Crack-Enhanced Microfluidic Stretchable E-Skin Sensor - ACS

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Cite This: ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Crack-Enhanced Microfluidic Stretchable E‑Skin Sensor Dong Hae Ho,†,∇ Ryungeun Song,‡,∇ Qijun Sun,∥ Won-Hyeong Park,⊥ So Young Kim,# Changhyun Pang,†,§ Do Hwan Kim,# Sang-Youn Kim,⊥ Jinkee Lee,*,‡ and Jeong Ho Cho*,†,§ †

SKKU Advanced Institute of Nanotechnology (SAINT), ‡School of Mechanical Engineering, and §School of Chemical Engineering, Sungkyunkwan University, Suwon 440-746, Republic of Korea ∥ Beijing Institute of Nanoenergy and Nanosystems, Chinese Academy of Sciences, Beijing 100083, P. R. China ⊥ Interaction Lab, ATRC, Korea University of Technology and Education, Cheonan 31253, Republic of Korea # Department of Organic Materials and Fiber Engineering, Soongsil University, Seoul 156-743, Republic of Korea S Supporting Information *

ABSTRACT: We reported the development of a transparent stretchable crack-enhanced microfluidic capacitive sensor array for use in E-skin applications. The microfluidic sensor was fabricated through a simple lamination process involving two silver nanowire (AgNW)-embedded rubbery microfluidic channels arranged in a crisscross fashion. The sensing performance was optimized by testing a variety of sensing liquids injected into the channels. External mechanical stimuli applied to the sensor induced the liquid to penetrate the deformed microcracks on the rubber channel surface. The increased interfacial contact area between the liquid and the nanowire electrodes increased the capacitance of the sensor. The device sensitivity was strongly related to both the initial fluid interface between the liquid and crack wall and the change in the contact length of the liquid and crack wall, which were simulated using the finite element method. The microfluidic sensor was shown to detect a wide range of pressures, 0.1−140 kPa. Ordinary human motions, including substantial as well as slight muscle movements, could be successively detected, and 2D color mappings of simultaneous external load sensing were collected. Our simple method of fabricating the microfluidic channels and the application of these channels to stretchable e-skin sensors offers an excellent sensing platform that is highly compatible with emerging medical and electronic applications. KEYWORDS: electronic skin, sensor, microfluidic, crack, pressure

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INTRODUCTION

however, the high cost, lack of optical transparency, and toxicity of the liquid metal severely limited the potential applications of E-skin platform. The second microfluidic pressure sensor involved a nonconductive liquid medium sandwiched between two plastic films bearing micropatterned electrodes that formed the sensing element.36 Variations in the liquid impedance induced by the deformations of the sensor could be detected with considerable sensitivity and a fast response time; however, the device fabrication process remained complicated, as they required four complex steps, including the vacuum deposition of electrodes and the construction of a spacing layer via photolithography, bonding, and microfluidic loading. The plastic film and indium-tin-oxide electrodes unfortunately lacked flexibility, precluding the incorporation of this sensor platform into wearable devices that require conformal contact with the skin. There remains an unmet need for a simple and low-cost fabrication process capable of producing transparent and stretchable microfluidic sensors for use in E-skin

Human skin has attracted attention in both industry and academia due to its excellent sensing properties and mechanical durability over long periods of time.1−3 The development of high-performance imitation skin technologies presents one of the biggest challenges and the greatest unmet need in the field of soft electronics. Significant efforts have been applied toward realizing artificial electronic skins (E-skins) with various sensory capabilities.4−23 Mechanical signal sensing, including the detection of pressures or strains, is a primary requirement of E-skin sensors for use in haptic interactions with electronic devices, human action feedback, and health monitoring.24−28 The recent introduction of microfluidics into the field of mechanical sensing has expanded the scope of possible capabilities due to their simple, scalable, and eco-friendly fabrication, their excellent adaptability to complex target shapes, and their facile integration with other biological sensors.29−32 Two microfluidic pressure sensor designs were recently developed. One involved a liquid metal (i.e., a eutectic gallium−indium alloy) embedded in an elastomer to form an active sensing component.33−35 This sensor device exhibited excellent accuracy and reliability for measuring large strains; © XXXX American Chemical Society

Received: October 24, 2017 Accepted: December 5, 2017 Published: December 5, 2017 A

DOI: 10.1021/acsami.7b15999 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces

