Highly Conformable, Transparent Electrodes for Epidermal Electronics

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Highly Conformable, Transparent Electrodes for Epidermal Electronics Jin-Hoon Kim, Seung-Rok Kim, Hye-Jun Kil, Yu-Chan Kim, and Jin-Woo Park Nano Lett., Just Accepted Manuscript • DOI: 10.1021/acs.nanolett.8b01743 • Publication Date (Web): 20 Jun 2018 Downloaded from http://pubs.acs.org on June 21, 2018

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Highly Conformable, Transparent Electrodes for Epidermal Electronics Jin-Hoon Kim†, Seung-Rok Kim†, Hye-Jun Kil‡, Yu-Chan Kim‡, and Jin-Woo Park*,† †

Department of Materials Science and Engineering, Yonsei University, Seoul, 03722, Korea

‡

Biomedical Research Institute, Korea Institute of Science and Technology, Seoul, 02792, Korea

We present a highly conformable, stretchable, and transparent electrode for application in epidermal electronics based on polydimethylsiloxane (PDMS) and Ag nanowire (AgNW) networks. With the addition of a small amount of a commercially available non-ionic surfactant, Triton X, PDMS became highly adhesive and mechanically compliant, which are key factors for the development of conformable and stretchable substrates. The polar functional groups present in Triton X interacted with the Pt catalyst present in the PDMS curing agent, thereby hindering the crosslinking reaction of PDMS and modulating the mechanical properties of the polymer. Due to the strong interactions that occur between the polar functional groups of Triton X and AgNWs, AgNWs were effectively embedded in the adhesive PDMS (a-PDMS) matrix, and the highly enhanced conformability, mechanical stretchability, and transparency of the a-PDMS matrix were maintained in the resulting AgNW-embedded a-PDMS matrix. Finally, wearable strain and electrocardiogram (ECG) sensors were fabricated from the AgNW-embedded a-

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PDMS. The a-PDMS-based strain and ECG sensors exhibited significantly improved sensing performances compared to those of the bare PDMS-based sensors because of the better stretchability and conformability to skin of the former sensors.

KEYWORDS: conformability, adhesive electrodes, AgNWs, strain sensor, ECG

Wearable and implantable biosensors for the real-time monitoring of biosignals, such as body motion, blood pressure, electrocardiogram (ECG), electroencephalography (EEG), and electromyogram (EMG), have become one of the most important elements of wearable devices for healthcare and the internet of things (IoT).1, 2 These biosensors are attached to human skin or organs to detect various biosignals. Hence, the maintenance of good conformal contact between the sensor and skin is essential for obtaining precise biosignals because the human skin and organs are curved surfaces that move dynamically. That is, slippage, delamination, and breakage of the biosensors can occur if the adhesion between the biosensors and the skin surface is poor.2-4 Furthermore, wearable biosensors are preferred to be optically transparent for aesthetic reasons and due to their integration with optoelectronic devices.5 Various materials, such as hydrogels and adhesive polymers, have been used to improve conformal contact by increasing the adhesion force between the biosensors and the skin or organs.4, 6-8 Hydrogels are good candidates for adhesive electrodes due to not only their high optical transparency and biocompatibility but also their good ionic conductivity, which is one of the essential properties in the measurement of biosignals.9 However, the adhesion force of hydrogels is not sufficient to maintain conformal contact between the biosensors and the skin or organs.4 Additionally, the performance of hydrogels decreases significantly with time as the


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water in the hydrogels evaporates.10 Furthermore, most hydrogels are mechanically weak and thus cannot be applied to the dynamically moving parts of the human body, such as the elbow, fingers, and heart. Elastomeric polymers with Young’s moduli similar to that of human skin, such as polyacrylate, hydrocolloids, rubber, polyurethane and silicone, have been used as adhesive materials between sensors and human skin.11 Among these polymers, polyacrylate, hydrocolloids, rubber, and polyurethane are called aggressive adhesives because they cause skin trauma, such as ripping, irritation, maceration and allergic reactions.11 Furthermore, polymer residues are left on the skin or organs following detachment.12 In contrast, the use of siliconebased polymers can minimize the skin trauma and residue associated with skin adhesives.11, 13 Consequently, silicone-based polymers not only are highly biocompatible (non-toxic, nonirritating, and non-sensitizing) but also have superior mechanical stretchability, temperature stability, chemical inertness, and environmental stability.11 The adhesion force of silicone polymers can be easily enhanced by mixing tackifier or additives, such as a polyethylenimine ethoxylated (PEIE), into the formulation.6,

