Hygroscopic Auxetic On-Skin Sensors for Easy-to-handle Repeated

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Hygroscopic Auxetic On-Skin Sensors for Easy-to-handle Repeated Daily Uses Hyun Woo Kim, Tae Yeong Kim, Hyung Keun Park, Insang You, Junghyeok Kwak, Jong Chan Kim, Heeseon Hwang, Hyoung Seop Kim, and Unyong Jeong ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.8b13857 • Publication Date (Web): 15 Oct 2018 Downloaded from http://pubs.acs.org on October 15, 2018

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ACS Applied Materials & Interfaces

Hygroscopic Auxetic On-Skin Sensors for Easy-to-handle Repeated Daily Uses Hyun Woo Kim,1‡ Tae Yeong Kim,1‡ Hyung Keun Park,1 Insang You,1 Junghyeok Kwak,1 Jong Chan Kim,2 Heeseon Hwang,2 Hyoung Seop Kim,1 and Unyong Jeong1*

1

Department of Materials Science and Engineering, Pohang University of Science and

Technology (POSTECH), 77 Cheongam-ro, Nam-gu, Pohang-si, Gyeongsangbuk-do, 37673, Republic of Korea 2

Marine Robotics R&D Division, Korea Institute of Robot and Convergence (KIRO), 39 Jigok-

ro, Nam-gu, Pohang-si, Gyeongsangbuk-do, 37666, Republic of Korea

‡

These authors contributed equally to this work

* Corresponding author: [email protected]

ACS Paragon Plus Environment

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ABSTRACT Despite the advance of the on-skin sensors over the last decade, a sensor that solves simultaneously the critical issues for using in everyday life, such as stable performance in various environments, use over a long period of time, and repeated use by easy handling, has not yet been achieved. Here, we introduce an auxetic hygroscopic sensor that simultaneously meets all the conditions. The auxetic structure with a negative Poisson’s ratio matches with deformation of the skin in ankles, hence conformal contact between the sensor and the skin could be maintained during repeated movements. Sweat was absorbed in the auxetic electrode made of a hydrogel pattern coated with Ag nanowires and it evaporated quickly, the hygroscopic characteristic led to excellent breathability. An electrocardiogram (ECG) sensor and a haptic device were fabricated according to the proposed design for sensor electrode. The sensors provide stable detecting performance in various environments, such as exercising, submersion in water, exposure to concentrated salt water, and continuous wearing for long time (7 days). And, the sensors could be manually attached repeatedly without degrading the performance. This study provides new structural insights for on-skin sensors and presents the future research directions.

KEYWORDS: On-skin sensor, Healthcare, Auxetic structure, Reusable sensor, Stretchable electronics, Electrocardiogram


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Introduction Healthcare devices are evolving into on-skin sensors attaching to the skin.1,2 The on-skin sensors detect mechanical, electrical, chemical, and sonic signals from body or sense the external stimulations like pressure, strain, and torque as human skin does. Many on-skin sensors have been reported during the past decade, including electrocardiogram (ECG),3,4 apexcardiogram (ACG),5,6 electromyogram (EMG),7,8 electroencephalography (EEG),9,10 tactile,11–15 chemical sensors,16 and temperature sensors.17 The main direction of the on-skin sensors is comfortable daily-life use. There are a number of technical issues for the daily-life use of the on-skin sensors; high mechanical compliance required for imperceptible wearing, dry contact to the skin to maintain the same electrical performance during sweating and wearing in rain, breathability for long-term use without skin irritation, chemical stability in high salinity condition, easy to handle for user’s convenience, and repeated uses without performance change. So far, a number of deformable on-skin sensors with excellent mechanical compliance to dynamic body motions have been demonstrated using the stretchable interconnections or using ultrathin polymer substrates.18,19 Consistent contact resistance of the electrode has been obtained using a conducting polymer or hybrid electrodes.20,21 Very recently, Someya and coworkers presented a breathable thin film on-skin sensor that can be used for a week.22 A practically meaningful onskin sensor must have a certain thickness for easy handling while satisfying other needs simultaneously. Because human skin under bending, exhalation, and muscle tension expands biaxially, its Poisson’s ratio is negative depending on the body parts. Due to the Poisson’s ratio (typically 0.5) of polymer elastomers,23,24 elastomer films with a thickness manageable by hand cannot maintain its conformal contact with skin under large body motions. In this study, we introduced a thick


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auxetic structure to maintain the conformal contact of the sensor to the body parts. And we fabricated a Ag nanowire (NW) electrode on a network of hygroscopic gel serpentine interconnections to achieve dry contact in wet environments and the excellent breathability for long-term use.

