330
A
during ~510 ms (168
C).
Several studies were performed to validate the robust operation
of the fully integrated system. The three-panel FTENGs were acti-
vated via sliding motion at frequencies ranging from 2 to 1 Hz to
simulate human arm swing during exercise (
40
). The resultant
charging and discharging cycles of the storage capacitors are shown
in Fig. 4D. Furthermore, to validate the operation of the low-power
wireless sensor circuitry, the potentiometric inputs were simulated
by using a DC power supply to apply voltages ranging from 100 to
300 mV (charged every 300 s) across the reference and working
electrode pins (Fig. 4E). These simulated sensor inputs were accurately
on October 1, 2020
http://advances.sciencemag.org/
Downloaded from
Song
et al
.,
Sci. Adv.
2020;
6
: eaay9842 30 September 2020
SCIENCE ADVANCES
|
RESEARCH ARTICLE
5 of 10
measured and transmitted by the FPCB platform while being pow-
ered by a three-panel FTENG actuated at various working frequencies.
The long-term stability of the whole FWS
3
system is demonstrated
by using the FTENG to power the FPCB for more than 4 hours,
during which pH and Na
+
concentrations in collected human sweat
were measured for an hour (Fig. 4, F and G). Furthermore, the long-
term durability of the FPCB-based FTENG was tested by compar-
ing its ability to power the whole platform after 1 month of its initial
use (fig. S22). Improvements in wireless data transmission in terms
of transmission interval can be attained by further advancing the
power density and efficiency of the FTENG.
On-body evaluation of
the
FWS
3
in
prolonged
physical activities
Common cardiovascular exercises such as running, rowing, and
elliptical training induce a sliding motion between the side of the
torso and the inner arm. Taking advantage of this mechanical
motion, the stator of FTENG can be fixed on the side torso, and the
slider of FTENG can be attached to the inner part of the arm. For
on-body evaluation, a six-panel stator FTENG-based FWS
3
was
used for increased power output (as illustrated in fig. S23). The
FTENG power output waveforms during various exercises are
shown in Fig. 5A. Treadmill running was chosen as the exercise to
perform on-body validation experiments for the entire system. The
FPCB’s storage capacitor charging and discharging curve during a
60-min constant speed running session shows that up to 18 operation
cycles can be achieved (Fig. 5B). The length of charging/discharging
cycle ranges from 2.1 to 3.7 min (Fig. 5C). It should be noted that
the system generates power when the stator and slider physically
rub against each other; the charge in the capacitor will accumulate
without discharge whenever there is rubbing motion; when the
capacitor is charged to a threshold voltage, the capacitor will dis-
charge and power a single measurement event. Despite duration
variations in capacitor charging/discharging cycles caused by varied
rubbing area, forces, and frequencies, the FWS
3
system proved itself
to be fully functional during normal physical exercises (fig. S24).
02468
10
12
0.20
0.24
0.28
0.32
[N
a
+
]
pH
Potential (V)
Time (min
)
02468
10
12
0.20
0.24
0.28
0.32
[N
a
+
]
pH
Potential (V)
Time (min)
0
100
2003
00
400
0.27
0.29
0.31
0.33
–1
–1
–1
Potential (V)
Time (s)
45678
0.0
0.1
0.2
0.3
Potential
(
V
)
pH
03
0609
0
120
150
0.0
0.1
0.2
0.3
Potential (V)
Time (s)
03
0609
0
120
150
0.24
0.27
0.30
0.33
Potential (V)
Time (s)
1.0
1.5
2.0
2.5
0.24
0.27
0.30
0.33
Potential
(
V
)
10
[[Na
+
] (mM)]
0
600
1200
1800
0.27
0.28
0.29
0.30
0.31
0.32
Potential (V)
Time (s)
G
D
A
H
B
F
E
I
C
pH
4
5
6
7
8
[Na
+
] (mM)
200
100
50
25
12.
5
200 cycles
200 cycle
s
100
mM
4
5
50
mM
100
mM
4
5
50
mM
[Na
+
] (mM)
50
200
Ref.
pH
Na
+
Biosensor
arra
y
[Na
+
] (mM)
50
200
To
flexible
circuitr
y
Ref
.
N
a
Outl
et
PEDOT
ISE
Sw
eat
2
min
–1
M-
tape
(sw
eat inle
ts)
PDMS
(channels)
M-
tape
(sensors)
Microf
luidic
sensor patch
Fig. 3. Characterization of the microfluidic-based biosensor patch in
vitro.
(
A
) Schematic of a flexible biosensor array containing a pH sensor and a Na
+
sensor pat-
terned on a flexible PET substrate. (
B
and
C
) Open-circuit potential responses of a pH sensor in standard Mcllvaine’s buffer solutions (B) and a Na
+
sensor in NaCl solutions
(C). Insets show the corresponding calibration plots of each sensor. Error bars represent the SDs from six independent tests. (
D
) Schematic of the microfluidic design for
dynamic sweat sampling. M-tape, medical tape. (
E
) Dynamic response of a Na
+
sensor under different flow rates upon switching the solution concentrations. (
F
) Repeat-
ability of dynamic response of a Na
+
sensor by continuously switching the inflow solutions at a flow rate of 2
l min
−1
. (
G
) Schematic of microfluidic sensor patch con-
formally attached to human skin. Inset image, optical image of a microfluidic sensor patch under mechanical deformation. Scale bar, 5
mm. (
H
and
I
) Responses of a Na
+
and pH sensor array after 0, 200, and 400 cycles of bending (H) and during bending states (I) (radius of curvature, 2 cm). Data recording was paused for 30 s to change the
conditions and settings. Photo credit: Yu Song, California Institute of Technology.
on October 1, 2020
http://advances.sciencemag.org/
Downloaded from