pH calibration allows accurate glucose detection in interstitial fluid via reverse iontophoresis

Domain: physics
Published: 2025-11-24
Editorial outcome: major_revision

Abstract
Reverse iontophoresis (RI) is a promising non-invasive, wearable technology for the transdermal extraction of interstitial fluid (ISF), which contains rich biomarkers relevant to health status. Despite the advancement of wearable sensors, this technology is still restricted for accurate non-invasive biomarkers detection. The main challenge lies in the instability of ISF extraction during RI. We found that this instability is primarily caused by the skin surface pH variations because of the interaction between the RI-induced H+ movement and the skin recovery ability. Here, we investigated how the skin surface pH affected RI, theoretically and experimentally; and developed a wearable device and a calibration method to enable accurate non-invasive ISF glucose detection, accordingly. The results showed that glucose prediction accuracy was markedly improved, with mean absolute relative difference (MARD) decreased from 34.44% to 14.78% across both healthy and diabetic volunteers.

Introduction
With rising living standards and economic progress, personalized diagnostic technologies are gaining increasing attention for their potential in health assessment and early disease diagnosis1–3. These technologies typically involve two critical dimensions and trends: wearable, non-invasive monitoring and accurate prediction of physiological status. The breakthroughs of flexible electronics and nanotechnology in recent years may address the requirements for wearables4,5, however, it is still a big challenge to achieve accurate and non-invasive prediction6. Biomarkers, such as metabolites7–9, hormones10,11, electrolytes12, and proteins13,14, serve as pivotal roles in achieving accurate prediction of physiological status6,15,16. These biomarkers, derived from biofluids such as sweat17–20, tears21, saliva22, urine23, and interstitial fluid (ISF)24–27, have already been explored for wearable and non-invasive health monitoring. Notably, among these biofluids, ISF contains the most categories of biomarkers, and its composition is closely correlated with blood-derived information28,29. Therefore, we think that ISF may be the best choice to achieve accurate and non-invasive biomarker detection.

Non-invasive ISF detection is currently achievable through radiofrequency30, ultrasound31, optics32, metabolic heat conformation33, and reverse iontophoresis (RI)34–37. The former four methods face challenges for accuracy due to indirect measurement and interferences from complex tissue, especially from omnipresent blood38. In contrast, RI provides direct extraction and detection of ISF, holding great potential for accurate analysis of biomarkers24. RI, first introduced by Richard H. Guy’s group in 1989, has undergone significant advancements over the years39. In 2001, Cygnus Inc. launched the GlucoWatch® Biographer, the first and only commercial device based on RI, demonstrating the potential for wearable and non-invasive glucose detection40. Unfortunately, the device was discontinued in 2007 due to several limitations, including skin irritation, sweat-induced interference, discomfort of wearing, measurement inaccuracies from sensor-skin displacement by rigid components, and instability of ISF extraction41. Recent advancements in flexible electronics and nanotechnology have addressed many of these challenges, such as skin irritation42, sweat-induced interference38, and discomfort of wearing5,43–45, reigniting interest in RI for wearable and non-invasive biomarkers detection46–49. Despite these innovations, the issue with accurate detection remains. We found that the key issue lies in the fluctuations of ISF extraction34.

Under the influence of the skin’s physiological negative charge, RI facilitates the transdermal extraction of ISF by applying a mild current to the skin50. This process is driven primarily by the electromigration of cations in the electrical double layer (EDL), thereby making the electroosmotic flow of small neutral biomarkers, for example, glucose51. Here, we found that the EDL at the interface between the skin and the RI electrodes significantly affects this electroosmotic flow. Specifically, the extracted H⁺, along with the release of additional H⁺ from biochemical reactions, collectively lead to changes in the pH of the RI-induced fluid accumulated at the skin surface and within the epidermal electroosmotic microchannel (referred to as the skin surface pH in this study). The skin recovery mechanisms, including natural barriers such as sweat glands and the stratum corneum that maintain homeostasis and prevent excessive polarization52–54, will hinder the change. The interaction leads to the fluctuations of skin surface pH, which changes the zeta potential (ZP) of the EDL55,56, thereby causing instability in the electroosmotic flow of ISF accordingly.

To validate this hypothesis, here we provide a detailed theoretical analysis of how skin surface pH influences ISF extraction by modulating the ZP of keratin, thereby directly affecting ISF extraction and subsequent detection accuracy. We conducted a numerical simulation to elucidate these effects; we designed and fabricated an epidermal electronic device integrating symmetric ISF extraction electrodes, glucose, Na+, and pH sensors for experimental investigation. Theoretical experiments were conducted firstly to validate our hypothesis. Then, a skin surface pH calibration method was developed and validated to enable accurate and non-invasive ISF glucose detection through in-vitro and in-vivo experiments.

Results
Theoretical analysis of effect of skin surface pH on RI

Theoretical analysis. Human skin is composed of dermis, epidermis, stratum corneum (SC), and hydrolipid film (HF) that overlay the SC (Fig. 1a). At the interface between the HF and the SC, charge separation occurs due to the interaction of solid and liquid phases. This phenomenon leads to a spatial distribution of potential differences near the interface, thereby establishing the EDL structure, which comprises a Stern layer and a diffusion layer. RI extraction of ISF hinges on two key mechanisms: the electromigration of ions and the electroosmosis of small neutral molecules. The electromigration involves ionic fluxes originating from both the bulk ISF and the diffusion layer. In contrast, the driving force for the electroosmotic transport of small neutral molecules primarily arises from the migration of counterions (carrying charges opposite to the skin surface) within the diffuse layer under an applied electric field. Consequently, the concentration of these counterions governs the electroosmotic velocity (Fig. 1b, Supplementary Note 1).

The skin’s inherent negative charge under physiological conditions results in cation-dominated ion flux (counterions) within the EDL during RI. This cation flux facilitates the movement of small neutral molecules, e.g., glucose, towards the cathode. In this process, Na+ dominates the ion flux and acts as the primary carrier in the EDL, while extracted H+ and released H+ from biochemical reactions at the epidermal interface alter the skin surface pH. Additionally, the skin recovery mechanisms, including natural barriers such as sweat glands and the stratum corneum, help maintain homeostasis and prevent excessive polarization52–54, thereby mitigating the pH changes induced by RI. The interaction between the H+ movement and the skin recovery ability make the fluctuations of skin surface pH. Fluctuations in skin surface pH indeed modulate the charge density of the skin, affecting the quantity of counterions (mainly Na+) in the EDL diffusion layer, which in turn lead to the fluctuations of electroosmosis.

The charge density of the skin can be reflected by the ZP (the potential at the slip plane)55. Particularly, the negative charge of skin originates from the acidic nature of the HF, which is induced by the sweat secretion, sebum degradation, and microbial metabolites57,58. The ionization state of carboxylic groups within the keratin of SC varies with pH, exhibiting a neutral state at the isoelectric point (pI) where the pH ranges from 4 to 4.5, a positive charge state below the pI while a negative charge state exceeds the pI59,60 (Supplementary Fig. 1a). We preliminarily explored the influence of skin surface pH on ZP using keratin (main components of SC), and the results revealed the great influence and a negative correlation between ZP and pH (Fig. 1c, Supplementary Table 1).

Numerical simulation. Here, we use glucose as an example to investigate the impact of pH variations on electroosmotic flow, an ISF extraction model was developed in COMSOL, incorporating ZP measurements across different pH levels (Fig. 1c, Supplementary Note 2 and Supplementary Table 1). Simulation results revealed that at pH 4, the glucose concentration at the anode is higher than at the cathode. Conversely, when the pH exceeds 4, the glucose concentration at the cathode becomes significantly higher than that at the anode, and both electrodes exhibit a rise in glucose concentration as pH increased, aligning with the ZP-dependent extract mechanism by RI (Fig. 1d, e, Supplementary Fig. 1b–h).

Design and characterization of the epidermal electronic device

Design and fabrication of the wearable device. To validate our hypothesis and theoretical analysis, we designed a wearable electronic device accordingly, which incorporated a pair of RI extraction electrodes for ISF extraction, glucose sensors, Na⁺ sensors and pH sensors for fluxes measurement and skin surface pH monitoring. Specifically, the device is divided into two symmetrical sensing regions, Part A and Part B. Each part comprises a semi-circular RI electrode, a glucose sensor, a Na⁺ sensor, and a pH sensor (Fig. 2b). The layer assembly of the device consists of a flexible polyimide (PI) substrate layer, enabling seamless conformity to skin deformations for comfortable wearing; an Ag conductor layer to ensure stable electrical connections; a sensors layer for simultaneous ISF extraction, glucose, Na⁺, and pH detection; a patterned 3M adhesive layer for attachment to the skin and precise hydrogel patterning; and a hydrogel layer serving as the ISF collection reservoir and microreaction chamber (Fig. 2c, d). The epidermal device was fabricated by screen printing, followed by layer-by-layer surface functionalization (Supplementary Fig. 4).

Characterization of the wearable device. We firstly conducted comprehensive characterizations of the proposed device to validate its functionality. Scanning Electron Microscopy (SEM) and X-ray Photoelectron Spectroscopy (XPS) of the electrodes confirm the successful fabrication and functionalization of the device (Supplementary Figs. 6, 7). The detection principle of this device involves two electrochemical mechanisms: an enzyme-catalyzed reaction that converts glucose concentration into amperometric current (i-t) under the applied bias voltage; and a Nernstian response that converts Na+ concentration and pH into potentiometric voltage (Fig. 2a).

H₂O₂, an intermediate product of glucose oxidation, was initially measured to evaluate the performance of the electrodes, which exhibited excellent performance with a sensitivity of 0.44 μA mM⁻¹ (Supplementary Fig. 8). Then, a series of comprehensive tests were conducted on the glucose sensor, including stepwise, sensitivity, consistency, selectivity, and storage stability. Stepwise testing revealed that the steady-state current of the i-t response increased with glucose concentration (Fig. 2e). The glucose sensor exhibited a highly linear response, with a sensitivity of 236.42 nA mM⁻¹ from 0–3 mM (Fig. 2f) and 28.22 nA mM⁻¹ from 3 to 30 mM (Supplementary Fig. 9d). Sensor-to-sensor consistency was well maintained with minimal variation in sensitivity across three devices (Supplementary Fig. 9h). The sensor also demonstrated high selectivity against a range of potential interfering species (Fig. 2g, Supplementary Fig. 9e–g). After 14 days of storage, the sensor retained ~75% of its original sensitivity (Fig. 2h, Supplementary Fig. 9i). Considering that the fluctuations in skin surface pH during RI process affect the performance of the glucose sensor, sensitivity testing was performed under different pH conditions (Fig. 2i, Supplementary Fig. 10e). The relationship between the sensitivity and pH presented as a quadratic curve. At the pH of 7, normal pH of ISF, the sensitivity reached 298.56 nA mM⁻¹ (Fig. 2j). The ultraviolet (UV) absorbance (ABS) of H₂O₂ generated by glucose oxidase (GOx) at varying pH values exhibited a comparable trend in sensitivity (Supplementary Fig. 10a–d), revealing that pH impacts the performance of glucose sensor through diminishing the activity of GOx. Consequently, prior to calculating the glucose concentration through sensitivity, a pH correction for sensitivity is required first, which is essential for accurate glucose detection. The test results for different pH values and glucose concentrations demonstrated that higher glucose detection accuracy was attained after pH correction (Fig. 2k, l, Supplementary Fig. 11). Similar to the glucose sensor, the Na⁺ sensor exhibited robust performance across the testing of step-response (Fig. 2m), sensitivity (52 mV decade⁻¹) (Fig. 2n), stability (Supplementary Fig. 12a), consistency (Supplementary Fig. 12d), selectivity (Fig. 2o, Supplementary Fig. 12c), and storage life (Fig. 2p, Supplementary Fig. 12e). The pH sensor demonstrated excellent outcomes in multiple evaluations: step-response (Fig. 2q, Supplementary Fig. 13c), sensitivity (−52 mV decade⁻¹) (Fig. 2r, Supplementary Fig. 13a), stability (Supplementary Fig. 13b), consistency (Supplementary Fig. 13e), selectivity (Fig. 2s, Supplementary Fig. 13d), and storage life (Fig. 2t, Supplementary Fig. 13f). Utilizing an environment-controlled flexible electronics system (FE-SDT), the device maintained functionality during mechanical deformation, thus making it ideal for wearable applications (Supplementary Figs. 14, 15). These results confirm the robust performance of the proposed wearable device.

Experimental validation of effect of skin surface pH on RI

Na+ flux under different skin surface pH. To further corroborate the hypothesis regarding the influence of skin surface pH on RI, validation experiments were carried out employing a Franz diffusion cell system in conjunction with a Strat-M® membrane to simulate the ISF extraction process (Fig. 3a, Supplementary Fig. 16). Tris-HCl buffer solutions at varying pH levels were added into the receiver chamber to simulate different skin surface pH environments, while NaCl solutions containing glucose were placed in the donor chamber to simulate ISF. Since the extraction process of RI involves the Na+ flux by electromigration and the glucose flux by electroosmosis, and the primary counterions in a physiological state are Na+, the impact of skin surface pH on the Na+ flux was explored firstly. The significant difference in potential response of the Na⁺ sensor in the receiver chamber before and after 15 min of RI extraction clearly confirmed the effective Na⁺ extraction (Fig. 3c). Notably, the extracted Na⁺ flux exhibited minimal variation across different pH levels, confirming that changes in skin surface pH has a negligible impact on the Na+ flux (Fig. 3d, Supplementary Table 3). The reason may be that the electromigration of Na+ predominantly occurs in the bulk solution, while the pH value only affects the Na+ flux in the diffusion layer of the EDL.

Glucose flux under different skin surface pH. Experiments were also conducted to explore the impact of skin surface pH on the electroosmotic extraction of glucose. The i-t responses of the glucose sensor before and after extraction showed a substantial increase in the steady-state current, underscoring the effective glucose extraction (Fig. 3e, Supplementary Figs. 17–22, and Supplementary Table 4). After converting this current difference into the extracted glucose flux, it was discovered that the extracted glucose flux rose progressively as the increasement of skin surface pH. Moreover, higher glucose levels in the subcutaneous ISF at 30 mM (Fig. 3g, Supplementary Fig. 23b) led to greater extracted glucose flux compared to the 15 mM one (Fig. 3f, Supplementary Fig. 23a). The linear regression analysis across the entire pH range 5–10 showed a slope significantly different from zero (p=0.00207 at 15 mM and p=0.01343 at 30 mM) and a high determination coefficient (R²=0.927 at 15 mM and R²=0.817 at 30 mM), indicating a clear overall upward trend in extracted glucose with increasing pH. And the results demonstrated that, in the cases of subcutaneous glucose levels at 15 mM and 30 mM, the extracted glucose concentration was increased by 0.013 mM and 0.027 mM per each 1 pH unit increment, respectively (Fig. 3h). Interestingly, the increase ratios are both ~33%/pH, confirming the uncorrelation between the electroosmotic velocity and the subdermal glucose concentrations. These results verify that skin surface pH affects the electroosmotic flow of glucose significantly, requiring a calibration model to achieve accurate glucose prediction.

Skin surface pH calibration for subdermal glucose prediction

The skin surface pH calibration method presented in this work tackles two crucial aspects: the correction of glucose sensor sensitivity and the calibration of glucose extraction flux (Fig. 3b). The former rectifies the change in enzyme activity resulting from pH variations, which directly influence the performance of the glucose sensor. The sensitivity of the glucose sensor exhibits a quadratic relationship with pH as described by equation (1.10). The latter centers on mitigating the deviations in glucose extraction flux linearly induced by skin surface pH changes. In this RI process, Na+ flux induces glucose flux; therefore, it serves as the base of subdermal glucose prediction, and it also indicates the different skin properties of different person38,61. The comprehensive calibration and subdermal glucose prediction equation is presented as equation (1.25), with detailed derivations in the Supplementary Note 3.

