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Nov 20, 2015 - touch screens, and user health monitoring.1−6 Strain sensors in .... the PDMS microfluidic strain sensor subjected to cyclic deformat...
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Highly Stretchable and Transparent Microfluidic Strain Sensors for Monitoring Human Body Motions Sun Geun Yoon,† Hyung-Jun Koo,‡ and Suk Tai Chang*,† †

School of Chemical Engineering and Materials Science, Chung-Ang University, Seoul 156-756, Republic of Korea Department of Chemical and Biomolecular Engineering, Seoul National University of Science and Technology, Seoul 139-743, Republic of Korea

‡

S Supporting Information *

ABSTRACT: We report a new class of simple microfluidic strain sensors with high stretchability, transparency, sensitivity, and long-term stability with no considerable hysteresis and a fast response to various deformations by combining the merits of microfluidic techniques and ionic liquids. The high optical transparency of the strain sensors was achieved by introducing refractive-index matched ionic liquids into microfluidic networks or channels embedded in an elastomeric matrix. The microfluidic strain sensors offer the outstanding sensor performance under a variety of deformations induced by stretching, bending, pressing, and twisting of the microfluidic strain sensors. The principle of our microfluidic strain sensor is explained by a theoretical model based on the elastic channel deformation. In order to demonstrate its capability of practical usage, the simple-structured microfluidic strain sensors were performed onto a finger, wrist, and arm. The highly stretchable and transparent microfluidic strain sensors were successfully applied as potential platforms for distinctively monitoring a wide range of human body motions in real time. Our novel microfluidic strain sensors show great promise for making future stretchable electronic devices. KEYWORDS: microfluidics, transparent strain sensors, ionic liquids, stretchable devices, human motion monitoring

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INTRODUCTION While conventional strain sensors based on the rigid metals have a long and established foundation, the topic of strain sensors has attracted extensive recent attention for various purposes such as wearable devices, electronic skins, flexible touch screens, and user health monitoring.1−6 Strain sensors in recent research efforts have been fabricated by transferring or casting various conductive nanomaterials onto or within elastomeric substrates such as polydimethylsiloxane (PDMS) or Eco-flex.2,6−8 Strain sensors based on carbon nanomaterials have been widely studied in the forms of buckling,9 mesh,10,11 interlocking,12 foam,13 and film of graphene or carbon nanotube.14,15 In other efforts, random network thin films of inorganic conductive materials including silver nanowire,16−18 zinc oxide nanowire,19,20 and metal nanoparticles2,21 have been utilized to produce the strain sensors. Although these strain sensors formed with solid-state conductive nanomaterials exhibited high sensitivity to deformation from external stresses, most of them required complicated fabrication procedures or additional transferring processes to create the conductive nanomaterial networks on the elastomeric matrix. This can create a strain-dependent electrical response from the sensors. Since their sensing performance depends upon a network above the percolation threshold of the conductive solid nanomaterials, a large hysteresis in the electrical response is exhibited during the stretching and releasing of the strain sensors. This is due to © XXXX American Chemical Society

imperfect recovery of original contact positions between the conducting nanomaterials or a breakdown of the network. In addition, some of the above-mentioned sensors have the inability to capture more than one deformation.9,22,23 Microfluidics has been the focus of intense research in chemical and biological fields due to their many advantages, which include easy fabrication with plastics or elastomers, compact size, rapid analysis, and low energy consumption.24−26 Microfluidic systems have already been successfully applied in a wide range of analytical and chemical processes such as biosensing,27,28 fabrication of various types of colloidal particles,29−31 and chemical synthesis.32,33 Another promising research area, with investigations occurring in recent years, is the use of microfluidics in electronic devices. The soft, stretchable, and deformable nature of the elastomer-based microfluidic devices is very suitable to develop elastic electronics, unlike conventional rigid and stiff solid-state electronic components. Microchannels or microvascular networks in the elastomeric PDMS matrix filled with liquid metals, ionic solution, or hydrogels have been used to fabricate microfluidics-based electronics such as stretchable electroReceived: September 7, 2015 Accepted: November 20, 2015

