A PEDOT:PSS‐Based Composite Hydrogel as a Versatile Electrode for Wearable Microneedle Sensing Platforms

Advances in biomarker detection have acclaimed a new era of biosensors that enable continuous monitoring of health status, device miniaturization, and wearability. This transition toward integrated, wearable biosensors has necessitated the co‐development of novel materials that can adequately support the operation of these devices. In this study, a novel type of electrode is presented that is suitable for use in wearable electrochemical biosensors. The electrode is constructed using a biocompatible composite hydrogel and takes the form of a hydrogel microneedle (HMN) patch. It is specifically designed for analyzing interstitial fluid. The HMN electrode is a combination of poly(3,4‐ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS), a highly conductive polymer, and graphene oxide, incorporated into a crosslinked hydrogel network of methacrylated hyaluronic acid. To ensure the successful penetration of the skin, the fabrication process is carefully optimized to create sharp needles. To assess the performance of the HMN electrode, electrochemical tests are conducted using an ex vivo porcine skin model. Additionally, HMN electrode's suitability is demonstrated as the working electrode of a wearable electrochemical biosensor for in vivo measurement using a rat model. The findings highlight the advancement of the HMN electrode array as an alternative to solid microneedles, representing the next generation of polymeric electrodes.


Introduction
The detection of biomarkers within biofluids has played a critical role in the prognosis, diagnosis, and recurrence monitoring of numerous diseases and health conditions. [1,2]Biomarker sensing, or biosensing, offers valuable insights into the existence of diseases, as well as individualized information regarding underlying or subclinical conditions, morbidity, and overall trends in the health status of a patient. [2]As such, biosensors have been increasingly developed for the purposes of personalized medicine and health tracking and a shift in the approach to biosensor development has emerged toward device miniaturization and wearability.
The current state of wearable, transdermal biosensing has seen the development of electrochemical-based sensors.Electrochemical sensors are expected to dominate due to their potential for portability, sensitivity, low cost, rapid and selective analysis, simple operation, and high diversity of electroanalytical techniques. [3]lectrochemical sensors have been demonstrated for monitoring of electrolytes, metabolites, pathogens, and hormones in a variety of biofluids. [4]The miniaturization of these sensors has enabled their usage in transdermal, wearable technologies.
The most common biological fluids that are analyzed for disease detection include blood, urine, and saliva. [1]However, in common practice, these biofluids are extracted and then analyzed externally in a laboratory setting. [5]Dermal interstitial fluid (ISF) has been gaining traction as an alternative biofluid for biomarker detection due to its highly similar composition compared to that of blood plasma. [6]8][9] Additionally, ISF analysis allows relatively painless detection of these analytes in the body due to its extraction at skin depths < 1 mm. [1,5,10]However, one enduring challenge of ISF analysis is the longer extraction times and lower volume yields when compared to blood. [5]To surmount this issue, wearable microneedle (MN) devices have been introduced for direct ISF extraction and in situ analysis.
MNs enable direct access to the dermal interstitium, the fluidrich layer beneath the outer layers of skin, in which ISF is readily available. [5]][12][13] Commonly, solid MNs, hollow MNs, and coated MNs have been employed for electrochemical biosensing applications.However, these MNs suffer from costly and heavily involved fabrication processes, as well as performance constraints.Solid and hollow MNs require sophisticated procedures, such as reactive-ion etching (RIE), photolithography, and laser micromachining. [14]Furthermore, solid MN sensing is confined to the surface of the needles in contact with ISF, and hollow and coated MNs are limited in the amount of ISF they can extract.There are additional issues of solid MNs causing skin irritation, swelling and discoloration, [15] as well as hollow MNs being susceptible to clogging. [16]The utilization of conductive hydrogelbased MNs (HMNs) as the working electrode (WE) in these electrochemical sensors can address these shortcomings.