Figure 1. (a) Schematic diagram showing the procedure used to fabricate the crack-enhanced microfluidic stretchable E-skin sensor matrix. (b) OM images of the PDMS channel surfaces before and after peel-off from the SU-8 pattern surface. (c) SEM images of the microcracks on the PDMS channel surface. (d) Photographic images of the sensing liquid injection (left) and OM image of the sensing liquid propagation into the channel (right). (e) Schematic illustration of the deformation of the microcracks and liquid penetration into the cracks. (f) 4 × 4 Stretchable sensor matrix attached to a metal rod 1 cm in diameter. (g) Stretchable sensor attached to various parts of a human subject (cheek, hand, neck, and knee).

human skin to demonstrate sensing of substantial as well as slight muscles movements. The sensor array was shown to map the distribution of external loads applied across different shapes. Our stretchable crack-enhanced microfluidic sensor array offers an excellent sensing interface that is highly compatible with emerging medical and electronic applications.

applications. Silver nanowires (AgNWs) are good transparent and stretchable conductor materials for use in microfluidic sensors: in comparison to vacuum-deposited metals, their conductivity and transparency are high and their solution manufacturing processes are simple.37−40 In addition, AgNW films may be readily patterned using selective transfer based on interfacial adhesion differences and they are easily embedded in elastomeric matrices to provide the desired mechanical robustness.41−43 Here, we demonstrated the fabrication of a transparent and stretchable capacitive sensor using a microfluidic system enhanced by introducing microcracks into the walls of a AgNW-embedded rubber channel for use in E-skin applications. The sensor array consisted of two laminated poly(dimethylsiloxane) (PDMS) films bearing AgNW-embedded microfluidic channels. The two microfluidic channels were arranged in a crisscross manner to form a crossbar-type capacitive sensor, in which an air gap was sandwiched between the two embedded AgNW electrodes. Various sensing liquids were injected into the channels to optimize the sensing performance. External mechanical stimuli (pressures or strains) were applied to the sensor, deforming the microcracks on the PDMS channel surfaces and permitting penetration of the sensing liquid into the cracks. The increased effective interfacial area between the sensing liquid and the AgNW electrodes changed the capacitance of the sensor. The microfluidic sensor sensitivity depended strongly on the wettability of the sensing liquid on the PDMS surface. The initial fluid interface between the liquid and crack wall and the pressure-induced variations in the contact length of the liquid and crack wall were simulated using COMSOL Multiphysics (COMSOL) based on the finite element method. The simulation results agreed well with the experimental results. The stretchable microfluidic sensor was brought into conformal contact with different areas of the

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RESULTS AND DISCUSSION Figure 1a shows a schematic diagram of the procedure used to fabricate the crack-enhanced microfluidic stretchable E-skin pressure sensor array (a 4 × 4 sensor array was fabricated in this work, and a simple 2 × 2 sensors array was used for demonstration purposes). First, the negative photoresist (SU8) pattern used to fabricate the microfluidic channels was applied to a SiO2/Si wafer using conventional photolithography. The AgNWs were then spray-coated uniformly onto the wafer, which was subsequently used as an electrode in the microfluidic capacitive sensors. The sheet resistance of the AgNWs film was around 198 Ω/sq. Th e poly(dimethylsiloxane) (PDMS) liquid precursor was drop cast onto the Si wafer supporting the SU-8 patterns and then peeled off slightly after thermally curing the PDMS at 100 °C for 20 min. Interestingly, only the AgNWs deposited onto the SU-8 line patterns were transferred and embedded within the surfaces of the PDMS microfluidic channels. The selective transfer of AgNWs occurred because the adhesion between the AgNWs and the hydrophobic SU-8 was much poorer than between the AgNWs and the hydrophilic plasma-treated SiO2. The surfaces of the PDMS regions that were peeled away from the AgNW-coated SiO2 substrate were smooth and flat (Figure S1), but microcracks were generated uniformly across the AgNW-embedded PDMS regions during the peel-off process (compare the left and right panels in Figure 1b). The B

DOI: 10.1021/acsami.7b15999 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces

Figure 2. (a) Capacitance of the crack-enhanced microfluidic pressure sensors prepared with five sensing liquids as a function of the applied pressure over the range 0−140 kPa. (b) Normalized capacitance vs pressure plot obtained from all sensors. Solid lines represent the simulated values. (c) Normalized capacitance of the sensors prepared with EG or water as a function of the stretching strain, from 0 to 9%. (d) Schematic illustration of the channel and crack modeled in the COMSOL simulation. (e) COMSOL simulation results obtained under the initial conditions, modeling the interface in the crack after filling the channel in each liquid case. (f) Contact angles of the liquids on the PDMS surface. (g) COMSOL simulation results obtained for the water and PC cases, under several pressure conditions. (h) Simulated Δd/d0 values for each liquid as a function of the applied pressure. (i) Plot of the wettability parameters (k) vs contact angle. (j) Plot of the simulated normalized capacitance vs the pressure.

