Advances in Ultrathin Soft Sensors, Integrated Materials, and Manufacturing Technologies for Enhanced Monitoring of Human Physiological Signals

Recent advances in soft sensors and flexible electronics offer various applications in detecting physical, electrical, and chemical signals. However, there are still technical barriers in current mechanical, electrical, and material properties for enhanced signal sensing. When measuring signals from the human skin, minimizing the skin‐sensor contact impedance is still challenging while maximizing sensitivity through optimized materials and soft electronics. Here, this review summarizes recent advances in materials, manufacturing, and integration technologies to develop ultrathin soft sensors for monitoring various human physiological signals. The enhancements in soft and compliant structures and mechanical properties are critical to making reliable wearable electronic systems. This article shares the details of soft sensors, integration processes, manufacturing methods, and their applications to target physical, electrical, and chemical signals. In addition, the limitations and current trends in developing multifunctional sensors, self‐powered devices, and integration with external stimuli systems are discussed.


Introduction
Recently studies about monitoring human physiological signals have attracted substantial research interest because of their potential for application in human-machine interface, [1] disease diagnosis, [2] patient monitoring in clinic, telemedicine, and mental state monitoring. [3] However, the sensors that are devices [8] since the extremely thin and flexible form factor of sensors provides skin conformability and user's comfort. [8a,9] In human physiological signal monitoring, the sensors with various thicknesses from the nanometer scale [10] to the millimeter scale [11] are called ultrathin sensors. The thickness standard is not precise. However, the ultrathin dimension of sensor increases the tolerance to mechanical deformations due to peeling from substrate or physical strain. [9,12] In addition, the ultrathin properties allow the sensor to be attached to the skin via van der Waals forces without adhesives that can attenuate the signal and cause irritation. [8a,9,13] Furthermore, by combining the open mesh, serpentine layouts, and neutral mechanical plane construction, ultrathin sensors can minimize local strain even under large system-level deformation. [9] This allows conformal skin attached sensors, and prevents from interfering with the user's movement involving both the wrinkling and tensile strains. [14] Over the past several years, considerable effort has been put into developing ultrathin soft sensors for the enhanced monitoring of human physiological signals, and interesting developments have been achieved in sensor structure, integrated materials, and manufacturing technologies.
In this review paper, we summarized recent advances in ultrathin soft sensors by classifying the types of sensors into three types according to the types of human physiological signals to be monitored, as shown in Figure 1: 1) physical signals, 2) electrophysiological signals, and 3) chemical signals. In Section 2, we introduced general sensing mechanisms and sensor structure for each type of ultrathin soft sensor. Then, we reviewed various strategies to increase sensing performance while maintaining ultrathin properties. Interestingly, porous structures have been applied to ultrathin soft sensors regardless of sensor types by introducing microstructure or mesh-structure. After that, integrated materials and manufacturing technologies were summarized in Sections 3 and 4, respectively. We noticed that common materials and manufacturing technologies were applied to fabricate the ultrathin soft sensor, despite the differences in the signals they measure, the working principles, and the sensor structure. Lastly, in Section 5, we discussed the future direction of ultrathin soft sensors for enhancing human physiological signal monitoring in terms of integration of wireless system, multifunctional, selfpowered sensors, etc.

Ultrathin Soft Sensors
As mentioned above, the ultrathin soft sensors can be classified into three types: 1) sensors of monitoring physical signals, 2) sensors for monitoring electrophysiological signal, and 3) sensors for monitoring chemical signals. First, the physical signal sensors can detect the deformation of physical features such as mechanical change and temperature change and convert them into the change of electrical characteristics such capacitance, resistance, and potential. Another type of sensors and electrode can directly detect changes in electrical potential and conductivity. The last type of sensor provides the change of electrical characteristics by transducing the concentration changes of molecule and/or ion. In this section, we organized the basic working mechanisms, general structure, and strategies for improving the performance while reducing thickness. Ultrathin soft sensors were developed in the direction of reducing the number of layers of multilayered sensors and the sensitivity of the sensor could be improved by forming a porous structure. In this process, gas permeability (or breathability), which is important for long-term use of the sensor, could also be improved.

Physical Signal Monitoring
Human physiological and physical signals indicate mechanical, kinematic, and/or temperature signals such as muscle movement, pulse, breathing, vocal cord vibration, and skin temperature. These physical signals can be transduced to the change of electrical features. Depending on the principle of transducing, ultrathin soft sensors for monitoring physical signals can be categorized in detail as capacitive, resistive, piezoelectric, and triboelectric sensors (Figure 2). Generally, the ultrathin soft sensors consist of two conductive layers: functional layers and substrate. To improve the performance while maintaining the ultrathin structure, porous structure has been applied to each layer regardless of measurement principles by introducing microstructure and using nanofiber membrane. Furthermore, new strategies that can further reduce the number of layers are being applied such as using interdigitated electrode (IDE).

Capacitive Sensors
A capacitive sensor is one whose capacitance changes in response to a physical change, such as pressure or strain. Most conventional pressure sensors are capacitive, due to the lower probability of mechanical failure and higher stability under practical working conditions. [15] Capacitor sensors generally consist of two conductive plates sandwiching a dielectric layer (Figure 2a). The main mechanism underlying the sensor response is the capacitance fluctuation where different levels of physical deflection of the sensor result in a corresponding capacitance change. The capacitance can be calculated by C = ε 0 ε r A/d, where ε 0 is the vacuum permittivity, ε r is the relative dielectric permittivity, A is the overlapping area of the two conductive layers, and d is how far between the two conductive layers. [8c,16,17] Table 1 summarizes the materials and characteristics of ultrathin soft capacitive sensors for monitoring physical signals.
For the capacitive sensor with high sensitivity, both the conductive and dielectric layers are required to be as flexible as possible so that the dielectric layer and electrodes can be deformed upon slight pressure stimulation. Sensor with this structure enables significant change in capacitance, [16] and ultrathin architecture with great flexibility is an ideal candidate for the highly sensitive capacitive sensor. [18] Howe et al. [19] developed capacitive ring-type flow-sensor that was composed of ultrathin metal layers (300 nm thick nitinol (NiTi) or magnesium (Mg) or 100 nm thick gold (Au)) and an ultrathin dielectric layer (1.4 µm thick polyimide (PI)) using lithography and etching technologies. The ultrathin flow-sensor was placed at the center of the flow diverter and the compressive force from fluid flow against the sensor results in a compressed PI dielectric layer, allowing for capacitive response upon varying intra-aneurysmal flow ( Figure 2a). This ultrathin soft sensor could detect mean vessel velocities as small as 0.032 m s −1 , which shows a very high sensitivity to monitor the alteration of the intra-aneurysmal flow. Herbert et al. [20] also introduced a fully passive wireless, low-profile capacitive sensor (total thickness: 100 µm), and coil inductor for monitoring hemodynamics using aerosol jet printing (AJP). This ultrathin soft sensor consists of multilayers with silver (Ag; 4.1 µm thick), PI (2.7 µm thick), and soft elastomer and the sensitivity of this sensor was 0.29 pF m −1 s. However, this mechanism, based on the Poisson effect, theoretically limits the sensitivity to around a gauge factor 1. [21] In order to overcome this limitation, Herbert et al. proposed ultrathin soft capacitive sensors with employing a sliding mechanism where overlapping plates slide relative to each other (Figure 2b). [22] The overlapping plates slide in opposing directions when the sensor is strained. The sliding of plates decreases the overlapping area and causes a decrease in capacitance.
Capacitive sensors used on unstructured flat elastomeric dielectrics with small moduli have limited compressibility and strong viscoelasticity, resulting in low sensitivity, slow reaction time, and relaxation time. [23] Utilizing microstructures such as pyramids, [24] and microconvex [25] as both conductive and dielectric layers is a popular method in order to improve the sensitivity of conductive sensors. [16] The porous structures increase the electric volume by introducing air gaps between conductive layers, which in turn reduces Young's modulus of the dielectric layer and maximized the capacitance change induced by physical stimulus. [15] However, a very pricy and challenging photolithography process is needed to create microstructured dielectric. [17,26] Instead of the conventional microfabrication method, a simple and cost-effective manufacturing www.advelectronicmat.de technologies such as self-assembly, and printing were applied to fabricate microstructure to the ultrathin capacitive sensors. Herbert et al. [27] enhanced the pressure sensitivity of ultrathin soft capacitive sensor for hemodynamic monitoring by add a polydimethylsiloxane (PDMS) microstructure layer (total thickness: 28 µm). Moreover, due to the thin, flexible layers, and patterned PDMS, the fully printed capacitive sensors with microstructured features significantly improve pressure Reproduced with permission. [16] Copyright 2020, Elsevier. b) Illustration of an implantable device and aerosol jet printed strain sensor with its schematic under strain. Reproduced with permission. [22] Copyright 2022, Elsevier. c) Schematics of the piezoresistive sensor patch working mechanisms. Reproduced with permission. [35] Copyright 2022, American Chemical Society. d) Exploded view of a conformable facial code extrapolation sensor (cFaCES). Reproduced with permission. [45] Copyright 2020, Springer Nature. e) Schematics of the working mechanism of an (i) IDE-based and (ii) MIMbased piezoelectric device during the poling process. Reproduced under the terms of Creative Commons Attribution (CC-BY). [8b] Copyright 2021, The Authors, Published by American Chemical Society. f) Schematic diagram of bio-degradable triboelectric nanogenerator and its working principle. Reproduced with permission. [50b] Copyright 2017, Wiley-VCH GmbH. g) Schematic structure of the all-fiber triboelectric nanogenerator-based electronic skin and its working principle. Reproduced under the term of CC-BY license. [55] Copyright 2022, The Authors, published by MDPI.
www.advelectronicmat.de sensing during bending. Similarly, Xiong et al. [16] developed an ultrathin (8.2 µm) and ultra-highly sensitive (30.2 kPa −1 ) capacitive sensor integrated by two PDMS-Au conductive layers with convex microarrays and an ultrathin polyvinylidene fluoride (PVDF) dielectric layer (Figure 2a). The use of the IDE structure decreases sensor thickness and simplifies the fabrication process because only one conductive layer is required. [8b] When external pressure strain was applied to this sensor, the thickness of the IDE will decrease, and the area of IDE will expand outward and increase due to the Poisson effect. The deformation of IDE causes change of the relative dielectric constant of the dielectric layer. Qiu et al. [7] proposed a capacitive sensor composed of silk hydrogel film, silver nanowire-based IDE (20 nm), and polyurethane (PU) film as dielectric layer, conductive layer, and encapsulating layer, respectively. Although the IDE itself had the pressure sensing function, the author could fabricate higher sensitive sensors through ultrathin silk hydrogel film (60 µm) since it has the self-patterned microstructure on the surface and the microstructure can increase the amount of deformation.

