Self‐Powered Real‐Time Arterial Pulse Monitoring Using...

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Self-Powered Real-Time Arterial Pulse Monitoring Using Ultrathin Epidermal Piezoelectric Sensors

Dae Yong Park, Daniel J. Joe, Dong Hyun Kim, Hyewon Park, Jae Hyun Han, Chang Kyu Jeong, Hyelim Park, Jung Gyu Park, Boyoung Joung, and Keon Jae Lee*

D. Y. Park, Dr. D. J. Joe, D. H. Kim, J. H. Han, Prof. K. J. LeeDepartment of Materials Science and EngineeringKorea Advanced Institute of Science and Technology (KAIST)291 Daehak-ro, Yuseong-gu, Daejeon 34141, Republic of KoreaE-mail: [email protected]. Park, H. Park, Prof. B. JoungDivision of CardiologySeverance Cardiovascular HospitalYonsei University Health SystemYonsei University College of Medicine50 Yonsei-ro, Seodaemun-gu, Seoul 03722, Republic of KoreaDr. C. K. JeongKAIST Institute for NanoCentury (KINC)Daejeon 34141, Republic of KoreaDr. C. K. JeongDepartment of Materials Science and EngineeringThe Pennsylvania State UniversityUniversity Park, PA 16802, USAJ. G. ParkROBOPRINT Co., Ltd.75 Nowon-ro, Buk-gu, Daegu 41496, Republic of Korea

DOI: 10.1002/adma.201702308

early disease warning, and therapeutic treatments to patients. However, conven-tional real-time monitoring devices have critical demerits such as rigid structures,[7] high power consumption,[8] and limited functionality,[9] causing inconvenience in daily life and restricted medical applica-tions. An attractive approach to solve these issues is to develop flexible biomedical sensors that offer conformal and ultrasen-sitive properties for an imperceptible link between the human body and electronic devices.[10–14]

Many research groups have reported flexible pressure sensors based on var-ious transduction mechanisms, including piezoresistive,[15,16] capacitive,[17–19] elec-tromagnetic,[20] and optical types.[21] Although these pressure sensors have detected physiological signals produced by human activities, voltage supply to operate

these sensor devices is inevitable,[22] which is one of critical obstacles for wearable/Internet of Things (IoTs) sensors.[23,24] Frequent replacement of power sources also restricts the devel-opment of implantable and wearable electronics for both in vitro and in vivo environments.[7] Alternatively, a self-powered pressure sensor has been considered as a promising candidate to solve the power consumption issue for the upcoming era of wearable healthcare sensors. A triboelectric[25–28] and piezoelec-tric[29,30] pressure sensor can directly generate electrical signals in response to mechanical force, facilitating the realization of a self-powered sensor system. Nevertheless, the triboelectric sen-sors have some drawbacks such as vulnerability to humidity, abrasion resistance, and limitation in detection of biosignals because two counterpart substrates need to effectively interact with each other.[31] Polymer-based piezoelectric sensors have also been reported,[32] but they exhibit low sensitivity because of intrinsically poor piezoelectric properties, which makes it dif-ficult to distinguish the subtle change of vital signals from envi-ronmental noise in real time.

Inorganic piezoelectric materials such as BaTiO3 and Pb[Zrx,Ti1−x]O3 (PZT) have been studied as active materials for microelectromechanical devices,[33] actuators,[34] and strain sensors[35] owing to their large piezoelectric and excellent elec-tromechanical coupling coefficients, as well as high dielectric permittivity.[36] Our group has also reported a flexible perovskite piezoelectric energy harvester[37,38] and acoustic sensor[39] uti-lizing mechanical exfoliation and inorganic-based laser lift-off

Continuous monitoring of an arterial pulse using a pressure sensor attached on the epidermis is an important technology for detecting the early onset of cardiovascular disease and assessing personal health status. Conventional pulse sensors have the capability of detecting human biosignals, but have significant drawbacks of power consumption issues that limit sustainable operation of wearable medical devices. Here, a self-powered piezoelectric pulse sensor is demonstrated to enable in vivo measurement of radial/carotid pulse signals in near-surface arteries. The inorganic piezoelectric sensor on an ultrathin plastic achieves conformal contact with the complex texture of the rugged skin, which allows to respond to the tiny pulse changes arising on the surface of epidermis. Experimental studies provide characteristics of the sensor with a sensitivity (≈0.018 kPa−1), response time (≈60 ms), and good mechanical stability. Wireless transmission of detected arterial pressure signals to a smart phone demonstrates the possibility of self-powered and real-time pulse monitoring system.

