Skin-Attachable, Stretchable Electrochemical Sweat Sensor for

Research Article Cite This: ACS Appl. Mater. Interfaces 2018, 10, 13729−13740

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Skin-Attachable, Stretchable Electrochemical Sweat Sensor for Glucose and pH Detection Seung Yun Oh,† Soo Yeong Hong,‡ Yu Ra Jeong,‡ Junyeong Yun,‡ Heun Park,‡ Sang Woo Jin,† Geumbee Lee,† Ju Hyun Oh,‡ Hanchan Lee,‡ Sang-Soo Lee,†,§ and Jeong Sook Ha*,†,‡ KU-KIST Graduate School of Converging Science and Technology and ‡Department of Chemical and Biological Engineering, Korea University, Seoul 02841, Republic of Korea § Photoelectronic Hybrids Research Center, Korea Institute of Science and Technology, Seoul 02792, Republic of Korea Downloaded via MIDWESTERN UNIV on January 24, 2019 at 11:27:05 (UTC). See https://pubs.acs.org/sharingguidelines for options on how to legitimately share published articles.

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ABSTRACT: As part of increased efforts to develop wearable healthcare devices for monitoring and managing physiological and metabolic information, stretchable electrochemical sweat sensors have been investigated. In this study, we report on the fabrication of a stretchable and skin-attachable electrochemical sensor for detecting glucose and pH in sweat. A patterned stretchable electrode was fabricated via layer-by-layer deposition of carbon nanotubes (CNTs) on top of patterned Au nanosheets (AuNS) prepared by filtration onto stretchable substrate. For the detection of glucose and pH, CoWO4/CNT and polyaniline/CNT nanocomposites were coated onto the CNT−AuNS electrodes, respectively. A reference electrode was prepared via chlorination of silver nanowires. Encapsulation of the stretchable sensor with sticky silbione led to a skin-attachable sweat sensor. Our sensor showed high performance with sensitivities of 10.89 μA mM−1 cm−2 and 71.44 mV pH−1 for glucose and pH, respectively, with mechanical stability up to 30% stretching and air stability for 10 days. The sensor also showed good adhesion even to wet skin, allowing the detection of glucose and pH in sweat from running while being attached onto the skin. This work suggests the application of our stretchable and skinattachable electrochemical sensor to health management as a high-performance healthcare wearable device. KEYWORDS: stretchable electrochemical sensor, sweat sensor, wearable electronics, nanomaterial-based electrode, filtration, layer-by-layer

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INTRODUCTION In accordance with greatly increased interest in healthcare management, fitness, and biomedicine, there have been extensive efforts to develop advanced wearable electronic devices related to energy harvesting, energy storage, and biosensors, which can be worn or attached onto skin.1−5 In particular, great advances in wearable biosensors to monitor the individual’s physiological state have been reported.6 Among the various wearable biosensors, electrochemical sensors for monitoring vital signals via the detection of physiological components in sweat have been paid special attention because they are noninvasive, with the advantages of minimizing skin irritation and ease of measurement.7 Human sweat, an important body fluid that contains metabolites and electrolytes, provides rich information about our health and fitness condition.8 Therefore, sensors capable of accurate and facile measurements of sweat components can serve as an important disease diagnosis device. Diabetes is a chronic disease caused by insufficient insulin production or secretion that many people suffer from.9 As diabetes causes many critical complications, the patient’s © 2018 American Chemical Society

glucose level needs to be continuously managed. Usually, diabetes can be diagnosed from analysis of the glucose level in blood. Also, hyperglycemia occurs when glucose concentration in blood is above 126 mg dL−1 under fasting.10 However, detection of blood glucose level is irritating because it requires bleeding with a needle. A correlation of the glucose level between blood and sweat has been reported that 0.3 mM glucose in sweat corresponds to 300 mg dL−1 glucose in blood.11 Thus, it is expected that detecting the glucose level in sweat can be used for diabetes diagnosis. In addition, pH gives further important information on health conditions.12 For example, skin diseases, such as dermatitis and fungal infections, can cause a change of pH in sweat.13,14 Therefore, measurement of the glucose concentration and pH in sweat can provide an important assessment to determine individual health conditions. Received: February 26, 2018 Accepted: April 6, 2018 Published: April 6, 2018 13729