AgNW electrode changed the device capacitance. The prepared crack-enhanced microfluidic capacitive sensor array was conformally contacted with a steel bar of diameter 1 cm (Figure 1f) and was readily attached to different regions of the human skin (Figure 1g). First, we characterized the liquid enhancement effect using a unit-sensing cell based on microfluidic channels filled with five different sensing liquids. A poly(tetrafluoroethylene) (PTFE) rod was used to apply an external pressure onto the sensor cell. Real-time variations in the capacitance were monitored using an LCR meter. Figure 2a plots the capacitances of the microfluidic sensors as a function of the applied pressure. Injection of the sensing liquids significantly increased the device capacitance due to the aligned dipoles of the liquid molecules at the liquid− electrode interface. The capacitance values measured in the absence of an applied pressure depended on the dielectric constant of the sensing liquid. The terpineol, EG, DMSO, PC, and water liquids exhibited room-temperature dielectric constants of 2.8,49 37,50 47,50 64,51 and 88,50 respectively. The capacitance values were linearly proportional to the dielectric constant, as shown in Figure S3. As the applied pressure was increased from 0 to 140 kPa, the capacitances of the sensors increased and negligible hysteresis was observed upon release of the pressure. Figure 2b shows the normalized

microcracks formed on the AgNW-embedded PDMS surface may have originated from the large difference between the moduli of the metallic AgNWs (>10 GPa)44 and the rubbery PDMS (∼1 MPa).45,46 Another PDMS mold with the same structure was fabricated and then laminated in a crisscross fashion onto the first PDMS mold to form microfluidic channels. A small amount of the PDMS precursor liquid was applied on top of the cured PDMS relief prior to lamination to prevent leakage of the sensing liquid. Figure 1c shows scanning electron microscopy (SEM) images of the AgNW-embedded PDMS channels, indicating the uniform formation of microcracks on the channel surface. Five different sensing liquids (water, dimethyl sulfoxide (DMSO), ethylene glycol (EG), propylene carbonate (PC), and terpineol) were injected into the channels using a syringe connected to a liquid reservoir (Figure 1d). Liquids that did not induce swelling or other damage to the PDMS were selected as sensing liquids47,48 (Table S1). Figures 1e and S2 illustrate the proposed sensing mechanism of the microfluidic capacitive sensor enabled by microcracks formed on the channel surfaces. Under the application of pressure onto the cell, the injected sensing liquids penetrated the microcracks upon deformation of the channel (Figure 1e). The pressure-induced variations in the interfacial contact area between the sensing liquid and the C

