A Multifunctional Bioelectronic Device with Switchable Rigidity and Reconfigurable Shapes for Comprehensive Diagnosis

Soft bioelectronic devices have potential in medical applications, especially comprehensive diagnoses that rely on immovable equipment to acquire multiple physiological signals. Challenges for bioelectronic devices lie in the opposite modulus demand for multifunctional integration and biocompatibility, and shape compatibility for reducing implantation trauma and covering target tissue. Here, this work reports a multifunctional bioelectronic device endowed with switchable rigidity and reconfigurable shapes by a fast thermal response shape memory polymer substrate. The customizable substrate has a melting temperature around body temperature (≈38 °C) and an ≈100 times modulus drop (from ≈100 MPa to ≈700 kPa). The switchable rigidity realizes stacked layers fabrication in the rigid state and self‐adaptive contact in the soft state. Shape reconfiguration allows the device to be implanted through small incisions and recover in limited spaces to envelop biosurfaces. Multiple physical, biochemical, and electrophysiological signal sensing functions are integrated in the device, realized by temperature and pH electrodes with linear, stable, and fast responses and a low‐impedance potential electrode. For some epilepsy and pericardial effusion types that a single physiological parameter fails to distinguish, animal experiments demonstrate efficiency in their comprehensive diagnosis with electrocorticogram assisted by brain temperature and electrocardiogram assisted by pericardial fluid pH, respectively.

therefore, the higher rigidity of the device may cause interface mismatches and even tissue damage. On the other hand, biocompatibility entails soft substrates that are similar to human tissue to reduce interface constraints and form conformal interfaces, which is unsuitable for fabricating and implanting integrated devices. In addition, bioelectronic devices with fixed shapes still have difficulty in both being implanted through small surgical incisions and unfolded to envelop 3D biosurfaces. Previous works have attempted to provide potential solutions (Table S1, Supporting Information). Stents without substrates [9] can reduce surgical trauma, especially for tubular targets such as vessels. Unfortunately, their mesh structures are unsuitable for multifunctional integration and biosurfaces with arbitrary topographies. Researchers have developed bioelectronic devices integrated with passive soft substrates to improve biocompatibility. [10] Although the substrates have large deformation behaviors, it is challenging for users to maintain the device compressed throughout the minimally invasive implantation process, and adjust the device to form a conformal interface after implantation. Methods such as 3D printing directly reconstruct passive soft substrates into the shapes of target organs; [11] however, the customization limits free deformation behaviors during implantation and the general use of these devices in other patients. Bioelectronic devices with active substrates [12] have shape and/or modulus responses to stimuli, and their temporary shapes can be fixed after reconfiguration, which can be applied to address these trade-offs. Currently, adverse properties of active substrates such as complex synthesis procedures, harmful stimuli, and slow responses remain to improve, and bioelectronic devices for comprehensive diagnosis applications have rarely been studied.
Here, based on a polyurethane shape memory polymer (SMP) substrate, we report a multifunctional bioelectronic device with switchable rigidity and reconfigurable shapes for comprehensive diagnosis, as shown in Figure 1a. The intelligent substrate has customizable rigidity switch properties, i.e., a melting temperature (T m ) near body temperature (≈38 °C) and an ≈100 times modulus drop from ≈100 MPa to ≈700 kPa, which is achieved by one-stage polymerization. At room temperature (T < T m ), the SMP substrate in the rigid state enables the bioelectronic device to be stably high-precision fabricated through 2D methods. The device is fixed as a compressed temporary shape to reduce surgical trauma during implantation. Driven by body temperature (T > T m ), the device softens in 2 s and conformally contacts the 3D target biosurfaces. The cellular structure of the substrate reduces interface constraints and prompts biofluid exchange, maintaining the internal environment of the human body. [10d] Multiple sensing functions of physical (temperature), electrophysiological (electrocorticogram, ECoG, and electrocardiogram, ECG), and biochemical (pH) signals are integrated in the bioelectronic device. We demonstrate detailed comprehensive diagnosis studies through animal experiments, including diagnosis of ischemic, primary, and hemorrhagic epilepsy with ECoG and temperature signals, and of inflammatory, bloody, and noninflammatory pericardial effusion with ECG and pH signals. The proposed device is a portable medical tool that obtains accurate diagnoses by combining multiple physiological signals, offering a novel option for digital medicine. The application of the SMP substrate demonstrates a solution to balance multifunctional integration and biocompatibility, and enhances the diversity of experimental and surgical operations in limited spaces to reduce surgical trauma.