The main advantage of using conductive HMNs is the ease of ISF extraction and their ability to retain the extracted fluid within their conductive polymeric network for in situ electrochemical sensing.The highly absorbent property of hydrogel MNs [17] can maximize the amount of ISF available for interfacing with the sensor due to the uptake of fluid directly into their porous structure.In addition, conductive HMNs use straightforward and inexpensive fabrication techniques, are highly biocompatible, and have tunable strength based on their degree of crosslinking. [18,19]MNs can furthermore offer improved wearability due to their flexibility upon application and ISF extraction.This flexibility is beneficial for the compliance of the wearer and can better accommodate bodily movements.
Here, we present a flexible, biocompatible, conductive hydrogel composite, which consists of poly (3,4ethylenedioxythiophene):polystyrene sulfonate (PEDOT:PSS), a highly conductive polymer, and graphene oxide (GO) entrapped in a methacrylated hyaluronic acid (MeHA)-crosslinked hydrogel network (Figure 1).MeHA HMNs have been shown to exhibit excellent swelling ability, mechanical strength and can maintain structural integrity after achieving swelling saturation. [20]Compared to other hydrogel-based microneedles such as poly(methylvinylether co.maleic acid) (PMVE/MA), MeHA shows excellent biocompatibility and swelling rates for extracting and sensing biomarkers within ISF. [17] The addition of PEDOT:PSS makes the hydrogel composite conductive and suitable as the WE of a three-electrode electrochemical biosensor (Figure 1a). [21]It is a blend of the two polymers, PEDOT and PSS.[24] Charge transport across entangled PEDOT:PSS chains are enabled by the polaron and bipolaron states of the thiophene rings populating the PEDOT backbone. [25]The presence of GO further improves the mechanical strength of the patch [4] while also enabling the option of surface functionalization due to the abundance of chemical functional groups. [26]The reported material (hereafter referred to as MeHA PP-GO) can provide key benefits to electrochemical microneedle systems.These benefits include WE flexibility, and thus improved wearability; good conductivity, and therefore, reliable signal transduction; low fabrication cost; efficient ISF extraction for timely sensing; and ease of surface functionalization for immobilization of biorecognition elements, such as aptamer probes and antibodies. [27]The MeHA PP-GO hydrogel composite requires inexpensive and rapid synthesis procedures (Figure 1b; see Table S1, Supporting Information, for cost estimation).Hyaluronic acid is methacrylated by addition of methacrylic anhydride in an alkaline condition to form MeHA. PEDOT:PSS is synthesized at various ratios via metal-catalyzed oxidative polymerization.By simply mixing PEDOT:PSS and GO with pre-synthesized MeHA and subsequently crosslinking the MeHA network under UV radiation, the PEDOT:PSS and GO become effectively entrapped within its matrix (Figure 1c).PEDOT + oligomers can interact with PSS − chains as well as with deprotonated carboxylic acid groups along the MeHA backbone and formed ionic bonds while PEDOT + oligomers coordinate with longer PSS − chains.GO sheets form pi-pi stacking with PEDOT and PSS while carboxyl groups of GO form hydrogen bonds with the amine and/or hydroxyl groups present in the MeHA hydrogel [27] (Figure 1d).We studied different PEDOT to PSS ratios to fabricate HMN electrodes.The electrochemical performance of the HMN electrode was examined using an ex vivo skin model.Bending and twisting experiments showed that the HMN electrode can retain its performance underbody motion, highlighting the wearability of the electrodes.Finally, we performed a proof-of-concept animal experiment and demonstrated that the HMN electrodes can be potentially employed as the next-generation flexible wearable biosensors for in vivo monitoring.