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DOI: 10.1021/acsami.5b08404 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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ACS Applied Materials & Interfaces des,34−36 reconfigurable antennas,37,38 sensors,39−44 memristors,45 and regenerable photovoltaics.46 Here, we demonstrate highly stretchable transparent strain sensors by combining the merits of microfluidic systems and ionic liquids. An ionic liquid is a salt with a low vapor pressure, thereby existing in a liquid state at room temperature. Ionic liquids possess the properties of low volatility, ionic conductivity, thermal and electrical stability, and the ability to solvate.47 Due to these properties, they are widely used as organic solvents,47,48 electrolytes in electrochemical sensors,49−51 actuators,52,53 ionic conducting energy devices,54,55 and transistors.56 The microfluidic-channel networks embedded in the thin elastomeric PDMS layer are filled with a binary mixture of ionic liquids, with a refractive index that is matched with the PDMS in order to produce the transparent strain sensors. When this transparent microfluidic strain sensor is deformed from external forces, it produces distinct electrical signal changes from diverse motions such as stretching, bending, pressing, and torsion. Unlike the strain sensors made of the solid conducting materials, our microfluidic approach exhibits excellent sensor performance with almost negligible hysteresis. This contrasts with solid-state strain sensors because of their intrinsic mechanical mismatch between conductive solid nanomaterials and an elastomeric matrix.57 Although there have been a couple of papers that reported ionic liquid based strain sensors,42,58−60 optically transparent strain sensors by refractive-index tuning of ionic liquids and its capability for measuring the diverse deformations are unprecedented. Furthermore, our simple-structured microfluidic strain sensor was attached to a finger, wrist, and arm to confirm its capability of detecting various human motions in real time.

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The microfluidic strain sensor was also prepared with an Eco-flex elastomer (ECOFLEX 0010, Smooth-On) to deliver higher strain measurements. A straight circular tube (PEEK tubing, IDEX Health & Science) with an outer diameter of 360 μm was placed on a square Petri dish, and the Eco-flex prepolymer was poured over the tube, followed by curing on a hot plate at 80 °C for 15 min. The resulting Eco-flex matrix had a thickness of 5 mm. After removing the circular tube from the polymerized Eco-flex matrix, the binary mixture of ionic liquids was injected into the circular channel. The Pt wires were brought into both ends of the channel with a separation of 17 mm, and the channel ends were sealed with a drop of the Eco-flex prepolymer. Optical Characterization of the Microfluidic Strain Sensor. Two PDMS slabs containing the serpentine microchannel were sealed to each other with orthogonal orientation of the channels using the airplasma treatment. The resulting microchannel network had a length of 36 mm and a width of 24 mm. Each pure ionic liquid and their binary mixture with 51:49 molar ratio between [BMIM][Ntf2] and [BMIM][Ac] were injected into the microchannel on a hot plate at 80 °C. After sealing the holes for the fluid injection, the optical properties of the transparent microfluidic strain sensor were characterized with a UV−vis−NIR spectrophotometer (V-670, Jasco). Sensor Performance. Two ends of the microfluidic strain sensor were attached to a motorized moving stage (AL1-1515-3S, Micro Motion Technology). The uniform stretching and releasing cycles were applied to the sensor at constant speeds of 2.9−13.58 mm s−1 for characterizing its tensile strain detection performance. For sensing the bending angle, both ends of the strain sensor mounted on two glass plates with a gap distance of 5 mm were pushed at a constant speed of 2.9 mm s−1 by the motorized stage to change the bending angle (θ) from 0° to 150.4°. To demonstrate the capability of our strain sensors for monitoring human body motions in real time, the transparent microfluidic strain sensors were mounted on a forefinger, wrist, and triceps using a double-sided adhesive silicon thin film (VHB 4910, 3M). The electrical signals of the microfluidic strain sensors under the various deformations were measured from the application of a DC voltage of 1 V using Keithley 2400 source/measure units combined with a custom-built LabVIEW program.