In-vitro evaluation of the calibration method and wearable device

Dynamic changes in glucose and skin surface pH during RI extraction. Capitalizing on the established linear relationship between skin surface pH and glucose extraction flux, we performed in-vitro experiments to appraise the precision of glucose prediction both with and without pH calibration (Fig. 4a, Supplementary Fig. 24). Artificial skin was used in the in-vitro experiments, it was assumed that the physical properties of the skin remained constant during the RI process, thus the Na+ flux would be constant at a certain extraction condition. Therefore, the Na+ flux was not tested here. To minimize errors from repeated ISF extractions, glucose and pH measurements from five consecutive cycles were used for validation. The subcutaneous simulated ISF was a solution containing 140 mM NaCl with glucose concentrations of 6, 12, 18, 24, and 30 mM, respectively. The in-vitro testing findings of this device were categorized into three segments: Part A, Part B, and the overall test results. When the subcutaneous glucose concentration was 12 mM, the steady-state current of the i-t responses acquired from Part A and Part B increased significantly over the five consecutive RI extractions (Fig. 4b, e, Supplementary Figs. 30, 31). Simultaneously, the skin surface pH detected by the pH sensors in both parts exhibited significant fluctuations (Fig. 4c, f, Supplementary Figs. 30, 31). The extracted glucose concentrations were linearly fitted against the extraction times of RI. Herein, the slope (denoted as Slope 1) derived from the regression represents the extracted glucose by a single instance of RI. Notably, after applying pH calibration, the slope 1 for Part A and Part B exhibited a marked increase, with the Slope 1 for Part A augmenting by 8.1% (Fig. 4d) and that for Part B growing by 7.5% (Fig. 4g). The different increase between the Part A and Part B may be caused by the different changes of electrical property at the interface of artificial skin and electrodes. The increase reflects the suppressive effect of declining skin surface pH on glucose flux during consecutive RI cycles, which is essential for subsequent prediction of subcutaneous glucose concentrations. The extraction results at subdermal glucose concentrations of 6 mM (Supplementary Figs. 28, 29), 18 mM (Supplementary Figs. 32, 33), 24 mM (Supplementary Figs. 34, 35), and 30 mM (Supplementary Figs. 36, 37) were also recorded and analyzed (Supplementary Table 4).

Relationship between extracted and subcutaneous glucose concentrations. A linear fit of the extracted glucose (represented by the fitted slope 1) against the subcutaneous glucose concentrations revealed a consistent positive linear correlation (denoted as slope 2) for both Part A and Part B of the device (Supplementary Fig. 38). The Slope 2 serves as the foundation for subsequent subcutaneous glucose prediction in the experiments. The results showed that following pH calibration, the Slope 2 exhibited an increment, with an increase of 12.1% in Part A and 6.2% in Part B of the device (Supplementary Fig. 38). To achieve higher-frequency measurements, which are of great importance for practical wearable applications, the measurements from Part A and Part B are combined to obtain the overall results (Supplementary Fig. 27). The results further confirmed the linear relationship with 9.9% increase of slop after pH calibration (Fig. 4h), underlining the impact of pH calibration on subsequent glucose prediction.

Subcutaneous glucose prediction using the calibration method and wearable device. After obtaining the relationships of extracted glucose to simulated subcutaneous glucose with and without pH calibration, we performed RI extraction and detection on additional solutions with pre-determined subcutaneous glucose concentrations. Specifically, solutions with subcutaneous glucose concentrations of 5 mM, 8 mM, 12 mM, 18 mM, and 25 mM were individually subjected to our experimental procedures. The results demonstrated a substantial elevation in glucose prediction accuracy after pH calibration, with the mean absolute percentage error (MAPE) decreased markedly from 11.2% to 2.9% (Fig. 4i). This improvement highlights the efficacy of skin surface pH calibration in reducing the inaccuracies caused by pH fluctuations during RI. These experiments serve as a preliminary validation of the proposed skin surface pH calibration method and device.

In-vivo evaluation of the calibration method and wearable device

To further validate the applicability of the wearable device and the pH calibration method in practice, oral glucose tolerance tests (OGTT) were performed on 6 healthy participants and 15 Type Ⅱ diabetic patients (Fig. 5a). During the 2.5-h experiment, fluctuations were observed in the extracted glucose flux, skin surface pH, and Na⁺ flux. The slight fluctuations in Na⁺ extraction are likely attributable to changes in skin permeability induced by prolonged RI extraction38. These three curves showed abrupt changes between consecutive data points throughout the test, which were attributed to the alternating polarity of ISF extraction (Fig. 5b, c left, Supplementary Fig. 39). Glucose prediction accuracy was evaluated under various calibration conditions: without calibration, with Na⁺ calibration only, and with cooperative Na⁺ and pH calibration (Fig. 5b, c right, Supplementary Fig. 39). The results showed that after cooperative calibration, the glucose prediction curve closely aligned with the reference curve. Clarke’s error grid analysis of healthy participants demonstrated a progressive increase in the percentage of data points falling within zone A: from 56.67% without calibration to 58.34% with Na⁺ calibration (61.54% reported by Pu et. al.38), and 78.33% with cooperative Na⁺ and pH calibration (Fig. 5d, g, Supplementary Table 6). The results of diabetic patients showed that the percentage of data points falling within zone A and B: from 70% without calibration to 90% with cooperative Na⁺ and pH calibration (Fig. 5h, k, Supplementary Table 6). Mean absolute relative difference (MARD) of healthy individuals further corroborated this trend, significantly decreasing from 25.49% (without calibration) to 20.04% (Na⁺ calibration) and 10.91% (cooperative calibration) (MARD≤20% by FDA) (Fig. 5g). The results of diabetic patients showed that MARD was changed from 42.11% (without calibration) to 33.24% (Na⁺ calibration) and 18.10% (cooperative calibration) (Fig. 5k). Following cooperative calibration, the MARD observed across all volunteers was 14.78% (Supplementary Fig. 41). These results indicate that skin surface pH calibration plays a vital role in accurate predication of subdermal glucose based on RI, especially for prolonged extraction. The results also highlight the potential of the developed device and the proposed pH calibration method as a wearable and non-invasive tool for continuous biomarkers monitoring.

Methods
Materials
Glucose, sodium chloride (NaCl), potassium chloride (KCl), calcium chloride (CaCl₂), ascorbic acid (AA), uric acid (UA), dopamine (DA), Nafion solution, polyimide (PI) film, keratin azure (K8500), glucose oxidase (GOx), polyvinyl butyral (PVB), poly(3,4-ethylenedioxythiophene)-polystyrenesulfonate (PEDOT:PSS), and platinum nanoparticles (PtNPs) were purchased from Sigma-Aldrich (Shanghai, China). Chitosan (CS), poly(vinyl chloride) (PVC), industrial-grade multi-walled carbon nanotubes (MWCNTs), glacial acetic acid, glycerol, Na⁺ carrier X, sodium tetrakis[3,5-bis(trifluoromethyl)phenyl] borate (Na-TFPB), dioctyl sebacate (DOS), Hydrogen Ionophore I (Tridodecylamine), potassium tetraphenylborate (KTClPB), cyclodextrin (CYC), Polyvinyl butyral (PVB), Pluronic F-127, and carbinol were prepared by HEOWNS (Tianjin, China). Deionized water (DI), phosphate-buffered saline (PBS, pH 7.2–7.4), and tris(hydroxymethyl) aminomethane (Tris)-HCl buffer were prepared by Huaying Prosperous Technology Co., Ltd. (Tianjin, China). All reagents were of analytical grade and used without further purification. Polyimide (PI) film was purchased from Leviathan Technology Co., Ltd (Tianjin, China). Ag/AgCl ink was supplied by Sun Chemical Co., Ltd. (UK). Prussian blue-doped carbon (PBC) ink and Ag conductive ink were provided by Gaintzor Electronic Technology Co., Ltd. (Qingdao, China). Medical adhesive tape (3M-2477P) was acquired from 3M Medical Materials & Technologies (USA). Connectors (65801-008LF) were purchased from Amphenol (USA). Franz diffusion cells were purchased from Shanghai Chemical Laboratory Equipment Co., Ltd. (Shanghai, China). Strat-M® membranes (47 mm disks, 300 μm thickness) for transdermal ISF extraction were obtained from EMD Millipore Corporation (Billerica, USA). The TPY-2 transdermal diffusion apparatus was purchased from Huanghai Pharmaceutical Testing Instruments Co., Ltd. (Shanghai, China) for precise temperature control and magnetic stirring of the ISF. A constant current source (ZF-9, Shanghai Zhengfang Electronic & Electrical Co., Ltd.) was used to provide a 150 μA cm⁻² current density for RI extraction. All electrochemical characterizations were performed with an electrochemical workstation (CHI 660E, Shanghai Chenhua Instrument Co., Ltd.).

Fabrication and preparation of the wearable device
Screen printing process. The fabrication process of the device is depicted in Supplementary Fig. 4. Firstly, the surface of the PI film (50 μm thickness) was cleaned using an alcohol wipe to remove surface contaminants. After cleaning, the PI substrate was subjected to O₂ plasma treatment (TS-PL10MA, Shenzhen Tonson Tech Automation Equipment Co., Ltd.) to eliminate organic impurities and enhance the surface’s hydrophilicity. This treatment was carried out under radio frequency power of 300 W, with a flow rate of 100 mL min⁻¹ for 90 s. The treatment resulted in improved surface hydrophilicity (Supplementary Fig. 2), which is critical for subsequent electrode printing and functional modification. Electrodes were printed on the treated PI substrate using a screen-printing method (420 mesh screen, HP-AL series, Shanghai Xuanting Screen Printing Equipment Co., Ltd.), which was performed in a layer-by-layer process. The initial step was printing an Ag conducting layer on the PI substrate, which was essential for ensuring good electrical conductivity. Then an Ag/AgCl layer was deposited on the Ag conducting layer, serving as the ISF extraction electrode and, importantly, the reference electrode (RE) for the glucose, Na⁺, and pH sensors. Finally, a PBC layer was added to function as the working electrode (WE) for the glucose, Na⁺, and pH sensors, thus completing the construction of this screen-printed electrodes fabrication. Each printed layer was cured at 120°C for 3 min to ensure adequate adhesion and functionality. Non-sensing areas were encapsulated using a waterproof, skin-adhering, double-sided acrylic medical adhesive tape (3M-2477P, USA). The patterned structure of the sensor was realized by employing laser engraving technology using the Universal VLS2.30 Laser System (USA). The operation parameters were set as follows: a power level of 80%, a speed of 30%, and the laser engraving process was repeated three times to enable subsequent Functional modification.

GOx membrane solution preparation. GOx lyophilized powder (10 KU, Sigma) was dissolved in 555 μL of PBS to prepare a GOx solution with a concentration of approximately 100 mg mL⁻¹. This solution was then diluted to obtain a final concentration of 10 mg mL⁻¹. To prepare the chitosan precursor solution for embedding GOx, 10 mL of DI was added to 0.1 g of chitosan (1 wt%), followed by the addition of 30 μL of glacial acetic acid (0.3 v/v%) and 1 mL of glycerol (10 v/v%). The mixture was heated at 80°C and stirred for 2 h. Subsequently, 0.1 mL of PtNPs aqueous solution was incorporated to complete the chitosan precursor solution. The prepared GOx solution (10 mg mL⁻¹) was mixed with the chitosan precursor solution in a 2:1 volume ratio to yield the GOx membrane solution, which was stored at 4°C for future utilization.

Na+ and H+ selective membrane solution preparation. The Na+ selective membrane solution was prepared by dissolving 2 mg of Na+ carrier X (97%), 1 mg of Na-TFPB (97%), and 131 mg of DOS in 708 μL of CYC, ensuring complete dissolution. Afterward, 66 mg of PVC was added and ultrasonicated for 30 min to achieve a homogeneous solution, forming the Na+ selective membrane solution. For the pH sensor selective membrane, the Na+ carrier was replaced with Tridodecylamine and the Na-TFPB was replaced with KTClPB, while the other steps and ratios remained unchanged.

Preparation of reference membrane solution for Na+ and pH sensors. The reference membrane solution for both the Na+ and pH sensors was prepared by dissolving 0.2 mg of MWCNTs, 2 mg of pluronic F-127, 79.1 mg of PVB, and 50 mg of NaCl in 1 mL of anhydrous methanol. MWCNTs, F127, and NaCl were first dissolved in methanol to form a uniform solution, with NaCl reaching saturation. Following this, PVB was added to the mixture and ultrasonicated for 20 min to ensure complete dissolution of PVB. The final reference membrane solution was stored at 4°C to ensure adequate stabilization prior to use.

Modification process of the functional layer. A 1 μL aliquot of the GOx membrane solution was individually applied to the WE on both Part A and Part B of the glucose sensors. The solution was left to naturally solidify at room temperature for 30 min. Subsequently, a 1 μL aliquot of PEDOT:PSS aqueous solution was separately applied to the WE on both Part A and Part B of the Na+ and pH sensors. The solution was allowed to solidify at room temperature for 1 h, thereby forming a conductive polymer layer. A 1 μL aliquot of Na⁺ selective membrane solution was subsequently applied to the WE of the Na⁺ sensors, while a 1 μL aliquot of H⁺ selective membrane solution was applied to the WE of the pH sensors, both on Part A and Part B of the device. The electrodes were then allowed to dry at room temperature for 4 h. Finally, a 30 μL aliquot of Nafion (5 wt%, Sigma) was diluted with alcohol to a 1 wt% concentration and applied to both part A and part B of the prepared device. After curing at room temperature for 1 h, the device was fully prepared and stored at 4°C for preservation. Structurally, each part comprises a semi-circular Ag/AgCl electrode for RI extraction, a glucose sensor with a PBC WE and a quarter-circular Ag/AgCl RE, a circular PBC WE for the Na⁺ sensor, a circular PBC WE for the pH sensor, and a shared quarter-circular Ag/AgCl RE for both the Na⁺ and pH sensors. The detailed structure of the electrode is shown in Supplementary Fig. 3.

Activation process of the device. To ensure stable potential measurements of the ion sensor and eliminate errors caused by baseline drift, the Na+ and pH sensors were immersed for 2 h in a Tris-HCl buffer solution at pH 7 containing 0.1 mM NaCl for activation prior to in vivo testing.