Piezoelectric Sensors

Real-time biomedical monitoring systems have provided a tremendous medical breakthrough in our modern lifestyle because they can recognize bilateral and instantaneous com-munication of physiological signals and medical diagnosis.[1–3] For example, continuous monitoring of heart rate,[4] blood pres-sure,[5] and respiration rate[6] can provide personal health status,

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(ILLO) technique. Despite their outstanding performance, there has been a substantial challenge to demonstrate self-powered piezoelectric biomedical sensors for real-time healthcare moni-toring implemented on a human interface.

Herein, we demonstrated a self-powered flexible piezoelec-tric pulse sensor based on PZT thin film for a real-time health-care monitoring system. ILLO process was used to transfer a high-quality piezoelectric thin film onto an ultrathin plastic (thickness of 4.8 µm). An epidermal piezoelectric sensor ena-bles complete conformal attachment on a rugged skin to respond the tiny human pulse. To investigate the driving mech-anism of the piezoelectric sensor, a pressure-induced piezopo-tential was theoretically calculated by a finite element analysis (FEA). Material properties and output performance of the flex-ible pressure sensor were characterized depending on various force/frequency modes and process conditions. Ultrathin pulse sensors on human wrist and neck detected radial/carotid artery signals, respiration rates, and trachea movements. A self-pow-ered piezoelectric pulse sensor was integrated with a signal

processing circuit composed of amplifiers, comparators, and output modules to identify the tiny arterial pulse. Finally, the pulse signal was wirelessly transmitted to a smart phone for a real-time monitoring system using a microcontroller and a Bluetooth transmitter.

Figure 1a schematically shows the fabrication process of self-powered flexible pressure sensor on an ultrathin plastic and its application to a real-time pulse monitoring system. The fabri-cation process of the flexible piezoelectric sensor is as follows: i) A high-quality PZT thin film was coated and annealed on a sapphire substrate, followed by attaching thermal release tape served as a temporary substrate. Subsequently, a XeCl excimer laser (wavelength of 308 nm) was irradiated to the backside of the sapphire substrate to exfoliate the PZT thin film from rigid substrate. The exfoliation was due to partial vaporization and decomposition of the PZT at the interface between the sapphire and PZT films. ii) The exfoliated PZT thin film was transferred to an ultrathin polyethylene terephthalate (PET) substrate (thickness of 4.8 µm, mass per unit area of 5 g m−2) utilizing an

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Figure 1. a) Schematic illustration of the fabrication process for self-powered pressure sensor: i) PZT thin film (2 µm) deposited on a sapphire wafer, supported by a thermal release tape, exfoliated via the ILLO process. ii) Transferred onto an ultrathin PET substrate (4.8 µm), thermal release tape is removed, and Cr and Au are deposited for IDEs (width of 200 µm and an inter-electrode gap of 100 µm). iii) Conformal contact of sensor on human wrist, and wireless transmission of detected pulse signals to a smart phone. b) Photographs of PZT thin film on an ultrathin flexible substrate floating on soap bubbles. The inset indicates a cross-sectional SEM image of the flexible pressure sensor showing the ultrathin adhesive layer between the ultrathin flexible substrate and PZT thin film (Scale bar, 5 µm). c) The calculated piezopotential distribution inside the PZT film under normal pressure of 10 kPa in the direction of the z-axis.



ultraviolet (UV)-cured adhesive polymer, and then the transfer medium was detached by thermal treatment. Gold interdigi-tated electrodes (IDEs with a width of 200 µm and interelec-trode gap of 100 µm) and a photocurable passivation layer were formed on the PZT thin film. The thickness of passivation layer to be placed on a theoretical mechanical neutral plane inside the PZT thin film was selected to protect the device from mechanically exerted force and the high voltage poling process. iii) The completed ultrathin piezoelectric sensor was confor-mally attached on a human wrist using a widely used biocom-patible liquid bandage (Nexcare, 3M), thus making it possible to detect minute biosignal changes in the epidermis.[40,41] Finally, the self-powered piezoelectric sensor was applied as a real-time pulse monitoring system through wireless transmis-sion of the detected signals to a smart phone (see Figure S1 in the Supporting Information for details about the entire fabrica-tion process).