Design and Fabrication
The bioelectronic device integrates multiple physical, electrophysiological, and biochemical signal sensing functions, which are realized by the electrical components of a temperature electrode (TE), a potential electrode (PE), a pH reference electrode (pH RE), and a pH working electrode (pH WE). Accordingly, the entire device is formed by stacking multiple layers, as shown in Figure 1b. The conductive layer (Cr/Au) is packaged by polyimide (PI) films on both sides, but has exposed pads for chemical modification, signal acquisition, and connection to outer wires. Modified with sensitive layers, the multifunctional device works as follows. i) The resistance of the totally packaged TE has a linear response to temperature, and the realtime temperature is monitored by its resistance change. ii) The exposed pad of the PE establishes an electrochemical interface responding to electrophysiological activities, and the potential difference between the PE and the selected reference point corresponds to the electrophysiological signal. iii) The exposed pad of the pH WE modified by a polyaniline (PANI) film linearly responds to the H + concentration, while that of the pH RE modified by a composite film (Ag/AgCl, electrolyte paste, ion selective membrane) is rarely influenced by H + , so the potential difference between the pH WE and RE corresponds to the pH value. v) The cellular SMP substrate supports the layers with electrical functions. The SMP has customizable melting temperature and modulus of the soft state, which are close to those of the human body (≈38 °C; ≈700 kPa). These properties endow the device with thermally driven switchable rigidity and reconfigurable shapes, allowing the device to conformally contact to target organs and form specific temporary shapes. Moreover, the cellular structure facilitates less constraint and normal biofluid diffusion for in vivo applications. Figure 1c presents the robust, general fabrication process utilizing the thermally driven rigidity switch of the SMP substrate. Two photolitho graphy steps define the Cr/Au conductive layer and Cu mask into patterns. The substrate achieves stronger adhesion in the soft state than in the rigid state ( Figure S1a, Supporting Information), and the modulus effect and the shapelocking effect of the SMP material enhance the adhesion for transfer printing. [13] The next step is patterning encapsulation under the protection of Cu mask, which promotes the deformability of the cellular substrate. Finally, the pH RE and WE are chemically modified. An Ag/AgCl film is screen printed and covered by a UV-curable paste containing KCl electrolyte and an ion selective film of Nafion to stabilize the pH RE potential in biofluid. The rigid substrate at room temperature ensures that the device does not deform in response to experimental operations, which contributes to high precision chemical modification steps, e.g., alignment for screen printing Ag/AgCl paste. PANI is electroplated on the pH WE for a linear response to H + . Experimental details are described in the Methods section. Figure 1. Overview of the multifunctional bioelectronic device with switchable rigidity and reconfigurable shapes for comprehensive diagnosis. a) Schematic showing integrated components, thermally driven rigidity switch and shape reconfiguration for a conformal interface (left) and potential intracranial and cardiac applications (right). b) Layout of the stacked layers in an exploded view. c) Robust, general fabrication process: i) laser cutting the SMP substrate into a cellular structure; ii) patterning Cr/Au by photolithography; iii) patterning Cu mask by photolithography; iv) picking up multilayers with a polydimethylsiloxane (PDMS) stamp; v) printing multilayers on the SMP substrate in the soft state; vi) patterning encapsulation by RIE; vii) removing Cu mask by wet etching; viii) screen printing Ag/AgCl for the pH RE; ix) electroplating PANI for the pH WE. Photographs of the d) initial planar shape and e) reconfigured wave temporary shape in the rigid state. Scale bar, 1 cm. f) Conformal contact with an ellipsoid nondevelopable surface in the soft state.

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The fabricated device initially has a planar shape (Figure 1d). To minimize implantation trauma, a specific temporary shape (e.g., the wave shape in Figure 1e) can be realized by taking advantage of the shape reconfiguration. Driven by body temperature, the device in the soft state forms a self-adaptive conformal interface for nondevelopable surfaces in the human body (e.g., the ellipsoid in Figure 1f). Figure S1b (Supporting Information) demonstrates this thermally driven process in detail. When heated to a temperature greater than T m , the SMP substrate switches from the rigid state to the soft state. Thus, the entire device fits into a mold to be reconfigured into a specific temporary shape, which is fixed when the SMP substrate cools and returns to the rigid state. The temporary shapes render the device compatible for various operations, such as implantation through incisions of different sizes. During the shape recovery process, the SMP recovers to the initial shape after heating, which is self-adaptive for conformally contacting any surface due to its low modulus. The device has a fast thermal response time of ≈2 s, and this property is maintained after several shape memory cycles (Movie S1, Supporting Information).