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As most

biosignals are interpreted as electrical signals, biosensors should be electrically conductive.1, 2, 6 However, silicone polymers are electrically insulating; hence, conductive electrodes need to be coated or printed on the biosensor to establish electrical contact between the biosensors and the skin or organs for biosignal detection. Solid conductive fillers such as conducting polymers, metal nanoparticles, carbon powder, and carbon nanotubes (CNT) have been integrated into adhesive elastomers for biosensor applications.14-18 However, to achieve a low electrical resistance for excellent biosignal detection, high concentrations of the conductive filler materials


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are needed, which results in very low optical transmittance and changes in the mechanical compliance of the polymer matrices. In this paper, we fabricated highly conformable and transparent electrodes by embedding Ag nanowires (AgNWs) in an adhesive PDMS (a-PDMS) matrix. The physical properties, such as the adhesiveness, Young’s modulus, and viscoelasticity of the PDMS, were tuned by simply adding a commercially available non-ionic surfactant, Triton X. Triton X could inhibit the PDMS crosslinking reaction and thus resulted in heterogeneously crosslinked networks of polymer chains. To verify the applicability of the AgNW-embedded a-PDMS-based conformable electrodes in epidermal electronics, strain sensors and ECG sensors were fabricated from the conformable electrodes and compared with sensors prepared from bare PDMS with embedded AgNWs.


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Figure 1. (a) Schematic images of the fabrication processes of the transparent and adhesive electrodes. (b) Chemical structures of Triton X, PDMS base, and PDMS crosslinker. (c) Conceptual illustration of the adhesive and transparent electrodes attached to human skin. (d) Optical photographs of the adhesive electrodes attached to the wrist and forearm for strain and ECG sensors, respectively, and the electrical signals expected from the sensors. Figure 1a shows the fabrication process of the conformable electrodes using a-PDMS with embedded AgNWs. To modify the mechanical properties, such as Young’s modulus, failure strain, viscoelasticity, and adhesion force, of PDMS, Triton X was added to the liquid mixture of PDMS base and crosslinker. Figure 1b shows the chemical structures of Triton X, PDMS base, and PDMS crosslinker. To embed AgNWs in the a-PDMS film, AgNWs were spin-coated on a glass release substrate prior to coating of the liquid a-PDMS. After the curing of a-PDMS, the electrode film was released from the glass substrate. Figure 1c shows actual images of the conformable, stretchable, and transparent electrodes developed for epidermal electronics based on a-PDMS and AgNWs in this work. Owing to the high adhesion force and low Young’s modulus of a-PDMS, the electrode attached to the forefinger showed good conformal contact with the skin surface and was not delaminated under repeated bending of the finger, as shown in Figure 1c. As shown in the field emission scanning electron microscopy (FE-SEM) image in Figure 1c, AgNWs were effectively embedded in the aPDMS matrix to introduce electrically conducting properties in the otherwise insulating a-PDMS. In Figure 1d, the a-PDMS-based electrodes are directly attached to the human body for use as strain and ECG sensors. The purpose of strain sensors is to monitor the motion of the human body, while ECG sensors are utilized to monitor the function and structure of the heart, as schematically illustrated in Figure 1d.


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Figure 2. (a) Optical transmittance of a-PDMS as a function of the weight ratio of Triton X. (b) Stress-strain curves of a-PDMS films with various weight fractions of Triton X and Tc. (c) Viscoelasticity measurements of a-PDMS films with various weight fractions of Triton X and Tc. (d) Adhesion test results for bare PDMS and a-PDMS as a function of the weight ratio of Triton X and Tc. The inset image shows a schematic of the peel test. (e) Optical transmittance and (f) cyclic stretching test results for the adhesive and transparent electrodes with embedded AgNWs. (g) Cell viability and (h) cell proliferation rate analysis results for a-PDMS and PDMS. Table 1. Detailed descriptions of the samples and varying parameters.

Sample name

wt% of Triton X in the PDMS base (%)

PDMS_40 PDMS_70

40 0

a3-PDMS_40 a3-PDMS_50

Tc (oC)

