Quantitative Assessment of Signal Quality and Usability of EEG and EMG recordings with PEDOT:PSS-Coated Microneedle Electrodes

Quantitative Assessment of Signal Quality and Usability of EEG and EMG recordings with PEDOT:PSS-Coated Microneedle Electrodes | Research Square window.SnipcartSettings = { analytics: { enabled: false } }; (function() { var accessVector = localStorage.getItem('access_vector') || ''; window.dataLayer = window.dataLayer || []; if (accessVector) { window.dataLayer.push({ user: { profile: { profileInfo: { snid: accessVector } } } }); } })(); (function(w,d,s,l,i){w[l]=w[l]||[];w[l].push({'gtm.start':new Date().getTime(),event:'gtm.js'});var f=d.getElementsByTagName(s)[0],j=d.createElement(s),dl=l!='dataLayer'?'&l='+l:'';j.async=true;j.src='https://www.googletagmanager.com/gtm.js?id='+i+dl;f.parentNode.insertBefore(j,f);})(window,document,'script','dataLayer','GTM-K279D39R'); Browse Preprints In Review Journals COVID-19 Preprints AJE Video Bytes Research Tools Research Promotion AJE Professional Editing AJE Rubriq About Preprint Platform In Review Editorial Policies Our Team Advisory Board Help Center Sign In Submit a Preprint Cite Share Download PDF Research Article Quantitative Assessment of Signal Quality and Usability of EEG and EMG recordings with PEDOT:PSS-Coated Microneedle Electrodes Tomoya Yamaguchi, Yuta Kurashina, Eiji Nagahara, Naotsugu Kaneko, and 3 more This is a preprint; it has not been peer reviewed by a journal. https://doi.org/ 10.21203/rs.3.rs-7343232/v1 This work is licensed under a CC BY 4.0 License Status: Posted Version 1 posted You are reading this latest preprint version Abstract Background: Bioelectrical signals are vital indicators of physiological function, psychological status, and clinical conditions. However, traditional wet electrodes using conductive gel are time-consuming, labor-intensive, and degrade over time due to gel drying. Recently, microneedle (MN) electrodes coated with poly(3,4‐ethylene dioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) have been developed for high-quality signal acquisition. Although low electrode–skin impedance has been demonstrated on hairless regions, their use in electroencephalography (EEG) recordings on hairy scalp areas and quantitative comparisons with conventional electrodes in both EEG and electromyography (EMG) remain limited. Methods: We fabricated two MN electrode types of different lengths—one for EMG on hairless skin and one for EEG on the hairy scalp—and compared them with conventional wet and dry electrodes. EMG was recorded during a force-matching task; EEG was assessed via somatosensory evoked potentials. We also evaluated setup/cleanup time, comfort, and pain during EEG measurement. Results: For EMG, MN electrodes achieved significantly higher signal-to-noise ratios than conventional electrodes. For EEG, they outperformed dry electrodes and matched wet electrodes in signal quality—without using conductive gel. Although additional time was required to part hair, cleanup was faster due to the absence of gel. Pain was comparable to dry electrodes. Conclusions: PEDOT:PSS-coated MN electrodes provided superior EMG signal quality and high-quality EEG signals comparable to wet electrodes even without gel. These findings suggest that PEDOT:PSS-coated MN electrodes offer a compelling balance of signal quality and user convenience, making them especially advantageous for real-world and clinical applications where time efficiency, minimal discomfort, and gel-free operation are critical. Electroencephalography (EEG) Electromyography (EMG) Brain Muscle Biosignal Figures Figure 1 Figure 2 Figure 3 Figure 4 Figure 5 Figure 6 Figure 7 1. Background Bioelectrical signals are crucial indicators of bodily function, psychological status, and clinical pathologies. Thanks to significant advancements in bioelectric acquisition technology, electrocardiography (ECG), electromyography (EMG), and electroencephalography (EEG) have become essential tools in both clinical and neuroscience fields [ 1 – 3 ]. These bioelectrical signals can be recorded using two approaches: invasive and non-invasive recordings. Invasive recordings, which involve inserting electrodes directly into the body near the target site, yield high-quality signals but are generally restricted to severe clinical interventions due to the invasiveness of the procedures involved [ 4 ]. In contrast, surface bioelectrical signals are now more frequently employed as a non-invasive alternative [ 5 – 7 ]. For instance, surface EEG signals, which are most widely used modality in various fields in the above-mentioned three types of bioelectrical signals, are employed to diagnose epilepsy [ 8 ], to brain–machine interface (BMI) technologies that compensate for impaired bodily functions [ 9 ], and to further neuroscientific research on human cognitive and motor functions [ 10 , 11 ]. Because surface electrodes are placed on the skin to capture bioelectric signals, they essentially function as transducers, converting ionic currents from the human body into electronic currents that are then transmitted to external electronic systems made of various metals [ 12 , 13 ]. For the signal transmission at the electrode–skin interface, the electrode–skin interface impedance (EII) is a crucial parameter. Typically, decreased EII leads to better signal quality, an increased signal-to-noise ratio, and diminished baseline drift [ 14 ]. To lower this impedance during EEG recordings with surface electrodes, a conductive gel containing electrolytes has traditionally been applied between the skin and the electrode [ 13 ]. However, the wet electrode approach is both time-consuming and labor-intensive—requiring gel application and hair washing before or after measurements—and it also limits recording time due to gel drying [ 13 ]. Consequently, balancing ease of use and high-quality signals remains challenging with current wet electrodes, which in turn hinders their widespread adoption for routine clinical use and in real-world applications for health monitoring. In recent years, numerous research efforts have focused on developing dry electrodes to address this problem [ 15 , 16 ]. Dry electrodes without conductive gel are better suited for long-term brain activity monitoring thanks to no gel drying problem and for daily use and real-world applications because of the no need for gel application and hair washing. Because of the high-resistance stratum corneum (SC) and insufficient contact between the skin and the metal electrode surface, dry electrodes generally have higher skin–electrode impedance compared to wet electrodes [ 15 , 17 ]. Various new materials and structures have been proposed to improve the electrode–skin contact [ 18 – 20 ], but the trade-off is the large electrode volume or footprint because the electrode size is critical for high density EEG/EMG recordings, which are required for source estimation of the EEG and EMG signals [ 21 – 23 ]. Recently developed microneedle (MN) electrodes can penetrate the high-resistance stratum corneum by using tiny needles that penetrate the skin with minimal discomfort [ 24 , 25 ], thus enabling high-quality measurements without the need for extensive skin preparation and conductive gel (Fig. 1 ). Because the needles on these electrodes are very small, causing no/slight pain [ 26 ], they are known as micro-invasive electrodes. Recently, it was reported that depositing a conductive polymer, poly (3,4-ethylene dioxythiophene):poly(styrenesulfonate) (PEDOT:PSS), onto metal MNs dramatically reduces electrode impedance [ 27 ]. Thus, these electrodes can be smaller than—thanks to the low impedance per size—conventional wet electrodes. PEDOT:PSS exhibits mixed conductivity, which can conducts both electronic and ionic charges, it functions as a transducer that facilitates smooth current flow between the body’s ionic environment and the electrode’s electron-conductive metal [ 28 ]. A study developing microneedle electrodes coated with PEDOT:PSS found that the electrode–skin impedance in hairless areas (such as the forehead or arms) was much lower compared to wet electrodes [ 27 ]. However, these electrodes have not yet been applied to the scalp, which is covered with hair. In addition, there has been no comprehensive quantitative comparison in EEG and EMG signals with conventional dry and wet electrodes. Although the study by Li et al, (2022)[ 27 ] evaluated EMG signal amplitude and signal-to-noise ratio (SNR) against conventional electrodes, several methodological issues arose. For instance, electrode placement differed between electrode types—even though EMG signals vary by location [ 29 ]—and no standardized muscle contraction level was used, potentially causing inconsistent muscle activity across tasks. In this study, we aimed to quantitatively compare the performance of MN electrodes coated with PEDOT:PSS against conventional wet and dry electrodes. First, to assess electrode performance under hairless conditions, we measured EMG signals and impedance on the arms. Next, we evaluated their performance in EEG signals measured from hairy regions by somatosensory evoked potentials (SEPs). During these measurements, we recorded setup and cleanup times, as well as any pain or discomfort reported by participants, to assess usability across electrode types. 2. Methods Overview of MN electrodes Figure 2 . Fabrication process of PEDOT:PSS-coated microneedle electrodes for EEG and EMG recordings. (A) A polyimide layer formed by thermal curing of polyamic acid. (B) An Au layer deposited via vacuum evaporation. (C) A PEDOT:PSS layer formed by electrodeposition from a monomer solution. (D) Needle height was adjusted according to application—1000µm was adopted for EEG and 600µm for EMG to account for hair thickness. (E) Example of a fabricated microneedle electrode; the one shown is designed for EMG recordings. 2.1 Materials A polyamic acid solution (Pyre-M.L) was purchased from IST Corp. (Japan). The PDMS-based silicone mold for MN electrode (Mpatch Microneedle Template) was obtained from Micropoint Technologies Pte Ltd (Singapore). Au and chromium for vapor deposition were purchased from Niraco inc. (Japan), and PSS (molecular weight of ~ 70,000) along with EDOT (97%, the monomer for PEDOT) were purchased from Sigma-Aldrich LLC (USA). 2.2 Fabrication process Two types of PDMS molds were used to fabricate MN arrays. For EMG experiments, the mold had a needle length of 600 µm, a base diameter of 200 µm, a pitch of 500 µm, and a 10×10 needle array. For EEG experiments, the mold had a needle length of 1,000 µm, a base diameter of 250 µm, a pitch of 500 µm, and a 15×15 needle array, considering hair layer thickness. 2.2.1 Polyimide layer (Fig. 2 A) Each PDMS mold was ultrasonically cleaned (120 kHz with a Ultrasonic Cleaner HFC-3D, AS ONE, Japan) for 30 minutes and then dried. Next, polyamic acid solution was poured into the PDMS mold and centrifuged at 3,000 rpm for 15 minutes to ensure complete filling with a centrifugal machine (H-19a, KOKUSAN, Japan). Additional polyamic acid solution was then applied, and a scraper was used to level it to the mold’s edges. The mold was placed on a hot plate at 100°C for 30 minutes to evaporate the solvent. The temperature was gradually raised to 200°C over a 30-minute period, and then held at 200°C for another 30 minutes to induce imidization and cure the film. Because the PDMS mold can only tolerate up to 200°C, the polyimide substrate was removed from the mold before an additional 5-minute heating step at 300°C, completing the imidization process and yielding the final polyimide layer. 2.2.2 Au layer (Fig. 2 B) Using a vacuum deposition device (VE-2013, Vacuum Device Inc., Japan), we coated both sides of the polyimide layer with a 10 nm titanium (Ti) adhesion layer and a 200 nm Au conductive layer, thereby producing conductive MN electrodes. Deposition was performed separately on each side to ensure conductivity on both the front and back surfaces. 2.2.3 PEDOT:PSS layer (Fig. 2 C) A 2.5 wt% PSS solution and 0.01 M EDOT were dissolved in pure water, and a PEDOT:PSS layer was deposited onto the Au surface via electrochemical polymerization. In this setup, a platinum rod served as the cathode, while the MN electrode (with its gold layer) was used as the anode. Because excessive current can accelerate polymerization too quickly—leading to cracks or poor adhesion in the PEDOT:PSS layer [ 27 ]—the current was increased stepwise at 20, 40, 60, and 80 µA for 150 seconds each, and then 100 µA was applied for 600 seconds. Matlab (MathWorks, USA) was used to control the current, and an NI-9265 (National Instruments, USA) functioned as the constant current source. 2.3 EMG experiment 2.3.1 Participant Eight healthy male adults were participated in this experiment. All participants provided written informed consent for participation in the study, which was conducted in accordance with the Declaration of Helsinki and approved by the Ethics Review Committee of The University of Tokyo (reference number, 23–396). 2.3.2 EMG acquisition Because MN electrodes create small punctures in the skin, measurements were performed in the following order: wet electrode, dry electrode, and then MN electrode. Before each measurement, the skin was cleaned with an alcohol swab and dried to maintain consistent skin conditions. EMG signals were recorded using an electromyography amplifier (BrainAmp-ExG MR, Brain Products, Germany) at a sampling rate of 2,500 Hz. The data were then processed with a 4th-order Butterworth bandpass filter (20–350 Hz), and 50 Hz powerline noise was removed with a notch filter. Two electrodes were placed on muscle belly of the biceps brachii of the dominant arm with a 2 cm center-to-center spacing, and a ground electrode was positioned on the acromion. To minimize external noise, the wires were twisted. The positions of the electrodes were marked to ensure consistent placement for each electrode condition. After attaching each electrode to the arm, it was fixed with an elastic underwrap. All electrodes were standardized to a 5 mm × 5 mm square area. The dry electrode consisted of a round, gold-plated metal body onto which a thin insulating film with a 5 mm × 5 mm square opening was affixed; a rod on the back allowed for clip-connector wiring. To form the wet electrode, we attached a double-sided adhesive urethane foam—also featuring a 5 mm × 5 mm square opening—to the same round, gold-plated metal base used for the dry electrode, then filled the opening with conductive paste (AC cream, OT Bio Elettronica, Italy). The MN electrode was fabricated as a 5 mm × 5 mm MN array consisting of 10×10 needles. A silver-paste-based conductive epoxy (CW2400, Chemtronics, USA) was applied to the back of the needle surface to enable wiring. 