A Sensitive and Biodegradable Pressure Sensor Array for

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A Sensitive and Biodegradable Pressure Sensor Array for
Cardiovascular Monitoring
Clementine M. Boutry, Amanda Nguyen, Qudus Omotayo Lawal, Alex Chortos,
Simon Rondeau-Gagné, and Zhenan Bao*
A new class of sensors made from completely biodegradable
materials is currently emerging as environmentally friendly
and biocompatible with the human body.[1–4] These devices
improve the efficacy of implantable systems and reduce waste
associated with wearable and point-of-care diagnostics. Highly
sensitive pressure sensors can measure pulse, arterial waveform, and pulse wave velocity (PWV), which provide important
information about cardiac health.[5–13] In this work, we report
the development of a single-use biodegradable pressure sensor
patch for application in cardiovascular monitoring. The sensors
have a very high sensitivity of 0.76 kPa−1 and are capable of distinguishing a weight of 5 mg. Each patch contains an array of
sensors to facilitate fast and easy positioning for cardiac health
measurements.
Only a single example of biodegradable pressure sensor was
reported so far by Luo et al.[14] in which the microfabricated
wireless RF pressure sensor consisted of a parallel-plate capacitor filled with air, connected to an inductor coil.[14] Despite its
simplicity, this air gap approach suffers from poor sensitivity
in the low pressure regime (<≈10 kPa).[15–17] Instead, the biodegradable and biocompatible pressure sensor described in
this work was designed using the microstructured dielectric
approach that was previously described by our group.[17–19] The
dielectric consists of a thin, regularly microstructured rubber
that is highly compressible, which enables the device to have
high pressure sensitivity and fast response time, which are
required to measure pulse waveforms and produce accurate
PWV measurements.
A schematic image of the device is shown in Figure 1a. It
consists of a flexible capacitor, with two substrates assembled
via lamination. The key element of our material design is
the elastic dielectric between the top and bottom electrodes.
Poly(glycerol sebacate) (PGS) is a biodegradable elastomer first
reported by Wang et al. for tissue engineering applications.[20]
It is an excellent candidate for this application in terms of the
degradation characteristics, the mechanical properties, and the
processing characteristics. The degradation of PGS has been
confirmed and characterized in a range of different conditions.
It has been approved by the US Food and Drug Administration (FDA) for biomedical use, which reduces the timeline for
Dr. C. M. Boutry, A. Nguyen, Q. O. Lawal,
A. Chortos, Dr. S. Rondeau-Gagné, Prof. Z. Bao
Shriram Center Chemical Engineering
Stanford University
443 Via Ortega, Stanford, CA 94305-4125, USA
E-mail: [email protected]
DOI: 10.1002/adma.201502535
Adv. Mater. 2015,
DOI: 10.1002/adma.201502535
implementation. Furthermore, both reactants used in the synthesis of PGS (sebacic acid and glycerol) are inexpensive, which
is favorable for large-scale and low-cost applications.[21,22] The
mechanical properties of PGS can be tuned by controlling
various parameters such as the reaction and curing temperature and time. The elongation at break is larger than 260% and
the Young's modulus can be controlled within the range of
0.05–2 MPa, allowing tunable pressure sensor characteristics,
i.e., the sensor sensitivity can be adjusted to fit the pressure
range of interest for a particular application. In comparison, the
modulus of PDMS used in previous reports[17,18] on high sensitivity pressure sensors is ≈1 MPa. The similar modulus for PGS
ensures similar high-performance sensing characteristics. We
show that PGS has minimal viscoelastic response, which enables fast time response and good reproducibility of the devices
with cycling. The simple synthesis mechanism allows PGS to be
readily molded into structured dielectrics. Figure 1b shows an
scanning electon microscope (SEM) image of the microstructured surface of PGS that was fabricated by molding a 2D array
of square pyramids into the PGS prepolymer film before the
polymer was physically cross-linked. Pyramids are reproduced
with high quality even on very thin (<100 µm) and highly flexible PGS sheets. These microstructures allow the PGS dielectric to elastically deform when an external pressure is applied,
storing and releasing the energy reversibly, thus minimizing
undesirable viscoelastic behavior.[17] For the electrodes, several
biocompatible corrodible metals can potentially be selected.