Resistive Sensors
A resistive sensor is a sensor that detects the change in resistance of sensitive materials and converts it into an electrical signal output. [8c] Conventional resistive sensors consist of functional layer sandwiched by substrate layer and electrode ( Figure 2c). If the functional layer is piezoresistive, the resistive sensor can monitor mechanical change. If thermoresistive material was used as functional layer, the resistive sensor could monitor the temperature change. Total resistance of ultrathin resistive soft sensor is sum of the electrode, resistive material, and material/electrode contact resistance. [29] For example, if the applied pressure increase, the resistance of the piezoresistive material is changed. At the same time, the contact areas between the resistive material and the electrode are also increase for recued contact resistance upon further increase pressure. Table 2 summarizes the materials and characteristics of ultrathin soft resistive sensors for monitoring physical signals.
Herbert et al. [6] introduces printed, nanostructured strain sensors based on the direct patterning of nanowires and integration with soft materials, with applications in structural health monitoring and human physiology detection. AJP is employed for patterning miniaturized strain sensors and highly stretchable strain sensors. The thickness of this ultrathin sensor was 8.5 µm and highly stretchable strain sensor shows high wearability and unobtrusively detects human motion, pulse, and breathing with a high sensitivity comparable to a clinical-grade system. Meanwhile, Sang et al. introduced an ultrathin (3.5 µm) and ultrahigh sensitive (−37270.72 ppm °C −1 ) Au-doped silicon nanomembrane temperature sensor assay. [30] They reduced strain applied to active array by transferring an ultrathin layer of Au-doped silicon nanomembrane to a PI substrate and pattering the entire device, including interconnections, into a Polyimide; b) Silver nanoparticle; c) Poly(styrene-isoprene-styrene); d) Polydimethylsiloxane; e) Polyvinylidene fluoride; f) Poly(methyl methacrylate); g) Poly(3,4-ethylenedioxythiophene) polystyrene sulfonate; h) poly(vinylidene fluoride-trifluoroethylene); i) Silver Nanowire; j) Polyacrylamide; k) Polyvinyl alcohol. www.advelectronicmat.de serpentine mesh. As a result, they were able to monitor skin temperature using sensors with a superhigh precision and making a conformal contact to the skin.
Introducing porous structure to enhance the sensing performance was equally utilized in the ultrathin resistive sensor. Cai et al. developed an ultrathin strain sensor (<2 µm) by layering MXene and carbon nanotube (CNT). [31] Islands and gaps induced by stretching decreased conductive pathways, which is responsible for the increasing of the overall electrical resistance of the MXene/CNT film. The water-based assembled sandwichlike microstructure provides the ultrathin devices with a low limit of detection as small as 0.1%, high sensitivity (a gauge factors up to 772.60), and tunable sensing range (30% to 130% strain). The extraordinary sensing performances enabled successful detecting of both tiny deformations such as phonations and substantial movements such as walking, running, and jumping. Besides, Jeong et al. also developed a flexible pressure sensor with an ultrathin thickness (31.3 µm) and wide Activated carbon fiber; b) Gauge factor; c) Laser-induced graphene; d) 3,4-ethylene dioxythiophene; e) Polystyrene-block-poly(ethylene-ran-butylene)-block-polystyrene; f) Carbon nanotube; g) Gold nanowire; h) Multiwalled carbon nanotube; i) Reduced graphene oxide; j) Single walled nanotube; k) Polyethylene terephthalate; l) Printed circuit board. www.advelectronicmat.de pressure sensing range (10-500 kPa) by microstructuring the surface of the conductive film into a pyramid shape. [32] Wu et al. [33] fabricated an ultrathin (<40 µm) aramid nanofibers film-based flexible sensor by combining inkjet printing method and compressible buckled microstructures. The flexible sensor possesses sensitive subtle-pressure perception (38.4 Pa) as well as rapid response and recovery speed (20 ms). For configuring porous structures between functional materials and electrode, various methods of combining a porous substrate such as paper or fabric with a resistive material was utilized. Sun et al. fabricated a highly stretchable (up to 300 strain) and ultrathin (≈100 m) epidermal strain sensor from a flexible multiwalled CNT (MWCNT) nanopaper and PDMS resin using an ultrasonication resin impregnation process. [34] The flexible MWCNT nanopaper was fabricated using a vacuum filtration method and the PDMS/MWCNT nanopaper composite with homogeneously dispersed MWCNTs at high loading (>10 wt%) is able to provide both excellent electrical and superior mechanical properties.
Meanwhile, Zhao et al. fabricate a textile-based piezoresistive sensor using gold nanowire-impregnated fabric as the functional layer (Figure 2c). [35] In the pressure-sensing tests, the gold nanowires (AuNWs)/textile wearable sensor shows not only long-term stability and ultrahigh sensitivity (i.e., 914.970 kPa −1 in the range of 0-0.1 kPa) but also a fast response and recovery time (<40 ms) and a low detection limit (i.e., 0.49 Pa). Although the thickness of the sensor was relatively thick at 1 mm, the AuNWs/textile wearable sensor was shown by wearing the sensor on a human body or sportswear to monitor the physical parameters of breathing, pulse, heart rate, and human movements during the activities. Similarly, Li et al. fabricated pressure sensor using a conductive cotton fibers modified by reduced graphene oxide nanosheets. [36] The conductive cotton fibers can be easily prepared through a simple dipping and annealing process. Furthermore, Liu et al. demonstrated a facile and novel approach for fabricating all-textile-based pressure sensors and large-area sensor arrays. [37] The resistive textile sensor unit is formed with CNT-coated cotton fabric on a top bridge and interdigitated textile electrode on a bottom bridge. The intertwined conductive electrodes on textile materials (polyester, nylon, etc.) were fabricated by laser-scribing masking and electroless deposition of conformal nickel (Ni) coatings. The resulting sensor on textile substrate showed high sensitivity (14.4 kPa −1 for a pressure range below 3.5 kPa, 7.8 kPa −1 for a pressure range of 3.5-15 kPa), stable cycle performances (1000 cycles), a fast response time (≈24 ms), and a low detection limit (2 Pa). Moreover, the fabricated textile sensor could be attached to human skin to detect various forces, vibrations, and monitor real-time pulse waves with low power consumption (<6 µW). These ultrathin textile-based pressure sensors are perfect for monitoring human physiological signal because they can easily be incorporated into clothing without significantly sacrificing its comfort.
The IDE structure was also applied to the ultrathin resistive sensor to decrease sensor thickness and simplify the fabrication process. Zhan et al. [38] have proposed a novel strategy to fabricate a flexible and wearable pressure sensor by impregnating a single-walled CNT (SWCNT) into tissue paper (SWCNT/tissue paper) and sandwiching them between a bare PDMS sheet and a PI sheet patterned with interdigitated Au electrodes. A typical sensitivity of 2.2 kPa −1 in a wide range of 35-2500 Pa and a sensitivity of 1.3 kPa −1 in the range of 2500-11 700 Pa could be achieved, which is comparable with the record of organic transistor pressure sensors reported recently. Notably, the active elements, SWCNT/tissue paper, could be easily fabricated at low cost and a large scale. Yang et al. [29] designed and demonstrated a flexible paper-based pressure sensing platform that features the MXene-coated tissue paper (MTP) sandwiched between an encapsulation layer and a printing paper with interdigital electrodes. Because of the highly porous 3D structure of the MTP and high specific surface area of MXene, the MTP pressure sensor encapsulated by the 35 µm-thick PI film exhibits an ultrahigh sensitivity of 509.5 kPa −1 , a low limit (1 Pa), and a broad range (100 kPa) of detection. More importantly, replacing the PI film with the weighing paper further results in a recyclable paper-based MTP pressure sensor with a sensitivity of 344.0 kPa −1 over the same sensing range of up to 100 kPa.
Although ultrathin sensors with complex structures such as multilayer and microstructure showed high performance, the complex structures usually require more expensive and complex fabrication steps, which limit the scalability and reproducibility. [39] Xi et al. reported an ultrathin microtubular resistive sensor for pulse monitoring unlike general multilayer structures consisting of substrates, conductive layers, and resistive layers. [40] This microtubular sensor was 120 µm thickness (approximately the cross-section of a strand of hair) and could detect external mechanical forces based on the change of cross-sectional area. The electrical resistance of liquid metal increases when the sensor is compressed, and the sensitivity of the sensor was 68 N −1 .

Piezoelectric Sensors
Various ultrathin capacitive sensors and resistive sensors have been developed for the enhanced monitoring of physical signals with high sensitivity. However, these two types of sensors have a critical disadvantage that needs an external power source limiting the use of sensors in the various fields such as wearable monitoring system and in vivo physiological signals. [41] Table 3 summarizes the materials and characteristics of ultrathin soft piezoelectric sensors for monitoring physical signals. Piezoelectric effect indicates a phenomenon in which mechanical stimuli deform some anisotropic crystalline materials, polarize internal dipoles, and produce potential differences between the crystals' two opposing surfaces. [8c] These piezoelectric materials are more favorable than resistive materials due to their low power, [42] high sensitivity and easy readout features. [43] Piezoelectric sensors generally consist of two electrodes sandwiching a piezoelectric material and substrate: electrode-piezoelectric layer-electrode (EPE) structure ( Figure 2d). Dagdeviren et al. [43] reported an ultrathin piezoelectric sensors (thickness: 77.4 µm) for monitoring physiological signals and ingestion within the gastrointestinal tract using a lead zirconate titanate (PZT) sandwiched by two parallel metal plate (Au and platinum (Pt)). Using same materials, Lü et al. [44] fabricated piezoelectric sensors for monitoring eye fatigue. The entire thickness was 10 µm so that the sensor can maintain conformal contact with www.advelectronicmat.de the eyelid skin during eye blinking. In addition, Sun et al. [45] introduced a conformable facial code extrapolation sensor (thickness: 46.9 µm) based on aluminum nitride (AlN) this film, which is low cost, [46] suitable for mass manufacturing and leadfree nature. Furthermore, Ha et al. [47] showed that an ultrathin and skin-soft sensor for monitoring seismocardiography could be developed using commercially available metalized PVDF sheets without microfabrication progress such as lithography. The total thickness including piezoelectric layer, electrode, and substrate was 128 µm, which was relatively thicker than other sensors, but they demonstrated that an ultrathin sensor could be made quickly (20 min) and at low cost through commercial materials and cost-effective cut-and-paste method.
As mentioned above, utilizing an IDE structure (Figure 2e) decrease the sensor thickness compared to the EPE structure as only electrode layer is required, and it also improves fabrication yield by preventing the shorting of electrodes with different polarities. [8b] Moreover, compared to a similar thickness EPE structure, the use of an IDE structure in the piezoelectric sensors increases sensitivity because of its reduced capacitance. [41] Park et al. developed a self-powered ultrathin piezoelectric pulse sensor (6.8 µm) based on PZT thin film, Au-IDE, and ultrathin plastic. The inorganic piezoelectric sensor on an ultrathin plastic makes conformal contact with the complex texture of the rugged skin, allowing it to respond to tiny pulse changes (sensitivity: ≈ 0.018 kPa −1 ) on the epidermis' surface. [41] Likewise, Montero et al. [8b] and Laurila et al. [48] developed a fully printed ultrathin piezoelectric electronic tattoo (e-tattoo) pulse sensors with the thickness of 7 and 4.2 µm, respectively. The sensitivity of two different sensors was 275.8 mV N −1 and 1703 pC N −1 , respectively. The ultrathin form factor, due to the IDE structure, enabled these sensors to bend during the pulse measurement leading to higher device sensitivity. [48]