DOI: 10.1021/acsami.8b03342 ACS Appl. Mater. Interfaces 2018, 10, 13729−13740

Research Article

ACS Applied Materials & Interfaces To accurately measure the glucose and pH in sweat, sensors are required to conformally attach onto skin and also to be stretchable for mechanical stability upon deformation due to skin movement.15 Recently, various stretchable devices have been reported, where the mechanical stability was obtained via adopting island-bridge,16 serpentine,17 and wavy structures18 and using materials of liquid metal19 and carbon nanotube (CNT).20−22 In particular, a highly stretchable electrode could be fabricated using a percolation network of one-dimensional (1D) Ag nanowires (AgNW).23 However, AgNW has the disadvantage of being easily oxidized in air or water,24 so they are difficult to use as electrode materials. On the other hand, two-dimensional (2D) gold nanomaterials not only have excellent electrical and chemical stability, but can also be overlapped to form percolation networks.25 Stretchable sweat sensors could be fabricated via conventional evaporation and photolithography techniques.1,26 However, these techniques have the demerits of requiring high vacuum and are timeconsuming and expensive. Unlike these techniques, the filtration method has the advantages of low cost, low temperature, simplicity, and fast process.27 Electrochemical glucose sensors can be categorized into two types: enzyme and nonenzyme sensors.28 Typically, enzyme sensors composed of modified glucose oxidase are widely used for glucose determination. However, several drawbacks exist of being sensitive to environmental condition, as well as the high cost of enzyme, the immobilization process, and poor stability.29 To resolve these problems, nonenzyme sensors using various nanomaterials have recently been investigated.30 However, nonenzyme sensors using metals or alloys, such as Pt and Pt−Pb alloy, are known to be expensive and poisonous.31,32 Transition-metal oxides,33−36 such as NiO or Co3O4, can solve these problems and show high sensitivity. Among these transition-metal oxides, mixed transition-metal tungstate has attracted much attention because of its extraordinary properties, which are due to the various valence states of the W atom.37 In addition, tungstate material has the advantages of facile synthesis, low price, low toxicity, and high stability.38 A nonenzyme sensor based on cobalt tungstate nanomaterials, such as CoWO4 nanosphere39 and CoWO4 microring,40 was reported to detect glucose and exhibit good electrocatalytic activity. However, in these previous studies, glucose was detected using CoWO4 nanomaterials deposited on a rigid glassy carbon electrode in the base electrolyte. Therefore, these sensors are difficult to apply to the skin-attachable sweat sensor. Also, it is known that nonenzymatic glucose sensors using metals such as Pt and Ni show poor selectivity.30,41 Therefore, various studies have been reported to solve these problems. Biosensors based on nanomaterials or nanoporous structure have been proposed to enhance selectivity by increasing the signal due to their large surface area.42 It is also known that electroactive species, such as ascorbic acid (AA), dopamine, and uric acid, in body fluid are oxidized above 0.3 V in neutral electrolyte.43,44 Therefore, selectivity can be improved by reducing the interference current when glucose detection is performed at low potential. Generally, ion-selective electrodes (ISEs) are most widely used for pH detection. Polyaniline (PANI) is known as a pH-sensitive material because it can detect hydrogen ion by deprotonation on its surface.13 In addition, PANI-based ISEs are expected to be used as wearable pH sensors due to their low cytotoxicity and low skin irritation. Recently, Kim’s research group has reported on a stretchable electrochemical device for diabetes monitoring and therapy,26

where a conductive electrode was fabricated using conventional photolithography and chemical vapor deposition (CVD) and an enzyme like glucose oxidase was used for glucose detection. For stretchability of the device, serpentine interconnections were adopted, whereas the sensor itself was not flexible as a fixed island. In this work, we report the fabrication of a nanomaterialbased skin-attachable and stretchable electrochemical sensor that is capable of detecting glucose and pH in human perspiration. A percolation network was formed on the stretchable substrate via filtering AuNS to make intrinsically stretchable sensor without serpentine or island-bridge structures. In this way, the stretchable electrode can be easily fabricated at low temperature without use of complicated and multistep processes, such as CVD and photolithography. In addition, layer-by-layer (LbL) deposition of CNT was performed to modify working electrodes because this process has the advantages of simplicity, low cost, and control of the film thickness according to the number of LbL cycles.45,46 Therefore, the LbL-CNT film was deposited on the AuNS electrode. The working electrode for glucose detection was prepared by depositing CoWO4/CNT nanocomposite on top of LbL-CNT/AuNS via LbL deposition without use of an enzyme. CoWO4 nanoparticles were hydrothermally synthesized on the surface of CNT with high conductivity and large surface area. The PANI-based electrode for pH detection was prepared via electropolymerization on LbL-CNT film. In addition, AgNW synthesized by the polyol method was dropped on the AuNS electrode to prepare the solid-state stretchable reference electrode, and chlorination was then performed. The fabricated electrochemical sensor encapsulated with sticky polymer, silbione, was conformally attached onto skin and showed highly selective sensitivity to both glucose and pH in sweat without any degradation, even under repetitive deformations of stretching by 30%. The stretchability of 30% is sufficient for the sweat sensor attached to the skin and to operate stably against the movement of the skin. Continuous monitoring of the change in glucose concentration and pH due to meal ingestion and running could be performed for long duration with our skin-attached sweat sensor. Therefore, glucose and pH values could be stably detected in sweat after exercising when the sensor was attached onto skin, regardless of the skin movements. These results indicate the high potential application of our sweat sensor to personal physiological monitoring, especially of glucose and pH, when attached onto skin for long duration.

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EXPERIMENTAL PROCEDURES

Fabrication of a Stretchable Patterned AuNS Electrode. For the synthesis of AuNS, 1.36 mg of L-arginine (Aldrich) was dissolved in 10 mL of distilled water and the solution was heated to 95 °C. Meanwhile, 4 mL of 17 mM HAuCl4·3H2O (Aldrich) was quickly injected into the L-arginine solution. The mixture solution was maintained at 95 °C for 2 h, cooled down to room temperature in air, and AuNS solution was then obtained. Polydimethylsiloxane (PDMS, Sylgard 184, Dow Corning) mask was prepared by mixing a base and curer in a weight ratio of 10:1. The mixture was degassed in a vacuum desiccator to eliminate air bubbles and then cured in an oven at 65 °C for 2 h. The cured PDMS was cut into the desired shape. To fabricate a patterned AuNS electrode, the PDMS mask was put on the polycarbonate membrane filter (Sterlitech; pore size, 0.2 μm; diameter, 47 mm) and AuNS solution was filtered through it so that the patterned AuNS electrode with percolation network structure could be prepared. The number of filtrations of the synthesized AuNS solution 13730


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ACS Applied Materials & Interfaces