Pure Mg, pure Fe, and their alloys have already been in use for
more than a decade in clinical trials in applications such as biodegradable stents and bone scaffolds.[23,24] In addition to investigations on dissolution rate, change in electrical properties,
chemical properties, and morphology, Mg, Fe, and the common
used alloying elements are also being investigated from the
point of view of pathophysiology, toxicology, and daily allowance in the body, providing design rules for human safety: in
particular Mg and Fe are essential nutrients with a daily allowance of 0.7 g and 10–20 mg, respectively.[25] In previous works,
resistor–inductor–capacitor (RLC) resonators made of Mg, Fe,
and their alloys were investigated as RF components for wireless biosensor applications.[4] Luo et al. used electrodeposited
Zn/Fe bilayers as conductor material for their sensor,[14] while
Yin et al., investigated the dissolution properties of thin films
of Mg, Mg alloy, Zn, Fe, W, and Mo, to assess their potential for
use in water-soluble, physically transient electronics.[26] In this
project, Mg is used for the top and bottom electrodes due to its
easy processing, biocompatibility, and rapid rates of hydrolysis.
A thin layer of Fe is used as an adhesion layer to promote good
adhesion of Mg with the biodegradable polymer substrate.
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Figure 1. Schematic design and fabrication of a fully biodegradable and flexible pressure sensor array from microstructured poly(glycerol sebacate)
PGS films. a) Schematic of the final fabrication step of our fully biodegradable pressure sensor and its device structure. b) SEM images of the microstructured PGS films. 2D arrays of square pyramids are formed into PGS from a PDMS mold. c) PGS biodegradable elastomer film.
The sensitivity of the biodegradable pressure sensor is investigated for a 4 × 5 sensor array, in which the sensor area is
defined by the intersection of two orthogonal electrodes (2.42 ×
2.42 mm2). The pyramids on the surface of PGS have a height
of 2 µm, a pyramid base width of 4 µm, and a spacing between
pyramids of 5.6 µm (Figure 1b). Figure 2a shows the pressure
response from three sensors. The pressure sensitivity S is
defined as the slope of the traces in Figure 2a (S = δ(∆C/C0)/δp,
where C and C0 are the capacitances with and without applied
pressure, and p is the applied pressure). The pressure sensitivity, averaged over 20 sensors, is 0.76 ± 0.14 kPa−1 in low
pressure regime (p < 2 kPa) and 0.11 ± 0.07 kPa−1 at higher
pressures (2 < p < 10 kPa). These values are in good agreement
with results previously reported by our group for nonresorbable
PDMS-based sensors with a similar design.[17] Moreover, these
pressure sensitivities are ≈60 times higher (10 < p < 20 kPa),
two orders of magnitude higher (×180, 2 < p < 5 kPa), and
three orders of magnitude higher (×1260, p < 2 kPa) as compared to previously published biodegradable pressure sensors
(0.0006 kPa−1, 0 < p < 20 kPa).[14] Our obtained high sensitivity
should enable new applications, especially in the field of cardiovascular monitoring. And for applications where a high sensitivity is required at higher pressures, the design can be adapted
with larger pyramids, as described by Tee et al.[27] There is nonnegligible variability from one sensor to the other within the
array, as illustrated by: (i) the standard deviations associated
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with pressure sensitivities (0.76 ± 0.14 and 0.11 ± 0.07 kPa−1),
and (ii) the variations of C0 (6.22 ± 2.05 pF) when comparing
20 sensors in the array. However, in the context of cardiovascular monitoring, this variability between sensors is not as critical since the signal of interest is the shape of the pulse wave,
which is measured as a relative variation of pressure. Cardiologists are usually more interested in peaks location and relative
peak area, together with specific time markers such as the foot
of the wave used to calculate the pulse wave velocity (PWV).[7]
In this work, the differences between sensors can be explained
by: (i) slight variations in the thickness of the PGS at different
locations in the array, (ii) slight variations in the area of capacitors parallel plates (the metallization lines of top and bottom
electrodes are defined with a shadow mask), and (iii) additional
air gaps that form between the top electrode and the PGS dielectric. The presence of the air gaps, due to small curvatures in
the flexible substrate, was previously confirmed by comparing
the experimental results with calculated values.[18]
In Figure 2a, there is a reduction in sensitivity for pressures
p > 2 kPa. Our current sensor has a favorable sensing response
as it shows a progressive damping in sensitivity, which will further extend the range of detectable pressures to higher loads