One of the drawbacks of bioelectronic devices with fixed modulus and shape is that the pursuit of soft mechanical properties close to those of human tissue sacrifices the holding convenience for implantation. Here, the SMP substrate is sufficiently rigid to support its temporary shape at room temperature and sufficiently soft to envelop the target biosurfaces driven by body temperature ( Figure S1c, Supporting Information). Furthermore, shape reconfiguration is beneficial in minimally invasive surgery. In vitro experiments show three types of reconfigured shapes implanted through small incisions in pig skin, which ranging from simple to complex shapes, i.e., compressing the device along its width (Movie S2, Supporting Information), along its length (Movie S3, Supporting Information) and as a spiral (Movie S4, Supporting Information). The device maintains the reconfigured shapes throughout implantation, and then recovers its initial planar shape due to the ≈38 °C water, which simulates subcutaneous tissue. Compared to passive elastomeric substrates, the SMP substrate enables complex shape reconfiguration for implantation (Movie S5, Supporting Information). Thin PDMS is inappropriate because it tends to roll. The PDMS substrate that is as thick as the SMP cannot maintain the reconfigured shape along its length, especially in devices with large length to width ratios. Moreover, it's almost impossible for passive elastomeric substrates to be reconfigured as spirals.

Material and Mechanical Optimization for Conformal Interface
Strong interface constraints induce inflammation, immune responses and permanent damage to human tissue. [14] Rigidity switch realizes self-adaptive contact with surfaces with complex morphologies. Furthermore, we optimize the customizable SMP material and its cellular structure to achieve a self-sustainable, less-constrained conformal interface for soft human tissue in dynamic deformation.
The SMP has a simple synthesis procedure of onestage polymerization of polycaprolactone diol (PCL), poly(hexamethylene diisocyanate) (PHMD), and hexamethylene diisocyanate (HDI). The microscopic mechanism of the rigidity switch is the phase transition of crystals in the SMP. The rigidity switch is a macroscopic reflection of crystal melting. As the SMP cools, its shape can be reconfigured until the crystallization process is completed. In experiments, the main rigidity switch properties (T m and the modulus of the soft state) can be optimized by adjusting the molar ratio of HDI and PHMD. The rigidity switch properties have a wide range of values, so we can customize the substrate according to the physiological properties of the target organs ( Figure S2a,b, Supporting Information). We choose the SMP sample with a T m value near body temperature (≈38 °C), as shown by the differential scanning calorimetry (DSC) result in Figure 2a. The dynamic mechanical analysis (DMA) result in Figure 2b demonstrates a drop of ≈100 times in the storage modulus-temperature curve, indicating that the elasticity of the SMP undergoes a switch from rigid (≈100 MPa) to soft (≈700 kPa). The synthesis details are described in the Methods section.
In addition to self-adaptive contact of the device in its soft state with target tissue, the conformal interface is maintained without an external load after the switch from rigid to soft, which enhances the accurate acquisition of biosignals. From the energy point of view, two key factors are taken into consideration for qualitative analysis, namely deformation energy U deformation and adhesion energy U adhesion . For bending and stretching deformation of the device from the initial planar shape to conformal contact, U deformation can be described as U deformation = U bending + U stretching , where U bending and U stretching scale with the effective modulus E and thickness h with the forms of U bending ∼ Eh 3 and U stretching ∼ Eh. Meanwhile, U adhesion increases when the SMP undergoes rigidity switch, and our experimental results in Figure S1a (Supporting Information) prove that the interface exhibits stronger adhesion. In addition, the influence of environmental conditions on adhesion is negligible due to the temperature range of room temperature to body temperature and biofluid for in vivo experiments. For quantitative analysis, previous works [15] have developed theories to discuss the conformability of the epidermal electronics contacting to soft tissues from the energy point of view. The schematic in Figure 2b exhibits the dominant energy in the thermally driven rigidity switch of the SMP. For a substrate with fixed thickness, the rigidity switch decreases U deformation and increases U adhesion ; thus, the relationship U deformation > U adhesion changes to U deformation < U adhesion , which demonstrates the achievement of a self-sustainable conformal interface.
