Protein-Based Electronic Skin Akin to Biological Tissues - ACS Nano

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Protein-Based Electronic Skin Akin to Biological Tissues Minsik Jo, Kyungtaek Min, Biswajit Roy, Sookyoung Kim, Sangmin Lee, Ji-Yong Park, and Sunghwan Kim ACS Nano, Just Accepted Manuscript • DOI: 10.1021/acsnano.8b01435 • Publication Date (Web): 24 May 2018 Downloaded from http://pubs.acs.org on May 24, 2018

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is published by the American Chemical Society. 1155 Sixteenth Street N.W., Washington, DC 20036 Published by American Chemical Society. Copyright © American Chemical Society. However, no copyright claim is made to original U.S. Government works, or works produced by employees of any Commonwealth realm Crown government in the course of their duties.

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Protein-Based Electronic Skin Akin to Biological Tissues Minsik Jo,†,‡ Kyungtaek Min,†,§,‡ Biswajit Roy,† Sookyoung Kim,† Sangmin Lee,# Ji-Yong Park,†,# and Sunghwan Kim†,#,* †

Department of Energy Systems Research, Ajou University, Suwon 16499, Republic of Korea

§

Department of Nano-Optical Engineering, Korea Polytechnic University, Siheung 15073,

Republic of Korea #

Department of Physics, Ajou University, Suwon 16499, Republic of Korea

*

To whom correspondence should be addressed. E-mail: [email protected]

‡

These authors contributed equally to this work.

KEYWORDS: electronic skin, silk protein, hydrogel electronic device, stretchable, water permeability

ACS Paragon Plus Environment

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ABSTRACT

Human skin provides an interface that transduces external stimuli into electrical signals for communication with the brain. There has been considerable effort to produce soft, flexible, and stretchable electronic skin (E-skin) devices. However, common polymers cannot imitate human skin perfectly due to their poor biocompatibility, biofunctionality, and permeability to many chemicals and biomolecules. Herein, we report on highly flexible, stretchable, conformal, molecule-permeable, and skin-adhering E-skins that combine a metallic nanowire (NW) network and silk protein hydrogel. The silk protein hydrogels offer high stretchability and stability under hydration through the addition of Ca2+ ions and glycerol. The NW electrodes exhibit stable operation when subjected to large deformations and hydration. Meanwhile, the hydrogel window provides water and biomolecules to the electrodes (communication between the environment and the electrode). These favorable characteristics allow the E-skin to be capable of sensing strain, electrochemical, and electrophysiological signals.


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Electronic skin (E-skin), an artificial mimic of human skin and capable of transducing external stimuli such as pressure, heat, and humidity to electrical signals, has been the subject of intense study for decades with the ultimate goal of enabling applications such as human/computer interfaces, continuous health monitoring, and sensory skin for robotics.1-5 To obtain reliable signals from the skin or other biological tissues subjected to large deformation under mechanical loading, E-skin devices should ideally be capable of being conformally attached to threedimensionally curved tissue surfaces.6,7 The most common approach to realizing E-skin devices involves the embedding of conductive fillers or rigid electrical devices into an elastomeric matrix formed from the likes of polydimethylsiloxane (PDMS) and Ecoflex elastomer. Through the integration of micro- or nanostructures into the elastomeric matrix by means of lithographic, molding, and/or printing methods, various revolutionary multifunctional and high-sensitivity Eskin devices have been realized.3-5,7,8 Further, the engineered synthetic elastomeric matrices, for better permeability and skin-adhesion, have been adapted to epidermal electronics on human skin.9-11 Nevertheless, it is still challenging to develop materials that is biocompatible, adhering well to biological tissues, and having high permeability to water, which has many essential functions in the human body, including the transport of nutrients and enabling chemical and metabolic reactions. Hence, the fabrication of stretchable and conformal E-skins by using protein hydrogels, which are biocompatible polymer networks filled with water or an aqueous solution, can provide a seamless interface between the human and E-skins.12-15 In the food and medical industries, protein-based biopolymers are attracting interest as a green alternative to petroleum-based polymers due to their biocompatibility, processing versatility, and low cost.16,17 Among proteins, silk fibroin protein has been identified as being particularly suitable for biopolymer applications due to its mechanical durability, tunable