Figure 1. Schematic of the fabrication processes of the skin-attachable, stretchable electrochemical sweat sensor for glucose and pH detection. (a) Fabrication of the stretchable AuNS electrode via filtration. (b) Fabrication of the CNT film on the AuNS electrode via the LbL method. (c) Fabrication of the skin-attachable, stretchable electrochemical sweat sensor. was controlled to optimize the electrode. After peeling off the PDMS mask, the filtered membrane filter was dried in an oven. An uncured PDMS mixture (10:1) was spin-coated on the filtered membrane filter at 800 rpm for 30 s and cured in an oven at 65 °C for 1 h. Finally, the filtered membrane filter was peeled off from the PDMS to get the stretchable patterned AuNS electrode. Functionalization of CNT. To functionalize CNT with negatively charged carboxyl group (−COOH), CNTs (Aldrich; length, 5−9 μm; outer diameter, 110−170 nm) were dispersed in a solution of sulfuric acid (Aldrich) and nitric acid (Aldrich) in a mixed volume ratio of 3:1 and the mixture solution was refluxed at 60 °C for 3 h at 200 rpm. The refluxed solution was cooled down to room temperature and filtered to remove excess acid using mixed cellulose ester membrane filter (Advantec; pore size, 0.2 μm; diameter, 47 mm). Then, osmosis filtration was carried out through a tube cellulose membrane (Aldrich; average flat width, 33 mm; average diameter, 21 mm) for a week. After 1 week, the mixture solution was dried and negatively charged CNTs (CNT-COOH) were prepared. To functionalize the CNTs with

positively charged amine group (−NH 2 ), 1-ethyl-3-(3dimethylaminopropyl)carbodiimide (Alfa Aesar) was added into the CNT-COOH solution. Then, ethylenediamine (Aldrich) was injected and stirred for 6 h. After that, the solution was washed by the filtering process and the positively charged CNT (CNT-NH2) was prepared. Preparation of Patterned LbL-CNT Film for Working Electrode. The patterned AuNS electrode was dipped into 18 mM cysteamine (Aldrich) for 4 h to make it positively charged with a selfassembled monolayer (SAM). The electrode was then rinsed with distilled water and dried with nitrogen gas. To deposit active materials on the electrode pad, a poly(ethylene terephthalate) (PET) mask with the desired pattern was prepared and attached to the SAM-treated AuNS electrode. The AuNS electrode was immersed in the CNTCOOH solution for 20 min and rinsed with distilled water. The same procedure was followed in the CNT-NH2 solution. After five cycles of the LbL process, the CNT film for working electrode was successfully prepared. Finally, the patterned LbL-CNT film could be obtained by peeling off the mask. 13731


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UV-1800) was used to measure the absorbance of LbL-CNT film according to the number of cycles of LbL deposition. ζ-Potential of nanomaterial was measured by a ζ-potential analyzer (Otsuka Electronics, ELSZ-1000). The electrochemical performance was analyzed by probe measurement (HP 4140B, Hewlett Packard) and potentiostat (Ivium Technologies, CompactStat). The performance of the electrochemical sensor for data collection was measured after stabilization period of equilibrium time (10 s) in electrolyte. Similarly, data collection was done after equilibrium time when the concentration of the electrolyte was changed. The mechanical stability of the fabricated sensor was measured using computer software (PMC1HS, Autonics Corp., Korea) and a custom-built stretching stage. Photographic images were obtained using a cell phone camera.

Synthesis of CoWO4 Nanoparticles and CoWO4/CNT Nanocomposite. To synthesize CoWO4/CNT nanocomposite, 20 mg of CNT-COOH was dispersed in a Teflon-lined stainless steel autoclave, containing 40 mL of distilled water. Co(NO3)2·6H2O (0.291 g, Alfa Aesar) was dispersed in the CNT-COOH solution. After adding 0.37 g of Na2WO4·2H2O (Aldrich) in the mixture solution with vigorous stirring, it was heated in an oven at 180 °C for 8 h. After cooling to room temperature, the dark blue precipitates were washed with distilled water several times and dried in an oven at 60 °C. Finally, the CoWO4/CNT nanocomposite was obtained. In addition, CoWO4 nanoparticles were obtained by the same method without CNT. Synthesis of AgNW. AgNW was synthesized by a polyol method. CuCl2·2H2O (40 μL, 4 mM; Alfa Aesar) was dispersed in 5 mL of ethylene glycol (Aldrich) in an oil bath at 152 °C with stirring at 260 rpm for 15 min. After sequential injection of 1.5 mL of 0.147 M poly(vinylpyrrolidone) (Aldrich) and 1.5 mL of 0.094 M AgNO3 (Aldrich) to the solution, the mixture solution was heated at 152 °C with stirring at 260 rpm for 1 h and then cooled down to room temperature. Finally, the mixture solution was washed with distilled water three times, yielding AgNW dispersed in distilled water. Fabrication of a Stretchable Electrochemical Sensor. The electrochemical sensor consists of a two-electrode system. The working electrode is composed of two parts to detect glucose and pH. To fabricate the working electrode for glucose sensing, PET mask with desired pattern was attached onto the LbL-CNT film. The LbLCNT film was immersed in 2 mg mL−1 CoWO4/CNT nanocomposite solution for 20 min, rinsed with distilled water, and dried with nitrogen gas. Finally, patterned electrode with the deposited CoWO4/CNT could be obtained by peeling off the PET mask. Another working electrode for the pH sensor was fabricated by electrodeposition of polyaniline (PANI) via a potentiodynamic deposition technique. Potentiodynamic deposition was carried out for 30 cycles at a scan rate of 100 mV s−1 in a solution of 0.1 M aniline (Aldrich) and 0.5 M H2SO4 (Aldrich) between 0 and 0.85 V. Here, Pt wire electrode and Ag/AgCl (3 M NaCl) electrode were used as the counter and reference electrodes, respectively. Finally, the PANI-based electrode for the pH sensor could be fabricated. To fabricate the solid-state Ag/ AgCl reference electrode, 40 μL of AgNW synthesized via the polyol method was dropped onto the electrode pad and annealing was carried out at 90 °C for 5 min. The Ag/AgCl electrode was prepared by dipping in 0.05 M FeCl3 (Aldrich) for 10 s through a chlorination process. Poly(vinyl butyral) (PVB) stock solution was prepared by dissolving 79.1 mg of PVB (BUTVAR B-98, Aldrich) and 50 mg of NaCl (Aldrich) in 1 mL of methanol (Aldrich).47 The solid-state reference electrode was prepared by drop-casting 40 μL of the stock solution on the Ag/AgCl electrode. A mask was put on the electrode pad of the stretchable electrochemical sensor. Silbione (Silbione RT Gel 4717 A&B, Bluestar Silicones) mixed at a 1:1 ratio was spin-coated (2000 rpm, 30 s) and cured on a hot plate at 65 °C for 5 min. After peeling off the mask, it was possible to cover the electrical contact to enable the attachment to the skin, except for the working and reference electrode pads for sensing. To connect prepared sensors with electrochemical analyzer, the bottom end of each electrode made a contact with a copper wire using silver paste. And then, a polyimide tape was used to prevent electrical contact between the skin and silver paste. Characterization. Optical images were obtained by optical microscopy (Olympus, eXcope T300). The surface morphology and cross-sectional image were investigated by scanning electron microscopy (SEM, Hitachi, S-4800). The chemical composition of the electrodes was analyzed by energy-dispersive spectrometry (EDS, Horiba, EX-200). The size and morphology of the nanomaterials were characterized by transmission electron microscopy (TEM, Hitachi, H7100). The dimensions and morphology of the AuNS were measured using atomic force microscopy (AFM, XE-100, PSIA). Characterization of the synthesized CoWO4, CoWO4/CNT, and PANI/CNT was performed by X-ray diffraction (XRD, Rigaku, SmartLab), Raman spectroscopy (laser (Omicron) at 532 nm) (HORIBA, LabRam ARAMIS IR2), and X-ray photoelectron spectroscopy (XPS, ULVACPHI, X-TOOL). UV/visible spectrophotometry (UV/vis, Shimadzu,