The SMP substrate is modified to have a cellular structure to alleviate the interface constraint, [16] and finite element analysis (FEA) provides evidence for this optimization. A planar SMP and a cellular SMP (porosity φ = 40% and angle θ = 60°) with the same outer contour (7 mm × 7 mm) and thickness (0.3 mm) are perfectly bonded to the top surface of a solid skin (outer contour 10 mm × 10 mm, 1 mm thick). The material parameters and FEA settings are described in the Methods section. For 15% stretching strain applied to the skin along the x direction, Figure 2c shows the interfacial shear stress τ xz between the skin and the SMP substrates with planar and cellular structures, shown in the left and middle views, respectively. The τ xz between the planar SMP substrate and skin exceeds the normal skin sensitivity of 20 kPa. [8b,17] In contrast, the τ xz of the cellular www.advelectronicmat.de substrate is well below 20 kPa, which yields less constraint on the skin than the planar substrate, especially at the contact boundary of the inner unit (zoomed-in view on the right). The skin under 120° bending strain around the y axis is also analyzed. Figure 2d demonstrates larger regions affected by interfacial shear stress compared to the 15% stretching strain, but the cellular structure (middle view) yields a smaller interfacial shear stress value and region than the planar structure (left view). The zoomed-in view shows almost no constraints on the skin. Meanwhile, FEA is also used to evaluate the tensile stiffness EA of the SMP with different structures ( Figure S3a, Supporting Information). The slope of the tension-strain curve in Figure S3b (Supporting Information) is the tensile stiffness EA. Remarkably, the cellular structure significantly decreases the EA of the SMP even under integration with serpentine multilayers. The softer substrate with a smaller tensile stiffness responds better to stretching deformation. Furthermore, the open cellular substrate retains the natural diffusive and convective flows of biofluids, which is important for the normal functions of tissue and organs. [10d]

Intracranial Comprehensive Diagnosis of Epilepsy
Epilepsy is a common neurological disease with various causes: complications of stroke (poststroke epilepsy, PSE), including ischemic and hemorrhagic types, which are related to changes in blood flow in the brain; factors such as genes (primary epilepsy); and symptoms of brain lesions, injuries, etc. [18] As abnormal neuronal activity in the brain, epilepsy can intuitively be visualized through electrographic methods such as ECoG, but we fail to distinguish the types with their similar electrical discharges. Almost all cerebral processes are sensitive to temperature fluctuations, and cerebral blood flow mainly influences brain temperature. [19] We propose the principle of intracranial comprehensive diagnosis of epilepsy in which the ECoG is monitored with the assistance of temperature signals. Thus, TE and PE are integrated onto the cellular SMP substrate, as shown in Figure 3a The device is modeled under deformation far beyond actual service conditions to estimate its structural stability. The FEA results demonstrate that the maximum principal strain of is far less than its yield strain (0.3%) under 180° bending strain (Figure 3c). In contrast, max Device ε under the same deformation is ≈200 times larger than max Au ε ( Figure S4a, Supporting Information). Stretching deformation is indispensable for a developable device fabricated in a 2D process to conformally contact the nondevelopable 3D cerebral cortex, and FEA results similar to those under bending strain are observed under stretching strain ( Figure S4b,c, Supporting Information).
The integrated TE has a linear, stable, and fast response to temperature variations. Figure 3d shows its positive temperature coefficient behavior with a sensitivity of 0.183% °C −1 . The TE has a slight resistance drift around body temperature in the range of 30-45 °C, indicating stability during the service term ( Figure 3e). The TE is highly sensitive to temperature changes, with a response time within 3 s when contacting a cold or hot source (Figure 3f). The interfacial impedance is a vital parameter for electrophysiological recording in the brain, typically the value at 1 kHz. [20] Electrochemical impedance spectroscopy (EIS) in the frequency range of 0.1-100 000 Hz is applied. As shown in Figure 3g, the integrated PE has a competitive impedance modulus of 1.35 kΩ at 1 kHz. Compared to electrodes with chemical modification, [21] our work offers a simple method for electrophysiological signal recording. Meanwhile, the real and imaginary parts of the complex impedance can be calculated with the impedance spectrum ( Figure 3g) and phase spectrum (Figure 3h) to quantitatively characterize the electrochemical interface. After ten bending cycles, the integrated TE and PE properties are maintained ( Figure S5, Supporting Information), indicating excellent electrical and mechanical stabilities for intracranial comprehensive diagnosis of epilepsy.