in which high sensitivity is not required, hence resulting in a
more versatile pressure sensor.[17] The sensitivity and sensing
range can easily be adjusted by choosing the appropriate pyramid height and density on the PGS film.[18] Another important
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Figure 2. Response characteristics of the pressure sensor. a) A sensor array of 4 × 5 elements is characterized by measuring the pressure response
curve of the 20 individual sensor elements of the array (curves for three elements are displayed). The average sensitivity and standard deviation over
all the 20 sensors of the array are calculated for the pressure ranges p < 2 kPa and 2 kPa < p < 10 kPa. A small glass plate (208 mm2) is placed between
the force gauge tip and the array surface. It allows a uniform repartition of the pressure over all the arrays.[17] b) Response and relaxation times of
sensors with structured and unstructured films (PGS with and without pyramids). When a constant pressure is applied (here 20 and 80 kPa for about
8 s), both structured and unstructured PGS films show immediate response to the application of pressure. However, after the pressure is released, the
response time to return to the original baseline is much lower for structured PGS films as compared to unstructured films, being in the millisecond
range instead of tens of seconds. c) An enlarged view on sensors response to pressure release.
benefit of the molded dielectric is that the response time is
greatly improved for structured PGS films as compared to
unstructured films, being in the millisecond range instead of
a few seconds, respectively (Figure 2b). Moreover, the pressure response curves of the sensors show minimal hysteresis
(Figure S1, Supporting Information) and can be reproducibly
cycled thousands of times (Figure S2, Supporting Information).
After applying a constant pressure of 80 kPa and releasing the
force for more than 8000 cycles (duration of one cycle: 4 s),
C0 and Cmax increase by 16% and 4%, respectively, while the
variation of pressure sensitivity is less than 10% (Figure S2,
Supporting Information). The robustness of the sensor when
applying high pressures is illustrated in Figure S3 (Supporting
Information). It is shown that for several consecutive runs of
pressures applied in the range 0–270 kPa, the sensor output is
highly reproducible and the characteristic of the device is not
altered by applying such high pressures.
To demonstrate the ability of the sensors for object mapping, a sensor array is fabricated and illustrated in Figure 3a.
Such a sensor array is used to detect the presence of small
weights, e.g., a single grain of salt (weight: 5 mg), with excellent time response. Similarly, the sensor array is able to detect
and localize very light objects, e.g., a grain of rice or a single
small insect (weight of 21.8 and 22.5 mg, respectively), which
corresponds to very low pressures down to 3 Pa. In Figure 3b
(right), the response of the device to the placement of a bipod
object (weight 5 g, Dagoma 3D printing), illustrates the crosstalk effect between several pixels located close-by in the array.
While this does not constitute a problem in the present context
of arterial pulse wave velocity measurement, this effect might
be undesirable in other applications, such as electronic skin,
and could be circumvented by implementing stretchable top
and bottom electrodes.[19]
The sensor performance is also investigated while it is bent
at various convex curvatures. As illustrated in Figure S4a (Supporting Information), the pressure response is very stable at all
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bending radii down to 27 mm, which is sufficient in the context of external cardiovascular monitoring (e.g., if the device is
attached at the wrist, or on the skin above the femoral or carotid
artery). The pressure sensitivity has a slight decrease for applied
pressures in the ranges 5–10 and 0.1–2 kPa when maintained at
a bending radius of 27 mm (Figure S4b, Supporting Information). This is ≈10% and ≈20% lower, respectively, as compared
to operation without any bending. At pressures p >10 kPa, no
significant change in the pressure sensitivity is observed. As
previously described, in the context of the present application,
such a change in sensitivity is not especially critical, since the
signal of interest is the relative variation of pressure over time,
i.e., arterial pulse wave shape.[7]
In vitro biodegradation studies are performed to further
investigate device resorption characteristics. The objective is to
demonstrate that these devices will indeed degrade (if needed)
after their period of usage. The mass-loss kinetics of the device,
which is composed primarily of the polyhydroxybutyrate/polyhydroxyvalerate (PHB/PHV) top and bottom substrates with
metallization layers and PGS dielectric, is shown in Figure 4a.