RESULTS AND DISCUSSION

Figure 1. Schematic illustration describing the fabrication of the hygroscopic auxetic electrodes. A PEG-DA gel network structure was formed by UV crosslinking through a mask and Ag nanowires (NWs) were spray-coated. After washing, the electrode was peeled off and turned over to attach to the skin.

Schematic illustration in Figure 1 shows the process of fabricating the hygroscopic auxetic electrode. A solution containing poly(ethylene glycol)-diacrylate (PEG-DA), 2-hydroxy-2methylpropiophenone (HMOPP), and acrylic acid (AA) was dropped on a glass substrate with


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100 µm-thick spacers. A photomask was placed on the mixture liquid, and UV (λ = 365 nm, 10 mW) was irradiated for 6 s to form a network of centroids with six serpentine interconnections. The uncrosslinked liquid was washed with DI water. The thickness, width, and curvature of the serpentine arms were 100 μm, 100 μm, and 300 μm, respectively. In this study, we demonstrate an ECG electrode and a haptic sensor. The ECG electrode should be stable under deformation, but the haptic sensor should be sensitive to strain. We spray-coated the Ag NW suspension in ethanol (0.05 wt%) on the PEG gel pattern by 8 mL for the fabrication of the strain-insensitive ECG electrodes and by 4 mL for production of strain-sensitive haptic sensors. After annealing at 160 oC for 2 h, the conductive pattern was peeled off and transferred onto the body parts of interest. The mechanical property of the gel was dependent on the weight fraction of AA (φAA) in the gel. Figure 2A exhibits the stress-strain curves of the gel films with different φAA. All the gel films had a well-defined elastic behavior below the fracture stress (σf) regardless of φAA. The pristine PEG gel film had a tensile modulus (E) of 15.7 MPa and σf of 3.0 MPa. E and σf decreased at higher φA, E = 11.9 MPa and σf = 1.8 MPa at φAA = 7.0 wt%, and E = 11.4 MPa and σf = 1.2 MPa at φAA = 9.0 wt%. The finite element method (FEM) simulations shows the stress distribution in the gel network at ε = 40% and the maximum local stress (σmax) as a function of ε (Supporting Information, Figures S1, S2). Fracture of the network electrode was observed in φAA = 9.0 wt% at ε = 40% (red-colored area) but not in the other conditions (φAA = 0, 5.0, 7.0 wt%). σmax for the gel electrode with φAA = 7.0 wt% was 1.7 MPa at ε = 30% which was less than σf of the corresponding film, hence the gel network was stable at ε = 30%. We chose φAA = 7.0 wt% as the optimal condition to achieve high mechanical deformability. Figure 2B compares the changes in the relative electrical resistance of the serpentine network electrodes as a function of ε,


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depending on φAA. A relatively large amount (8 mL) of Ag NWs solution was coated to obtain strain-insensitive electrode. The resistance of the gel pattern without AA increased with strain due to low adhesion between Ag NWs and the gel.25

Figure 2. (A) Stress-strain curves of the continuous gel films according to the weight fraction of acrylic acid (AA) in the PEG-DA gel. (B) Relative resistance changes under uniaxial stretching of the network-structured electrodes having different AA weight fractions.