Our work applies an animal experiment to determine the characteristics of epilepsy in the ECoG. The experimental setup is shown in Figure S6a (Supporting Information). Commercial Characterization of the temperature electrode: d) calibration plot, e) stability, and f) dynamic response. Characterization of the potential electrode: g) impedance spectrum and h) phase spectrum. Error bars represent the standard deviation of three samples.
www.advelectronicmat.de potential electrodes, as a comparison, are inserted into cells encircled by the corresponding PEs. Kainic acid (KA) solution is used to induce epilepsy, [22] and the cellular structure of the SMP substrate shows its advantage in drug delivery. In general, the ECoG signals recorded by the commercial potential electrode and PE in this work have similar waveforms ( Figure  S6b,c, Supporting Information) and frequency spectra ( Figure  S6d,e, Supporting Information), but the signals of the latter are slightly weaker than those of the former because of the depth of electrode implantation. For ECoG signals in the time domain, epileptic waves have an intermittent discharge with high amplitude and frequency compared to normal waves. The frequency spectrum ranging from 5 to 30 Hz proves this enhancement, especially around the frequencies of 12.5 and 25 Hz.
In the in vivo experiment for intracranial comprehensive diagnosis of epilepsy shown in Figure 4a, we establish three groups of animal models: ischemic type with a temperature decrease, primary type without any temperature change, and hemorrhagic type with a temperature increase. The key steps in the animal experiment are described in Figure 4b, and all groups undergo a KA injection step. Carotid arteries on both sides are ligated to reduce blood flow in the brain to form the ischemic type, while ruptured vessels on the surface of the cerebral cortex simulate the hemorrhagic type. Note that the two parameters, electrical discharge and temperature change, are established in sequence by drug injection and operations on the relative vessels. Rather than invading brain tissue, the implantable device balances signal quality and biocompatibility. The device is convenient to apply to the surgical window due to its initial rigid planar shape. After that, saline at 38 °C drives the SMP substrate to switch from rigid to soft, constructing a conformal interface to perfectly cover the cerebral cortex (Figure 4c; Movie S6, Supporting Information).
As shown in Figure 4d, we can hardly distinguish epilepsies of ischemic, primary and hemorrhagic types from ECoG signals, because they are all composed of electrical discharges with high amplitude and frequency. The detailed evidence includes similar waveforms in a zoomed-in view ( Figure 4e) and average amplitudes of each period (Figure 4f) in the time domain, as well as similar characteristic peaks in the frequency spectra ( Figure 4g). Comprehensive diagnosis of epilepsy assisted by brain temperature can overcome this limitation (Figure 4h). When we apply vessel ligation, which results in ischemic type epilepsy, the brain temperature drops to a low level. For the primary type, the brain temperature remains almost constant. Vessel rupture on the surface of the cerebral cortex, which simulates hemorrhagic type epilepsy, causes the brain temperature to increase, and then a decline occurs due to blood clotting. The raw data of the potential of the TE are shown in Figure S7 (Supporting Information).

Cardiac Comprehensive Diagnosis of Pericardial Effusion
The electrical activity of cardiomyocytes, which generate ECG signals, indicates the state of the heart. [23] As a component of the internal environment of the heart, pericardial fluid is sensitive to heart disease, and its abnormal biochemical indicators, such as pH, are clinical evidence for diagnosis. [24] Although pericardial effusions, especially those leading to severe damage such as cardiac tamponade, result in abnormal ECG waveforms, [25] we can hardly conclude nonspecific and insensitive indicators from ECG variables to identify different types of pericardial effusions. The clinical utility of pericardial effusion pH is remarkable for this classification: inflammatory effusions have an acid pH of less than 7.1 (e.g., bacterial); noninflammatory reasons such as hypothyroidism lead to an alkaline pH of more than 7.4; and bloody effusions have a pH of ≈7.3. [26] Based on the above aspects, the principle we propose for cardiac comprehensive diagnosis of pericardial effusion is to analyze the pH while monitoring ECG signals ( Figure 6a).
As shown in Figure 5a,b, ECG and pH functions are integrated on a customized fan shape to envelop the conical heart surface. We rotate the electrode patterns by 90° so that wires connect to the external device on one side, which facilitates the implantation process. The conformal interface of the planar device with the nondevelopable heart surface requires flexibility, including stretching and bending deformations. The FEA result demonstrates that max Au ε is far less than its yield strain (0.3%) under bending strain ( Figure 5c) and stretching strain ( Figure S8b We record the potential difference between the pH WE and pH RE as raw pH signals. PANI serves as the sensitive layer of the integrated pH WE because of the influence of the H + concentration on the state of PANI reflected in potential signals, as well as its features of simple fabrication, controllable structure, and stable electrochemical response. [27] A linear correlation between the potential output of the pH WE and H + concentration is obtained with a negative coefficient of −60.58 mV per pH (Figure 5d). The pH WE also retains a stable potential under pH values ranging from 4.01 to 9.37, which are far beyond the abnormal range of pericardial fluid (Figure 5e). A fast response to pH changes is exhibited in Figure 5f, whether it goes toward or away from a more acidic solution.