MeHA PP-GO HMN Fabrication and Characterization
PEDOT:PSS synthesis was carried out based on metal-catalyzed oxidative polymerization and using previously reported protocols [22,28] (Figure 2a).Three different ratios of PEDOT to PSS were prepared based on weight percentage (0.5:0.25, 0.5:0.8,0.5:3.0wt%).Briefly, PSS and DI water were combined, followed by the addition of aqueous solutions of sodium persulfate and iron (III) sulfate at molar ratios to EDOT of 2:1 and 98:1, respectively.The solution was then mixed for 24 h.The purified PEDOT:PSS was then used to prepare MeHA PP-GO HMN electrode.
UV-vis absorbance measurement was performed to characterize the synthesized PEDOT:PSS.The UV and visible absorbance spectra are shown in Figure S1a,b (Supporting Information), respectively.The peak present at ≈ 230 nm corresponds to the presence of aromatic rings in the PSS structure (Figure S1a, Supporting Information), while PEDOT presence is observed in the visible range [29] (Figure S1b, Supporting Information).
We next moved forward and fabricated the HMN electrodes by combing at home synthesized PEDOT:PSS, as the conductive polymer, MeHA, as the hydrogel backbone, and GO, as the filler to improve the mechanical strength of the electrodes.Figure 2b demonstrates the fabrication process using the micromolding technique.Briefly, MeHA, PEDOT:PSS, GO, photo initiator, and the crosslinking agent, N,N′-methylenebisacrylamide (MBA), were mixed to make a homogenous mixture.The mixture was then placed into the mold and centrifugation was done to force the composite into the needle cavities.This was followed by two UV crosslinking steps (wet and dry form) which ensures formation of needles with good integrity.Finally, the HMN patch was attached to a laser induced graphene (LIG), fabricated on polyimide substrate for electrical connection.HMN electrodes fabricated with different PEDOT:PSS ratio was observed and imaged before (Figure 2c, top) and after swelling (Figure 2c, bottom).The needles were able to maintain their integrity upon swelling that confirm the capability of HMN electrodes for in ISF measurement.Sharp needles are important for successful skin penetration.We found that the addition of GO as a filler along with the wet-crosslinking protocol improves formation of sharp needles for all three ratios as shown in scanning electron microscopy (SEM) images (Figure 2d).The HMN electrodes are 850 m in height and have a conical shape.
An important characteristic of HMN patches that enable increased ISF extraction on their swelled needles is their porous structure.Composite hydrogels with higher porous structures have higher swelling capability, allowing increased ISF extraction.We studied the porous structure of MeHA PP-GO samples fabricated with three different PEDOT:PSS ratios.The SEM images of the porous structure of these samples have been shown in Figure 2e.The porosity of 0.25% PSS, 0.8% PSS, and 3% PSS samples were calculated 19%, 32%, and 34%, respectively.The 0.25% PSS sample showed a more dense and less porous network compared to 0.8% and 3% PSS samples.This can be attributed to the lower amount of PSS which is the hydrophilic component.The swelling capability of HMN electrodes fabricated with MeHA PP-GO composite hydrogel and different PEDOT:PSS ratio was also examined using agarose parafilm skin model (Figure 2f).After insertion, HMN arrays were kept on the agarose hydrogel for 15, 30, 60, 90, or 120 min and the swelling ratio was measured based on the weight of patch before and after insertion.The 0.8% and 3% PSS HMN patches showed similar swelling behavior and almost reached to their maximum swelling upon 90 min insertion, however, 0.25% PSS HMN patch showed a lower swelling ratio compared to the other two samples.We also applied 0.8% PSS HMN on the porcine skin and observed a much lower swelling rate of only 64% after 120 min (Figure S2, Supporting Information).The lower swelling property of 0.25% is in agreement with porosity observation (Figure 2e).The capacity of HMN electrodes consisting of various PSS ratios in penetrating the skin was evaluated through compression testing using the dynamic mechanical analyzer (DMA) technique (Figure 2g).The mechanical compression tests performed on all the HMN patches confirmed their robustness, with a mechanical strength exceeding 0.62 N per needle, which is considered sufficient for successful insertion into the skin. [30]

Ex Vivo Electrochemical Characterization of MeHA PP-GO HMN Electrodes
Upon successful fabrication and characterization of HMN electrodes, we moved forward and studied their electrochemical properties.PSS ratio (0.8%) was selected for the fabrication of HMN electrodes due to its ability to produce high-quality needles suitable for skin penetration.Additionally, this ratio offers sufficient swelling capacity for extracting ISF and facilitating sensing.The electrochemical properties of HMN electrodes were examined using an ex vivo porcine skin model.Porcine skins were cut and incubated overnight in a buffer solution containing ferro/ferricyanide as the redox probe.The MN electrodes, including HMN electrode as the WE, Ag/AgCl MN as the reference electrode (MN-RE) and gold coated epoxy MNs as the counter electrode (MN-CE) were punched through the skin loaded with ferro/ferricyanide and the electrochemical measurement was performed using portable potentiostat (Figure 3a).Cyclic voltammetry (CV) scans showed oxidation and reduction peaks at ≈ 0.2 and ≈ 0.1 V, respectively, which aligns with the ferro/ferricyanide redox potentials reported for other PEDOTcontaining electrode materials [31,32] (Figure 3b).CV scans were performed after 30, 45, and 60 min of HMN electrode insertion into the porcine skin and we observed that the scanning reaches to a stable current after 45 min insertion.Based on our previous studies, the increase in the current can be attributed to decrease in the charge transfer resistance during the swelling stage. [33]Next the HMN electrodes were applied to the skins loaded with different concentrations of ferro/ferricyanide redox couple (1-10 mm) and square wave voltammetry (SWV) scans were performed.As shown in Figure 3c, by increasing the redox analyte concentration, the SWV peak was linearly increased, demonstrating the capability of the HMN electrodes to measure the redox analytes effectively.To evaluate the reliability of measuring the MeHA PP-GO HMN signal in the presence of real skin movement and mechanical pressures, we conducted bending and twisting experiments.The effectiveness of the three MN electrode system (WE, RE, CE) was assessed using porcine skin that was loaded with 10 mm ferro/ferricyanide and subjected to 10 cycles of bending (Figure 3d) or twisting (Figure 3e).We found that the deformations had a minimal impact on the average electrochemical signals, indicating that the measurements remained stable despite the skin's movement.We also assessed the biocompatibility of the HMN electrodes' components by exposing them to NIH-3T3 fibroblast cells and conducting an MTT assay (Figure S3, Supporting Information).The findings indicated that the presence of MeHA, GO, or PE-DOT:PSS did not have a significant impact on cell viability, implying that the components of HMN were compatible with biological systems.

