Body-attachable multifunctional electronic skins for bio-signal monitoring and therapeutic applications

Electronic skin and healthcare applications

Organ-attached e-skin systems are strategically applied to monitor significant bio-signals for personal healthcare [Figure 1]. These systems collect bio-signals from the equipped sensors and then convert them into personal medical data. This data is wirelessly transmitted to various medical services including real-time health monitoring, internet network-based telemedicine for fast disease diagnosis, and emergency medical alerts for acute diseases^[21,22]. For example, when the cardiovascular patients are in an acute heart attack, the following order is carried out in a series: (i) the e-skin systems, which have been continuously monitoring the heart rate, transmit the recorded information thus far to a nearby medical institution; (ii) Medical team checks the biosignals of heart rate, saturation of partial pressure oxygen (SpO₂) level and blood pressure and (iii) sends the medical alert to caregivers and emergency services. This process is not only limited to cardiovascular patients but is also critical to other patients requiring real-time vital sign monitoring.

Figure 1. Schematic illustration of skin-attachable multifunctional electronics for personal healthcare.

Figure 2 provides a concise overview of the essential factors for realizing multifunctional and flexible sensing devices, including e-skin components, integrated sensing devices, bio-signal sensing applications, and their medical applications. Since the components of skin-attached e-skins should be soft enough to withstand mechanical stress from the physical movement of the human body, the fabrication processes for flexible and stretchable devices have been actively developed. To apply the device to the e-skin as a component, flexible photodetectors (PDs) have been realized with polymer composite materials, organic materials, and p-n junction semiconductors^[23]. Micro-scale light-emitting diodes (µLEDs) have been manufactured to earn flexibility for integration with wearable substrates^[24]. Nevertheless, these miniaturized unit devices can only detect a single signal in a limited small surface region with low sensing reliability. To overcome this issue, the integrated circuit array has great merit in multifunctional devices with large-scale sensing, efficient power consumption, and reduced manufacturing costs^[25].

Figure 2. Developing sequence of multifunctional e-skins for biomedical applications. Figure “Photodetector” Reproduced with permission from ref^[23]. Copyright 2022, John Wiley and Sons; Figure “Flexible µLED” Reproduced with permission from ref^[24]. Copyright 2022, American Chemical Society; Figure “Sensing device array” Reproduced with permission from ref^[25]. Copyright 2019, Springer Nature; Figure “Multi-functional sensing device” Reproduced with permission from ref^[26]. Copyright 2007, Elsevier; Figure “Skin-adhesive device” Reproduced with permission from ref^[27]. Copyright 2016, Elsevier. Figure “Therapeutic application” Reproduced with permission from ref^[28]. Copyright 2020, Springer Nature; Figure “PPG sensor” Reproduced with permission from ref^[29]. Copyright 2022, AAAS. µLED: Micro-scale light-emitting diode; PPG: photoplethysmography.

For reliable bio-signal sensing, although commercial ‘smart watches’ have been commercialized with sensing features (e.g., PPG, stress score, and nighttime sleep patterns), the intrinsic gap between these devices and human skin can result in inaccurate vital levels or limited accessibility^[26,27]. Skin-adhesive devices, which can be conformally attached onto the human skin surface, have recently risen in popularity for providing more precise bio-signal sensing data, as they can send more health vital signals than conventional devices^[28]. In particular, the sensing accuracy of e-skin has been enhanced due to advancements in materials engineering and fabrication processes, enabling the application of e-skin systems with several sensors and actuators in the medical field. For example, these systems regulate therapeutic actuation in real-time such as administering medication, providing heat transfer treatments, delivering electrical stimulation, and drug delivery systems (DDSs), all accomplished by signal information from sensors adhered to the skin^[29-31]. Recently, the closed-loop systems with the microneedle-based chemical sensors and drug delivery devices have been designed for accurate drug dose control with immediate feedback. Microneedle arrays enabled the detection of the ion concentration in the interstitial fluid and the control of the DDS for the glucose level monitoring and sustenance of diabetic patients^[32,33]. Although these systems were originally designed for diabetes management, they have been applied to overcome several other illness models, such as seizures, stroke, and aneurysms, in recent years.

Piezoelectric and capacitance-based mechanical sensors

Strain sensors are considered as a crucial requisite for skin-attachable devices because the skin, the largest organ in the human body, continuously receives various forms of mechanical strains (e.g., bending, stretching, compressing, and twisting) during everyday movements. Wearable strain sensors are usually divided into three main categories based on their working principles: piezoresistive, capacitive, and piezoelectric effects [Figure 3A]^[34]. Piezoresistive type sensors comprise conductive or semiconductor layers between two electrodes and transform the applied force into a resistance signal. The resistance (R) in resistive type pressure sensor is expressed as follows:

Figure 3. Fabrication process and signal measurements of piezoelectric and capacitance-based mechanical strain sensors. (A) Schematic illustrations of mechanical sensor operating mechanisms: piezo-resistivity, capacitance, and piezoelectricity. Reprinted with permission from ref^[34]. Copyright 2020, Elsevier; (B) Fabrication method of PEDOT on a PS fiber for realizing textile sensors; (C) Optical images of linear- and zigzag-type textile strain sensors embedded in fabrics. Reproduced with permission from ref^[36]. Copyright 2017, American Chemical Society; (D) 3D images of the CSF strain sensor fabrication; (E) Photograph of the wearable sensors weaved into a stocking. The inset is a magnified image; (F) Resistance changes graph of the wearable sensor at the movements of flexing/extending, marching, jogging, jumping, and squatting-jumping. Reproduced with permission from ref^[37]. Copyright 2016, John Wiley and Sons; (G) Schematic of experimental setup for fabricating the flexible strain gauge through the spraying process. Reproduced with permission from ref^[39]. Copyright 2022, American Chemical Society; (H) Photograph of e-3DP for a planar array of soft strain sensors. Reproduced with permission from ref^[40]. Copyright 2014, John Wiley and Sons; (I) Microscopic images of the PDMS/hexane nanomesh with various weight ratios 1/40 (left) and 1/80 (right); (J) Photograph of nanomesh based-sensor (left side) and control device array (right side) attached on the face during speech of “a” (K) Strain distribution images during speech of “a” with nanomesh-based sensor (left) and control one (right). Reproduced with permission from ref^[41]. Copyright 2020, AAAS. PEDOT: Poly(3,4-ethylenedioxythiophene); PS: polyester; 3D: three-dimensional; CSF: carbonized silk fabric; e-3DP: embedded 3D printing; PDMS: polydimethylsiloxane.

where ρ is the resistivity, L is the length, and A is the contact area with the electrode. Generally, contact area/resistivity values are constant, and the length of the active layer is the only variant of resistance determined by applied force. The only difference between capacitive and piezoelectric type sensors is an active layer with dielectric and piezoelectric materials. Capacitance C of the capacitive type sensors is conformed to the capacitance of parallel plate capacitors calculated by:

where ε represents the dielectric constant, A is the electrode-active layer overlapping area, and d is the interval between two plates. With the same active layer material, interval length changes of the two electrodes only affect the capacitance difference which measures the pressure to electrical capacitance difference. Piezoelectric materials generate deformation in a certain direction by external force to electricity. Polarization of sensitive material creates the positive and negative charges to different surfaces, forming a potential difference enhanced by external force^[35].