DOI: 10.1021/acsami.7b15999 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces capacitances (ΔC/C0) of the sensors, where ΔC is the capacitance change under an applied pressure (C − C0) and C0 is the initial capacitance. The sensitivity of each sensor was evaluated from the ratio between (ΔC/C0) and the applied pressure (P). The air-gap sensor showed the lowest sensitivity, 0.002 kPa−1. Injection of the sensing liquids into the channel dramatically enhanced the sensitivity. The maximum sensitivities of the devices prepared with PC, DMSO, EG, or water were 0.009, 0.011, 0.012, or 0.021 kPa−1, respectively. By contrast, terpineol injection did not improve the sensitivity of the sensor. Note that the water sensor exhibited a 10-fold increase in sensitivity over that of the air-gap sensor and the minimum detectable pressure was as low as 0.1 kPa (Figure S4). Unlike the air-gap device, the capacitance of the microfluidic sensor resulted from the alignment of the liquid molecule dipoles and thus this value was not affected by the variation of the distance between the two electrodes induced by the structural deformation of the channel (as discussed below). The variations in the device capacitance were monitored as the sensor was stretched from 0 to 9%, as shown in Figure 2c. For both the EG and water sensors, ΔC/C0 increased with the stretching strain. The application of an external load, such as a pressure or a stretching strain, to the sensing cell deformed the microfluidic layer between the two electrodes, and the sensing liquid in the channel penetrated the microcracks. As a result, the interfacial contact area between the sensing liquid and the AgNW electrode increased, producing detectable variations in the device capacitance. Optical microscopy images of the microchannel (inset of Figure 2c) revealed that the microfluidic layer deformed and the microcracks enlarged as the applied stretching strain increased, in synchrony with the strain-induced capacitance increase displayed by the sensors. The results of durability tests performed on the sensors prepared with EG are shown in Figure S5. A pressure of 140 kPa was applied and released repeatedly to the sensing membrane. The capacitance remained stable over 200 cycles, which was attributed to the excellent mechanical properties of the AgNW network embedded in the rubber PDMS. The effects of the crack deformation and liquid properties on the PDMS microfluidic channel upon application of pressure were examined by simulating the movement of the air−liquid interface within a simplified device model. Figure 2d show a schematic diagram of the channel and crack used in the simulation. The computational load was reduced and the problem simplified by considering a two-dimensional symmetric geometry with a single crack. The microcracks on the PDMS channel surface were short in width and sufficiently long that the longitudinal direction was assumed to be infinite in length. Under these circumstances, the problem could be modeled as a 2D symmetric system. The position of the interface within the channel was simulated according to the following two steps. The first step involved finding the initial fluid interface once the sensing liquid had filled the PDMS channel and crack. As the channel began to fill up with liquid from the left inlet, air escaped via the right outlet. If the liquid and PDMS were highly wettable, the liquid fully filled the crack, whereas the liquids with a lower wettability toward PMDS did not fully penetrate the crack, and some air remained trapped. Figure 2e shows the position of the interface in the crack after channel filling for all liquid cases. The blue and red areas represent liquid and air, respectively. The smaller the contact angle was (see Figure 2f), the more liquid could penetrate the crack. For a large contact

angle (as in the water case), only a small amount of water penetrated the crack, whereas for a low contact angle (as in the terpineol case), the crack filled completely. Therefore, the crack was expected to fill completely in the presence of liquids with low contact angles. These results were used as the initial conditions in the second step. In the second step, we calculated the contact area between the liquid and the crack wall as the pressure applied to the PDMS chip was increased. In the absence of pressure, the interface remained unchanged at the initial contact length (d0). The application of a load to the top of the PDMS chip deformed the channel and, ultimately, the crack. The air and liquid were assumed to be incompressible, and the interface in the crack moved downward to conserve the volume of trapped air as the PDMS chip was deformed by pressure. Figures 2g and S6 show the simulation results obtained in all liquid cases under several pressure conditions. The effects of the liquid properties on the device capacitance level were characterized and compared by normalizing the contact length at a specific pressure (d). Figure 2h presents the normalized contact length (Δd/d0) as a function of the pressure for all liquids. All liquids exhibited a nearly linear relationship between the pressure and the Δd/d0. Compared with the initial contact length, the contact length of water increased steeply, whereas the contact length of terpineol increased slightly due to a small elongation of the crack wall. The slope of the Δd/d0 was used to calculate a new parameter of the “wettability parameter (k)”, which was equal to 1 for a fully filling liquid such as terpineol. This parameter increased as the slope of the graph increased, as shown in Figure 2i. The k was used as a correction factor to represent the differences in the sensitivities of the devices. On the basis of the experimental results (see Figure 2b), it was assumed that the normalized capacitance of terpineol had the same value with that of air, which was inversely proportional to the distance between two electrodes. The capacitance of terpineol was used to define a standard value because its wettability parameter was almost 1. The capacitances of the other liquids were calculated according to Ct d P ⎡ ⎤ C C = C0 d = C0⎣(k − 1) 140 + 1⎦ C t , w h e r e Ct ,0 0 t ,0 the subscript t was used to denote the value for terpineol. The effects of the k for each liquid were confirmed by calculating the normalized capacitance value using this equation and comparing these values to the experimental results. Figure 2j presents the normalized capacitances calculated from the wettability parameters. The calculated values agreed well with the experimental results for all liquids (see the solid lines in Figure 2b). This approach was also applicable to the devices placed under a stretching strain (Figure S7). These results indicated that the variation of the contact area between the liquid and the crack wall dominated the sensor sensitivity obtained from each liquid. Our simplified theoretical model predicted the normalized capacitance values and the contact angles of the liquid/PDMS, providing an appropriate approach to predicting the microfluidic sensor sensitivity. After characterizing the sensor performance and analyzing the sensing mechanisms, we demonstrated the sensing abilities of the crack-enhanced microfluidic capacitive sensors on human skin subjected to ordinary human motions, including substantial as well as slight muscle movements. To monitor substantial actions, the sensor was attached to the wrist and joint parts of the arm, which require a high degree of stretchability during movement. To sense slight muscles