2.3.3 Experimental Task Participants performed a force match task of flexion of the elbow joint (Fig. 3 A). Before the main task, maximum voluntary contraction (MVC) force was measured for 3 s. The MVC task was performed before attaching electrode. Thus, the same contraction level was used across all the electrode conditions. Then we recorded 10 seconds of EMG signals under relaxed (rest) conditions. Next, the participant performed elbow flexion at 20% MVC for 10 s. We selected this contraction level because 20% MVC has been frequently used in high-density EMG studies to examine motor neuron firing[ 23 , 30 ], which is a potential application of the MN electrodes. The EMG recordings of rest and the force match task were iterated for each electrode type. 2.3.4 Evaluation of signals quality of EMG and skin-electrode impedance The root mean square (RMS) of EMG amplitude was computed for each condition, and the ratio of these RMS values served as the SNR, in accordance with Eq. (1): $$\:\begin{array}{c}SNR=20{\text{log}}_{10}\frac{\text{R}\text{M}{\text{S}}_{\text{s}\text{i}\text{g}\text{n}\text{a}\text{l}}}{\text{R}\text{M}{\text{S}}_{\text{n}\text{o}\text{i}\text{s}\text{e}}}\#\left(1\right)\end{array}$$ , where RMS signal ​is the RMS value during the force match task and RMS noise​ is the RMS value during the rest task. Additionally, before measuring EMG signals, the EII spectra of the electrodes were measured by a two-electrode system with an Analog Discovery 3 (Digilent, USA) and its control software WaveForms (Digilent, USA). 2.4 EEG experiment 2.4.1 Participant Nine healthy male adults were participated in this experiment. All participants provided written informed consent for participation in the study, which was conducted in accordance with the Declaration of Helsinki and approved by the Ethics Review Committee of The University of Tokyo (reference number, 23–396). 2.4.2 EEG acquisition EEG signals were recorded at a sampling rate of 5,000 Hz using an EEG amplifier (actiCHamp Plus, Brain Products, Germany). The data were then processed with a 4th-order Butterworth bandpass filter (30–150 Hz) [ 31 ], and 50 Hz power-line noise was removed using a notch filter. Following the International 10–20 System, Cz was used as the ground electrode, Fz as the reference, and Cp3 for measuring somatosensory evoked potentials. In this study, three different types of EEG electrodes (Dry, Wet, and MN) were compared. Measurements under these three conditions were conducted by applying gel to a Dry EEG system (actiCap Xpress Twist system) used in neuroscience research [ 32 , 33 ] to create the Wet condition, and by attaching microneedle electrodes to the tip of the Dry electrode to create the Microneedle condition. The details are as follows: 1)Wet Electrode (Fig. A) A slightly convex, dish-shaped, gold-plated metal electrode (actiCap Xpress Twist with Flat QuickBit, Brain Products, Germany). Conductive gel (SuperVisc, Easycap GmbH, Germany) was applied between the metal surface and the skin. 2)Dry Electrode (Fig. B) The same actiCap Xpress Twist system, but equipped with Long QuickBit tips (length: 12 mm) designed for hairy areas. Used without conductive gel. 3)MN Electrode (Fig. C) The same actiCap Xpress Twist system with Flat QuickBit, as in the wet electrode condition. A MN array was affixed to the electrode using conductive double-sided tape (T-9426, ESD EMI Engineering, Japan), and no conductive gel was applied. In the above comparison of electrodes during EMG measurements, the area of skin contacted by the electrodes could be controlled. However, it should be noted that in EEG measurements, such control is difficult due to the influence of hair. Since MN electrodes can create small punctures in the skin, measurements were performed in this order: wet electrode, dry electrode, then MN electrode. Before applying the wet electrode, the skin at the electrode site was cleaned with an alcohol swab. After wet electrode measurements, the conductive gel was washed off, and the scalp and hair were thoroughly dried with a hairdryer before proceeding to the dry electrode measurement. All electrodes were then secured to an elastic cap specifically designed for actiCap Xpress Twist. 2.4.3 Electrical Stimulation for somatosensory evoked potential We administered 200 ms monophasic square-wave constant-current pulses using a clinical stimulator (USE-100, UNIQUE MEDICAL, Japan) to stimulate the right median nerve. We placed a pair of electrodes in positions on the right wrist where the visible muscle twitch was most clearly observed. The current intensity was set to evoke a 1–2 cm thumb movement, following the recommendations of the International Federation of Clinical Neurophysiology [ 34 ]. Stimulation was applied 300 times at a repetition rate of 2.1 Hz. To maintain the participants’ attention on the stimulus, pauses were introduced every 30–50 seconds. Specifically, between 3 and 4 of these pauses—each lasting about 4 seconds—were inserted at random during each session, and participants were instructed to count the pauses and report their totals at the end of the session. All participants accurately reported the number of pauses in every session. To avoid variations in posture or hand position that could affect the experimental results, each participant’s hand and arm placements were marked on the support platform, ensuring a consistent posture across all conditions. 2.4.4 Evaluation of Electrode quality for EEG To assess the quality of EEG signals recorded by MN electrodes compared to conventional (dry and wet) electrodes, we calculated the SNR of somatosensory evoked potentials (SEPs). The SEPs were obtained using a stimulation triggered averaging method, in which signal segments around the timing of electrical stimulation were extracted and averaged. The amplitude of the evoked response was defined as the N20 peak (observed between 19.0 and 21.0 ms following stimulation), and noise was defined as the standard deviation of the pre-stimulation signal from − 70 to -5 ms. The ratio of these two values was taken as the SNR, in accordance with Eq. (2): $$\:\begin{array}{c}SNR=20{\text{log}}_{10}\frac{{\text{A}\text{m}\text{p}\text{l}\text{i}\text{t}\text{u}\text{d}\text{e}}_{\text{s}\text{i}\text{g}\text{n}\text{a}\text{l}}}{{\text{S}\text{D}}_{\text{b}\text{a}\text{c}\text{k}\text{g}\text{r}\text{o}\text{u}\text{n}\text{d}}}\#\left(2\right)\end{array}$$ , where Amplitude signal is a peak amplitude of N20 and SD background is the standard deviation of the pre-stimulation period (from − 70 to -5 ms). Additionally, before measuring EEG signals, the EII at 10Hz of the electrodes were measured with a impedance measurement function of the EEG amplifier (actiCHamp Plus) and its control software (Brain Products, Germany). The analyzed frequency was predetermined by the EEG system. Since 10 Hz corresponds to the alpha wave in EEGs, impedance at 10 Hz has been used to evaluate the biopotential electrodes[ 35 , 36 ]. Due to equipment limitations, any value exceeding 500 kΩ was recorded as 500 kΩ. For many participants in the dry electrode, the impedance did not fall below 500 kΩ. 2.4.5 Assessment of MN Electrode Usability We evaluated (1) preparation and cleanup time, (2) participant comfort, and (3) pain intensity during EEG measurements. Preparation time was measured from the moment the experimenter began fitting the participant with an EEG cap until the electrode–skin impedance fell below 20 kΩ for the wet and MN electrodes, and below 400 kΩ for the dry electrode. In all experiments, the same experimenter (T.Y.) was responsible for both the preparation and cleanup. If impedance for the dry electrode remained above 400 kΩ and did not improve even after parting the hair under the electrode, we stopped recording the preparation time. Cleanup time included removing the electrodes after each measurement. For wet electrodes, it also included the time needed to wash out the conductive gel and thoroughly dry the hair. Comfort and pain were rated for each electrode condition during setup, recording, and removal using a visual analogue scale (VAS). For comfort, the highest value (10) indicated very uncomfortable and the lowest value (0) very comfortable. For pain, the highest value (10) represented the worst pain imaginable and the lowest (0) no pain at all. Participants were asked to consider overall duration, pain, and any discomfort when rating the comfort. 2.5 Statistics We compared the frequency characteristics of the impedance among the electrode types on the hairless skin during the EMG experiment. we first performed a logarithmic transformation considering homoscedasticity. One-dimensional statistical parametric mapping (SPM1d) was used to determine whether impedance differed among the electrodes [ 37 ]. The SPM analysis can find statistically significant frequency portions regarding the impedance difference among the electrodes rather than compare discrete frequency. For multiple comparisons across the three electrodes, p-values were adjusted using the Bonferroni method. We also examined differences among the electrode types for the following parameters: 1) SNR during the EMG experiment, 2) SNR during the EEG experiment, 3) impedance at 10 Hz during the EEG experiment, 4) setup time, 5) removal time, 6) total time (setup + removal),7) comfort ratings during setup, recording, and removal, and 8) pain level. Normality was tested using the Shapiro-Wilk test. The test indicated that normality was not rejected in all parameters except SNR of EMG, impedance during the EEG experiment, and pain level. Therefore, we performed a one-way analysis of variance (ANOVA) with a parametric approach to evaluate the main effect of electrode type for normally distributed data. Post-hoc comparisons were conducted using Student’s t-test if equal variances were confirmed, or Welch’s t-test otherwise. For the parameters, which did not follow the normality, we employed the Fisher–Pitman permutation test of matched pairs with 1,000 permutations for comparisons between electrode types [ 10 , 38 ]. All p-values were corrected for multiple comparisons using the Bonferroni correction. 3. Results 3.1 Overview of fabricated MN electrodes and experiments for comparisons among electrode types Figure 2 illustrates the flowchart of the fabrication process for PEDOT:PSS-coated MN electrodes. Based on a previous study [ 27 ], we fabricated the MN electrode composed of three layers:1) a polyimide substrate layer providing flexibility, mechanical strength; 2) a Au layer to provide electrical conductivity (Fig. 2 B); and a PEDOT:PSS conductive polymer layer (Fig. 2 C), which decrease skin-electrode contact impedance. We then compared the performance—both in terms of signal quality and usability—of the MN electrode with conventional wet and dry electrodes. First, to assess performance under hairless conditions, we recorded EMG signals and electrode impedance on the forearm (Fig. 3 A). Next, we evaluated their performance for EEG signals from hairy scalp regions using somatosensory evoked potentials (SEPs) (Fig. 3 B). During the EEG recordings, we also recorded setup and cleanup times, as well as any pain or discomfort reported by participants, to assess usability across electrode types. 3.2 EMG experiment Eight healthy participants performed an elbow flexion force-matching task (Fig. 3 A). Before measuring EMG signals, the EII spectra of the electrodes was measured. Then, prior to the main task, maximum voluntary contraction (MVC) force was measured. We then recorded 10 seconds of EMG signals under both relaxed (rest) conditions and during elbow flexion at 20% MVC. Figure 5 A shows an example of EMG signals recorded during elbow flexion at 20% MVC (top row) and at rest (bottom row). In this participant, the MN electrode produced a larger amplitude during muscle contraction and a lower noise level at rest compared to the wet and dry electrodes. Consistent with this finding, the averaged data across participants (Fig. 5 B) revealed that the MN electrodes had statistically lower impedance than the dry electrodes below 750 Hz, and lower than the wet electrodes below about 150 Hz (p < 0.05). In Fig. 5 C, the signal-to-noise ratio (SNR) of EMG—calculated as the EMG amplitude during elbow flexion relative to that during rest—was highest on average with the MN electrodes, followed by the wet electrodes and then the dry electrodes. Statistical analysis showed that the SNR with the MN electrodes was significantly higher (p < 0.05) than that observed with both the wet and dry electrodes. 3.3 EEG experiment Nine healthy male adults participated in this experiment. Before measuring EEG signals, the EII at 10Hz were measured with a impedance measurement function of the EEG amplifier (actiCHamp Plus, Brain Products, Germany). The analyzed frequency was predetermined by the EEG system. Since 10 Hz corresponds to the alpha wave in EEGs, impedance at 10 Hz has been used to evaluate the biopotential electrodes[ 35 , 36 ]. We applied 200 ms monophasic constant-current pulses to the right median nerve using an electrical stimulator (USE-100, UNIQUE MEDICAL, Japan). Stimulation (300 trials at 2.1 Hz) was delivered via wrist electrodes at an intensity sufficient to evoke a clear thumb twitch (~ 1–2 cm). As shown in Fig. 6 A, the impedance at 10 Hz was significantly lower in the wet (7.4 ± 4.6 kΩ) and MN electrodes (6.0 ± 5.3 kΩ) compared to the dry electrode (424.8 ± 140.7 kΩ, p < 0.05). It should be noted that, in 2 of the 9 participants, the impedance of the dry electrode exceeded the EEG system’s 500 kΩ measurement limit, even after parting the hair. There was no significant difference between the wet and MN electrodes. The Fig. 6 B demonstrates typical example SEP response from a participant. All the electrode showed clear response around 20 ms after the stimulation, the N20 peak was larger in the MN electrode compared to the other two electrodes and the noise level (activity during the pre-stimulus period) seemed lower in the wet electrode compared to the other two electrodes. To evaluate quality of the EEG signals among the electrodes, we evaluated the SNR of SEP (Fig. 6 C). The SNR was higher in the wet and MN electrodes compared to that in the dry electrode (p < 0.05). It did not significantly differ between the wet and MN electrodes. 