A sandpaper-inspired flexible and stretchable resistive sensor for pressure and strain measurement

A sandpaper-inspired flexible and stretchable resistive sensor for pressure and strain measurement

Organic Electronics xxx (xxxx) xxx–xxx Contents lists available at ScienceDirect Organic Electronics journal homepage: www.elsevier.com/locate/orgel...

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Organic Electronics xxx (xxxx) xxx–xxx

Contents lists available at ScienceDirect

Organic Electronics journal homepage: www.elsevier.com/locate/orgel

A sandpaper-inspired flexible and stretchable resistive sensor for pressure and strain measurement Ashok Chhetry, Partha Sarati Das, Hyosang Yoon, Jae Yeong Park∗ Department of Electronic Engineering, Kwangwoon University, 447-1 Wolgye-dong, Nowon-gu, Seoul, 01897, Republic of Korea



Keywords: Sandpaper Shape factor Microstructures Contact spots

We report very small shape-factored microstructures developed via a simple and cost-effective approach, enabling a high degree of sensitivity in a low-pressure regime (< 2.67 kPa). The surface intertexture on the counter electrode and irregular microstructures with a high surface area developed on the base electrode help reduce the shape factor, allowing the device to deform more easily under pressure. Moreover, the irregular patterns with higher unloaded surface area strengthen the tunneling current sufficiently at low pressure. Furthermore, the fabricated features enable the device to perform as a flexible and stretchable sensing mechanism; the outstanding performance was achieved through a novel and feasible fabrication from a low-surface-energy template without surfactant coating. An ultra-low hysteresis of 3.17%, a high sensitivity of 0.3954 kPa−1, a fast response time of 0.49 s and stability over 6000 cycles were achieved. Finally, the sensing capability was examined by gentle finger tapping and arbitrary movement of the sensor placed on the forefinger. The current platform can be a key component for diverse applications such as muscle movement, speech detection, and health monitoring systems.

1. Introduction The design and development of wearable sensor systems for physiological measurements that interface with human skin and/or body organs are of importance for disease diagnosis, health status monitoring, electronic skin (e-skin), and human-machine interfaces. Wearable sensors with unique properties such as low cost, ultra-lightweight, highly flexible, and stretchable, which can be conformally attached to the human skin have tremendous opportunities for use in personal healthcare and physiological signal monitoring. In particular, many sensing devices based on various structures, transduction principles, and fabrication processes have been introduced to enable ultrathin, low-modulus, flexible and stretchable devices. Depending on the active material and device structure, each of the sensor systems has its own sensing mechanism and characteristics. Recently, great achievements have been made in the development of flexible and stretchable electronics based on resistive [1–6], capacitive [1,7,8], piezoelectric [9,10] and optical [11] sensing technologies. Among these, resistive sensors have been more dominantly used because of their attractive advantages such as simple acquisition, low power consumption, and low susceptibility to electromagnetic interference [6,12]. These sensors hold substantial promise for the realization of human-machine interface [5,13,14], health monitoring [15,16], e-skin [1,15] due to their inherent flexibility and stretchability, simplicity in

fabrication process and structure. In addition, the advent of new piezoresistive materials such as carbon nanotubes (CNTs), graphene, poly (3,4-ethylenedioxythiophene)-poly(styrene-sulfonate) (PEDOT:PSS), and metallic nanoparticles, has inspired exploration of their use in pressure sensing mechanisms, because such flexible and stretchable materials can be embedded in structural materials with three-dimensional deformability [4,14,17]. The most common materials utilized for novel sensing are carbon-based composites, graphene foams, or polyurethane (PU)-based composite foams [11]. Because of their superior electrical properties, they often show excellent linear pressure/strain response. However, despite the success of these approaches, many challenges remain to meet the requirements of more accurate, reliable, and efficient pressure sensors for flexible electronics applications. For example, higher sensitivity in low-pressure regime, challenge in packaging and stability issues potentially limit their broad applications for health-status monitoring and human-machine interface [13]. Typically, polymer films decorated with carbon black particles and conductive fillers such as silver nanowires (AgNWs), and reduced graphene oxide (rGO) are used as a sensing material for piezoresistive sensors [11,17,18]. However, these techniques suffer from poor sensitivity and repeatability, limiting their application in e-skins [19]. Microstructured polymer films covered by metallic materials are also employed in the fabrication of strain-gauge and piezoresistive sensors, exploiting the variation in resistance induced by changes in the

Corresponding author. E-mail address: [email protected] (J.Y. Park).

https://doi.org/10.1016/j.orgel.2018.06.034 Received 15 December 2017; Received in revised form 14 May 2018; Accepted 23 June 2018 1566-1199/ © 2018 Elsevier B.V. All rights reserved.