Figure 1b presents a photograph of PZT thin film transferred onto an ultrathin substrate conformally floating on soap bub-bles. The extreme flexibility of the device is due to decreased bending rigidity of the ultrathin plastic. The cross-sectional scanning electron microscopy (SEM) image indicates the PZT thin film attached onto ultrathin substrate with the 1 µm thick-ness UV-cured adhesive polymer, as shown in the inset of Figure 1b. The adhesive layer thickness was selected to balance out between flexibility of the completed device and the adhesion of the PZT thin film to the ultrathin substrate (the flexibility of the device is inversely proportional to the film thickness, while the adhesion between PZT film and ultrathin substrate is pro-portional to the thickness of adhesive layer). This outstanding flexibility allows the sensor device to achieve conformal contact to the complex human skin topography, which is essential to improve sensing ability of tiny pressure arising near the surface region of epidermis such as that produced by arterial pulse. In order to demonstrate the sensing mechanism of piezoelectric pressure sensor, piezopotential distribution was simulated by a FEA using the COMSOL Multiphysics software. Figure 1c shows the calculated piezopotential distribution of the PZT film induced by pressure of 10 kPa normal to the device surface. The PZT thin film on ultrathin PET substrate exhibited a well-distributed piezopotential difference (∆V) of 26.7 V between adjacent electrodes under normal pressure of 10 kPa, as shown in Figure S2 (Supporting Information). According to previous studies,[8] the pressure produced by human-body movements is mainly distributed to a low-pressure (1–10 kPa) and medium-pressure (10–100 kPa) regime. In particular, the medium-pres-sure regime is formed by vital signs including arterial pulse, human respiration, and phonation vibrations. The simulation results strongly support that our piezoelectric pressure sensor can identify the biosignals arising from the medium-pressure regime by converting them into electric energy immediately.

Maintaining a morphotropic composition and crystal struc-ture of PZT thin film during the transfer process is essential to maximize the performance of piezoelectric sensors. Accord-ingly, we characterized the PZT film before and after the transfer process using X-ray photoelectron spectroscopy (XPS), X-ray diffraction (XRD), energy-dispersive spectroscopy (EDS), and Raman spectroscopy. Figure S3a (Supporting Information) shows the XPS survey spectra of PZT surfaces on the sapphire

(bottom, red) and ultrathin PET substrate (top, blue). The ele-mental binding energy of the PZT thin film shows negligible change after the transfer process, indicating that the ILLO method does not cause any compositional degradation of the PZT thin film. The XRD analysis also demonstrated that poly-crystalline perovskite PZT film on the sapphire (bottom, red) was successfully transferred to the ultrathin PET substrate (top, blue) without any structural damage, as appeared in Figure S3b (Supporting Information). The crystallinity of the PZT film before and after transfer was examined by the rocking curve of the (200) peak. The measured full width at half maximum (FWHM) of the rocking curve of the PZT film was 0.391° (on the sapphire) and 0.395° (on the ultrathin PET), respectively, representing good crystalline quality and orientation of the PZT thin film. An elemental mapping and compositional analysis of the PZT thin film from EDS, as shown in Figure S4 (Sup-porting Information), indicated that the PZT on the ultrathin PET consisted of all appropriate elements, and Pb, Zr, Ti, and O atoms were properly distributed in a measured area. Raman spectroscopy measurement was also conducted to determine the phase of the PZT film using a 514.5 nm Ar+ laser beam as an excitation source. The peaks marked by yellow arrows in the Raman spectra express the typical features of perovskite phase that do not change during the ILLO process (Figure S5, Supporting Information). These results clearly confirmed that the structural damage and compositional transition of PZT thin film were negligible during ILLO process because of rapid recrystallization of the PZT interface layer after short pulse laser irradiation.