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secondary structure, all-aqueous processing, and in vivo biocompatibility.18 The excellent oxygen permeability of silk films in the wet state, which is of a level similar to that of human skin, also points to their potential for application to artificial skin systems.19 Additionally, over the last decade, silk has been the subject of high-technology reinvention in the fields of optics and electronics. Electronic and optical micro/nanodevices such as transistors, electrodes, and photonic crystals have been successfully integrated onto silk films to realize the functions of the devices (controlling the behaviors of two quanta: photons and electrons) in a biological context.20-26 However, pure silk films tend to be stiff and brittle in the dry state. Although the blending of suitable plasticizers like glycerol is one means of making silk films flexible and water-insoluble, their lack of stretchability still hinders their E-skin applications.27,28 Recently, Ling et al. reported on silk films that had been made stretchable through the addition of Ca2+ ions, but these films proved unstable in humid environments and lost their optical transparency, both when immersed in water and when subjected to stretching.29 In the present study, we successfully fabricated skin-compatible E-skin that is a stretchable, conforming, molecule-permeable, and skin-adhering electronic device fabricated using silk fibroin protein and a conductive silver nanowire filler (AgNW). The incorporation of glycerol (inducer of the β-sheets of silk molecules) and Ca2+ ions (bonder of silk molecules via chelation and charge interaction) into the silk matrix simultaneously offers a useful route to generate polymorphically regenerated protein membranes with the desirable physical properties for the E-skin application, and this enables the silk E-skin to offer high stretchability (up to 400%) and on-skin adhesion in a living environment. A nanowire network, of an arbitrary pattern, was buried in the stretchable silk membrane, allowing it to be applied as an electrode for electrochemical sensing, electrocardiograms (ECGs), radio-frequency (RF) antennas, and the


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integration of opto-electronic devices. Even when subjected to mechanical strain and hydration, the electrodes were able to operate stably. Sensing electrodes fabricated in this way proved capable of being activated by analytes/water that permeated through the protein membrane, indicating that the silk E-skin devices communicate with their environments through the silk skin in the same way as real human skin. Our protein-based E-skin devices can be used for the continuous monitoring of vital signals under normal, everyday living conditions.

RESULTS AND DISCUSSION The preparation of stretchable silk E-skin is illustrated in Figure 1a. Briefly, an adhesive tape layer with patterns was attached to a silicone-coated polyethylene terephthalate (PET) film to allow the better release of the silk E-skin. An aqueous NW solution was poured into the patterns and then the adhesive tape layer was peeled off, leaving the NW network. The device was completed by releasing the silk film with the patterned NW network after the assembly of the silk fibroin, Ca2+ ions, and glycerol. The resulting silk film was optically transparent and highly stretchable when subjected to mechanical strain. To mimic human skin, high conformability, on-skin adhesion, and permeability are highly desirable features. Figure 1b shows the silk E-skin attached to a human wrist before and after wrist-bending to demonstrate the high conformability and on-skin adhesion of the device. In addition, when the human skin was wrinkled, the silk membrane accurately followed the irregular surface of the skin, whereas the PDMS membrane became detached from the skin (Figure S1). The skin-adhesion of our silk membrane could be estimated quantitatively by measuring peel-forces for detachment of the silk film from the pig skin, the mimic of the human skin (Figure S2).30 The obtained peel-forces (~ 10 N/m) were comparable to those of commercial