3.4 Usability Preparation time was measured from the moment the experimenter began fitting the participant with an EEG cap until the electrode–skin impedance fell below 20 kΩ for the wet and MN electrodes, and below 400 kΩ for the dry electrode. In all experiments, the same experimenter (T.Y.) was responsible for both the preparation and cleanup. Cleanup time included removing the electrodes after each measurement. For wet electrodes, it also included the time needed to wash out the conductive gel and thoroughly dry the hair. We found that the wet electrode had a significantly shorter preparation time than the other electrode types (p < 0.05, Fig. 7 A). However, it required the longest cleanup time (p < 0.05) compared with the other electrodes. Consequently, total time (preparation + cleanup) was significantly longer for the wet electrode than for the other two electrode types (p < 0.05, Fig. 7 C). Regarding comfort, participants reported mild comfort levels (approximately 2 to 4 on the visual analog scale (VAS), where 0 is “most comfortable,” 5 is “neutral,” and 10 is “most uncomfortable”) across all three electrode types, with no significant differences during either the preparation (Fig. 7 D) or measurement phases (Fig. 7 E). In contrast, during cleanup (Fig. 7 F), participants indicated greater comfort with the dry (1.65 ± 1.5 [mean ± SD]) and MN (1.83 ± 1.5) electrodes compared to the wet electrode (3.41 ± 1.2, p < 0.05). For pain (Fig. 7 G), participants reported significantly less pain with the wet electrode (0.60 ± 1.0) compared to the dry (2.93 ± 2.1, p < 0.05) and MN electrodes (1.86 ± 1.6, p < 0.05) on the VAS scale (0 = no pain, 10 = worst imaginable pain). There was no significant difference between the dry and MN electrodes. 4. Discussion We quantitatively assessed the signal quality and comfort level of newly developed PEDOT:PSS-coated MN electrodes, comparing them with conventional wet and dry electrodes. Our findings indicate that the MN electrode can deliver superior signal quality in hairless regions—surpassing even the wet electrode—and comparable performance in hairy areas. The remarkable aspect is that, despite the MN electrode being able to measure high-quality signals equal to or better than those of wet electrodes, it is a dry electrode and does not require gel. This balance of signal quality and ease of use makes the MN electrode a promising technology for both EEG and EMG applications. 4.1 MN electrodes on hairless region: EMG recording In the EMG experiment, the MN electrode showed significantly lower impedance below approximately 150 Hz compared with conventional electrodes (Fig. 5 B), as well as a significantly higher SNR than either type of conventional electrode (Fig. 5 C). This suggests that MN electrodes, despite being dry-type, can record high-quality signals when used on hairless skin. Two key factors would contribute to this high signal quality: (1) the penetration of the high-resistance stratum corneum by tiny needles and (2) the mixed conductivity of PEDOT:PSS. Regarding the first factor, MNs can indeed reduce impedance; however, MNs made solely of traditional metallic coating show higher impedance than wet electrodes [ 39 ]. In this study, following Li et al [ 27 ], we coated our MNs with PEDOT:PSS on Au, which led to a low impedance in the low-frequency band, even below that of wet electrodes (Fig. 5 B). This improvement occurs because the ionic/electronic mixed conductivity of PEDOT:PSS acts as a transducer, facilitating smooth current flow between the body’s ionic environment and the electrode’s electron-conductive metal [ 28 ]. While Li et al. [ 27 ] did measure EMG signals using PEDOT:PSS-coated MN electrodes, their study did not standardize the measurement site and contraction force, resulting in a less robust quantitative comparison. In contrast, by controlling for the measurement site and contraction force on EMG signals, our results indicate that MN electrodes yield superior signal quality compared with widely used wet and dry electrodes (Fig. 5 C). Therefore, under ideal conditions—namely, in hairless regions—the MN electrode, despite being a dry type, can surpass even conventional gel-assisted electrodes in terms of signal quality. In the comparison between conventional dry and wet electrodes, a significant difference in impedance was observed (Fig. 5 B). However, there was no significant difference in the SNR of the recorded EMG signals between them. We assume that this is because the difference in impedance magnitude between the dry and wet electrodes was smaller than that between the microneedle electrode and the other electrode types (Fig. 5 B). Consequently, a significant difference in SNR was observed between the microneedle electrode and the other electrodes, whereas no significant difference was likely to be observed between the dry and wet electrodes (Fig. 5 C). 4.2 MN electrodes on hairy region: EEG recording Although a previous study on PEDOT:PSS-coated MN electrodes evaluated EEG signals in hairless areas such as the forehead, their usefulness in hairy regions—the most common targets of EEG measurements—remains largely unexplored. In this study, we attempted to measure EEG over a hairy area and found that although the MN electrode achieved better impedance and SNR than the dry electrode, it was comparable to the wet electrode (Figs. 6 B and 6 C). This is likely because hair interferes with the electrode–skin contact, thereby reducing MN electrode performance. When using dry-type EEG electrodes, measuring in hairy areas has long been recognized as challenging [ 40 ]. Moreover, research on MN electrodes has shown that greater hair coverage caused higher impedance [ 41 ]. To address this, Kawana et al. [ 41 ] developed a small shutter mechanism to separate the hair and ensure reliable electrode–scalp contact. Employing such an approach may enable PEDOT:PSS-coated MN electrodes to realize their full potential for EEG measurements in hairy regions. Although the performance of MN electrode was similar to the wet electrode, it is noteworthy that the wet electrode used in this study was a high-grade, commercially available research system. Achieving comparable performance with a dry-type MN electrode is therefore highly significant in terms of balancing signal quality and convenience. Moreover, because MN electrodes do not require gel, they offer greater feasibility for long-term stable measurements [ 25 ]. Indeed, a study using Au-based MN electrodes for EEG showed less impedance drift over extended recordings compared to wet electrodes, leading to more stable signals and improved accuracy in decoding movement intention from brain activity [ 42 ]. Hence, even if PEDOT:PSS-coated MN electrodes do not exceed the signal quality of wet electrodes, their potential for prolonged, stable recording makes them promising for applications such as sleep research or brain–machine interface (BMI) studies, where long wearing times are common. 4.3 Usability The MN and dry electrodes had a shorter cleanup time than the wet electrode (Fig. 7 B), because they do not require a conductive gel, allowing participants to go home immediately after removing the electrodes. On the other hand, preparation time of the MN electrodes exceeded that of the wet electrode (Fig. 7 A), as positioning the MN electrode took longer due to parting the hair. Incorporating a hair-separating mechanism[ 41 ] (as mentioned previously) might automate this step and reduce setup time. In terms of pain, the MN electrodes induced only mild pain (1.86 ± 1.6 on the VAS scale, Fig. 7 G). Although this was higher than that was reported for the wet electrode (0.60 ± 1.0), it did not significantly differ from the dry electrode (2.93 ± 2.1). Because the pain level for the MN electrode remained relatively low, no significant differences were observed among the three electrode types regarding comfort/discomfort level during measurement (Fig. 7 E). Overall, despite having tiny needles, the MN electrode does not cause severe pain, and the user experience is similar to that of a conventional dry electrode. 4.4 Methodological consideration There is a thing to note regarding our impedance results. In the hairless region, we used a frame to standardize the contact area for all electrode types, but in the hairy region, no frame was used. Consequently, gel could have spread under the electrode in the wet electrode condition, potentially creating a larger contact area than that of the MN electrode and, in turn, overestimating the wet electrode’s impedance performance. Since the MN electrode penetrates the skin, we fixed the measurement order to wet, dry, and then MN, which may introduce order effects. In the EMG experiment, there was a concern on muscle fatigue, but previous research suggests that at 20% MVC, the biceps brachii typically becomes fatigued after about 10 min of sustained contraction [ 43 ]. Because our contractions lasted only 10 seconds, the impact of fatigue was likely minimal. In the EEG experiment, we repeated SEP sessions in a fixed order, as in the EMG experiment. To check order effects on SEP response size, we conducted a supplemental experiment of three repeated SEP sessions using wet electrodes for all sessions. The session interval was 8 minutes, which is similar to the total time spent on cleanup and preparation in the main experiment (Fig. 7 C). The supplemental experiment showed no clear order effects on SEP response (Figure S1). 4.5 Future direction The MN electrodes have a potential to provide high-quality yet convenient EEG measurements, they may be very useful for clinical applications. For instance, in a previous study [ 44 ], walking intentions extracted from the EEG of paraplegic patients triggered electrical stimulation of leg muscles, achieving partial restoration of walking function. However, the authors noted that using a conventional gel-based wet EEG system in everyday settings was challenging, mainly due to time and effort constraints, which could limit its widespread adoption in gait rehabilitation. Although a faster setup in hairy regions remains needed, MN electrodes could simplify EEG recording while maintaining high signal quality, thereby advancing routine clinical use. In EMG measurements, MN electrodes showed better signal quality than conventional electrodes (Fig. 4 C). Additionally, their low impedance suggests that reducing electrode size might still yield good-quality signals (Fig. 4 B). Consequently, MN electrodes could be highly valuable for high-density surface electromyography (HDsEMG), which has been actively studied recently [ 23 , 45 ]. Achieving high density EMG recordings requires miniaturizing electrodes, and the MN electrodes are well suited to that purpose. HDsEMG gains attention when used with blind source estimation, since it enables non-invasive investigation of spinal motor neuron firing, and the number of neurons that can be identified increases with higher SNR [ 29 , 46 ]. Therefore, applying MN electrodes to HDsEMG could increase the ability to evaluate motoneuron activity via surface EMG. 5. Conclusion We fabricated PEDOT:PSS coated MN electrodes and demonstrated experimentally that they can measure EMG signals with higher quality than conventional electrodes, and in EEG recordings they can match the signal quality of wet electrodes while offering dry-electrode convenience. These findings highlight the practicality of MN electrodes and indicate their strong potential for use in clinical EMG and EEG applications. It is expected that future studies apply the MN electrodes in rehabilitation BMI systems and high-density EMG recording, expanding their utility across a wide range of bio-signal measurement contexts. Declarations Ethics approval and consent to participate All participants provided written informed consent for participation in the study, which was conducted in accordance with the Declaration of Helsinki and approved by the Ethics Review Committee of The University of Tokyo (reference number, 23-396). Consent for publication Not applicable Availability of data and materials Data and code reported in this article will be shared by the lead contact upon request. All original code reported in this article will be shared by the lead contact upon request. Any additional information required to reanalyze the data reported in this article is available from the lead contact upon request. Competing interests The authors declare no competing interests. Funding This work was supported by JSPS KAKENHI Grant in Aid for Scientific Research (B) to H.Y.(21H03340), JSPS KAKENHI WAKATE to N.K. (50969285), Nakatani Foundation for Advancement of Measuring Technologies in Biomedical Engineering to H.Y. and N.K., JKA through its promotion funds from KEIRIN RACE to H.Y., and Tateisi Science and Technology Foundation to H.Y. and N.K. Authors' contributions T.Y., H.Y., and K.N. designed the research. T.Y., N.K., and Y.K. performed device fabrication. H.T. and H.Y. constructed the experimental setups. T.Y., E.N., and H.Y. collected data. T.Y. performed analyses and drafted the manuscript. All authors interpreted the data, discussed the findings, and approved the final version of the manuscript. Acknowledgements Not applicable References Teplan M. Fundamentals of EEG measurement. Measurement science review. 2002;2:1–11. Grosse P, Cassidy MJ, Brown P. EEG-EMG, MEG-EMG and EMG-EMG frequency analysis: Physiological principles and clinical applications. Clinical Neurophysiology. 2002;113:1523–31. Anbalagan T, Nath MK, Vijayalakshmi D, Anbalagan A. 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Microneedle Array Electrodes Fabricated With 3D Printing Technology for High-Quality Electrophysiological Acquisition. IEEE Transactions on Neural Systems and Rehabilitation Engineering. 2024; 32:2460-2469. Wang R, Jiang X, Wang W, Li Z. A microneedle electrode array on flexible substrate for long-term EEG monitoring. Sens Actuators B Chem. 2017;244:750–8. Jeong HR, Lee HS, Choi IJ, Park JH. Considerations in the use of microneedles: pain, convenience, anxiety and safety. J Drug Target. 2017;25:29–40. Li J, Ma Y, Huang D, Wang Z, Zhang Z, Ren Y, et al. High-performance flexible microneedle array as a low-impedance surface biopotential dry electrode for wearable electrophysiological recording and polysomnography. Nanomicro Lett. 2022;14:132. Bianchi M, De Salvo A, Asplund M, Carli S, Di Lauro M, Schulze‐Bonhage A, et al. Poly (3, 4‐ethylenedioxythiophene)‐based neural interfaces for recording and stimulation: fundamental aspects and in vivo applications. Advanced Science. 