Triboelectric Sensors
The triboelectric effect is a type of contact-induced electrification in which a substance becomes electrically charged after coming into contact with another material, and the sign of the charges carried by a material is determined by its relative polarity to its counterpart. [23,49] Generally, triboelectric sensors are composed of two triboelectric layers, electrodes, and substrate. As shown in Figure 2f, when two materials with different electron affinities come into contact with each other under external force, they will obtain surface charges of opposite polarities based on the triboelectric effect. [50] During the separation process, the corresponding back electrodes induce the flow of electrons under the effect of electrostatic induction. [51] Triboelectric sensors have the advantages of low cost, simple structure, easy to access, diverse material option, and high conversion efficiency. [52] In addition, triboelectric sensors have potential in both sensing and wearable power supply, making it a promising candidate for the enhanced monitoring of human physiological signals. When driven by a low-pressure stimulus, electrostatic self-powered pressure sensors, including triboelectric sensors and electret/piezoelectric sensors, can generate higher voltages than piezoelectric self-powered pressure sensors. This is because the former can hold rich megascopic electric dipoles/monopoles and are easily deformed under a low-pressure stimulus, resulting in excellent mechanicalelectrical conversion efficiency. [53] As in other types of sensors, recent ultrathin triboelectric sensors are developing in the direction of applying a porous structure to improve performance and using a textile-based substrate that is easy to manufacture. Furthermore, triboelectric sensors composed of only single polarity are being developed as a strategy to reduce the entire thickness. Table 4 summarizes the materials and characteristics of ultrathin soft triboelectric sensors for monitoring physical signals.
Introducing microstructure was also used to enhance the performance of triboelectric sensors. Ouyang et al. [50b] proposed a triboelectric thin sensor (300 µm) with nanostructured PI thin film and nanostructured copper (Cu) film ( Figure 2f) for the pulse signal detection. Nanostructured Cu film was obtained by depositing ultrathin CU film (50 nm) to the nanostructured PI film. In addition, as an electrode, ultrathin Cu film was also deposited on the back side of the PI film. Because of the flexible

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nanostructures, the real contact area between two triboelectric layers is enlarged, and charge transfer is aided by the ultrathin electrode's tip discharge effect. In addition, by methodically examining the mechanical-electrical-sensing behaviors of electret nanogenerators, Chen et al. proposed self-powered pressure sensors. The performance of the sensor can be changed by the spatial densities of elastomer structures. [54] They also reported a hierarchical elastomer microstructure based self-powered pressure sensor that can operate continuous monitoring of cardiac and vascular conditions with a comfortable multimodal. The pressure sensor showed high sensitivity (7.989 V kPa −1 ), fast response (40 ms), wide operation range (0.1-60 kPa), and high signal-to-noise ratio (38 dB) that can monitor various biosignals such as artery, heart condition, and blood pressure monitoring. In a similar form, Fang et al. presented an ultrathin (255 µm) scalable, conformable, waterproof, and machine-learningassisted textile triboelectric pulse sensor that can be serviceable in a sweaty state along with body motion artifacts ( Figure 2g). [55] A hierarchically structured CNTs triboelectric layer facilitates the textile triboelectric sensor to achieve high sensitivity with a signal-to-noise ratio (SNR) of 23.3 dB, a response time (τ) of 40 ms, and a sensitivity up to 0.21 µA kPa −1 . Shi et al. fabricated a stretchable triboelectric sensor using the electrospinning and spraying method. [56] The sensor consists of silver nanowire network that are tightly wrapped by double-layered thermoplastic polyurethane (TPU) fibers ( Figure 2g). Because this ultrathin soft sensor with 120 µm thickness is single-electrode mode, the surface charging due to the triboelectric effect is obtained between human skin and a TPU layer that is contacted to the skin. The resulting sensor exhibits exceptional deformation recovery ability, working stability, and durability, and has outstanding mechanical qualities with a low-resistance electrode of 257.3 at a strain of 150%. An intelligent electronic glove with a 0.1539 kPa −1 pressure sensitivity was developed to be able to recognize various movements.
Similarly, Peng et al. also designed all-nanofiber triboelectric sensors by sandwiching the silver nanowire electrode between the top polylactic-co-glycolic acid (PLGA) triboelectric layer and the bottom polyvinyl alcohol (PVA) substrate. [52b] These nanofiber networks and numerous 3D micro-to-nanohierarchical pores provide high specific surface area for contact electrification and pressure sensing. This single-electrode mode triboelectric sensor was 120 µm and the voltage response pressure sensitivity was 0.011 kPa −1 , which enables to monitor wholebody physiological signals.

Electrophysiological Signal Monitoring
Human electrophysiological signals indicate the change of electrical potential occurring in human body such as electrocardiogram (ECG), electroencephalogram (EEG), electrooculogram (EOG), electromyogram (EMG), and galvanic skin response (GSR). Unlike the ultrathin soft sensors for physical signal monitoring, the sensor for monitoring electrophysiological signals, referred to as electrode, directly detects changes in electrical potential and conductivity from human skin. Electrodes are optimally attached to the body, depending on the targeted signals. Typically, two electrodes are placed on the skin to monitor the electrical potential, and the difference in potential between two electrodes is recorded. By contrast, to detect the conductivity of human skin, a potential is applied across the two electrodes attached to the skin and the current passing through the body is recorded. Generally, the electrode consists of a conductive material and substrate. The wrinkles of human skin make it difficult to measure electrophysiological signals with conventional electrodes, which typically use thick metals and strong adhesives. [5] Therefore, ultrathin soft electrodes that can be conformally attached to the skin without other adhesives have been actively developed. By introducing stretchability a) Fluorinated ethylene propylene; b) Ethyl cellulose; c) Poly(l-lactide); d) Thermoplastic polyurethane; e) Po(lylactic-co-glycolic acid); f) Gauge factor.
www.advelectronicmat.de through serpentine pattern to thin electrode, the ultrathin soft electrode can be conformably attached to curved surfaces such as human body, while not disturbing human motion. [10] In addition, by making the substrate soft, thin, and porous, the electrode was able to achieve strong adhesion to the skin. Table 5 summarizes the materials and characteristics of ultrathin electrodes for monitoring electrophysiological signals.

Planar Substrate
Yeo's group fabricated serpentine patterned ultrathin Au electrode (1.2 µm) through microfabrication methods such as photolithography, wet/dry etching, metal deposition, and material transfer printing. [3,5,57] The open-mesh serpentines provided mechanical flexibility and stretchability for conformal contact on the skin. This conformality increases the contact area between the skin and the electrode, lowering the skin-toelectrode impedance. Therefore, this ultrathin electrode could offer higher SNR than the conventional ones in detecting electrophysiological signals, such as ECG and GSR. Kireev et al. developed multipurpose and reusable electronic tattoos using platinum diselenide (PtSe 2 ) and platinum ditelluride (PtTe 2 ) [58] through chemical vapor deposition (CVD) process (Figure 3a). The thickness was 25.025 µm, thicker than the gold electrode manufactured through microfabrication, but the skin impedance was 4 kΩ, which was lower than that of the electrode. Moreover, the author demonstrated EMG, EOG, and ECG can be monitored continuously without fault using this ultrathin electrode. However, time-consuming techniques and the requirement such as conventional microfabrication method for pricey specialist equipment severely limit its widespread implementation. To leverage the potential for scalable manufacturing that bypasses costly microfabrication processes, Yeo's group fabricated the ultrathin gold electrode more readily by combining electron beam (e-beam) evaporation and laser cutting. Although the thickness of new gold electrode was thicker than previous gold electrode as 12.7 µm, the Au electrode could be attached to skin conformally and monitor the ECG signal successfully. Mishra et al. [59] and Kwon et al. [60] used the AJP for electrode fabrication (Figure 3b). With AJP enabling highthroughput, efficient, and rapid manufacturing of nanomembrane sensors and electronic platforms, they fabricated ultrathin electrodes based on silver nanoparticle (2 µm) and graphene (11 µm), respectively. Similarly, Behfar et al. [61] fabricated ultrathin electrodes by screen-printing conductive ink on ultrathin elastic and breathable substrate. Yang et al. prepared a freestanding nanomembrane through novel onestep bubble blowing method. [62] Then, by simply transferring a silver nanowire (AgNW) conductive film onto the thermoplastic elastomer (TPE) nanomembrane, they could develop an ultrathin (150 nm), ultralightweight (0.24 mg cm −2 ), self-adhesive, and transparent AgNW-TPE electrode. The patterned AgNW-TPE electrode has a high sensitivity to electromyogram signals and simultaneously offers better stretchability (62%), better breathability (water vapor transmission rate = 580.18 g m −2 d −1 ), higher SNR, and a mild peel strength compared to those of its commercial counterpart. Lastly, Zhang et al. fabricated laser-induced graphene (LIG)/MXene-Ti 3 C 2 T x @EDOT (LME) electrode with simple step: 1) laser engraving of a commercial Kapton PI layer and 2) transfer of patterned conductive layer to elastic substrate. [63] The LME electrodes have low skin-electrode impedance and produce high-quality ECG readings.

Nanofiber Substrate
Planar substrate-based ultrathin soft electrodes can fully conform to human without extra adhesive. However, most of the epidermal electrodes reported previously do not allow the penetration of gas as well as liquid, which poses a challenge for long-time applications due to the evaporation of sweat (Figure 3a). For allowing the penetration of liquid and gas, ultrathin soft electrodes based on porous substrates are being developed. Liu et al. reported a type of ultrathin, conformable, and highly robust nanonetwork epidermal electrode (NEE), which was used to collect ECG signals in a stable and robust way. [64] It was fabricated through a novel method in which the Reproduced with permission. [58] Copyright 2021, American Chemical Society. b) Schematic and optical image of an electronic tattoo (E-tattoo) and SEM image of a CNT/SNF membrane. Reproduced with permission. [65] Copyright 2021, Wiley-VCH GmbH. c) Picture of a temporary tattoo electrode (TTE) on the scalp and the schematic of TTE structure. Reproduced under the terms of Creative Commons Attribution (CC-BY). [66] Copyright 2020, The Authors. Published by Springer Nature.
www.advelectronicmat.de electrospinning of polyamide nanofibers and electrospraying of AgNWs was simultaneously conducted. As a result, two kinds of networks were formed by the nanofibers and the AgNWs mutually and homogenously convoluted in a nonwoven way. For the NEE with a thickness of 125 nm, a low sheet resistance of 4.14 Ω sq −1 with an optical transmittance of 82.3% at a wavelength of 550 nm was achieved. The NEE could fully conform to human skin via van der Waals force and showed a contact impedance that was over 50% lower than what was obtained for commercial gel electrodes. Because of the high conductivity and the conformal contact, the NEE was confirmed to stably record an ECG signal, especially when a human body was in motion. These features make the NEE promising for use in epidermal electronics that perform ambulatory measurements of physiological signals for healthcare applications. Gogurla et al. [65] developed multifunctional, biocompatible, malleable, lightweight, and ultrathin e-tattoo devices that may be directly applied to biological tissues and are easily removed after use. The e-tattoo platform is made up of an ultrathin and porous silk nanofiber (SNF) network that is functionalized by CNTs. CNT/SNF e-tattoos adhere to uneven skin surfaces, such as those on the fingertips and forehand. Because of the methanol solvent, the CNT ink could also promote crosslinking SNFs during functionalization. Uniformly coated CNTs on the SNF membrane have the potential to be used as an ultrathin, single-sided-conductive e-tattoo sticker that is biocompatible, light, malleable, and skin-adhesive. Furthermore, Ferrari et al. developed conductive polymer tattoo electrodes using commercially available tattoo paper and organic materials. [66] These tattoo electrodes (TTEs) are made by depositing PEDOT:PSS onto commercially available tattoo paper by inkjet printing. TTEs can conformably attach to the skin and provide an unnoticeable touch between the human body and the electronics due to their 0.36 µm thickness and softness (Figure 3c). TTEs have been successfully used in EMG and ECG in numerous cases. Their performance has been demonstrated not only by their mechanical resilience, but also by their long-term electrical stability in monitoring biosignals. The latter was assessed utilizing skin-contact impedance measurements and electrophysiological recordings for a period of up to 48 h.

Water-Soluble Substrate
Planar substrate of traditional electrode have always been one of major limiting factors to further improving gas permeability, weight, and thickness of on-skin electronics. [10] Miyamoto et al. demonstrated substrate-free electronics by developing inflammation-free, extremely gas-permeable, ultrathin, lightweight, and flexible sensors using a nanomesh structure. [10] The biocompatible PVA mesh structure allows for great gas permeability without obstructing sweat glands, as well as stretchability without producing discomfort even when glued to the skin for an extended period of time. Because of the dissolved PVA layer on the bottom of the mesh, the nanomesh conductors cling to the skin. PVA nanofibers that have been dissolved generate a superthin adhesion layer with a thickness of many tens of nanometers. This level of conformability has not previously been recorded, and it permits people to forget that the nanomesh conductors are attached to their skin.