Characterization of the wearable device
Electrochemical characterization of the glucose sensor. All electrochemical tests were conducted using the CHI660E electrochemical workstation. Cyclic voltammetry (CV) scans were performed within the voltage range of −0.6 V to +0.6 V in 1× PBS (0.01 M, pH=7.2–7.4), with a scan rate of 0.1 V s⁻¹. H₂O₂, an intermediate product of glucose oxidation, was first measured to assess the performance of PBC WE. Typically, the WE requires CV scans to remove surface impurities and identify the optimal bias voltage, which was found to be −0.1 V (Supplementary Fig. 8a, b). CV scans at different scan rates showed a linear relationship between peak current and the square root of the scan rate, confirming that the H₂O₂ reaction is diffusion-controlled (Supplementary Fig. 8c, d). For H₂O₂ and glucose testing, the i-t curves were carried out for 60 s, with the steady-state current recorded at the 60-s mark. Stepwise testing revealed a good linear response between sensor current and H₂O₂ concentration (Supplementary Fig. 8e), with linear sensitivity of 441 nA mM⁻¹ observed ranging the 0–5 mM (Supplementary Fig. 8f). The glucose sensor testing follows a similar procedure as H₂O₂, with an initial CV scan determining the optimal bias voltage of −0.1 V (Supplementary Fig. 9a, b). CV results at different scan rates also showed a linear correlation between peak current and the square root of the scan rate, confirming that the glucose reaction is diffusion-controlled (Supplementary Fig. 9c). Therefore, the current can be linearly fitted to concentration using the Cottrell equation. The sensitivity of glucose within the range of 3–30 mM is 28.22 nA mM⁻¹ (Supplementary Fig. 9d). The selectivity of the glucose sensor was assessed by immersing the glucose sensor into a 10 mL Tris-HCl solution at pH 7.4 (Fig. 2g, Supplementary Fig. 9e, f, g). Starting at 100 s, glucose (0.5/5 mM), NaCl (14/140 mM), KCl (0.4/4 mM), CaCl₂ (0.25/2.5 mM), uric acid (0.01/0.1 mM), dopamine (0.01/0.1 mM), ascorbic acid (0.01/0.1 mM), fructose (0.1/1 mM) and glucose (1/10 mM) were sequentially added every 60 s. The solution was stirred continuously for 1 min to facilitate stability before each detection.

The enzyme activity of GOx was evaluated by measuring the concentration of H₂O₂ generated by GOx-catalyzed glucose at different pH values. The reaction mixture consisted of glucose oxidase (10 mg mL⁻¹) and glucose solutions (10 mM) in a 1:1 v/v ratio (prepared in Tris-HCl buffer), incubated for 1 h before measurement The UV ABS test was based on the Trinder’s reaction for H₂O₂, where a standard curve for H₂O₂ concentration was obtained by measuring UV ABS at 500 nm (Supplementary Fig. 10a). A solution of 0.2 mM phenol and 0.2 mM 4-AAP in PBS buffer, along with 0.25 mg mL⁻¹ HRP solution, were prepared for testing. Different concentrations of H₂O₂ (10 mM, 20 mM, 30 mM, 40 mM, 50 mM) were added to the reaction mixture, followed by 500 μL of HRP solution. The mixture was shaken for 5 min at 200 rpm using a shaking incubator (SK-L180-S, DLAB). The reaction solution was transferred to a microplate for analysis using an enzyme-linked immunosorbent assay reader (Feyond-300, Hangzhou Aosheng Instrument Co., Ltd). The UV ABS at 500 nm for different H₂O₂ concentrations was plotted to obtain the standard curve (Supplementary Fig. 10b, c). The fitting results of the ABS of H₂O₂ and pH value also show a quadratic function curve, which is similar to the relationship between glucose sensitivity and pH, demonstrating its effectiveness (Supplementary Fig. 10d).

Electrochemical characterization of the Na+ and pH sensors. The Na⁺ and pH sensors, based on open circuit potential (OCP), were measured over a duration of 60 s, with the steady-state voltage taken at the 60-s mark. The selectivity of the Na⁺ sensor was evaluated by immersing the Na⁺ sensor into a 10 mL Tris-HCl solution at pH 7.4 (Fig. 2o, Supplementary Fig. 12c). Starting at 100 s, various substances were sequentially added every 60 s, including NaCl (50 mM), KCl (4 mM), CaCl₂ (2.5 mM), uric acid (0.1 mM), dopamine (0.1 mM), ascorbic acid (0.1 mM), glucose (5 mM), fructose (1 mM) and NaCl (200 mM) were sequentially added every 60 s. The selectivity of the pH sensor was assessed by immersing the pH sensor into a 10 mL Tris-HCl solution at pH 6.6 (Fig. 2s, Supplementary Fig. 13d). Starting at 100 s, NaCl (140 mM), KCl (4 mM), CaCl₂ (2.5 mM), uric acid (0.1 mM), dopamine (0.1 mM), ascorbic acid (0.1 mM), glucose (5 mM), fructose (1 mM) and Tris-HCl buffer (pH 7.6) were sequentially added every 60 s. Prior to each test, the solution was continuously for 1 min to facilitate the stability of all substances added to the reaction system.

Mechanical evaluation of the wearable device. The fabricated wearable device was evaluated using an environment-controlled flexible electronic testing system (FE-SDT, SINOAGG, Beijing). This testing included measurements of the resistance of the RI electrodes before and after bending, as well as electrochemical performance tests of the sensors under different bending conditions. The bending test consisted of two phases: the bending process and the recovery process, which together formed a complete cycle. The bending limit was set to 50% of the sensor’s flexibility (approximately 90°), with a waiting time of 10 s once the limit was reached. The resistance testing during the bending process was conducted over two cycles (Supplementary Fig. 14). To evaluate the stability of electrochemical performance, the device was tested for the steady-state current of glucose, the steady-state voltage of Na⁺, and the steady-state voltage of pH after undergoing different numbers of bending cycles (0, 10, 30, 50, and 100 bends) (Supplementary Fig. 15). These tests allowed for a comprehensive assessment of the sensor’s mechanical performance and its ability to maintain stability under wearable applications.

Experimental validation of effect of skin surface pH on RI
Given the challenges presented by uncontrollable factors, including human skin temperature, skin impedance, and ISF flow rate, experiments were employed to exclusively evaluate the impact of skin surface pH on RI. The ISF extraction and testing system was constructed using Strat-M® transdermal membranes (47 mm in diameter, 300 μm in thickness) to simulate human skin. A Franz diffusion cell setup was employed, with the donor and receiver compartments designed to mimic the ISF extraction process. The receiver chamber contained 3 mL of Tris-HCl buffer solution at varying pH levels (ranging from pH 5 to 10). Commercial Ag/AgCl electrodes were used as the RI electrodes to minimize ISF extraction fluctuations related to electrode degradation during RI. A constant current source was applied, delivering a 0.5 mA current for 15 min to facilitate the ISF extraction (Supplementary Fig. 16). Firstly, we assessed the effect of varying skin surface pH on Na+ extraction flux. The donor chamber in the Franz cell contained 15 mL of Tris-HCl buffer solution with 140 mM NaCl and either 15 mM or 30 mM glucose at pH 7.4. Then we investigated the influence of skin surface pH on glucose extraction flux. This experimental setup provided a controlled environment to evaluate the pH-dependent changes in RI.

Preparation of the in-vitro ISF extraction and testing system
Due to the interference of uncontrollable factors such as skin temperature, skin impedance, and interstitial fluid flow rate, the effect of pH as a single variable on glucose extraction flux was investigated through in vitro simulation experiments. Figure 4a, Supplementary Fig. 24 illustrates the setup of the system, which includes a constant current source to supply the RI extraction current, a TPY-2 intelligent transdermal diffusion instrument to maintain a constant temperature of 37 ± 0.1°C, and a magnetic stirrer to simulate ISF flow. The Strat-M® transdermal membrane (47 mm diameter, 300 μm thickness) mimics human skin, while two electrochemical workstations (CHI660E) are used to measure glucose concentration, Na+ concentration, and pH on the membrane surface. A switch controls the ISF extraction and detection process automatically, and a computer is used for data acquisition and processing. RI is set as 150 μA cm⁻² for 3 min for one extraction cycle, which is selected for subsequent glucose extraction testing.

To reduce fluctuations in ISF extraction induced by electrode degradation, alternating current polarity was utilized between consecutive extractions (Supplementary Figs. 25, 26). This approach helped maintain the performance of RI electrodes and ensured more stable extraction conditions. To further minimize potential errors from repeated ISF extractions, glucose and pH measurements from five consecutive cycles were selected for validation, providing a more reliable dataset for assessing extraction consistency and the accuracy of glucose predictions.

OGTT Procedure
The OGTT data collection process is illustrated in Fig. 5a. After attaching the prepared device to the skin, a 10-min warm-up was initially conducted to ensure tight adhesion of the device to the skin and stabilize the RI current. Following this, glucose, Na⁺, and pH are sequentially measured in both parts of the device; subsequent measurements follow the same protocol. After the initial measurements, ISF is extracted for 3 min, with the extraction direction set from Part B to Part A. Measurements are then repeated, with the first set of valid data labeled as Datapoint 1. The polarity is then switched to obtain subsequent data points, such as Datapoint 2, Datapoint 3, and so on. The time lag between blood glucose and ISF glucose is established as 10 min, consistent with commonly reported physiological lags62,63 (Supplementary Table 8).

After an 8–14 h fasting period, 6 healthy volunteers and 15 diabetic patients participated in a 3-h OGTT. Participants consumed a 250 mL bottle of glucose solution containing 75 g glucose over a 5-min period. Blood glucose levels were measured using a commercial glucometer (ACCU-CHEK, Switzerland) for healthy volunteers and venous blood sampling points for diabetic patients as the reference. The experimental setup included the fabricated epidermal device, a constant current source, electrochemical workstations, a switch circuit, and a laptop. A wearable device was applied to the participants’ forearm for continuous glucose, Na+, and pH detection, with RI performed at a current density of 150 μA cm⁻² for 3 min for one extraction cycle. After each extraction cycle, glucose, Na+, and pH levels were measured for one minute. The polarity of the RI current was alternated between successive cycles to reduce the consumption of RI electrodes. The performance of the proposed skin surface pH calibration method was evaluated by comparing the predicted blood glucose levels with reference measurements using Clarke’s error grid analysis.