Characteristics of the flexible piezoelectric sensor were inves-tigated utilizing a dynamic pushing stage system capable of modulating normal force and frequency rate. Figure 2a reveals open-circuit voltage of the sensor under different normal pres-sure after fast Fourier transform (FFT) analysis by MATLAB. Compared to raw data (Figure S6, Supporting Information), the ±100 mV fluctuation signal was almost filtered after the FFT process, which clearly showed that this fluctuation was due to over 60 Hz high frequency noise arising from sur-rounding environments. The output voltage generated by the piezoelectric sensor consistently increased from 0.3 to 1.85 V under exerted pressure ranging from 3.75 to 25 kPa. In addi-tion, characteristic of the sensor under various bending radius was demonstrated, as shown in Figure S7 (Supporting Infor-mation). As the mechanical neutral plane (MNP) is located near the middle plane of the PZT film, the output voltage is almost the same regardless of the bending radius from flat to 1.0 cm (see Figure S8 in the Supporting Information for details about MNP). A pressure sensitivity, response time, and durability of the piezoelectric pressure sensor were characterized to verify the sensor’s widespread applicability in different types of vital signal sensing. As the applied pressure increases, the gener-ated output voltage increases and finally saturates at 60 kPa, as shown in Figure 2b. The pressure sensitivity of the piezoelec-tric sensor is defined as the slope of normalized voltage (VN) versus pressure curve (S = dVN/dP = d(∆V/VS)/dP, where P is the applied pressure, VS is the saturation voltage, and ∆V is the relative change in output voltage).[26] The flexible pressure sensor exhibits almost a linear relationship within a pressure of 30 kPa with a pressure sensitivity of 0.018 kPa−1. The pressure

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was measured under our system as low as 1 kPa, which is the minimum value of our pushing stage system. The linear prop-erty between applied pressure and output voltage of the piezo-electric sensor was theoretically consistent with the following equation[42]

V g hP ks

hPj jj

j3 3

1

3ε= − = −

(1)

where g3j is the piezoelectric voltage constant, h is the thickness of the PZT film, P is the normal pressure, k3j is an electrome-chanical coupling factor, s is the elastic compliance, and ε is the permittivity of PZT film, respectively. The first number and second j subscripts indicate the direction of the poled dipoles and applied pressure. The output voltage is proportionally increased in the low- and medium-pressure region, while it is slightly decreased after pressure of 30 kPa. This is presumably due to the theoretical limits of effective strain in piezo-materials at high pressure region.[43–45] The piezoelectric pressure sensor exhibited a response time of 60 ms (in the inset of Figure 2b), whose results are comparable to the recently reported self-powered pressure sensor.[26] Moreover, the mechanical dura-bility of the sensor was verified by a dynamic pushing test for 5000 cycles under 20 kPa pressure and 2 Hz frequency, as shown in Figure 2c. The generated output voltage displays no noticeable fluctuation during the repetitive pushing tests (see

in the inset of the Figure 2c). This remarkable durability is estimated because the MNP is located near the middle plane of the PZT layer. Furthermore, the piezoelectric sensor also responded to various ranges of frequency vibrations. Figure S9 (Supporting Information) indicates output voltage of the piezo-electric sensor in response to applied pressure (25 kPa) with different frequencies (0.2–5.0 Hz). In spite of its intrinsic insen-sitivity to a sustained static pressure, the piezoelectric sensor can instantaneously generate the output voltage in response to rapid variation of dynamic pressure. Figure 2d shows a sche-matic diagram of sound wave detection setup by a piezoelectric sensor (top panel) and output voltage (bottom panel) generated by sound wave at audible frequency whose intensity changes at regular intervals. The peak-to-peak voltage (Vpp) of 100 mV was measured from the sound pressure with even a low inten-sity level of 80 dB.[46] The frequency-response characteristics of piezoelectric pressure sensor can be utilized for self-powered electronics such as accelerometers, vibrometers, and wearable voice monitoring.