In particular, organic materials have been extensively studied for flexible piezoresistive and capacitive sensors because of their simple fabrication process, mechanical stability, easy scalability, and applicability to versatile substrates. Figure 3B and C shows textile strain sensors enabled by polymerized conductive poly(3,4-ethylenedioxythiophene) (PEDOT)-coated polystyrene (PS) fibers^[36]. The PEDOT/PS-based strain sensor exhibited excellent identification capabilities for physical states, resulting in a ΔR/R₀ of ~20% at a knee angle difference of approximately 90°. This lightweight and ultrathin PEDOT/PS strain sensor had a superior strain gauge factor (GF) of over 0.76, easily integrating with conventional clothes. Combined with a user-interface (UI) device, the device distinguished hand gestures and translated sign language into letters.

Wang et al. demonstrated a highly stretchable and stable carbonized silk fabric (CSF)-based strain sensor with a 9.6 GF under 250% strain and 37.5 GF from 250% to 500% strain, enduring > 500% stretching strain^[37]. Figure 3D illustrates a fabrication schematic of the CSF textile, which was thermally carbonized under an Ar-rich atmosphere and encapsulated by elastic silicone (Ecoflex). This CSF textile-based strain sensor was stably integrated with clothing such as gloves and stockings, detecting various human motions [Figure 3E]. The CSF-based device monitored and specified knee joint motions such as extending/flexing, marching, jogging, jumping, and combinations of squatting/jumping [Figure 3F]. The results from the carbonated silk-based strain sensors suggest that textile devices can be easily and simply produced using other fabrics including wool, cotton, and other artificial fiber fabrics.

With the increased demand for mass production of strain sensors, several research teams have explored cost-effective and scalable production methods using the fluidic properties of organic materials, such as spraying, ink-jet printing, and 3D printing techniques^[38]. Among these methods, printing-on-substrate techniques, including 3D printing and ink-jet printing, promise advances in high resolution, sophisticated structure manufacturing, and mass production. Figure 3G displays a schematic illustration of a flexible strain gauge utilizing a 2D material-percolated thin film, realized by spraying a graphene flake-distributed 1-methylpyrrolidone (NMP) solution and instantly annealing it^[39]. The applied strain degree was evaluated by measuring the resistance change caused by the conduction pathway breakdown of graphene flakes in the thin film. The sensitivity of the graphene flake network-based device was reliable under various strain values and easily modulated by tuning the 2D material ratio in the sprayed solution. The solution spraying-on-substrate method has the potential for commercializing wearable strain sensors using other 2D materials with low-cost production and easy scalability.

Figure 3H is a process image of an embedded 3D printing (e-3DP) method on elastomeric matrices using ink-form sensing materials, enhancing the stretchability and tunability of the active material by controlling the printing path. Muth et al. printed carbon conductive ink at the interface between Ecoflex and silicone thinner to demonstrate highly conformal and extensible elastomeric electronics^[40]. The printing parameters, including nozzle size, pressure, and printing speed, were sophisticatedly optimized to realize the specific shape “hairpin”. The resistance of the printed ink increased from 11 to 60 kΩ by decreasing the cross-sectional areas of the features (0.71 to 0.066 mm²) with the printing speed increase. Furthermore, the structure and electrical properties of the conductive ink were stably maintained on the cured elastomer under stretching/bending conditions.

Wang et al. reported an ultrathin, mechanically stable, and durable nanomesh-based strain gauge enabled by an electrospinning and dipping process of polyurethane (PU) nanomesh conductors^[41]. Figure 3I indicates the fabricated nanomesh structures with different polydimethylsiloxane (PDMS)/hexane weight ratios (1/40 and 1/80) to verify the optimized condition for the nanomesh-based strain sensor. The 1/40 ratio strain sensor had the highest GF of 46.3, enabling it to map facial movements during speech onto the cheek skin surface [Figure 3J]. Figure 3K presents strain distribution mapping images of the nanomesh devices compared to a PDMS thin-film device during the speech of “a”. The nanomesh-based strain sensor in this work exhibited advantages of being lightweight, having a large detectable strain range, and being gas-permeable.

Transistor-based mechanical sensors

Organic material-based wearable mechanical sensors, despite their advantages of being ultrathin, lightweight, and having a simple fabrication process, still suffer from low electronic mobility, heat/humidity stability, and robustness in the air^[42]. To address these challenges, a promising approach involves applying inorganic material to transistor-based sensors through microfabrication processes. In particular, inorganic thin-film transistors (TFTs) have received much attention due to their cost-effectiveness, simplicity, and rapid response times^[43].

Figure 4 explains the technical development of transistor-based mechanical sensors to satisfy the mechanical and electrical properties suitable for an e-skin system. Figure 4A presents the differences between the conventional structures of metal-oxide-semiconductor field effect transistors (MOSFETs) and TFTs^[44]. In terms of operation mechanism, both devices rely on the field effect in inorganic semiconductor materials. While MOSFETs offer high electrical performance due to the single crystalline material, the device must be fabricated on a limited rigid wafer. However, TFTs can be realized on an insulating substrate at relatively low temperatures, which limits the active material phase to polycrystalline or amorphous semiconductors. Figure 4B shows the typical output curve of an n-type TFT, focusing on mobility (μ) and the on/off ratio. The mobility of the TFT is the most important factor for determining the carrier transport efficiency of the semiconductor material, directly affecting the maximum drain current (I_D). The on/off ratio is simply defined as the ratio of the maximum to the minimum I_D, which should be enhanced at a constant voltage to increase the sensitivity of TFT-based sensors.

Figure 4. TFT array fabrication process with organic/inorganic materials. (A) Comparison between the typical structures of a MOSFET and a TFT; (B) Output curves of an n-type oxide TFT. Reproduced with permission from ref^[44]. Copyright 2012, John Wiley and Sons; (C) Photograph (left) and optical microscopic images (right) of transparent and flexible NL-FET arrays; (D) The switching operation and (E) voltage gain characteristics of the NL-FET array under pristine and bending (r = 2.5 mm) conditions. Reproduced with permission from ref^[46]. Copyright 2020, American Chemical Society; (F) Photograph and pressure mapping image using CNT-based TFT array. Reproduced with permission from ref^[47]. Copyright 2018, American Chemical Society; (G) Cross-sectional illustrations of the pressure (left) and thermal (right) sensor cells using organic transistors; (H) Pressure and (I) temperature mapping images using an organic transistor-based sensor array. Reproduced with permission from ref^[48]. Copyright 2015, PNAS; (J) Schematic illustration of a transparent oxide TFT array fabricated by the LLO process; (K) The stable operation of skin-like oxide TFTs under negative bias temperature stress (left) and positive bias temperature stress (right). Reproduced with permission from ref^[50]. Copyright 2016, John Wiley and Sons. MOSFET: Metal-oxide-semiconductor field effect transistor; TFT: thin-film transistor; NL-FET: nanoline-based field effect transistor; CNT: carbon nanotube; LLO: laser lift-off.