Please cite this article as: Chhetry, A., Organic Electronics (2018), https://doi.org/10.1016/j.orgel.2018.06.034

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conductive fabric was prepared by absorption of Ag precursor and subsequent reduction of the precursor into AgNPs. All the chemicals/ reagents were purchased from Sigma-Aldrich. First, poly(styrene butadiene styrene) (SBS) polymer suspension was prepared by dissolving 1.6 g of SBS pellets (15 wt%) in a 10 mL mixed solvent of tetrahydrofuran (THF) and dimethylformamide (DMF) (3:1 weight ratio). The suspension was mechanically stirred for 15 min and then ultra-sonicated for another 30 min. A thin nano-layer of SBS polymer was coated on the cotton patch (2 cm × 2 cm) by dipping it into the SBS suspension for 15 s. After 5 min of evaporation in the air, the SBScoated fabric was immersed into silver trifluoroacetate (AgCF3COO) solution in ethanol (15 wt%) for 30 min to absorb Ag ions into the SBS polymer. Ag precursor was reduced into AgNPs by a solution of 50% hydrazine hydrate (N2H4·4H2O) in a mixture of ethanol and deionized (DI) water (1:1 vol ratio). After 5 min, the residual reduction agent was rinsed out several times with DI water. Sufficiently low surface resistance was achieved after repeating absorption of Ag precursor in the SBS polymer and reduction into AgNPs cycle eight times. The as-prepared conductive fabric has excellent electrical conductivity and is resistant to water washing.

geometry of the structure [20]. Some metals such as nickel and platinum alloys, and silver nanoparticles (AgNPs) show higher resistivity variation with respect to resistance changes induced by geometrical changes. For these reasons, many flexible sensors exploit metal piezoresistors either with simple geometry or as strain gauges, thus combining both resistivity and geometrical variation. For the piezoresistive transduction mechanism, various structures, such as micro-pyramids, micro hair, and porous materials or fracture on the porous structure are utilized as an active layer for improving the sensitivity of the device. A cost-effective and large-area-compatible technology is an indispensable aspect for the fabrication of flexible pressure-sensitive materials with sufficient sensitivity for wider pressure ranges. One approach to enhancing the sensitivity as well as deformability is to introduce the porous structure in the sensing material, for example, conductive polymer hollow spheres, porous CNT sponge, and metalpolymer hybrid sponge. Recently, Sungmook Jung et al. used reverse micelles solution to introduce pores into a conventional pressure-sensitive rubber [4]. However, the difficulties and complexities in controlling the pore morphology restrict the repeatability and sensing capability. On the other hand, stretchable resistive pressure sensors based on micro-pyramidal polydimethylsiloxane (PDMS) arrays [2,3], and interlocked micro-dome arrays [4,21] are reported in the literature. However, these devices have limitations on responsivity to pressure, and the composite elastomer coating has an effect on temperature variation and inter-filler distance [4]. In addition, fabrication of these structures involves complex fabrication, high cost, and lower reproducibility. In this work, we present a flexible resistive pressure sensor-design approach that employs a microstructured platform to reduce the shape factor (the ratio of an area in compression to an unloaded surface area free to expand) to achieve excellent sensitivity with a simple and costeffective fabrication technique. The sensing mechanism is based on the resistance change caused by variation in contact area between the fabric intertexture and conductive microstructures on PDMS. The assembly altogether incurred minimum contact area with maximum unloaded area free to expand during compression. The contact spots between the fabric intertexture and microstructured PDMS create the minimum area in compression. Thus, a high degree of resistance gradient at sufficiently low-pressure change was achieved and also the process was reversible with almost negligible viscoelastic effect. An ultra-low hysteresis of 3.17%, a higher sensitivity of 0.3954 kPa−1 for the low-pressure region (< 2.67 kPa), a fast response time of 0.49 s, and stability > 6000 cycles were achieved.