After being immersed for seven weeks in a phosphate-buffered
saline (PBS) solution at 37 °C, the pressure sensor weight
remains at ≈85% of its initial value. As shown in Figure 4b, the
first layer to be dissolved is the metal layer (Fe Mg), in which
the degradation occurs as soon as the PBS buffer penetrates
into the PHB/PHV bonded substrates. The complete resorption of both PHB/PHV and PGS films requires a duration of
at least a few months and depends greatly on its biological
environment.[28,29] For PGS, Wang et al. have shown that films
immersed in a PBS solution at 37 °C lost 17.6% of their dry
weight after 60 days.[20] The degradation rate was found to be
higher in vivo, with surface erosion being typically in the range
of 0.2–1.5 mm month−1.[28] For PHB/PHV, Pouton and Akhtar
have shown that the complete weight loss of a 50 µm thick
PHB/PHV film happened after ten weeks in soil at 25 °C, and
after 50 weeks in sea water at 15 °C.[30]
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Figure 3. Pressure sensor array performances. a) The sensor array of 4 × 5 pressure-sensitive elements is able to detect the presence of small weights
such as a grain of salt with fast time response (weights: 55, 9, and 5 mg). b) Capacitance change distribution on the sensor array upon placing light
weights such as a bee (22.5 mg), a grain of rice (21.8 mg), and a bipod object (5 g). The array allows to localize and sense the application of pressures
as small as 3 Pa. The measurement with bipod object illustrates the crosstalk effect between several pixels located close-by in the array, which could
be minimized in the future by designing stretchable top and bottom electrodes.
In the present study, we are also particularly interested in
the mechanical properties of the PGS dielectric layer upon
degradation in order to understand how well this key element
of the sensor would retain its performance after prolonged
exposure to humid environments, e.g., during transport or
storage. It is also desirable that the elastomer has very limited
viscoelastic behavior and hysteresis, even after being exposed
to a degrading environment. Figure S5 (Supporting Information) shows stress–strain curves for PGS. It is observed that
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very little hysteresis is present both in tension and compression regime, even after seven weeks in PBS solution at 37 °C.
The tensile modulus and compression modulus of PGS as a
function of incubation duration are given in Figure S6 (Supporting Information). It is observed that both moduli decrease
when the incubation duration is increased. A simple linear fit
indicates a reduction of about 10% and 4% week−1 for the tensile and compression moduli, respectively. Modulus degradation will affect the device performance mainly by changing the
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Figure 4. In vitro degradation of pressure sensor and PGS dielectric layer. a) A plot of mass remaining demonstrates that the biodegradable pressure
sensor starts to lose weight already after one week being immersed in phosphate-buffered saline (PBS) solution at 37 °C, the weight being ≈85% of
its initial value after seven weeks of incubation (study performed on four devices). b) Photographs from a representative device before in vitro degradation and after seven weeks of incubation. The first layer to be dissolved is the metal layer (Fe Mg). The degradation occurs as soon as the PBS
solution penetrates through the PHB/PHV bonded substrates. c) In vitro degradation study of PGS, where the viscoelastic properties are investigated
through relaxation tests. The samples were immersed in phosphate buffered saline solution (pH 7.4, Medicago, USA) at 37 °C. The PBS solutions
were renewed every week to prevent significant changes of pH. After 1, 2, 3, 4, 5, 6, and 7 weeks, seven PGS cylinders were removed from incubator,
washed in deionized water, dried with towel and air jet, and dried in vacuum during 24 h. For each cylinder, a relaxation test was performed at five
different strain levels (ε = 0.05–0.25), and the Findley model was used to extract the initial load F0 and the relaxation coefficient B from the measured
relaxation curves. d) Relaxation coefficient B and e) initial load F0 as a function of the incubation duration.
output response sensitivity. Additionally, the viscoelasticity of
PGS is investigated by relaxation measurements, as illustrated
in Figure 4c,d. The measured load curves are evaluated using
the modified superposition model of Findley adapted from the
Fung model[31,32] (Figure 4c). The nonlinear viscoelastic model
of Findley was chosen because it is observed in the measurements that the relaxation function depends on the strain level.