70 40

0.3

50

a3-PDMS_70

70

a4-PDMS_40

40

a4-PDMS_50

0.4

a4-PDMS_70

50 70

a5-PDMS_40

0.5

40

a8-PDMS_40

0.8

40


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To change the mechanical properties such as Young’s modulus, failure strain, viscoelasticity, and adhesion force of PDMS, both the concentration of Triton X in the PDMS matrix and the curing temperature (Tc) were varied. Detailed descriptions of the samples and the control parameters are summarized in Table 1. Figure 2a shows the optical transmittance (T) of the aPDMS films (thickness of 2 mm) with various contents of Triton X prepared at a fixed Tc of 40 °C. Figure S1 (Supporting Information) shows photographs of the a-PDMS films listed in Figure 2a. As shown in Figure 2a, with increasing content of Triton X, the T of a-PDMS decreases. Furthermore, according to Figure S2 (Supporting Information), the effect of Tc on the T of the aPDMS was almost negligible. The PDMS film without Triton X (PDMS_40) exhibits the highest T with a value of approximately 94.3% at a 550-nm wavelength. The T of the a-PDMS films containing 0.3 wt% Triton X (a3-PDMS_40) and 0.4 wt% Triton X (a4-PDMS_40) decrease to 91.4% and 84.7%, respectively. To fabricate transparent electrodes based on aPDMS, the T of a-PDMS should be higher than 80%. Hence, a3-PDMS_40 and a4-PDMS_40 were selected for further analysis. The reason for the decrease in T with increasing content of Triton X is the scattering of light by the micelle structures formed in the PDMS matrix with the introduction of Triton X.19 Since the PDMS matrix is highly hydrophobic, the alkyl part at the Triton X chain forms the shell, while the corresponding polyethylene glycol (PEG) portion forms the core of the micelle structure.19, 20 As the concentration of the surfactant increases, the light scattering intensity is well known to also increase because of the increase in the size of the micelles.21 Figure 2b compares the stress-strain curves of the bare PDMS and a-PDMS films listed in Table 1. Based on the results shown in Figure 2a, Young’s modulus and the failure strains were


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measured in only the samples with 0.3 or 0.4 wt% Triton X prepared at varying Tc values. The mechanical properties of the samples plotted in Figure 2b are summarized in Table 2. According to Figure 2b and Table 2, a3-PDMS_40 and a4-PDMS_40 (in Table 1) show the lowest Young’s modulus and the highest failure strain among all the samples. Young’s modulus and the failure strain of PDMS_40 were 500 kPa and 230%, respectively. In contrast, Young’s moduli of a3PDMS_40 and a4-PDMS_40 were 38 and 40 kPa, respectively, and both films showed failure strains over 400%. Young’s moduli of a3-PDMS_40 and a4-PDMS_40 were much lower than that of human skin (between 500 kPa to 1 MPa), which is good for epidermal electronics because the conformability of the electrodes increases with decreasing Young’s modulus of the polymer matrix.4, 6, 22 When the polymer matrix has a lower Young’s modulus than the skin, the contact area between the electrode and the skin is enlarged and the users feel less discomfort. Photographs of the a4-PDMS_40 film during a uniaxial stretching test are shown in Figure S3 (Supporting Information). According to Figure 2b, Young’s modulus tends to increase with increasing Tc of a-PDMS. According to the previous results obtained by others, Young’s modulus of PDMS could be decreased by reducing the amount of crosslinker added in the PDMS base.6 However, the failure strain of PDMS also decreased.6 In contrast, the a3-PDMS_40 and a4-PDMS_40 films presented in Figure 2b show highly enhanced failure strains, which indicates that these two a-PDMS matrices are good candidates for application as the transparent and stretchable polymer matrix of epidermal electronics.

Table 2. Mechanical properties of the PDMS and a-PDMS films analyzed in Figure 2b.

Sample name

Young’s modulus (kPa)

Failure strain (%)


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PDMS_40

480 ± 30

230

a3-PDMS_40

38 ± 6.3

> 400

a4-PDMS_40

40 ± 5

> 400

a3-PDMS_50

194 ± 7.2

> 300

a4-PDMS_50

162 ± 20

> 300

a3-PDMS_70

1000 ± 50

220

a4-PDMS_70

810 ± 20

210

The viscoelasticities of the bare PDMS and a-PDMS films were measured using a dynamic mechanical analyzer.23 The loss tangent (tan δ) is the tangent of the phase angle (δ) and is defined as:

(1)

where E` and E`` indicate the elastic and viscoelastic properties of the elastomer, respectively.11 Hence, an elastomer shows more viscoelastic behavior when it has a higher tan δ value. The measured tan δ values of the PDMS_40, a4-PDMS_40, a4-PDMS_50, and a4PDMS_70 films are compared in Figure 2c.23 Since the a3-PDMS_40 and a4-PDMS_40 films showed similar results in the uniaxial tensile test (Figure 2b), only the a-PDMS film with 0. 4 wt% Triton X was analyzed. According to Figure 2c, as Tc decreases, tan δ increases for the a-PDMS film with 0. 4 wt% Triton X, which indicates that this film showed more viscoelastic behavior