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EXPERIMENTAL SECTION

RESULTS AND DISCUSSIONS The fabrication process of the transparent microfluidic strain sensor is illustrated in Figure S1. The PDMS sheets with microfluidic channel patterns were fabricated via conventional soft lithography.24 Two serpentine-patterned PDMS sheets were brought into conformal contact with each other with orthogonal orientation after the air plasma treatment. The resulting double-layered PDMS slab had a grid channel network. The mixture of ionic liquids was injected into the network with a syringe on the hot plate at 80 °C. The slight heating of the PDMS slab facilitated the injection of ionic liquids by decreasing their viscosity.61,62 The void channels in the PDMS sheet were visible because of the different refractive indices in PDMS and air. Filling the channels with the ionic liquid mixture having the same refractive index as PDMS enabled the microfluidic sensor to be perfectly transparent. The cured PDMS had the refractive index of 1.44−1.46.63 To match this value, two ionic liquids were chosen: 1-butyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imide ([BMIM][Ntf2]) and 1butyl-3-methylimidazolium acetate ([BMIM][Ac]). The refractive indices of [BMIM][Ntf2] and [BMIM][Ac] are ∼1.42 and ∼1.49 for the pure component, respectively.64 The refractive index (n12) of a binary mixture is given by the Lorentz−Lorenz equation65,66

Characterization of Ionic Liquids. The binary mixture of ionic liquids was composed of 1-butyl-3-methylimidazolium bis(trifluoromethanesulfonyl)imide ([BMIM][Ntf2], Sigma-Aldrich) and 1-butyl-3-methylimidazolium acetate ([BMIM][Ac], Sigma-Aldrich) with various molar ratios. Measuring the refractive indices of the ionic liquids was performed with a refractometer (DR301-95, Kruss). The conductivities of pure [BMIM][Ntf2] and [BMIM][Ac] are 0.39 and 0.19 S m−1, respectively. Conductivity of the binary mixture of ionic liquids was obtained from a conductivity meter (B-771, Horiba). Fabrication of Microfluidic Strain Sensors. Linear and serpentine types of microfluidic channels were fabricated in a PDMS elastomer (Sylgard 184, Dow Corning) using soft lithography.24 Channel masters were prepared by spin-coating the SU-8 2050 photoresist (MicroChem) on a silicon wafer. UV light (UV CURE 60PH, Lichtzen) was exposed to the photoresist films through the transparency photomasks with the channel printout, and the uncured photoresist was removed by an SU-8 developer solution (MicroChem). The PDMS prepolymer was cast on the masters and cured at 80 °C. Two holes were punched at both ends of the channels using a 16-gauge needle to inject the ionic liquids and connect Pt wire electrodes. The PDMS replicas containing the microchannels were irreversibly sealed to a PDMS slab by an air-plasma treatment (model PDC-32G, Harrick Plasma). The width and height of the prepared microchannels in the resulting PDMS matrix with a thickness of 1.0− 1.5 mm were 400 and 70 μm, respectively. Tubings (Tygon) were inserted into each hole. A drop of PDMS prepolymer was cast at the interface between the tube surface and the PDMS and cured to prevent the leakage of the fluids. The binary mixture of ionic liquids was introduced into the microchannels using a syringe on a hot plate at 80 °C, and the Pt wires were inserted through the tubings to apply an external DC voltage. Since the ionic liquid, [BMIM][Ntf2], is corrosive to skin or eyes, it should be handled carefully.