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medical-grade adhesives31 and enhanced as increasing the humidity. Moisture on an arbitrary surface could sustainably hydrate the silk membrane and infiltrate through the network of silk molecules (which is not possible with PDMS), such that the buried or attached electronic devices would be able to communicate with the analytes in the body fluid (Figure 1c). Figures 1d and 1e show light emitting diodes (LEDs) that are integrated into the AgNW network in the silk. All the blue, green, and red LEDs, which were connected in series, successfully turned on by an electric current flowing through the AgNW network even when attached to moist tissue subject to stretching. The mechanical properties of the stretchable silk membrane can be tuned by adjusting the Ca2+ ion and glycerol contents. We measured the stress–strain relationship of the silk membrane to find the optimum combination of silk protein, Ca2+ ions, and glycerol (Figure 2a). Figure 2b shows the typical stress–strain curves for the silk membranes with different silk/Ca2+ weight ratios when the silk solution was mixed with glycerol at a fixed weight ratio of 10:6 (silk:glycerol). The Ca2+ ions in the silk membrane bonded to the silk molecules via chelation and charge interaction, and captured water molecules from the air. Therefore, as the number of Ca2+ ions in the silk membrane increased, more water molecules could be captured.29 This resulted in a much softer and stretchable membrane. The addition of Ca2+ ions reduces the Young’s modulus from 69.5 kPa to 10.5 kPa when the silk/CaCl2 weight ratio was changed from 83:17 to 75:25, as shown in Figure 2b. At a weight ratio of 67:33, the resulting silk membrane was mechanically too weak and soft (like a soft gel cream) to be used in an E-skin application. Interestingly, the obtained Young’s moduli are much lower than those of both PDMS and the human epidermis (140 to 600 kPa),32,33 indicating that our silk E-skin devices would have less of a feeling of “tightness” as the wearer’s skin naturally deformed. The glycerol content also


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influences the mechanical properties of the silk membrane since glycerol can be hydrogenbonded to the peptide matrix and then induce β-sheet formation.27 As the glycerol content increased, the Young’s modulus decreased (Figure 2c). Note that the stretchability could be tuned by the humidity condition as shown in Figure S3. Additionally, we could observe a hysteresis behavior in the stress-strain curve that indicated the viscoelasticity of the silk membrane (Figure S4). This provides a useful information to estimate the fatigue lifetime of the silk E-skin device as the absorbed energy for a cycle is obtained.34 In the present study, those silk membranes with a silk:CaCl2 weight ratio of 75:25 and silk:glycerol weight ratio of 10:6 were used for E-skin applications. Unlike the as-cast silk membrane prepared from a pure silk solution, the redesigned stretchable silk membranes were highly stable and robust even after soaking in water. Fouriertransform infrared (FTIR) spectroscopy was performed to investigate their molecular structures. Figure 2d shows that our silk membranes contained extensive β-sheets, a secondary proteinbased structure, indicating that silk molecules were crosslinked and stable in solvents.35 The βsheet formation in the stretchable silk was induced by the addition of glycerol,27,28 whereas the random coil was the dominant protein structure in the self-assembled silk. In addition, there was no significant difference between the spectra before and after stretching. This proved that the unstretched silk membrane already had the maximum β-sheet content and therefore the high stretchability originated from the interplay between the highly crosslinked silk molecules and the captured water molecules.29 As shown in Figure 2e, the silk films exhibited good optical transparency (75% at 550 nm for the unstretched film and 90% for the stretched film) over a broad spectral range. The transmission increases as the thickness of the silk film decreases as a result of the elongation of the film (thus obeying the Beer-Lambert law T = exp(− ), where α


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is the absorption coefficient and t is the thickness of the silk film). In addition, we could obtain a Poisson’s ratio (ν) = 0.33, which is comparable to those of rubbery materials, using the equation

εz = - νεx, where εz (εx) is the strain along the direction normal to the film surface (elongation). The percolated network of metal NWs provides a continuous electron transport pathway with high conductivity and good mechanical flexibility.36-40 Especially, the infiltrated silk protein can strongly anchor the NWs and bury the NW network, thus forming mechanically robust and flexible electrodes in the silk protein (Figure S5). To understand the electrical properties of the NW electrodes in the stretchable silk membranes, we generated rectangular electrodes of different widths (3, 5, 7 mm) but a fixed length (10 mm) using AgNWs, which had an average diameter of 40 nm and an average length of 20 µm. Figure 3a shows the measured resistances as a function of the strain. At zero strain, as shown in Figure S6, the AgNW/silk electrodes, regardless of their width, all exhibited a similar electrical resistivity ρ of around 3 × 10-6 Ω·m (200 times greater than that of bulk Ag but comparable to that of AgNW/elastomer electrodes),41

as obtained from Pouillet’s law = , where R is the resistance, L is the length, and A is the cross-sectional area of the electrode. This indicates that the topology of the percolating AgNW network is homogeneous. Under strain, the number of junctions maintaining contact in the stretched AgNW network fell, and the distance between the NWs increased, thereby increasing the value of ρ.36 Hence, the value of R for all the electrodes increased significantly as a result of the strain caused by the increased ρ and L, as well as the decreased A. Additionally, we obtained the tunable gauge factors, defined as