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with a lower Tc. The a-PDMS _40 film showed a high tan δ value of approximately 0.5 even at a low frequency (0.1 Hz). In the previous work of other researchers, the tan δ value of silicone rubber was shown to be less than 0.4 or even below 0.1,24, 25 which indicates that a-PDMS _40 possesses highly viscoelastic behavior comparable to that of other silicone-based elastomers. Based on a uniaxial tension test and dynamic mechanical analysis, PDMS becomes substantially more compliant and viscoelastic than bare PDMS when 0.3 or 0.4 wt% Triton X is added at a fixed Tc of 40 °C. The adhesion forces of the a-PDMS and bare PDMS films were measured by a peel test following ASTM D3330 standards, and the results are compared in Figure 2d.26 The peeling angle was 90°, and the corresponding test setup is schematically illustrated in the inset image of Figure 2d. As shown in Figure 2d, the a3-PDMS_40 and the a4-PDMS_40 films showed the highest adhesion force among all the samples. The adhesion force of the a3-PDMS_40 film is 35 N m-1, which is seven times higher than that of the bare PDMS film according to Figure 2d. The force-distance curves obtained from the peel test are shown in Figure S4 (Supporting Information). The peel test results shown in Figure 2d indicate that the Tc significantly affects the adhesion force of a-PDMS. The very low Young’s modulus and viscoelasticity of the a-PDMS film cured at 40 °C led to enhanced wetting and spreading of the polymer chains in a-PDMS, which resulted in improved surface contact, that is, a higher adhesion force of the a-PDMS film.6, 27, 28 As shown in Figure S5 and Movie S1 (Supporting Information), a4-PDMS_40 could even withstand a load of a 50 g weight and remain perfectly attached to the forefinger without delamination. In summary, according to Figure 2, the optical properties of a-PDMS are dramatically affected by the content of Triton X in PDMS, and the mechanical properties (Young’s modulus, failure


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strain, viscoelasticity, and adhesion force) are significantly affected by the Tc of a-PDMS. According to the previous study by other researchers, Triton X have been used to modify the properties of the PDMS.29 The Triton X was added to the PDMS to modify its surface properties such as the surface energy and protein affinity.29 In contrast, in our current work, it is shown that the Triton X could modify the bulk properties of the PDMS such as the Young`s modulus, elongation at break, viscoelasticity, and adhesion force. Hence, although the methods used in here are similar to the previously reported work by Holczer et al., our approaches and results were quite different from the previously published work in which we focused on the modification of the bulk properties of the PDMS and not of its surface properties. In order to analyze the time dependent properties of the a4-PDMS_40, uniaxial tensile and adhesion tests were done on the a4-PDMS_40 stored for a week, and 2 months under ambient atmospheric conditions. As shown in Figure S6a and Table S1 (Supporting Information), the Young`s modulus of the a4-PDMS_40 did not change after being stored for a week after its fabrication. The Young`s modulus of the 2 month-aged a4-PDMS_40 increased from 40 to 50 kPa or equivalent to about 25% increase. The adhesion test was conducted by suspending 50 g weight from a forefinger using a4-PDMS_40 as the adhesive. As shown in Figure S6b to c, and Movie S2 to S3 (Supporting Information), both a4-PDMS_40 stored for a week, and 2 months held the 50 g weight attached by them to the forefinger without any delamination. Based on these results, it was verified that the a4-PDMS_40 could maintain its initial properties for a long period of time. To explain the effects of the Triton X content on Young’s modulus, the viscoelasticity, and the adhesion force of PDMS, the swelling ratios30 and gel fractions31 of the a-PDMS and the PDMS_40 films were measured and compared using chloroform and toluene as solvents. As