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DOI: 10.1021/acsami.7b15999 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces

Figure 3. Variations in the capacitance of the crack-enhanced microfluidic sensor attached to various parts of the human body during the performance of simple motions. (a) Top of the wrist, repetitive bending and releasing motion. (b) Arm joint, forearm bending motion through four different bending angles (30, 60, 90, and 100°). (c) Knee, bending legs through two different angles (45 and 90°). (d) Neck, stretching motion involving pointing the chin upward. (e) Neck, deep inhalation and exhalation motions. (f) Cheek, opening and closing the mouth.

motions, the sensor must detect sufficiently small detachable strains. EG was selected as the channel liquid, as it had a relatively low vapor pressure and good sensitivity compared to that of the other liquids, as discussed above. Optimal packaging considerations require that the liquid evaporation be minimized. The microfluidic sensor was first attached to the top of the wrist between the opisthenar and the arm. As shown in Figure 3a, the capacitance increased from 2.9 to 3.8 pF upon bending of the wrist supporting the sensor, as the EG penetrated the cracks, which increased the effective contact area. After release, the capacitance recovered its initial value (∼2.9 pF), demonstrating the reliable sensing properties of our microfluidic sensor. The microfluidic sensor performance properties were tested by mounting them on the most bendable parts of human body, that is, the joint parts of the arms and legs. As shown in Figure 3b, the arm joint was bent from a relaxed state to different angles (30, 60, 90, and 100°) and then returned to the relaxed state, such that each state was held for 6 s. The capacitance of the sensor was measured to be 2.92 pF in the relaxed state, and it increased to 3.1, 3.6, 4.0, and 4.4 pF at four bending states with different angles (30, 60, 90, and 100°), respectively. Similarly, the sensor attached to the knee was used to characterize the extension state of the legs (Figure 3c). The extension action involved release from a bending state. The

capacitance was measured to be 2.9 pF in the initial state, with an increase to 3.5 and 3.9 pF in the two bending states characterized by 45 and 90° angles, respectively. In addition to demonstrating our microfluidic sensor performance during substantial actions at the wrist and joint parts of the arms and legs, slight muscle movements, such as those detectable on the neck and face, were also captured using the microfluidic sensor. Muscle movements on the neck and face involving slight deformations cannot be detected without a highly sensitive sensor. The microfluidic sensor was attached to the neck and used to capture signals induced during neck stretching or a deep inhalation. Figure 3d shows the capacitance signals obtained during head raising and lowering motions, which represented the stretching and release of the microfluidic sensor. The inhalation and exhalation actions were also characterized. During a deep inhalation, the collarbone muscles contracted, which stretched the microfluidic sensor attached to the neck. The relevant capacitance increased to 3.04 pF, as shown in Figure 3e. After exhalation, the muscles relaxed, and the attached sensor recovered its initial state, exhibiting a stable capacitance of around 2.9 pF. Finally, the microfluidic sensor was attached to the cheek to capture facial muscle movements during expansion of the oral cavity (Figure 3f). Opening the mouth increased the capacitance of the microfluidic sensor from 2.93 to 3.00 pF, and closing the mouth decreased the E

DOI: 10.1021/acsami.7b15999 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces

Figure 4. Two-dimensional mapping of the pressure distribution across the sensor matrix. The gray and red bars represent the output signal distribution on the sensor arrays prepared using air and EG.