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The relative resistance change decreased as φAA increased from 0 wt% to 7.0 wt%, which resulted from the enhanced adhesion between the Ag NWs and the gel in addition to the improved softness of the gel. The Ag NW layer on the gel with φAA = 0 wt% was damaged by five times peel-off tests with a 3M tape (Magic tape 810) (Figure 3A), whereas the Ag NW layer on the gel with φAA = 7.0 wt% was not damaged during the peel-off tests (Figure 3B). As a result, the resistance of the Ag NW layer on the gel with φAA = 7.0 wt% maintained its initial resistance during repeated tape test, while the Ag NW layer on the pure gel showed sharp increase of the resistance (Figure 3C). Excessive amount of AA (φAA = 9.0 wt%) degraded the mechanical durability of the serpentine network, hence the condition of φAA = 7.0 wt% was used to produce reliable electrodes.


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Figure 3. (A, B) SEM images of the Ag NW layers on the PEG-DA gel after the five times peeling tests; PEG-DA gel film with 0 wt% AA fraction (A) ad 7.0 wt% AA fraction (B). (C) Relative resistance changes in the auxetic network gel electrode by repeated peeling tests. The electrodes on the PEG-DA gel network with 0.0 wt% AA fraction and the PEG-DA gel film with 7.0 wt% AA fraction were tested.


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Figure 4. (A) Optical microscope (OM) images of the network gel electrode at a unstretched state ( ε = 0%) and at a uniaxially stretched state ( ε = 30%). (B) Simulated Poisson’s ratio of the gel network electrode at a unstretched state ( ε = 0%) and at a uniaxially stretched state ( ε = 30%). (C) Simulated Poisson’s ratio of the gel network electrode as a function of uniaxial strain ( ε). (D) Camera images of the electrode on the knee at the initial state ( ε = 0%) and at the bended state (ε = 40%).

The structure of the electrode suggested in this study has several unique advantages. In spite of its thickness (100 µm), the electrode was deformable and ultralight (4.93 mg/cm2) because of the open voids between the gel serpentines, hence the resulting sensor was


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imperceptible in wearing. The optical microscope (OM) image in Figure 4A shows the deformation of the serpentine network electrode before and after uniaxial stretching at ε = 30% (indicated by black arrows). The electrode expanded also in the perpendicular direction to the uniaxial stretching. Based on the OM image, the Poisson’s ratio of the serpentine network was measured to be -0.32. Figure 4B shows the FEM simulation of the Poisson’s ratio as a function of uniaxial tensile strain. The calculated Poisson’s ratio of the serpentine network electrode was 0.31 at ε = 30% which was consistent with the measured value. Figure 4C exhibits the Poisson’s ratio as a function of the uniaxial strain from ε = 2.5% to ε = 40%. In all the strain range, the Poisson’s ratio was about -0.3. Figure 4D shows the dimensional change of the network electrode attached to a knee at the initial state and the bended state. The bending corresponds to ε = 40%. This auxetic deformation made the electrode compliant to the biaxial expansion of the human skin without any delamination. The relationship between the design parameters of the auxetic structure and the corresponding Poisson’s ratio has been studied in recent reports.26,27 The network electrode attains breathability which is critical for long-term use without skin irritation. As illustrated in Figure 5A, the voids and the hygroscopic serpentine pattern are the routes to air penetration and fast evaporation of sweat. The relative evaporation rates through the voids and the gel pattern were measured quantitatively. Figure 5B compares the evaporation rates of a PEG gel/Ag NW hybrid film without voids (red dotted line), the serpentine network structure used in this study (blue dotted line), and no sealing (black dotted line). For comparison, we added the evaporation rates of a polyimide (PI) film (yellow dotted line) and a PDMS film (cyan dotted line) that are widely used as substrates. The thickness of the specimens was the same to be 100 µm. A vial (20 mL) containing water (10.0 g) was sure sealed with the films, and the loss of weight was monitored for 48 h. The evaporation rate of the PEG/Ag hybrid gel film


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(eh) was 3.6 mg/h, which was about a half the evaporation rate of the ‘no sealing’ case (ev = 7.2 mg/h). It is notable that the evaporation rate of the PEG gel film was the 3.3 mg/h, indicating that the existence of Ag NWs on the surface of the gel does not affect the evaporation rate of the gel. The evaporation rate of the serpentine network structure (en) was 5.0 mg/h. This value was predicted by a linear function (thick magenta dotted line) between ev and eh in relation to the surface fraction (fc = 0.57) of the network, en = ev (1-fc)+ eh fc. Because water or sweat is absorbed in the gel serpentines and evaporates from the top surface of the gel pattern, the Ag NWs at the bottom of the gel pattern can maintain dry contact to the skin layer.