Peer review rounds
[{"round": 1, "reviewer_comments": {"reviewer_1": "The authors present an intriguing wearable, non-invasive device designed for clinical-level detection of interstitial fluid biomarkers by leveraging pH-calibrated reverse iontophoresis. This approach is innovative, particularly as it addresses the longstanding issue of extraction instability due to variations in skin surface pH. While the work undoubtedly pushes the frontier in biosensing technology, several critical issues must be addressed before the manuscript can be considered for publication in Nature Communications.\n1. Zeta Potential Measurements: In supplementary Table 1, the standard deviations for the zeta potential (ZP) measurements of keratin across different pH levels are alarmingly high relative to the mean values. This level of uncertainty impedes a confident assignment of positive or negative charges, especially at lower pH values, and undermines the assertion presented in Fig. 1d regarding a pronounced negative correlation between ZP and pH. A re-evaluation of these uncertainties, along with a more robust statistical treatment or additional replicates where possible, would strengthen this aspect of the study.\n2. Glucose Concentration Discrepancy (Lines 105–107 vs. Fig. 1f): The manuscript states that glucose concentrations at the anode are higher at low pH compared to high pH (Lines 105–107), yet Figure 1f does not appear to corroborate this claim. The authors must reconcile these differences by either adjusting the text or providing a thorough explanation that accounts for the observed data trends. It is crucial to clarify whether this discrepancy stems from experimental variability, calibration issues, or another underlying phenomenon.\n3. Rationale for Glucose Concentration in Figure 2g: The selection of a 5 mM glucose concentration for the experiments depicted in Figure 2g is not adequately justified, particularly since Figure 2f employs a much lower concentration range (0– 300 µM). The authors should clarify the rationale behind these differing concentration regimes and explain their relevance to the overall experimental design and clinical applicability.\n4. Absence of Error Bars in Key Figures: Figures 2h, 2p, and 2t currently lack error bars, which are indispensable for assessing the reliability and reproducibility of the presented data. Including error bars will provide insights into the variability and statistical confidence of the measurements, thereby bolstering the study's credibility.\n5. Inconsistency in Sodium Concentration Ranges: The experimental design utilizes sodium concentrations in the range of 10–200 mM, yet Figure 3d shows extracted sodium concentrations in the micromolar range. This apparent inconsistency warrants a detailed explanation—potentially involving differences in extraction efficiency, calibration protocols, or inherent sample dilution effects—to ensure that readers can clearly follow the experimental logic.\n6. Anomalous Electrode Current Response: In Figure 8b, the Prussian Blue-modified electrode exhibits a simultaneous increase in both anodic and cathodic currents in the presence of hydrogen peroxide. This dual response is unexpected and remains unexplained by the current discussion. A more detailed analysis or additional experiments to elucidate this phenomenon would greatly enhance the manuscript’s mechanistic insights.\n7. Statistical Validity of pH-Dependent Results: The overlapping standard deviation ranges depicted in Figures 3f–h call into question the conclusion that the results at pH 7–10 are distinct. The authors should either perform a more rigorous statistical analysis to substantiate this claim or reconsider the interpretation of these data points.\n8. Correction of Supplementary Figure 37 Axis Labels: The glucose concentration axis labels in Supplementary Figure 37 appear to be misrepresented. Correcting these unit labels is essential for the clarity and accuracy of the supplementary data.\n9. Clarification of Extraction Time for Figure 4g: Figure 4g presents glucose concentration values derived from an extraction process, but the specific time interval used for this extraction is not indicated. Including this detail is necessary for reproducibility and for a proper understanding of the extraction kinetics.\n10. Sodium Flux versus pH Inconsistency: Lastly, the manuscript (Lines 335–336) asserts that sodium flux decreases with increasing pH; however, the trend in Figure 3d does not consistently support this claim. The authors are encouraged to re- examine the data trend, provide additional analysis, or revise their conclusion accordingly", "reviewer_2": "Comments to Authors on the paper\nAbstract\nLines 12-15. These 2 sentences are quite difficult to understand. Is 'instability' defined/explained anywhere in the paper? Is H+ flux ever measured, and what is the skin self-recovering from, exactly?\nLines 15-17. There is a sensing of pH by the device created, but is it really measuring skin 'surface' pH, or rather the pH of whatever medium is present adjacent to the pH sensor?\nIntroduction\nLines 52-54. At pH 7, the physiological concentration of Na+ is about 133 mM, while that of H+ is 10-4 mM. Even though H+ is a very mobile ion, it seems to me that the more than 106-fold difference in concentration between Na+ and H+ means that the former will be do the \"heavy-lifting\" in terms of carrying current across the skin during RI. This should be the case even if the pH falls to 6 or 5 (and we know that by pH 4 the net negative charge on the skin has been neutralised (or even reversed) and the EOF switches to the opposite direction).\nLines 56-57. Reference 52 is incomplete and is a paper about wound healing. It does not appear to say very much of relevance to the self-recovery mentioned in the abstract associated with application of RI.\nResults & Discussion\nLines 79-82. RI applies a fixed current, so the charge flowing between the electrodes must be matched by the charge carried by ions in the 'in vivo' part of the circuit. This will involve all available charge carriers with the distribution between them being determined by their mobilities and concentrations.\nLines 85-87. I'm not sure that anyone has measured this proton flux or shown that it changes skin surface pH. Incidentally, what is meant by skin surface pH in this context? Without introducing some water on the surface, how is a measurement of pH performed? Without water, what does pH \"of the surface\" even mean?\nLines 87-88. Another statement for which there is no experimental evidence as far as I know. What 'self-recovery' mechanism is invoked here, exactly?\nLines 89-91. It is true that experiments ex vivo where it has been possible to change the pH of solutions on either side of the membrane have shown that lowering the pH from 7.4 to that equal to the isoelectric point of the membrane, and then further to below this value, have first reduced electroosmotic flow and the reversed it. But given the ISF pH of ~7.4, the available [H+] is really small and it's hard to envisage that the flux will be able to substantially alter the surface pH (whatever that means - see earlier comment).\nLines 93-94. Characterization of this \"hydrolipid film\" (HF) would be helpful... What is its nature when a conductive hydrogel is placed between the RI electrodes and the skin? How is its pH chosen and how - if at all - does this pH change with time with or without application of a RI current? Are the changes the same or different under anode and cathode? Do they reverse when the polarity of the electrodes is switched between extractions to prevent depletion/build-up of AgCl?\nLines 105-107. This is all \"old news\", surely? See comment in S.I.\nLines 118-120. More 'old news' here - the advantages of using this approach were identified by Cygnus more than 30 years ago!\nLines 122-127. So, do I understand correctly that the pH sensor is detecting the value in the gel near the electrode surface? Is this not near where glucose is converted to gluconic acid (releasing H+ locally)?\nFig. 1c. Presumably, the regions labelled as high and low pH cannot be ISF, as the pH there must be fixed. The figure needs to be more clearly drawn.\nFig. 2k and 2l. Difficult to assess the relevance of some of this characterisation work performed in the pH range of 5-10. Yes, it's indeed possible that glucose detection is influenced over this range, but surely this does not occur in vivo? So, the more important question is 'over what range does [glucose] vary in a typical RI extraction-detection use of the device?'\nLine 177. A more objective assessment of performance would be better!\nLines 178-180. But are you really measuring skin \"surface\" changes in pH here, as opposed to changes in the bulk pH of the hydrogel?\nLines 204-206. This is not a widely accepted model for mammalian skin so data obtained using it must be treated with caution and scepticism. The experiment has been done before, in essence, using mammalian skin so why repeat it here using an artificial membrane?\nFig. 3. There is really nothing new here except that EO flow can occur across a charged membrane when an electric field is applied. More flow will occur in the direction of counterion movement as the charge on the membrane increases. The flux of a water-soluble, unionised species present in the solvent will be proportional to its concentration, all other factors being the same.\nLines 226-228. Again, no surprise here. Na+ flux is driven principally by electromigration, far more than by EO. It's the major positive ion current/charge carrier in the system, so it's obvious that its flux will be insignificantly affected by pH.\nLines 244-246. This is well-known as has been reported multiple times before. However, as we do not know the extent to which the charge on a (real) skin membrane in vivo changes during the use of an RI device, there is no basis for concluding that a pH-based correction is needed to ensure that accurate deductions about the glucose levels in the ISF are being made.\nLines 255-257. How does it do that? As Na+ is the principal charge carrier and RI is done under constant current conditions, one expects (and it's true) that the sodium flux per unit current per unit time is actually quite constant, and this has been shown in vitro and in vivo in man. So how can JNa+ be \"at the base of subdermal glucose prediction\"?\nLine 261. Sub-heading does not make complete sense.\nLines 272-276. The system is symmetrical so why are the i-t responses different in Figs. 4b and 4e, and why are the pH fluctuations different (Figs. 4c and 4f), both in absolute values and the extent of change of pH - less than 0.1 unit in Part A, more than 0.2 unit in Part B? These don't seem to be massive changes in pH in any case and the impact as reported in Figs. 4d, 4g, 4h and 4i confirm that this is the case. The \"explanation\" offered in this paragraph is not convincing, to say the least (mostly, because it does not explain anything!).\nLines 295-297. Exactly how were the data from Parts A and B combined?\nLines 298-300. The data in Fig. 4h show a decent correlation between extracted glucose and subdermal glucose both with and without pH 'calibration'. It would be good to see confidence intervals on the slopes and to test whether they are significantly different. In a practical sense, it seems unlikely that the differences in slopes would dramatically impact on the clinical value of the data when one adds in the pH correction.\nFig. 5. [1] Are the units of extracted glucose correct in 5b and 5c? Should this be micromolar? [2] In the same panels, pH is increasing from about 6 to 10, the latter likely to cause some skin irritation, no? What is causing the increase in alkalinity? Why does the pH change in a sawtooth manner for the normal volunteer but not for the person living with T2D? The fluctuation in predicted glucose levels is more sawtooth for the diabetic patient (whose pH is not varying like that) than that for the normal patient (whose pH is up-and-down throughout) How (in)consistent was that behaviour?\nMethods\nLine 512. I don't believe that this sensitivity can be measured to an accuracy of 8 picoAmps (the 3rd decimal place here)!\nComments to the authors on the Supplementary Information\nLines 30-31. How do you know this as a fact? Based on what evidence, exactly?\nLines 38-39. Which is which? Whatever is the phase on the skin surface, it is hard to describe it as liquid...\nLines 59-60. Same statement as above but does not obviously flow from what is written here which appears to be standard, accepted dogma.\nGeneral remark on Suppl. Note 2. OK, not really anything terribly new or controversial here, it seems. The bottom line appears to be that EO flow depends on the charge on the membrane and that the latter depends on the pH of the 'liquid' phases on either side. The premise, therefore, appears to be that, if one can control the pH of the 'liquid' phase on the skin surface, one can be sure that EO flow will be fixed during RI and that the glucose detected in the 'receiving' gel will therefore be proportional to that in the ISF.\nLine 121. The data from which this empirical equation is derived was acquired over a very wide - and non-physiological - range of pH, so its relevance to the 'real world' is not established. What is the form of the relationship when the range of pH values is limited to something sensible/realistic?\nLine 124. And the electroosmotic component? Missing?\nLines 142-146. I’m not sure that I would write the equation like this. The units of uglu are velocity. The product {A x uglu} is therefore equivalent to the EO volume flow rate induced by the imposition of the electric field across the charged skin membrane. It seems to me that {A x uglu} must have the same value for all small, neutral molecules present in the system - how can it be specific to glucose?\nGeneral remark on Suppl. Note 3. So, at the end of the day, the approach used requires measurements in the sensor gel of glucose, Na+ and pH. Ref 22 in the SI used Na+ as an 'internal standard', effectively to correct for how the charge on the skin may vary from person to person and hence how EO flow would be affected as a result. Did the authors check to see if the approach outlined here, without considering the impact of small changes in pH, would work just as well? If the answer is no, then they probably should...\nS.I Fig. 10. On many graphs, using stronger colours for the datapoints and the lines connecting them would be helpful.\nS.I Fig. 10. Why are dependencies so different as a function of concentration at different pH values? I'm not sure I'd want a sensor on my arm that led to a pH of 9 or 10...\nS.I Fig. 13b. Impossible to see data on this panel.\nS.I Fig. 24. Some explanation of what one is supposed to be seeing here would be helpful... significance of colouring?\nS.I. Fig. 26. Impossible to see bars in the \"Data point 1\" and \"Data point 3\" boxes.\nS.I. Fig. 27. Minimal variation in pH here…\nS.I. Fig. 29. Again difficult to understand the manner in which pH varies over extraction times.\nSame comment on S.I. Figs, 30, 31, 32, 34, 35.\nS.I. Fig. 32. But what happened in panel g?\nS.I. Fig. 37. Makes more sense to use the same y-axis scale for these two graphs. The correlations between glucose extracted and subdermal glucose are pretty good with or without pH calibration, it seems to me. I'd like to see confidence intervals on the slopes and whether the with and without pH calibration values are significantly different or not.\nS.I. Fig. 38. Would be helpful if the data for the healthy volunteer were shown using the same y-axes for extracted glucose, Na+ and pH. Similarly for the diabetic subject. Please use bolder colours to facilitate visualisation. Why are the sawtooth variations prominent in some curves but not others?\nS.I. Fig. 42. What skin pH values were recorded for this individual?\nSuppl. Table 2. References for the source of the values in this table should be provided.\nSuppl. Table 5. Why do the slopes sometimes agree in Parts A and B, and sometimes they do not? Why is the % increment with pH calibration not consistent?", "reviewer_3": "This manuscript is an excellent addition to the literature. The authors have identified a pain point in the development of on-body iontophoretic devices and have provided an innovative solution for it. Polarity switching to minimize current instability is a notable feature of the proposed device design to mitigate skin irritation. Moreover, the proof of concept of pH fluctuation as well as the establishment of a solution via calibration is commendable.\nWhile I suggest that the manuscript be accepted for publication, certain aspects can be included in the discussion section, which can help delineate the relevance and practicality of the designed device:\n1. Is there any lag or delay in the physiological status of markers in blood versus their effect in ISF? Can you please provide some references to this aspect?\n2. In relation to my first question, is this lag (if any) different for different markers, hence making such a device more suitable for certain biomarkers (such as glucose) more than others? Essentially, can this same device be used for any biomarker, or will the calibration or other aspects be specific to a single biomarker?\n3. A major drawback of iontophoretic devices is the risk of local skin burns due to electric current itself, or trips in the device due to non-ideal contact with high conductance fluids for example: accidental water contact. Can you please discuss or suggest any potential safety features that can be embedded in the on-body device, which will break the circuit and arrest the device before a current overload takes place?\n4. How did the authors arrive at the glucose range of 6 to 30 mM for the in vitro testing? Can you please provide a correlation of the tested ranges with blood glucose levels in humans? In addition, can you provide some information about the quantitation limits of the developed device?\n5. Did the authors make several of these devices and if yes, what was the variability observed across different devices in terms of calibration and sensitivity to skin pH changes? If some discussion on this can be added, it will improve the robustness of the proposed model/device."}, "author_response": ""}, {"round": 2, "reviewer_comments": {"reviewer_1": "General Comments: The authors present an intriguing wearable, non-invasive device designed for clinical-level detection of interstitial fluid biomarkers by leveraging pH- calibrated reverse iontophoresis. This approach is innovative, particularly as it addresses the longstanding issue of extraction instability due to variations in skin surface pH. While the work undoubtedly pushes the frontier in biosensing technology, several critical issues must be addressed before the manuscript can be considered for publication in Nature Communications.", "reviewer_2": "Authors have made detailed responses to my comments but have not, in my opinion, provided compelling evidence that the addition of pH monitoring represents a true step-change in the use of reverse iontophoresis for glucose monitoring. Specific remarks follow:\n1. Authors have accepted that they nmeasure a bulk pH in the fluid on the skin surface; they do not measure the flux of H+ ions. Relatedly, the authors now state that the pH measured is the same as that in the \"epidermal electroosmotic microchannel\" but how do they know this, given that they cannot actually measure what is going on in that space?\n2. Authors also accept that they do not know the quantitative contribution that H+ makes to carrying the applied current across the skin. They acknowledge that it must small (given the very low concentration of H+ ions available to do so) and this - by extrapolation - challeges their premise that correcting for the amount of H+ around is needed.\n3. Authors' definition of 'skin recovery' remains vague. Imposing an electric field across a membrane drives in ions and causes membrane impedance to decrease; when the field is removed, ions will diffuse out down the created concentration gradient and impedance returns to its original value... is this recovery or just physics?