Direct detection of physiological signs including arterial pulse and respiration rate using a self-powered piezoelectric sensor has significance for human medical health information. Figure 3a shows a photograph of a piezoelectric pulse sensor conformally attached on a human wrist with a biocompatible liquid bandage spray. The sensor device was stably deformed to effectively respond to blood vessel movements due to the

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Figure 2. a) The generated output voltage of the piezoelectric pressure sensor under different pressure conditions. b) The normalized output voltage as a function of pressure and the enlarged output peak representing the response time of the pressure sensor (inset). The pressure sensor shows good linearity and sensitivity in the range from 1 to 30 kPa. c) Mechanical durability test of the piezoelectric sensor under repeated pushing/unpushing conditions with frequency of 2 Hz and pressure of 20 kPa. The insets show a partially magnified output voltage versus time. d) Schematic illustration of sound pressure detection by flexible pressure sensor (top) and output voltage in response to single sinusoidal wave with frequency of 240 Hz and intensity of 80 dB (bottom).



excellent flexibility of the ultrathin plastic, as appeared in the inset of Figure 3a. After detaching the device from the wrist, optical and SEM images indicate that there were some adhe-sive residues but no significant cracks (Figure S10, Supporting Information). Radial artery pulse signals detected in situ by the flexible pulse sensor before (red line) and after (blue line) physical exercise (running for 10 min) are shown in Figure 3b.

The demonstrations were performed a healthy male in his late twenties, and no allergic reactions, skin wounds, or ailments were observed during our studies. The radial artery pulse generated an average Vpp of 65 mV and ≈73 beats per minute (BPM) before exercise, 81.5 mV and ≈100 BPM after exercise, respectively. These values represent increases of approximately by 25% and 37%, respectively, due to the increased rate of heart

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Figure 3. a) Photograph of a piezoelectric pulse sensor conformally attached on human wrist using a biocompatible liquid bandage. The inset shows the deformed sensor consistent with blood vessel movement. b) Radial artery pulse signals detected by the self-powered pulse sensor, showing different heart rates and generated output voltages before and after physical exercise. The inset indicates the magnified voltage signal by the radial artery pulse before exercise, clearly representing pulse pressure (P1) and late systolic augmentation (P2). c) Photograph of piezoelectric pulse sensor conformally attached on a carotid artery position (top) and the middle of the throat (bottom). d) The generated output voltage in response to carotid arterial pres-sure (top) and saliva swallowing actions (bottom). e) Photograph of medical mask integrated with a flexible pressure sensor for monitoring human respirations (left) and output voltage response of the pressure sensor to normal (right bottom, blue) and deep oral breathing (right top, red) during periodic oral breathing. f) Immunofluorescence images of HEK 293 and HL-1 cells cultured on the piezoelectric sensor device after a day (left paler) and 5 days (right panel). The scale bar is 100 µm (blue: DAPI, green: α-SMA).



pumping to maintain the oxygen supply delivered to muscles during the exercise. As shown in the inset of Figure 3b, the enlarged pulse signal (before exercise) clearly exhibits the char-acteristic peaks of peripheral artery waveforms, which contain important biomedical and physiological information such as arterial stiffness, coronary artery disease, and myocardial infrac-tion. Under normal conditions, peaks P1 and P2 denote the sum of the forward travelling wave and reflected wave (from the hand), and the reflected wave from the lower body subtracted by the end-diastolic pressure, respectively. The radial artery augmentation index (AIr) and ∆TDVP, defined as P2/P1 and the time difference between the P1 and P2 peaks, respectively, are strongly related to arterial stiffness.[47] The measured average values of AIr and ∆TDVP were 0.54 and 0.23 s, respectively, which are compatible with a person in his or her late twenties, as reported by Nichols.[48] After the exercise, the average value of AIr dropped to 0.22 due to decrease in the late systolic aug-mentation (P2), resulting from changed heart rate/ventricular ejection, large artery stiffness, or alteration in tone of muscular arteries affecting pressure wave reflection.[47] Figure S11 (Sup-porting Information) indicates arterial pulse signals detected from mid-twenties skinny and obese males. A little differ-ence in output voltage is estimated to be due to individual variations in each person’s physiological parameters such as the epidermal thickness and Young’s modulus. Furthermore, other human physiological activities including carotid artery pulse and muscle movements were noninvasively monitored by the piezoelectric sensor conformally attached on a human neck (Figure 3c). The generated Vpp from the carotid artery pulse (top panel) and saliva swallowing action (bottom panel) were 400 and 1000 mV, respectively, as shown in Figure 3d. The output voltage measured by the carotid artery is roughly six times higher than that by radial artery due to high carotid artery blood pressure. A regular pattern of trachea movements was identified by the consistent output voltage waveform during saliva gathering and swallowing. Human respiratory activities were also continuously monitored by attaching the flexible pie-zoelectric sensor onto a conventional medical mask, as shown in Figure 3e. To eliminate any influences by the humidity on the device properties, a complete biocompatible encapsulation layer was deposited.[49,50] The sensor distinguished between deep (red) and normal (blue) respiration mode with consist-ently obtained output voltage waveforms under repeated oral breathing. Because of the temperature changes during human respiration, the obtained voltage also might be affected by a pyroelectric effect, whereby temperature variations are con-verted into electric potential. The pyroelectric Vpp generated by temperature variations can be theoretically expressed as the fol-lowing equation