The stacked sensitive layers for the pH RE include Ag/AgCl, electrolyte paste and an ion-selective membrane from bottom to top. Ag/AgCl is an easily available, environmentally compatible material for reference electrodes in electrochemical fields. [28] We use KCl to form an electrolyte paste stabilizing the potential of the pH RE. Nafion is chosen as the ion selective membrane to prevent Cl − leakage into the biofluid. [29] The integrated pH RE has excellent stability, minimal potential drift, and high ion selectivity. For pH values ranging from 5.10 to 8.09, as shown in Figure 5g, the potential of the pH RE is maintained at 0 mV relative to a commercial Ag/AgCl RE. Figure 5h shows minimal potential drift over the long term (at least 500 s) under various pH values. We apply different interfering ions to characterize the ion selectivity of the pH RE (Figure 5i). Considering that Cl − determines the potential of the pH RE in accordance with the Nernst equation, we evaluate Cl − first. There is a supersaturated KCl electrolyte in the sensitive layer, so Cl − in the test solution hardly affects the potential. The cation Na + and a different anion SO 4 2− are tested, and the same results as Cl − are obtained.
To record ECG signals, we use the same PE as in the intracranial application, considering the similar electrophysiological mechanisms, stronger amplitude and simpler frequencies of ECG than www.advelectronicmat.de

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ECoG. After ten bending cycles, the properties of the integrated pH WE and RE are maintained ( Figure S9, Supporting Information), showing their excellent electrical and mechanical stabilities for cardiac comprehensive diagnosis of pericardial effusion.
For the in vivo experiment, three groups of animal models of pH 6.8, 7.3 and 7.8 are established to simulate the different types of pericardial effusions mentioned above (Figure 6a). The key steps in the animal experiment are described in Figure 6b. The bioelectronic device is implanted onto the heart surface but under the pericardium. After draining pericardial fluids, a pH buffer is injected to prepare artificial effusions at specific pH values. Finally, the pericardium is sutured.
In the intracranial application, an excellent conformal interface between the implantable device and the target organ is observed. Furthermore, with the reconfigurable shape of the intelligent substrate, we demonstrate that the device can be customized into specific temporary shapes, thereby minimizing surgical trauma. As shown in Figure 6c, a device with a wave shape is implanted into the surgical window. The reconfigured shape is half the size of the initial planar shape, contributing to minimally invasive surgery. The SMP substrate then recovers its initial planar shape driven by 38 °C saline, and constructs a conformal interface that perfectly envelops the heart (Movie S7, Supporting Information). The implantable device is beneficial to patients with relapsing pericardial effusion for reducing the frequency of invasive examinations such as pericardiocentesis.
As shown in Figure 6d, there is no clear evidence in ECG signals to distinguish pericardial effusions of inflammatory, bloody and noninflammatory types due to their similar waveforms ( Figure 6e) and average amplitudes (Figure 6f). ECG www.advelectronicmat.de signals have simple frequency characteristics that are consistent with the heart rate. We count the heart rate for six epochs, and the length of each epoch is 10 s. The results in Figure 6g exhibit similar heart rates of ≈3.6 Hz (216 bpm). The limitation of the ECG renders accurate diagnosis with a single parameter impossible. Assisted by stable pH signals, we can easily distinguish the various types of pericardial effusions with their obviously different values (Figure 6h). Raw data of the potential difference between the pH WE and RE are shown in Figure S10 (Supporting Information).