Nanomaterials have been spotlighted for application to TFT electrodes to enhance electrical conductivity, robustness, and transparency, making them suitable for next-generation displays^[45]. As demonstrated in Figure 4C, Kim et al. reported a transparent and flexible nanoline-based field effect transistor (NL-FET) array of 254 pixels-per-inch (PPI) with high transparency (~90%), flexibility (stable at a 2.5 mm bending radius), uniformity (~90%), and high electrical performances (μ of 0.52 cm²/V·s, on/off ratio of 7.0 × 10⁶)^[46]. Figure 4D and E provides the sharp and stable switching operation of the flexible NL-FET array, maintaining its voltage gain characteristics for a complementary inverter circuit in both pristine and bent states (bending curvature radius of 2.5 mm).

To map mechanical pressure or stress changes over a large area, the common method is to create an active-matrix array and integrate TFTs into each pixel of the array. Figure 4F shows how each unit device demonstrates variations in current (I_ON) between contacted and non-contacted pixels, exhibiting the capabilities of transistor arrays in mapping pressure. Nela et al. realized a carbon nanotube (CNT)-based TFT array on a flexible polyimide (PI) substrate, characterized by rapid response time and superior sensitivity^[47]. This 4-inch CNT TFT array operated at 3 V with a pressure response time under 30 ms, surpassing the responsiveness of human skin. Moreover, the flexible CNT TFTs exhibited exceptional flexibility and robustness, maintaining their device structure and electrical properties at a bending curvature radius of 5 mm.

The TFT array is utilized as a read-out device for operating multifunctional network systems with various sensors. As illustrated in Figure 4G, Someya et al. developed an organic transistor-integrated pressure and thermal sensor array^[48]. The designed net-shaped structure with PI and poly(ethylenenaphthalate) (PEN) substrates demonstrated stable operation under uniaxial extension of 25%, adaptable for attachment to the human skin surface. Figure 4H and I shows the pressure and thermal sensor mapping images with the 12 × 12 device array that simultaneously measured the distribution of pressure and temperature. However, these organic-inorganic hybrid material-based devices still confront low electrical properties such as low output current due to charge transport efficiency, material stability and interface quality between organic-inorganic materials, directly affecting sensitivity^[49].

Although previous works have been actively conducted to develop transparent electrodes, transferring complete transparent TFTs to target substrates remains challenging for wearable electronic systems. Direct fabrication of the TFT on a flexible substrate is considered an unresolved limitation due to the essential high-temperature annealing process. The laser lift-off (LLO) technique has arisen as a novel transfer method for flexible inorganic TFTs. Lee et al. reported transparent and skin-like oxide TFT array through the LLO method [Figure 4J]^[50]. When the exfoliation layer on the transparent rigid substrate is exposed to a high-energy laser pulse, the melted layer instantly forms nanostructures, reducing the adhesion force between the device layer and the substrate. By attaching the delaminated devices onto an ultrathin PET substrate, the flexible and transparent oxide TFTs were achieved, demonstrating outstanding characteristics with mobility as high as 38.5 cm²·V^-1·s^-1, an on/off ratio of ~ 8 × 10⁷, a threshold voltage of -0.3 V, and optical transmittance of 83% in the visible light region. To confirm the usability of the skin-like oxide TFTs for practical transparent displays, several stability tests, including negative (NBTS) and positive bias temperature stress (PBTS) tests, have been conducted to perform excellent stability under direct illumination [Figure 4K].

Optoelectronics-based sensors

Optoelectronic devices, including light-emitting diodes (LEDs), laser, optic fiber, and PDs, have been spotlighted as a rising technology for biosensors and visualizing devices of bio-signals owing to their tunability, precise detection capabilities, and superior power efficiency^[51]. Figure 5 presents state-of-the-art studies on flexible LED fabrication processes for application in e-skin systems. Figure 5A displays a flexible organic LED (OLED) on a 6 μm-thick flexible substrate for a phototherapeutic wound dressing patch, emitting 5 mW/cm² (670 nm) radiation in air conditions^[52]. In addition, despite the instability of the organic material in water, the OLED demonstrated excellent reliability under a mechanical stress of 1,000 bending cycles (bending curvature radius of 1 mm) and humid conditions for an hour due to the superior passivating effect of a SiO₂-based polymer.

Figure 5. Flexible optoelectronics for soft electronics. (A) Photographs of an OLED photonic skin under skin-attached (left) and wrinkled state (right). Reproduced with permission from ref^[52]. Copyright 2020, John Wiley and Sons; (B) Optical images of a wearable photoplethysmograph sensor with QD-LED/PD (left) and real-time PPG signal graph measured by QD-stretchable LED and -conventional LED (right). Reproduced with permission from ref^[53]. Copyright 2017, American Chemical Society; (C) Photographs of a flexible PPG sensor based on PLED and OPD. Reproduced with permission from ref^[54]. Copyright 2016, AAAS; (D) Photograph of a wireless WμLED, showing various alphabets; (E) Forward voltage/irradiance changes of the WμLED during 100,000 bending operations (r = 2.5 mm) (left), and I-V curves of the WμLED with various device sizes (right). Reproduced with permission from ref^[56]. Copyright 2019, Elsevier; (F) Photograph of an RGB micro-LED display, attached onto a glass rod (r = 3 mm); (G) Current-voltage curve of the µLED when bent with r = 3 mm. Reproduced with permission from ref^[57]. Copyright 2022, Springer Nature. OLED: Organic light-emitting diode; QD: quantum dot; LED: light-emitting diode; PD: photodetector; PPG: photoplethysmography; PLED: polymer light-emitting diode; OPD: organic photodetector; WμLED: wearable micro-scale light-emitting diode; µLED: micro-scale light-emitting diode.

Recently, with the development of integrated LED/PD systems, several studies have been actively conducted to develop them as wearable PPG systems, showing high sensitivity for monitoring heart rate and SpO₂. As shown in Figure 5B, Kim et al. developed fully stretchable and foldable PPG sensors constructed with colloidal quantum dot LEDs (QLEDs), accomplishing a photocurrent/dark current ratio of 1.3 × 10⁴ at 0 V and a brightness of 1,000 cd·m^-2[53]. The QLEDs were transfer-printed onto pre-strained elastomers with a uniformly curved structure, allowing easy deformation without device breakdown under tensile/compressive strain. The developed stretchable QLEDs were stretched up to 70% strain without performance degradation and folded under a 35 μm bending radius with high brightness and luminous efficiency, implying that these thin QLEDs can be attached to curved surfaces or stretchable elastic substrates. Furthermore, according to the PPG recording tests, the developed stretchable QLEDs measured accurate PPG signals compared to the conventional rigid LEDs.

Yokota et al. reported an ultra-flexible PPG sensor based on polymer LEDs (PLEDs) and organic PDs (OPDs), demonstrating multifunctionality of simultaneous bio-signal sensing and displaying on the human skin surface [Figure 5C]^[54]. The passivation layer, which provides air and water resistance, was mainly investigated to extend the lifetime of the organic material-based optoelectronic devices. Particularly, the water vapor transmission rate (WVTR) was reduced to 10^-5 g/m² per day by multi-stacking SiON and organic layers on a thick plastic substrate (125 μm). The developed 3 µm-thick PPG sensor maintained its performance even under a bending radius of 100 µm and measured the PPG signal on the fingertip surface.