\n4. Authors' responses refer repeatedly to H+ changes being due to \"extracted H+\" (which cannot be very much from a physiological fluid like the ISF that has a pH close to neutral) and \"released H+\" from 'biochemical reactions' at the epidermal interface. What are these reactions and how much H+ do they release? Can they really be sufficient to cause the local pH to increase to 10?\n5.The response about the hydrolipid film is obscure and quickly digresses to discuss the mysterious \"epidermal electroosmotic microchannel\"...\n6. Authors state without any 'back-up' or verification that any contribution to H+ presence resulting from the enymatic reaction between glucose and glucose oxidase is small compared to H+ extraction in the EOF. Is the Gox reaction the \"biochemical reactions\" referred to in previous and other answers, or are there others (and, if yes, then what are they)? Either way, the authors need to be up-front here.\n7. It's quite difficult to come up with an explanation for pH increasing to 10 during a short period of reverse iontophoresis without some undesirable electrochemistry occurring that results in water electrolysis...\n8. This reviewer remains unconvinced about the relevance of the artificial membrane experiments.\n9. It remains unclearas to whether the addition of a pH correction, in addition to that achievable with the Na+ correction, provides any real-world added-value in terms of accuracy/prediction given that - except perhaps in very extreme cases as the Authors acknowledge - the pH variation possible is really quite small and unlikely to affect EOF. New Figures R3/R4 support this point of view.\n10. This reviewer's concerns about Figure 5 are not assuaged by the Authors' reply which is highly speculative and not supportd by any hard facts.\n11. The term \"reverse ions\" is unusual. I think the Authors mean 'counterions', i.e., ions with a charge opposite to that of the skin.\n12. I think that maintaining a pH of 10 on the skin for very long will typically cause visible irritation.", "reviewer_3": "Reviewer #1:\nGeneral Comments: The authors present an intriguing wearable, non-invasive device\ndesigned for clinical-level detection of interstitial fluid biomarkers by leveraging pH-\ncalibrated reverse iontophoresis. This approach is innovative, particularly as it\naddresses the longstanding issue of extraction instability due to variations in skin\nsurface pH. While the work undoubtedly pushes the frontier in biosensing technology,\nseveral critical issues must be addressed before the manuscript can be considered for\npublication in Nature Communications.\nReviewer #2:"}, "author_response": "Reviewer #2:\nResponse: Thank you for your detailed and rigorous comments of our work. We have carefully reviewed each of your concerns and have made substantial revisions to the manuscript accordingly. Below, we provide detailed point-by-point responses to your comments.\nComments to Authors on the paper\nAbstract\n1. Lines 12-15. These 2 sentences are quite difficult to understand. Is 'instability' defined/explained anywhere in the paper? Is H+ flux ever measured, and what is the skin self-recovering from, exactly?\nResponse: Thank you for your valuable comments. We acknowledge that our original phrasing lacked clarity and failed to fully define key terms.\nIn this work, “instability” refers to the fluctuations in the volume and composition of electroosmotic flow (EOF) of interstitial fluid (ISF) during reverse iontophoresis (RI). The instability leads to a dynamic changed relationship between the extracted analyte concentrations and those in subcutaneous ISF, for example, glucose.\nFor the H⁺ flux, we apologize for the inaccurate description. Exactly, we did not measure the H⁺ flux directly. We measured the pH of the entire microfluid at the skin surface and within the epidermal electroosmotic microchannel throughout the RI extraction process (Fig. R1). The pH changes are primarily due to the movement of H⁺, which have high mobility and is one of the earliest migrating species during the RI process 1. And the pH changes at the skin surface and in the epidermal electroosmotic microchannel affect the skin’s zeta potential (ZP) and consequently influence the flux of the entire EOF.\nThe skin exhibits a natural resistance to RI-induced ISF extraction, which we refer to as “skin self-recovery”. This encompasses the skin’s ability to buffer pH fluctuations and regulate fluid loss through its structural and physiological barriers, such as the stratum corneum and sweat glands 2-4. These self-regulating responses can suppress the accumulation of extracted ISF and mitigate the extent of local acid-base imbalance caused by RI.\nReferences:\nFig. R1 Schematic of RI-induced electroosmosis flow and extracted fluid at the skin surface and in the epidermal electroosmotic microchannel.\nChanges: We have replaced “H⁺ flux” with “H⁺ movement” throughout the text. Additionally, we have further updated on the relevant description of “instability” and “skin’s self-recovery”. Fig. R1 has been added to Supplementary Fig. 5a. For your convenience, we included the revised text here:\n“The interaction leads to the fluctuations of skin surface pH, which changes the zeta potential (ZP) of the EDL 55, 56, thereby causing instability in the electroosmotic flow of ISF accordingly.” (Lines 61-63 in the revised Manuscript)\n“Specifically, the extracted H⁺, along with the release of additional H⁺ from biochemical reactions, collectively leads to changes in the pH of the RI-induced fluid accumulated at the skin surface and within the epidermal electroosmotic microchannel (referred to as the skin surface pH in this study).” (Lines 56-59 in the revised Manuscript)\n“The skin’s self-recovery mechanisms, including natural barriers such as sweat glands and the stratum corneum that maintain homeostasis and prevent excessive polarization 52, 53, 54, will hinder the change.” (Lines 59-61 in the revised Manuscript)\n2. Lines 15-17. There is a sensing of pH by the device created, but is it really measuring skin 'surface' pH, or rather the pH of whatever medium is present adjacent to the pH sensor?\nResponse: Thank you again for your comment regarding the pH sensing in our system. We acknowledge that our initial phrasing of skin surface pH did not convey the intended meaning.\nAs stated in the response to Comment 1, the term “skin surface pH” in this study refers to the pH of the entire microfluid at the skin surface and within the epidermal electroosmotic microchannel throughout the RI extraction process (Fig. R1). We refer to this measured value as “skin surface pH” in the manuscript. In our opinion, the extracted microfluid at the skin surface is the same as that in the epidermal electroosmotic microchannel, therefore, we think that the measured pH also reflects the conditions within the epidermal electroosmotic microchannel. This pH value influences the skin’s ZP, thereby affecting the electroosmotic flux of glucose.\nChanges: We have provided explanations and definitions for the first appearance of “skin surface pH” and have made corresponding revisions in other relevant sections. For your convenience, we added the text here:\n“Specifically, the extracted H⁺, along with the release of additional protons from biochemical reactions at the epidermal interface, collectively leads to changes in the pH of the RI-induced fluid accumulated at the skin surface and within the epidermal electroosmotic microchannel (referred to as the skin surface pH in this study).” (Lines 56-59 in the revised Manuscript)\nIntroduction\n3. Lines 52-54. At pH 7, the physiological concentration of Na+ is about 133 mM, while that of H+ is 10-4 mM. Even though H+ is a very mobile ion, it seems to me that the more than 106-fold difference in concentration between Na+ and H+ means that the former will be do the \"heavy-lifting\" in terms of carrying current across the skin during RI. This should be the case even if the pH falls to 6 or 5 (and we know that by pH 4 the net negative charge on the skin has been neutralised (or even reversed) and the EOF switches to the opposite direction).\nResponse: We fully agree with the reviewer’s quantitative assessments of ionic flow in the RI process. We acknowledge that our description of EOF was not accurate. We have revised the relevant sentences in the “Introduction” section. For your convenience, we added the text here:\n“This process is driven primarily by the electromigration of cations in the electrical double layer (EDL), thereby making the electroosmotic flow of small neutral biomarkers, for example, glucose.” (Lines 52-54 in the revised Manuscript)\n4. Lines 56-57. Reference 52 is incomplete and is a paper about wound healing. It does not appear to say very much of relevance to the self-recovery mentioned in the abstract associated with application of RI.\nResponse: We appreciate the reviewer’s comments regarding Reference 52 and the relevant description of skin self-recovery. In the revised manuscript, we have replaced Reference 52 with more appropriate three literatures that directly addresses the physiological recovery mechanisms of normal skin under iontophoretic or electric stimulation. For instance, Catherine Curdy et al. 2 and Prashanth Manda et al. 3 reported that after cessation of iontophoretic current, skin impedance gradually decreases and stabilizes to prevent further polarization. This recovery is likely due to passive ion diffusion and the re-establishment of normal ion gradients, including H+, indicating a form of physiological regulation and self-recovery. Specifically, the skin’s natural barriers, such as sweat glands and the stratum corneum, play a stabilizing role in maintaining ion homeostasis and preventing excessive polarization 4.\nReferences:\nChanges: The relevant sentences were revised for clarity in the manuscript and more accurate references were added to support the mechanisms of the skin’s self-recovery ability. For your convenience, we added the text here:\n“The skin’s self-recovery mechanisms, including natural barriers such as sweat glands and the stratum corneum that maintain homeostasis and prevent excessive polarization 52, 53, 54, will hinder the change.” (Lines 59-61 in the revised Manuscript)\nResults & Discussion\n5. Lines 79-82. RI applies a fixed current, so the charge flowing between the electrodes must be matched by the charge carried by ions in the 'in vivo' part of the circuit. This will involve all available charge carriers with the distribution between them being determined by their mobilities and concentrations.\nResponse: Thank you for your insightful comment. We fully agree that the charge carried by all ions should match the applied RI current. Here, we aim to highlight that electroosmosis is mainly driven by the presence of “reverse ions” (such as Na⁺) in the EDL that carry a positive charge opposite to the negatively charged skin. We may not clearly describe the process, making the misunderstanding, therefore, for the guidance to readers, we adjusted the relevant descriptions to make them more accurate.\nChanges: The sentences below were revised in section “Theoretical analysis” to reflect the changes explained above. For your convenience, we added the text here:\n“The electromigration involves ionic fluxes originating from both the bulk ISF and the diffusion layer. In contrast, the driving force for the electroosmotic transport of small neutral molecules primarily arises from the migration of reverse ions (carrying charges opposite to skin surface) within the diffuse layer under an applied electric field. Consequently, the concentration of these reverse ions governs the electroosmotic velocity (Fig. 1c and Supplementary Note 1).” (Lines 82-86 in the revised Manuscript)\nThe relevant sentences were also revised in Supplementary Information. For your convenience, we added the text here:\n“The electromigration includes the ions flux both in bulk ISF and diffusion layer, whereas the electroosmosis is primarily driven by the migration of reverse ions in the diffusion layer 13-18.” (Lines 30-32 in the revised Supplementary Information)\n“The electroosmotic flow of glucose is mainly induced by the migration of reverse ions in the diffusion layer 14-16,18,19.” (Lines 63-64 in the revised Supplementary Information)\n6. Lines 85-87. I'm not sure that anyone has measured this proton flux or shown that it changes skin surface pH. Incidentally, what is meant by skin surface pH in this context? Without introducing some water on the surface, how is a measurement of pH performed? Without water, what does pH \"of the surface\" even mean?\nResponse: Thank you again for raising the important point regarding H⁺ flow and skin surface pH. As stated in the responses to Comments 1 and 2, our original description of H⁺ flux was inaccurate and we have revised the relevant descriptions. Additionally, the term “skin surface pH” refers to the pH of the entire microfluid at the skin surface and within the epidermal electroosmotic microchannel throughout the RI extraction process (Fig. R1). We refer to this measured value as “skin surface pH” in the manuscript. In our opinion, the extracted microfluid at the skin surface is the same as that in the epidermal electroosmotic microchannel, therefore, we think that the measured pH also reflects the conditions within the epidermal electroosmotic microchannel. This pH value influences the skin’s ZP, thereby affecting the electroosmotic flux of glucose.\nChanges: We have revised the relevant manuscript to clarify the meaning of “skin surface pH”. For your convenience, we added the text here:\n“Specifically, the extracted H⁺, along with the release of additional protons from biochemical reactions at the epidermal interface, collectively leads to changes in the pH of the RI-induced fluid accumulated at the skin surface and within the epidermal electroosmotic microchannel (referred to as the skin surface pH in this study).” (Lines 56-59 in the revised Manuscript)\n“In this process, Na+ dominates the ion flux and acts as the primary carrier in the EDL, while extracted H+ and released H+ from biochemical reactions at the epidermal interface alter the skin surface pH.” (Lines 89-91 in the revised Manuscript)\n7. Lines 87-88. Another statement for which there is no experimental evidence as far as I know. What 'self-recovery' mechanism is invoked here, exactly?\nResponse: Thank you for your detailed observation. As mentioned in Comment 4, we have updated the relevant descriptions and added more appropriate literatures. For your convenience, we added the text here:\n“Additionally, the skin’s self-recovery mechanisms, including natural barriers such as sweat glands and the stratum corneum, help maintain homeostasis and prevent excessive polarization 52, 53, 54, thereby mitigating the pH changes induced by RI.” (Lines 91-94 in the revised Manuscript)\n8. Lines 89-91. It is true that experiments ex vivo where it has been possible to change the pH of solutions on either side of the membrane have shown that lowering the pH from 7.4 to that equal to the isoelectric point of the membrane, and then further to below this value, have first reduced electroosmotic flow and the reversed it. But given the ISF pH of ~7.4, the available [H+] is really small and it's hard to envisage that the flux will be able to substantially alter the surface pH (whatever that means - see earlier comment).\nResponse: Thank you for your insightful comments. We fully agree that the pH of subdermal ISF is approximately 7.4 and that the absolute concentration of free H⁺ under such conditions is indeed low. As you mentioned in Comment 3, H⁺ is not the primary charge carrier compared to Na+; however, due to the high mobility of H⁺ 1 and prolonged RI extraction, these H⁺ can rapidly move at the skin surface and in the epidermal electroosmotic microchannel, leading to pH fluctuations (Fig. R1).\nAt the same time, the skin’s physiological self-recovery mechanisms (such as barrier function restoration, ionic reabsorption, and sweat-mediated buffering) oppose ISF extraction and act to neutralize pH disturbances at the skin surface and in the epidermal electroosmotic microchannel. The interaction between H⁺ movement and the skin’s self-regulatory capacity gives rise to dynamic changes in the skin surface pH, which translate into modulations in ZP of the skin. These modulations, in turn, directly impact the EOF that governs neutral molecule transport, for example, glucose. The in vivo results further demonstrate that the skin surface pH values fluctuate dynamically during the RI extraction process.\nChanges: We have revised the relevant sections to provide a more detailed analysis in our manuscript. For your convenience, we added the text here:\n“Specifically, the extracted H⁺, along with the release of additional protons from biochemical reactions at the epidermal interface, collectively leads to changes in the pH of the RI-induced fluid accumulated at the skin-sensor interface (referred to as the skin surface pH in this study).” (Lines 56-59 in the revised Manuscript)\n“In this process, Na+ dominates the ion flux and acts as the primary carrier in the EDL, while extracted H+ and released H+ from biochemical reactions at the epidermal interface alter the skin surface pH.” (Lines 89-91 in the revised Manuscript)\n“The interaction between the H+ movement and the skin’s self-recovery ability makes the fluctuations of skin surface pH.” (Lines 94-95 in the revised Manuscript)\nReference:\n9. Lines 93-94. Characterization of this \"hydrolipid film\" (HF) would be helpful... What is its nature when a conductive hydrogel is placed between the RI electrodes and the skin? How is its pH chosen and how - if at all - does this pH change with time with or without application of a RI current? Are the changes the same or different under anode and cathode? Do they reverse when the polarity of the electrodes is switched between extractions to prevent depletion/build-up of AgCl?\nResponse: Thank you for valuable comments. In fact, the “hydrolipid film” (HF) is a physiological layer covering the surface of skin and has minimal direct relevance to the entire RI extraction and detection process. In this work, we focus on the impact of skin surface pH, which refers to the pH of the entire microfluid at the skin surface and within the epidermal electroosmotic microchannel throughout the RI extraction process (Fig. R1). In our opinion, the extracted microfluid at the skin surface is the same as that in the epidermal electroosmotic microchannel, therefore, we think that the measured pH also reflects the conditions within the epidermal electroosmotic microchannel. This pH value influences the skin’s ZP, thereby affecting the electroosmotic flux of glucose.\nTherefore, we did not investigate the changes in the HF during the extraction process.\n10. Lines 105-107. This is all \"old news\", surely? See comment in S.I.\nResponse: We appreciate the reviewer’s comment and fully acknowledge that the fundamental principles regarding EOF of glucose under different pH conditions were explored in the field. However, the aim of our simulation was not to rediscover these classical electrokinetic concepts, but rather to model the changes in EOF of glucose under different skin surface pH conditions during RI extraction. Importantly, the simulation is consistent with our experimental observations (Fig. 3h) and provides a reference for the subsequent pH calibration of glucose extraction flux.\n11. Lines 118-120. More 'old news' here - the advantages of using this approach were identified by Cygnus more than 30 years ago!\nResponse: Thank you for your comment regarding the symmetrical design and polarity switching in our RI system. We fully agree with your perspective and continue to utilize this valuable method in our RI extraction process. We have delated these sentences and revised the relevant descriptions to be more concise.\n12. Lines 122-127. So, do I understand correctly that the pH sensor is detecting the value in the gel near the electrode surface? Is this not near where glucose is converted to gluconic acid (releasing H+ locally)?