Chemical Signal Monitoring
Chemical signals indicate the concentration change of molecules and/or ions measurable in our body such as water molecules, glucose, pH, and electrolyte. Likewise monitoring physical signals, chemical signals can be converted into electrical properties such as potential, current, and resistance, and analyzing the concentration of chemical substances using these three electrical properties is called potentiometry, [11] amperometry, [67] and conductometry, [68] respectively. Therefore, ultrathin soft sensors for monitoring chemical signals consist of substrate, electrode, and functional layer. In this section, we reviewed ultrathin soft sensors according to the type of chemical signal to be monitored. Table 6 summarizes the materials and characteristics of ultrathin soft electrochemical sensors for monitoring chemical signals.

Molecule Monitoring
Monitoring humidity of breathing and skin surface can be key to estimating and monitoring individual health status. [68] The ultrathin sensors that can monitor humidity by inducing a change in resistance through a functional layer that absorbs water molecules have been developed. Chen et al. fabricated a robust ultrathin (3.091 µm) humidity sensor with high sensitivity and ultrafast response based on polydopamine-cross-linked-gold nanoparticles (AuNPs) monolayer (Figure 4a). [68] Through dynamical hydrogen bond interactions, the hydrophilic groups of this monolayer can rapidly grab water molecules. If this ultrathin soft sensor absorbs water molecules, the hydrogen bonds are formed between water molecules and hydrophilic groups. This hydrogen bonds allow protons to hop between water molecules (Grotthuss mechanism), forming electrical channels between adjacent AuNP, resulting in increased conductive pathways and conductivity. Simultaneously, the sensors can easily release water molecules with high speed by illuminating nearinfrared laser light. Due to the excellent light-to-heat conversion properties of AuNPs, the sensor is quickly heated up by light and the absorbed water is rapidly evaporated, resulting in ultrafast response and consistent use. Contrary to this sensor, Qiu et al. fabricated an ultrathin humidity sensor whose resistance rises due to the increased humidity. [7] The author fabricated an transparent ultrathin sensor (100 µm) by synthesized double networked hydrogel and silk fibroin, which can absorb water molecule.
Monitoring and managing glucose levels are essential for diagnosing and managing diabetes. A sensor that facilitates long-term, continuous monitoring without causing discomfort to the user is highly desirable. [69] Liu et al. developed 0.23 µm as-assembled glucose and lactate sensors [67] that exhibit excellent sensitivity, good cyclic repeatability, and especially longterm stability of up to 25 and 23.6 h, respectively (Figure 4b). These glucose sensors have a 1.25-times longer working life www.advelectronicmat.de than the best available sensors reported so far. For fabricating this sensor, they presented a hybridization material strategy for immobilizing enzymes in a porous enzymatic network membrane connected to an ultrathin platinum nanoparticles/graphene nanocomposite film. This tactic is effective in decreasing enzyme escape and increasing enzyme-analyte reactions to ensure the integrity of the glucose and lactate sensors.

Ion Monitoring
Consuming sodium is crucial for maintaining blood pressure and healthy nerve and muscle function. [70] Severe conditions like kidney failure, [71] heart disease, [72] cancer, [73] and osteoporosis [74] can be brought on by abnormal sodium levels in the body. Several ultrathin soft sensors with an ion-selective membrane were developed to monitor sodium concentration, and to improve user comfort, the ultrathin sensors were integrated with a wireless system. [11,75] Lee et al. developed ultrathin, lowprofile, and soft electronic platform along with miniaturized sodium sensors. [11] Total thickness was 2 mm and the ultrathin soft platform was conformally laminated on an oral retainer. Then, they demonstrated the microstructured, ion-selective sodium sensor with functionalized polymer membranes demonstrates a highly sensitive and selective detection of sodium ions in a food mixture. Lim et al. also developed a fully integrated wireless ultrathin sodium detection system with improved long-term signal stability by applying nanostructures to the electronic transducers, and this ultrathin soft system (thickness: 2 mm) was able to make conformal contact even on uneven hand surfaces (Figure 4c). [75] In a wide range of applications, including environmental, industrial, and biomedical conditions, pH, which is defined as a logarithmic measure of the hydrogen ion concentration, [76] is a crucial analytical measure of the mechanism of chemical or biological reactions. To monitor health, pH sensors are essential in the majority of continuous chemical processes. [77] In recent, two research introduced ultrathin soft sensors for monitoring pH using a same pH responsive polyaniline (PANI) membrane. By applying IDE structure or porous structure, they could acquire high performance with ultrathin dimension of sensor. Li et al. developed a conductimetric-type pH sensor using PANI membrane fabricated on an ultrathin soft substrate (Figure 4d). [78] The sensitive element of one sensor is a PANI membrane layer, the transmission layer is an interdigital electrode, and the substrate film is a PI membrane. A membrane with a thickness of less than 100 µm still makes up the entire sensor because each layer of the sensor is less than 30 µm thick. Etching was used to pattern the electrodes on the PI film. The pH sensor's responsiveness is enhanced by the contact area between the PANI and interdigital electrodes. And over the entire pH range of 5.45 to 8.62, the sensor showed a high sensitivity of 58.57 mV pH −1 . Meanwhile, using same material, Liu et al. developed a potentiometric pH sensor with high-performance, portable, gas permeable, transparent based on a self-supporting carbon micromesh electrode (CME). [79] The CME was made up of interlaced carbon lines with width ranging from 6 to 30 µm, which was   Reproduced with permission. [68] Copyright 2019, American Chemical Society. b) Schematic of the working electrode and the sensing mechanism of a glucose sensor. Reproduced with permission. [67] Copyright 2021, Wiley-VCH GmbH. c) Photo of an integrated, flexible device and illustration of an SS-ISE, reference electrode, and transducer preparation step. Reproduced with permission.
[75b] Copyright 2021, Elsevier. d) Schematic of a pH sensor describing a transformation of PANI in acid and basic solutions. Reproduced with permission. [78] Copyright 2020, Royal Society of Chemistry. www.advelectronicmat.de

Integrated Materials
Ultrathin soft sensors commonly are composed of the following three materials: conductive materials, functional materials, and substrate materials. Conductive materials are used in all types of sensors. Conductive materials played a direct role in sensing electrodes, capacitive sensors, and resistive sensors. In other types of sensors, conductive materials played a part in transferring the externally induced changes in electrical properties. Functional materials play a role in converting mechanical and chemical change into electrical change can be subdivided into dielectric materials, resistive materials piezoelectric materials, triboelectric materials, and chemically reactive materials. The substrate materials maintain the sensor's shape and determine the sensor's mechanical properties. In addition, the substrate prevents functional materials from escaping such as chemical molecules. [68] As a substrate of ultrathin soft sensors, PI and PDMS, which have been traditionally used a lot, are still being used, and recently fabrics and papers with porous structures and cost-effective are being newly used. In this section, among the various conductive materials and functional materials that are the core of sensing, we reviewed the materials commonly used in various ultrathin soft sensors.

Conductive Materials
Most of these ultrathin sensors are constructed from metal, metal-nanowire/nanoparticle, carbon material, and conductive polymers, and possess complex micro-and nanostructures, such as pyramids, semispheres, and cylinders. [23] Among various conductive materials, Au is the most used for the ultrathin soft sensor reviewed in this paper. [3,5,10,16,19,38,41,43,44,47,57,67,68,75a,80] Au has been widely used for biomedical devices with a great biocompatibility [19,81] unlike Ag and Cu, which require additional chemical treatment due to poor biocompatibility despite high conductivity. However, the downside is that gold is expensive. Howe et al. investigated two different materials of NiTi, [82] and Mg as a material used for sensors implanted in the body. [19] As a result, they fabricated ultrathin capacitive sensor using Mg because it can be safely and naturally absorbed in the body for transient sensing. [83] Conductive materials used in epidermal or tattoo-like ultrathin soft sensors must be compatible with highly deformable tissues and easily functionalized in order to create seamless biotic-abiotic interfaces. [65] Due to their high aspect ratios, which lower the percolation threshold for conducting charges, networks of 1D/2D nanomaterials have proven to be reliable candidates. Among various nanomaterials such as AgNWs/ AgNPs, AuNWs, [35,84] CNT, and graphene, AgNWs/AgNPs have been used most frequently to fabricate ultrathin soft sensors. [6,7,15,20,22,27,[60][61][62]64] However, in biological environments, AgNWs have a strong propensity to oxidize and corrode. [65] Instead of AuNWs, stable and biocompatible but expensive and unwieldy, the carbon-based materials [34,[36][37][38]60,65] have been utilized as promising for ultrathin soft sensor due to certain distinctive qualities, such as high photo/electrothermal conversion efficiency, high mechanical strength, chemical stability, and biocompatibility. In addition, it was confirmed that these materials were combined with a porous substrate and used to fabricate a high-performance ultrathin soft resistive sensor.
A metal-free material without the need for high temperatures, [85] PEDOT:PSS can be produced inexpensively and on a large scale. PEDOT:PSS is the most popular conducting polymer, and this polymer has been successfully used in electrophysiological signal recording and an electrode that transmits a change in an electrical property such as resi stance. [8b,17,48,66] Although metal-based materials are still being used more, PEDOT:PSS has the potential to be used more in the development of sensors for enhanced monitoring of human physiological signals due to its low mechanical stiffness that is compatible with skin mechanics as well as fabric, tattoo releasable polymer film, and nanofibrous membrane. [66]