Despite novel demonstrations, PLEDs and OLEDs have faced challenges, including low efficiency, short lifetimes, and instability in heat/humidity conditions^[55]. Micro LEDs (µLEDs) offer significant advantages, such as high luminous efficacy, long lifespan, and robust endurance, under high temperatures and humid conditions. Lee et al. demonstrated a wirelessly powered wearable μLED (WμLED) array with high stability, enabled by transferring it onto a fabric coat [Figure 5D]^[56]. The 30 × 30 WμLED array was operated by receiving radio frequency (RF) electrical power, transmitted from a power supply to the wearable antenna. Figure 5E represents that the WμLED array had significant mechanical durability without device breakdown in the fatigue test with 100,000 bending motions with a 2.5 mm bending radius. The forward voltage (V_f) and irradiance (E) of the device increased by 0.76 V and decreased by only 1.33mW·mm^-2, respectively. In addition, the WμLED showed extreme stability even when exposed to high temperature/humidity (85 °C/85% relative humidity) detergent solutions and artificial sunlight. These results pave the road for fully functional outdoor environments with stable optoelectrical component-based sensors. Regarding the μLED size, the WμLED had negligible degradation in electrical properties [Figure 5E].

For achieving wireless sensing systems, light-based communication, including light fidelity (Li-Fi), was realized by flexible µLEDs with high bandwidth and high transmission rates. In Figure 5F, Hu et al. provided gallium nitride (GaN)-based µLEDs enhanced in light efficiency by adding quantum dots (QDs) and TiO₂ nanoparticles^[57]. Figure 5G is the I-V curve of the flexible µLEDs with a bending radius of 3 mm, indicating a slight change in V_f regardless of the flatness of the device. The QD-coated µLED panel demonstrated a high data transmission rate of 1.9 Gbps, which also fully operated on a flexible substrate. According to these results, flexible µLEDs among diverse optoelectronic devices are suitable for various wearable sensing applications due to their superior stability, high electrical performance, and application expandability.

Human-device interfaces

Although numerous wearable systems with various optical or electrical devices have been developed, they have all shown integration with bulky conventional chips onto a wearable substrate^[58]. In this case, the device delamination from the attached body occurs from the body movement-induced mechanical stresses, caused by the Young’s modulus difference between the bulk chip and the wearable substrate^[59]. Furthermore, when conventional flexible devices are utilized for long-term wearable sensing applications, they can lead to critical dermal problems such as skin irritation, itchiness, and inflammation due to the accumulation of skin by-products (e.g., dead cells, sebum, and sweat)^[60]. In Figure 6, we introduce various human-sensing device interfaces designed to minimize skin-related issues. Kim et al. developed a multifunctional, chip-less, and wireless e-skin system based on a surface acoustic wave (SAW) sensor with a freestanding single-crystalline piezoelectric thin film of GaN^[61]. Figure 6A compares the flexibility of the demonstrated SAW e-skin system with conventional chip-based ones. Since the SAW e-skin utilized the nanoscale GaN thin-film with an ultrathin elastomer, the device showed excellent stretchability and suitability for skin-adhesive e-skin applications with living organisms. Figure 6B presents wireless pulse measurements using the SAW e-skin strain sensors on the human wrist, highlighting their high strain sensitivity and extremely low minimum detectable strain of 0.048%. The wearable SAW sensor could monitor not only skin-applied strain but also Na⁺ ions in perspired sweat by coating the device with ion-selective membranes [Figure 6C]. The membrane-penetrated Na⁺ ions attached to the SAW sensor surface changed the resonant frequency of the acoustic wave. Therefore, the SAW device compares the changed frequency value, quantifying the Na⁺ concentration which is a biomarker for organ failures and diseases related to psychological conditions [Figure 6D].

Figure 6. Perforated human-device interfaces for e-skin system. (A) Wireless e-skin based on conventional IC chips (left) and the SAW sensor-integrated wireless e-skin (right) fabricated with GaN freestanding membranes; (B) Photograph of the SAW e-skin with a strain sensor attached onto the human wrist (left) and a strain measurement graph (right); (C and D) Optical microscopic image and schematic illustration of a GaN SAW device-based wireless ion sensor. Reproduced with permission from ref^[61]. Copyright 2022, AAAS; (E) Comparison of by-product removal effects in conventional and wearable patches. Reproduced with permission from ref^[62]. Copyright 2023, John Wiley and Sons; (F and G) Wearable e-skin with an eye pattern for enhancing the sweat removal property; (H) Sweat removal monitoring by eye-patterned e-skin and conventional one. Reproduced with permission from ref^[63]. Copyright 2023, Springer Nature; (I) Schematic illustration of the auxetic dumbbell hole pattern design for perforated e-skins; (J) 3D image of the sandwich-type e-skin fabricated with a PDMS-based adhesive layer, PI substrate layer, Au electrode, sensor semiconductor layer, and upper PI top layer; (K and L) Schematic illustrations of (I) auxetic dumbbell-patterned patch, (II) resistance-type thermometer, (III) Au-based capacitive hydration sensor, and ZnO thin-film-based (IV) strain and (V) UV sensor; (M) TEWL change after various e-skin lamination. Comparison of skin suitability of e-skin for one week. Strain sensor characteristics of the perforated e-skins and the conventional one according to skin dehydration and sweating. Reproduced with permission from ref^[64]. Copyright 2021, AAAS. IC: Integrated circuit; SAW: surface acoustic wave; 3D: three-dimensional; PDMS: polydimethylsiloxane; PI: polyimide; UV: ultraviolet; TEWL: transepidermal water loss.

Recently, wearable substrates with perforated patterns have been actively investigated to prevent dermal issues by mechanically removing skin by-products over a long-term period. Kim et al. demonstrated a wearable and biocompatible 3D porous patch combined with InGaN/GaN-based micro-PDs (μPD) as an ultraviolet (UV) monitoring sensor^[62]. The porous PDMS patch was fabricated by compressing sugar microparticles (diameter of 420~480 μm), filling them with PDMS, and removing the sugar-based mold. To verify the by-product removal property of the developed porous patch, sweat permeability tests between the 3D porous skin and the conventional one were carried out on the perspiring hand, showing rapid color changes of litmus paper within 60 minutes [Figure 6E]. Kim et al. suggested a mechanically stable e-skin to improve mechanical properties and sweat permeability^[62]. Figure 6F displays a newly designed e-skin with porous eye patterns, incorporating a resistive-type body temperature sensor. The mechanical property of the eye pattern was analyzed with finite element analysis (FEA) simulation compared to that of other patterns, presenting lower maximum stress than the auxetic kirigami pattern, about 48.7%. Unlike the conventional e-skin, the continuous sweat removal to the outside was fulfilled by numerous holes inside the eye-pattern patch [Figure 6G]. According to the results of sweat removal tests, there was no visual change in a conventional flexible patch, but in the perforated e-skin, a large area of litmus paper (45.04 %) showed a color shift due to sweat absorption [Figure 6H]. Finally, thanks to the novel porous eye-pattern substrate, temperature sensing of the developed device was stably conducted, showing a similar measuring trend to a conventional infrared (IR) thermometer during intense exercise^[63].