In Vivo Characterization of MeHA PP-GO HMN Electrodes
To demonstrate the functionality of MeHA PP-GO HMN electrodes for in vivo monitoring, a rat model was used.The successful insertion of HMN electrodes was verified by conducting hematoxylin and eosin (H&E) staining and visualizing the needle's cavity, revealing an approximate depth of 80 m (Figure 4a).We then applied the HMN electrodes, along with MN-RE and MN-CE on the shaved back of the rat (Figure 4b) and performed electrochemical impedance spectroscopy (EIS) (Figure 4c).We fitted the EIS data to an equivalent circuit as shown in Figure 4d.We employed Cole-Cole bioimpedance model for fitting. [34]In this model, high frequency resistance (R ∞ ) is a composition of both extracellular and intracellular resistance and low-frequency resistance (R 0 ) is made of extracellular resistance. [35]As our microneedles reach to dermis layer which is composed of ISF and connective tissues, R ∞ can be attributed to dermis resistance. [36]hese results are summarized in Table 1.This experiment demonstrates the capability of the flexible HMN electrodes for in vivo monitoring.

Conclusion
In this study, we present a flexible, biocompatible, and conductive HMN array that exhibits excellent swelling and mechanical strength.These electrodes are designed for use in wearable elec- trochemical biosensors to analyze ISF.The fabrication process of the HMN electrodes was carefully optimized to ensure the creation of sharp needles, enabling effective penetration into the skin.By incorporating GO into the composite hydrogel matrix, we not only enhance the mechanical strength of the electrode patch but also enable the immobilization of biomolecules.This immobilization facilitates signal amplification and improves the overall sensitivity of the biosensor.Furthermore, our proof-ofconcept animal testing demonstrated the potential of the HMN electrodes for in vivo measurements, validating their utility.
The developed MN biosensor employs a polymeric structure, which offers several advantages when compared to traditional solid-based MNs.These advantages include reduced manufacturing costs and enhanced compatibility with soft and curved skin.Moreover, the newly developed HMN electrode can absorb a larger volume of ISF within their swollen needles, potentially improving biosensing capabilities.Additionally, it is important to highlight that solid MNs can lead to issues such as skin irritation and discoloration, [15] whereas this HMN biosensor utilizes highly biocompatible materials, further underlining its significance.
It is worth noting that conductive HMNs possess multifunctional capabilities, as indicated by their ability to respond to various stimuli.This feature positions them as promising candidates for controlled drug delivery systems, should there be a future need to integrate therapeutic functions into the sensor.Overall, our study showcases the development of a highly versatile HMN electrode array, which holds promise for advancing the field of wearable electrochemical biosensors for ISF analysis.
MeHA Synthesis: MeHA was synthesized following the previous protocols. [20,27]Initially, 2 g of HA was dissolved in 100 mL of millipore water and stirred at 4 °C overnight until complete dissolution was achieved.Subsequently, 1.6 mL of methacrylate anhydride was added to the HA solution, followed by the addition of 3.6 mL of 5 n NaOH solution to adjust the pH of the solution to the range of 8-9.The mixture was stirred overnight at 4 °C to ensure the completion of the reaction.Next, MeHA was precipitated using acetone and underwent three washes with ethanol.The resulting precipitated MeHA was then redissolved in millipore water and subjected to dialysis for a period of two days to remove impurities.To further purify the MeHA, it was subsequently lyophilized for three days.MeHA (2-5 mg) was dissolved in 1 mL of deuterium oxide (D 2 O) and subjected to testing using a 300 MHz 1HNMR instrument to confirm methacrylate modification.
PEDOT:PSS Synthesis: For synthesizing 0.5:0.25,0.5:0.8, and 0.5:3.0wt% PEDOT:PSS, a 30 wt% stock solution of PSS in water was added to a vial at volumes of 59, 189, and 725 L, respectively.At this stage, additional DI water is added to the vial, which was precalculated in order to bring the final solution volume to 10 mL.Sodium persulfate (Na 2 S 2 O 8 ) was then added at a molar ratio to EDOT of 2:1.Following this, iron (III) sulfate was added at a molar ratio to EDOT of 98:1.EDOT was then added at 40 L and the mixture was left to mix vigorously for 24 h.The resultant crude PEDOT:PSS was purified with anion and cation exchange resins and then vacuum-filtered.The purified PEDOT:PSS was then used to prepare MeHA PP-GO HMNs.