sensing, and the other realized sensing with an injected EG, as shown in Figure 4. A plastic rod was positioned diagonally on the sensory array. A load of 14 kPa was applied perpendicularly to the top of the plastic rod using a pressure gauge. Before the sensing liquid was injected, as shown in the gray graph in Figure 4, the ΔC/C0 value of the pressed cells increased to 0.005. The small variation in ΔC/C0 in the pressed sensing cells displayed certain noise peaks in the unpressed pixels. This result suggested that the base capacitance signal of the sensor with an air-gap interlayer was so small that it could be easily disturbed by the environment, generating noise signals. After liquid injection, the ΔC/C0 value increased to 0.027 at a maximum (red graph in Figure 4), over 5 times the value obtained from the air-gap capacitive sensor. Notably, the pressed pixels in the liquid injection device displayed ΔC/C0 values that were much higher than the noise signals obtained from the unpressed pixels, providing a high signal-to-noise ratio. Next, a rectangular plastic block was placed atop the sensory array with an applied pressure of 17 kPa. Prior to liquid injection, the ΔC/C0 value was 0.01 and noise signals were observed. After injecting the sensing liquid, the ΔC/C0 value increased by a factor of 4−0.04, yielding a high noise-to-signal ratio. Similar enhanced sensitivity and noise-to-signal ratio results were also observed in the case of the ring-shaped PTFE load, in which the ΔC/C0 value increased by more than a factor of 4. These results demonstrated that sensing liquid injection

capacitance to the base level (2.93 pF). The small capacitance variation of 0.07 pF, observed during the expansion of the oral cavity, was attributed to the very slight movements of the facial muscles. This test demonstrated that our microfluidic capacitive sensor could detect both substantial actions and slight muscle movements, indicating a significant potential for use in sensing both large-scale and small-scale motions. External stress distributions were monitored by integrating the microfluidic capacitive sensors into a sensory array. Two PDMS films bearing four microchannels were laminated in a crisscross fashion to fabricate a crossbar-type capacitive sensor array (4 × 4 microfluidic sensors), in which AgNWs were embedded in the microfluidic channels to form parallel working electrodes and the sandwiched air gap formed a capacitive layer. The sensing liquid was injected into the sensory array to improve the sensing performance. The sensing ability of the microfluidic sensor array was tested by attaching the termini of the AgNWs in each microchannel to a copper wire connected to the measurement system. After making the connection, plastic objects in three different shapes (rods, blocks, and rings) were positioned on the sensor array and subjected to an applied pressure using a pressure gauge. The stress distributions of the target PTFE objects on two sensor configurations were characterized to visualize the effects of the enhanced sensitivity after injecting the sensing liquid. One sensor configuration utilized an air gap for capacitive F

DOI: 10.1021/acsami.7b15999 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces

Figure 5. Remote-sensing system of the microfluidic sensor. (a) Diagram of the remote-sensing system. (b) Photographic images of all electronic components. (c) Real-time remote-sensing characteristics during repetitive finger pushing.

improve the applicability of the device to real-life E-skin devices in the future.

into the microfluidic capacitive sensor significantly enhanced the device sensing properties. Finally, we demonstrated the remote-sensing application. The remote-sensing system consisted of five components of an LCR meter, an interface device, a microprocessor, a wireless communication module (a bluetooth module), and a mobile device. As shown in the diagram of Figure 5a, our microfluidic sensor was connected to the LCR meter and the signals measured by the LCR meter were sent to the microprocessor after converting to those compatible with the microprocessor through the interface device. The bluetooth module, connected to the microprocessor, conveyed the measured signals to the mobile device. The photographic images of all elements for the remote-sensing system are shown in Figure 5b. Figure 5c displays the real-time pressure sensing characteristics using the remote-sensing system. As the five different pressures from (i) to (v) states were directly engaged onto the sensor by a bare finger, the output signal increased step by step because the sensing liquid penetrated the microcracks upon deformation of the channel. The weak pressure that was similar to that in (i) state was then applied sequentially, exhibiting that the output signals were stable during repetitive finger pushing.