Yeon et al. engineered a sweat pore-inspired porous e-skin capable of effectively preventing sweat accumulation, thereby allowing the integrated inorganic sensors to accurately gather physical health information without malfunction^[64]. Figure 6I shows the design rule of auxetic dumbbell patterns for applying conventional complementary metal-oxide-semiconductor (CMOS) fabrication processes to the perforated e-skin. The size and distribution of the holes in the dumbbell patterns were constructed to those of real human sweat pores, enabling mechanical properties similar to those of human skin. The e-skin structure was configured with a PDMS-based dry adhesive layer at the bottom, a PI substrate layer, an Au electrode, a semiconductor layer for the sensor, and a sandwich-like structure as the top PI [Figure 6J]. Furthermore, the electrode and sensor were completely embedded in PI, unaffected by electrical shorts or external moisture. Figure 6K and L explains the sensor parts in detail: (I) auxetic dumbbell-patterned patch; (II) resistance-type thermometer; (III) Au-based capacitive hydration sensor; and ZnO thin-film-based (IV) strain and (V) UV sensor. The ZnO strain sensors were realized with a freestanding structure to reduce strain damping by the PDMS layer, while the ZnO UV sensor was made with the hole-patterned top PI layer to be exposed to external UV radiation. The biocompatibility and sweat-removing ability of the developed e-skin were analyzed by transepidermal water loss (TEWL) and a 1-week device lamination test, indicating that the developed device has excellent retaining capability of the normal skin condition without any allergic reaction. In addition, the skin-attached e-skin with dumbbell patterns exhibited superior monitoring ability of external humidity and strain in a skin perspiring condition. These results suggest that the perforated e-skin with multifunctional sensors is reliable for long-term bio-signal monitoring compared to the conventional wearable patches.

Electrical power sources

The ultimate objective of innovative wearable systems is to emulate real human skin, achieving imperceptibility and high comfort levels. However, the electrical wiring between the device and the external power supply is still regarded as a significant restriction for user movements^[65]. To resolve this issue, several types of energy harvesters have been reported to attain self-powered electronic systems without any external power sources such as solar cells, thermoelectric power generators, and piezoelectric nanogenerators (NGs)^[66].

In particular, triboelectric NGs (TENGs) have been spotlighted as a novel energy-harvesting technology due to their simple device structure, easy/fast fabrication processes, inexpensive materials, and high-power generation^[67]. Figure 7A shows four fundamental TENG working modes: contact-separation, contact-sliding, single-electrode, and freestanding triboelectric layer modes^[68]. The TENG, consisting of two types of materials, generates electrical power through the charge redistribution, temporarily or continuously creating a potential difference through contact and subsequent separation (contact-separation mode) or sliding motion (contact-sliding mode), respectively. Conversely, the TENG with single triboelectric material makes a potential difference with a grounded or oppositely charged object (single-electrode mode). In the freestanding triboelectric layer mode, the device generates electricity by utilizing a generated electric potential difference across the device with deformation. The TENG working mechanism combines contact electrification and electrostatic induction. The polarized triboelectric charges contribute to the voltage V_OC(x), which depends on distance x. The transferred charges Q will also influence the electric potential difference. If the device has no triboelectric charges, it acts as a typical capacitor which is affected by capacitance and charges between the two electrodes, denoted by -Q/C(x). Thus, the total voltage difference V is determined by:

Figure 7. TENG module for e-skin systems. (A) Four TENG modes: vertical contact-separation mode, lateral sliding mode, single-electrode mode, freestanding triboelectric-layer mode. Reproduced with permission from ref^[69]. Copyright 2021, Elsevier; Schematic illustration of (B) the TENG and (C) the EMG operating mechanism; (D) V_oc, and I_sc graph of the HNGM integrated with TENG and EMG. Reproduced with permission from ref^[70]. Copyright 2022, John Wiley and Sons; (E) Schematic illustration of the polymeric membrane-based energy harvester; (F) Graph of output power against external load resistance at 5 Hz. Reproduced with permission from ref^[71]. Copyright 2022, Springer Nature; (G) Schematic illustration of A-PGSS with TENG and f-SC; (H) Operation principle of the A-PGSS; (I) Scheme of the all-in-one PGSS with low-power conventional LP-SCC, LP-MCU, and BLE. Reproduced with permission from ref^[72]. Copyright 2023, Elsevier. TENG: Triboelectric nanogenerator; EMG: electromagnetic generator; HNGM: hybrid nanogenerator module; A-PGSS: all-in-one power-generating and sensing system; PGSS: power-generating and sensing system; f-SC: flexible solar cell; LP-SCC: low-power conventional signal conditioning circuit; LP-MCU: low-power microcontroller unit; BLE: Bluetooth low energy.

which is the fundamental relationship among Q_SC, C, and V_OC^[69].

Under the short circuit (SC) conditions, the transferred charges (Q_SC) are completely generated from polarized triboelectric charges where V is 0:

Gai et al. reported a self-powered wearable sweat analysis system (SWSAS) by hybridizing TENG and an electromagnetic generator (EMG), which harvested electrical energy from human motion through two hybrid NG modules (HNGMs)^[70]. The SWSAS selectively monitored biomarkers of Na⁺ and K⁺ in human sweat, wirelessly transmitting the sensing data via a Bluetooth module. By applying external force, the polytetrafluoroethylene (PTFE) film at the TENG center contacts and makes sliding friction with both sides of the nylon films, resulting in negative charge accumulation due to the difference in triboelectric polarity. On the other hand, the surface of the nylon film accumulates an equivalent number of positive charges [Figure 7B (i-iii)]. The repetitive, vertical movement of the PTFE film generates an electron flow due to the potential difference, resulting in a continuous alternating current (AC) output [Figure 7B (iv)]. Figure 7C shows the working principle of the EMG. Initially, the magnet is fixed to the upper coil, showing no current in the coil. After the downward movement of the magnet, the change in magnetic flux across the coil generates an electrical current to equalize the charge of the upper coil with that of the lower one. As the magnet moves upward, the magnetic flux crossing the coils is induced in the opposite direction, resulting in a reverse current. Figure 7D exhibits the output power for charging a 330 µF capacitor by TENG, EMG, and HNGM. The HNGM, consisting of TENG and EMG, solved the shortages of each device by integrating/engineering the high output voltage (V_oc) of the TENG and the superior output current of the EMG.

To apply the power source to e-skins, the TENG requests properties of flexibility, stretchability, and durability corresponding with human skin. As shown in Figure 7E, a polymeric membrane-based energy harvester was reported to realize highly flexible e-skin with TENG, fabricated by PEDOT: poly(styrenesulfonic acid) (PSSa)-naphthalene sulfonated PI [(PPNSP)-EMI.BF₄]^[71]. The PPNSP e-skin with a porous network continuously transported ions through nano-channels, producing strong molecular attracting-repulsing interactions with intrinsic ions. In addition, V_oc (109 V) and high output current (I_scs) (2.35 µA) were induced by doping EMI.BF₄ ionic liquid (IL) into the PPNSP skin due to active inter-exchange of ions. In the TENG application, the PPNSP-EMI.BF₄ e-skin and the PTFE film with aluminum electrodes repeatedly contacted and separated to generate a maximum output power of 33.2 mW [Figure 7F].