V PA

CT P

hTQ

eQ

r 0ε ε∆ = ∆ = ∆

(2)

where ∆V is the generated pyroelectric Vpp, PQ is the pyroelec-tric charge coefficient, A is the area of the sensor, Ce is the capacitance of the sensor, ∆T is the temperature variations, h is the thickness of the PZT, εr is the relative permittivity of the PZT, and ε0 is the vacuum permittivity. At ∆T ≈ 10 °C, corresponding with the temperature difference between inhale

and exhale, the calculated pyroelectric Vpp of the sensor was ≈0.24 V, which accounts for only 25% (normal mode) and 10% (deep mode) of the obtained output voltage, respec-tively. These theoretical results indicate that the signal peaks obtained from respiratory activities are dominated by tiny mechanical movements (see the Supporting Information for details). In addition, we believe that that the pyroelectric effect can be neglected in respiratory activities because phase transi-tion of the PZT film is small in low temperature region below 50 °C.[51,52]

A reliable evaluation of biocompatibility is essential to apply piezoelectric sensor directly on human skin without inducing any forms of biological damage. Figure 3f shows cell cytotoxicity tests performed by assessing the viability of human embryonic kidney cells (HEK293) and mouse cardiac muscle cell line (HL-1). Each cell was stained using 4′,6-diamidino-2-phenylin-dole (DAPI) and alpha smooth muscle actin (α-SMA) utilized as a marker for nuclei (blue) and myofibroblast (green) forma-tions, respectively (more detailed procedures are described in Supporting Information). The HEK293 and HL-1 cells with undamaged nuclei were well cultured on the device surface after a day (left panel) and 5 days (right panel) in an incubator because the device was fully passivated by biocompatible epoxy (SU-8). Cells grown on the device showed effectively no differ-ences from those grown on standard culture dishes, as shown in Figure S12 (Supporting Information). These experimental results demonstrate that the well-passivated piezoelectric sensor offers negligible cytotoxicity to human beings, facili-tating application as a wearable healthcare sensor.

To validate its applicability to a real-time pulse monitoring system operated by our self-powered flexible pulse sensor, it is essential to realize the circuit operation composed of signal amplification, frequency filtering, and decision making module. Here, we implemented a complementary signal pro-cessing circuit to easily identify arterial pulse signals. Figure 4a presents the complementary amplifier circuit diagram com-posed of several electronic parts including conventional operation amplifier chips (LM358, Texas Instruments) with pas-sive band-pass filters (BPFs), a commercial red light emitting diode (LED), and a speaker module (Figure S13, Supporting Information). Resistance and capacitance values for the BPFs are specifically chosen to eliminate unwanted DC noise that disturbs the desired artery signals, and high frequency noise components (see Figure S14 in the Supporting Information for details). A comparator unit generates a constant voltage output of either 5 or −5 V depending on the output from the second noninverting amplifier unit compared to a threshold voltage. If it exceeds the threshold, the comparator output is set to 5 V, which makes the LED blink and turn on the speaker module. The LED and speaker unit were synchronously operated cor-responding to the radial artery pulse waveform (see Video S1 in the Supporting Information). The inset of Figure 4b shows the voltage output from the first and second amplifier stage in each cycle of repetitive attachments/detachments, indicating that the flexible piezoelectric sensor has capability for long term and multiple uses, which may reduce the medical expenses of customers. Furthermore, we expect that complete self-powered complementary circuit system can be implemented in real time by utilizing other flexible energy harvester.[53]