2022;9:2104701. Yokoyama H, Sasaki A, Kaneko N, Saito A, Nakazawa K. Robust Identification of Motor Unit Discharges from High-Density Surface EMG in Dynamic Muscle Contractions of the Tibialis Anterior. IEEE Access. 2021;9:123901–11. Lundsberg J, Björkman A, Malesevic N, Antfolk C. Inferring position of motor units from high-density surface EMG. Sci Rep. 2024;14:3858. Rossini PM, Cracco RQ, Cracco JB, House WJ. Short latency somatosensory evoked potentials to peroneal nerve stimulation: scalp topography and the effect of different frequency filters. Electroencephalogr Clin Neurophysiol. 1981;52:540–52. Doerrfuss JI, Kilic T, Ahmadi M, Weber JE, Holtkamp M. Predictive value of acute EEG measurements for seizures and epilepsy after stroke using a dry cap electrode EEG system—Study design and proof of concept. Epilepsy & Behavior. 2020;104:106486. Diaz-Piedra C, Gianfranchi E, Catena A, Di Stasi LL. Electrophysiological correlates of the reverse Stroop effect: Results from a simulated handgun task. International journal of psychophysiology. 2022;175:32–42. Nuwer MR, Aminoff San Francisco M, Desmedt Brussels J, Eisen AA, Matsuoka Kitakyushu S, Maugui F, et al. IFCN recommended standards for short latency somatosensory evoked potentials. Report of an IFCN committee Introduction Patient-related factors. Electroencephalogr Clin Neurophysiol. 1994;91:6-11 Hsieh J-C, Alawieh H, Li Y, Iwane F, Zhao L, Anderson R, et al. A highly stable electrode with low electrode-skin impedance for wearable brain-computer interface. Biosens Bioelectron. 2022;218:114756. Li G, Liu Y, Chen Y, Xia Y, Qi X, Wan X, et al. Robust, self‐adhesive, and low‐contact impedance polyvinyl alcohol/polyacrylamide dual‐network hydrogel semidry electrode for biopotential signal acquisition. SmartMat. 2024;5:e1173. Pataky TC, Robinson MA, Vanrenterghem J. Region-of-interest analyses of one-dimensional biomechanical trajectories: bridging 0D and 1D theory, augmenting statistical power. PeerJ. 2016;4:e2652. Boik RJ. The Fisher‐Pitman permutation test: A non‐robust alternative to the normal theory F test when variances are heterogeneous. British Journal of Mathematical and Statistical Psychology. 1987;40:26–42. O’Mahony C, Grygoryev K, Ciarlone A, Giannoni G, Kenthao A, Galvin P. Design, fabrication and skin-electrode contact analysis of polymer microneedle-based ECG electrodes. Journal of Micromechanics and Microengineering. 2016;26:084005. Lopez-Gordo MA, Sanchez-Morillo D, Valle FP. Dry EEG electrodes. Sensors. 2014;14:12847–70. Kawana T, Yoshida Y, Kudo Y, Iwatani C, Miki N. Design and characterization of an EEG-hat for reliable EEG measurements. Micromachines (Basel). 2020;11:635. Mahmood M, Kwon S, Kim H, Kim YS, Siriaraya P, Choi J, et al. Wireless Soft Scalp Electronics and Virtual Reality System for Motor Imagery-Based Brain–Machine Interfaces. Advanced Science. 2021;8. Hagberg M. Muscular endurance and surface electromyogram in isometric and dynamic exercise. J Appl Physiol. 1981;51:1–7. King CE, Wang PT, McCrimmon CM, Chou CC, Do AH, Nenadic Z. The feasibility of a brain-computer interface functional electrical stimulation system for the restoration of overground walking after paraplegia. J Neuroeng Rehabil. 2015;12:1–11. Del Vecchio AD, Sylos-Labini F, Mondì V, Paolillo P, Ivanenko Y, Lacquaniti F, et al. Spinal motoneurons of the human newborn are highly synchronized during leg movements. Sci Adv. 2020;6:2–4. Dai C, Hu X. Independent component analysis based algorithms for high-density electromyogram decomposition: Systematic evaluation through simulation. Comput Biol Med. 2019;109:171–81. Additional Declarations No competing interests reported. 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Also discoverable on Platform About Our Team In Review Editorial Policies Advisory Board Help Center Resources Author Services Accessibility API Access RSS feed Manage Cookie Preferences © Research Square 2026 | ISSN 2693-5015 (online) Privacy Policy Terms of Service Do Not Sell My Personal Information {"props":{"pageProps":{"initialData":{"identity":"rs-7343232","acceptedTermsAndConditions":true,"allowDirectSubmit":true,"archivedVersions":[],"articleType":"Research Article","associatedPublications":[],"authors":[{"id":501539288,"identity":"9bec9aa2-3fda-4ddf-8731-770a3f710fb6","order_by":0,"name":"Tomoya Yamaguchi","email":"","orcid":"","institution":"Tokyo University of Agriculture and Technology","correspondingAuthor":false,"prefix":"","firstName":"Tomoya","middleName":"","lastName":"Yamaguchi","suffix":""},{"id":501539289,"identity":"313c7202-84e4-4e37-bfd5-7be9e3d736ec","order_by":1,"name":"Yuta Kurashina","email":"","orcid":"","institution":"Tokyo University of Agriculture and Technology","correspondingAuthor":false,"prefix":"","firstName":"Yuta","middleName":"","lastName":"Kurashina","suffix":""},{"id":501539290,"identity":"c3430247-75f7-441e-b4e5-d9d65966c8a4","order_by":2,"name":"Eiji Nagahara","email":"","orcid":"","institution":"Tokyo University of Agriculture and Technology","correspondingAuthor":false,"prefix":"","firstName":"Eiji","middleName":"","lastName":"Nagahara","suffix":""},{"id":501539291,"identity":"a7488b2c-c68b-4629-bb94-a389d91fe1b0","order_by":3,"name":"Naotsugu Kaneko","email":"","orcid":"","institution":"The University of Tokyo","correspondingAuthor":false,"prefix":"","firstName":"Naotsugu","middleName":"","lastName":"Kaneko","suffix":""},{"id":501539292,"identity":"463b0854-f99f-42a4-a796-cca391574754","order_by":4,"name":"Kimitaka Nakazawa","email":"","orcid":"","institution":"The University of Tokyo","correspondingAuthor":false,"prefix":"","firstName":"Kimitaka","middleName":"","lastName":"Nakazawa","suffix":""},{"id":501539293,"identity":"df7b76b5-5517-4f73-b433-ffff8cb2ae7f","order_by":5,"name":"Hideyuki Tanaka","email":"","orcid":"","institution":"Tokyo University of Agriculture and Technology","correspondingAuthor":false,"prefix":"","firstName":"Hideyuki","middleName":"","lastName":"Tanaka","suffix":""},{"id":501539294,"identity":"6104f9e6-c5de-46a0-b110-88483ccd15e8","order_by":6,"name":"Hikaru Yokoyama","email":"data:image/png;base64,iVBORw0KGgoAAAANSUhEUgAAAZAAAAAyAQMAAABI0h/eAAAABlBMVEX///8AAABVwtN+AAAACXBIWXMAAA7EAAAOxAGVKw4bAAABC0lEQVRIiWNgGAWjYBACAyBmbAAzmQ8gS7ARo4UtAcJnI14LjwGyFtzAnIHH8OGMisPyuu1nPr/4wPAnsX9+A+OHHwx8ebi0WDbwGBtuOHPYcNuZ3G2WMxgMEmccY2CW7GFgK8bpsPtvzCQfth1m3HYgd5sxD1BLwzEGBmmg8xIbcGk5wAPWYr/t/JtnYC3zgbb8JqhlY9vhxG03cpgfg7RsOMbARsAWtmLDGWfSk7fdeGbGOMPA2HjjscQ2yx4DPH45wLzxYU+Fte2288mPP3yokJOdd/jw4Rs/Ko7hDDEGBg5QdDSDWGwS4GgCx5PBsQTcWtgfAIk6EIv5A5JwDR4to2AUjIJRMMIAAP9vW0EtzD81AAAAAElFTkSuQmCC","orcid":"","institution":"Tokyo University of Agriculture and Technology","correspondingAuthor":true,"prefix":"","firstName":"Hikaru","middleName":"","lastName":"Yokoyama","suffix":""}],"badges":[],"createdAt":"2025-08-11 07:08:13","currentVersionCode":1,"declarations":"","doi":"10.21203/rs.3.rs-7343232/v1","doiUrl":"https://doi.org/10.21203/rs.3.rs-7343232/v1","draftVersion":[],"editorialEvents":[],"editorialNote":"","failedWorkflow":false,"files":[{"id":89391263,"identity":"c2260e94-7ba4-4e66-b34b-77d4cb13262f","added_by":"auto","created_at":"2025-08-19 13:05:46","extension":"png","order_by":1,"title":"Figure 1","display":"","copyAsset":false,"role":"figure","size":92296,"visible":true,"origin":"","legend":"\u003cp\u003eSchematic illustration of microneedle electrodes applied to human skin. The tiny needles penetrate the stratum corneum, which has high impedance, resulting in low skin-electrode impedance.\u003c/p\u003e","description":"","filename":"floatimage1.png","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/89242e777b540edc7b692bfd.png"},{"id":89391266,"identity":"55d82c5d-0bca-4cac-8967-08c065935f17","added_by":"auto","created_at":"2025-08-19 13:05:46","extension":"png","order_by":2,"title":"Figure 2","display":"","copyAsset":false,"role":"figure","size":260049,"visible":true,"origin":"","legend":"\u003cp\u003eFabrication process of PEDOT:PSS-coated microneedle electrodes for EEG and EMG recordings. (A) A polyimide layer formed by thermal curing of polyamic acid. (B) An Au layer deposited via vacuum evaporation. (C) A PEDOT:PSS layer formed by electrodeposition from a monomer solution. (D) Needle height was adjusted according to application—1000μm was adopted for EEG and 600μm for EMG to account for hair thickness. (E) Example of a fabricated microneedle electrode; the one shown is designed for EMG recordings.\u003c/p\u003e","description":"","filename":"floatimage2.png","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/7bf64fbce7de6afb49df0532.png"},{"id":89391265,"identity":"0a34cdb0-1209-4a02-a81f-fc363451b282","added_by":"auto","created_at":"2025-08-19 13:05:46","extension":"png","order_by":3,"title":"Figure 3","display":"","copyAsset":false,"role":"figure","size":120592,"visible":true,"origin":"","legend":"\u003cp\u003eSchematic illustration of EMG and EEG experiment. (A) In EMG experiment, participants performed a force match task of the elbow flexion. The exerted force was measured by a force sensor (A1) and visually feedbacked to the participant. Two electrodes were placed on muscle belly of the biceps brachii (A2), and a ground electrode was positioned on the acromion (A3). (B) In EEG experiment, somatosensory evoked potentials were evaluated. Electrical stimulation was delivered to the right median nerve (B1) and the sensory information was propagated and reached to the somatosensory cortex. The EEG electrodes measured the sensory evoked potentials (B2).\u003c/p\u003e","description":"","filename":"floatimage3.png","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/b1cae0dd4aac2e9a590fabc6.png"},{"id":89392305,"identity":"01c29af1-e21a-4506-8032-c49772a2e013","added_by":"auto","created_at":"2025-08-19 13:13:46","extension":"png","order_by":4,"title":"Figure 4","display":"","copyAsset":false,"role":"figure","size":106349,"visible":true,"origin":"","legend":"\u003cp\u003eSchematic illustration of EEG recordings by different electrode types.\u003c/p\u003e\n\u003cp\u003e(A) Wet Electrode. (B) Dry Electrode. (C) Microneedle electrode.\u003c/p\u003e","description":"","filename":"floatimage4.png","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/9e4f4ae2c2196a4b16dd6271.png"},{"id":89391269,"identity":"7bdbca2e-7637-489d-8557-63d24b0fd4f0","added_by":"auto","created_at":"2025-08-19 13:05:46","extension":"png","order_by":5,"title":"Figure 5","display":"","copyAsset":false,"role":"figure","size":630705,"visible":true,"origin":"","legend":"\u003cp\u003eResults of three types of electrodes in the EMG experiment. (A) Representative examples of recorded EMG signals during elbow flexion (top row) and at rest (bottom row). (B) Electrode-skin interface impedance spectra for different electrodes across a frequency range from 5 Hz to 1000 Hz (mean ± standard error). The black horizontal bars indicate frequency bands where impedance differences were significant. Transparent areas represent the standard error. (C) Signal-to-noise ratio (SNR) of the EMG signals. Error bars and asterisks denote the standard error and significant differences, respectively. *: p \u0026lt; 0.05.\u003c/p\u003e","description":"","filename":"floatimage5.png","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/bef1f6a5ae10edd2498a0411.png"},{"id":89391270,"identity":"df1d9616-4158-4f57-a05c-f1b531cf6aa4","added_by":"auto","created_at":"2025-08-19 13:05:46","extension":"png","order_by":6,"title":"Figure 6","display":"","copyAsset":false,"role":"figure","size":280619,"visible":true,"origin":"","legend":"\u003cp\u003eResults from three types of electrodes in the EEG experiment. (A): Electrode-skin impedance at 10 Hz. (B): Typical examples of somatosensory evoked potentials (SEP); N20 responses are highlighted by a dashed-line circle. (C): Signal-to-noise ratio (SNR) of the EEG signals based on the SEP responses. Error bars and asterisks denote the standard error and significant differences, respectively. *: p \u0026lt; 0.05, **: p \u0026lt; 0.01. Each black plot represents data from a participant.\u003c/p\u003e","description":"","filename":"floatimage6.png","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/c9b37835b09aadd9a7a43133.png"},{"id":89391275,"identity":"102859ea-391b-4a51-8589-2ff0628d2fa6","added_by":"auto","created_at":"2025-08-19 13:05:46","extension":"png","order_by":7,"title":"Figure 7","display":"","copyAsset":false,"role":"figure","size":265079,"visible":true,"origin":"","legend":"\u003cp\u003eUsability scores for three types of electrodes in EEG measurements. (A) Preparation time until the electrode–skin impedance dropped below the target impedance (20 kΩ for wet and microneedle electrodes, 400 kΩ for dry electrodes). (B) Cleanup time, which included washing the hair only in the wet electrode. (C) Total time (preparation + cleanup). (D–F) Visual analog scales (VAS) assessing comfort levels during preparation (D), measurement (E), and cleanup (F). On the VAS, 0 represents “most comfortable”, 5 “neutral”, and 10 “most uncomfortable”. (G) VAS for pain. 0 represents “no pain”, 10 “the worst pain imaginable”. Error bars and asterisks indicate the standard error and significant differences, respectively. Each black plot represents data from a participant. *: p \u0026lt; 0.05, **: p \u0026lt; 0.01.