\nResponse: Thank you for your thoughtful comments. As previously addressed in Comments 1, 2, 3, and 6, the term “skin surface pH” in this study refers to the pH of the extracted microfluid at the skin surface and within the epidermal electroosmotic microchannel throughout the RI extraction process (Fig. R1). In practice, the electrochemical conversion of glucose to gluconic acid occurs near the chitosan- embedded glucose working electrode where GOx is localized. The pH measurements, however, more broadly reflect the pH of the entire fluid at the skin surface and in the electroosmotic microchannel. The released H+ from gluconic acid may also be one factor of skin surface pH fluctuations. Fluctuations in this overall fluid pH can, in turn, affect the skin’s ZP and the EOF of glucose. Thank you again for your valuable comments and reminder.\nChanges: The following sentences were revised in the “Introduction” sention.\n“Specifically, the extracted H⁺, along with the release of additional H⁺ from biochemical reactions, collectively leads to changes in the pH of the RI-induced fluid accumulated at the skin surface and within the epidermal electroosmotic microchannel (referred to as the skin surface pH in this study).” (Lines 56-59 in the revised Manuscript)\n13. Fig. 1c. Presumably, the regions labelled as high and low pH cannot be ISF, as the pH there must be fixed. The figure needs to be more clearly drawn.\nResponse: Thank you for your comment regarding the clarity of the figure. We appreciate your attention to this detail. We understand that the labels “high pH” and “low pH” in the figure might have caused confusion. To clarify, these regions refer to the accumulated fluid via RI at the skin surface, rather than the ISF itself. The skin surface pH values mentioned are indicator of the local skin surface and epidermal electroosmotic microchannel conditions.\nTo address your suggestion, Fig. 1c has been updated with Fig. R2 as follows. We have added annotations and a brief explanation in the figure legend to explicitly state that the high and low pH regions refer to the skin surface conditions. This should help readers better understand the context and avoid any potential misinterpretation. Thank you again for your valuable comments.\nFig. R2. The EDL structure of the skin and the electroosmosis flow under low and high pH Conditions, with the ZP at the shear plane. SNM, small neutral molecules; ZP, zeta potential.\n14. Fig. 2k and 2l. Difficult to assess the relevance of some of this characterisation work performed in the pH range of 5-10. Yes, it's indeed possible that glucose detection is influenced over this range, but surely this does not occur in vivo? So, the more important question is 'over what range does [glucose] vary in a typical RI extraction-detection use of the device?'\nResponse: Thank you for your insightful observation. Indeed, the physiological relevance of pH-dependent glucose sensor performance must be grounded in the actual pH fluctuations observed during RI in vivo. While we initially characterized the glucose sensor across a broad pH range (pH 5-10) to fully establish its operational boundaries, our in vivo data from both healthy and diabetic volunteers confirm that the skin surface pH during RI typically fluctuates within a relevant window-approximately between pH 6 and 10 (Fig. 5b, 5c, and Supplementary Fig. 38).\n15. Line 177. A more objective assessment of performance would be better!\nResponse and changes: We appreciate the reviewer’s request for a more objective and quantitative evaluation of sensor performance. In response, we have revised the relevant paragraph in the manuscript and S.I. to provide a more balanced and objective evaluation. For your convenience, we added the text here:\n“Stepwise testing revealed that the steady-state current of the i-t response increased with glucose concentration (Fig. 2e). The glucose sensor exhibited a highly linear response, with a sensitivity of 236.42 nA mM-1 from 0-3 mM (Fig. 2f) and 28.22 nA mM- 1 from 3 to 30 mM (Supplementary Fig. 9e). Sensor-to-sensor consistency was well maintained with minimal variation in sensitivity across three devices (Supplementary Fig. 9h). The sensor also demonstrated high selectivity against a range of potential interfering species (Fig. 2g and Supplementary Fig. 9e-g). After 14 days of storage, the sensor retained ~75% of its original sensitivity (Fig. 2h and Supplementary Fig. 9i).” (Lines 176-183 in the revised Manuscript)\n16. Lines 178-180. But are you really measuring skin \"surface\" changes in pH here, as opposed to changes in the bulk pH of the hydrogel?\nResponse: We appreciate the reviewer’s important clarification. As noted in our response to Comments 1, 2, 3, 6, 8, and 12, due to the hydrophilic nature of the Nafion hydrogel and the thin layer of fluid that accumulates directly at the skin surface, the measured pH within the hydrogel more accurately reflects the entire microfluid at the skin surface and within the epidermal electroosmotic microchannel throughout the RI extraction process (Fig. R1). We have defined “skin surface pH” upon its first appearance in the manuscript. We hope this revision provides greater clarity and aligns with the reviewer’s concern about the precise origin and interpretation of the pH signal in our system. Thank you again for your detailed comments.\nChanges: The following sentences were revised in the “Introduction” section.\n“Specifically, the extracted H⁺, along with the release of additional H⁺ from biochemical reactions, collectively leads to changes in the pH of the RI-induced fluid accumulated at the skin surface and within the epidermal electroosmotic microchannel (referred to as the skin surface pH in this study).” (Lines 56-59 in the revised Manuscript)\n17. Lines 204-206. This is not a widely accepted model for mammalian skin so data obtained using it must be treated with caution and scepticism. The experiment has been done before, in essence, using mammalian skin so why repeat it here using an artificial membrane?\nResponse: We appreciate the reviewer’s concern and the opportunity to clarify our rationale. While we acknowledge that artificial membranes are not widely accepted as full biological equivalents to mammalian skin, they offer significant advantages in terms of structural and physicochemical stability. In this study, we aim to isolate and investigate the specific influence of skin surface pH on RI-based ISF extraction. Given the inherent variability in mammalian or human skin, such as fluctuations in skin temperature, impedance, hydration, and ISF flow rate, it is difficult to decouple these interdependent factors during in vivo or ex vivo experiments. Otherwise, the artificial membrane holds stable and reproducible properties to be able to control these confounding variables and focus exclusively on the role of pH. This model thus allowed us to validate the theoretical framework of skin surface pH and its impact on electroosmotic transport during RI. The in vitro results serve as a controlled, mechanistic validation of our pH calibration strategy.\n18. Fig. 3. There is really nothing new here except that EO flow can occur across a charged membrane when an electric field is applied. More flow will occur in the direction of counterion movement as the charge on the membrane increases. The flux of a water-soluble, unionised species present in the solvent will be proportional to its concentration, all other factors being the same.\nResponse: We appreciate the reviewer’s comment and fully agree that EOF increases with the surface charge density of a charged membrane under an applied electric field, and that the transport of neutral, water-soluble species is proportional to their concentration when other variables are held constant.\nThe novelty of our work lies not in re-establishing these fundamental principles, but in experimentally demonstrating how dynamic physiological conditions-specifically, changes in skin surface pH during RI-modulate electroosmotic glucose transport. We hypothesize that variations in skin surface pH alter the distribution of surface charges on the epidermal membrane, thereby shifting the zeta potential at the slipping plane within the electric double layer. This, in turn, affects the magnitude of electroosmotic flow. To support this, we experimentally show a near-linear relationship between surface pH and zeta potential (Fig. 1d), as well as between surface pH and glucose flux via electroosmotic transport (Fig. 3h). These findings extend classical EO theory into a biologically relevant, dynamic context and provide a mechanistic basis for our proposed pH calibration strategy. Thank you again for your comment.\n19. Lines 226-228. Again, no surprise here. Na+ flux is driven principally by electromigration, far more than by EO. It's the major positive ion current/charge carrier in the system, so it's obvious that its flux will be insignificantly affected by pH.\nResponse: We appreciate the reviewer’s comments and fully agree that Na⁺ flux is primarily governed by electromigration, and it is the dominant charge carrier in the RI process. Theoretically, the transport of Na⁺ includes both electromigration and electroosmosis. Although EOF may contribute to the overall ionic transport in RI, the Na⁺ flux caused by EOF is negligible compared to direct field-driven migration. Thank you again for your detailed comments.\n20. Lines 244-246. This is well-known as has been reported multiple times before. However, as we do not know the extent to which the charge on a (real) skin membrane in vivo changes during the use of an RI device, there is no basis for concluding that a pH-based correction is needed to ensure that accurate deductions about the glucose levels in the ISF are being made.\nResponse: We thank the reviewer for the comment. Our study tried to quantitatively examine how fluctuations in skin surface pH-measured dynamically during RI-affect glucose extraction efficiency. Importantly, our results revealed a clear, linearly positive correlation between extracted glucose concentrations and skin surface pH, with an average increase of ~33% in extracted glucose per unit increase in pH (Fig. 3h). This empirical observation provides strong evidence that pH-dependent modulation of skin surface charge-and by extension, zeta potential affects electroosmotic glucose transport under RI.\nWhile we agree that the precise dynamics of surface charge change on real skin in vivo remain complex, our data suggest that skin surface pH can serve as an accessible and physiologically relevant indicator for such charge variations. Therefore, incorporating pH into a correction algorithm is not only a theoretical precaution, but also a data-driven strategy to improve the accuracy and reliability of glucose quantification in RI-based sensing systems.\n21. Lines 255-257. How does it do that? As Na+ is the principal charge carrier and RI is done under constant current conditions, one expects (and it's true) that the sodium flux per unit current per unit time is actually quite constant, and this has been shown in vitro and in vivo in man. So how can JNa+ be \"at the base of subdermal glucose prediction\"?\nResponse: We appreciate the reviewer’s insightful question regarding the role of Na⁺ flux in glucose prediction under RI, particularly given the constant current conditions and the established stability of Na⁺ flux per unit current over time. Our rationale is based on the electroosmotic mechanism of glucose transport by RI. Na⁺ flux is not only a charge carrier but also the driving force behind electroosmotic solvent flow, which carries glucose from the dermal layers to the skin surface.\nMoreover, although the system is operated under constant current, physiological factors such as skin temperature, local impedance, hydration state, and sweat gland activity can subtly alter the effective ion transport dynamics across the skin. These variations can introduce fluctuations in Na⁺ flux, which in turn affect the electroosmotic flow and the amount of glucose extracted. Therefore, even under constant current, Na+ flux still reflects real-time extraction conditions and serves as a useful internal reference for correcting glucose signal drift. This concept has been experimentally validated in prior studies, demonstrating that Na⁺ calibration significantly improved blood glucose prediction accuracy in wearable RI devices, underscoring its predictive value in practical applications 5-7. Importantly, in our pH-calibrated glucose prediction algorithm, we continue to incorporate Na⁺ calibration as a foundational element.\nTogether, Na⁺ and pH calibration address both the driving force and modulation factors of RI-based glucose extraction, yielding enhanced predictive accuracy.\nReferences:\n22. Line 261. Sub-heading does not make complete sense.\nResponse and changes: Thank you for your valuable comments. The section titled “In- vitro evaluation of the calibration method and wearable device” is indeed divided into three parts. The first sub-heading here is intended to highlight the original measurements of glucose i-t current and pH voltage during the extraction process. To improve clarity and more accurately reflect the content, we have revised the sub- heading to more precisely capture the core focus of this section, which describes how glucose levels and skin surface pH change throughout the continuous RI process. For your convenience, we have added the text here:\n“Dynamic changes in glucose and skin surface pH during RI extraction” (Line 269 in the revised Manuscript)\n23. Lines 272-276. The system is symmetrical so why are the i-t responses different in Figs. 4b and 4e, and why are the pH fluctuations different (Figs. 4c and 4f), both in absolute values and the extent of change of pH - less than 0.1 unit in Part A, more than 0.2 unit in Part B? These don't seem to be massive changes in pH in any case and the impact as reported in Figs. 4d, 4g, 4h and 4i confirm that this is the case. The \"explanation\" offered in this paragraph is not convincing, to say the least (mostly, because it does not explain anything!).\nResponse: Thank you for your careful evaluation and for pointing out the need for more explanations.\nAlthough the device is structurally symmetrical, with identical configurations for Part A and Part B, the microenvironment of the skin-sensor interface introduces unavoidable asymmetries that can affect ion mobility, extraction efficiency, and electrochemical reaction rates, resulting in the observed differences in i-t responses (Figs. 4b and 4e) and pH fluctuations (Figs. 4c and 4f). In particular, the difference in the extent of pH variation, less than 0.1 unit in Part A compared to more than 0.2 units in Part B, likely reflects unequal H+ movement, localized buffering capacity, and dynamic ionic exchange during RI.\nWhile the absolute magnitude of pH change appears small, it is important to note that even subtle variations in skin surface pH can lead to measurable shifts in ZP, which in turn modulate EOF of glucose. Our experimental findings and simulations both support the conclusion that these pH fluctuations could affect the EOF of glucose by RI. This is evident in the correlation trends shown in Figs. 3h and 4h, and in the improved predictive accuracy after pH calibration presented in Fig. 4i.\n24. Lines 295-297. Exactly how were the data from Parts A and B combined?\nResponse: Thank you for your question. We apologize for the lack of clarity in the original description and appreciate the opportunity to elaborate on our data processing methodology.\nAs illustrated in S.I. Fig. 26, the in vitro experimental protocol involved alternating ISF extraction and detection between two symmetrical sensing regions (Part A and Part B). For example, in the first cycle of forward extraction and detection, “Data point 1” (Part A) was generated by calculating the difference of glucose concentration measured, and the average of pH values recorded, both before and after a 3-minute RI extraction cycle. In the second cycle, the polarity was switched, so “Data point 2” (Part B) was generated similarly by calculating the difference in glucose concentration measured, and the average of the pH values recorded, both before and after the extraction. Subsequently, this process was repeated by alternating between Part A and Part B, achieving continuous monitoring.\nChanges: The following sentences were added to the caption of S.I. Fig.27 to clarify this methodology. For your convenience, we added the text here:\n“The in vitro experimental protocol involved alternating ISF extraction and detection between two symmetrical sensing regions (Part A and Part B). For example, in the first cycle of forward extraction and detection, “Data point 1” (Part A) was generated by calculating the difference of glucose concentration measured, and the average of pH values recorded, both before and after a 3-minute RI extraction cycle. In the second cycle, the polarity was switched, so “Data point 2” (Part B) was generated similarly by calculating the difference in glucose concentration measured, and the average of the pH values recorded, both before and after the extraction. Subsequently, this process was repeated by alternating between Part A and Part B, achieving continuous monitoring.” (Lines 432-440 in the revised Supplementary Information)\n25. Lines 298-300. The data in Fig. 4h show a decent correlation between extracted glucose and subdermal glucose both with and without pH 'calibration'. It would be good to see confidence intervals on the slopes and to test whether they are significantly different. In a practical sense, it seems unlikely that the differences in slopes would dramatically impact on the clinical value of the data when one adds in the pH correction.\nResponse: We appreciate the reviewer’s thoughtful suggestion and have conducted a statistical comparison of the slopes in Fig. 4h with and without pH calibration. As noted, the difference between the two slopes was not statistically significant (Fig. R3).\nWe believe this result is primarily due to the fact that the artificial membrane holds stable and reproducible properties but lacks the self-recovery function, thus leading to minimal pH fluctuations. In contrast, normal human skin, when subjected to RI stimulation, has a self-defense behavior against polarization. The skin’s natural barriers, such as sweat glands and the stratum corneum, also play a stabilizing role in maintaining ion homeostasis and preventing excessive polarization. Additionally, to our knowledge, the slope of the correlation curve reflects the properties of the skin. Variations in slop between different body of the same individual or between different individuals can significantly affect the relationship between skin surface and subcutaneous glucose concentrations. Therefore, calibration is necessary to overcome the RI extraction fluctuations caused by these differences.\nWe performed a statistical analysis of the Mean Absolute Relative Difference (MARD) values for in vivo glucose prediction with and without calibration. The results show that MARD values were significantly reduced after applying the combined Na⁺ and pH calibration, confirming the practical utility of the calibrated RI for real-time, noninvasive glucose monitoring (Fig. R4).\nFig. R3. Correlation curves of extracted glucose and subcutaneous glucose with and without pH calibration of the overall test results (n = 6).\nFig. R4. Mard values of the predicted glucose and blood glucose without calibration, with Na+ calibration, and with combined Na+ and pH calibration. *P < 0.05, **P < 0.01, and ***P < 0.001 (one-way ANOVA with Tukey’s multiple comparison)\nChanges: Fig. R4 was added to the Supplementary Fig. 41b. (Lines 534-537 in the revised Supplementary Information)\n26. Fig. 5. [1] Are the units of extracted glucose correct in 5b and 5c? Should this be micromolar? [2] In the same panels, pH is increasing from about 6 to 10, the latter likely to cause some skin irritation, no? What is causing the increase in alkalinity? Why does the pH change in a sawtooth manner for the normal volunteer but not for the person living with T2D? The fluctuation in predicted glucose levels is more sawtooth for the diabetic patient (whose pH is not varying like that) than that for the normal patient (whose pH is up-and-down throughout) How (in)consistent was that behaviour?\nResponse: Thank you for your careful and thoughtful comments. We address each point below to provide a more comprehensive explanation.