Conclusion
The multifunctional bioelectronic device with switchable rigidity and reconfigurable shape properties proposed in this work enables comprehensive diagnosis of diseases that can hardly be distinguished by single parameters. The SMP with fast thermal response possesses a customizable rigidity switch, i.e., a melting temperature and a modulus of the soft state close to those of the human body, balancing multifunctional integration and biocompatibility. On the one hand, the patterns on the rigid substrate are stable enough to fabricate stacked sensitive layers for various functions, and the soft substrate produces stronger adhesion in transfer printing. On the other hand, the switch from the rigid to the soft state is selfadaptive and self-sustainable, allowing the bioelectronic device to form conformal interfaces with undevelopable biosurfaces, and the rigid substrate is convenient for surgical and experimental operations. The reconfigurable shape of the bioelectronic device expands its usage in minimally invasive surgery, as shown by in vitro experiments. Moreover, the optimized cellular structure reduces its interface constraints with the target organs and promotes biofluid exchange. In vivo animal experiments demonstrate applications in intracranial and cardiac comprehensive diagnosis. A conformal interface is constructed by the rigidity switch of the proposed bioelectronic device, and we distinguish ischemic, primary, and hemorrhagic epilepsies from their similar ECoG signals with temperature signals reflecting blood flow in the brain. In addition to conformal contact, the reconfigured wave shape of the implantable device reduces surgical trauma when distinguishing inflammatory, bloody, and noninflammatory pericardial effusions with ECG and pH signals.
The bioelectronic device in our work is effective in acquiring multiple physiological signals for comprehensive diagnosis, but for urgent situations, prompt treatment is equally important. Thus, functions such as drug delivery and electrical stimulation need to be integrated to form a closed-loop bioelectronic system. [30] Future work will also include biodegradable materials to further enhance the biocompatibility of the implantable device for long-term use. [31]

Experimental Section
Synthesis and Characterization of the SMP Substrate: PCL (2 g, 2000 g mol −1 ; Sigma-Aldrich, USA) was melted in an oven at 80 °C and then dissolved in butyl acetate (Sinopharm Chemical Reagent Co., Ltd., China). Then, PHMD (0.207 g, Sigma-Aldrich, USA) and HDI (0.059 g, Aladdin, China) were added, followed by dibutyltin dilaurate (0.13 g, 10 wt%, TCI, Japan). After stirring for several minutes, the mixture was poured into an aluminum mold and cured at 80 °C overnight to form a film. Finally, the synthesized SMP film was cut into the cellular structure by an automatic laser cutting machine (VLS2.30, Universal, USA). Other formulas are listed in Figure S2c (Supporting Information).
DSC (Q2000, TA Instruments, USA) analyses were conducted at a heating rate of 10 °C min −1 . Before the heating process, a cooling process from 90 to 0 °C was performed at the same rate, thereby minimizing the differences in the thermal history during the synthesis of different samples. DMA (Q800, TA Instruments, USA) was applied under multifrequency, strain mode (parameters: frequency 1 Hz; strain 0.2%; heating rate 3 °C min −1 ).
Fabrication of Patterned PI-Metal-PI-Mask Multilayers: A Si wafer served as a temporary substrate for PI-metal-PI multilayers, and poly(methyl methacrylate) (PMMA) (MicroChem, USA) was spin-coated on it at 3000 rpm for 40 s and cured at 180 °C for 20 min to form a sacrificial layer. The fabrication process for the bottom PI encapsulation (Furunte, China) included spin-coating at 3000 rpm and step-heating from 80 to 250 °C. Conductive layers of Cr/Au (10 nm/150 nm) were deposited on the bottom PI by electron beam evaporation. The conductive layers were patterned by photolithography and wet etching. Top PI encapsulation involved a similar fabrication process of spincoating at 5500 rpm and then step-heating from 80 to 195 °C. A Cu layer (100 nm) was deposited and patterned as a mask for etching the PI layers. Then, the cellular SMP substrate was heated to the soft state, and the multilayers were transfer printed from the Si wafer onto it by a polydimethylsiloxane (PDMS) stamp. When the SMP recovered its rigidity, reactive ion etching (100 W, 2 h, oxygen flow) was conducted to pattern the top and bottom PI layers. After that, the Cu mask was removed by wet etching. Finally, the exposed metal pads were connected to anisotropic conductive film (ACF) wires.
Fabrication of the Sensitive Layers for pH Signals: Stacked sensitive layers were used for the pH reference electrode. Ag/AgCl paste (5874, DuPont, USA) was screen printed on the exposed metal pad and cured at 120 °C for 15 min. Then, a mixture of UV-curable glue (3105, Loctite, USA) and KCl powder at a weight ratio of 1:1 was pasted and cured under UV light for 2 s. [32] A drop of Nafion (5 wt%, Sigma-Aldrich, USA) was used to cover on the paste and dried at room temperature overnight.