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shown in Figure S7 (Supporting Information), the swelling ratio of the a4-PDMS_40 film is 2.5 and 2 times higher than that of the PDMS_40 film in chloroform and toluene, respectively. Furthermore, the gel fractions of the a4-PDMS_40 film are 0.77 and 0.8 in toluene and chloroform, respectively, whereas PDMS_40 has values of 1.0 and 0.95 in toluene and chloroform, respectively. These results indicate that the a4-PDMS_40 film has a lower crosslinking density or a more heterogeneous crosslinking distribution than PDMS, which led to the changes in the mechanical properties of a-PDMS shown in Figure 2b, c, and d.6 The crosslinking of PDMS occurs through hydrosilylation with a Pt catalyst.32 The Pt catalyst diffuses through the PDMS matrix to complete the crosslinking reaction. Since the Pt catalyst is coordinatively unsaturated, it forms complexes with other polar functional groups, such as the PEG chain of Triton X.6, 33 Moreover, Triton X forms core-shell structures inside the PDMS matrix. Hence, depletion of the amount of active Pt catalyst present in the PDMS matrix occurs, as the non-polar functional groups enclose the Pt-interacting polar groups of the Triton X coreshell structure inside the PDMS matrix. Hence, the crosslinking reaction is inhibited by the addition of a small amount of Triton X into the PDMS mixture through deactivation of the Pt catalyst, which is the main mechanism for the formation of heterogeneously crosslinked networks in PDMS.6 By the Triton X molecules, a heterogeneously crosslinked network composed of crosslinked and non-crosslinked PDMS forms in the a-PDMS matrix. Such a composite-like structure results in modulation of the mechanical properties, such as Young’s modulus, failure strain, viscoelasticity, and adhesion force, to form a soft and adhesive PDMS matrix.6 The depletion of the amount of active Pt catalyst during curing of the PDMS mixture reduced or hindered the crosslinking reaction of PDMS and thus resulted in a heterogeneously


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crosslinked PDMS matrix. As previously mentioned, the Tc is the most important factor that determines the mechanical properties (Young’s modulus, failure strain, viscoelasticity and adhesion force) of a-PDMS. The polarity of Triton X decreases as the temperature increases, which results in disruption of the micelle structure and weakening of the interactions between Triton X and the Pt catalyst.34 Hence, as Tc increases, the Pt catalyst is more likely to diffuse back into the PDMS matrix and outside the micelle structure formed by Triton X, resulting in an increase in the number of crosslinking reactions that occur in a-PDMS, which indicates that the Tc of a-PDMS should be low. Adhesive electrodes based on a-PDMS embedded with AgNWs were successfully fabricated. These electrodes have a sheet resistance (Rs) of 35 Ω⋅sq-1 and a T of 75%, as shown in Figure 2e. Detailed descriptions of the embedding process and the shapes of the electrode samples are shown in Figure 1a and 1b. A cyclic stretching test was done for the a-PDMS and PDMS-based electrodes (a4-PDMS_40 and PDMS_40) to evaluate their stretchability. Since the Rs of the a3PDMS_40-based electrodes (a3-PDMS_40NW) was too high for application in transparent electrodes (150 Ω ⋅sq-1), only the a4-PDMS_40-based electrodes (a4-PDMS_40NW) were used as the specimens for the stretching tests. The cyclic stretching test was conducted at a strain of 15% for 10000 cycles. Interestingly, as shown in Figure 2f, the relative resistance change (∆R ⋅R0-1) of the a4-PDMS_40NW measured after 2000 cycles in the cyclic stretching test decreased slightly to -0.15 and reached 0.32 after 10000 cycles of stretching, whereas the ∆R⋅R0-1 of the PDMS_40based electrodes (PDMS_40NW) kept increasing from the beginning of the stretching test and reached up to 1.3 after 10000 stretching cycles. The mechanism for the stable stretchability of the a4-PDMS_40NW electrode was analyzed using FE-SEM (Figure S8 in the Supporting Information). As shown in Figure S8a,


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PDMS_40NW exhibited severe delamination of the AgNWs even when not stretched, which indicates that the AgNWs were not embedded well in the PDMS matrix due to the low adhesion force between PDMS and the AgNWs.35 Moreover, individual AgNWs were fractured after the cyclic stretching tests, as shown in Figure S8b (Supporting Information). On the other hand, a4PDMS_40NW did not show any delamination of the AgNWs (Figure S8c in the Supporting Information). Furthermore, the AgNWs were not fractured after the cyclic stretching test of the a4-PDMS_40NW electrode (Figure S8d in the Supporting Information), which indicates that a4PDMS_40NW has better adhesion than PDMS_40NW with the AgNWs. As the PEG chain in Triton X can participate in electrostatic interactions with the AgNWs, the adhesion between a4-PDMS_40 and the AgNWs was stronger than that between PDMS_40 and the AgNWs.35 The better adhesion between a4-PDMS_40 and the AgNWs resulted in improved stretchability for the a4-PDMS_40NW electrode.36 The surface of the glass substrates after releasing the a4-PDMS_40NW and PDMS_40NW was analyzed using the FE-SEM to verify the strong electrostatic interactions between Triton X and the AgNWs. As shown in Figure S9 (Supporting Information), many AgNWs remained on the glass substrate after releasing the PDMS_40NW, whereas, no observable AgNWs remained on the glass substrate after releasing the a4-PDMS_40NW. These FE-SEM analysis results further verify the strong electrostatic interaction between the AgNWs and Triton X. The other reason for the decrease in Rs after 2000 cycles of stretching was a change in the alignment of the AgNWs along the stretching direction. According to previous studies by other researchers, AgNWs can become aligned along the stretching direction during mechanical stretching.37, 38 Because the adhesion between the AgNWs and a4-PDMS_40 is strong due to the