2 n12 −1 2 n12 +2

B

= ϕ1

n12 − 1 n12 + 2

+ ϕ2

n22 − 1 n22 + 2

(1)

DOI: 10.1021/acsami.5b08404 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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Figure 1. Optical characterization of the microfluidic strain sensor. (a) Refractive index of a binary mixture of [BMIM][Ntf2] and [BMIM][Ac] as a function of molar ratio. The solid line is a fit to the Lorentz−Lorenz equation (eq 1). (b) Absorbance spectra of air-filled and ionic liquid-filled microfluidic networks embedded in the PDMS matrix. Photographs of the PDMS microfluidic networks filled with (c) an air and (d) a binary mixture of ionic liquids with 51:49 of molar ratio between [BMIM][Ntf2] and [BMIM][Ac]. The presence of refractive-index-matched ionic liquids in the microchannels gives rise to better transmittance and allows clear observation of texts on the background paper.

Figure 2. (a) Schematic illustration of microfluidic strain sensor with a single linear microchannel embedded in elastomers. (b) Current response of the PDMS microfluidic strain sensor subjected to cyclic deformation with various tensile strains (ε = 10−25%) at a speed of 4.9 mm s−1. (c) Current response and applied strain variation as a function of time for the microfluidic strain sensor subjected to cyclic deformation with tensile strain up to ε = 25% at a speed of 4.9 mm s−1. (d) Relative electrical resistance change as a function of tensile strain up to ε = 25% for the PDMS microfluidic strain sensor at different stretching/releasing speeds. The dashed line in (d) is theoretical predictions of the resistance change obtained from eq 3.

C

DOI: 10.1021/acsami.5b08404 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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ACS Applied Materials & Interfaces where n1 and n2 are the refractive indices of components 1 and 2 in binary mixtures, respectively. The volume fractions of the components are ϕ1 and ϕ2. These volume fractions were calculated from ϕ1 = x1v1/Σxivi, where x and v are the mole fraction and the molar volume of component i, respectively. Figure 1a shows the theoretical refractive-index values of the binary mixture of [BMIM][Ntf2] and [BMIM][Ac] obtained from eq 1 with experimentally measured values, which are in close agreement with each other. On the basis of the numerical estimation from the Lorentz−Lorenz equation, we found the optimum molar ratio of the binary ionic liquid mixture to be approximately 51 mol % of [BMIM][Ntf2] in order to match its refractive index with that of PDMS. The UV absorbance of the microfluidic PDMS sheets filled with the pure ionic liquid and binary mixture are compared in Figure 1b. It is observed that the absorbance is close to zero when the channel is filled with an optimum mixing ratio of ionic liquid mixture. Figure 1c and d visually demonstrate that the microfluidic channels become invisible after injection of the ionic liquid mixture with a refractive index that matches that of the PDMS. Because the channel dimensions can be easily distorted under the applied external forces to the elastomeric PDMS layer, the conductive ionic liquids confined inside the microchannel are readily deformed. As a result, the electrical resistance across the channel changes, and the transparent microfluidic network acts as a strain sensor. To minimize the channel geometry complexity of our transparent microfluidic system as an effective strain sensor, we chose the simple linear channel instead of the previously mentioned grid network. The straight microfluidic channel with 400 μm width, 70 μm height, and 30 mm length was sealed and injected with the refractiveindex-matched ionic liquid mixtures. Platinum (Pt) wires were connected to each edge of the channel as electrodes for resistance measurement at a constant DC voltage of 1 V, as shown in Figure 2a. The electrical current change across the microfluidic channel upon repetitive stretching (i.e., tensile strain) cycles is shown in Figure 2b. The ionic liquid-filled PDMS linear channel was elongated up to 25% of its original length at a speed of 4.9 mm s−1. Figure 2b shows the current changes through the ionic liquid in the channel under different tensile strains of 10, 15, 20, and 25%. As the tensile strain increased, the current decreased proportionally because the cross-sectional area and the length of the elongated channel become smaller and longer, respectively, corresponding to the higher strain. The magnified current signal at the various strains in Figure 2c and Figure S2 displayed almost bilateral symmetry from the stretching and releasing states with a good response speed. This confirmed the reliability of the strain-dependent current response over multiple cycles. We further investigated the effects of applied strain speed on the performance of our microfluidic strain sensor. The resistance change as a function of strain at different stretching and releasing speeds during the multiple cycles is shown in Figure 2d and Figure S3. While little hysteresis between the stretching and releasing was observed at relatively low strain speeds (2.9 and 4.9 mm s−1), the hysteresis became noticeable as the strain speed increased. For quantitative comparison of the hysteresis magnitude, we defined the degree of hysteresis (DH) as follows DH =