/

, where Δ/ is the fraction change in the resistance

and ε is the strain, over a range of 1 to 15, which were larger than those of conventional strain sensors. As such, we can confidently state that the membrane could be used for strain sensor applications.33,42,43 The mechanical robustness of the AgNW networks buried in the silk film can


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also expand the applicability of our approach. Our previous study showed that the AgNW/silk electrodes was very stable under cyclic bending tests.40 As shown in Figure S7, cyclic loading and unloading tests exhibited that resistances of the AgNW/silk electrodes were increased as increasing the number of cycles. Nevertheless, high electrical conductivity was maintained during cycles of strains up to 20%. To induce communication at the biological–electrical interface and prevent tissue damage, the on-skin (or implanted) devices must permit the transmission of water vapor, oxygen, and other chemical compounds. These properties are often overlooked for bioelectronic devices. Accordingly, as shown in Figure 3b, we investigated the water vapor transmission rate (WVTR) of three types of membranes: silk (660 µm thickness), silk incorporating AgNW (700 µm thickness), and PDMS (630 µm thickness). All the WVTRs were measured at 23 °C/40% RH. The water was heated on a hotplate at 50 °C. The bulk silk and AgNW-buried silk membranes were found to have WVTRs ranging from a high of 4046 g·m-2·d-1 to a low of 1720 g·m-2·d-1. These values are comparable to those of common bandages and wound dressings.44 On the other hand, the WVTR of a PDMS membrane was negligible, which would be detrimental to epidermal and biological tissues over prolonged usage. The electrode also exhibited a stable resistance when smeared with water and saline, indicating that it would operate reliably on wet skin and tissue, even in humid environments (Figure 3c). Even if the wettability of the electrode surface worsened due to the presence of the buried AgNWs (Figure S8), the droplets of water and saline could fully infiltrate the membrane in a few minutes. Additionally, we examined the diffusion properties of the silk hydrogel by using rhodamine B dye solution (Figure S9). Based on a simplification of Fick’s law, diffusion coefficients of 9.7 × 10-3 mm2‧s-1 for the silk membrane and 7.1 × 10-3 mm2‧s-1 for the silk membrane incorporating AgNW, both of which are


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comparable to those of porous silica gel,45 were obtained. These results indicate that the silk hydrogel membrane provides a reliable and speedy hydrogel window between the E-skin device and the environment. Using the skin-like traits of the silk protein membrane, the buried electrode and device can communicate with the environment via the transportation of molecules through the silk hydrogel window (Figure 3d). As a proof-of-concept experiment (Figure 3e), we investigated the resistance between two separated line-electrodes when a small amount of water was dropped onto the opposite surface of the hydrogel window.12,46,47 Upon reaching the electrodes, the infiltrated water induces a decrease in the resistance, such that water electrolysis can start. Due to the high surface-area-to-volume ratio of the AgNW network, the bubbling of the electrolysis rapidly damages the AgNW electrode such that the resistance level is restored. Meanwhile, the bulk electrodes exhibited only a low level of electrolysis (Figure S10). And the same electrolysis reaction occurred when a small amount of saline was dropped instead of water (Figure S11). It should be noted that the small amount of ions in the permeated liquid (water or saline) had an even weak influence on the electrical conductivity of the AgNW electrode, as already shown in Figure 3c. This concept could be applied to the electrochemical recording of dopamine (DA) levels,

which

is

an

electroactive

and

critical

neurotransmitter

enabling

neuronal

communication.48-50 We generated a silk insulating gap (around 500 µm in width) between the electrodes such that the electrical current could not flow in the dry state. As shown in Figures 3f and 3g, a droplet of DA aqueous solution infiltrated the silk insulating gap, making it conductive. The current flow increased with the DA concentration, indicating that this configuration could be used as an electrochemical sensor. It should be noted, however, that even pure water in the silk gap would induce a current due to the presence of ions. In this case, however, the current