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presence of Triton X, the AgNWs embedded in a4-PDMS_40 can align along the stretching direction during cyclic stretching. The biocompatibilities of the a-PDMS (a4-PDMS_40 and a4-PDMS_70) and PDMS (PDMS_40 and PDMS_70) matrices were evaluated using cell viability tests (Figure 2g)39 and the proliferation rate of fibroblast cells (Figure 2h).31 The samples were selected to verify the effect of Triton X and the Tc on the biocompatibility of the a-PDMS matrix. The cell viability of the samples is defined as the ratio of the number of cells grown on the sample surface to the number of cells grown on a control sample (the environment in which cells grow under the best conditions). If the cell viability is over 80%, the sample is considered biocompatible. As shown in Figure 2g, the four tested samples (Table 1) have cell viabilities over 80%, which indicates that Triton X does not affect the biocompatibility of PDMS. Optical microscopy (OM) images of the fibroblast cells on each sample are shown in Figure S10 (Supporting Information). A low cell proliferation rate is especially desirable for biosensor applications because if cells grow on the biosensor or at the interface between the biosensor and organ, the biosensor cannot accurately detect the biosignals.40 Accordingly, the low proliferation rate of a-PDMS was expected to be beneficial for biosignal recording purposes. The proliferation rates of fibroblast cells

41, 42

on the samples are presented in Figure 2h. The absorbances shown in Figure 2h were

measured using the CCK-8 kit.39 The absorbance is proportional to the proliferation rate of the cells because the CCK-8 kit detects the signal of living cells. According to Figure 2h, the addition of Triton X and lower Tc values independently lead to a decrease in the cell proliferation rate. In particular, a4-PDMS_40 shows a significantly lower cell proliferation rate than the other samples. The low cell proliferation rate of the a4-PDMS_40 sample is attributed to the high surface roughness of a-PDMS.43 The surface roughness (Rrms) of the a4-PDMS_40 and a4-


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PDMS_70 was 878 and 79 nm, respectively. The surface profile images of the a4-PDMS_40 and a4-PDMS_70 are presented in Figure S11 (Supporting Information), which shows the significantly higher surface roughness of the a4-PDMS_40 as compared with the a4-PDMS_70. For cells to proliferate, the surface needs to be smooth enough to allow the spreading and growth of cells.44


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Figure 3. Photographs of strain sensors based on (a) a4-PDMS_40NW and (b) PDMS_40NW attached to the wrist in the straight and bent states. The electrical resistance changes in the (c) a4PDMS_40NW- and (d) PDMS_40NW-based strain sensors during their operation on the human wrist. (e) Photographs of the a4-PDMS_40NW-based sensor attached to the forearm for skin impedance measurements. (f) Electrode-skin impedance measurements of a4-PDMS_40 and PDMS_40 as a function of the frequency. ECG signals recorded from (g) commercial gel, (h) PDMS_40NW-based and (i) a4-PDMS_40NW-based ECG sensors.

Using a4-PDMS_40NW and PDMS_40NW, strain and ECG sensors were fabricated. The strain sensor used in this study was a resistive-type sensor. The sensing mechanism of the resistive-type strain sensor is based on the the change of the electrical resistance of the a-PDMS


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based electrodes under deformational strain.45 As shown in Figure S12a to b (Supporting Information), the gauge factor of the a4-PDMS_40NW-based strain sensor was 1.4 in the 0 to 5% strain range, and 13 in the 10 to 20% strain range. The minimum limit of detection for the a4PDMS_40NW-based strain sensors was determined to be 1% strain (as shown in Figure S12c in the Supporting Information). Based on Figure S12d (Supporting Information), the sensing response time was around 140 ms. These sensing characteristics of the a4-PDMS_40NW-based strain sensors were comparable to the AgNWs-based strain sensors reported by other researchers.46, 47 The a4-PDMS_40NW demonstrates perfect conformal contact with the skin and is not delaminated following 10 repeated bending motions of the wrist, as shown in Figure 3a and Movie S4 (Supporting Information). However, the PDMS_40_NW sensor delaminated from the skin even on the first bending of the wrist, as shown in Figure 3b and Movie S5 (Supporting Information). The motion detection of the wrist using the a4-PDMS_40NW- and PDMS_40NW-based strain sensors are shown in Figure 3c and d, respectively. As shown in Figure 3c, a4-PDMS_40NW exhibits a larger change in electrical resistance than PDMS_40NW for the same applied deformation, which indicates that the sensitivity of a4-PDMS_40NW is significantly higher than that of PDMS_40NW at a fixed strain; this difference is due to the highly enhanced conformability and stretchability of a4-PDMS_40NW on the skin compared to PDMS_40NW. Since the a4-PDMS_40NW adhered well on the wrist, it could follow the entire movement of the wrist. On the contrary, due to the poor adhesion of the PDMS_40NW, it could not mimic the whole motion of the wrist. Furthermore, the a4-PDMS_40NW-based strain sensor shows low hysteresis, where the ∆R ⋅R0-1 remains at zero during cyclic bending of the wrist, as shown in