\u003c/p\u003e","description":"","filename":"floatimage7.png","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/fe79f11dd6a31cb83a8cadac.png"},{"id":91115935,"identity":"b52aa7aa-6c40-4404-bd87-5b410d1d8016","added_by":"auto","created_at":"2025-09-11 17:31:42","extension":"pdf","order_by":0,"title":"","display":"","copyAsset":false,"role":"manuscript-pdf","size":2560951,"visible":true,"origin":"","legend":"","description":"","filename":"manuscript.pdf","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/f2cb8012-5dba-433e-8a25-117675bae9fa.pdf"},{"id":89391264,"identity":"c9a7f4b7-2fdb-4666-b01a-1fc9ed89b60b","added_by":"auto","created_at":"2025-08-19 13:05:46","extension":"docx","order_by":1,"title":"","display":"","copyAsset":false,"role":"supplement","size":36724,"visible":true,"origin":"","legend":"","description":"","filename":"Supplementaryfigure.docx","url":"https://assets-eu.researchsquare.com/files/rs-7343232/v1/5c413d318026959b19299545.docx"}],"financialInterests":"No competing interests reported.","formattedTitle":"Quantitative Assessment of Signal Quality and Usability of EEG and EMG recordings with PEDOT:PSS-Coated Microneedle Electrodes","fulltext":[{"header":"1. Background","content":"\u003cp\u003eBioelectrical signals are crucial indicators of bodily function, psychological status, and clinical pathologies. Thanks to significant advancements in bioelectric acquisition technology, electrocardiography (ECG), electromyography (EMG), and electroencephalography (EEG) have become essential tools in both clinical and neuroscience fields [\u003cspan additionalcitationids=\"CR2\" citationid=\"CR1\" class=\"CitationRef\"\u003e1\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR3\" class=\"CitationRef\"\u003e3\u003c/span\u003e]. These bioelectrical signals can be recorded using two approaches: invasive and non-invasive recordings. Invasive recordings, which involve inserting electrodes directly into the body near the target site, yield high-quality signals but are generally restricted to severe clinical interventions due to the invasiveness of the procedures involved [\u003cspan citationid=\"CR4\" class=\"CitationRef\"\u003e4\u003c/span\u003e]. In contrast, surface bioelectrical signals are now more frequently employed as a non-invasive alternative [\u003cspan additionalcitationids=\"CR6\" citationid=\"CR5\" class=\"CitationRef\"\u003e5\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR7\" class=\"CitationRef\"\u003e7\u003c/span\u003e]. For instance, surface EEG signals, which are most widely used modality in various fields in the above-mentioned three types of bioelectrical signals, are employed to diagnose epilepsy [\u003cspan citationid=\"CR8\" class=\"CitationRef\"\u003e8\u003c/span\u003e], to brain\u0026ndash;machine interface (BMI) technologies that compensate for impaired bodily functions [\u003cspan citationid=\"CR9\" class=\"CitationRef\"\u003e9\u003c/span\u003e], and to further neuroscientific research on human cognitive and motor functions [\u003cspan citationid=\"CR10\" class=\"CitationRef\"\u003e10\u003c/span\u003e, \u003cspan citationid=\"CR11\" class=\"CitationRef\"\u003e11\u003c/span\u003e].\u003c/p\u003e\u003cp\u003eBecause surface electrodes are placed on the skin to capture bioelectric signals, they essentially function as transducers, converting ionic currents from the human body into electronic currents that are then transmitted to external electronic systems made of various metals [\u003cspan citationid=\"CR12\" class=\"CitationRef\"\u003e12\u003c/span\u003e, \u003cspan citationid=\"CR13\" class=\"CitationRef\"\u003e13\u003c/span\u003e]. For the signal transmission at the electrode\u0026ndash;skin interface, the electrode\u0026ndash;skin interface impedance (EII) is a crucial parameter. Typically, decreased EII leads to better signal quality, an increased signal-to-noise ratio, and diminished baseline drift [\u003cspan citationid=\"CR14\" class=\"CitationRef\"\u003e14\u003c/span\u003e]. To lower this impedance during EEG recordings with surface electrodes, a conductive gel containing electrolytes has traditionally been applied between the skin and the electrode [\u003cspan citationid=\"CR13\" class=\"CitationRef\"\u003e13\u003c/span\u003e]. However, the wet electrode approach is both time-consuming and labor-intensive\u0026mdash;requiring gel application and hair washing before or after measurements\u0026mdash;and it also limits recording time due to gel drying [\u003cspan citationid=\"CR13\" class=\"CitationRef\"\u003e13\u003c/span\u003e]. Consequently, balancing ease of use and high-quality signals remains challenging with current wet electrodes, which in turn hinders their widespread adoption for routine clinical use and in real-world applications for health monitoring.\u003c/p\u003e\u003cp\u003eIn recent years, numerous research efforts have focused on developing dry electrodes to address this problem [\u003cspan citationid=\"CR15\" class=\"CitationRef\"\u003e15\u003c/span\u003e, \u003cspan citationid=\"CR16\" class=\"CitationRef\"\u003e16\u003c/span\u003e]. Dry electrodes without conductive gel are better suited for long-term brain activity monitoring thanks to no gel drying problem and for daily use and real-world applications because of the no need for gel application and hair washing. Because of the high-resistance stratum corneum (SC) and insufficient contact between the skin and the metal electrode surface, dry electrodes generally have higher skin\u0026ndash;electrode impedance compared to wet electrodes [\u003cspan citationid=\"CR15\" class=\"CitationRef\"\u003e15\u003c/span\u003e, \u003cspan citationid=\"CR17\" class=\"CitationRef\"\u003e17\u003c/span\u003e]. Various new materials and structures have been proposed to improve the electrode\u0026ndash;skin contact [\u003cspan additionalcitationids=\"CR19\" citationid=\"CR18\" class=\"CitationRef\"\u003e18\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR20\" class=\"CitationRef\"\u003e20\u003c/span\u003e], but the trade-off is the large electrode volume or footprint because the electrode size is critical for high density EEG/EMG recordings, which are required for source estimation of the EEG and EMG signals [\u003cspan additionalcitationids=\"CR22\" citationid=\"CR21\" class=\"CitationRef\"\u003e21\u003c/span\u003e\u0026ndash;\u003cspan citationid=\"CR23\" class=\"CitationRef\"\u003e23\u003c/span\u003e].\u003c/p\u003e\u003cp\u003eRecently developed microneedle (MN) electrodes can penetrate the high-resistance stratum corneum by using tiny needles that penetrate the skin with minimal discomfort [\u003cspan citationid=\"CR24\" class=\"CitationRef\"\u003e24\u003c/span\u003e, \u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e], thus enabling high-quality measurements without the need for extensive skin preparation and conductive gel (Fig.\u0026nbsp;\u003cspan refid=\"Fig1\" class=\"InternalRef\"\u003e1\u003c/span\u003e). Because the needles on these electrodes are very small, causing no/slight pain [\u003cspan citationid=\"CR26\" class=\"CitationRef\"\u003e26\u003c/span\u003e], they are known as micro-invasive electrodes. Recently, it was reported that depositing a conductive polymer, poly (3,4-ethylene dioxythiophene):poly(styrenesulfonate) (PEDOT:PSS), onto metal MNs dramatically reduces electrode impedance [\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]. Thus, these electrodes can be smaller than\u0026mdash;thanks to the low impedance per size\u0026mdash;conventional wet electrodes. PEDOT:PSS exhibits mixed conductivity, which can conducts both electronic and ionic charges, it functions as a transducer that facilitates smooth current flow between the body\u0026rsquo;s ionic environment and the electrode\u0026rsquo;s electron-conductive metal [\u003cspan citationid=\"CR28\" class=\"CitationRef\"\u003e28\u003c/span\u003e].\u003c/p\u003e\u003cp\u003e\u003c/p\u003e\u003cp\u003eA study developing microneedle electrodes coated with PEDOT:PSS found that the electrode\u0026ndash;skin impedance in hairless areas (such as the forehead or arms) was much lower compared to wet electrodes [\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]. However, these electrodes have not yet been applied to the scalp, which is covered with hair. In addition, there has been no comprehensive quantitative comparison in EEG and EMG signals with conventional dry and wet electrodes. Although the study by Li et al, (2022)[\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e] evaluated EMG signal amplitude and signal-to-noise ratio (SNR) against conventional electrodes, several methodological issues arose. For instance, electrode placement differed between electrode types\u0026mdash;even though EMG signals vary by location [\u003cspan citationid=\"CR29\" class=\"CitationRef\"\u003e29\u003c/span\u003e]\u0026mdash;and no standardized muscle contraction level was used, potentially causing inconsistent muscle activity across tasks. In this study, we aimed to quantitatively compare the performance of MN electrodes coated with PEDOT:PSS against conventional wet and dry electrodes. First, to assess electrode performance under hairless conditions, we measured EMG signals and impedance on the arms. Next, we evaluated their performance in EEG signals measured from hairy regions by somatosensory evoked potentials (SEPs). During these measurements, we recorded setup and cleanup times, as well as any pain or discomfort reported by participants, to assess usability across electrode types.\u003c/p\u003e"},{"header":"2. Methods","content":"\u003cp\u003e\u003cb\u003eOverview of MN electrodes\u003c/b\u003e\u003c/p\u003e\u003cp\u003eFigure \u003cspan refid=\"Fig2\" class=\"InternalRef\"\u003e2\u003c/span\u003e. Fabrication process of PEDOT:PSS-coated microneedle electrodes for EEG and EMG recordings. (A) A polyimide layer formed by thermal curing of polyamic acid. (B) An Au layer deposited via vacuum evaporation. (C) A PEDOT:PSS layer formed by electrodeposition from a monomer solution. (D) Needle height was adjusted according to application\u0026mdash;1000\u0026micro;m was adopted for EEG and 600\u0026micro;m for EMG to account for hair thickness. (E) Example of a fabricated microneedle electrode; the one shown is designed for EMG recordings.\u003c/p\u003e\u003cdiv id=\"Sec3\" class=\"Section2\"\u003e\u003ch2\u003e2.1 Materials\u003c/h2\u003e\u003cp\u003eA polyamic acid solution (Pyre-M.L) was purchased from IST Corp. (Japan). The PDMS-based silicone mold for MN electrode (Mpatch Microneedle Template) was obtained from Micropoint Technologies Pte Ltd (Singapore). Au and chromium for vapor deposition were purchased from Niraco inc. (Japan), and PSS (molecular weight of ~\u0026thinsp;70,000) along with EDOT (97%, the monomer for PEDOT) were purchased from Sigma-Aldrich LLC (USA).\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec4\" class=\"Section2\"\u003e\u003ch2\u003e2.2 Fabrication process\u003c/h2\u003e\u003cp\u003eTwo types of PDMS molds were used to fabricate MN arrays. For EMG experiments, the mold had a needle length of 600 \u0026micro;m, a base diameter of 200 \u0026micro;m, a pitch of 500 \u0026micro;m, and a 10\u0026times;10 needle array. For EEG experiments, the mold had a needle length of 1,000 \u0026micro;m, a base diameter of 250 \u0026micro;m, a pitch of 500 \u0026micro;m, and a 15\u0026times;15 needle array, considering hair layer thickness.\u003c/p\u003e\u003cdiv id=\"Sec5\" class=\"Section3\"\u003e\u003ch2\u003e2.2.1 Polyimide layer (Fig.\u0026nbsp;\u003cspan refid=\"Fig2\" class=\"InternalRef\"\u003e2\u003c/span\u003eA)\u003c/h2\u003e\u003cp\u003eEach PDMS mold was ultrasonically cleaned (120 kHz with a Ultrasonic Cleaner HFC-3D, AS ONE, Japan) for 30 minutes and then dried. Next, polyamic acid solution was poured into the PDMS mold and centrifuged at 3,000 rpm for 15 minutes to ensure complete filling with a centrifugal machine (H-19a, KOKUSAN, Japan). Additional polyamic acid solution was then applied, and a scraper was used to level it to the mold\u0026rsquo;s edges. The mold was placed on a hot plate at 100\u0026deg;C for 30 minutes to evaporate the solvent. The temperature was gradually raised to 200\u0026deg;C over a 30-minute period, and then held at 200\u0026deg;C for another 30 minutes to induce imidization and cure the film. Because the PDMS mold can only tolerate up to 200\u0026deg;C, the polyimide substrate was removed from the mold before an additional 5-minute heating step at 300\u0026deg;C, completing the imidization process and yielding the final polyimide layer.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec6\" class=\"Section3\"\u003e\u003ch2\u003e2.2.2 Au layer (Fig.\u0026nbsp;\u003cspan refid=\"Fig2\" class=\"InternalRef\"\u003e2\u003c/span\u003eB)\u003c/h2\u003e\u003cp\u003eUsing a vacuum deposition device (VE-2013, Vacuum Device Inc., Japan), we coated both sides of the polyimide layer with a 10 nm titanium (Ti) adhesion layer and a 200 nm Au conductive layer, thereby producing conductive MN electrodes. Deposition was performed separately on each side to ensure conductivity on both the front and back surfaces.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec7\" class=\"Section3\"\u003e\u003ch2\u003e2.2.3 PEDOT:PSS layer (Fig.\u0026nbsp;\u003cspan refid=\"Fig2\" class=\"InternalRef\"\u003e2\u003c/span\u003eC)\u003c/h2\u003e\u003cp\u003eA 2.5 wt% PSS solution and 0.01 M EDOT were dissolved in pure water, and a PEDOT:PSS layer was deposited onto the Au surface via electrochemical polymerization. In this setup, a platinum rod served as the cathode, while the MN electrode (with its gold layer) was used as the anode. Because excessive current can accelerate polymerization too quickly\u0026mdash;leading to cracks or poor adhesion in the PEDOT:PSS layer [\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e]\u0026mdash;the current was increased stepwise at 20, 40, 60, and 80 \u0026micro;A for 150 seconds each, and then 100 \u0026micro;A was applied for 600 seconds. Matlab (MathWorks, USA) was used to control the current, and an NI-9265 (National Instruments, USA) functioned as the constant current source.\u003c/p\u003e\u003c/div\u003e\u003c/div\u003e\u003cdiv id=\"Sec8\" class=\"Section2\"\u003e\u003ch2\u003e2.3 EMG experiment\u003c/h2\u003e\u003cdiv id=\"Sec9\" class=\"Section3\"\u003e\u003ch2\u003e2.3.1 Participant\u003c/h2\u003e\u003cp\u003eEight healthy male adults were participated in this experiment. All participants provided written informed consent for participation in the study, which was conducted in accordance with the Declaration of Helsinki and approved by the Ethics Review Committee of The University of Tokyo (reference number, 23\u0026ndash;396).\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec10\" class=\"Section3\"\u003e\u003ch2\u003e2.3.2 EMG acquisition\u003c/h2\u003e\u003cp\u003eBecause MN electrodes create small punctures in the skin, measurements were performed in the following order: wet electrode, dry electrode, and then MN electrode. Before each measurement, the skin was cleaned with an alcohol swab and dried to maintain consistent skin conditions.\u003c/p\u003e\u003cp\u003eEMG signals were recorded using an electromyography amplifier (BrainAmp-ExG MR, Brain Products, Germany) at a sampling rate of 2,500 Hz. The data were then processed with a 4th-order Butterworth bandpass filter (20\u0026ndash;350 Hz), and 50 Hz powerline noise was removed with a notch filter. Two electrodes were placed on muscle belly of the biceps brachii of the dominant arm with a 2 cm center-to-center spacing, and a ground electrode was positioned on the acromion. To minimize external noise, the wires were twisted. The positions of the electrodes were marked to ensure consistent placement for each electrode condition. After attaching each electrode to the arm, it was fixed with an elastic underwrap.