2.3. Characterization and measurements In order to examine the surface morphologies of microstructured PDMS and AgNPs on the conductive fabric, field emission scanning electron microscopy (FE-SEM) was used. The surface resistance of the conductive fabric was measured by a four-point meter (RC3175, EDTM). A moving stage (JSV-H1000, Japan Instrumentation System Co., Ltd.) in combination with a force gauge (HF-1, Japan Instrumentation system Co., Ltd.) was used to apply precise load. An acrylic spacer of 2 cm × 2 cm was also used between the sensor and force gauge fixture to apply uniform load on the sensor. The use of spacer offers good insulation between the sensor and fixture, enabling measurement of very low pressure. Resistance measurements were performed using an LCR meter (Hioki, IM 3536) in combination with a computer to record compression distances, forces, and resistances. 3. Results and discussion 3.1. Working mechanism In our sensor design, the AgNPs-reinforced conductive fabric was placed on a Ti/Au-coated microstructured PDMS electrode face-to-face in order to form an intimate contact between the fabric intertexture and conductive features of the microstructured PDMS. A schematic illustration of the assembly and its operation is shown in Fig. 2(a). The bottom layer is microstructured PDMS film with an average pattern size of ∼270 μm. In order to complete the base electrode, the microstructures were covered by a thin layer of Ti/Au. The top counter electrode had the intertextures of AgNPs-coated on cotton fabric. Because the microstructures are random with unequal feature sizes, two kinds of intimate contact spots exist; one intrinsic connection between the conductive fabric and taller features of the base electrodes, and another between the fabric intertexture and shorter features, which are analogous to resistors in a parallel configuration. As shown in the inset of Fig. 2(a), very few contact spots with smaller contact area are achieved during no-pressure condition because of the presence of intertexture on the fabric and Au-coated microstructures on the PDMS. Contact resistances associated with each contact spots are represented by variable resistors (R1 – Rn), best illustrated by Fig. 2(b). Whenever an external pressure is applied, the local stress field widens and the stress intensity increases, which in turn causes an increase in contact area. The degree of increase in contact area depends on the degree of free area to expand and the elastic modulus of the structure. An induced deformation in both of the structures increases the contact area of previously connected contact spots, thereby experiencing a decrease in

2. Materials and methods 2.1. Fabrication of micro-structured base electrode For the generation of microstructures on the PDMS film, readily available silicon carbide abrasive sandpaper with an average grit size of ∼270 μm was used as a template. A suspension containing PDMS base and curing agent (Sylgard® 184, Sigma-Aldrich) in a 10:1 weight ratio was mechanically stirred for 5 min followed by ultra-sonication for 20 min. After degasification for 30 min, the suspension was spin-coated (SCS 6800 spin coater) at 1000 rpm for 30 s to achieve ∼300-μm-thick film. The structure inverse to the sandpaper pattern was obtained by curing the film at 70 °C for 3 h and then peeling off carefully from the template. Finally, the base electrode was grafted on to the PDMS patterns by magnetron sputtering of a thin film of ∼600/1000 Å Ti/Au on the microstructured PDMS. The ohmic contacts were made by attaching a conductive adhesive tape. A schematic illustration of the detailed fabrication sequence is shown in Fig. 1(a). 2.2. Fabrication of the conductive fabric For the purpose of the counter electrode, the AgNP-reinforced 2

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Fig. 1. (a) Schematic illustration of the fabrication sequence. Photograph of the (b) Ti/Au (600/1000 Å) sputtered on the microstructured PDMS, (c) stretchability of the base electrode and (d) fabricated sensor unit.