The electrochemical experiments were conducted on an electrochemical workstation (CS350H, CorrTest, China). A threeelectrode setup, including the prepared electrode in this paper as the working electrode, a Pt counter electrode, and a commercial Ag/AgCl reference electrode (218, INESA Instrument, China), was used. PANI was electroplated on the pH working electrode by cyclic voltammetry (CV). The plating bath contained HCl (1 m) and aniline (0.1 m). The CV experiment was conducted from −0.2 V to 1.2 V at a scanning rate of 50 mV s −1 for 10 cycles. The electroplating layer was heated at 60 °C for 1 h.
Characterization of the integrated electrodes: Characterization data of the temperature electrode were obtained in a thermostatic water bath. EIS experiments of the potential electrode were performed in a phosphate buffered saline (PBS) solution with a three-electrode setup, which represented the electrochemical interface in body fluid. The test frequency ranged from 0.1 Hz to 100 000 Hz. Characterization data of the pH working/reference electrode were obtained in a closed loop with a commercial Ag/AgCl reference electrode in standard pH solutions, which were prepared by adjusting PBS solutions with NaH 2 PO 4 and Na 2 HPO 4 powders except for the potassium hydrogen phthalate solution to obtain a pH of ≈4 and sodium tetraborate solution to obtain a pH of ≈9. The solutions in the ion selectivity experiment were KCl (0.1 m), NaCl (0.1 m), and Na 2 SO 4 (0.1 m) in sequence.
Finite Element Analyses: FEA was performed with the commercial software Abaqus. In a previous work, [4c] the SMP was described as an isotropic hyperelastic material obeying the Neo-Hooke law (C 10 = 0.05, D 1 = 0). The device was modeled as S4R elements, and the PI-metal-PI multilayers and sensitive layers were simplified as a composite layup of PI-Au-PI due to the other layers being negligible in thickness or area.

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The elastic modulus, Poisson's ratio and thickness of the isotropic linear elastic PI were given by E PI = 2.5 GPa, ν PI = 0.34, h PI bottom = 6 µm, and h PI top = 2 µm. The isotropic elastic-plastic Au had material parameters of E Au = 78 GPa, ν Au = 0.44, σ yield = 234 MPa, and h Au = 150 nm. [4c] Displacement was applied to the boundaries of the SMP substrate, realizing stretching and bending deformations of the entire device. In the comparison of the interfacial shear stress between planar and cellular SMP substrates, they were perfectly bonded to skin, which was described as a linear elastic material (E skin = 130 kPa, ν skin = 0.48) and modeled as C3D8H elements. [16] Displacement was applied on the boundaries of the skin.
In Vitro and In Vivo Experiments: A petri dish full of ≈38 °C water was covered with pig skin, on which an incision ≈1 cm in length was prepared, and then the device with a reconfigured shape was implanted into the incision.
In vivo animal experiments were performed at Medical Services Biotechnology Co., Ltd. (Beijing, China) and approved by the Ethics Committee of Medical Services Biotechnology Co., Ltd. (MDSW-2022-017C for intracranial application; MDSW-2022-029C for cardiac application).
For the animal model of epilepsy, New Zealand rabbits were anesthetized by intraperitoneal injection of xylazine hydrochloride (5 mg kg −1 ), and anesthesia was maintained with isoflurane (2%)/oxygen, followed by skin preparation and disinfection before the operation. The rabbit underwent skin incision and dissection through the subcutaneous tissues until the skull was visible. Part of the skull was removed to form a window for implanting the device. After a conformal interface between brain tissue and the implantable device was established, KA solution (30 µL, 5 µg µL −1 , Shanghai Yuanye Bio-Technology Co., Ltd, China) was injected into the brain tissue to induce epilepsy. For the ischemic group, carotid arteries on both sides were ligated to reduce the brain temperature. For the hemorrhagic group, vessels on the brain surface were ruptured to increase the brain temperature.
A similar operation was applied for the surgical window on the chest. After opening the pericardium, the original biofluid was drained. A device with an area-saving wave shape was implanted, followed by a thermally driven process to construct the conformal interface. A standard pH solution was dropped into the surgical window, and then, the pericardium was sewn.
ECoG and ECG signals were recorded by an amplifier (UEA-16BZ, Symtop, China). A ground electrode was placed around the surgical window, and a reference electrode was placed at the shaven ear. Temperature signals were recorded by a DC resistance meter (TH2515, Tonghui, China). pH signals were recorded by a portable electrochemical analyzer (CS100, CorrTest, China).

Supporting Information
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