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continuously decayed, while there was no current flow once the device had stabilized. A stable current flowed when the gap was bridged by the infiltrated DA. Our electrode is capable of conformal and intimate contact with human skin, making it a viable

alternative

to

conventional

Ag/AgCl

wet

electrodes

for

the

recording

of

electrophysiological signals. Although Ag/AgCl wet electrodes are of high quality and suitable for short-term clinical use, they cannot be used for long-term wearable applications due to need for an electrolytic gel layer, which irritates the skin and causes signal degradation.51 By adding a metal cap to the patterned AgNW network and performing silk-curing, a wearable silk protein electrode for measuring ECG signals was produced (Figure 4a). Figure 4b shows the ECG signals captured with a commercial Ag/AgCl electrode and the silk/AgNW electrode, both being applied to a forearm. Although the ECG signal from the silk/AgNW electrode is slightly noisier and weaker, each wave (P, QRS complex, and T) is clearly defined and the absence of a wandering baseline shows that our electrode is well-attached to the skin, as shown in Figure 4c. The noise can be reduced by adding a post-process filter to the Arduino-based signal processing circuit used here. For long-term and daily use, the capture of ECG signals must be possible despite increasing degrees of movement. The conformally attached silk/AgNW electrode could record ECG signals even when attached to the sweaty skin of a subject who had been running, producing clearly visible P, QRS complex, and T signals (Figure 4b). The silk hydrogel film could be better adhered to the hydrated or moistened skin (Figure S2), which led to reduction of the skin-silk interface impedance. Therefore, the noise of the ECG signal from the sweaty skin after exercise was reduced. Meanwhile, the commercial electrode lost its adhesion (Figure S12). A device that it incapable of conformal contact cannot acquire a clear ECG signal on highly curvilinear skin surfaces (e.g. the skin covering protruding bones). Figure 4d shows


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electrodes attached to the bump of the wrist. We attempted to apply a commercial electrode to the skin, but it was inevitable that some part of the electrode was released from the skin. This led to there being a higher signal-to-noise ratio and a weaker voltage of the ECG signal (Figure 4e). The silk/AgNW electrode perfectly conformed to the protrusion and therefore was able to capture a relatively clear ECG signal. These results show the remarkable potential of the silk/AgNW electrode as a skin-compatible, skin-conforming, and skin-adhering electrode for long-term daily health monitoring. A useful component that could be incorporated into a functional E-skin device is an RF antenna for wireless communications. For telemedicine applications, mechanically deformable RF antennas are incorporated into wearable systems.52 Using our method, a stretchable loop RF antenna was successfully fabricated (Figure 5a). At zero strain, the fabricated RF antenna exhibited resonance at the frequency of 3.6 GHz (corresponding to a half-wavelength) with a return loss (S11) of -16.1 dB, indicating that more than 97.3% power could be transmitted at the resonant frequency, as shown in Figure 5b. As the applied strain increased, S11 increased due to the deterioration in the electrical conductivity of the AgNW network, which was accompanied by a shift in the resonant frequency. Additionally, at large strains of over 100%, resonant behavior was successfully obtained from the stretched RF antenna (Figure 5c).


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CONCLUSION In summary, we have successfully fabricated stretchable, conformal, and waterpermeable E-skins by using silk protein and a conductive nanomesh. The NW electrodes buried in the protein membranes exhibited high and stable conductivity despite the application of mechanical forces and the infiltration of water/saline in the environments in which E-skin devices are applied, and could be designed to sense strain, electrochemical, and electrophysiological signals by recording the changes in the electrical conductivity. Like real skin, the stretchable silk hydrogel membranes provided a communicating channel to interface with vital biological signals from the soft tissues, while also having other bio-friendly traits. Our approach has opened up the possibility of implementing skin-like E-skin devices in important applications such as the monitoring of human health and environmental conditions, and the treatment and prevention of diseases.