With the increasing requirement of wearable energy harvesters for daily use, Liu et al. demonstrated an all-in-one power-generating and sensing system (A-PGSS) based on TENG and flexible solar cell (f-SC)^[72]. Figure 7G shows a structural design of the A-PGSS with two rabbit fur-based TENGs and one amorphous Si-based SC. When the rabbit fur with good tribo-positivity and low friction resistance moves by external force, electrons flow to the opposite electrode through an external circuit, maintaining electrostatic balance between the two electrodes and generating a brief current [Figure 7H (i) and (ii)]. Negative charges accumulate on the fluorinated ethylene propylene (FEP) membrane surface, while the rabbit fur surface accumulates an equal number of positive charges. When the rabbit fur slider moves back, electrostatic equilibrium is disrupted in the fourth stage, causing positive charges on electrode-2 to flow in the opposite direction, towards electrode-1, resulting in a reverse transient current in the external circuit [Figure 7H (iii) and (iv)]. The fabricated A-PGSS produced a maximum power density of 0.072 W/m² (load resistance: 3 MΩ) at a 2.5 Hz contact frequency. Finally, the authors demonstrated the usability of A-PGSS with low-power conventional signal conditioning circuit (LP-SCC), low-power microcontroller unit (LP-MCU) and Bluetooth low energy (BLE) under daily life conditions such as daytime, nighttime and emergencies [Figure 7I].

Biodegradable sensors

E-skins are expected to be mainly used as disposable products. However, considering the environmental destruction caused by electronic waste, which is a severe issue due to the exponential development of electronic devices, the development of biodegradability is an essential factor for e-skins in healthcare applications^[73]. Figure 8A shows biodegradable materials for substrates, electrodes, semiconductors, dielectrics, encapsulants, and adhesives according to their applications in implantable and non-invasive sensors^[74]. The biodegradable properties of these materials are suited to monitor temporary phenomena such as wound healing, aftereffects, and side effects. Veeralingam et al. reported a NiSe₂-grown cellulose paper with silver paste for a low-cost and disposable sensing platform that can be applied to pH meters, breath analysis, and strain sensors^[75]. The cellulose paper-based flexible substrate was dipped in a seed solution with selenium powder, sodium borohydride, and NiCl₂ in deionized (DI) water for about one hour. The seed-coated cellulose paper was transferred to a Teflon autoclave for a hydrothermal reaction at 200 °C for 20 h. Thereafter, the NiSe₂/Cellulose paper was dried at 70 °C and integrated with an Ag electrode to realize a breath sensor [Figure 8B]. Figure 8C presents the breath sensing results for one minute. The current of the sensor increased during exhalation, whereas the device current decreased during inhalation.

Figure 8. Biodegradable materials and sensor device for soft electronics. (A) Circular diagram of biodegradable materials for realizing a health monitoring device. Reproduced with permission from ref^[74]. Copyright 2020, American Chemical Society; (B) Fabrication diagram of a NiSe₂/Cellulose paper-based breath sensor; (C) Breath pattern monitoring by the NiSe₂/Cellulose paper-based breath sensor. Reproduced with permission from ref^[75]. Copyright 2019, American Chemical Society; (D) Image of a biodegradable pressure sensor, which was fabricated by SF hydrogels and SF/PVA hybrid films; (E) Electrical properties of the biodegradable SF-based pressure sensor; (F) SF-based e-skin with 64 pixels for mapping the pressure distribution; (G) Pressure distribution using the SF-based e-skin; (H) Exploded view of a transient n-channel MOSFET array on a biodegradable Fe foil. Reproduced with permission from ref^[76]. Copyright 2022, American Chemical Society; (I) Measured dissolution kinetics of metal foils in pH 7.4 PBS at 37 °C (left). Line scan profiles of Mo thickness changes (right). Reproduced with permission from ref^[80]. Copyright 2015, John Wiley and Sons. SF: Silk fibroin; PVA: polyvinyl alcohol; MOSFET: metal-oxide-semiconductor field effect transistor; PBS: phosphate-buffered saline.

Wang et al. developed a highly strain-sensitive, fast-responding, stable, and biodegradable e-skin using silk fibroin (SF) as a structural basis, combined with a hydrogel and a flexible conductive film electrode^[76]. The biodegradable e-skin was fabricated with a SF/polyvinyl alcohol (PVA) hybrid film and a SF-based hydrogel. The SF/PVA film was manufactured with SF/PVA mixed solution using the solution casting method, and the hydrogel was obtained with horseradish peroxidase (HRP)-blended. Figure 8D illustrates the fabrication process of the SF-based pressure sensor using SF/PVA hybrid film and SF hydrogel with an enzymatic crosslinking method. The hierarchical structure of SF enabled the design of elastic hydrogels with resilience and flexible films with a porous structure with high deformability. The fabricated SF sensor could accurately sense a wide surface area of pressure due to its variable capacitance depending on pressure. Figure 8E shows excellent properties of the biocompatible and biodegradable sensor such as a high strain sensitivity of 4.78. The SF sensor also demonstrates fast responsiveness of < 0.1 s and superior durability for 20,000 cycles of 20% stretching. Figure 8F and G presents an 8 × 8 array of SF-based sensors for mapping the pressure by gathering the capacitance change of each unit sensor. The detailed vertical and horizontal pressure data was used to generate a mapping image of the pressure distribution in 3D [Figure 8G].

Despite the fast advancements in the biodegradable and mechanical properties of polymer substrates, the electrode materials have faced significant issues, such as swelling, cracking, and dissolving, due to their poor stability in the external environment^[77]. Therefore, several researchers have investigated biodegradable metal films (e.g., Ti and Al) with notable mechanical stability for application to biomedical devices as electrode materials^[78,79]. Kang et al. reported the transient electronics with biodegradable metal foils (Fe, Mo, W, and Zn)^[80]. Figure 8H displays an exploded view of a transient n-MOSFET array with Mg electrodes, SiO₂ gate dielectric, and Si nanomembrane active layer on a Fe foil. Figure 8I shows the dissolution kinetics of various biodegradable metal foils in phosphate-buffered saline (PBS) solution with a pH of 7.4. The changes of each film thickness in PBS were 0.02 (Mo), 0.08 (Fe), 0.15 (W), and 3.5 (Zn) µm/day, respectively. The Mo foil particularly showed a uniform thickness decrease in PBS, as shown in the right graph of Figure 8I. Because of the uniform removal of the Mo foil in the PBS solution, the fabricated MOSFET exhibited stable operation for about 5 h.

Vital monitoring devices

Thin film-based flexible sensors have opened the possibility of vital sign monitoring of skin, organs, and body movements due to their conformal attachability to body organs^[81-83]. For detecting biological problems, several research groups have demonstrated wearable sensors for analyzing various homeostatic biomarkers such as heartbeat, sweat ion concentration, tears, and saliva^[84-88]. Kang et al. reported a low-power heterojunction phototransistor to monitor the PPG signal [Figure 9A]^[89]. The PPG sensor was fabricated using the poly{3-[(2,2’:5’,2’’-terthiophen)-5-yl]-2,5-bis(2-octyldodecyl)-2,5-dihydropyrrolo[3,4-c]pyrrole-1,4-dione-6,5’’-diyl} (DPP2ODT2-T) layer with chlorobenzene (CB) and toluene (Tol), accomplishing a stable voltage level and low power consumption (5 V) for accurate sensing under various heart rates. The superior electrical/optical properties of the sensor provide long-term and accurate sensing capabilities under different heart rate conditions in cardiovascular checks.