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Figure 5. (A) Schematic illustration showing the breathability of the gel network electrode on skin. Sweat can evaporate quickly through the voids among the serpentines. It is also absorbed in the gel pattern and evaporates on the surface of the gel pattern. (B) Evaporation rates of water through various films and the network gel electrodes: without capping (black), the gel network electrode (blue), gel/Ag NW hybrid electrode film (red), polydimethylsiloxane (cyan) film, polyimide (yellow) film. The thickness of the films was fixed at 100 µm. The magenta line is a theoretically calculated result.


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Figure 6. Mechanical chemical stability of the electrode and use as an electrocardiogram (ECG) sensor. (A) Relative resistance changes of the gel network electrode under dynamic mechanical stretching at ε = 30%. The weight fraction of AA in the gel pattern was varied to be 0.0, 5.0, 7.0, 9.0 wt%. (B) Relative resistance change of the gel network electrode (7.0 wt% AA fraction) under repeated dipping test in water. Resistance of electrode was measured right after taking out from water and after completely drying. (C) Relative resistance change as a function of NaCl concentration (wt%). The resistance in pure water (RO,W) was taken as the reference resistance. The electrode had 7.0 wt% AA fraction. (D) Relative resistance change of the gel electrode (7.0 wt% AA fraction) under daily life condition for long time. Electrode was exposed to the condition of 50 % humidity at 25°C for 7 days.


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Figure 6A shows the relative resistance change of the strain-insensitive auxetic electrode during repeated stretch cycles between ε = 0% and ε = 30%. The electrode showed stable dynamic response with a small change. The on-skin sensors to be used in everyday life should be stable in sweating condition. The sensor electrode with 7.0 wt% AA fraction showed a stable performance during repeated wetting tests. The resistance increased by about 20 % when the electrode was dipped in water, but it returned to its initial value when the electrode was dried (Figure 6B). Since the salt concentration increases by evaporation of sweat,28 the sensing electrode must have good stability in high salt concentration. Figure 6C shows the change in the relative resistance in relation to the salt concentration in water. Small amount (1 mL) of aqueous solutions with different NaCl concentrations (0, 5, 10, 15 wt%) were dropped on the electrode. Measurement was carried out after 90 s to ensure penetration of the salt solution in the electrode. The network electrode with 7.0 wt% AA fraction showed negligible change of the relative resistance, indicating it is not affected by high salinity environment. The human skin humidity is about 50 % in the condition of 70 ~ 80 % relative air humidity at 25°C.29 The PEG gel electrode with 7.0 wt% AA fraction showed stable electrical performance for 7 days, having only 10 % resistance increase, when it was continuously monitored in a thermohygrostat at 25°C and 50 % humidity (Figure 6D). To investigate the validity of the hygroscopic auxetic on-skin sensor, we used it as an ECG sensor electrode. The sensor was attached to a stretchable nylon fabric band for easy wearing on arms (Figure 7A). The performance of the sensor was monitored in consecutive situations. A male user was bending and spinning his arm for 1 h (Figure 7A) and sweating in exercise for 50 min (Figure 7B). The sensor was tested after dipped in water for 5 min (Figure 7C) and after wearing continuously for 7 days (Figure 7D). In Figure 7E, the ECG signals corresponding to


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Figures 7A-D are shown together with enlarged ECG signals for each situation. During the whole tests, the ECG signals were stable without increase in the signal-to-noise ratio or change of the base line. The quality of the signals was comparable to the signal obtained with a commercial Ag/AgCl electrode (Supporting Information, Figure S3). We did not find any skin irritation while wearing for 7 days (Figure 7D) and the sensor could be used again later without performance degradation.