Piezoelectric Materials
Inorganic piezoelectric materials such as PZT have been utilized as active materials for realizing microelectromechanical devices, actuators, and strain sensors due to their high electromechanical coupling coefficients, [41] high Curie temperature, and high piezoelectric coefficient, which allow large output voltage signal with small mechanical deformation. [44] Lu et al used nanoscale PZT ribbon based piezoelectric sensor to monitor the real-time eyelid motion for the eye fatigue recognition through the eyelid motion. [44] Due to the small size of PZT nanoribbons, the entire sensor had 10 µm of thickness conformally attaching to the eyelid skin. However, despite the excellent performance of the PZT-based piezoelectric sensors, it has been proposed that the PZT lacks cost-effectiveness, limited scalability, as well as biocompatibility due to its nature of containing materials such as lead. Therefore, it was investigated that the material selection trend for piezoelectric sensors had shifted toward a low cost, biocompatible materials such as AlN, PVDF, P(VDF-TrFE), and poly-l-lactic acid (PLLA).
Recently, it was discovered that PLLA, a biodegradable polymer widely used in FDA-approved implants, exhibits piezoelectricity when properly treated. Moreover, it is worth noting that there exists a controversy about the biocompatibility of the PVDF. Some papers had insisted the toxicity and nondegradable character of PVDF, [86] on the other hand, PVDF is considered as biocompatible and mechanical robust material suitable as material for piezoelectric sensor. [86] Sun et al. used an AlN based in vivo spatiotemporal epidermal strain monitoring and decoding facial deformation signatures device. [45] The usage of AlN piezoelectric layer had possibility of enabling mass manufacturability, while having biocompatibility with its lead-free nature which will make the clinical transition much smoother. Moreover, newly discovered microfabrication technique from this research realized the flexible AlN sensor while not compromising its performance. Ha et al. utilized PVDFbased piezoelectric e-tattoo for seismocardiography (SCG) measurement. [47] Here, the metalized PVDF sheets were patterned into a filamentary serpentine network as an ultrathin and stretchable mechano-acoustic sensor. A low-cost, ultrathin, and skin-soft 28-µm-thick SCG sensor based on PVDF was developed. Laurila et al. propose a piezoelectric arterial pulse wave sensor using P(VDF-TrFE) as piezoelectric material. [48] The www.advelectronicmat.de performance of the ultrathin P(VDF-TrFE) layer was improved by the use of (3-glycidyloxypropyl)trimethoxysilane crosslinked PEDOT:PSS electrodes. This resulted in ≈70% improvement in remanent polarization and ≈34% increase in coercive field. Moreover, the sensitive of the sensor was directly increased due to a specific set of direct piezoelectric coefficient calculated optimized by experimental result and statistical analysis and finite element modeling. Montero et al. demonstrated a PVDF-TrFE and PEDOT:PSS based self-powered ultrathin pressure sensor used for the arterial pulse waveform signal and limb movement detection.
[8b] P(VDF-TrFE) is a material that is being extensively researched as an active material for sensing applications. It exhibits great biocompatibility, flexibility, cost effectiveness, and ease of processing over a vast region. P(VDF-TrFE) work well with additive manufacturing techniques such as inkjet printing, screen printing, and bar coating. The P(VDF-TrFE)coated IDE structure was compared with the conventional metal-insulator-metal (MIM) structure. First, the uniformity of the P(VDF-TrFE) layer in IDE-based sensors proved effectiveness by not having a critical impact during the fabrication process while the MIM-based sensors, had extreme thickness variations which had possible danger to have short circuits between the electrodes. Moreover, piezoelectric and voltage sensitivity of three distinct P(VDF-TrFE) thickness IDE sensors was compared to that of MIM structure. Higher voltage sensitivity was achieved by the IDE-based structure because the output voltage (V) of the sensor is dictated by the equation V = Q/C, where Q is the generated surface charge, and C is the capacitance. According to the findings, the MIM device has a seven times lower piezoelectric sensitivity than the IDE structure with a comparable thickness (12.8 µm sample), resulting in a 4.7 times higher voltage sensitivity.

Triboelectric Materials
When fabricating triboelectric sensors, selecting the appropriate paired materials with opposite triboelectric polarities from the triboelectric series is one of the most effective strategies for improving the performance of triboelectric sensor. [23,49] As a pair of triboelectric layers, fluorinated ethylene propylene (FEP) and CNTs can be employed. [55] Fang et al. used the nonwoven FEP textile as negative charge layer since it is a highly negative electron affinity material that is very table and lightweight. [87] An SWCNTs conductive network was used as the opposite positive charge layer, and this network was also used as a lead-out electrode. FEP can be used as triboelectric layer after treated by corona charging. Chen et al. used polyethylene terephthalate (PET) and FEP as a pair of triboelectric layers. [54] In the fabrication process, the FEP was treated by negative corona changing and became an electret device assembling. According to their experiment, after 10 days of rapid decay, the surface potential and charge density stabilized at ≈0.8 kV and 1.1 mC m −2 , respectively.
Similarly, Teflon amorphous fluoropolymer (Teflon AF) could be used as one of triboelectric layer. [53] Teflon AF is a biocompatible material [88] that can be easily spin-coated on flexible substrates, [89] and can hold electrostatic charges for a significantly long time after corona charging treatment. [90] Zhong et al. generated net positive and negative charges on the surface of two different Teflon AF layers by corona charging method, and then utilized them as a pair of triboelectric layers. Another candidate material is a TPU. TPU, which is widely used in clothing fabrics due to its excellent fit, biocompatibility, and nontoxicity, has widely used for the fabrication of stretchable triboelectric nanogenerator (TENG) and its potential for energy harvesting and self-powered sensing have been demonstrated. [56] Therefore it can be used for the top encapsulation layer and the bottom sensing layer. [91] According to Zhao et al., [92] in the past, TENG-based electrospun fiber mats were mostly made of PVDF or composite fiber mats. Recently, however, TENG sensors based on biocompatible and biodegradable electrosun fiber mats have been reported. A PLLA is a type of biocompatible and biodegradable polymer that also has benefits like high strength, excellent plasticity, and ease of molding. Thus, it has been used in the biomedical field, and in particular, the micro/nanostructured electrospun PLLA fibers can be inserted into the human body as a bioscaffold. [93] Zhao et al. fabricated ultrathin triboelectric sensor using biocompatible PLLA (negative), and ethyl cellulose (positive) micro-nanofibers and it was able to sensitively detect human gait and limb movements. A PLGA is well-known synthetic biodegradable materials. [52b] PLGA has a strong resistance to weight loss and water absorption in the initial stage, but slightly shrinks and curls due to hydrolytic cleavage of the polymer backbone. [94] Using this material, Peng et al. developed the biodegradable triboelectric sensors.
Finally, the surface nanoarchitectures of two friction layers are important for the output performance. Ouyang et al. [50b] proposed a triboelectric thin sensor (300 µm) with nanostructured PI thin film and nanostructured Cu film. They compared the electrical output according to the structure of the friction layer. As a result, the performance was higher when nanostructured was applied even though it was the same material. [50b,55] These changes could be explained by the fact that the friction layers' actual contact area has been enlarged even more by the flexible nanostructures, and that the charge transfer was facilitated by the tip discharge effect of the nanoelectrodes.

Chemical Reactive Materials
PANI is one of the conductive polymers that are of considerable interest in chemical sensing due to its conjugated electronic structure and excellent electrical conductivity. [78,95] In addition, PANI is an intrinsically pH-sensitive polymer with good environmental stability. Except the reviewed ultrathin pH sensors, PANI has been integrated in a microfluidic device for dynamic pH imaging and mapping. [78,96]

Manufacturing Technologies
The main microfabrication processes include thin film deposition, layer doping, patterning by photolithography, etching to create the desired design, lastly polishing, and bonding. [97] Conventional microfabrication methods derived from the semiconductor industry have been adapted to ultrathin sensors. The www.advelectronicmat.de main techniques of microfabrication are thin films deposition, layers doping, patterning via photolithography, etching to obtain the required design, polishing, and bonding. Devices miniaturization is also used in various research areas: chemistry, physics, materials, computer sciences, etc., allowing the manufacture of portable, hand-held, or even implantable structures. Utilizing conventional microfabrication processes including deposition and lithography techniques are still proposed as a way to fabricate ultrathin layers and film of sensors, which is complicated, expensive, and time-consuming. [23] Recent advances in manufacturing techniques such as laser cutting, [37,80] printing methods, [29] coating methods, [7] and electrospinning [15] are used in fabrications drastically because of cost-effective and customizable processes. The fabricating ultrathin sensors have become more novel with the new printing methods and electrospinning that can make porous structures easily. Newly proposed costeffective, fast, and customizable materials expand the capability of the ultrathin sensors to various microscale sensor applications. In addition, new methods such as self-assembly, [16,37] and bubble blowing [62] have been applied to produce superfine fabrications.

Deposition
There has been a recent emerging interest in electrochemical deposition, especially deposition of multilayer structures. Studies have shown that this multilayer technology can create systems with nanoscale structure and composition variations, with capability to modify material properties for various applications. [98] Regarding wearable devices, which are typically not planar and have sub-millimeter features, deposition process is physically essential for the fabrication device to meet the biocompatibility goals. Vacuum-based deposition techniques have been further developed to produce thin films and layers. The physical vapor deposition technique was employed in ultrathin sensor fabrication studies including thermal evaporation, e-beam, vacuum evaporation, and sputtering method. All possible deposition methods have inherent advantages and disadvantages regarding the quality and characteristics of the multilayers they create and the ease of their production. Financial and temporal restrictions should be also considered to find manufacturing feasibility.
Recent studies have proposed thermal evaporation techniques to control the thickness and strength of ultrathin sensor's layer, among the various evaporation techniques. The materials with a low melting point like Ag, Cu, Au are favorable with this technique. Zhong et al. [53] performed thermal evaporation on the patterned chromium (Cr)/Au bonding pads to enhance the strength between Au and the Teflon AF/parylene layer. Sang et al. [30] used thermal evaporation technique to dope the gold ions on top of silicon nanomembrane by controlling the deposition thickness. This technique allows researchers to create an ultrathin self-powered pressure sensor by depositing two Au electrodes on the outside of the device surface. Nevertheless, it is difficult to use thermal evaporation technique to some materials with high melting point and impurities can be evaporated with the deposition material because it uses tungsten wire as a heating source inside of the evaporator. Also, the strength between the layers is weaker than other techniques. E-beam technique was attempted to reduce the possibility of impurities getting evaporated with the material in recent studies. Jeong et al. [32] performed e-beam metal deposition method to fabricate the electrode. Kim et al. [80] used e-beam deposition tool to deposit Cr and Au layers in thicknesses of 5 and 200 nm, respectively, on top of the polyimide sheet. Miyamoto et al. [10] used vacuum evaporation technique to form a gold layer on the surface of the PVA nanomesh sheet with a nominal thickness 70-100 nm. Fang et al. [55] also used vacuum evaporation chamber to deposit Al electrode (100 nm) on top of clean FEP film at a rate of 0.5 nm s −1 and a base pressure of 3 µTorr.
Since thermal evaporation, e-beam, vacuum evaporation proceeds in vacuum and high temperature, they are uneconomical in terms of power supply. An alternative technique, sputtering is used to deposit various types of materials in lower vacuum conditions and lower temperatures compared to evaporation techniques. Recent studies from Chen et al. [68] used sputter to interdigitate a Ni/Au IDE on the polyimide substrate of the humidity sensor. Hua et al. [99] deposited Au, aluminum thin film, and sensitive elements using sputtercoating. The study forms a patterned mask for a skin-inspired highly stretchable and conformable matrix network. In addition to enhance strength between films, CVD technique is recommended to evenly deposit and form a very thin layer at the same time. Researchers introduce strong successful cases of CVD to deposit the substrate of the sensor. Montero et al. [8b] performed CVD to deposit a Parylene-C layer in order to fabricate a highly uniform thickness ultrathin substrate of sensor. Lee et al. [100] also deposited 2 µm thick parylene film on top of the F-polymer/glass layer, which was used as the substrate of the organic electrochemical transistor-based sensor. Li et al. [36] fabricated reduced graphene oxide used as a base of pressure sensor using CVD under 250 °C for 5 h to make graphene oxide as reduced graphene oxide under nitrogen protection. Thermal-assisted conversion (TAC)-based CVD was performed in Kireev et al. [58] because an advantage of PtSe 2 and PtTe 2 is their low-temperature growth. TAC-based CVD process allows to form a high quality PtSe 2 and PtTe 2 at temperature as low as 400 °C, making it compatible with polymeric substrates.

Photolithography and Etching
Photolithography is a process used in microfabrication to selectively remove areas in thin film or the bulk of a substrate. Photolithography uses light or other optical sources on the substrate to transfer a geometric pattern in the photomask to light reactive chemical photoresist. Since photolithography transfers shape from a template, this is used in micromanufacturing applications including sensor fabrications. The etching process removes the uppermost layer substrate that is not protected by the photoresist with a chemical agent. This reveals the layer to the unprotected areas. Fabricating ultrathin sensors with these techniques has been used widely, but is expensive, and time-consuming. However, recent studies have attempted to fabricate various ultrathin sensors with new photolithography applications to fabricate optimized layers, electrode, and www.advelectronicmat.de Figure 5. Manufacturing technologies based on laser and printing methods. a) Image of single-step removal of the nonfunctional region of the laser-cut film and leaving the functional electrodes and schematic of the electrode-interconnect system. Reproduced with permission. [80] Copyright 2022, Wiley-VCH GmbH. b) Schematic diagram of laser cutting process. Reproduced with permission. [17] Copyright 2020, American Chemical Society.