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METHODS Silk fibroin solution preparation. The Bombyx mori cocoon fibers were boiled in 0.02-M Na2CO3 solution for 30 min to remove all sericin protein. The remaining fibroin fibers were rinsed with distilled water > 3 times and then dried in air for 24 h. To prepare silk solutions with the required silk:Ca2+:glycerol ratios for fabricating the stretchable silk films, the degummed silk fibers were dissolved in CaCl2/glycerol/formic acid solution. In a typical experiment, 1 g of CaCl2 (Sigma-Aldrich, C1016–100G) was first dissolved into 20 g of formic acid (SigmaAldrich, F0507–1L), and then 3 g of fibroin fibers were also dissolved into the solution. After these processes, glycerol (Sigma-Aldrich, A16205) was added at a ratio of 60% by weight to make the silk films highly stable in water. Fabrication process. To generate a patterned AgNW electrode in a silk film, a stencil mask of an arbitrary pattern was prepared using commercial adhesive tape and then attached to a siliconecoated PET substrate (Silicone release film, i-ONE FILM). The patterns in the tape mask were generated using a medical scalpel and razor blade. Next, a AgNW aqueous solution (Nanopyxis Co., Ltd., 0.3 wt%, I27D-KNS6KL) was poured onto the PET film with the patterned mask and then dried at 50 °C on a hotplate. After releasing the adhesive tape mask to yield the patterned AgNW on the PET substrate, the prepared silk solution was poured onto the PET substrate with the AgNW pattern and then dried under ambient conditions for > 24 h. To connect the silk E-skin to other electronics, fine copper wires (Nilaco, CU-111267) were attached to the buried AgNWs using Ag paste. Red light-emitting diodes (LEDs, WELED, 1000001407), green LEDs (WELED, 1000001408), and blue LEDs (WELED, 1000001410) were applied to the silk E-skin, again using Ag paste.


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Tensile strain measurement. Stretchable silk films measuring 20 mm (W) × 30 mm (L) × 0.5 mm (H) were prepared. The films were suspended vertically relative to the ground. Then, force was applied by hanging weights from the films, and the stretched length was measured. Optical characterization of silk films. Fourier-transform infrared spectroscopy (FTIR, Thermo Scientific, Nicolet iS50) was used to analyze the molecular structure of the chemically crosslinked silk. To measure the transmission, white light from a halogen lamp was used to illuminate the silk films. The transmitted light was captured by the tip of a multimode fiber with a 200-µm core diameter placed on the opposite side of the film and sent to a spectrometer (Ocean Optics, USB-2000). A reference signal was collected using an aluminum mirror. The silk film was mounted in a holder and stretched while the transmission was being measured. Electrocardiography. A AgNW circular pattern with a 2-cm diameter was generated on the PET substrate, and then a metal cap, taken from a commercial ECG electrode (3M, Monitoring Electrode) was superposed on the AgNW pattern using Ag paste. ECG signals were obtained from the electrodes which were connected to the ECG amplifier of an Arduino Uno system and an ECG sensor (Single Lead Heart Rate Monitor - AD8232, SparkFun). Three prepared electrodes were attached to the skin: the negative electrode to the right wrist, the positive electrode to the left wrist, and the ground electrode to the right thigh. Radio frequency antenna. A loop antenna was designed with a resonant frequency of 3.5 GHz, corresponding to a half-wavelength. To connect the RF antenna to the measurement set-up, a copper wire was indium-soldered to the AgNW pattern in the silk film, and then the other end of the copper wire was connected to the SMA connector. Using a network analyzer (Anritsu, 37247D), the electromagnetic reflection of the AgNW/silk RF antenna was analyzed by observing the impedance matching. Prior to the measurements, the analyzer was calibrated to


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minimize errors in the reflection measurements, including the directivity, source matching, and frequency tracking, while ensuring a better frequency response from the RF antenna.


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Figure 1. Skin-adhering, stretchable, conforming silk protein-based hydrogel electronic skin (Eskin). (a) Fabrication process and working principle of the skin-like hydrogel E-skin. To make the silk film stretchable and stable under hydration, Ca2+ ions and glycerol are added to the silk aqueous solution. The solidified silk film with the silver nanowire (AgNW) electrode can act as an E-skin device under mechanical strain and hydration. (b) Silk E-skin adhered to a wrist. No detachment is observed under bending/unbending of the wrist. (c) Polydimethylsiloxane (PDMS) and silk membranes. A drop of water is trapped between the PDMS membrane and the substrate due to the low water permeability, whereas the water is absorbed by the silk membrane and subsequently evaporates. (d,e) LED chips integrated into the AgNW electrode are turned on when the silk E-skin is attached to biological tissue, even when the E-skin is stretched.