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RESULTS AND DISCUSSION Figure 1 shows a schematic of the fabrication processes of the stretchable and skin-attachable electrochemical sensor for detecting glucose and pH in sweat. Figure 1a shows the fabrication of stretchable Au nanosheets (AuNS) electrodes via a filtration method. The synthesized AuNS was vacuum-filtered using a patterned mask. Scanning electron microscopy (SEM) (Figure S1a) and optical microscopy (Figure S1b) images taken from the AuNS synthesized by wet process show that the AuNS has a lateral size of 15 μm. Figure S1c shows that the thickness of the synthesized AuNS estimated using atomic force microscopy (AFM) is 60 nm. These measurements confirm that the AuNS has a two-dimensional structure. A stretchable patterned AuNS electrode was fabricated via vacuum filtration because the filtration process is advantageous in that it is simple, fast, and cost-effective. After putting the patterned PDMS mask on top of the polycarbonate membrane filter, the AuNS solution was dropped under vacuum suction for filtration of the AuNS. Vacuum suction during the filtration process improves the resolution of the pattern because of the intimate contact between the mask and membrane filter. As the number of filtrations increases, the filtered AuNS film becomes thicker on the membrane filter. After removal of the PDMS mask, uncured PDMS (10:1) is spin-coated on the filtered membrane and then cured. Finally, the membrane filter is peeled off, resulting in a patterned AuNS electrode with percolation network structure on the stretchable substrate. After curing of the spin-coated PDMS on the patterned AuNS electrodes and subsequent peeling off of the membrane filter, an array of stretchable AuNS electrodes is obtained. With increase of the filtration number of AuNS, the sheet resistance decreases, as shown in the top graph of Figure S2a. One-time filtration gives a low density of AuNS to provide large sheet resistance with a large error bar, but after multiple filtration processes, dramatic decrease in resistance is observed due to the formation of a high-density percolation network of AuNS uniformly covering the PDMS substrate. After filtration four times, the sheet resistance is 1.4 ± 0.58 ohm sq−1, sufficiently conductive to act as a current conductor. The inset SEM image of the top graph in Figure S2a shows that the AuNS forms a percolation network structure after filtering four times. The thickness of AuNS gradually increases with increase of the filtration frequency reaching 13.4 ± 3.6 μm after filtering four times (bottom graph in Figure S2a), which is confirmed in the crosssectional SEM image. With the AuNS film formed after filtration four times, repetitive stretching cycles up to 50% were performed, as shown in Figure S2b. The resistance increased by 6 and 11% compared to the initial resistance under 30 and 50% stretching, respectively. After 1000 cycles of repetitive stretching by 30%, no notable change of resistance was 13732


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Figure 2. Characteristics of active materials for the working electrodes. (a) (Top) TEM image of CoWO4 (inset: size distribution of CoWO4) and (bottom) XRD pattern of CoWO4. (b) (Top) TEM image of CoWO4/CNT (inset: high magnification) and (bottom) XPS survey scans of CoWO4 and CoWO4/CNT. (c) (Top) SEM image of PANI/CNT (inset: high magnification) and (bottom) Raman spectra of PANI/CNT and CNT.