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patterning for their applications, such as capacitive, resistive, piezosensors. Such a process raises the intriguing possibility of a simple monolithic process forming a customized surface for optimal biocompatibility.
For capacitive sensor, [19] Howe et al. fabricated a soft flowdiverter system that allows active monitoring of intra-aneurysmal hemodynamics using photolithography. The flow sensor, in the form of capacitance sensor, had serpentine patterned top and bottom metal layer fabricated using photolithography and wet etching and PI as spin-coated dielectric layer. The sensor was able to allow 500% of strains, bend 180° of with 0.75 mm radius of curvature, and have sensitivity as low as 0.032 m s −1 . For resistive sensor, [101] Chang et al. used photolithography to fabricate the interfacial layer of the ultrasensitive flexible piezoresistive sensor, which its sensitivity was enhanced by the conductive network of an SWNT interfacial layer. The mold for the interfacial layer had shape of pyramidal microstructure and was fabricated using photolithography and anisotropic wet etching process. Then, the SWNT/M13 mixture was poured into the mold to form a top layer of the resistive sensor. M13 was used to overcome the issue of weak adhesion. This sensor showed the resistance change by 5 orders for a pressure range of 0.1-100 kPa, high stability and quick response (20 ms).
For multilayer applications, [44] Lu et al. introduced in vivo eye fatigue monitoring sensor fabricated by standard photolithography of a multiple layer of silicon dioxide, Pt, PZT, and Au/Cr, where silicon dioxide being the sacrificial layer. To transfer the printing, PDMS stamp was used according to the law of fracture competition. The printed sensor proves its durability and high sensitivity. This work could be used for detection of abnormal eyelid motions, such as overlong closure, high blinking frequency or weak gazing strength. For clinical grade stretchable electrode, [57a] Kim et al. applied photolithography fabrication technique for the stretchable electrodes with optimized structure that generates clinical-grade electrocardiogram. The three gold electrodes that directly contact with the skin were photolithographically patterned into a network which is a circular island bridged with meaner lines for enhancement of stretchability and conformal contact to the skin. Moreover, the photolithography was used to fabricate the thin-film structure as a standard microfabrication process. The skin-conformal electrode system was capable of comfortable and ambulatory monitoring of ECG signal as well as notifying clinically significant cardia event.
Multiple scholars attempted to fabricate skin-conformal and stretchable electrode that propose new trend of biocompatible applications. Kim et al. [3a] used photolithography to pattern the Au island and meander interconnect in the skin-conformal electronics system (SKINTRONICS) as well as the thin film electronic circuit for portable and continuous stress monitoring. The SKINTRONICS could offer ergonomic and conformal lamination on the skin. Both circuit and the electrode were developed on a silicon wafer spin-coated with PDMS/PI layer. For the circuit, PI-Cu-PI-Cu-PI layer was deposited on the substrate with photolithography patterning conducted for each layer. For the electrode, photolithography was done on Cr/Au. The flexible circuit and electrode system were able to allow sensitive monitoring of GSR with minimized reduction of SNR compared to the SNR of BioRadio. This showed that the stretchable nanomembrane gold electrode was proven capable of seamless and conformal contact with the skin without the usage of conductive gel. Recent researcher shows a full device development using this advanced fabrication for clinical applications. Zavanelli et al. [57b] used photolithography fabricated stretchable electrode and flexible circuit for the at-home wireless monitoring of acute hemodynamic disturbance for sleep apnea and sleep stages detection. As shown in previous research, Au/Cr was patterned for stretchable electrode and Cu and PI multilayer was stacked and patterned for the flexible circuit. This system showed 100% of sensitivity and 95% of precision at detecting apneas and hypopneas in at-home trials compared to the data analyzed by professional sleep clinicians. In addition, Kwon et al. [57c] demonstrated a nanomembrane electrode fabrication using photolithography for EMG) activity recording during operant conditioning of H-reflex. Notably, the researchers proposed capability of breathable, large-area epidermal electronic systems for recording electromyographic activity from new electrode fabrication. The Au/Cr metal layer was deposited on PI, patterned using photomask and etchant. Then, PI was spin-coated on top of patterned Au/Cr layer. Then the PI with same photomask was etched to get the electrode exposed. The patterning of electrode enabled large-area epidermal EMG recording with high breathability due to the electrode patterning.

Laser
Advanced nanomanufacturing based on laser machining enables rapid, facile, and cost-effective fabrication of ultrathin soft sensors. Laser machining techniques enable versatile, high-precision processing such as cutting, [37,80] lifting, [41] and carbonization [63,79] on a variety of materials while minimizing thermal effects. Recent studies have shown selective cutting, chemical and physical reaction cutting, and edge cutting. For selective cutting, Kim et al. [80] fabricated ultrathin Au electrode with a single-step removal method using a femtosecond infrared (IR) laser micromachining system (WS-Flex, Optec). The ultrathin electrode was fabricated by cutting the ultrathin Au film and removing the nonfunctional region (Figure 5a). Then, remaining functional electrodes were directly incorporated with a highly breathable substrate. In addition, the laser cutting can be adjusted to cut only the top layer without damaging the layer underneath by simply adjusting the intensity of the laser (Figure 5b). Liu et al. [37] fabricated the interdigitated conductive electrode on commercial textiles by forming a mask using a laser technique. The PI tape, sealing commercial polyester c) Schematic of the screen-printing process of a paper-based wireless wearable pressure sensor. Reproduced with permission. [29] Copyright 2021, American Chemical Society. d) Illustration of aerosol jet printing, cross-sectional SEM image and measured cross-sectional profile. Reproduced with permission. [60a] Copyright 2020, American Chemical Society. e) Screen printing process and configuration of flexible sensing electronics. Reproduced with permission. [33] Copyright 2021, Elsevier. www.advelectronicmat.de textiles, was cut into predesigned patterns without damaging the textile underneath using a computer-controlled commercial CO2 laser cutter system. Then, the author could fabricate interdigitated electrode by depositing Ni on the textile with PI mask and removing PI mask. Research about selective laser cutting proposed highly efficient and controllable processing techniques.
Laser machining technique can also be used to ultrathin sensor's fabrication process. The unique laser lift-off (LLO) technique removes the ultrathin soft film from the rigid substrate, which selectively removes the response layer. A series of chemical reactions and physical changes are induced at the interface when the responsive material is exposed to laser light through a transparent carrier. [102] Park et al. [41] employed inorganic-based LLO method to fabricate PZT based-ultrathin soft sensors. In the process of fabrication, to exfoliate the PZT thin film off the hard substrate, a 308 nm XeCl excimer laser was irradiated to the backside of the sapphire substrate. The PZT was partially vaporized and decomposed at the interface between the sapphire and PZT layers, which caused the exfoliation. Compared with other methods, LLO process has become an emerging key technology for electronics manufacturing due to its large-area, non-contacting, high efficiency, reliability, controllability, and low-damage advantages.
Moreover, carbon-based material can be cost-effectively fabricated using laser technique into various patterns and porous structures. Researchers have shown that allowing laser scribing and local carbonization during the patterning process, saves time and energy. Zhang et al. [63] obtained conductive LIG as a supporting layer-by-laser engraving of PI film using a 10.6 µm CO2 microlaser. Through this laser scribing technique that is scalable, facile, and cost-effective, the author could easily fabricate various patterned LIG layer, resulting in the fabrication of high-performance ultrathin multifunctional sensors. Generally, the carbon-based materials generated by laser scribing methods are restricted to the spatial resolution of the focused laser. Additionally, due to beam optics and diffraction restrictions, laser-engraved carbon features formed directly from lasing are bigger than hundreds of micrometers. [103] Liu et al. [79] introduced fabricating the carbon micromesh electrode with a minimum width of 6 µm by combining photolithography and laser scribing technologies. First, wet etching and photolithography were used to pattern the PI micromesh. A 450 nm laser scribing micromachining system (Diatools K6, Shanghai Diaotu Industrial Co., Ltd., China) was then used to carbonize and turn the PI micromesh into carbon micromesh. The surface morphologies of the CMEs revealed a porous network with numerous active edge-plane locations. This was due to the rapid evaporation of the gas product during the laser-induced PI film burning and carbonization process. [104] This high control over the scribing process results in an excellent, micro-cracking-free technique.

Printing Method
Printing in manufacturing has recently provided significant cost, time, and quality benefits across various industries.
Particularly printed electronics offer novel applications, including sensor arrays and energy harvesting systems. Printed electronics lead the development of new materials in conjunction with diverse fabrication processes utilized in various applications. Recent research on printing methods in manufacturing highlights the benefits of removing costly hard tooling, masks, and vertical/horizontal integration, which reduces overall manufacturing steps. Flexible manufacturing with printing, researchers optimized geometrics and material composite to create new sensors and unique layers in electronics design. Reduced manufacturing lead-time and agile process allow flexibility in mass customization, low-volume production, and diverse optimization. The printing technique is categorized as additive manufacturing, in which the material is deposited layer by layer to build structures or features. Traditional subtractive manufacturing requires masking and etching to remove material to get to the final form. In comparison, microfabrication with printing shows the capabilities of processing thin and thick film. This advanced printing method enables to create screen-printed electrodes and biosensors, printable electronic devices, and polymeric microseparation devices. Recent microfabrication integrated printing techniques such as screen printing, aerosol jet, and inkjet printing methods to show a cost-effective, simple step, and expandable manufacturing process.

Screen Printing
Screen printing is a widely used, convenient, and cheap process (Figure 5c). Screen printing can be done by screen printer or even by hand using liquid phase materials. Using well-known industrial screen printers, layers are deposited onto a flat substrate to produce screen-printed electrodes. [105] Because of their small size, low cost, high integration, and capacity in electrochemical (bio)sensing, screen-printed electrodes stand out from other transducers-based electrodes. [105] In addition, screen printing allows flexibility to produce reliable and repeatable electrodes for sensors due to material compatibility and customization of electrode design. [106] Researchers introduced how the printing process is utilized in sensor fabrications that bring essential analytical capabilities for human physiological signal monitoring. The two electrodes from Yang et al. [29] and Zhao et al. [35] introduced hand screen printing, and Behfar et al. [61] used screen printer for printing process.
Yang et al. [29] made a flexible paper-based wireless wearable pressure sensor using printing paper as a substrate with an interdigital electrode. The interdigitated electrode was screen-printed using silver ink on printing paper in an area of 14.7 mm × 10 mm. It used a screen-printing board that had an electrode shape with a finger width of 0.8 mm and spacing of 0.6 mm. MXene/tissue paper was sandwiched after between an encapsulation layer and a printed silver interdigitated electrode. Zhao et al. [35] used silver ink to screen-print a pair of flexible electrodes on top of nylon fabric and sealed it with parafilm. The shape of the electrode was based on a screen-printing board. Two nylon fabrics were cut into 2 cm www.advelectronicmat.de by 2 cm and then the Ag ink was screen-printed on top of the nylon fabric using a scraper. All layers were printed onto the nylon fabric substrate after six screen-print passes. After the process, it was heat-cured at 135 °C for 5 min. Behfar et al. [61] screen-printed the circuit layout using stretchable conductive ink on an ultrathin TPU. The start of printing process began by preheating TPU film in oven for 30 min at 120 °C to avoid shrinkage during the process. Next, the screen-printing process was done by an automatic offline sheet-fed screen printer (EKRA XH). After the printing, the screen-printed circuit layout was cured at the same temperature and time that was used to preheat the TPU film. Component assembly was carried out by a sheet-fed and roll-to-roll bonding machine (Datacon 2200 EVO). Researchers utilized screen printing to produce optimized electrode and sensor designs, considering material composite and compatibility. The integration of screen-printed electrodes and integrated biosensors provides cost-effective and alternative analytical tools for clinical monitoring solutions. [105]