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Figure 2. Structural and mechanical characterizations of stretchable silk hydrogel membranes. (a) Silk membranes before/after application of mechanical tension. (b,c) Stress–strain plots of silk membranes at various silk:Ca2+ and silk:glycerol weight ratios, at a constant humidity (50%). (d) Fourier-transform infrared (FTIR) spectra of silk membranes from 1750 to 1450 cm-1. This region includes two bands, corresponding to β-sheet (1630 cm-1) and random coil (1655 cm-1). (e) Optical transmission spectra of the stretched silk membrane. Elongation induces contraction in the vertical direction (inset), thereby making the membrane more transparent.


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Figure 3. Electronic characterization of permeable silk E-skin. (a) Resistance of rectangular AgNW electrodes of various widths buried in a silk film under tensile strain. Tensile strain induces an increase in the length of the electrode and therefore increases the resistance. (b) Measurements of water vapor transmission rates (WVTRs) for PDMS membranes (black), silk membranes (red), and silk membranes incorporating AgNW (blue). WVTRs were measured at 23 °C/40% RH. (c) Short-circuit current of a rectangular electrode as a function of time. No change in the current is observed when a drop of water (at 50 s) and saline (at 1600 s) infiltrated to the AgNW/silk layer. (d) Schematic of permeable silk E-skin device. (e) Short-circuit resistance between the two buried electrodes as a function of time. The permeated water bridges the gap between the electrodes and reduces the resistance. After racking the electrode by electrolysis reaction, the resistance is recovered. (f) Schematic of dopamine (DA) sensor using two disconnected electrodes. DA molecules combined with permeated water induce current flow. (g) Electrical current responses upon addition of DA solutions with increasing concentrations as a function of time.


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Figure 4. Wearable AgNW/silk electrodes for electrocardiograms (ECGs). (a) Schematic diagram of fabrication of AgNW/silk electrode. (b) ECG recording with commercial Ag/AgCl and AgNW/silk electrodes. In the case of the AgNW/silk electrode, ECGs were captured before and after the subject exercised. The records show that the heart rate increased after exercising. (c) Enlarged ECG signal traces, comparing P, Q, R, S, and T waves obtained by commercial Ag/AgCl and AgNW/silk electrodes. (d) Schematics and photos of conformal attachment of AgNW/silk electrode. (e) ECG signals captured by electrodes attached to bump of the wrist. The non-conforming nature of the commercial electrode induces a large amount of noise in the signal.


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Figure 5. Stretchable radio-frequency (RF) antenna. (a) Loop RF antenna before (top) and after (bottom) stretching. (b,c) Return loss (S11) of antenna for different strains. For (b) small and (c) large strains, the resonant frequency increases with the strain.


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ASSOCIATED CONTENT Supporting Information. Adhesion-to-skin performance (S1); Adhesion properties of a silk hydrogel on skin (S2); Mechanical properties of a silk hydrogel film at different humidity conditions (S3); Viscoelasticity of a silk hydrogel membrane (S4); Morphology of silk film incorporating AgNW (S5); Electromechanical response of AgNW/silk nanocomposite (S6); Mechanical durability of AgNW/silk electrodes under cyclic strains (S7); Wettability of stretchable silk hydrogel membrane (S8); Diffusion characteristics of permeable silk hydrogel membranes (S9); Waterpermeable silk hydrogel membrane (S10); Electrolysis reaction by the permeated saline (S11); Peel-off adhesion of commercial ECG electrode (S12).

Conflict of Interest. The authors declare no competing financial interest.

AUTHOR INFORMATION Corresponding Author *E-mail: [email protected] Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. ‡These authors contributed equally to this work.


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ACKNOWLEDGMENT The authors acknowledge support from the National Research Foundation (NRF) of Korea (no. 2017R1A2B4010807), the GRRC program of Gyeonggi province (GRRC-AJOU-2016-B01, Photonics-Medical Convergence Technology Research Center), and the Korea Institute of Energy Technology Evaluation and Planning (no. 20164030201380, Human Resources Program in Energy Technology).


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Table of Contents artwork


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