This nonlinear behavior requires a more general description
than the separable quasilinear viscoelasticity theory.[32] In the
model of Findley, the relaxation depends on time and strain
level as shown below in Equation (1)[31]
F(ε , t ) = F0 (ε )t −B( ε )
(1)
where F is the time and strain-dependent load, F0 is the initial
load (load immediately measured when the strain is applied),
and B is the relaxation coefficient that describes the straindependent rate of load relaxation. The averaged B and F0 for
PGS samples as a function of incubation duration are plotted in
Figure 4d,e. After seven weeks, a slight decrease is observed for
F0 (rate −6% week−1), while no significant change is observed
for B. This observation indicates that for samples evaluated
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after one week of incubation (under a constant applied strain of
0.05, which corresponds to an applied pressure of 34 kPa), the
measured initial load is around 2 N, and it remains at 90% of
this initial value after 3 min. After seven weeks of incubation,
the measured initial load is around 1.4 N and again it is ≈90%
of this initial value after 3 min. In conclusion, our in vitro biodegradation study of PGS shows that there is no significant
change in the viscoelastic behavior of PGS after seven weeks
of incubation in PBS solution. This is attributed to the fact that
PGS undergoes surface erosion, in which the main mechanism of degradation is the cleavage of ester linkages. Unlike
bulk degradation mechanisms present in most biodegradable
polymers, PGS is characterized by a slow loss of mechanical
strength relative to mass loss,[28] which is an asset to our present sensor application.
The storage stability is tested for pressure sensors that are
stored for one week in the following two conditions: (1) Storage
in an oven at a temperature of 45 °C, a large open water container placed in the oven allowing reaching 40% relative
humidity and (2) storage in a freezer at a temperature of −6 °C
with 56% relative humidity. After 7 d storage, the response
characteristics of the pressure sensors are measured and
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compared to characteristics prior storage (see Figure S7, Supporting Information). As shown in Figure S7a,b (Supporting
Information), the sensitivities at low and high pressure regime
are not affected by the harsh storing conditions, and as illustrated in Figure S7c (Supporting Information), the sensors are
still able to successfully measure the wrist pulse wave after one
week storage in extreme conditions.
The high sensitivity and fast response time of the sensors
make them suitable for cardiovascular monitoring such as
arterial tonometry and PWV measurements. A critical, global
concern is the prevalence of cardiovascular diseases (CVD),
which is the primary cause of premature death in the world
today,[5] and hypertension was identified as the main risk factor
and leading cause of CVD.[6] The early detection and regular
monitoring of hypertension patients improves the prognosis
and reduces societal cost.[5,6] Measurement of carotid-femoral
pulse wave velocity (cf-PWV) is currently the “gold standard”
to measure aortic arterial stiffness, which is directly correlated
to hypertension.[7,8] Pulse wave velocity is the speed at which
the blood pulse wave travels in arteries.[9–12] In elastic (healthy)
arteries, the blood pulse wave travels slowly, while in stiff
arteries (arteries tend to become stiffer with increasing age and
diseases, e.g., hypertension), the pulse waves will travel faster.
Due to its ability to independently predict cardiovascular risk
and mortality,[9,11] cf-PWV measurement is a reference in international cardiology/hypertension treatment guidelines.[7,8] A
major challenge in developing countries[5,13] is the cost-effective
screening of large populations in order to diagnose hypertensive patients early. Having a low-cost single-use biodegradable
sensor patch for point-of-care measurement of pulse wave
signal would be an asset for improving patient throughput and
environmental impact.