AS − AR × 100% AS

where AS and AR are the area of stretching and releasing state curve, respectively. A DH value close to 0% would indicate less hysteresis of the electrical signal. The transparent microfluidic strain sensors stretched and released at 2.9 and 4.9 mm s−1 strain speeds showed an almost negligible average DH of 3.73% and 1.89%, respectively. Although an increment of hysteresis was observed in the electrical response from our strain sensor at the faster strain speed, it is still relatively less or comparable to that from strain sensors based on random connections between the solid-state nanomaterials (Table S1). Therefore, our microfluidic strain sensor showed outstanding signal recovery in the tensile strain, with no significant hysteresis. This is because the conducting ionic liquids can readily flow through the stretched and relaxed microchannels without creating disconnections. Due to the limited PDMS strain ranges,67 the microfluidic strain sensor fabricated with a highly stretchable Eco-flex material was tested to larger tensile strains with a strain speed of 2.9 mm s−1. This highly stretchable microfluidic sensor was formed by introducing the binary mixture of ionic liquids into a circular channel (360 μm diameter and 17 mm length) in the Eco-flex. Under the repetitive cycles of 200% tensile strain, the resistance increased linearly up to 50% strain with a coefficient of determination of 0.98. It then produced a nonlinear response. After reliably measuring up to approximately 20 times of its original value, the resistance was fully recovered. This demonstrated the outstanding stretchability of our microfluidic strain sensor (Figure 3a). The strain−resistance

Figure 3. (a) Relative electrical resistance change as a function of tensile strain up to ε = 200% for the Eco-flex microfluidic strain sensor cycled three times at a speed of 2.9 mm s−1. The dashed line in (a) is theoretical predictions of the resistance change obtained from eq 3. (b) Gauge factors of the strain sensor as a function of tensile strain.

(2) D

DOI: 10.1021/acsami.5b08404 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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ACS Applied Materials & Interfaces ratio curves in Figure 3a during the stretching and releasing cycles also showed negligible hysteresis, with an average DH of 2.41%, even with such high strain conditions. In addition to the hysteresis characteristics, the performance of strain sensors can be evaluated by a gauge factor (GF). This is the ratio of the relative resistance change and the applied strain (GF = (dR/ R0)/(dL/L0), where R0 and L0 are the resistance and the length of unstretched channel, respectively, and dR and dL are the change of resistance and length before and after the application of tensile strain, respectively.).13,16 The highly stretchable microfluidic strain sensor exhibited GF = 2 for the linear region ( 120%). This discrepancy could be from the irregular distortions of rectangular and circular cross-sectional channels in these high strains. For strain sensors to have practical application with human body motions, they should not be limited to measuring tensile strain. Therefore, we further demonstrated the sensing performance of the transparent microfluidic strain sensors for other types of deformation related to human body movement such as bending, torsion, and pressing. First, we measured the electrical resistance variation from point bending on the microfluidic strain sensor. This could be applied to capture the motion of joints in the human body. Two glass plates with a separation of 5 mm were attached to the bottom of the PDMS slab containing a single linear channel filled with the binary mixture of ionic liquids. The plates were continuously pushed at a speed of 2.9 mm s−1 by a motorized stage to produce point bending of the microfluidic strain sensor, as shown in the inset of Figure 4a. The relative resistance changes of the strain sensor under the multiple bending cycles up to an angle (θ) of 150.4° are shown as a function of bending angle in Figure 4a. The resistance behavior changed noticeably after θ ∼ 50°, likely due to significant deformation of the channel cross-section at the bending position. The electrical resistance reached a value four times higher than that of the relaxed sensor. Remarkably, the resistance recovered during the bending cycles without any