Figure 9. Healthcare monitoring wearable devices with various bio-signals. (A) Image of a skin-conformable PPG sensor and its heart rate monitoring. Reproduced with permission from ref^[89]. Copyright 2021, Elsevier; (B) Sweat sensor system for monitoring ions (left) and real-time monitoring results of K⁺ and Na⁺ in the perspiration (right). Reproduced with permission from ref^[95]. Copyright 2020, Elsevier; (C) Photograph of the Ag-DS/CF sensor under finger bending strain; (D) Resistance changes of Ag-DS/CF in bending and releasing cycles. Reproduced with permission from ref^[102]. Copyright 2018, Springer Nature; (E) Optical image of theranostic smart contact lenses (scale bar, 5.5 mm); (F) I-t curve of the flexible DDS with different three drug containers. (G) IOP monitoring with timolol release from Day 1 to Day 5. Reproduced with permission from ref^[108]. Copyright 2022, Springer Nature; (H) Photograph of the wireless saliva sensor; (I and J) Various solution responsivity of wireless saliva sensor. Reproduced with permission from ref^[109]. Copyright 2018, John Wiley and Sons; (K) Schematic illustration of a skin-attachable, stretchable TSFET device with a suspended gate and thermochromic display; (L) Absorbance spectra of the thermochromic leuco dye film at 25 and 45 °C. Reproduced with permission from ref^[111]. Copyright 2018, Elsevier; (M) Motion detecting characteristics of the ADAM device on the dorsal hand, compared to conventional devices. Reproduced with permission from ref^[112]. Copyright 2022, AAAS. PPG: Photoplethysmogram; Ag-DS/CF: Ag-dragon skin with conductive fiber; DDS: drug delivery system; IOP: intraocular pressure; TSFET: temperature sensor based on field effect transistor; ADAM: advanced acousto-mechanic.

Invasive blood tests for analyzing ion concentrations, such as Na⁺, K⁺, and Cl^- or glucose, are applied in various fields (electrolyte analysis, pH tests, and diabetes), but they have significant issues such as skin trouble, wounds, and pain^[90-92]. The potentiometric sensors measure the potential difference between a reference electrode and a working electrode to track the change of the ion concentrations against time. With their simplicity, diverse ion detection capabilities and rapid response times, potentiometers have been spotlighted in diverse fields, especially in ion sensing of body fluid^[93,94]. Sweat analysis with potentiometric sensors has been considered as an invasive vital monitoring method; however, several obstacles remain, such as limitations of sweat collection, materials toxicity, and low multi-sensing capabilities.

Criscuolo et al. developed a multifunctional sweat analyzing system, integrating ion-selective membranes and cotton microfluidics for stable and continuous monitoring [Figure 9B]^[95]. The system consisted of four ion-selective electrodes for each ion: Na⁺, Li⁺ (therapeutic drug), Pb²⁺ (heavy metals exposure), and K⁺ (hydration and activity). The multifunctional sweat sensors were fabricated using conventional micro-electro-mechanical systems (MEMS) processes, demonstrating stable sensing properties in water and artificial sweat, even under mechanical stress (180° bending). To prove real-time monitoring applications, the sodium and potassium levels were tracked by the wearable system, showing high Pearson correlation coefficients of about 0.97 (Na⁺) and 0.81 (K⁺) during physical exercise of five human volunteers. These results mean that the developed wearable sweat sensing system has the potential for the non-invasive vital sign monitoring technology.

Since various mechanical oscillations/movements (e.g., voices, heartbeats, and tremors) occur in our body, monitoring them is important to check bio-stability. To distinguish small bodily oscillations, high sensitivity and signal separation properties of the devices are required. Although several researchers have reported the wearable strain sensor with high sensitivity over 1 kPa^-1 and GF of ~10³ under 2%~6% strain, these strain sensors have various issues including high-cost/complex manufacturing processes, and low stability under harsh environmental conditions^[96-101]. Tolvanen et al. developed a stretchable and washable crack sensor using silver ink-patterned silicone elastomer [Ag-dragon skin with conductive fiber (Ag-DS/CF)] [Figure 9C]^[102]. This strain sensor was fabricated in the pre-stretching state to withstand the dynamic loading-induced strain by forming the wrinkle of Ag-ink. Depending on the intensity of the applied strain, the change of the device sensitivity was inversely proportional, providing action detectivity regardless of its size. The fabricated strain sensor exhibited a high GF of 1.2 × 10⁵ and negligible hysteresis loss under the repeated stretching cycles. This crack-based strain sensor with the maximum pressure sensitivity of 0.82 kPa^-1 at 0.035 kPa demonstrated the notable sensing properties under body movement and steel ruler oscillation [Figure 9D].

Intraocular pressure (IOP), the pressure within the eye, must be consistently monitored to prevent significant ocular diseases such as glaucoma and ocular hypertension^[103-105]. Smart contact lenses with DDSs have been reported for IOP response but faced some issues such as low sensitivity, biocompatibility, and stability for long-term IOP monitoring^[106,107]. As displayed in Figure 9E, Kim et al. developed a theranostic smart contact lens composed of a sensitive gold hollow nanowire-based IOP sensor, a flexible DDS, wireless power/communication systems, and an integrated circuit chip for controlling IOP to treat glaucoma^[108]. The system achieved superior IOP sensitivity and stability with high transparency of 84% for long-term usable lens. Furthermore, the flexible DDS system had biocompatibility and high drug loading efficiency (~85%), enabling practical IOP control. Figure 9F shows the timolol releasing properties from the smart lens, which is the medicine for controlling IOP. The smart lens was inserted at the eye of a glaucoma-infected rabbit, measuring the IOP and treating high IOP state via timolol release [Figure 9G]. The developed smart lens showed the potential for expandability of wearable biomedical sensors.

Tseng et al. reported a wireless saliva sensor consisting of a silk film and a poly(N-isopropylacrylamide) (PNIPAM) responsive hydrogel^[109]. This wearable saliva sensor was attached to tooth enamel, enabling the determination of the kind of food being eaten by monitoring changes of the resonant frequency [Figure 9H]. The PNIPAM hydrogel was responsive to a wide range of fluid properties, including alcohol content, salinity, sugars, pH, and temperature. The fabricated sensor distinguished the materials by comparing RF change under DI water, artificial saliva, 50% alcohol, methanol, and high salinity saliva [Figure 9I]. Compared to DI water, artificial saliva revealed a lower amplitude (and slightly lower resonant frequency) owing to its higher ionic strength, while alcohol caused an increase of the resonant frequency due to a lower net permittivity. Four kinds of foods were eaten by the subject, showing different resonance frequencies [Figure 9J]. The reported wearable saliva sensor opened the opportunity to provide immediate treatment to patients who have ingested hazardous food via its real-time pH sensing characteristics.