\u003c/p\u003e\u003cp\u003eAll electrodes were standardized to a 5 mm \u0026times; 5 mm square area. The dry electrode consisted of a round, gold-plated metal body onto which a thin insulating film with a 5 mm \u0026times; 5 mm square opening was affixed; a rod on the back allowed for clip-connector wiring. To form the wet electrode, we attached a double-sided adhesive urethane foam\u0026mdash;also featuring a 5 mm \u0026times; 5 mm square opening\u0026mdash;to the same round, gold-plated metal base used for the dry electrode, then filled the opening with conductive paste (AC cream, OT Bio Elettronica, Italy). The MN electrode was fabricated as a 5 mm \u0026times; 5 mm MN array consisting of 10\u0026times;10 needles. A silver-paste-based conductive epoxy (CW2400, Chemtronics, USA) was applied to the back of the needle surface to enable wiring.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec11\" class=\"Section3\"\u003e\u003ch2\u003e2.3.3 Experimental Task\u003c/h2\u003e\u003cp\u003eParticipants performed a force match task of flexion of the elbow joint (Fig.\u0026nbsp;\u003cspan refid=\"Fig3\" class=\"InternalRef\"\u003e3\u003c/span\u003eA). Before the main task, maximum voluntary contraction (MVC) force was measured for 3 s. The MVC task was performed before attaching electrode. Thus, the same contraction level was used across all the electrode conditions. Then we recorded 10 seconds of EMG signals under relaxed (rest) conditions. Next, the participant performed elbow flexion at 20% MVC for 10 s. We selected this contraction level because 20% MVC has been frequently used in high-density EMG studies to examine motor neuron firing[\u003cspan citationid=\"CR23\" class=\"CitationRef\"\u003e23\u003c/span\u003e, \u003cspan citationid=\"CR30\" class=\"CitationRef\"\u003e30\u003c/span\u003e], which is a potential application of the MN electrodes. The EMG recordings of rest and the force match task were iterated for each electrode type.\u003c/p\u003e\u003cp\u003e\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec12\" class=\"Section3\"\u003e\u003ch2\u003e2.3.4 Evaluation of signals quality of EMG and skin-electrode impedance\u003c/h2\u003e\u003cp\u003eThe root mean square (RMS) of EMG amplitude was computed for each condition, and the ratio of these RMS values served as the SNR, in accordance with Eq.\u0026nbsp;(1):\u003cdiv id=\"Equa\" class=\"Equation\"\u003e\u003cdiv format=\"TEX\" class=\"mathdisplay\" id=\"FileID_Equa\" name=\"EquationSource\"\u003e\n$$\\:\\begin{array}{c}SNR=20{\\text{log}}_{10}\\frac{\\text{R}\\text{M}{\\text{S}}_{\\text{s}\\text{i}\\text{g}\\text{n}\\text{a}\\text{l}}}{\\text{R}\\text{M}{\\text{S}}_{\\text{n}\\text{o}\\text{i}\\text{s}\\text{e}}}\\#\\left(1\\right)\\end{array}$$\u003c/div\u003e\u003c/div\u003e,\u003c/p\u003e\u003cp\u003ewhere RMS\u003csub\u003esignal\u003c/sub\u003e ​is the RMS value during the force match task and RMS\u003csub\u003enoise​\u003c/sub\u003e is the RMS value during the rest task.\u003c/p\u003e\u003cp\u003eAdditionally, before measuring EMG signals, the EII spectra of the electrodes were measured by a two-electrode system with an Analog Discovery 3 (Digilent, USA) and its control software WaveForms (Digilent, USA).\u003c/p\u003e\u003c/div\u003e\u003c/div\u003e\u003cdiv id=\"Sec13\" class=\"Section2\"\u003e\u003ch2\u003e2.4 EEG experiment\u003c/h2\u003e\u003cdiv id=\"Sec14\" class=\"Section3\"\u003e\u003ch2\u003e2.4.1 Participant\u003c/h2\u003e\u003cp\u003eNine healthy male adults were participated in this experiment. All participants provided written informed consent for participation in the study, which was conducted in accordance with the Declaration of Helsinki and approved by the Ethics Review Committee of The University of Tokyo (reference number, 23\u0026ndash;396).\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec15\" class=\"Section3\"\u003e\u003ch2\u003e2.4.2 EEG acquisition\u003c/h2\u003e\u003cp\u003eEEG signals were recorded at a sampling rate of 5,000 Hz using an EEG amplifier (actiCHamp Plus, Brain Products, Germany). The data were then processed with a 4th-order Butterworth bandpass filter (30\u0026ndash;150 Hz) [\u003cspan citationid=\"CR31\" class=\"CitationRef\"\u003e31\u003c/span\u003e], and 50 Hz power-line noise was removed using a notch filter. Following the International 10\u0026ndash;20 System, Cz was used as the ground electrode, Fz as the reference, and Cp3 for measuring somatosensory evoked potentials.\u003c/p\u003e\u003cp\u003eIn this study, three different types of EEG electrodes (Dry, Wet, and MN) were compared. Measurements under these three conditions were conducted by applying gel to a Dry EEG system (actiCap Xpress Twist system) used in neuroscience research [\u003cspan citationid=\"CR32\" class=\"CitationRef\"\u003e32\u003c/span\u003e, \u003cspan citationid=\"CR33\" class=\"CitationRef\"\u003e33\u003c/span\u003e] to create the Wet condition, and by attaching microneedle electrodes to the tip of the Dry electrode to create the Microneedle condition. The details are as follows:\u003c/p\u003e\u003c/div\u003e\u003c/div\u003e\n\u003ch3\u003e1)Wet Electrode (Fig. A)\u003c/h3\u003e\n\u003cp\u003eA slightly convex, dish-shaped, gold-plated metal electrode (actiCap Xpress Twist with Flat QuickBit, Brain Products, Germany). Conductive gel (SuperVisc, Easycap GmbH, Germany) was applied between the metal surface and the skin.\u003c/p\u003e\n\u003ch3\u003e2)Dry Electrode (Fig. B)\u003c/h3\u003e\n\u003cp\u003eThe same actiCap Xpress Twist system, but equipped with Long QuickBit tips (length: 12 mm) designed for hairy areas. Used without conductive gel.\u003c/p\u003e\n\u003ch3\u003e3)MN Electrode (Fig. C)\u003c/h3\u003e\n\u003cp\u003eThe same actiCap Xpress Twist system with Flat QuickBit, as in the wet electrode condition. A MN array was affixed to the electrode using conductive double-sided tape (T-9426, ESD EMI Engineering, Japan), and no conductive gel was applied.\u003c/p\u003e\u003cp\u003eIn the above comparison of electrodes during EMG measurements, the area of skin contacted by the electrodes could be controlled. However, it should be noted that in EEG measurements, such control is difficult due to the influence of hair.\u003c/p\u003e\u003cp\u003eSince MN electrodes can create small punctures in the skin, measurements were performed in this order: wet electrode, dry electrode, then MN electrode. Before applying the wet electrode, the skin at the electrode site was cleaned with an alcohol swab. After wet electrode measurements, the conductive gel was washed off, and the scalp and hair were thoroughly dried with a hairdryer before proceeding to the dry electrode measurement. All electrodes were then secured to an elastic cap specifically designed for actiCap Xpress Twist.\u003c/p\u003e\u003cdiv id=\"Sec19\" class=\"Section3\"\u003e\u003cdiv class=\"Heading\"\u003e2.4.3 Electrical Stimulation for somatosensory evoked potential\u003c/div\u003e\u003cp\u003eWe administered 200 ms monophasic square-wave constant-current pulses using a clinical stimulator (USE-100, UNIQUE MEDICAL, Japan) to stimulate the right median nerve. We placed a pair of electrodes in positions on the right wrist where the visible muscle twitch was most clearly observed. The current intensity was set to evoke a 1\u0026ndash;2 cm thumb movement, following the recommendations of the International Federation of Clinical Neurophysiology [\u003cspan citationid=\"CR34\" class=\"CitationRef\"\u003e34\u003c/span\u003e]. Stimulation was applied 300 times at a repetition rate of 2.1 Hz. To maintain the participants\u0026rsquo; attention on the stimulus, pauses were introduced every 30\u0026ndash;50 seconds. Specifically, between 3 and 4 of these pauses\u0026mdash;each lasting about 4 seconds\u0026mdash;were inserted at random during each session, and participants were instructed to count the pauses and report their totals at the end of the session. All participants accurately reported the number of pauses in every session. To avoid variations in posture or hand position that could affect the experimental results, each participant\u0026rsquo;s hand and arm placements were marked on the support platform, ensuring a consistent posture across all conditions.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec20\" class=\"Section3\"\u003e\u003cdiv class=\"Heading\"\u003e2.4.4 Evaluation of Electrode quality for EEG\u003c/div\u003e\u003cp\u003eTo assess the quality of EEG signals recorded by MN electrodes compared to conventional (dry and wet) electrodes, we calculated the SNR of somatosensory evoked potentials (SEPs). The SEPs were obtained using a stimulation triggered averaging method, in which signal segments around the timing of electrical stimulation were extracted and averaged. The amplitude of the evoked response was defined as the N20 peak (observed between 19.0 and 21.0 ms following stimulation), and noise was defined as the standard deviation of the pre-stimulation signal from \u0026minus;\u0026thinsp;70 to -5 ms. The ratio of these two values was taken as the SNR, in accordance with Eq.\u0026nbsp;(2):\u003cdiv id=\"Equb\" class=\"Equation\"\u003e\u003cdiv format=\"TEX\" class=\"mathdisplay\" id=\"FileID_Equb\" name=\"EquationSource\"\u003e\n$$\\:\\begin{array}{c}SNR=20{\\text{log}}_{10}\\frac{{\\text{A}\\text{m}\\text{p}\\text{l}\\text{i}\\text{t}\\text{u}\\text{d}\\text{e}}_{\\text{s}\\text{i}\\text{g}\\text{n}\\text{a}\\text{l}}}{{\\text{S}\\text{D}}_{\\text{b}\\text{a}\\text{c}\\text{k}\\text{g}\\text{r}\\text{o}\\text{u}\\text{n}\\text{d}}}\\#\\left(2\\right)\\end{array}$$\u003c/div\u003e\u003c/div\u003e,\u003c/p\u003e\u003cp\u003ewhere Amplitude\u003csub\u003esignal\u003c/sub\u003e is a peak amplitude of N20 and SD\u003csub\u003ebackground\u003c/sub\u003e is the standard deviation of the pre-stimulation period (from \u0026minus;\u0026thinsp;70 to -5 ms).\u003c/p\u003e\u003cp\u003eAdditionally, before measuring EEG signals, the EII at 10Hz of the electrodes were measured with a impedance measurement function of the EEG amplifier (actiCHamp Plus) and its control software (Brain Products, Germany). The analyzed frequency was predetermined by the EEG system. Since 10 Hz corresponds to the alpha wave in EEGs, impedance at 10 Hz has been used to evaluate the biopotential electrodes[\u003cspan citationid=\"CR35\" class=\"CitationRef\"\u003e35\u003c/span\u003e, \u003cspan citationid=\"CR36\" class=\"CitationRef\"\u003e36\u003c/span\u003e]. Due to equipment limitations, any value exceeding 500 kΩ was recorded as 500 kΩ. For many participants in the dry electrode, the impedance did not fall below 500 kΩ.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec21\" class=\"Section3\"\u003e\u003cdiv class=\"Heading\"\u003e2.4.5 Assessment of MN Electrode Usability\u003c/div\u003e\u003cp\u003eWe evaluated (1) preparation and cleanup time, (2) participant comfort, and (3) pain intensity during EEG measurements.\u003c/p\u003e\u003cp\u003ePreparation time was measured from the moment the experimenter began fitting the participant with an EEG cap until the electrode\u0026ndash;skin impedance fell below 20 kΩ for the wet and MN electrodes, and below 400 kΩ for the dry electrode. In all experiments, the same experimenter (T.Y.) was responsible for both the preparation and cleanup. If impedance for the dry electrode remained above 400 kΩ and did not improve even after parting the hair under the electrode, we stopped recording the preparation time. Cleanup time included removing the electrodes after each measurement. For wet electrodes, it also included the time needed to wash out the conductive gel and thoroughly dry the hair.\u003c/p\u003e\u003cp\u003eComfort and pain were rated for each electrode condition during setup, recording, and removal using a visual analogue scale (VAS). For comfort, the highest value (10) indicated very uncomfortable and the lowest value (0) very comfortable. For pain, the highest value (10) represented the worst pain imaginable and the lowest (0) no pain at all. Participants were asked to consider overall duration, pain, and any discomfort when rating the comfort.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec22\" class=\"Section2\"\u003e\u003ch2\u003e2.5 Statistics\u003c/h2\u003e\u003cp\u003eWe compared the frequency characteristics of the impedance among the electrode types on the hairless skin during the EMG experiment. we first performed a logarithmic transformation considering homoscedasticity. One-dimensional statistical parametric mapping (SPM1d) was used to determine whether impedance differed among the electrodes [\u003cspan citationid=\"CR37\" class=\"CitationRef\"\u003e37\u003c/span\u003e]. The SPM analysis can find statistically significant frequency portions regarding the impedance difference among the electrodes rather than compare discrete frequency. For multiple comparisons across the three electrodes, p-values were adjusted using the Bonferroni method.\u003c/p\u003e\u003cp\u003eWe also examined differences among the electrode types for the following parameters: 1) SNR during the EMG experiment, 2) SNR during the EEG experiment, 3) impedance at 10 Hz during the EEG experiment, 4) setup time, 5) removal time, 6) total time (setup\u0026thinsp;+\u0026thinsp;removal),7) comfort ratings during setup, recording, and removal, and 8) pain level. Normality was tested using the Shapiro-Wilk test. The test indicated that normality was not rejected in all parameters except SNR of EMG, impedance during the EEG experiment, and pain level. Therefore, we performed a one-way analysis of variance (ANOVA) with a parametric approach to evaluate the main effect of electrode type for normally distributed data. Post-hoc comparisons were conducted using Student\u0026rsquo;s t-test if equal variances were confirmed, or Welch\u0026rsquo;s t-test otherwise. For the parameters, which did not follow the normality, we employed the Fisher\u0026ndash;Pitman permutation test of matched pairs with 1,000 permutations for comparisons between electrode types [\u003cspan citationid=\"CR10\" class=\"CitationRef\"\u003e10\u003c/span\u003e, \u003cspan citationid=\"CR38\" class=\"CitationRef\"\u003e38\u003c/span\u003e]. All p-values were corrected for multiple comparisons using the Bonferroni correction.\u003c/p\u003e\u003c/div\u003e"},{"header":"3. Results","content":"\u003cdiv id=\"Sec24\" class=\"Section2\"\u003e\u003ch2\u003e3.1 Overview of fabricated MN electrodes and experiments for comparisons among electrode types\u003c/h2\u003e\u003cp\u003eFigure \u003cspan refid=\"Fig2\" class=\"InternalRef\"\u003e2\u003c/span\u003e illustrates the flowchart of the fabrication process for PEDOT:PSS-coated MN electrodes. Based on a previous study [\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e], we fabricated the MN electrode composed of three layers:1) a polyimide substrate layer providing flexibility, mechanical strength; 2) a Au layer to provide electrical conductivity (Fig.