\nAs shown in Fig. 5b, 5c, and Supplementary Fig. 39, the extracted glucose concentrations are in the range of tens of µM to mM grade. For consistency, the y-axis is labeled in mM.\nIn our study, we indeed observed that the pH at the skin surface can transiently reach up to 10 during the RI process. But no adverse skin reactions were observed during or after the in vivo measurements in either healthy or diabetic participants. We think the reason may be that the pH changes are transient and occur within a localized microenvironment, simultaneously, the skin’s natural buffering capacity and the short duration of the pH elevation help mitigate the potential irritation. We will continue to monitor and refine these aspects in future studies to further enhance the safety and comfort of our device.\nThe pH changes observed during the extraction process exhibit stochasticity across different individuals. Supplementary Figs. 39d and 39e show relatively stable changes, which may stem from differences in the skin’s natural buffering capacity and self- recovery mechanism. As shown in Supplementary Fig. 39d, the measured pH changes in type 2 diabetes (T2D) patients also exhibit a sawtooth pattern, and these changes are stochastic. The fluctuations in predicted glucose levels also show no consistent pattern between healthy individuals (Fig. 5b and Supplementary Fig. 39, a to c) and diabetic patients (Fig. 5c and Supplementary Fig. 39, d to f), with no obvious regularity. This inconsistency likely stems from differences in skin properties among individuals. This irregular phenomenon underscores the complexity of RI-based sensing in vivo and further confirms the necessity of pH calibration.\nMethods\n27. Line 512. I don't believe that this sensitivity can be measured to an accuracy of 8 picoAmps (the 3rd decimal place here)!\nResponse: Thank you for raising this important point regarding the reported current sensitivity precision. In our experiments, the current was measured using a CHI 660E electrochemical workstation, which has a minimum measurable current resolution of ±10 picoamperes. However, when calculating the sensitivity of the glucose sensor, we inadvertently retained three decimal places in the unit of μA mM-1, corresponding to sub-picoampere levels-beyond the actual resolution of the measurement system.\nWe have revised the sensitivity values in the manuscript. The updated values are now reported in nA mM-1 with one decimal place, in accordance with the instrument’s resolution and the meaningful significant figures. All related descriptions, figures, and data tables throughout the manuscript have also been corrected accordingly. Thank you again for your detailed comment. For your convenience, we added the text here:\n“The sensitivity of glucose within the range of 3-30 mM is 28.22 nA mM-1 (Supplementary Fig. 9d).” (Lines 529-530 in the revised Manuscript)\nComments to the authors on the Supplementary Information\n28. Lines 30-31. How do you know this as a fact? Based on what evidence, exactly?\nResponse: Thank you for the detailed comments. We acknowledge that the original description was inaccurate and missed relevant references in this section.\nElectromigration refers to the movement of charged species under an applied electric field. Since the electric field can penetrate into the bulk of the ISF, ions throughout the conductive medium (both in the diffusion layer and bulk) can experience a driving force, resulting in ionic electromigration. In contrast, electroosmosis arises from the electrical double layer (EDL) near charged skin, and the driving force of electroosmosis originates from the drag effect of reverse ions (such as Na+) within the EDL 8-13. Thank you again for your detailed reminder.\nReferences:\nChanges: The sentences were revised and the relevant references were added accordingly. For your convenience, we added the text here:\n“The electromigration includes the ions flux both in bulk ISF and diffusion layer, whereas the electroosmosis is primarily driven by the migration of reverse ions in the diffusion layer 13-18.” (Lines 30-32 in the revised Supplementary Information)\n29. Lines 38-39. Which is which? Whatever is the phase on the skin surface, it is hard to describe it as liquid...\nResponse: Thank you for your insightful comment regarding the description of the phases at the skin surface. We apologize for the inaccurate phrasing in our original description. Here, we intend to convey that charge separation occurs at the interface between the epidermal electroosmotic microchannel (solid phase) and the extracted ISF (liquid phase), leading to the formation of an EDL.\nChanges: To clarify this point, we have revised the sentences to more accurately describe the formation of EDL at different phases. For your convenience, we added the text here:\n“Charge separation occurs at the interface between the epidermal electroosmotic microchannel (solid phase) and the extracted ISF (liquid phase), leading to the formation of an electric double layer (EDL) structure.” (Lines 38-40 in the revised Supplementary Information)\n30. Lines 59-60. Same statement as above but does not obviously flow from what is written here which appears to be standard, accepted dogma.\nResponse: Thank you again for your comments. We have corrected the original description and added the appropriate references to support this statement and clarify its scientific basis. Furthermore, we have reviewed the descriptions related to electroosmosis to ensure accuracy. For your convenience, we added the text here:\nChanges: “The electroosmotic flow of glucose is mainly induced by the migration of reverse ions in the diffusion layer 13-18.” (Lines 63-64 in the revised Supplementary Information)\n31. General remark on Suppl. Note 2. OK, not really anything terribly new or controversial here, it seems. The bottom line appears to be that EO flow depends on the charge on the membrane and that the latter depends on the pH of the 'liquid' phases on either side. The premise, therefore, appears to be that, if one can control the pH of the 'liquid' phase on the skin surface, one can be sure that EO flow will be fixed during RI and that the glucose detected in the 'receiving' gel will therefore be proportional to that in the ISF.\nResponse: Thank you for your accurate summary of the premise. We fully agree that EOF during RI is governed by the surface charge of the skin, which in turn is sensitive to the pH of the liquid phases. As the reviewer correctly noted, controlling the skin surface pH can stabilize the EOF by RI, thereby ensuring consistent analyte extraction. While this conceptual framework aligns with accepted electrokinetic theory, in practice, maintaining the skin’s native charge under electrical stimulation is quite difficult. To our knowledge, no existing methods have been proposed to date. Our work therefore explores pH compensation as a strategy to calibrate the EOF changes dynamically. We are grateful for the reviewer’s comment, which allowed us to better articulate this point.\n32. Line 121. The data from which this empirical equation is derived was acquired over a very wide - and non-physiological - range of pH, so its relevance to the 'real world' is not established. What is the form of the relationship when the range of pH values is limited to something sensible/realistic?\nResponse: Thank you for your important observation regarding the physiological relevance of the pH range. We fully agree that the applicability of any empirical model should be assessed within a physiologically relevant range. As mentioned in Comment 14, the physiological relevance of pH-dependent glucose sensor performance must be grounded in the actual pH fluctuations observed during RI in vivo. While we initially characterized the glucose sensor across a broad pH range (pH 5-10) to fully establish its operational boundaries, our in vivo data from both healthy and diabetic volunteers confirm that the skin surface pH during RI typically fluctuates within a relevant window-approximately between pH 6 and 10 (Fig. 5b, 5c, and Supplementary Fig. 38).\n33. Line 124. And the electroosmotic component? Missing?\nResponse: Thank you for your detailed comment regarding the electroosmotic component of Na+. As the reviewer correctly mentioned in Comment 19, Na+ flux is driven principally by electromigration, far more than by EOF; and it is the dominant charge carrier in the RI process. Theoretically, the transport of Na⁺ includes both electromigration and electroosmosis. Although EOF may contribute to the overall ionic transport in RI, the Na⁺ flux caused by EOF is negligible compared to direct field-driven migration. Thank you again for your detailed comments.\n34. Lines 142-146. I’m not sure that I would write the equation like this. The units of uglu are velocity. The product A x uglu is therefore equivalent to the EO volume flow rate induced by the imposition of the electric field across the charged skin membrane. It seems to me that A x uglu must have the same value for all small, neutral molecules present in the system - how can it be specific to glucose?\nResponse: Thank you for your insightful comment regarding the formulation of the electroosmotic velocity in Lines 142-146.\nWe fully agree that the product A * u represents the EO volume flow rate (v ) glu glu induced by the electric field across the charged skin membrane. This is indeed a macroscopic perspective that is evident from the results. However, our intention in this section is to emphasize the more microscopic expression of electroosmotic velocity, specifically its relationship with the electric field. The electroosmotic velocity is expressed as the product of the electroosmotic mobility (μ ) and the electric field glu strength (E), as shown in Line 146. The electroosmotic mobility, which is the velocity per unit electric field, is directly related to the ZP (ζ) within the EDL. This relationship is fundamental to our study, as the ZP is a critical parameter influencing the EOF.\nWe fully agree with your observation that the EOF should represent the movement of all small, neutral molecules by RI and is not specific to glucose. However, we are focusing on glucose as the target analyte in our RI-based system. Therefore, we use the electroosmotic velocity of glucose to represent the overall EOF of the ISF. Thank you again for your detailed comments.\nChanges: We have revised the text to clarify this point and ensure that the relationship between EO volume flow rate (v ), electroosmotic velocity (u ), electroosmotic glu glu mobility (μ ), and zeta potential (ζ) is clearly articulated. For your convenience, we glu added the text here:\n“In equation (1.14), represents the proportionality coefficient between the glucose 3 concentration in the extracted ISF and the glucose concentration in the subcutaneous 𝑘𝑘 ISF, v represents the electroosmotic volume flow rate which is equivalent to A glu multiplied by u (electroosmotic velocity of glucose), and c is the glucose glu iglu concentration in the subdermal ISF.” (Lines 158-161 in the revised Supplementary Information)\n35. General remark on Suppl. Note 3. So, at the end of the day, the approach used requires measurements in the sensor gel of glucose, Na+ and pH. Ref 22 in the SI used Na+ as an 'internal standard', effectively to correct for how the charge on the skin may vary from person to person and hence how EO flow would be affected as a result. Did the authors check to see if the approach outlined here, without considering the impact of small changes in pH, would work just as well? If the answer is no, then they probably should...\nResponse: Thank you for your general remark on Supplementary Note 3. We fully agree that skin surface pH calibration proposed in this work requires measurements of glucose, Na+ and pH. However, regarding Ref. 22 in the S.I., the use of Na⁺ as an “internal standard” was to provide a stable reference point for calibration, thereby minimizing the need for frequent blood draws. Additionally, the assumption of using Na+ as an internal standard is based on the premise that the extracted flux ratio (glucose/ Na+) would vary linearly with the subdermal glucose/ Na+ concentration ratio. However, we found that this relationship is not linear and can be influenced by various factors, including the electrical properties of the skin. To address this issue, we proposed pH calibration to compensate the extraction fluctuations caused by changes in skin electrical properties due to pH variations.\nWe have indeed evaluated the effectiveness of our method both with and without considering pH calibration. In Oral Glucose Tolerance Test (OGTT) experiments, we assessed the performance of our system using Na+ calibration alone. As shown in Fig. 5g (healthy) and 5k (diabetic), the Mean Absolute Relative Difference (MARD) values were 20.04% (healthy) and 33.24% (diabetic) when only Na+ calibration was applied. When we further incorporated pH calibration into the model, the MARD values were significantly reduced to 10.91% (healthy) and 18.10% (diabetic), respectively. These results clearly demonstrate that accounting for pH changes in addition to Na+ calibration significantly improves the accuracy of glucose predictions. Thank you again for your valuable comments.\n36. S.I Fig. 10. On many graphs, using stronger colours for the datapoints and the lines connecting them would be helpful.\nResponse: Thank you for your suggestion regarding the clarity of the graphs in Supplementary Information Fig. 10. We appreciate your comments and agree that enhancing the visibility of the data points and connecting lines can improve the readability of the figures.\nChanges: We have now revised the graphs in Supplementary Fig. 10 by using stronger colors for the data points and lines. We have also double-checked other figures (Supplementary Figs. 8, 9, 12, and 13) to ensure that the color choices are consistent across the graphs to maintain coherence and facilitate easier interpretation. We hope these revisions address your concerns and improve the overall presentation of our figures.\n37. S.I Fig. 10. Why are dependencies so different as a function of concentration at different pH values? I'm not sure I'd want a sensor on my arm that led to a pH of 9 or 10...\nResponse: Thank you for your valuable comments. As shown in S.I. Fig. 10e, the correlation curves for the glucose sensor exhibit different sensitivities at various pH levels. This variation is primarily due to the influence of pH on enzyme activity. Different pH conditions can lead to changes in enzyme conformation and activity, thereby affecting the sensor’s sensitivity 14.\nIn our study, we indeed observed that the pH at the skin surface can transiently reach up to 10 during the RI process. But no adverse skin reactions were observed during or after the in vivo measurements in either healthy or diabetic participants. We think the reason may be that the pH changes are transient and occur within a localized microenvironment, simultaneously, the skin’s natural buffering capacity and the short duration of the pH elevation help mitigate the potential irritation. We will continue to monitor and refine these aspects in future studies to further enhance the safety and comfort of our device. Thank you for raising this important point.\nReference:\n38. S.I Fig. 13b. Impossible to see data on this panel.\nResponse: Thank you for your detailed attention. We apologize for the poor visibility of the data in S.I. Fig. 13b. In response, we have revised the figure to enhance the visibility of the data points and connecting lines. We also reviewed and revised other figures (Supplementary Figs. 8, 9, 12, and 13) to ensure consistency and clarity across all supplementary materials. Each figure has been checked for resolution to ensure clarity. Thank you again for your valuable comments.\n39. S.I Fig. 24. Some explanation of what one is supposed to be seeing here would be helpful... significance of colouring?\nResponse: Thank you for your comment. We apologize for the oversight in providing a detailed explanation for this figure.\nWe observed that after a single-polarity extraction, the anode of the RI electrode would darken, while the cathode remained relatively unchanged. To investigate the impact of polarity switching on the anode extraction electrode, we conducted a series of experiments, and the results of Panels a, b, and c in S.I. Fig. 25 depict the surface morphology of the anode before extraction, after a single-polarity extraction cycle, and after polarity switching post single polarity extraction, respectively. In each panel, (ⅰ) represents the SEM image, while (ⅱ-ⅳ) show the elemental surface mapping for Ag, Cl, and O, respectively. The surface mapping illustrates the distribution of these elements on the anode surface, with brighter (or colored) areas indicating higher elemental concentrations.\nThe elemental mapping (ⅱ-ⅳ) shows an increase in Cl content after single-polarity extraction, which is restored upon polarity switching. This increase in Cl content leads to electrode darkening due to AgCl formation, which hinders efficient current transfer. However, polarity switching restores the surface and mitigates these effects.\nChanges: The figure captions for S.I. Fig. 25 have been updated, and relevant explanations have been added.\nSupplementary Fig. 25. (a) SEM images and elemental surface distribution of the anode electrode before RI extraction. (b) SEM images and elemental surface distribution of the anode electrode after a single-polarity extraction. (c) SEM images and elemental surface distribution of the anode electrode after polarity switching post single polarity extraction. Scar bars, 30 μm. Each RI cycle applied a current density of 150 μA cm-2 for 3 minutes. The SEM and elemental mapping results show that single-polarity extraction makes the anode surface rougher and increases Cl content, causing electrode darkening due to AgCl formation. This hinders efficient current transfer. Polarity switching restores the surface and mitigates these effects.\n40. S.I. Fig. 26. Impossible to see bars in the \"Data point 1\" and \"Data point 3\" boxes.\nResponse: Thank you for pointing out the issue with the visibility of the bars in the “Data point 1” and “Data point 3” boxes in S.I. Fig. 26. We apologize for the oversight and any confusion may have caused.\nOur original intention was to differentiate between Part A and Part B by using lighter colors for the positive y-axis values (Part A) to contrast with the negative y-axis values (Part B). We recognize that this choice of color scheme may have compromised the visibility of the data bars. To address this issue, we have revised S.I. Fig. 27 to enhance the visibility of the data bars.We have also updated Fig. 5a to maintain consistency across these figures. We hope these revisions improve the clarity and readability of our data presentation. Thank you again for your detailes.\nSupplementary Fig. 27. Schematic of the in-vitro procedure, illustrating the relationship between the measurements and each data point related to subdermal ISF glucose. Each “data point” was generated by calculating the difference in glucose concentration measured before and after a 3-minute RI extraction cycle. The corresponding pH value for that data point was taken as the average of the pH values recorded immediately before and after the extraction.\nFig. 5a Schematic of the OGTT procedure, illustrating the relationship between the measurements and each data point related to blood glucose. Inset: Optical images of a female healthy participant (left), a male diabetic patient (middle) wearing the device, and the device applied to the arm during the OGTT (right). Scar bars, 1 cm. OGTT, oral glucose tolerance test.\n41. S.I. Fig. 27. Minimal variation in pH here…\nResponse: Thank you for your detailed comments. As mentioned in Comment 25, the artificial membrane used in our in vitro experiments exhibits stable and reproducible properties, allowing us to control confounding variables and focus exclusively on the role of pH. We believe the minimal pH fluctuations observed are primarily due to the artificial membrane’s stable and reproducible nature. Unlike human skin, this membrane lacks self-recovery functions, thus leading to minimal pH fluctuations.