to the reflections of the (0 0 1), (−1 1 1), (2 0 0), (−1 2 2), and (−2 3 1) planes, respectively.48,49 The observed diffraction peaks agree well with the standard pattern of the CoWO4 structure phase (JCPDC card no. 15-0867). In addition, it can be confirmed that high-purity CoWO4 nanoparticles are synthesized because there are no peaks of impurities or other residuals. Figure 2b shows the TEM image taken from the CoWO4/CNT nanocomposite, which indicates that CoWO4 nanoparticles are synthesized on the surface of CNT. Here, the CoWO4/CNT nanocomposite was synthesized using the same method used for the synthesis of CoWO4 nanoparticles by adding CNT. During the hydrothermal process of the CoWO4/ CNT nanocomposite, Co2+ cations were adsorbed on the negatively charged CNT via simple electrostatic interaction and then CoWO4 nanoparticles could be hydrothermally crystallized. XPS measurements were performed to investigate the composition and chemical bonding state. XPS survey scans of the CoWO4/CNT nanocomposite and CoWO4 nanoparticles are shown. Co 2p, Co 3s, O 1s, W 4p, W 4d, and W 4f peaks are observed in both samples. Figure S4 shows XPS narrow scans for individual components taken from the CoWO4/CNT nanocomposite. Figure S4a shows peaks at 781.2 and 796.5 eV corresponding to Co 2p3/2 and Co 2p1/2, respectively. In addition, the spectrum of Co 2p exhibits two satellite peaks at 785.6 and 802.8 eV.50 Moreover, Figure S4b shows two peaks at 35.1 and 37.2 eV corresponding to W 4f7/2 and W 4f5/2.49 From these results, cobalt ions are considered to exist in the divalent state and tungstate ion exists as +6 valence state.49,51 Figure S4c shows two peaks of O 1s at 529.8 and 531.1 eV, corresponding to the metal−oxygen bond and −OH bond.52 In the element analysis, Co and W exist in the ratio of 1:1, which indicates the formation of CoWO4 nanoparticles. Figure S4d shows the C 1s core-level spectrum. The C 1s peak appearing at 284.4 eV is attributed to sp2-hybridized CC bond. Three peaks are additionally observed at 285.1, 286.2, and 287.5 eV, which correspond to sp3-hybridized C−C bond, C−O bond, and CO bond.48

observed. These results suggest that the patterned AuNS electrode with a percolation network can be used as a stretchable electrode. Figure S2c shows photographic images taken from the fabricated stretchable patterned AuNS electrode under stretching. The electrode pad has a diameter of 4 mm, and the substrate has a size of 1.5 cm × 1.8 cm × 100 μm. No cracking on the electrode appeared under stretching. Figure 1b shows that deposition of multiwalled CNT onto two AuNS electrodes is performed via a layer-by-layer (LbL) technique. Alternate dipping of the fabricated stretchable AuNS electrodes into negatively charged and positively charged CNT solutions results in multilayer CNT film on the AuNS electrode. The thickness of the CNT film can be controlled by adjusting the number of LbL cycles. The UV/vis absorbance at 250 nm was measured every LbL cycle and found to linearly increase by bilayer, as shown in Figure S3a. Figure S3b shows that the thickness of the film is estimated to be 1.8 μm from the top- and cross-sectional view SEM images taken of the fivebilayer CNT film. Figure 1c shows that next, two sensing electrodes and a reference electrode are fabricated following the sequential processes. First, a CoWO4/CNT nanocomposite film is coated on a CNT/AuNS electrode for glucose sensing. Second, PANI is electrodeposited onto the other CNT/AuNS electrode for pH sensing. Third, the solid-state Ag/AgCl reference electrode is prepared on the AuNS electrode using Ag nanowires synthesized by the polyol method and subsequent chlorination process using FeCl3, followed by PVB coating. Finally, the whole device, except for the three electrodes, is encapsulated with silbione, resulting in a stretchable and skin-attachable electrochemical sweat sensor. More detailed experimental procedures are given in Experimental Procedures. Figure 2a shows the TEM image taken of the hydrothermally synthesized CoWO4 nanoparticles, from which the average size of the particle is estimated to be 36 ± 6 nm. The XRD pattern shows that crystalline CoWO4 particles are synthesized; the diffraction peaks at 18.9, 30.5, 36.2, 54.1, and 65.1° are assigned 13733


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Figure 3. Electrochemical performance of the glucose sensor. (a) Chronoamperometry curves with different concentrations of glucose from 0.05 to 0.3 mM (inset: calibration plots for the glucose sensor). (b) Comparison of the sensitivity of glucose sensors fabricated by drop-casting (blue) and the LbL (red) method. (c) Selective detection of glucose. (d) Long-term stability and multiple use of the glucose sensor (applied potential: −0.2 V vs commercial Ag/AgCl reference electrode.).

(Figure S6a) was used to estimate the length and diameter of AgNW to be 22 ± 4 μm and 68 ± 13 nm, respectively. The AgNW coated on AuNS electrode clearly shows the percolation network in the SEM image of Figure S6b. Figure S6c shows the EDS images of AgNW on the AuNS electrode, uniformly covering AgNW on AuNS electrode. Figure S7a shows the SEM image taken from the Ag/AgCl electrode. The EDS image confirms the existence of chlorine, indicating the successful chlorination in Figure S7b. Such prepared Ag/AgCl electrode was tested to see if it can be used as a reference electrode. Figure S8a compares the open-circuit potential (OCP) value of AuNS electrode before and after Ag/AgCl deposition vs commercial Ag/AgCl (3 M NaCl) reference electrode. Unlike bare AuNS electrode, the Ag/AgCl-deposited electrode shows no potential difference from the commercial Ag/AgCl reference electrode. In addition, the potential stability is demonstrated with variation of ion concentration and ionic species before and after coating of PVB. It was reported that PVB coating on reference electrode plays a role in maintaining the stable potential of the electrode regardless of the concentration of various ions because the PVB membrane allows the exchange of electrolytes through the membrane.56 Figure S8b shows that the OCP value of Ag/AgCl electrode without PVB coating changes with the concentration of NaCl solution. Figure S8c shows that the PVB-coated Ag/AgCl electrode reveals stable potential in various electrolytes and different concentrations of NaCl solution, unlike the Ag/AgCl electrode without PVB coating, which indicates that the prepared solid-state Ag/AgCl electrode can be used as a reference electrode. The mechanical stability of each prepared electrode is investigated. Figure S9a shows that under stretching of up to 50%, the three different electrodes show almost no change of resistance. Moreover, Figure S9b confirms that each electrode is mechanically stable without change in resistance after 1000 cycles under 30% strain.