Figure 4. (a) Relative electrical resistance change as a function of bending angle (θ) from 0° to 150.4° for the microfluidic strain sensor cycled three times at a speed of 2.9 mm s−1. (b) Relative resistance change as a function of torsion angle up to one revolution. (c) Relative resistance change as a function of pressing depth (h). The microfluidic strain sensor containing a single linear channel (400 μm width, 70 μm height, and 30 mm length) in the PDMS matrix was used for bending and torsion tests in (a) and (b). The strain sensor with serpentine type of microchannel in the PDMS slab (36 × 24 mm2) was used for the pressing test in (c).

hysteresis. Similar sensor performance was observed when it was repetitively bent with other maximum bending angles (Figure S4). To apply our strain sensor for the torsion sensing, one end of the microfluidic strain sensor with a single straight channel was rotated about the axial axis of the channel, while the other end was fixed. As the strain sensor was twisted, the electrical resistance increased approximately 34% from the value of the E

DOI: 10.1021/acsami.5b08404 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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Figure 5. Monitoring of various human body motions in real time. (a) Electrical resistance change as a function of time for the microfluidic strain sensor with a single linear channel in the PDMS elastomer at different finger bending motions. Electrical resistance change as a function of time for the PDMS microfluidic strain sensors with two linear channels (b) at different wrist motions and (c) at various movements of the arm. Photographs in (a−c) are the corresponding motions of the forefinger, wrist, and arm attached with the strain sensors, which are indicated as numbers in each graph of the electrical responses.

results indicate that our microfluidic strain sensor show promise for measuring surface pressure. Since our transparent microfluidic strain sensors are highly flexible, stretchable, and sensitive with fast response to various deformations and no considerable hysteresis, they have a strong potential for wearable platforms to detect the movements of the human body with large strains and high bending angles. In order to further demonstrate the capability of the fluidic soft sensors for use on the human body, the sensors were installed on a forefinger joint knuckle, wrist, and triceps. The detection of the finger motion in real time was achieved with a single microchannel embedded in the PDMS slab. The resulting

untwisted sensor after one revolution (Figure 4b). The transparent microfluidic strain sensor was also constructed with a serpentine channel for the pressure sensing test. After securing two sides of the microfluidic strain sensor with an area of 36 × 24 mm2, a vertical force was applied to the center of this planar strain sensor, as shown in Figure 4c. As the sensor was pushed down with an increasing pressure force with depth h, the relative resistance gradually increased due to the distortion of the cross-sectional area and length of the channel (Figure 4c). Although further investigations are needed for the quantitative analysis of this pressure sensing performance, the F