was absorbed into the SBS polymer coated on the cotton patch. Percolation of the AgNPs inside the fibers of the patch led to a low resistivity with large deformations. The method employed for the fabrication of the conductive fabric patch is analogous to our previous work, where we have introduced the fabrication of AgNP-decorated conductive fibers [22]. When a patch of SBS-coated fiber is dipped into the AgCF3COO solution, the solvents are absorbed by the SBS-coated fabric. A representative image after immersing SBS-coated fabric into Ag precursor is shown in Fig. 4(a). After evaporation in the air for a few minutes, the Ag precursor was reduced into AgNPs by the solution of N2H4·4H2O, ethanol and DI water as shown in Fig. 4(b). During this process, an ion−dipole interaction is formed between the trifluoroacetate anions (CF3COO−) and hydroxyl groups (−OH) of the alcohols, which enables rapid absorption of both precursor and alcohols into the fabric. As prepared conductive fabric, drying after the repeated number of absorption and reduction cycle is shown in Fig. 4(c). The FESEM image of Fig. 4(d) illustrates the top view of the conductive fabric revealing the intertexture of the woven fabric. A representative highmagnification view of the microfibril surface morphology is shown in the inset of Fig. 4(d). Fig. 4(e) corresponds to the FE-SEM image of the cross-section of the conductive fabric patch. As seen from the figures, the fabric shows the uniform formation of AgNPs on the surface. The interconnected AgNPs provide the main contribution of electrical conductivity. The surface resistivity of the fabric was significantly lowered by a repeated absorption/reduction processes. The surface resistance profile of the conductive fabric against different cycles of absorption/ reduction of the Ag precursor is illustrated in Fig. 5. After eight cycles of the absorption/reduction processes, the fabric showed a surface resistance of 0.17 Ω sq−1. The electrical conductivity of the fabric was

resistance from R to R′. At the same time, new contact spots are also established for new current paths, which are represented by latent resistances (R′n+1 – R′m) as shown in Fig. 2(c). The conductance variation of the device is determined not only by the conducting capability of each bump but also the number of conducting paths. Since the microstructures are created randomly, the contact spots created will also be random in nature. When the pressure increases, the more contact area is accomplished and shorter features come in electrical contact, enabling a rapid increase in current. When the applied pressure is released, the structures recover their initial status because of the highly viscoelastic nature of PDMS. The pressure sensitivity of the device was compared with a sensor made of similar conductive fabric assembled with a base electrode grafted on plane PDMS. To facilitate feasible and cost-effective fabrication, readily available silicon carbide abrasive sandpaper was used to create microstructures on PDMS film. The advantage of using sandpaper is that it does not require any surfactant coating to avoid bonding between the polymer and template. The FE-SEM image of Fig. 3(a) depicts the morphology of the microstructured PDMS film. As seen from the figure it is clear that the structures are random with unequal feature sizes, which allow the device to deform more easily during compression because of the higher unloaded area free to expand. The experimental setup for the electrical characterization and measurement of the response of the sensor is presented in Fig. 3(b). 3.2. Characterization of the conductive fabric The conductive patch reinforced with AgNPs allows the formation of highly stretchable intertexture over a large-area. The Ag precursor 3

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Fig. 2. Schematic illustration of the working mechanism of the sensor: (a) Structural change in the behavior of the sensor under an external pressure; for increased pressure, the slightly touched contact spots are sufficiently touched and unconnected features connect, enhancing higher contact area (inset). Equivalent circuits of the sensor (b) without pressure and (c) with pressure. Each contact spots are represented by a parallel combination of resistors. Rex represents the external resistance of the electrode layer.

material always recovers completely even if the compression test is repeated. This is possibly because of the highly resilient nature of the PDMS material.

maintained by percolation of the AgNPs inside the patch. 3.3. Characterization of the sensing element 3.3.1. Hysteresis In order to analyze the hysteresis behavior of the sensor, the sensor was gradually loaded with compressive force and unloaded back to its original state. The stress-strain profile for the loading and unloading of various forces is shown in Fig. 6(a). We calculated the compressive stress (σ ) by applied force divided by the electrode area (3.8 cm2) and compressive strain (ε ) by compressed distance divided by the device thickness. by compressed distance divided by the device thickness. For the maximum compression of 1.15 mm, the force required was approximately 9.75 N corresponding to a strain of 71.87%. Hysteresis was calculated by,

(σL − σU ) Hysteresis (%) = × 100% σM

3.3.2. Sensitivity Among all of the features of a pressure sensor, sensitivity is considered as one of the most important parameters; it determines how well the sensor responds to subtle pressure changes. Here, we first present the characterization of pressure sensitivity (S) and the general analysis of pressure sensitivity. For quantitative analysis, the pressure sensitivity of a resistive pressure sensor is defined by,