CoWO4/CNT nanocomposite was deposited onto the LbLCNT film/AuNS to fabricate a working electrode for glucose detection. At the end of the LbL process, the LbL-CNT film has a positive charge and the CoWO4/CNT solution has negative ζ-potential, as shown in Table S1. When the LbLCNT film is dipped in CoWO4/CNT solution, negatively charged CoWO4/CNT composites can be bound via electrostatic interaction. Figure S5a shows the SEM morphology of the CoWO4/CNT nanocomposite deposited onto LbL-CNT film. In addition, Figure S5b shows the electrochemical performance of bare AuNS electrode and CoWO4, CNT, and CoWO4/CNT on the AuNS electrode by measuring the cyclic voltammetry (CV) curves for Fe(CN)63−/4− redox reaction. The prepared AuNS electrode with CoWO4/CNT shows the best electrochemical performance, probably owing to its large electrochemically active area. PANI was deposited via electrodeposition on LbL-CNT film for the pH sensor. Figure 2c shows the SEM image that reveals the morphology of the fabricated PANI/CNT. In the inset SEM image, PANI/CNT can be identified, showing the CNT wrapped with PANI. The Raman spectra of the CNT show the characteristics of D band (ID) and G band (IG) at 1330 and 1561 cm−1, respectively. The peak of ID indicates the vibrations of sp3-bonded carbon atom of defects and disorder, whereas the peak of IG indicates sp2bonded carbon atom.53 In addition, the ID/IG ratio is 0.38, which indicates the high crystallization of CNT. In the case of PANI/CNT, the Raman peaks at 1150, 1243, and 1467 cm−1 are attributed to the C−H bending vibration of quinoid rings, C−H bending vibration of benzene rings, and N−H bending vibrations, respectively.54,55 The results confirm that PANI/ CNT was successfully synthesized on the LbL-CNT film. To fabricate a solid-state Ag/AgCl reference electrode, AgNW synthesized by the polyol method was used with a chlorination process. The TEM image taken from the synthesized AgNW 13734


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Figure 4. Electrochemical performance of the pH sensor. (a) Open-circuit potential curves with different pHs from 4 to 8 (inset: calibration plots for the pH sensor). (b) Comparison of the sensitivity of pH sensors fabricated by PANI (blue) and PANI/CNT (red). (c) Selective sensing of pH. (d) Long-term stability and multiple use of the pH sensor (OCP vs commercial Ag/AgCl reference electrode).

the sensitivity of the glucose sensor fabricated via the LbL method is 2.9 times higher than that fabricated via the dropcasting method. This is attributed to the better electrochemical performance between the AuNS electrode and the active materials in the case of using the LbL method through electrostatic interactions. In addition, Figure 3c shows the selectivity of the sensor to detect 0.3 mM glucose over other interference agents, such as 10 μM ascorbic acid (AA), 50 μM uric acid (UA), 0.1 mM urea, and 0.1 mM acetaminophen. This figure clearly reveals that the current does not change after the addition of the interference agent. In general, nanoparticles such as CuWO4 and SnWO4 are known to have negative surface charge due to negative ζ-potential values.58,59 Likewise, CoWO4 nanoparticles also have negative surface charges due to negative ζ-potential, as shown in Table S1. As a result, the glucose sensor with CoWO4 nanoparticles has high selectivity by repelling negatively charged interference agents, such as AA and UA. Also, our sensor shows good selectivity probably because the glucose detection was performed at low potential to reduce the oxidation of interference agents. Figure 3d shows the long-term stability of our glucose sensor for up to 10 days. After each measurement per day, the glucose sensor was washed with distilled water and then reused the next day. It was possible to detect glucose without any deterioration in the sensitivity for 10 days, and the sensor could be reused multiple times. The performance of our glucose sweat sensor is compared to that of others previously reported in Table S2. Figure 4 analyzes the electrochemical performance of the fabricated PANI-based pH sensor at room temperature. pH detection was carried out by measuring the OCP value with variation of pH from 4 to 8 using the standard solution. The change of OCP with pH shown in Figure 4a is used to obtain the linear calibration curve shown in the inset figure. The sensitivity of 71.44 mV pH−1 is estimated from the slope of the calibration curve, as shown in Figure 4b, which exceeds the

These results demonstrate that each electrode with active material can be used as a stretchable sensing electrode. Such high mechanical stability is attributed to the percolation network formed in the 1D and 2D nanomaterial films. Figure 3 analyzes the electrochemical performance of the fabricated electrochemical sensor for glucose detection at room temperature. To determine the operation potential for glucose detection by chronoamperometry, CV curves were measured to locate the peak using CoWO4/CNT electrode without and with 0.3 mM glucose in phosphate-buffered saline (PBS, Aldrich) solution, as shown in Figure S10. In the absence of glucose, no peak was observed in the CV results, whereas the oxidation peak occurred at −0.2 V when glucose was added. Therefore, glucose detection was performed by choosing −0.2 V in chronoamperometry. It was reported that CoOOH is generated from CoWO4 by electrochemical reaction57 and subsequently CoO2 is produced from CoOOH.39,40 Then, glucose is oxidized by CoO2. The corresponding reactions can be explained in the following equations CoOOH + OH− ↔ CoO2 + H 2O + e− CoO2 + C6H12O6 → CoOOH + C6H10O6