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Figure 3c. However, the PDMS_40NW strain sensor shows high hysteresis behavior, which significantly degrades the performance of the strain sensor.48 As shown Figure 3d, the ∆R⋅R0-1 ratio of the straight state continuously increases during multiple bending cycles of the wrist. An ECG sensor was also fabricated using a4-PDMS_40NW, as shown in Figure 3e. First, the skin impedance was measured using PDMS_40NW and a4-PDMS_40NW. The PDMS_40NW and a4-PDMS_40NW electrodes were placed on the forearm, as shown in Figure 3e and Figure S13 (Supporting Information). The skin impedance measured using PDMS_40NW is very high (Figure 3f), which is due to the low degree of conformability based on weak adhesion (Figure 2d) to human skin, as shown in Figure S13a (Supporting Information). Hence, the skin impedance of PDMS_40NW decreased with the application of adhesive tape. The skin impedance measured using a4-PDMS_40NW is significantly lower than that of PDMS_40NW even with the adhesive tape, which indicates that a high degree of conformability is very important for the recording of ECG signals. Furthermore, as shown in Figure S14 (Supporting Information), no residues were left on the skin after detachment of the a4-PDMS_40NW electrodes, which is beneficial because residues can cause skin irritation or allergic reactions.11 ECG sensors were fabricated based on the PDMS_40NW, a4-PDMS_40NW, and commercially available gel-based electrodes. Three electrodes were attached to the right and left sides of the chest and under the right side of the rib cage (acting as the ground electrode), as illustrated in Figure S15a (Supporting Information). As shown in Figure 3g, each wave of the ECG signals (with the composition of the P wave, QRS complex, and T wave) is well defined for the commercially available gel electrodes.49, 50 However, as shown in Figure 3h, the ECG signal from the PDMS_40NW-based sensors shows severe noise relative to the reference commercially available gel-based ECG sensors such that the P wave cannot be observed in some signals from


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the measurements. Similar to the skin impedance measurements, the noise of the ECG signal decreases slightly when adhesive tape is applied to PDMS_40, as shown in Figure S15b (Supporting Information). In contrast, the signal from the a4-PDMS_40NW-based ECG sensors shows negligible noise (fluctuations in the graph of the ECG signal, Figure 3i) that is even lower than the noise from the commercially available gel-based ECG sensors (Figure 3g). Figure S16 (Supporting Information) shows the ECG signals measured between the forearms with a ground electrode attached to the left leg (a schematic image of the location of each electrode is shown in Figure S16a (Supporting Information)). The noise in the signals in Figure S16b to d (Supporting Information) is larger than the signals from the ECG sensors attached to the chest (Figure 3g to i) due to the large distance between the electrodes and the heart.51 The PDMS_40NW-based ECG sensors could not even detect the ECG signals due to the large background noise. In addition, the PDMS_40NW-based ECG sensors still showed large fluctuations in its ECG signals even with the attachment of adhesive tape (Figure S16c in the Supporting Information), which indicates that its contact with the forearm skin was very poor.49 In contrast, the a4-PDMS_40NW-based sensors showed clear and stable ECG signals (Figure S16d in the Supporting Information) similar to those of the gel-based sensors (Figure S16b) with low signal noise and distinguishable P waves. The low signal noise obtained from the a4-PDMS_40NW-based ECG sensors can be attributed to its high degree of conformability with the skin and high electrical conductivity of the AgNWs embedded in it.49,

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Based on the results shown in Figure 3, the strain and ECG sensors

fabricated using a4-PDMS_40NW are more highly conformable to the human skin than those based on PDMS-40_NW. The biosignals from the highly conformable a4-PDMS_40NW-based