δ (ΔR/ Ro) δP


Where ΔR is the change in resistance, Ro is the resistance without pressure, and P is the applied pressure. The slope of the trace in Fig. 6(b) represents the pressure sensitivity of the sensor. The sensor exhibited sensitivity of −0.3954 kPa−1 in the low-pressure regime (< 2.67 kPa), and it decreased to −0.0133 kPa−1 beyond 2.67 kPa. In the first region (< 2.67 kPa), the elastic microstructure deforms rapidly, leading to an abrupt increase in contact spots, thereby decreasing contact resistance. However, in the second region (> 2.67 kPa), the deformation of the microstructure would have saturated, such that a large additional force can induce a small change in effective contact spots and also low contact resistance. Meanwhile, the reference sensor fabricated with a plane structure showed little


where the subscripts L, U, and M represent the loading, unloading, and maximum values of the stress at a particular strain, respectively. An average value of hysteresis was calculated to be around 3.17% for entire range up to maximum pressure of 25.39 kPa, which is much lower than the hysteresis observed for conductive foams [23]. The hysteresis at lower values of stress-strain is much lower than the average value of the hysteresis, indicating the effectiveness of the sensor in the lowpressure region. The low value of hysteresis error shows that the 4

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Fig. 5. Surface resistance profile of the conductive fabric with dipping cycles. The surface resistance of 0.17 Ω sq−1 was achieved after a repeated absorption/ reduction of eight cycles.

exhibited sensitivity almost five times higher than that of its unstructured counterpart. The sensitivity achieved for the low-pressure region is significantly higher than the sensitivities reported for the graphene-PU sponge (0.26 kPa−1) [12], and foam-type pressure sensor (0.338 kPa−1) [24]. The sensitivity performance of the sensor is mainly influenced by the number of patterns, shape factor, and the extent of deformation of the features. The microstructure with a smaller shape factor has more free area to expand when compressed, and the material with low elastic modulus can deform more easily even under pressures. In addition, an advantage of such a geometry is the reduced temperature dependency, which is a more challenging issue in elastomeric materials in which the interfiller distances change during the thermal expansion of the composites [25]. Next, we analyzed the stretching behavior of our sensor by applying tensile strain. The device can detect a change in resistance due to stretching/contracting of the sensor attached; therefore, it is more

Fig. 3. (a) FE-SEM image of the micropatterned PDMS before Ti/Au sputtering. (b) Schematic of the measurement setup for the measurement of sensor response.

response compared to that of the microstructured type, in which pressure response originates mainly because of the compression of the conductive polymer film. It was observed that the microstructure sensor

Fig. 4. Photograph of the (a) SBS-coated cotton fabric immersed in the precursor solution of AgCF3COO, (b) corresponding image after reducing by N2H4·4H2O solution and (c) final image after a repeated number of absorption/reduction followed by drying in ambient temperature. (d) FE-SEM images of the top view showing the weaving intertexture of the conductive fabric. The inset shows the magnification view of the microfibril indicating a uniform formation and interconnection of the AgNPs. (e) FE−SEM image of the cross-sectional view of the conductive fabric patch. 5

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Fig. 6. Characterization of the microstructured PDMS-based resistive pressure sensor: (a) Stress-strain characteristic of the sensor for loading/unloading of pressures. (b) Comparison of the relative change in resistance of the fabricated sensor with that of the unstructured type. (c) Force vs tensile strain up to safer limit of strain. (d) Relative change in resistance with tensile strain representing the gauge factor. (e) Phase of the resistance with pressure. (f) I-V characteristic of the sensor for different values of the applied pressures. The current through the sensor linearly increased with the pressure up to 2.67 kPa.

stretch the device by 10% strain, the force required was as large as 89.5 N. The relative change in resistance as a function of tensile strain is illustrated in Fig. 6(d). With the increase in strain, the resistance decreased linearly and finally reached to approximately 298 mΩ from its initial value. The slope of the line represents the strain sensitivity of the device illustrating the GF of −3.25, higher than recently reported graphene/AgNWs porous sea sponge [17]. Whenever a higher strain is applied, the longitudinal distance between AgNPs was increased and crack formed between the AgNPs networks. In order to detect small strain such as the skin of face, eyes or neck, the sensor needs to have higher strain sensitivity (i.e. GF) and needs higher stretchability to

relevant to express the performance metric in terms of gauge factor (GF). The gauge factor of the strain sensor is defined by,