Figure 3a shows the chronoamperometry curve for successive addition of glucose to the PBS solution. The concentration of glucose was adjusted to 0−0.3 mM, as measured in sweat. The inset of Figure 3a shows the linear calibration curve obtained from the current density according to glucose concentration. The sensitivity is estimated to be 10.89 μA mM−1 cm−2 by linear least-squares fitting of the calibration data, as shown in Figure 3b. And the limit of detection (LOD) is estimated to be 1.3 μM according to the formula LOD = 3Sb/S,37 where Sb is the standard deviation of the blank signal and S is the sensitivity. Figure 3b shows the effect of the deposition method of CoWO4/CNT nanocomposite onto AuNS on the sensitivity; 13735


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Figure 5. (a) Photographic images taken of the sensor under stretching of up to 30%, and the definition of strain (ε). Characteristics of the (b) glucose sensor and (c) pH sensor, both under stretching deformation; (top) sensitivity at different applied strains and (bottom) stability under repetitive stretching cycles under 30% strain (applied potential: −0.2 V vs solid-state Ag/AgCl reference electrode; OCP vs solid-state Ag/AgCl reference electrode).

showing that the sensitivity of our sensors depended on temperature. For the reversibility test of the sensors, we measured performance of our sensors in a reversible way. The experimental conditions were kept the same as in Figures 3a and 4a. In the case of our pH sensor, the signal was measured along with the consecutive changes from pH 4 to 6 (forward directions) and then measured again from pH 6 to 4 (reverse directions). Figure S12 shows the similar signals measured in the forward and reverse directions. On the contrary, in the case of our glucose sensor, it was not reversible because the signal obtained in reverse direction, i.e., from high to low glucose concentration, was different from that in forward direction, i.e., from low to high glucose concentration, unless cleaning the sensor with water after each use. However, we think that it is not a serious problem at this moment because we can use our glucose sensor several times while attaching it onto skin and washing after each use. Figure 5a shows photographic images taken from the fabricated electrochemical sensor under uniaxial stretching of up to 30% and defines the applied strain. Figure 5b,c shows the sensing performance of the sensor to glucose and pH under repetitive stretching/releasing cycles, respectively. The inset figures show the relative sensitivity S/S0, where S0 and S are the sensitivity before and after the application of strain, respectively. As shown, our sensor demonstrates very stable performance, without any degradation of the initial sensitivity under stretching of up to 30%. In addition, 1000 repetition cycles of stretching/releasing by 30% does not cause any notable deterioration of either glucose or pH sensors. Moreover, Figure S13 shows the high-magnification optical images of the sensor under stretching deformation. In highmagnification optical images, microscopic cracks are observed unlike in photographic images. With increase of the strain, increase of cracks is observed. However, our sensor showed

Nernst limit of pH sensitivity. And it was reported that the sensitivity of the PANI-based pH sensor and conductivity of PANI depend on the phases of PANI and the manufacturing process.13,60 Electrodeposition process is reported to exceed the theoretical Nernstian response unlike sputtering process. Therefore, the reason for exceeding the Nernst limit of pH sensor in our study seems to originate from the superior orientation of PANI grown by electrodeposition. The sensitivity of the PANI grown on LbL-CNT film is higher than that of PANI on AuNS electrode, as compared in Figure 4b. This seems to be attributable to the enhanced electronic conductivity from the composite of PANI and CNT.61 Figure 4c shows the interference effect against other ions, which include Cu2+, NH4+, K+, and Na+. The pH sensor shows good selectivity to pH without any change in OCP for other ions, except for hydrogen ions. Hydrogen ions can be detected by deprotonation on the surface of PANI, so PANI-based pH sensor does not react with other ions and shows high selectivity only to pH.26 Moreover, Figure 4d shows that the pH sensor is stable for long duration and under multiple use. The performance of our pH sweat sensor is compared to the previously reported sensors in Table S3. Electrochemical sensor is known to be affected by the temperature, so we also investigated the temperature dependence of our sensors. The performance of sensor was measured by changing the solution temperature from 20 to 40 °C. For the measurement of glucose and pH, PBS solution containing 0.2 mM glucose and pH 5 standard solution were used, respectively. Current density of the glucose sensor was measured by chronoamperometry method at −0.2 V, and the OCP value of the pH sensor was measured at each temperature. In Figure S11, S0 and S are the measured responses of each sensor at 20 °C and at each temperature, respectively. With increase of temperature, the performance of both glucose and pH sensors decreased, 13736


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Figure 6. (a) Photographic image of the electrochemical sensor attached to the skin wet with sweat. (b) Correlation of glucose concentration (left) and pH (right) measured by our sensors (x-axis) with those by a commercial glucose assay kit and commercial pH meter (y-axis) in human sweat. Here, the solid and open symbols in glucose sensing are from sweat before and after meal, respectively. The solid and open symbols in pH sensing are from sweat after exercise and sauna, respectively. (c) Performance of our sensor attached to the skin in human sweat under mechanical deformation (inset: corresponding photographic images under mechanical deformation; scale bar: 1 cm). (d) Change of glucose concentration and pH measured by our sensor attached onto skin along with different activities of meal injection (green pulse) and running (pink pulse) for 10 h. All of the data were calibrated with the temperature-dependent sensitivity at 30 °C.