Aerosol Jet Printing
AJP (Aerosol Jet 200, Optomec) is an additive manufacturing technique using spraying out an aerosol (Figure 5d). Highspeed aerosolized ink is precisely deposited, and path printed to create printed electronics. In AJP method, maskless material deposition allows precise control of the dimensions. In AJP the ink is transported through N2 gas and gets atomized in the atomizer. Then AJP can print for printing sizes down to 10 µm. [107] AJP allows a wider range of ink viscosities (between 1 to 2500 cP) to be used, much wider than the inkjet printer. [20] It is reported that the AJP could be a highly sensitive method because of nozzle overspraying, which leads to the material being pressed out of the printing trace. However, it allows direct patterning of various nanomaterials to many surfaces, including patterning nanoscale materials and building a 3D structure by additive manufacturing.
Herbert et al. introduced a fully printed soft pressure sensor for multiplex sensing of hemodynamics. [27] Being a flexible pressure sensor printed by AJP, the sensor was able to display immediate response time even at high pressure and high bending at a radius of 0.25 mm. It is worth noting that the former flexible sensors could not sense while bending or degrading in such a large radius. The PDMS was used as dielectric layer on the bottom electrode layer. This sensor's total structure is composed of the top layer of PI-AgNP-PI and the bottom layer of structured PDMS-AgNP-PI. The printed PDMS structure showed uniformity and feasibility for continuous printing on numerous sensors. The printing parameter for PDMS trace also was able to be adjusted precisely. The printed PDMS layer allowed the sensor to achieve a sensitivity of 0.013 kPa −1 .
Moreover, a capacitive strain sensor was printed by researchers using AJP for arterial stiffness change monitoring from restenosis. [22] AJP method allowed a maskless form of nanomaterial printing enabling fast prototyping and customization of printed sensors. The printed sensor overcame the limitation of conventional capacitive strain sensors with a sliding mechanism to enhance the sensitivity for wireless continuous monitoring of restenosis. The sliding mechanism has overlapping plates sliding parallel to each other. [108] This mechanism was realized by AJP fabricating the top and bottom layers separately with the structure of PI-AgNP-PI. These electronics for arterial stiffness sensing showed an accurate 6.8% mm −1 with wireless ex vivo real-time detection. [22] It was also reported that the printing of AgNW and PI realized multifunctional strain sensors with adjustable resistance and gauge factor. [6] A printed multilayer of soft material with a structure of PI-AgNW-PI allowed strong adhesion and conformal lamination to the various surface type without the usage of external adhesion sources such as tape.
Researchers introduced nanomaterials fabrications with AJP for biocompatible skin recordings. Kwon et al. introduced a fully printed soft bioelectronic system with graphene nanomaterial for noninvasive biopotential recording. [60a] Utilizing graphene nanomaterial with AJP enabled the use of intrinsic electrical and morphological characteristics of graphene as well as its biocompatibility and sensitivity for skin-biopotential recording. The thickness graphene layer (top layer) was 1 µm, and the thickness of PI layer (bottom layer) was 10 µm showing ultrathin feature as stretchable nanomembrane electrode. As a printed flexible electrode, it was able to stretch 60% under the train test and proven its robustness by passing compressing, pinching, and pressing test without fracture. The repeatability of the electrode also proved no change of electrical resistance after being applied to 500 cycles of 60% tensile strain. Moreover, electrode with functionalized conductive graphene (FCG) and PI was also printed with AJP. [60c] The electrode consisted of 10.5 µm thick PI and 0.8 µm thick FCG. The additive manufacturing of the AJP showed improved signal quality. The SNR clearly showed that the 3D graphene electrode has improved its functionality from 2D sheet of Au and Ag layer electrode. The 3D graphene electrode even showed slightly low but comparable SNR noise against the gel electrodes because the difference was within the error range.
Another research proposed an optimized electrode system with consideration of human usage. Mishra et al. used AJP for fabricating soft electronic systems for highly sensitive tracking of eye movements (vergence). [59] Printing technique using AJP enabled skin-like electrodes and flexible hybrid circuit to comfortably integrate on human head. Optimization of the line width and thickness for each nozzle diameter was conducted with AJP setting in terms of different focusing ratios and sintering time with various temperatures. The target of the optimization was to achieve the highest focusing ratio that has lowest width-to-thickness ratio. This is desirable in the sense that it allows finer trace while keeping the low resistance. The precise printing of AJP allowed to achieve and print the optimal dimension of the sensor with a width of 30 µm and an entire dimension of 1 cm 2 . Multilayer printing of AgNPs showed lower resistance due to agglomeration and deification of AgNP. The average impedance of electrodes showed 7.6 kΩ for gel electrodes and 8.4 kΩ for the printed electrode. When the EOG was measured by both electrodes, printed electrode recorded slightly higher SNR in all cases. www.advelectronicmat.de

Inkjet Printing
Inkjet printing is a microfluidics fabrication process for the deposition of drop-on-demand liquid phase material (Figure 5e). It usually includes ejecting a unit amount of ink from the nozzle in pulses, resulting in drops sprayed onto the substrate. Typically, inkjet printing is less expensive than 3D printing for the same planar resolution, and it enables much thinner printable layers, which improve vertical resolution. [109] Inkjet printing is a useful method to print polymer. There are several examples that used inkjet printing method to deposit different types of polymers on top of a substrate.
To fabricate ultrathin aramid nanofiber films-based flexible sensing electronics, inkjet printing and buckling method were proposed by Wu et al. [33] The mixture of CNT, deionized water, PluronicF-127, and NMP was subjected to an ultrasonic process for 100 min, followed by centrifuging twice for 30 min at 3000 r min −1 . It was extracted and dried in an oven at 80 °C. The resulting CNTs ink was transferred to a cartridge and then printed on aramid nanofiber film using a household printer (HP Deskjet 1112). Montero et al. [8b] inkjet-printed an IDE pattern using PEDOT:PSS onto an ultrathin UV/ozone-treated Parylene-C layer substrate, and P(VDF-TrFE) layer was used for encapsulation. UV/ozone treatment was used to change the surface wettability. The exposure time started at 77 °C for a nontreated surface and decreased until 27 °C after 15 min exposure time. A set of lines were inkjet-printed on Parylene-C substrate to verify the impact of the UV/ozone treatment on the formation of uniform lines. The lines were printed with same printing parameters. Ferrari et al. [66] prepared electrodes in three steps: the temporary tattoo paper washing, the electrodes printing, and the lamination process for the TTEs assembly. The tattoo paper was attached on an acrylic transparent sheet (3 mm) with a tape, to avoid any water infiltration on the backside. After PVA layer is completely removed with deionized water, the tattoo paper was dried with compressed air. PEDOT:PSS was then inkjet-printed according to a customized electrodes design onto the washed tattoo. Inkjet printing was carried out with a Dimatix DMP-2800 system (Fujifilm Corp., Japan) with a 10 pL cartridge (DMC-11610). The printing process was carried out by jetting through three adjacent nozzles at 5 kHz; drop spacing is set at 20 µm. Such parameters are optimized to minimize coffee ring effect and to obtain consistent printing. The introduced inkjet-printing technology is effective at microfluidics and electronics fabrication, making it an ideal method for producing microfluidic sensors and integrated electronics. Their excellent flexibility and transparency enhance their extended use of wearable sensors and microfabrication. [109]

Electrospinning
Electrospinning technique is widely used to form a nanofiber layer for ultrathin sensors. Electrospinning happens when high electric field applies between the droplet coming out from the tip of needle and the surface. Electrospinning technique mainly depends on the material inside the syringe, followed by gauge and flow rate of the needle, working distance between the surface and needle. Also, the potential voltage of the working distance matters while electrospinning works. Under best condition, electrospinning technique can produce very thin fibers with large surface, which has high quality mechanical properties in an easy way. [110] Electrospinning technique that was used in the fabrication of ultrathin sensors could be categorized in three types: electrospinning the material in mixed condition (Figure 6a), electrospinning the materials separately, and electrospinning and spraying each material (Figure 6b).
Jin et al. [15] used method included in first type of electrospinning. The author used nanofiber solution of PVDF powder dissolved in dimethylacetamide and acetone with poly(methyl methacrylate) beads inside. This mixed solution was stirred strongly before electrospinning because of equal distribution of microbeads. Electrospinning method was used to fabricate nanofibrous membranes incorporating the insulating microbeads. A 10 kV voltage was applied between the syringe and the collector over a working distance of 10 cm. During the electrospinning process, the AgNWs/PVDF film was placed on the collector such that the nanofibrous membrane was deposited directly on the AgNW electrode. Sharma et al. [17] blended MXene with a P(VDF-TrFE), and A 21% (w/w) P(VDF-TrFE) in DMF/ACT (3:2) solution loaded to 10 mL plastic syringe connected to a 21-gauge needle and for electrospinning. CNSs were primed as a dielectric material by an electrospinning process. The polymer solution was injected with a continuous flow rate of 2 mL h −1 by applying a 20 kV potential over a collector-needle distance of 15 cm. All samples were electrospun with constant parameters, over a spinning area (5 × 5 cm 2 ) and in a controlled environment. For the best CNS thickness, the spinning time was controlled. Shi et al. [56] mixed N,N-dimethylformamide (DMF, ≥99.8%, ACS reagent) and tetrahydrofuran (THF, ≥99.5%, ACS reagent) with a volume ratio of 1:1. After stirring for 10 min, the TPU pellets were added to the resulting solution at a concentration of 20 wt%, followed by continuous stirring for 12 h, until it is completely dissolved to form a homogeneous electrospinning solution.
In the electrospinning process, the needle-collector distance, and the electrospinning voltage of the electrospinning device (DP30, Tianjin Yunfan Technology Co., Ltd., Tianjin, China) were 10 cm and 9 kV, respectively. The manufactured TPU substrate was dried overnight in a vacuum environment at 50 °C for further use. Miyamoto et al. [10] used electrospinning method to prepare a PVA nanomesh sheet. 13.5 mL of ultrapure water with 1.5 g of PVA powder was initially stirred at 70 °C for 2 h and then subsequently overnight at room temperature to make a 10 wt% PVA aqueous solution. A 20 mL syringe was used in this performance with PVA aqueous solution filled and placed in an electrospinning apparatus (ES-2000S, Fuence). The working distance was set to 20 cm and a voltage of 20 kV was applied. The PVA aqueous solution was ejected at a rate of 5 µL min −1 for ≈15 to 30 min in order to prepare a nanomesh sheet. For easy delamination of the fabricated nanomesh sheet, a silicone-coated paper was placed on the fiber collector, which was swung so that the thickness of the nanomesh sheet would be uniform. Gorgula et al. [65] added poly(ethylene oxide) (PEO) to the as-prepared silk solution to generate a mixture with a viscosity and surface tension suitable for electrospinning. For www.advelectronicmat.de making the SNF network, the silk/PEO solution was loaded into a syringe, which was mounted on the electrospinning machine. The author attached scotch tape on the glass slides placed on an Al foil and used as the collector. The tip and the collector were 20 cm far, and the flow rate of the silk bioink was set to 10 µL min −1 . An electric potential of 15 kV was applied between the 25 G nozzle tip and collector during electrospinning for a duration of 30 min. Afterward, the SNF/substrate samples were treated with methanol to make SNFs insoluble in water and then the used PEO was removed from SNFs during the treatment.
Zhao et al. [92] used second type of electrospinning method mentioned above. The author mounted a 27-gauge needle on a 10 mL syringe pump with precursor solution and the metal nozzle of syringe was attached to the output terminal of grounded high-voltage direct current (HVDC) power supply. A rotated disk wrapped by a stripe of Al foil cleaned by ethanol was used as the fiber collector. The collector was connected to the other output terminal of HVDC power supply. The flow rate of PLLA was 1 mL h −1 by applying total 12 kV potential over working distance of 10 cm and the rotation speed averaged 2 kRPM. The ethyl cellulose flow rate was 3 mL h −1 by applying a voltage of 13 kV over the working distance was 15 cm and the average rotation speed was 3 kRPM. Liu et al. [64] performed fabrication process for the NEE, which was conducted by an electrospinning device (Ucalery, ET-2535H) with a voltage of 19 kV. This technique is frequently used to fabricate functional nanofiber-based films blended with nanoparticles, which is part of third method of electrospinning technique. The polyamide-based nanofibers were electrospun. The needle-collector distance was set to be 16 cm and the pump rate for the PA6 electrospinning and AgNWs electrospraying was set to be 0.01 and 1.2 mm min −1 , respectively. The thickness was controlled by the spinning time. . Reproduced with permission. [64] Copyright 2019, Wiley-VCH GmbH. b) Illustration of the fabrication procedure of textile pressure sensor. Reproduced with permission. [37] Copyright 2017, Wiley-VCH GmbH. c) Customized extrusion technique of liquid-based microtubular sensor. Reproduced with permission. [40] Copyright 2017, Wiley-VCH GmbH. d) Schematic illustration of the freestanding TPE nanomembrane fabrication process. Reproduced with permission. [62] Copyright 2020, Wiley-VCH GmbH. www.advelectronicmat.de