To demonstrate its high functionality and fidelity for PWV
measurements, we fixed our flexible pressure sensor to an
adult human wrist, above the radial artery, as usually done
in arterial tonometry (Figure 5). Figure 5a shows a real-time
record over several pulse periods, together with a zoomed view
on a pulse having the characteristic peaks typically measured at
the radial artery.[32] The location of the pulse wave foot—that is
commonly used as a time reference for PWV measurements—
is also indicated.[7] Additionally, pulse wave velocity measurements are successfully performed using the same recording
setup as above. As illustrated in Figure 5b, the pressure sensor
is first attached above the carotid artery of an adult human, and
the arterial pulse wave is recorded. Furthermore, the ECG is
simultaneously recorded, being used for time reference. The
pressure sensor is then placed above the femoral artery and the
pulse wave and ECG are both again recorded. We then calculate the PVW to be 7.5 via: (i) the distance between the neck
and the groin, according to clinical practice, and (ii) the time
delay between the feet of the carotid and femoral pulse signals
(Figure 5b), in which this obtained result is typical for a healthy
subject.[7]
Our single-use sensors are intended to be placed on
human skin for typically several tens of minutes in a relatively dry environment for cardiovascular monitoring. However, it is important to verify that the device can withstand a
short period of exposure to sweat in case of sweaty patient.
As shown in Figure S8 (Supporting Information), the wrist
pulse wave is successfully measured in all sweating conditions (Figure S8a–d, Supporting Information), even when the
skin is entirely covered with artificial sweat (Artificial Eccrine
Perspiration 1700-0020, Pickering). Moreover, the PHB/PHV
top and bottom substrates offer an efficient protection against
the corrosion of the metallization layers. This is illustrated in
Figure S8e (Supporting Information), where a droplet of artificial sweat covers all the measurement area on top of the sensor.
The sensor output is recorded for several tens of minutes, and
the measured capacitance stays stable over time with less than
1% standard deviation.
In summary, we have described the design, fabrication, and
characterization of a highly sensitive flexible pressure sensor
entirely made of biocompatible and biodegradable materials.
Our device’s high sensitivity in the low pressure regime (0.76 ±
0.14 kPa−1 at p < 2 kPa and 0.11 ± 0.07 kPa−1 at 2 < p < 10 kPa),
in combination with its fast response time in the millisecond
Figure 5. Blood pulse wave of the radial, femoral, and carotid artery measured with our pressure sensor and pulse wave velocity measurement. a) Realtime transient signal recorded at the radial artery. The pressure sensor is connected to a custom-built wireless capacitance measurement system. A
24 bit capacitance to digital converter (Analog Devices AD7746) is used to convert sensor capacitance to a digital signal. The signal is sampled at
90 Hz and collected via an Atmega 328u4 microcontroller. The data are wirelessly transmitted to a computer for processing and display via an XBee
Series 2 radio module. b) Pulse wave velocity measurements performed using the same recording setup as above, combined with ECG measurements
for time reference. The real-time transient signals are recorded at the femoral and carotid artery.
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Experimental Section
Sensor Fabrication: PGS was synthesized based on previously
published methods,[20] where an equimolar mixture of glycerol and
sebacic acid was reacted at 120 °C under nitrogen for 24 h, resulting
in a viscous uncross-linked PGS prepolymer. This prepolymer was
further polymerized in oven at 130 °C in vacuum for 24 h. Afterward,
the highly viscous prepolymer was cured at 130 °C for 24 h between a
microstructured PDMS mold and a bare silicon wafer, both treated with
a nonadhesive layer evaporated in vacuum (trichlorosilane, Gelest, USA).
The PDMS mold (Sylgard 184 Silicone Elastomer, ratio 5:1) contains
2D arrays of square pyramids that are formed from arrays of pyramidal
recesses etched into the face of a (100)-cut Si-wafer mold, as described
in ref. [1]. After being cured at 130 °C in vacuum for 15 h, the ≈150 µm
thick PGS film was peeled off and laminated with the bottom and top
electrodes. The electrodes were fabricated by casting a thin polyvinyl
alcohol (PVA) adhesive layer (40-88, Fluka Analytical, USA) on top of a
25 µm thick PHB/PHV film (8%, Goodfellow, USA), and evaporating iron
(2 nm thick adhesive layer) and magnesium (100 µm thick electrode)
after having exposed the substrate surface to oxygen plasma. The sensor
was assembled by laminating the bottom electrode, intermediate PGS
dielectric layer, and top electrode in vacuum. The PHB/PHV top and
bottom electrodes were sealed previous to in vitro degradation study
using a nonsolvent assisted thermal bonding process.[38] For bonding,
the sensor was placed in a custom-built mold made of PDMS (to
equalize the applied pressure) and 50 µm thick Kapton polyimide films
(to prevent sticking, Katco, UK) and sealed in a hydraulic press (Carver,
USA, temperature: 155 °C, duration: 10 min, applied pressure: 1 MPa).