DOI: 10.1021/acsami.5b08404 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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ACS Applied Materials & Interfaces electrical resistance from the finger bending increased noticeably, similar to the previous point bending test. The sensor was able to clearly distinguish between the half- and full-finger bending motions, as shown in Figure 5a. In addition, rapid digit movements were also clearly detectable in the repetitive bending/relaxing cycles without delay or drifting of the signal. During step testing of the digit motion, the increased resistance in response to the forefinger bending was maintained with a nearly constant value while holding the bending load, followed by relaxation back to the original resistance. To further demonstrate human motion monitoring in real time, the microfluidic strain sensor containing two linear microchannels was mounted on the skin of the wrist and triceps on a left arm. The high transparency of our microfluidic strain sensors was demonstrated by a tattoo patch clearly visible through the attached sensor, as shown in Figure 5b. The output signal from the sensor on the wrist was able to distinguish the backward/forward bending and pressing modes (Figure 5b). In the backward bending of the wrist, the resistance change was distinctly observed for the full- and half-bending cycles. As the wrist was moved in a step motion after its complete bending, the resistance exhibited a corresponding stepwise reduction. Interestingly, when our sensor on the wrist was slightly bent in the forward direction, its electrical resistance decreased slightly below the nominal value. The microchannel of the sensor used in this wrist motion was placed in a vertical direction above the middle of the PDMS matrix (Figure S5). The reduction of the electrical resistance could be from the axial compression of the microchannel in this slight forward bending motion.69 Under rapid backward and forward bending of the wrist, the changes in electrical resistance were measured with a good response, showing oscillations over the nominal value. When a finger quickly pressed the strain sensor on the wrist, the electrical resistance sharply increased. It recovered well to the original value when the finger was removed. Various motions of human arms were also detectable from the transparent microfluidic strain sensor when attached to the triceps. The electrical resistance was influenced by expansion or contraction of the triceps muscle due to the movement of arms. The electrical signal responses of the strain sensor from the diverse movements of the arm are plotted in Figure 5c. When the arm connected to our strain sensor was bent due to the expansion of the triceps, the increased resistance was measured because of the deformation of the channels in the sensor. Extending the arm restored the sensor’s original resistance value. The microfluidic strain sensor was able to detect the rapid bending cycles and the step motion of the human arm. To further investigate the capability of the sensor, the extended arm was twisted repeatedly. This caused relatively small expansion and contraction of the triceps muscle. The corresponding resistance of the sensor was oscillatory around the nominal value. We further tested the performance of the strain sensor with a combination of two different motions. As the arm was bent and swayed from side to side simultaneously, the resistance was oscillated up and down across the increased resistance value and was rapidly recovered to the original resistance, once the swaying motion was stopped and the arm was extended. On the basis of these experiments, our transparent microfluidic strain sensor demonstrated a strong capability to distinctively measure a wide range of human body motions. In addition to being highly stretchable, bendable, and sensitive, its stability is also critical for practical usage of the strain sensors. We proved that our strain sensors have long-

term stability by monitoring their current response under a continuously applied voltage and measuring the resistance for more than 2 weeks (Figure 6). The electrical responses were

Figure 6. Long-term stability of the microfluidic strain sensor. (a) Typical current response of our strain sensors as a function of time under continuous application of constant DC voltage (1 V) across the channel. (b) Electrical resistance of the strain sensor as a function of time, measured with DC sweep scanning from −1 to +1 V. Our strain sensors showed stable electrical response for more than 2 weeks.

found to be stable without electrochemical reactions as long as the ionic liquids were completely sealed in the microchannels. It is also important to note that the microfluidic strain sensor retained its electrical response even after 1000 stretching/ releasing cycles (Figure 7). These results indicate that our strain sensors can be reusable for multiple occasions.

Figure 7. Electrical resistance of the microfluidic strain sensor. The strain sensor was cyclically stretched and released at a speed of 2.9 mm s−1 up to 27% tensile strain. The inset shows the resistance response of the strain sensor with time in 3 min. G

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CONCLUSION In summary, we developed a new class of simple fluidic strain sensors with high stretchability, flexibility, sensitivity, and longterm stability with a negligible hysteresis and a fast response by combining the merits of microfluidic techniques and ionic liquids. The high transparency of the strain sensors was achieved by introducing refractive-index-matched ionic liquids into the microfluidic network or channels embedded within an elastomeric matrix. The outstanding sensing performance was exhibited under a variety of deformations induced by stretching, bending, pressing, and twisting of the fluidic sensors. The electrical response of the sensors was well predicted by a theoretical model based on the elastic channel deformation. Such highly stretchable and transparent microfluidic strain sensors were successfully applied as potential platforms for distinguishably detecting a wide range of human body motions in real time. Our approach for fabricating high-performance microfluidic strain sensors is simple, low cost, and ergonomic; thus it shows great promise for creating future stretchable devices in such applications as wearable electronics, soft robotics, and artificial skins.

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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.5b08404. Comparison of the degree of hysteresis, schematic illustration of fabrication process, current response, and applied strain variation (PDF)

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This research was supported by Basic Science Research Program through the National Research Foundation of Korea (NRF) funded by the Ministry of Education (2013R1A1A2012177) and the Chung-Ang University Freshman Academic Record Excellent Scholarship Grants in 2015.

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REFERENCES

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