\u0026nbsp;\u003cspan refid=\"Fig2\" class=\"InternalRef\"\u003e2\u003c/span\u003eB); and a PEDOT:PSS conductive polymer layer (Fig.\u0026nbsp;\u003cspan refid=\"Fig2\" class=\"InternalRef\"\u003e2\u003c/span\u003eC), which decrease skin-electrode contact impedance.\u003c/p\u003e\u003cp\u003eWe then compared the performance\u0026mdash;both in terms of signal quality and usability\u0026mdash;of the MN electrode with conventional wet and dry electrodes. First, to assess performance under hairless conditions, we recorded EMG signals and electrode impedance on the forearm (Fig.\u0026nbsp;\u003cspan refid=\"Fig3\" class=\"InternalRef\"\u003e3\u003c/span\u003eA). Next, we evaluated their performance for EEG signals from hairy scalp regions using somatosensory evoked potentials (SEPs) (Fig.\u0026nbsp;\u003cspan refid=\"Fig3\" class=\"InternalRef\"\u003e3\u003c/span\u003eB). During the EEG recordings, we also recorded setup and cleanup times, as well as any pain or discomfort reported by participants, to assess usability across electrode types.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec25\" class=\"Section2\"\u003e\u003ch2\u003e3.2 EMG experiment\u003c/h2\u003e\u003cp\u003eEight healthy participants performed an elbow flexion force-matching task (Fig.\u0026nbsp;\u003cspan refid=\"Fig3\" class=\"InternalRef\"\u003e3\u003c/span\u003eA). Before measuring EMG signals, the EII spectra of the electrodes was measured. Then, prior to the main task, maximum voluntary contraction (MVC) force was measured. We then recorded 10 seconds of EMG signals under both relaxed (rest) conditions and during elbow flexion at 20% MVC.\u003c/p\u003e\u003cp\u003eFigure \u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eA shows an example of EMG signals recorded during elbow flexion at 20% MVC (top row) and at rest (bottom row). In this participant, the MN electrode produced a larger amplitude during muscle contraction and a lower noise level at rest compared to the wet and dry electrodes. Consistent with this finding, the averaged data across participants (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eB) revealed that the MN electrodes had statistically lower impedance than the dry electrodes below 750 Hz, and lower than the wet electrodes below about 150 Hz (p\u0026thinsp;\u0026lt;\u0026thinsp;0.05). In Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eC, the signal-to-noise ratio (SNR) of EMG\u0026mdash;calculated as the EMG amplitude during elbow flexion relative to that during rest\u0026mdash;was highest on average with the MN electrodes, followed by the wet electrodes and then the dry electrodes. Statistical analysis showed that the SNR with the MN electrodes was significantly higher (p\u0026thinsp;\u0026lt;\u0026thinsp;0.05) than that observed with both the wet and dry electrodes.\u003c/p\u003e\u003cp\u003e\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec26\" class=\"Section2\"\u003e\u003ch2\u003e3.3 EEG experiment\u003c/h2\u003e\u003cp\u003eNine healthy male adults participated in this experiment. Before measuring EEG signals, the EII at 10Hz were measured with a impedance measurement function of the EEG amplifier (actiCHamp Plus, Brain Products, Germany). The analyzed frequency was predetermined by the EEG system. Since 10 Hz corresponds to the alpha wave in EEGs, impedance at 10 Hz has been used to evaluate the biopotential electrodes[\u003cspan citationid=\"CR35\" class=\"CitationRef\"\u003e35\u003c/span\u003e, \u003cspan citationid=\"CR36\" class=\"CitationRef\"\u003e36\u003c/span\u003e]. We applied 200 ms monophasic constant-current pulses to the right median nerve using an electrical stimulator (USE-100, UNIQUE MEDICAL, Japan). Stimulation (300 trials at 2.1 Hz) was delivered via wrist electrodes at an intensity sufficient to evoke a clear thumb twitch (~\u0026thinsp;1\u0026ndash;2 cm).\u003c/p\u003e\u003cp\u003eAs shown in Fig.\u0026nbsp;\u003cspan refid=\"Fig6\" class=\"InternalRef\"\u003e6\u003c/span\u003eA, the impedance at 10 Hz was significantly lower in the wet (7.4\u0026thinsp;\u0026plusmn;\u0026thinsp;4.6 kΩ) and MN electrodes (6.0\u0026thinsp;\u0026plusmn;\u0026thinsp;5.3 kΩ) compared to the dry electrode (424.8\u0026thinsp;\u0026plusmn;\u0026thinsp;140.7 kΩ, p\u0026thinsp;\u0026lt;\u0026thinsp;0.05). It should be noted that, in 2 of the 9 participants, the impedance of the dry electrode exceeded the EEG system\u0026rsquo;s 500 kΩ measurement limit, even after parting the hair. There was no significant difference between the wet and MN electrodes.\u003c/p\u003e\u003cp\u003eThe Fig.\u0026nbsp;\u003cspan refid=\"Fig6\" class=\"InternalRef\"\u003e6\u003c/span\u003eB demonstrates typical example SEP response from a participant. All the electrode showed clear response around 20 ms after the stimulation, the N20 peak was larger in the MN electrode compared to the other two electrodes and the noise level (activity during the pre-stimulus period) seemed lower in the wet electrode compared to the other two electrodes. To evaluate quality of the EEG signals among the electrodes, we evaluated the SNR of SEP (Fig.\u0026nbsp;\u003cspan refid=\"Fig6\" class=\"InternalRef\"\u003e6\u003c/span\u003eC). The SNR was higher in the wet and MN electrodes compared to that in the dry electrode (p\u0026thinsp;\u0026lt;\u0026thinsp;0.05). It did not significantly differ between the wet and MN electrodes.\u003c/p\u003e\u003cp\u003e\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec27\" class=\"Section2\"\u003e\u003ch2\u003e3.4 Usability\u003c/h2\u003e\u003cp\u003ePreparation time was measured from the moment the experimenter began fitting the participant with an EEG cap until the electrode\u0026ndash;skin impedance fell below 20 kΩ for the wet and MN electrodes, and below 400 kΩ for the dry electrode. In all experiments, the same experimenter (T.Y.) was responsible for both the preparation and cleanup. Cleanup time included removing the electrodes after each measurement. For wet electrodes, it also included the time needed to wash out the conductive gel and thoroughly dry the hair. We found that the wet electrode had a significantly shorter preparation time than the other electrode types (p\u0026thinsp;\u0026lt;\u0026thinsp;0.05, Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eA). However, it required the longest cleanup time (p\u0026thinsp;\u0026lt;\u0026thinsp;0.05) compared with the other electrodes. Consequently, total time (preparation\u0026thinsp;+\u0026thinsp;cleanup) was significantly longer for the wet electrode than for the other two electrode types (p\u0026thinsp;\u0026lt;\u0026thinsp;0.05, Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eC).\u003c/p\u003e\u003cp\u003eRegarding comfort, participants reported mild comfort levels (approximately 2 to 4 on the visual analog scale (VAS), where 0 is \u0026ldquo;most comfortable,\u0026rdquo; 5 is \u0026ldquo;neutral,\u0026rdquo; and 10 is \u0026ldquo;most uncomfortable\u0026rdquo;) across all three electrode types, with no significant differences during either the preparation (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eD) or measurement phases (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eE). In contrast, during cleanup (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eF), participants indicated greater comfort with the dry (1.65\u0026thinsp;\u0026plusmn;\u0026thinsp;1.5 [mean\u0026thinsp;\u0026plusmn;\u0026thinsp;SD]) and MN (1.83\u0026thinsp;\u0026plusmn;\u0026thinsp;1.5) electrodes compared to the wet electrode (3.41\u0026thinsp;\u0026plusmn;\u0026thinsp;1.2, p\u0026thinsp;\u0026lt;\u0026thinsp;0.05).\u003c/p\u003e\u003cp\u003eFor pain (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eG), participants reported significantly less pain with the wet electrode (0.60\u0026thinsp;\u0026plusmn;\u0026thinsp;1.0) compared to the dry (2.93\u0026thinsp;\u0026plusmn;\u0026thinsp;2.1, p\u0026thinsp;\u0026lt;\u0026thinsp;0.05) and MN electrodes (1.86\u0026thinsp;\u0026plusmn;\u0026thinsp;1.6, p\u0026thinsp;\u0026lt;\u0026thinsp;0.05) on the VAS scale (0\u0026thinsp;=\u0026thinsp;no pain, 10\u0026thinsp;=\u0026thinsp;worst imaginable pain). There was no significant difference between the dry and MN electrodes.\u003c/p\u003e\u003cp\u003e\u003c/p\u003e\u003c/div\u003e"},{"header":"4. Discussion","content":"\u003cp\u003eWe quantitatively assessed the signal quality and comfort level of newly developed PEDOT:PSS-coated MN electrodes, comparing them with conventional wet and dry electrodes. Our findings indicate that the MN electrode can deliver superior signal quality in hairless regions\u0026mdash;surpassing even the wet electrode\u0026mdash;and comparable performance in hairy areas. The remarkable aspect is that, despite the MN electrode being able to measure high-quality signals equal to or better than those of wet electrodes, it is a dry electrode and does not require gel. This balance of signal quality and ease of use makes the MN electrode a promising technology for both EEG and EMG applications.\u003c/p\u003e\u003cdiv id=\"Sec29\" class=\"Section2\"\u003e\u003ch2\u003e4.1 MN electrodes on hairless region: EMG recording\u003c/h2\u003e\u003cp\u003eIn the EMG experiment, the MN electrode showed significantly lower impedance below approximately 150 Hz compared with conventional electrodes (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eB), as well as a significantly higher SNR than either type of conventional electrode (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eC). This suggests that MN electrodes, despite being dry-type, can record high-quality signals when used on hairless skin. Two key factors would contribute to this high signal quality: (1) the penetration of the high-resistance stratum corneum by tiny needles and (2) the mixed conductivity of PEDOT:PSS.\u003c/p\u003e\u003cp\u003eRegarding the first factor, MNs can indeed reduce impedance; however, MNs made solely of traditional metallic coating show higher impedance than wet electrodes [\u003cspan citationid=\"CR39\" class=\"CitationRef\"\u003e39\u003c/span\u003e]. In this study, following Li et al [\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e], we coated our MNs with PEDOT:PSS on Au, which led to a low impedance in the low-frequency band, even below that of wet electrodes (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eB). This improvement occurs because the ionic/electronic mixed conductivity of PEDOT:PSS acts as a transducer, facilitating smooth current flow between the body\u0026rsquo;s ionic environment and the electrode\u0026rsquo;s electron-conductive metal [\u003cspan citationid=\"CR28\" class=\"CitationRef\"\u003e28\u003c/span\u003e].\u003c/p\u003e\u003cp\u003eWhile Li et al. [\u003cspan citationid=\"CR27\" class=\"CitationRef\"\u003e27\u003c/span\u003e] did measure EMG signals using PEDOT:PSS-coated MN electrodes, their study did not standardize the measurement site and contraction force, resulting in a less robust quantitative comparison. In contrast, by controlling for the measurement site and contraction force on EMG signals, our results indicate that MN electrodes yield superior signal quality compared with widely used wet and dry electrodes (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eC). Therefore, under ideal conditions\u0026mdash;namely, in hairless regions\u0026mdash;the MN electrode, despite being a dry type, can surpass even conventional gel-assisted electrodes in terms of signal quality.\u003c/p\u003e\u003cp\u003eIn the comparison between conventional dry and wet electrodes, a significant difference in impedance was observed (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eB). However, there was no significant difference in the SNR of the recorded EMG signals between them. We assume that this is because the difference in impedance magnitude between the dry and wet electrodes was smaller than that between the microneedle electrode and the other electrode types (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eB). Consequently, a significant difference in SNR was observed between the microneedle electrode and the other electrodes, whereas no significant difference was likely to be observed between the dry and wet electrodes (Fig.\u0026nbsp;\u003cspan refid=\"Fig5\" class=\"InternalRef\"\u003e5\u003c/span\u003eC).\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec30\" class=\"Section2\"\u003e\u003ch2\u003e4.2 MN electrodes on hairy region: EEG recording\u003c/h2\u003e\u003cp\u003eAlthough a previous study on PEDOT:PSS-coated MN electrodes evaluated EEG signals in hairless areas such as the forehead, their usefulness in hairy regions\u0026mdash;the most common targets of EEG measurements\u0026mdash;remains largely unexplored. In this study, we attempted to measure EEG over a hairy area and found that although the MN electrode achieved better impedance and SNR than the dry electrode, it was comparable to the wet electrode (Figs.\u0026nbsp;\u003cspan refid=\"Fig6\" class=\"InternalRef\"\u003e6\u003c/span\u003eB and \u003cspan refid=\"Fig6\" class=\"InternalRef\"\u003e6\u003c/span\u003eC). This is likely because hair interferes with the electrode\u0026ndash;skin contact, thereby reducing MN electrode performance.\u003c/p\u003e\u003cp\u003eWhen using dry-type EEG electrodes, measuring in hairy areas has long been recognized as challenging [\u003cspan citationid=\"CR40\" class=\"CitationRef\"\u003e40\u003c/span\u003e]. Moreover, research on MN electrodes has shown that greater hair coverage caused higher impedance [\u003cspan citationid=\"CR41\" class=\"CitationRef\"\u003e41\u003c/span\u003e]. To address this, Kawana et al. [\u003cspan citationid=\"CR41\" class=\"CitationRef\"\u003e41\u003c/span\u003e] developed a small shutter mechanism to separate the hair and ensure reliable electrode\u0026ndash;scalp contact. Employing such an approach may enable PEDOT:PSS-coated MN electrodes to realize their full potential for EEG measurements in hairy regions.