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Pressure Response Measurement Setup: The measurement setup
consisted in a motorized vertical stage used in combination with a force
gauge, while the capacitance of each sensor was measured with an LCR
meter.
In Vitro Study Cylinder and Stripe Samples: A PGS film was polymerized
and cut with laser (Versa Laser, USA), in order to produce PGS cylinder
samples (diameter 7 mm, height 2.4 mm) and PGS stripe samples
(width 4.4 mm, length 5 mm, thickness 2.4 mm).
Mechanical Tests: All tensile, compression, and viscoelastic
mechanical tests were performed with an Instron 5565 Mechanical
tester, used in compression and tension modes. Measurements were
performed in controlled temperature and humidity atmosphere at 23 ±
1 °C and 50 ± 10% relative humidity. For tensile tests the stripes were
attached between two mechanical grips, while for compression tests the
cylinders were compressed between parallel platforms using a lubricant
between the sample and top/bottom plates to ensure proper mechanical
deformation of the samples. For all experiments, the strain rate was
1 mm/mm min−1.
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range, is unprecedented when compared to recent reports on
biodegradable pressure sensors. The pressure sensor performs
stably after thousands of cycles. We also demonstrate a sensor
array of 4 × 5 elements that allows the detection and localization of very light objects such as a grain of salt. An in vitro biodegradation study of PGS shows that this elastomer has a small
hysteresis and little viscoelasticity behavior, with no significant
changes even after seven weeks of incubation in a PBS buffer
solution. These characteristics allow our pressure sensor to be
useful for continuous cardiovascular monitoring, as illustrated
by the acquisition of blood pulse wave signal at the human
radial, carotid, and femoral arteries. The sensor is successfully
employed to assess the pulse wave velocity, the gold standard
for measuring subject’s aortic arterial stiffness and therefore
diagnosing hypertension.[7,8] This sensor represents the first
step toward realizing more complex biodegradable sensors that
could potentially be used inside the body for biomedical applications such as pressure-assisted wound healing,[33] pressureassisted hole drilling in the context of deep brain stimulation
electrode implantation,[34] intracranial pressure monitoring
after traumatic brain injury,[35] or monitoring heart tissues pulsatile movements with sensor arrays combined with cardiac
patches for the treatment of myocardial infarctions.[36,37] Developing electrodes made of biodegradable metals with appropriate degradation rates and tailoring the packaging of these
structures would permit preprogramming device lifetimes by
temporally mediating exposure of critical device components,
e.g., the metallization layers, to corrosive environments.
We envision that this work will contribute to fully biodegradable sensors for biomedical applications inside the body, with
high pressure and strain sensitivity, intrinsic stretchability,
and capable of distinguishing between pressure, strain, and
bending.
Supporting Information
Supporting Information is available from the Wiley Online Library or
from the author. A color version of Figure 1 to Figure 5 is also provided
in Supporting Information.
Acknowledgements
The authors thank B. Schroeder for support with elastomer
polymerization process, and J. B. Tok for help with paper proofreading.
C.M.B. acknowledges postdoctoral fellowship support by the Swiss
National Science Foundation (postdoc mobility fellowship No.
P2EZP2_152118) and the European Commission (Marie-Curie
international outgoing fellowship Grant No. 622362). Part of this work
was performed at the Stanford Nano Shared Facilities (SNSF). Z.B.
acknowledges funding support from the MSIP (Ministry of Science,
ICT and Future Planning), Korea, under the “IT Consilience Creative
Program” (NIPA-2014-H0201-14-1001) supervised by the NIPA
(National IT Industry Promotion Agency).
Prior to the beginning of the work described in this manuscript, an
IRB review waiver was obtained from the Stanford institutional review
board. The project does not involve human subjects as defined by
federal guidelines. The work is not about living individuals, only involves
the measurement of pulse waveforms, and does not include collection
of individually identifiable private information.
Received: May 27, 2015
Revised: August 6, 2015
Published online:
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DOI: 10.1002/adma.201502535