Figure 7. Stability of the electrocardiogram (ECG) sensor. (A-D) Camera images of wearing the ECG sensor in different situations; (A) wearing with a nylon band in dry state, (B) wearing while exercising for 1 h, (C) dipping in water, and (D) after wearing continuously for 7 days. No skin irritation was observed after wearing for 7 days. E) ECG signals obtained in the situations corresponding to the camera images shown above. The signals had negligible differences in the different situations.


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The electrode coated with a 4 mL of Ag NW suspension exhibited sensitive change in the relative resistance under mechanical deformation. When sensor was stretched, the distance between Ag NWs increased (Supporting Information, Figure S4). This increment made the part of the conduction paths disconnected as explained by the percolation theory.30 The gauge factors with respect to uniaxial strain were 6.17 for φAA = 0.0 wt%, 1.96 for φAA = 5.0 wt%, and 0.47 for φAA = 7.0 wt% (Supporting Information, Figure S5). Although the gauge factor decreased with the weight fraction of AA, the one with φAA = 7.0 wt% showed stable performance during repeated tests (Figure 6A) and maintained stable dynamic responses during repeated stretching cycles between the stretched states (ε = 10, 20, 30%) and the released state (ε = 0%) (Figure 8A). This strain-sensitive electrode was used as a haptic sensor monitoring the mechanical body motions (Figure 8B). The sensor attached to the knee showed signals that are clearly distinguished by the knee angles. A movie for the bending test is shown in the Supporting information (Movie S1). Figure 8C shows the corresponding block diagram of the device system. The relative resistance change in the sensor was transferred to a mobile phone through a portable wireless communication system fabricated on a flexible printed circuit board. The signal was calibrated according to the motions of the user and the movement was animated in real time. Figure 8D exhibits the user interface showing an animated human whose motion was synchronizing with the motion of the user (also see the Supporting Information Movie S2). The sensor could be managed by hands and used for a week without performance degradation.


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Figure 8. Use of the auxetic gel electrode as a haptic device. (A) Relative resistance changes of the strain sensor under dynamic stretching between ε = 0% and different uniaxial strains (ε = 10%, 20%, 30%). (B) Relative resistance changes of the strain sensor while a male user wore it on his knee and bended repeatedly at different bending angles (30o, 60o, 90o). (C) Block diagram of the haptic device based on the strain sensor made of the auxetic network gel electrode. The relative resistance change in the sensor was transferred to a mobile phone through a portable communication system. (D) Demonstration of the wireless user interface for the haptic device.


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CONCLUSION In summary, we proposed a new electrode structure for on-skin electronic devices. We fabricated hygroscopic auxetic electrodes by coating Ag nanowires on the surface of a serpentine gel network pattern. Water was absorbed by the hygroscopic gel pattern and evaporated rapidly through the voids and the surface of the gel pattern, hence the Ag nanowire electrode maintained the dry contact with the skin during sweating or in water. The auxetic electrode with a negative Poisson's ratio expanded biaxially during unidirectional stretching and deformed compliantly to the human skin. The advantages of this electrode structure were verified by an ECG sensor and a haptic device. The sensors could be used repeatedly, manually maneuvered, and worn continuously for a week without skin irritation. This study suggests a desirable electrode structure of the on-skin devices.

EXPERIMENTAL SECTION Materials. Poly(ethylene glycol) diacrylate (PEG-DA, Mn = 700 g/mol, Mw = 8000 g/mol), 2-hydroxy-2-methylpropiopheneone (HMOPP), and acrylic acid (AA, 99%) were purchased from Sigma Aldrich. Ag nanowire dispersed in ethanol (0.5 wt%) was purchased from Nanopyxis co. (Korea). Polydimethylsiloxane (PDMS, Sylgard 184) was purchased from Dow corning. Fabrication of auxetic PEG-DA gel substrate. Small-molecule PEG-DA (Mn = 700 g/mol) and large-molecule PEG-DA (Mn = 8000 g/mol) were mixed (19:1, w/w). AA was added to the mixed PEG-DA (AA = 5.0 wt%, 7.0 wt%, 9.0 wt%). HMOPP (10 wt%) was introduced to the blend solution of PEG-DA and AA, and used as a photoinitiator. The mixture solution was