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sensors are significantly enhanced compared to those from the PDMS_40NW-based sensors and even from the commercially available gel electrode-based sensors. In this study, we fabricated highly conformable and transparent electrodes for application as epidermal biosensors. The experimental results revealed that the conformability of the electrodes was enhanced by increasing the adhesion strength, compliance, and viscoelasticity of the PDMS matrix. These mechanical properties of the PDMS matrix could be optimized simply by controlling the amount of Triton X added and the Tc. The modulation of these mechanical properties of a-PDMS occurred through inhibition of the crosslinking reaction by Triton X. The a-PDMS matrix also proved to be biocompatible. The adhesive and transparent electrodes fabricated in this study will significantly improve the detection sensitivity of various biosignals, such as EMG, EEG, and glucose levels. We believe that the above-mentioned distinguishing features and versatile fabrication method of the electrode we developed here will be widely adopted by academic and industrial researchers and will provide a new route toward utilizing the transparent electrode for the various epidermal sensors with highly enhanced detection sensitivity. Furthermore, these electrodes can be used as electrode materials for various wearable electronic devices, such as triboelectric nanogenerators, optoelectronic devices, transparent thinfilm heaters, and wireless antennae.

ASSOCIATED CONTENT Supporting Information. Detailed experimental methods, figures, and movies of additional experimental data are included. The following files are available free of charge.


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Detailed experimental methods; photographs of the a-PDMS films with various contents of Triton X (Figure S1); T of the a-PDMS as the function of Tc (Figure S2); photographs of a4PDMS_40 during the uniaxial tension test (Figure S3); force-displacement curves of the samples obtained from the peeling test (Figure S4); photographs of various weights suspended through adhesion to a forefinger using a4-PDMS_40 (Figure S5); aging time dependence of the mechanical properties of the a4-PDMS_40 (Figure S6); the solvent swelling ratios and gel fractions of PDMS_40 and a4-PDMS (Figure S7); FE-SEM images of PDMS_40NW and a4PDMS_40NW before and after cyclic stretching tests (Figure S8); FE-SEM images of the glass substrate after releasing the PDMS_40NW and a4-PDMS_40NW (Figure S9); OM images of the fibroblast cells grown on PDMS and a-PDMS (Figure S10); optical surface profile images of the a4-PDMS_40 and a4-PDMS_70 (Figure S11); analysis of the a4-PDMS_40NW-based strain sensors (Figure S12); photographs of the PDMS_40-based electrodes attached to the forearm without and with adhesive tape for skin impedance measurements (Figure S13); photograph of the skin after removal of the a4-PDMS_40NW (Figure S14); Schematic image of the location of each electrode during the ECG measurements. ECG signals measured with the various electrodes (Figure S15 and S16); Mechanical properties of the aged a4-PDMS_40 (Table S1). (PDF) Movie S1: Video of the 50 g weight suspended from a forefinger using a4-PDMS_40 as the adhesive right after its fabrication. (AVI) Movie S2: Video of the 50 g weight suspended from a forefinger using a4-PDMS_40 as the adhesive after a week of storing under ambient atmospheric conditions. (AVI) Movie S3: Video of the 50 g weight suspended from a forefinger using a4-PDMS_40 as the adhesive after two months of storing under ambient atmospheric conditions. (AVI) Movie S4: Video of the a3-PDMS_40NW-based strain sensors attached to the wrist during repeated bending of the wrist. (AVI) Movie S5: Video of the PDMS_40-based strain sensors attached to the wrist during repeated bending of the wrist. (AVI)


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AUTHOR INFORMATION Corresponding Author *Corresponding author’s contact information: E-mail: [email protected], Phone: +82 221235834, Fax: +82 221235834 Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript.

ACKNOWLEDGMENT This research was supported by the Basic Science Research Program through the National Research Foundation of Korea (NRF) and the Ministry of Education (grant number 2015R1D1A1A01061340). This study was also funded by the Ministry of Science, ICT & Future Planning (grant number 2018R1A2B6001390) and the Joint Program for Samsung ElectronicsYonsei University.

ABBREVIATIONS PDMS, polydimethylsiloxane; a-PDMS, adhesive PDMS; AgNWs, Ag nanowire networks; ECG, electrocardiogram; EEG, electroencephalography; EMG, electromyogram; IoT, internet of things; PEIE, polyethyleneimine ethoxylated; CNT, carbon nanotube; FE-SEM, field emission scanning electron microscopy; Tc, curing temperature; T, optical transmittance; PEG, polyethylene glycol; tan δ, loss tangent; δ , phase angle; Rs, sheet resistance; a3-PDMS_40NW, a3-PDMS_40 based electrodes; a4-PDMS_40NW, a4-PDMS_40 based electrodes;


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ToC Figure


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Highly Conformable, Transparent Electrodes for Epidermal Electronics

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