GF =

δ (ΔR/ Ro) δε


Where ΔR is the change in resistance caused by applied stain ε and Ro is the resistance before straining. The sensor was subjected to tensile strain up to a safer limit of 10%, well enough to detect very small skin strain. The repeatability of the sensor response was well maintained within this limit of strain. As seen from Fig. 6(c), relatively very small force is sufficient to stretch the sensor for a lower strain. However, to 6

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Fig. 7. (a) Variation of resistance for 30 cycles of compression/release. Time response of the sensor for loading/unloading of (b) 7.67 kPa and (c) 0.69 kPa. Inset shows the response and relaxation time of the sensor. (d) The stability of the sensor for > 6000 compression/release cycles. Table 1 Performance comparison of the various resistive sensors. Sensing Material


Sensitivity (kPa−1)

Working Pressure range (kPa)

Response time (ms)


Multiwall CNT-PDMS Laser-scribed graphene rGO-PU Graphene/AgNWs Hybrid foam sensor Micropatterned PDMS Au-coated PU Metal-deposited PDMS

Sponge Pyramid Sponge Sea sponge Foam Pyramid Sponge Microstructures

< 0.025 0.005–0.96 0.26 0.016 0.001–0.338 0.01–0.31 0.02–0.31 0.3954

0−100 kPa 0–120 10 kPa 0–10 0–240 0–13 0–10 0–25.39

– 0.4–72 – 54 8.5–60 <5 < 10 0.49 s

[4] [8] [12] [17] [24] [26] [27] This work


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Fig. 8. (a) Photograph of the sensor during gentle finger tapping and no tapping. (b) The response of the sensor to gentle finger tapping (at the frequency of ∼1 Hz) and releasing each for 10 s. (c) The response of the sensor to tapping (at the frequency of ∼2 Hz) and releasing. (d) Photograph of the sensor fastened on the forefinger for folding and relaxing of the finger. (e) The response of the sensor to folding by the curvature of ∼0.85 cm−1 and relaxing of the finger each for 10 s. (f) The response of the sensor to folding by a higher curvature of ∼0.96 cm−1 with a slightly higher folding rate and relaxing.

with the pressure applied owing to the corresponding decrease in resistance.

make conformal contact to the skin with high signal quality [14]. For every pressure loading, the resistance along with phase was also measured. The phase-pressure plot of Fig. 6(e) shows that the phase of the device measured for every value of pressure was 0° indicating the purely resistive nature of the device.

3.3.4. Response time The resistance variations in both of the pressure regions were analyzed by compression/release of 0.69 and 7.67 kPa for 30 cycles. An average value of resistance in the no-pressure condition was measured to be approximately 2.33 Ω. When the pressure was increased from 0 to 0.69 kPa, the resistance decreased to 1.3 Ω from its initial value. Similarly, for the compression/release of 7.67 kPa, the resistance decreased to 1.1 Ω as illustrated in Fig. 7(a). The time-response of the resistance to compression/release of both 7.67 kPa and 0.69 kPa is shown in Fig. 7(b) and (c), respectively. As the moving stage fixture

3.3.3. Linearity The current-voltage (I-V) characteristics under four different pressure values (0.2, 0.5, 1, and 2 kPa) are shown in Fig. 6(f). The currents were measured for potential ranging from −1 to 1 V at each pressure values. The linear relationship between voltage and current indicates that the sensor obeys Ohm's law, and there is an obvious increase in current with increased pressure. The slope of the I-V curves increased 8