Others
For superfine ultrathin soft sensor fabrication, new and costeffective manufacturing technologies were proposed instead of the conventional microfabrication method, the latest laser technology, printing method, and electrospinning method. Yang et al. proposed the bubble blowing method (Figure 6c) to create a freestanding TPE nanomembrane that is highly transparent, stretchable, and breathable. [62] This method consists of injecting air into the solution using a syringe to generate semisolidified bubbles, transferring them to the substrate, and then solidifying the bubbles. The semi-solidified TPE bubble was formed by injecting 40 mL air into a TPE/cyclohexane solution through a medical syringe. The ultrathin bubble walls transferred to precut PET slices were fully solidified via natural solvent evaporation. This one-step process does not require specialized equipment or device released from a rigid substrate. By transferring an AgNW conductive film onto the TPE nanomembrane, the author created an ultrathin (150 nm), selfadhesive, and transparent electrode.
Meanwhile, a direct writing method was applied to fabricate ultrathin electrodes. [111] The method is writing ink directly on the substrate using a commercial refill pen. Gao et al. write a conductive nanosilver ink on top of nanocellulose papers using a refill pen with 0.5 mm width to fabricate electrodes. Following writing, the substrate was immediately placed in a drying oven to anneal the designed pattern at 140 °C for 10 min. This method allows electrodes to be fabricated easily and in various shapes on various substrates such as raw PET and printing paper at a general wiring speed. Lastly, Xi et al. fabricated an ultrathin resistive sensor by injecting conductive liquid metal (eGaIn) into the ultrathin microtube fabricated by a customized extrusion technique (Figure 6d). [40] When a metal filament is immersed in uncured PDMS and pulled out vertically, a uniform thin layer of PDMS is formed around the metal wire due to viscosity and surface tension. The wettability and viscosity can be controlled by adjusting the temperature of uncured PDMS. Then, the hollow tubular structure can be formed by removing the metal filament after fully curing the PDMS around the filament. This facile fabrication method has the advantage that the length of the sensor can be easily customized.

Conclusion and Outlook
In this paper, we summarized recent advances in developing ultrathin soft sensors in three aspects, including structures, integrated materials, and manufacturing technologies. First, ultrathin soft sensors have been developed to increase sensor performance by applying a porous structure. Microstructures such as pyramids and microconvex have been introduced to configure the porous structure. Meanwhile, the thickness of ultrathin soft sensors ranged from 0.1 to 2000 µm, and the average thickness was 157.09 ± 409.17 µm (median: 29.00 µm; interquartile range: 111.63 µm). Ultrathin sensors have been fabricated by using readily available materials like fabrics, papers, and nanofibers. In terms of materials, many sensors use textiles, which are easy to manufacture. Also, a biocompatible polymer has been used for a substrate. Finally, as the conductive material used in all sensors, the most stable Au is still widely used. Recently, studies to utilize new 2D materials such as MXene or liquid metals have been reported. From the point of view of manufacturing technologies, it is still suggested to produce ultrathin layers and films of sensors using traditional microfabrication methods like deposition and lithography techniques, but this is difficult, expensive, and time-consuming. However, recent advances in manufacturing techniques, such as laser cutting, printing, coating, and electrospinning, are increasingly being used in manufacturing because of costeffective and customizable processes. On the other hand, recent advances in ultrathin soft sensors are toward the development of various directions of gas permeability, [10] transmittance, [7,62] reusable, [29] and biodegradability [52b,111] for enhanced monitoring of human physiological signals.
Ultrathin soft sensors monitor human physiological signals and detect external stimuli applied to the human body. The ultrathin layer structures discussed in this review have a common feature of sense. The ultrathin layer system could detect human input, not even feeling the humans are wearing the sensor. Furthermore, it is also possible to detect external stimuli. Using a device with the same structure, it is possible to precisely detect external stimuli and raw signal acquisition from the body. Researchers introduced detecting finger pressure and an array of soft-touch with ultrathin sensor platforms that show scalability. Lee et al. developed ultrathin nanomesh pressure sensors that can monitor finger pressure without disturbing human sensation. [112] This nanomesh sensor is composed of two Au nanomesh conductive layer sandwiching polyurethane nanomesh intermediate layer (2.5 µm), which is the same structure as the capacitive sensors. This ultrathin soft sensor simultaneously achieves undetectable operation and increased durability, so pressure monitoring is now possible in applications requiring accurate, and continuous monitoring or motion in nature. Researchers attempted to minimize sensory interference from new ultrathin nanomesh sensors composed of compliant nanoporous structures. In addition, He et al. fabricated a semitransparent and highly stretching triboelectric sensor with graphene and PDMS. [113] Using this ultrathin stretchable sensor-based array, they could develop soft pressure sensing skin electronics that could distinguish the intensity and the distribution of the applied external pressure. Then, the author demonstrated that soft-touch interfaces could be applied to control an external device such as a computer. These scalabilities from advancements in sensor fabrication enhanced sensors' stability, enabled long-term biological object monitoring, and reduced physical interference.
One of the recent attractive, and essential research topics is the integration of various ultrathin soft sensors into one multifunctional sensor that can monitor human physiological signals or external stimuli. [99] The physiological signals and ambient conditions are integrated for a better smart system. Multifunctional sensors allow for monitoring key health indicators and create fast, personalized healthcare systems. [63,114] For example, Guo et al. [114] proposed a multifunctional smart contact lens system integrated with a temperature sensor, photodetector, and glucose sensor. Temperature sensors are used to monitor the ocular state, such as corneal temperature, www.advelectronicmat.de intraocular pressure, dryness, [115] and inflammation. [116] Simultaneously, a photodetector monitors external light intensity, eye blinking, and vision. [117] Furthermore, a glucose sensor embedded in the contact lens continuously monitors glucose levels for diabetic diagnostics by accessing tears. [118] Besides, multifunctional sensors have the advantage that they can be suitable for various biomedical needs and application purposes. Qiu et al. proposed a multifunctional ultrathin sensor that can monitor pressure, temperature, and humidity. [7] The sensor can monitor various physiological signals, such as facial muscle expression, human joint movement, breathing conditions, and throat. However, due to the decoupling interference of multiple signals, developing highly sensitive, and selective sensor are challenging in complex stimuli from an ambient environment. [99,119] Furthermore, integrating multifunctional sensors on an ultrathin soft substrate without interfering with each other is the most challenging task. [63] Han et al. integrated pressure, temperature sensors, and electrode into a multilayered multifunctional sensor. [120] Because the temperature, and pressure sensors were collected based on resistance changes of metal wire, the author minimized the interference across two different sensors by configuring temperature sensor arrays as a flat layout to avoid significant geometric changes under normal pressure. This ultrathin multilayer sensor array could be attached to an endocardial balloon catheter and utilized to map temperature, and pressure on the tissue surface. Based on this result, the author argued that the instruments incorporating multimode and multiplexing capabilities could perform better during surgery, and result in better patient outcomes. In other words, researchers attempt to minimize sensory interference and maximize sensitivity with optimized materials, sensors, and manufacturing technologies. This effort was possible with novel sensory applications with ultrathin-compliant structures, and unique mechanical structures against durability.
Many ultrathin sensors reviewed in this paper utilized conventional wired connection systems for human physiological signal monitoring. Some studies have tried integrating ultrathin soft sensors with ultrathin soft wireless transmission systems. [3,5,11,57,[59][60]75,80] They fabricated ultrathin soft circuits for wireless communication similar to the fabrication of ultrathin soft sensors using the microfabrication method [3,5,11,57,75] or the aerosol jet printing method. [60] By integrating ultrathin soft sensors and circuits, they could wirelessly monitor various human physiological signals such as ECG, EMG, GSR, sodium, etc. In addition, Behfar et al. reported a fully integrated, stretchable wireless ECG system developed by screen printing conductive ink on an ultrathin TPU (100 µm) sheet without a rigid or flexible interposer. [61] However, recent studies have shown replicated use of rigid components such as a Bluetooth-enabled MCU (nRF52832, ARM Cortex-M4), an analog-to-digital converter for biopotentials (ADS1299), singlechannel ECG frontend (MAX30003), and ceramic antenna for wireless monitoring. Flexible, stretchable transistors [100,121] and antennas [122] that have been rapidly developed can replace these rigid components. Lee et al. designed and fabricated a thin, flexible organic electrochemical transistor (OECT) for long-term electrophysiological monitoring. [100] The authors showed that the flexible OECT-based sensor could reliably measure ECG signals without rigid microchips, even in dry skin conditions. In addition, significant progress has been made in flexible and stretchable antennas through various strategies. [122] The methods include using flexible materials such as liquid metals, nanowires, and woven textiles or configuring 2D/3D structures such as serpentines and helical coils. Therefore, it is expected that a wireless ultrathin soft sensing system can be developed by incorporating these cutting-edge technologies.
Recently, several researchers have suggested self-powered ultrathin sensors based on piezoelectric [8b,41,43-45,47,48] or triboelectric effect. [50b,52b,53-56,92] However, these self-powered sensors cannot be applied to the wireless system based on Bluetooth and Wi-Fi that needs a constant external power supply. This is because these sensors cannot generate power constantly. Instead of a power generation strategy, Herbert at al. [20,22,27] demonstrated the information from the ultrathin capacitive sensors can be wirelessly transferred without a complex circuit and battery to the external computational device using an LCR coupling system. However, this method has a limitation that cannot be applied to other ultrathin soft sensors except capacitive ones. Meanwhile, Wu et al. [123] presented an ultrathin, soft, and biocompatible sweat-activated battery that can provide sufficient power to drive wearable electronics and wireless transmission to smartphones via Bluetooth. Therefore, by developing these various strategies for self-powered systems, the monitoring of human physiological signals will be further enhanced if the ultrathin soft sensors are combined with wireless communication and self-powered systems.

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Chanyeong Choi is currently a B.S. student in the George W. Woodruff School of Mechanical Engineering, Georgia Institute of Technology, Atlanta, GA, USA. His research interests include brain-machine interfaces and wearable devices.