HMN Electrode Fabrication: MeHA (50 mg), MBA (2 mg), PI (2 mg), PEDOT:PSS (192.3 uL), and GO (30 uL) were added to a vial and diluted with DI water (406 uL) for each HMN.After sufficient bath sonication, the MeHA PP-GO mixture was added to MN molds and centrifuged in custom 3D-printed holders to ensure thorough filling of the MN cavities.The wet MeHA PP-GO solution was then crosslinked under UV and dried overnight.The patches were then removed from their molds and exposed to UV again to ensure sufficient crosslinking of the MeHA network.The HMN patch was then attached to the laser-engraved graphene substrate for electrical connection.
Swelling Experiment: Agarose hydrogel (1.4 wt%) consisting was prepared using DI water.Prior to application, the dry mass (W0) of the HMN electrode was measured.The HMN patch was then inserted into the agarose layer through a parafilm barrier and allowed to swell different du-rations.After swelling, the wet mass (Wt) of the HMN patches was measured.The swelling ratio of the HMNs was subsequently calculated using the formula provided below: Dynamic Mechanical Analysis: The mechanical strength of the HMN patches was assessed using an Instron 5548 microtester equipped with a 500 n compression loading cell.To conduct the test, each HMN patch was positioned flat with its tips toward the ceiling.The distance between the two platens was adjusted to 1.5 mm.A vertical force was then applied at a constant speed of 0.5 mm min −1 by the opposing platen.The compression loading cell capacity was set to 70 n.The testing machine recorded the load (force in n) and displacement (distance in mm) every 0.1 s, resulting in the creation of a load-displacement curve for analysis.
Ex Vivo Electrochemical Measurement: To test the HMN electrodes ex vivo, the HMN-WE along with the MN-RE and MN-CE were penetrated through the pre-cut porcine skin model loaded with different concentrations of redox analyte.The CV or SWV measurements were then carried out using a portable potentiostat (PalmSens4).Measurements were repeated at least three times.
Evaluation of Biocompatibility: In vitro cytotoxicity was conducted to assess the biocompatibility of the HMN electrodes using mouse fibroblast cells (NIH-3T3).The cells were seeded in a 24-well plate in a total volume of 100 L.Subsequently, the wells were exposed to a 10 L sample solution of HMN electrodes for a duration of 24 h.As a control, 10 L of Dulbecco's Modified Eagle Medium (DMEM) was used.Following exposure, all wells were treated with 10 L of a 5 mg mL −1 stock solution of Methylthiazolyldiphenyl-tetrazolium bromide (MTT).The 96-well plate was then incubated away from light for a period of 3 h.To break up the cells and release the formazan crystals, 150 L of dimethyl sulfoxide (DMSO) was gently added to the treated wells.The absorbance of the samples was subsequently measured using a spectrophotometer at a wavelength of 540 nm.
In Vivo Animal Experiment: In vivo experiments were done following the Guidelines for the Care and Use of Laboratory Animals and the Ani-malWelfare Act Regulations; all protocols were approved by the University of Waterloo Institutional Animal Care and Use Committee.Male Sprague Dawley rats (Charles River,100-150 gr) were used for the experiment.Under isoflurane, the rat was sedated ad their back was shaved using a shaving machine and application of hair removal cream.Next, HMN electrode as the WE, MN-RE, and MN-CE electrode were applied, and the electrochemical measurement was performed.To perform H&E staining, HMN patches were applied to the shaved dorsal skin of rats for 5 min.After removing the patches, the rat was euthanized, and the skin sections were cut and washed with a NaCl solution.The samples were then fixed in formalin, stored in ethanol, and refrigerated.The fixed skin samples were cryopreserved in a sucrose/PBS solution overnight.Next, the samples were mounted on cork using OCT and frozen using isopentane cooled with liquid nitrogen, and stored at a very low temperature.Thin sections were cut from the frozen samples and placed on microscope slides.Hematoxylin and eosin (H&E) stains were used to examine the basic structure of the skin samples.Images of the stained samples were captured using a Cytation-5 multimode imager at 20× magnification and stitched together using Gen5 software.
Statistical Analysis: Data presented in Figures 2f, 3c