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dropped on a glass substrate (38 cm × 26 cm) having a 100 µm-thick spacer on the surface. Double-sided polyimide tape was used as the spacer on two side ends of the glass substrate. A printed photomask with a serpentine network pattern was placed on the solution and UV (λ = 365 nm, 10 mW) was irradiated for 6 s. The uncrosslinked solution was washed with D.I. water. Ag NW solution was diluted with ethanol to 0.05 wt%. The Ag NW solution was spraycoated on the gel pattern. The solution volume of the Ag NWs was adjusted according to the applications, 8 ml purpose of on-skin sensor. A solution of 8.0 mL was used for the ECG sensor electrode, while 4.0 mL was used for the haptic sensor. After spray-coating, the samples were ere thermally treated at 160 ℃ for 2 h in vacuum condition. Finite element method simulation. Simulations of stress distributions were performed using commercial FEM in ABAQUS (v6.9, Simulia) package. Failure stress was obtained from the fracture stress in the uniaxial tensile stress-strain curve. Von Mises stress and 4-node bilinear plane stress quadrilateral (CPS4) mesh were exploited for the simulation. Each component of stress failure criterion was assumed to be the same (σ11 = σ 22 = σ 12). The serpentine PEG-DA gel networks were stretched by 40 % uniaxial strain. Fabrication of the ECG electrodes and sensing of ECG. Three PEG gel electrodes were used as the ECG electrodes. The PEG gel electrodes were transferred to nylon fabric bands. A solution of two PEG-DA polymers (PEG-DA (Mn = 700 g/mol) : PEG-DA (Mn = 8000 g/mol) : HMOPP = 19 : 1 : 2) was used as a binder between nylon and the PEG electrodes. The PEG-DA solution was dropped on the edge of the gel electrodes and exposed to UV light (365 nm, 10 mW) for 30 s. Two electrodes were attached to the left forearm and another electrode was attached to the right side. All electrodes were connected to the physiograph (P800, Physiolab). ECG signals were measured by digital 8 channel physiograph (P800, Physiolab).


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Test of the haptic sensor. The PEG gel strain sensor was attached on the knee to measure the bending angle. Medical tape (Sportstape, Kinesiology Tape) was used at the edges of strain sensor to fix the sensor. Resistance data was measured by Analog-to-Digital converter (ADC) and Microcontroller unit (MCU) (KIRO). The data was delivered to a mobile device through RF communication. Mobile application (Wookyoung Information Technology) converted the electrical data to bending angle data, and it was displayed on portable device to monitor body motion in real time. Characterization. The stress-strain curves were obtained using a tensile stress tester (TST350E, Linkam Scientific Instruments). Electrical properties and mechanical properties were measured simultaneously with universal measurement probe system (UMP100, Teraleader, Inc.). The electrical performance of the devices was measured with a sourcemeter (Keithley-2450).

ASSOCIATED CONTENT Supporting Information. The magnified electrocardiogram signals in various conditions, Electrocardiogram signals measured with the conventional Ag/AgCl electrode (PDF). On-skin knee bending test (AVI). Wireless sensing of haptic device (AVI).

AUTHOR INFORMATION Corresponding Author


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*E-mail: [email protected] Author Contributions ‡

H.W.K and T.Y.K contributed to this work equally

Notes The authors declare no competing financial interest.

ACKNOWLEDGMENT This work was supported partially by the Center for Advanced Soft Electronics funded by the Ministry of Education, Science and Technology as a “Global Frontier Project” (CASE2015M3A6A5072945), by Basic Science Research Program (2017R1A4A1015811) through the National Research Foundation of Korea (NRF), and by the Sports Promotion Fund of Seoul Olympic Sports Promotion Foundation from Ministry of Culture, Sports and Tourism (s072016r2112016A0, 2016). We acknowledge Wookyoung Information Technology for allowing us to use their haptic motion analysis software.

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Hygroscopic Auxetic On-Skin Sensors for Easy-to-handle Repeated

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