Ag/AgCl reference electrode during glucose detection, our sensor was configured to perform consecutive measurements rather than simultaneous measurements. Temperature compensation is required to increase the accuracy of the sensor because the skin temperature is 30 °C. Therefore, we calibrated the original data taken from the sweat sensor attached to skin at 30 °C, following the measured temperature-dependent sensitivity curve (Figure S11). Figure 6c shows the detection of glucose and pH in sweat according to mechanical deformation, such as compression and stretching of the sensor attached to the skin. The insets in Figure 6c show photographic images of the sensor attached to the skin without deformation, stretching, or compression. The measured values of glucose and pH rarely change during compression or stretching of the sensor compared to those without any deformation. Additionally, the performance of our sensor was characterized under harsh mechanical deformations of torsion and pressing. As shown in Figure S15, our sensor showed strong adhesion to skin even in torsion and pressing modes. Furthermore, the performance of the sensor was not significantly changed. These results signify that our sensor attached to the skin exhibits very stable electrochemical performance even under skin deformation. Figure 6d shows long-term measurements of the sensor attached to the skin. The glucose concentration and pH value were measured in sufficient sweat after running for 20 min. After each measurement, the sensor was rinsed with distilled water and then reattached to the skin. It was observed that the concentration of glucose in sweat greatly increased at 30 min after meal ingestion and then gradually decreased with time, which is consistent with the behavior of glucose in blood; the previously study62 showed that the concentration of glucose in blood rapidly increased within 1 h after meal ingestion and thereafter it gradually decreased. These results show that the

very stable performance under stretching. Therefore, the microscopic cracks appearing on the sensor are considered not to deteriorate the sensor performance notably. These results indicate that our nanomaterial-based sensor can be used as a stretchable sensor. Figure 6a shows a photographic image of our stretchable electrochemical sensor attached onto the skin of an arm wet with sweat. The sensor encapsulated with silbione intimately adheres to the skin strongly. Moreover, Figure S14 shows that the sensor is still skin-attachable after 10 times attachment/ detachment cycles and washing. It is observed to conformally adhere to the skin, even during compression and stretching, enabling the sensor to detect the chemicals in human sweat. To calibrate our sensor, the electrochemical performance of the sensor was compared to that of commercial sensors. A glucose assay kit from Aldrich was purchased as a commercial glucose sensor. A commercial pH sensor (Sanxin, MP512) was also purchased. The graph on the left-hand side of Figure 6b measures the concentration of glucose in human sweat accumulated during exercise before (solid) and after (open) meals using both the commercial assay kit and our sensor. In addition, the graph on the right-hand side of Figure 6b shows the pH values of human sweat obtained from two different actions, after exercise (solid) and during the sauna (open) measured by both the commercial sensor and our pH sensor. The measured results appear to show different values in the human sweat by the matrix effect because the glucose and pH sensors were calibrated with PBS and standard solution, respectively. However, because the R2 values are close to 1, our sensor is qualified to detect glucose and pH in human sweat through compensation. Figure 6c,d shows the performance of the sensor attached to the skin. Because the performance of pH sensor can be affected by a shift in the reference potential of the 13737



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trend of glucose level in the sweat measured by our sensor is reasonable. Figure 6d clearly shows that such behavior is reproducible with additional meal ingestion. Unlike glucose, pH did not change significantly with meal ingestion, which is attributed to the maintenance of homeostasis. These results demonstrate that our sensor attached to the skin can be used for monitoring the glucose and pH in sweat for long duration even under mechanical deformation and can thus be applied to healthcare wearable devices.

AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. ORCID

Jeong Sook Ha: 0000-0003-1358-3295 Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was supported by the National Research Foundation of Korea (NRF) grant funded by the Korea government (MEST) (Grant No. NRF-2016R1A2A1A05004935). The authors acknowledge the KU-KIST graduate school program of Korea University.

CONCLUSIONS We have successfully demonstrated a skin-attachable, stretchable electrochemical sensor for the detection of glucose and pH in sweat. Stretchable electrodes of AuNS/CNT could be fabricated via use of a percolation network of 2D AuNS and 1D CNT prepared by vacuum filtration and the LbL method, respectively. Hydrothermally synthesized CoWO4/CNT nanocomposite and PANI/CNT were used for the detection of glucose and pH, respectively. The fabricated sensor exhibited high sensitivities of 10.89 μA mM−1 cm−2 and 71.44 mV pH−1 for glucose and pH, respectively. In addition, very selective sensing of glucose and pH was observed, without any interference from other chemical components and ions existing in human sweat. Long-term stability of the sensor for 10 days was guaranteed, without any notable deterioration of the sensitivity of either glucose or pH. Furthermore, the sensor exhibited mechanical stability upon 1000 repetitive stretching/ releasing cycles under 30% strain. Encapsulation of the sensor with sticky silbione enabled skin attachment and stable performance under multiple attachment/detachment cycles. The change of glucose and pH concentration in sweat upon meal ingestion and exercise could be monitored for long duration while the sensor was attached onto skin, which followed the conventionally observed behavior in blood under similar conditions. This work suggests the high potential of application of our skin-attachable and stretchable sensor for continuous monitoring of the important physiological components in sweat for personal healthcare.

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REFERENCES

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ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.8b03342. SEM, AFM, and optical images of AuNS (Figure S1); stretchability characteristics of AuNS electrode (Figure S2); UV/vis absorbance and SEM image of LbL-CNT film (Figure S3); XPS scans of CoWO4/CNT (Figure S4); SEM image and CV curves of CoWO4/CNT electrode (Figure S5); TEM, SEM, and EDS images of AgNW (Figure S6); SEM and EDS images of Ag/AgCl (Figure S7); OCP changes of Ag/AgCl electrode (Figure S8); stretchability characteristics of each electrode (Figure S9); CV curves of CoWO4/CNT electrode (Figure S10); temperature dependence of sensor (Figure S11); reversible test of pH sensor (Figure S12); zoomedin optical images of sensor (Figure S13); photographic images of sensor (Figure S14); mechanical deformation test of sensor (Figure S15); ζ-potential of nanomaterials (Table S1); comparison of sensor (Tables S2 and S3) (PDF) 13738


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Skin-Attachable, Stretchable Electrochemical Sweat Sensor for

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