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METHODS

Device Fabrication and Measurements. First, the SU-8 molding patterns were prepared from Microchem SU-8 2150 permanent epoxy negative photoresist. The SU-8 precursor solution was spin-coated on Si wafer, and then the film was soft-baked at 95 °C for 100 min. The soft-baked films were then exposed to UV with the wavelength of 350 nm through the photomask for 40 s. Finally, the film was developed and dried at 130 °C. The thickness of SU-8 was around 100 μm. Onto the Si wafer with the SU-8 patterns, the AgNWs (0.3 mL, 0.5 wt %) dispersion was spray-coated and then the films were dried at 70 °C for 10 min. The SYLGARD 184 silicon elastomer (the precursor and the curing agent were mixed at the weight of 10:1) was poured on the substrate and then cured at 100 °C for 30 min. Finally, the thermally cured PDMS films were peeled off from the mold substrate. Two microchannels were crisscrossed with each other to accomplish a crossbar-type capacitive sensor. Two AgNWembedded PDMS layers with the same geometry were crisscrossed with each other to accomplish a crossbar-type capacitive sensor. After the lamination process, various kinds of sensing liquids were injected into the microfluidic channel by a syringe. The poly(tetrafluoroethylene) (PTFE) rod with 3 mm diameter was utilized to apply the external pressure onto the sensor cell to minimize the fringing effect. The sensing properties of the device were measured by Agilent E4980A high precision LCR meter with mark-10 force gauge M5-05. For the remote-sensing system, a microprocessor (STMicroelectronics STM32F4), a bluetooth module (FirmTech-FB155BC), and an interface device (Silicon labs-CP2102) were used. Simulation. COMSOL MultiPhysics 5.2 (COMSOL Inc.) based on the finite element method was used for all simulations. In step 1, a laminar flow model and conservative level set model52 were coupled together and solved to track the interface between air and liquid. It is assumed that the flow is incompressible and laminar. The left, right, bottom, and other side wall have inlet, outlet, symmetry, and wetted wall boundary condition, respectively. The inlet velocity is 5 mm/s, and pressure of the outlet was 0. The contact angle for the wetted wall, surface tension coefficient, and other properties of liquids were taken from measured and reference data. In step 2, a laminar flow model and phase-field model53 were coupled to simulate the dynamics of the interface. Moreover, a solid mechanics model and moving mesh interface were coupled with laminar flow by a fluid−structure interaction model to consider the effect of deformation of PDMS chip. The solid mechanics model calculated the deformation of the

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CONCLUSIONS In conclusion, we demonstrated the fabrication of a transparent stretchable capacitive pressure sensor using a microfluidic system enhanced with microcracks within AgNW-embedded rubber channels. Five liquids were tested as the sensing liquids. The interfacial contact areas between the sensing liquids and the AgNW electrode increased with the applied pressure as the sensing liquid penetrated the microcracks deformed by the external stimuli. The liquid-dependent sensitivities of the sensors were strongly related to both the initial fluid interface between the liquid and the crack wall and the pressure-induced contact length change. A variety of ordinary human motions were successively detected, and the distributions of the pressure during selective pressing could be represented using 2D color maps obtained from a 4 × 4 sensor matrix. Further optimizations of the channel surface and the sensing liquid, preferably by identifying low vapor pressure liquids, would G

DOI: 10.1021/acsami.7b15999 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

Research Article

ACS Applied Materials & Interfaces PDMS chip, and the moving mesh interface automatically remeshed to apply geometrical change. Because the wall and liquid moved together, squeeze flow occurred and position of the interface was recalculated by the phase-field model. The pressure of the liquid could be calculated from the Navier−Stokes equation of the laminar flow model, and it was used as the boundary load for the solid mechanics model. These sequences were iterated until a steady state was achieved. The top, bottom, and side of the PDMS part were boundary load (0−141.5 kPa), boundary load from laminar flow model, and constraint condition, respectively. The channel had the same boundary condition with step 1 except for the left side that had the outlet boundary condition in step 2. The initial position of the interface was taken from the results of step 1.

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ASSOCIATED CONTENT

* Supporting Information S

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsami.7b15999. Swelling ratios and boiling points of the sensing liquids; OM image of the PDMS regions peeled off from SiO2 surface; schematic diagram of the crack-enhanced microfluidic pressure sensor-sensing mechanism; capacitance (at 0 kPa) of the crack-enhanced microfluidic sensors as a function of the dielectric constants of the sensing liquids; capacitance of the water sensors as a function of the applied pressure from 0 to 0.4 kPa; cyclic durability test of the crack-enhanced microfluidic sensors based on EG; COMSOL simulation results of EG, DMSO, terpineol cases for several pressure conditions; COMSOL simulation results of EG and water cases for several stretching strain conditions (PDF)

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected] (J.L.). *E-mail: [email protected] (J.H.C.). ORCID

Jeong Ho Cho: 0000-0002-1030-9920 Author Contributions ∇

D.H.H. and R.S. contributed equally to this work.

Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS J.H.C. and D.H.K. were supported financially by a grant from the Center for Advanced Soft Electronics (CASE) under the Global Frontier Research Program (2013M3A6A5073177 and 2014M3A6A5060932), Korea. J.L. is appreciative of the support provided by the Basic Science Research Program through the National Research Foundation of Korea (NRF) funded by the Korean Ministry of Science, ICT and Future Planning (2014M3C1B1033982).

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