Figure 1 .
Figure 1.Overview of the MeHA PP-GO HMN electrode.a) Schematic of MeHA PP-GO HMN as the WE in a three-electrode electrochemical sensor setup.The counter electrode (CE) is composed of gold-coated epoxy MNs, and the reference electrode (RE) is composed of Ag/AgCl-coated epoxy MNs.An optical close-up of the MeHA PP-GO HMN patch highlights its flexibility.b) A brief summary of the synthesis process for (i) MeHA and (ii) PEDOT:PSS.c) The entrapment of PEDOT:PSS polymer complexes and sheets of GO within the crosslinked MeHA hydrogel network.d) Chemical structures of PEDOT, PSS, MeHA, GO, and their interaction with each other.

Figure 2 .
Figure 2. MeHA PP-GO HMN fabrication and characterization.a) Schematic demonstrating the synthesis procedure of PEDOT:PSS.b) Schematic showing fabrication process of HMN electrodes.c) Optical images of HMN electrodes before swelling (top) and after swelling (bottom).d) SEM images of HMN arrays with different PSS ratios.Scale bar for array = 250 m; inset scale bar = 125 m.e) SEM images of different PSS ratios to characterize the porosity structure.In (c-e); (i-iii) show corresponding images for 0.25%, 0.8%, and 3% PSS ratios, respectively, Scale bar = 10 m.f) Swelling capability of HMN electrodes with different PSS ratios was studied via applying the patches in agarose hydrogel for different time points.Data is shown as the mean ± SD, n = 3. g) Mechanical test experiment for HMN electrodes fabricated with different PSS ratios.

Figure 3 .
Figure 3. Ex vivo electrochemical characterization of MeHA PP-GO HMN electrodes.a) Ex vivo porcine skin electrochemical setup.HMN-WE, MN-RE, and MN-CE were punched through the skin and the electrochemical measurement was performed.b) CV scans were performed at different time points with 100 mV s −1 scan speed and between −0.6 and 0.6 V versus Ag/AgCl MN-RE.c) SWV scans were performed on skin loaded with different concentrations of ferro/ferricyanide between −0.6 and 0.6 V versus Ag/AgCl MN-RE with 15 Hz frequency.Data is shown as the mean ± SD, n = 3. 10 cycles of bending d) or twisting e) were performed on a skin loaded with 10 mm ferro/ferricyanide.SWV scan before and after deformation test did not show significant difference.

Figure 4 .
Figure 4.In vivo electrochemical characterization of MeHA PP-GO HMN electrodes.a) Histology of rat skin after H&E staining.b) In vivo experimental setup demonstrating the MeHA PP-GO HMN, MN-RE and MN-CE applied on the shaved back of the rat.c) Electrochemical impedance spectroscopy of MeHA PP-GO HMN electrode after applying on the live rat skin.The scan was performed at open circuit voltage.d) The equivalent circuit applied to fit the results in c.
, and Figure S2 (Supporting Information) are presented with mean value± SD and n = 3.The data is analyzed via OriginPro 2018.

Table 1 .
Fitting results of impedance spectroscopy obtained from the in vivo experiment.