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without surfactant coatings. The conductive fabric was fabricated by reinforcing AgNPs through a repeated absorption/reduction cycles of Ag precursor. A relatively high sensitivity of 0.3954 kPa−1 in the lowpressure regime (< 2.67 kPa), an ultra-low hysteresis of 3.17%, and fast response time of 0.49 s were achieved. Moreover, to examine the sensing capability of the device, some application tests, which mimic the electronic skin, were also conducted. The facile and cost-effective fabrication approach, structure and mechanism introduced here present a robust technology platform to enhance the hysteresis, sensitivity, and linearity over the conventional methods for wearable and flexible sensor applications. Further optimization of the geometry and material will be of paramount interest to improve the sensitivity and fully utilize in electronic skins. Thus, the development of a pressure sensor that can mimic and distinguish subtle pressure sensing properties is important because it can be utilized widely for pressure sensors, strain sensors and impact sensors in various wearable electronic applications.

moved with constant speed of 1 mm/min, the time required to complete one compression/release cycle for two different pressures (controlled by compression distance) was different. The sensor's response and relaxation time is demonstrated in the inset of Fig. 7(c). A response time of 0.49 s and relaxation time of 0.56 s were measured. The dynamic response and relaxation time could be even lower than measured value however; we could not measure lower than 0.49 s due to the limitation of LCR meter. The response and relaxation time significantly increases from low to medium pressure values because of the viscoelastic creep of the material [24]. A performance comparison of the related works those used various sensing materials is presented in Table 1. 3.3.5. Stability Furthermore, to evaluate the multiple cycling stability of the sensor, continuous loading and unloading of the pressure of 140 Pa over 6000 cycles at 0.5 Hz was performed. As shown in Fig. 7(d), during the loading of the pressure of 140 Pa, the resistance decreased from an average value of 2.33 to 2.14 Ω. During the entire cycling test, the highest (unloading) and the lowest value (loading) of the resistance recorded were 2.39 Ω and 2.12 Ω, respectively. For each cycling test, the shape of the curve maintained well, indicating an excellent repeatability and reproducibility of the sensor.

Acknowledgements This research was supported by the Bio & Medical Technology Development Program of the NRF grant funded by the Korean government (MSIT) (NRF-2017M3A9F1031270). The authors are grateful to the MiNDaP group members of Kwangwoon University for their technical discussion and support.

3.4. Pressure and strain sensing


To demonstrate the applicability of the sensor for wearable electronics, pressure and strain sensing performances were investigated separately by examining the response of the sensor for subtle pressure, and tensile strain, respectively. For the purpose of subtle pressure detection, the response of the sensor was observed by gentle finger tapping and releasing as shown in Fig. 8(a). The response was first analyzed with a sequence of no-tapping and tapping at the frequency of ∼1 Hz each for 10 s and then a similar sequence of action at the higher frequency of ∼2 Hz. The relative change in resistance for the sequence of no-tapping and tapping at 1 Hz is shown in Fig. 8(b). Similarly, the relative change in resistance response for the sequence of no-tapping and tapping at the frequency of 2 Hz is shown in Fig. 8(c). In each response, the resistance decreased from an initial value of 2.33 Ω to approximately 470 mΩ with an appreciable transition in the pressures. In addition, our sensor has also the ability to detect tensile strain. For this purpose, the sensor was fastened on the forefinger to confirm its capability of detecting skin movements as shown in Fig. 8(d). The relaxing state of the finger corresponds to the zero curvature radius whereas bending by a positive curvature radius represents tensile strain. The relative change in resistance to the consecutive folding (curvature of ∼0.85 cm−1) and relaxing of the finger each for 10 s is illustrated in Fig. 8(e). Similarly, Fig. 8(f) depicts a similar response for the sequence of folding (by a higher curvature of ∼0.96 cm−1) at a slightly higher frequency and relaxing. As seen from the figures, the resistance of the finger-mounted sensor increases upon folding the finger and regains its original value when relaxed. The ability of the sensor to induce excellent resistance change for pressure as well as strain confirms that our sensor is well suited for artificial skins. Such sensors incorporated into clothing or even attached directly to the body are in imperative demands for the unprecedented applications such as human motion detection, wearable health-monitoring, and humanmachine interface.

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4. Conclusions In this study, a flexible and stretchable resistive sensor was fabricated, characterized, and examined the capability of pressure and strain sensing. To achieve this, an assembly of Au-sputtered microstructured PDMS as the base electrode and AgNPs-reinforced conductive fabric as the counter electrode was used. The microstructured PDMS film was replicated directly from the low-surface-energy sandpaper template 9

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