Among the five senses, visual information is the most powerful and memorable due to its high storability of ~78%^[110]. Therefore, numerous researchers have paid attention to the visually vital sign info-communication device with wearable patches. Hong et al. demonstrated a temperature sensor based on FET (TSFET) with visual communication via leuco dye^[111]. The TSFET utilized a suspended gate using PNIPAM, which changed the drain current as airgap decreased with increasing temperature, as shown in the upper right image of Figure 9K. Leuco, with an easy coating process and low cost, displayed various colors according to its structural change, resulting in the device sensitivity of 6.5%/°C in the temperature range between 25 and 45 °C. Furthermore, the thermochromic leuco dye showed a transmittance peak shift to the long wavelength under due to increased temperature [Figure 9L]. This suggested temperature sensor was an innovative approach demonstrating that the electrical signals can be replaced with visual signals.

Recently, strain sensors with high sensitivity, stability and robustness have been reported; detecting specific movements, such as finger, wrist and fine vibrations of skin, remained challenging. Chun et al. presented a skin-conformable wireless mechanical sensor to profile various hand and specific body movements using advanced acousto-mechanic (ADAM) sensors^[112]. The ADAM sensor was placed on the dorsum of the hand and captured acousto-mechanic signals from ambient noises. Figure 9M compares the obtained scratching movements by moving the arm and fingers. The comparison was conducted with two modes of scratching: one by the arm and the other by only the fingers. The blue line represented the time series of the raw data, and the red line indicated the final binary classification results, where 0 and 1 correspond to non-scratch and scratch, respectively. To apply the ADAM system to practical applications, an algorithm-based study was performed, and the ADAM system showed an accuracy of 99.0% against visual observation. This work highlighted the potential wide-ranging applications from assessing the efficacy of drugs for conditions causing itchiness to monitoring disease severity and treatment response.

Biomedical therapeutic devices

In recent decades, many researchers have developed e-skin-based treatment methods to address the medical inconveniences such as invasive treatments and in-hospital care. Innovative e-skin technologies using opto-devices, electric fields (EFs), and microneedles have attracted attention due to their non-invasiveness, convenience, and low side effects. Figure 10 explains biomedical treatment devices integrated into the e-skin systems. Since alopecia is a common disease with a wide age range of onset, treatment methods have been developed such as minoxidil (MNX), finasteride, and surgical hair transplants, but these approaches have low efficiency and side effects^[113-115]. Lee et al. reported a wearable photostimulator for hair growth applications through flexible red vertical µLEDs (f-VLEDs) with high output power (~30 mW/mm²) and low V_f bias (~2.8 V)^[116]. Figure 10A compares the hair growth rate of rat groups between f-VLED and MNX treatment. After 20 days of treatment, the hair length of the photostimulated mouse was 183.2 μm, approximately 76.8% longer than that of the MNX group (103.6 μm). Figure 10B compares stained tissues and hair follicles under each treatment condition to confirm the correlation between red light and hair follicle cell proliferation. Mice in the experimental group demonstrated local hair regrowth in light-irradiated regions, while control groups exhibited negligible changes.

Figure 10. Medical actuators for therapeutic applications. (A) Graphical figuration of trichogenic photostimulation with monolithic red f-VLEDs and mouse dorsal skin comparison about hair-regrowth effect (left). Comparison of the hair-grown skin area, the hair-growing speed, and the grown hair length after various skin stimulations (center). *P < 0.05, paired t test, #P < 0.05, two way ANOVA; **P < 0.01, ***P < 0.001, paired t test (center), *P < 0.05, paired t test (right); (B) Immunofluorescence analysis of the stimulated mouse dorsal skin. Reproduced with permission from ref^[116]. Copyright 2018, American Chemical Society; (C) Schematic illustration of an SID and a wearable EEG monitoring device with a wearable power transfer unit. The inset shows drug releasing mechanism of the SID with an activation signal from wireless power transfer; (D) Model drug releasing image of the SID which is attached onto porcine skin. Reproduced with permission from ref^[117]. Copyright 2021, AAAS; (E) Schematic image of an AgNW-MAA ink-based ePatch for realizing wound healing through pulsed EF; (F) Biological activities of the healing process with the ePatch attachment; (G) Stained live cell distribution with the ePatch (left) and cell grown with EF groups after applying EF for 24 h (right). Cells are aligned in lines indicated by the yellow arrows. Reproduced with permission from ref^[118]. Copyright 2022, Elsevier; (H) Schematic image of photostimulation for melanogenesis inhibition via the SµLED patch; (I) Diagram of photodamaged skin thickness, each treated with control, SµLED, and CLED. **P = 0.01, and ****P = 0.0001; (J) H&E-stained images of mouse skin after photostimulations in 20 days. Reproduced with permission from ref^[119]. Copyright 2022, John Wiley and Sons. f-VLEDs: Flexible red vertical micro-scale light-emitting diodes; SID: soft implantable drug delivery device; EEG: electroencephalography; AgNW-MAA: silver nanowire-methacrylated alginate; EF: electric field; SµLED: surface-lighting micro-light-emitting diode; H&E: hematoxylin and eosin.

Recently, accurate drug dosing methods have been developed using the implantable DDSs with a wearable structure for reducing the drug abuse risk. Figure 10C shows the RF-based wireless DDS for monitoring and treating epileptic seizures^[117]. The soft implantable drug delivery device (SID) consisted of a drug reservoir (parylene), epoxy to prevent drug leakage, and platinum pads for electrolysis electrodes. The drug delivery method followed these steps: (1) the electroencephalography (EEG) sensor detected the abnormal signal caused by an epileptic seizure; (2) the external RF sensor transmitted the electrical signal; (3) water in the SID was decomposed via electrochemical reaction; and (4) the bubbles increased the internal pressure, inducing the SID drug release [Figure 10C]. Figure 10D exhibits the drug releasing results by utilizing the implanted SID in pig skin tissue. The EEG and epileptic seizure signals were measured by the algorithm, and the model drug (Evans blue) was successfully released through the measured signals.

Since secondary infection from existing wounds can lead to various problems such as high fever, edema, and convulsion, several ointment-type drugs have been developed. However, these drugs must be applied multiple times according to the severity of the wound. To solve this issue, a wound treatment patch based on EF stimulation has been developed by Wang et al.^[118]. A flexible wound healing electrical patch (ePatch) applying the EF to the wound region was realized using conductive hydrogel electrodes [Figure 10E]. As shown in Figure 10F, the wound healing was assisted by cell migration and proliferation, induced by the EF. Figure 10G presents the comparative images of the fibroblast alignment with and without the EF. When the EF was not applied, the fibroblast migration and alignment were not observed. In contrast, when the EF was applied, both migration and alignment were evident.

In addition to treating diseases, phototherapeutic devices have recently gained popularity as skincare tools. However, most optical skincare devices for home-care use have been made with bulky LEDs, providing non-uniform light irradiation and beauty care effects. Lee et al. accomplished a surface-lighting micro-LED (SµLED)-based wearable patch for uniform melanogenesis inhibition [Figure 10H]^[119]. The SµLEDs with 630 nm of wavelength were adequate to prevent melanogenesis on the skin and were monolithically fabricated on a plastic substrate. Moreover, the SµLED exhibited outstanding mechanical stability under 10⁵ bending cycles at a 5 mm bending radius and endured a harsh environment within 100 days under 85 °C/85% relative humidity. According to the melanogenesis inhibition tests, the SµLED-based patch efficiently reduced the melanin cells without any skin photodamage [Figure 10 (i) and (j)]. These results confirmed that wearable patches with multifunctionality by integrating various electrical components could non-invasively treat various diseases and be applied in the beauty care field.