\u003c/p\u003e\u003cp\u003eAlthough the performance of MN electrode was similar to the wet electrode, it is noteworthy that the wet electrode used in this study was a high-grade, commercially available research system. Achieving comparable performance with a dry-type MN electrode is therefore highly significant in terms of balancing signal quality and convenience. Moreover, because MN electrodes do not require gel, they offer greater feasibility for long-term stable measurements [\u003cspan citationid=\"CR25\" class=\"CitationRef\"\u003e25\u003c/span\u003e]. Indeed, a study using Au-based MN electrodes for EEG showed less impedance drift over extended recordings compared to wet electrodes, leading to more stable signals and improved accuracy in decoding movement intention from brain activity [\u003cspan citationid=\"CR42\" class=\"CitationRef\"\u003e42\u003c/span\u003e]. Hence, even if PEDOT:PSS-coated MN electrodes do not exceed the signal quality of wet electrodes, their potential for prolonged, stable recording makes them promising for applications such as sleep research or brain\u0026ndash;machine interface (BMI) studies, where long wearing times are common.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec31\" class=\"Section2\"\u003e\u003ch2\u003e4.3 Usability\u003c/h2\u003e\u003cp\u003eThe MN and dry electrodes had a shorter cleanup time than the wet electrode (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eB), because they do not require a conductive gel, allowing participants to go home immediately after removing the electrodes. On the other hand, preparation time of the MN electrodes exceeded that of the wet electrode (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eA), as positioning the MN electrode took longer due to parting the hair. Incorporating a hair-separating mechanism[\u003cspan citationid=\"CR41\" class=\"CitationRef\"\u003e41\u003c/span\u003e] (as mentioned previously) might automate this step and reduce setup time.\u003c/p\u003e\u003cp\u003eIn terms of pain, the MN electrodes induced only mild pain (1.86\u0026thinsp;\u0026plusmn;\u0026thinsp;1.6 on the VAS scale, Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eG). Although this was higher than that was reported for the wet electrode (0.60\u0026thinsp;\u0026plusmn;\u0026thinsp;1.0), it did not significantly differ from the dry electrode (2.93\u0026thinsp;\u0026plusmn;\u0026thinsp;2.1). Because the pain level for the MN electrode remained relatively low, no significant differences were observed among the three electrode types regarding comfort/discomfort level during measurement (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eE). Overall, despite having tiny needles, the MN electrode does not cause severe pain, and the user experience is similar to that of a conventional dry electrode.\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec32\" class=\"Section2\"\u003e\u003ch2\u003e4.4 Methodological consideration\u003c/h2\u003e\u003cp\u003eThere is a thing to note regarding our impedance results. In the hairless region, we used a frame to standardize the contact area for all electrode types, but in the hairy region, no frame was used. Consequently, gel could have spread under the electrode in the wet electrode condition, potentially creating a larger contact area than that of the MN electrode and, in turn, overestimating the wet electrode\u0026rsquo;s impedance performance.\u003c/p\u003e\u003cp\u003eSince the MN electrode penetrates the skin, we fixed the measurement order to wet, dry, and then MN, which may introduce order effects. In the EMG experiment, there was a concern on muscle fatigue, but previous research suggests that at 20% MVC, the biceps brachii typically becomes fatigued after about 10 min of sustained contraction [\u003cspan citationid=\"CR43\" class=\"CitationRef\"\u003e43\u003c/span\u003e]. Because our contractions lasted only 10 seconds, the impact of fatigue was likely minimal. In the EEG experiment, we repeated SEP sessions in a fixed order, as in the EMG experiment. To check order effects on SEP response size, we conducted a supplemental experiment of three repeated SEP sessions using wet electrodes for all sessions. The session interval was 8 minutes, which is similar to the total time spent on cleanup and preparation in the main experiment (Fig.\u0026nbsp;\u003cspan refid=\"Fig7\" class=\"InternalRef\"\u003e7\u003c/span\u003eC). The supplemental experiment showed no clear order effects on SEP response (Figure S1).\u003c/p\u003e\u003cp\u003e\u003c/p\u003e\u003c/div\u003e\u003cdiv id=\"Sec33\" class=\"Section2\"\u003e\u003ch2\u003e4.5 Future direction\u003c/h2\u003e\u003cp\u003eThe MN electrodes have a potential to provide high-quality yet convenient EEG measurements, they may be very useful for clinical applications. For instance, in a previous study [\u003cspan citationid=\"CR44\" class=\"CitationRef\"\u003e44\u003c/span\u003e], walking intentions extracted from the EEG of paraplegic patients triggered electrical stimulation of leg muscles, achieving partial restoration of walking function. However, the authors noted that using a conventional gel-based wet EEG system in everyday settings was challenging, mainly due to time and effort constraints, which could limit its widespread adoption in gait rehabilitation. Although a faster setup in hairy regions remains needed, MN electrodes could simplify EEG recording while maintaining high signal quality, thereby advancing routine clinical use.\u003c/p\u003e\u003cp\u003eIn EMG measurements, MN electrodes showed better signal quality than conventional electrodes (Fig.\u0026nbsp;\u003cspan refid=\"Fig4\" class=\"InternalRef\"\u003e4\u003c/span\u003eC). Additionally, their low impedance suggests that reducing electrode size might still yield good-quality signals (Fig.\u0026nbsp;\u003cspan refid=\"Fig4\" class=\"InternalRef\"\u003e4\u003c/span\u003eB). Consequently, MN electrodes could be highly valuable for high-density surface electromyography (HDsEMG), which has been actively studied recently [\u003cspan citationid=\"CR23\" class=\"CitationRef\"\u003e23\u003c/span\u003e, \u003cspan citationid=\"CR45\" class=\"CitationRef\"\u003e45\u003c/span\u003e]. Achieving high density EMG recordings requires miniaturizing electrodes, and the MN electrodes are well suited to that purpose. HDsEMG gains attention when used with blind source estimation, since it enables non-invasive investigation of spinal motor neuron firing, and the number of neurons that can be identified increases with higher SNR [\u003cspan citationid=\"CR29\" class=\"CitationRef\"\u003e29\u003c/span\u003e, \u003cspan citationid=\"CR46\" class=\"CitationRef\"\u003e46\u003c/span\u003e]. Therefore, applying MN electrodes to HDsEMG could increase the ability to evaluate motoneuron activity via surface EMG.\u003c/p\u003e\u003c/div\u003e"},{"header":"5. Conclusion","content":"\u003cp\u003eWe fabricated PEDOT:PSS coated MN electrodes and demonstrated experimentally that they can measure EMG signals with higher quality than conventional electrodes, and in EEG recordings they can match the signal quality of wet electrodes while offering dry-electrode convenience. These findings highlight the practicality of MN electrodes and indicate their strong potential for use in clinical EMG and EEG applications. It is expected that future studies apply the MN electrodes in rehabilitation BMI systems and high-density EMG recording, expanding their utility across a wide range of bio-signal measurement contexts.\u003c/p\u003e"},{"header":"Declarations","content":"\u003cp\u003e\u003cstrong\u003eEthics approval and consent to participate\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eAll participants provided written informed consent for participation in the study, which was conducted in accordance with the Declaration of Helsinki and approved by the Ethics Review Committee of The University of Tokyo (reference number, 23-396).\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eConsent for publication\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eNot applicable\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAvailability of data and materials\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eData and code reported in this article will be shared by the lead contact upon request.\u003c/p\u003e\n\u003cp\u003eAll original code reported in this article will be shared by the lead contact upon request.\u003c/p\u003e\n\u003cp\u003eAny additional information required to reanalyze the data reported in this article is available from the lead contact upon request.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eCompeting interests\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eThe authors declare no competing interests.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eFunding\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eThis work was supported by JSPS KAKENHI Grant in Aid for Scientific Research (B) to H.Y.(21H03340), JSPS KAKENHI WAKATE to N.K. (50969285), Nakatani Foundation for Advancement of Measuring Technologies in Biomedical Engineering to H.Y. and N.K., JKA through its promotion funds from KEIRIN RACE to H.Y., and Tateisi Science and Technology Foundation to H.Y. and N.K.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAuthors' contributions\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eT.Y., H.Y., and K.N. designed the research. T.Y., N.K., and Y.K. performed device fabrication. H.T. and H.Y. constructed the experimental setups. T.Y., E.N., and H.Y. collected data. T.Y. performed analyses and drafted the manuscript. All authors interpreted the data, discussed the findings, and approved the final version of the manuscript.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eAcknowledgements\u003c/strong\u003e\u003c/p\u003e\n\u003cp\u003eNot applicable\u003c/p\u003e"},{"header":"References","content":"\u003col\u003e\n\u003cli\u003eTeplan M. Fundamentals of EEG measurement. 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Comput Biol Med. 2019;109:171\u0026ndash;81. \u003c/li\u003e\n\u003c/ol\u003e"}],"fulltextSource":"","fullText":"","funders":[],"hasAdminPriorityOnWorkflow":false,"hasManuscriptDocX":true,"hasOptedInToPreprint":true,"hasPassedJournalQc":"","hasAnyPriority":false,"hideJournal":true,"highlight":"","institution":"","isAcceptedByJournal":false,"isAuthorSuppliedPdf":false,"isDeskRejected":"","isHiddenFromSearch":false,"isInQc":false,"isInWorkflow":false,"isPdf":false,"isPdfUpToDate":true,"isWithdrawnOrRetracted":false,"journal":{"display":true,"email":"[email protected]","identity":"researchsquare","isNatureJournal":false,"hasQc":true,"allowDirectSubmit":true,"externalIdentity":"","sideBox":"","snPcode":"","submissionUrl":"/submission","title":"Research Square","twitterHandle":"researchsquare","acdcEnabled":true,"dfaEnabled":false,"editorialSystem":"","reportingPortfolio":"","inReviewEnabled":false,"inReviewRevisionsEnabled":true},"keywords":"Electroencephalography (EEG), Electromyography (EMG), Brain, Muscle, Biosignal","lastPublishedDoi":"10.21203/rs.3.rs-7343232/v1","lastPublishedDoiUrl":"https://doi.org/10.21203/rs.3.rs-7343232/v1","license":{"name":"CC BY 4.0","url":"https://creativecommons.org/licenses/by/4.0/"},"manuscriptAbstract":"\u003cp\u003e\u003cstrong\u003eBackground:\u003c/strong\u003e Bioelectrical signals are vital indicators of physiological function, psychological status, and clinical conditions. However, traditional wet electrodes using conductive gel are time-consuming, labor-intensive, and degrade over time due to gel drying. Recently, microneedle (MN) electrodes coated with poly(3,4‐ethylene dioxythiophene):poly(styrenesulfonate) (PEDOT:PSS) have been developed for high-quality signal acquisition. Although low electrode–skin impedance has been demonstrated on hairless regions, their use in electroencephalography (EEG) recordings on hairy scalp areas and quantitative comparisons with conventional electrodes in both EEG and electromyography (EMG) remain limited.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eMethods: \u003c/strong\u003eWe fabricated two MN electrode types of different lengths—one for EMG on hairless skin and one for EEG on the hairy scalp—and compared them with conventional wet and dry electrodes. EMG was recorded during a force-matching task; EEG was assessed via somatosensory evoked potentials. We also evaluated setup/cleanup time, comfort, and pain during EEG measurement.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eResults: \u003c/strong\u003eFor EMG, MN electrodes achieved significantly higher signal-to-noise ratios than conventional electrodes. For EEG, they outperformed dry electrodes and matched wet electrodes in signal quality—without using conductive gel. Although additional time was required to part hair, cleanup was faster due to the absence of gel. Pain was comparable to dry electrodes.\u003c/p\u003e\n\u003cp\u003e\u003cstrong\u003eConclusions: \u003c/strong\u003ePEDOT:PSS-coated MN electrodes provided superior EMG signal quality and high-quality EEG signals comparable to wet electrodes even without gel. These findings suggest that PEDOT:PSS-coated MN electrodes offer a compelling balance of signal quality and user convenience, making them especially advantageous for real-world and clinical applications where time efficiency, minimal discomfort, and gel-free operation are critical.\u003c/p\u003e","manuscriptTitle":"Quantitative Assessment of Signal Quality and Usability of EEG and EMG recordings with PEDOT:PSS-Coated Microneedle Electrodes","msid":"","msnumber":"","nonDraftVersions":[{"code":1,"date":"2025-08-19 13:05:42","doi":"10.21203/rs.3.rs-7343232/v1","editorialEvents":[{"type":"communityComments","content":0}],"status":"published","journal":{"display":true,"email":"[email protected]","identity":"researchsquare","isNatureJournal":false,"hasQc":true,"allowDirectSubmit":true,"externalIdentity":"","sideBox":"","snPcode":"","submissionUrl":"/submission","title":"Research Square","twitterHandle":"researchsquare","acdcEnabled":true,"dfaEnabled":false,"editorialSystem":"","reportingPortfolio":"","inReviewEnabled":false,"inReviewRevisionsEnabled":true}}],"origin":"","ownerIdentity":"e0d0af99-ef02-483f-8ec9-15fa6a1156bf","owner":[],"postedDate":"August 19th, 2025","published":true,"recentEditorialEvents":[],"rejectedJournal":[],"revision":"","amendment":"","status":"posted","subjectAreas":[],"tags":[],"updatedAt":"2025-09-11T17:23:33+00:00","versionOfRecord":[],"versionCreatedAt":"2025-08-19 13:05:42","video":"","vorDoi":"","vorDoiUrl":"","workflowStages":[]},"version":"v1","identity":"rs-7343232","journalConfig":"researchsquare"},"__N_SSP":true},"page":"/article/[identity]/[[...version]]","query":{"redirect":"/article/rs-7343232","identity":"rs-7343232","version":["v1"]},"buildId":"8U1c8b4HqxoKbykW_rLl7","isFallback":false,"isExperimentalCompile":false,"dynamicIds":[84888],"gssp":true,"scriptLoader":[]}