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Finally, the radial artery pulse detected by self-powered pulse sensor was wirelessly transmitted to a smart phone using an Arduino microcontroller unit (MCU) prototyping platform (Arduino-Nano ATMEGA 168). Figure 4c represents a photo-graph of the Arduino MCU board connected with a Bluetooth module (HC-05) and an Android-based smart phone that dis-plays the radial artery pulse in real-time (see Video S2 in the Supporting Information). A schematic diagram of the wireless transmission system is shown in Figure S15 (Supporting Infor-mation). The MCU board converts the artery pulse signal into 10 digital forms per second and transmits them to the smart phone through a Bluetooth transmitter. The received pulse signal presents a similar configuration compared to the output signal obtained from the measurement system (see Figure 3b). These results confirm that the self-powered pulse sensor can be utilized for a continuous real-time health/wellness monitoring system.

In conclusion, we have successfully demonstrated a self-powered piezoelectric pulse sensor on an ultrathin substrate (thickness of 4.8 µm) using an ILLO process. The PZT thin film obtained by a sol–gel method was transferred from a sac-rificial substrate onto an ultrathin PET substrate without any mechanical damage. The FEA simulation theoretically con-firmed that the self-powered piezoelectric sensor can generate electrical energy immediately from normal pressure. The flex-ible piezoelectric sensor exhibited a sensitivity of 0.018 kPa−1,

response time of 60 ms, and good mechanical stability under 5000 pushing cycles. The piezoelectric sensor also responded to lower frequency vibrations (0.2–5.0 Hz) and higher fre-quency sound waves (240 Hz). Ultrathin piezoelectric sensor was conformally attached on human epidermis and detected the radial/carotid artery pulse, respiratory activities, and trachea movements. The output voltage obtained by arterial pressure successfully operated the LED and speaker module of a signal processing circuit composed of amplifiers, BPFs, and com-parators to identify the arterial pulse. More interestingly, a self-powered real-time pulse monitoring system was demonstrated by wirelessly transmitting the pulse signal to a smart phone utilizing a MCU and a Bluetooth transmitter. We are currently investigating a self-powered healthcare sensor using biocom-patible piezoelectric materials for long-term clinical trials.

Experimental SectionDynamic Pushing Stage System: The dynamic pushing stage system

set-up was operated by utilizing a vibration stage (LW-140.141-110, Labworks Inc.) and function generator (33250A, Agilent), which can modulate force and frequency rate.

Measurement of Electrical Characteristics: The output performance of the flexible piezoelectric pressure sensor was conducted using a Keithley 6514 electrometer and then the signals were collected and analyzed by the Data Acquisition Card (NI PCIe-6321).

Figure 4. a) Schematic diagram of complementary signal processing circuits composed of amplifier chips, BPFs, a red LED, and a speaker module. b) Photograph of the LED and speaker unit operated synchronously corresponding to the radial artery pulse. The inset indicates the output voltage from the first (bottom, red) and second (top, blue) amplifier stage. c) Photograph of wireless transmission of the pulse to a smart phone, showing capability for a real-time arterial pulse monitoring system.



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Supporting InformationSupporting Information is available from the Wiley Online Library or from the author.

AcknowledgementsD.Y.P. and D.J.J. contributed equally to this work. This research was supported by the National Research Foundation of Korea (NRF) Grant funded by the Korean Government (MSIP) (NRF-2016R1A5A1009926). This work was also supported by Nano Material Technology Development Program through the National Research Foundation of Korea (NRF) funded by the Ministry of Science, ICT and Future Planning (MSIP) (grant code: NRF-2016M3A7B4905609, and 2016M3A7B4910636).

Conflict of InterestThe authors declare no conflict of interest.

Keywordsarterial pulse monitoring, laser lift-off, real-time monitoring, self-powered, ultrathin piezoelectric sensors

Received: April 25, 2017Revised: May 27, 2017

Published online:

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