\n42. S.I. Fig. 29. Again difficult to understand the manner in which pH varies over extraction times.\nResponse: Thank you for your detailed comments. As mentioned in Comment 16, the extracted H⁺, along with the release of additional H⁺ from biochemical reactions, collectively leads to changes in the pH of the RI-induced fluid accumulated at the skin surface and within the epidermal electroosmotic microchannel (referred to as the skin surface pH in this study). Therefore, as extraction progresses, the pH varies irregularly over extraction times.\n43. Same comment on S.I. Figs, 30, 31, 32, 34, 35.\nResponse: Thank you again for your detailed comments. As mentioned in Comment 25 and 41, the artificial membrane used in our in vitro experiments exhibits stable and reproducible properties, allowing us to control confounding variables and focus exclusively on the role of pH. We believe the minimal pH fluctuations observed are primarily due to the artificial membrane’s stable and reproducible nature. Unlike human skin, this membrane lacks self-recovery functions, thus leading to minimal pH fluctuations.\n44. S.I. Fig. 32. But what happened in panel g?\nResponse: Thank you for your comment. Although the device is structurally symmetrical, with identical configurations for Part A and Part B, the microenvironment of the skin-sensor interface introduces unavoidable asymmetries that can affect ion mobility, extraction efficiency, and electrochemical reaction rates, resulting in irregular pH fluctuations.\n45. S.I. Fig. 37. Makes more sense to use the same y-axis scale for these two graphs. The correlations between glucose extracted and subdermal glucose are pretty good with or without pH calibration, it seems to me. I'd like to see confidence intervals on the slopes and whether the with and without pH calibration values are significantly different or not.\nResponse: Thank for your valuable comments. S.I. Fig. 38 was revised to apply a consistent y-axis scale across both panels to facilitate direct comparison. We have conducted a statistical comparison of the slopes in Fig. 4h with and without pH calibration. As noted, the difference between the two slopes was not statistically significant (Fig. R5). We believe this result is primarily due to the fact that the artificial membrane holds stable and reproducible properties but lacks self-recovery function, thus leading to minimal pH fluctuations. Thank you again for your details.\nFig. R5. Calibration curves of extracted glucose and subcutaneous glucose with and without pH calibration of the overall test results (n = 6). (a) Part A of the device. (b) Part B of the device.\n46. S.I. Fig. 38. Would be helpful if the data for the healthy volunteer were shown using the same y-axes for extracted glucose, Na+ and pH. Similarly for the diabetic subject. Please use bolder colours to facilitate visualisation. Why are the sawtooth variations prominent in some curves but not others?\nResponse: Thank you for your valuable suggestion. We have revised Fig. 5b, 5c, and S.I. Fig. 39 by standardizing the y-axis ranges for extracted glucose, Na⁺, and pH across both healthy and diabetic subjects, and by enhancing the color contrast with bolder lines to improve visual clarity.\nThe pH changes observed during the extraction process exhibit stochasticity across different individuals. Supplementary Figs. 39d and 39e show relatively stable changes, which may stem from differences in the skin’s natural buffering capacity. As shown in Supplementary Fig. 39d, the measured pH changes in type 2 diabetes (T2D) patients also exhibit a sawtooth pattern, and these changes are stochastic. The fluctuations in predicted glucose levels also show no consistent pattern between healthy individuals (Fig. 5b and Supplementary Fig. 39, a to c) and diabetic patients (Fig. 5c and Supplementary Fig. 39, d to f), with no obvious regularity. This inconsistency likely stems from differences in skin properties among individuals. This irregular phenomenon underscores the complexity of RI-based sensing in vivo and further confirms the necessity of pH calibration.\nSupplementary Fig. 39 (a to c) Variations in extracted glucose concentration, Na⁺ concentration, and skin surface pH during a single OGTT (left); Blood glucose predictions from the fabricated device- without calibration, with Na⁺ calibration, and with cooperative Na⁺/pH calibration (right) for healthy participant. (d to f) Corresponding results for diabetic patients.\n47. S.I. Fig. 42. What skin pH values were recorded for this individual?\nResponse: Thank you for your question. The results shown in S.I. Fig. 43 correspond to the healthy individual depicted in Fig. 5b. In our earlier experiments, we observed that increasing the extraction current intensified the electrical stimulation on the skin, leading to skin darkening, which took about two weeks to resolve, specifically, the burning shape was consistent with that of the extraction electrodes. In this work, as shown in S.I. Fig. 43, the red marks were consistent with the shape of the whole Part A and Part B, indicating that they were caused by pressure rather than electrical stimulation. Importantly, S.I. Fig. 43 demonstrates that the skin returned to its normal state within 3 hours after the completion of the OGTT. Moreover, we conducted the OGTT on a total of 6 healthy individuals and 15 diabetic patients. No adverse skin reactions were observed during or after the in OGTT in either healthy or diabetic participants. Thank you again for your reminder.\nChanges: The following sentence was revised to the caption of Supplementary Fig. 43.\n“The images show that the red marks were consistent with the shape of the whole Part A and Part B, indicating that they were caused by pressure rather than electrical stimulation. The skin returned to its normal state after 3 hours, with no discomfort or adverse effects reported, indicating the safety and reversibility of the proposed wearable device. All volunteers, including 6 healthy individuals and 15 diabetic patients, reported no discomfort or adverse effects following the test.” (Lines 545-550 in the revised Supplementary Information)\n48. Suppl. Table 2. References for the source of the values in this table should be provided.\nResponse: Thank you for pointing this out. We apologize for the oversight and have now added the appropriate references for the values listed in Supplementary Table 2. The updated table includes citations for each source, ensuring full transparency and traceability of the data. We have also double-checked other sections that require references to ensure that all necessary citations are included and accurately reflect the sources of our data and methodologies. Thank you for your attention to detail.\nSupplementary Table 2. Parameters setting of the numerical simulation.\nParameter Value Units Ref.\nD 5.26E-10 m2 s-1 -\nNa+\nD 2.64E-10 m2 s-1 21\nglu\nz 1 - -\nR 8.314 J (mol*K) -1 -\nT 310.15 K -\nF 9.65e4 C mol-1 -\nε\n2.50e-3 - 22\np\na 2.9 nm 23\nμ 0.0035 Pa*s 24\nτ 10 - -\nε 7.1e-10 F m-1 -\nw\nc\n5 mM 25\niglu0\nc\n140 mM 25\niNa+0\nReferences:\n49. Suppl. Table 5. Why do the slopes sometimes agree in Parts A and B, and sometimes they do not? Why is the % increment with pH calibration not consistent?\nResponse: Thank you for your valuable comments. As mentioned in Comment 23, although the device is structurally symmetrical, with identical configurations for Part A and Part B, the microenvironment of the skin-sensor interface introduces unavoidable asymmetries. These asymmetries can affect ion mobility, extraction efficiency, and electrochemical reaction rates, resulting in the observed differences in i-t responses (Figs. 4b and 4e) and pH fluctuations (Figs. 4c and 4f). Consequently, the fitted slopes for Part A and Part B are also inconsistent.\nReviewer #3:\nGeneral Comments: This manuscript is an excellent addition to the literature. The authors have identified a pain point in the development of on-body iontophoretic devices and have provided an innovative solution for it. Polarity switching to minimize current instability is a notable feature of the proposed device design to mitigate skin irritation. Moreover, the proof of concept of pH fluctuation as well as the establishment of a solution via calibration is commendable.\nResponse: Thank you for your valuable and insightful comments of our work. We have carefully reviewed each of your concerns and have made substantial revisions to the manuscript accordingly. Below, we provide detailed point-by-point responses to your comments.\n1. While I suggest that the manuscript be accepted for publication, certain aspects can be included in the discussion section, which can help delineate the relevance and practicality of the designed device:\nResponse: Thank you for your positive and encouraging comments on our manuscript. We fully agree that further elaboration in the discussion section could enhance the clarity and practicality of the designed model/device. We have revised the discussion section to include further details on the potential applications, limitations, and future directions. We hope these additions enhance the clarity and contextual relevance of our findings. Thank you once again for your support.\nChanges: The “Results and Discussion” section has been divided into two distinct parts: “Results” and “Discussion”. And a detailed description outlining the limitations of this work and potential directions for future research were added to “Discussion” section. For your convenience, we added the text here:\n“Future work may focus on enhancing the personalization, generalizability, and safety of the pH-calibrated reverse iontophoresis (RI) sensing platform. More refined models of blood-to-ISF biomarkers kinetics may be developed to address physiological delays and achieve personalized calibration. In vivo results demonstrate accurate glucose prediction within physiological range of 4-30 mM across healthy and diabetic subjects. Future work should validate this capability by extending the tested range. Additionally, our pH-calibrated RI device can to be broadly applicable to a wide range of biomarkers, including other electroosmotically extractable analytes such as alcohol, ascorbic acid, caffeine, urea, and cortisol, which are also subject to extraction fluctuations influenced by pH. Therefore, the pH-calibrated RI model and device can be extended to other extractable biomarkers, enabling precise ISF-based multi-biomarkers detection. To ensure safety in long-term applications, future device iterations may explore methods to protect the skin under extreme pH conditions and incorporate circuit protection features such as real-time impedance monitoring, enabling automatic power cut-off in response to abnormal current or impedance changes.” (Lines 396-409 in the revised Manuscript)\n2. Is there any lag or delay in the physiological status of markers in blood versus their effect in ISF? Can you please provide some references to this aspect?\nResponse: Thank you for your valuable comment. As you correctly pointed out, there is indeed a physiological delay between biomarker levels in blood and those in interstitial fluid (ISF). For glucose, this time lag typically ranges from 5 to 15 minutes (Supplementary Table 8) ¹-⁸. This delay arises from the diffusion process of glucose from the bloodstream into the interstitial space, which is modulated by factors such as tissue metabolism and capillary permeability. One of the most widely accepted models for describing this relationship is the two-compartment model proposed by Kerstin Rebrin et al. 4. The physiological delay is influenced by the rate of glucose change and tends to be more pronounced during rapid fluctuations, such as those occurring with exercise or insulin administration 2.\nIn this work, we adopted a fixed delay time of 10 minutes for ISF glucose monitoring, consistent with commonly reported physiological lags. Accordingly, predicted blood glucose values are based on ISF measurements taken 10 minutes earlier. We agree that addressing this delay is crucial, and future work will focus on developing more personalized models of blood to ISF kinetics. Thank you once again for highlighting this important point.\nSupplementary Table 8. Summary of reported time lags between blood glucose (BG) and interstitial fluid (ISF) glucose levels.\nTest Time Lag Ref.\nHealthy, diabetic(T1DM) Microdialysis 9.1/11 min 1\nHealthy Microdialysis 5-6 min 2\nDiabetic (T1DM) Microdialysis 9.8 min 3\nHealthy, diabetic(T1DM) - 5-10 min 4\nDiabetic (T1DM) - 10 min 5\nDiabetic (T1DM) - 15 min 6\nDiabetic (T1/T2DM) CGM (Glucowatch) 5 min 7\nDiabetic (T1DM) CGM (FreeStyle Navigator) 12.5 min 8\nReferences:\nChanges: Supplementary Table 8 was added to the Supplementary Information. Corresponding text was also incorporated into the “Discussion” and “OGTT Procedure” sections to reflect and support these additions. For your convenience, we added the text here:\n“More refined models of blood-to-ISF biomarkers kinetics may be developed to address physiological delays and achieve personalized calibration.” (Lines 397-399 in the revised Manuscript)\n“The time lag between blood glucose and ISF glucose is established as 10 minutes, consistent with commonly reported physiological lags 62, 63 (Supplementary Table 8).” (Lines 630-632 in the revised Manuscript)\n3. In relation to my first question, is this lag (if any) different for different markers, hence making such a device more suitable for certain biomarkers (such as glucose) more than others? Essentially, can this same device be used for any biomarker, or will the calibration or other aspects be specific to a single biomarker?\nResponse: Thank you for your insightful question regarding the applicability of our pH-calibrated device to different biomarkers and the potential variations in lag times and calibration algorithms. You are correct that different biomarkers exhibit different lag times between blood and ISF due to differences in molecular transport and tissue distribution kinetics. For example, the time lag for alcohol between blood and ISF is approximately 8-12 minutes 9-10, while for caffeine, the time lag between blood and ISF concentrations is about 10 minutes 11.\nAlthough different biomarkers require distinct parameters for delay model calibration, such as diffusion coefficients and delay times, integrating the delay model with our pH- calibrated RI model remains feasible. Importantly, our pH-calibrated RI device can be broadly applicable to a range of biomarkers in ISF, including other electroosmotically extractable analytes such as alcohol, ascorbic acid, caffeine, urea, and cortisol. These uncharged substances are extracted based on electroosmotic flow and are also subject to extraction fluctuations influenced by pH. Previous studies have demonstrated the feasibility of extracting these biomarkers via RI 12-17, though few have addressed the challenge of fluctuating extraction yields. We believe our pH-calibrated method and device will advance the RI-based ISF biomarkers detection towards clinical precision. Thank you once again for your valuable comments.\nReferences:\nChanges: The following sentences were added to the “Discussion” section. For your convenience, we added the text here:\n“Additionally, our pH-calibrated RI device can be broadly applicable to a wide range of biomarkers, including other electroosmotically extractable analytes such as alcohol, ascorbic acid, caffeine, urea, and cortisol, which are also subject to extraction fluctuations influenced by pH. Therefore, the pH-calibrated RI model and device can be extended to other extractable biomarkers, enabling precise ISF-based multi- biomarkers detection.” (Lines 401-406 in the revised Manuscript)\n4. A major drawback of iontophoretic devices is the risk of local skin burns due to electric current itself, or trips in the device due to non-ideal contact with high conductance fluids for example: accidental water contact. Can you please discuss or suggest any potential safety features that can be embedded in the on-body device, which will break the circuit and arrest the device before a current overload takes place?\nResponse: Thank you for your insightful suggestion. We agree that current-induced skin burns and accidental water contact are indeed critical concerns in RI-based devices. Regarding skin burns caused by electric current, prior studies have reported the use of RI current densities ranging from 0 to 500 μA cm-2 17-20. Our previous research identified an optimized current density of 150 μA cm-2 applied for 3 minutes per extraction cycle, which minimizes the risk of skin irritation while ensuring efficient ISF extraction 20. Furthermore, the use of a Nafion coating on the sensor surface helps mitigate skin irritation by serving as a biocompatible protective layer. Across 21 in vivo trials involving 6 healthy and 15 diabetic individuals, no cases of skin redness, burns, or irritation were observed. To address potential risks from water exposure, the device is fully encapsulated with medical-grade 3M adhesive tape (2477P, USA), which adheres securely to the skin and effectively protects the sensor from accidental water contact. Under typical conditions, small amounts of water do not compromise the device’s functionality. In addition, the electrical contacts and wiring are sealed with connectors (6580-008LF, Amphenol, USA) to prevent short-circuiting caused by fluid ingress.\nNevertheless, sensor function may still be occasionally impaired by unexpected exposures, such as profuse sweating during exercise or mechanical impacts. In future iterations, circuit protection features such as real-time monitoring for abnormal current or sudden skin-electrode impedance changes, which automatically shut off power to ensure user safety, could be integrated to address these issues. This safety mechanism will be further explored in the development of our next-generation on-body system. Thank you again for your valuable comments.\nReferences:\nChanges: The following sentences were added to the “Discussion” section. For your convenience, we added the text here:\n“To ensure safety in long-term applications, future device iterations may explore methods to protect the skin under extreme pH conditions and incorporate circuit protection features such as real-time impedance monitoring, enabling automatic power cut-off in response to abnormal current or impedance changes.” (Lines 406-409 in the revised Manuscript)\n5. How did the authors arrive at the glucose range of 6 to 30 mM for the in vitro testing? Can you please provide a correlation of the tested ranges with blood glucose levels in humans? In addition, can you provide some information about the quantitation limits of the developed device?\nResponse: Thank you for your question regarding the glucose concentration range and its correlation with human blood glucose levels.\nIn healthy and diabetic individuals, blood glucose typically ranges from 3 to 30 mM 21- 23. While our goal was to cover the full 0-30 mM range, the 0 mM point represents baseline or background current rather than a true signal. Therefore, we selected five effective concentrations (6, 12, 18, 24, and 30 mM) for stepwise validation. Although this does not fully span the physiological range, the observed reduction in MAPE from 11.20% to 2.92% after pH calibration for the in vitro experiments supports the feasibility and effectiveness of the skin surface pH-calibrated strategy.\nRegarding the quantification limits of our device, in vivo results from both healthy and diabetic participants demonstrate accurate glucose prediction within the 4-30 mM range, covering a broad spectrum of physiological conditions. Theoretically, the proposed method and device are capable of detecting lower glucose concentrations. Future work will explore its applicability to more lower glucose concentrations by expanding the tested concentration range and assessing performance under relevant conditions. Thank you very much again for your valuable comments.\nReferences:\n23GIGNAC, T. et al. 1586-P: Post-OGTT Low Blood Glucose in Individuals with Metabolic Syndrome, Metabolic Phenotypes, and Metabolic-Associated Fatty Liver Disease. Diabetes 72 (2023).\nChanges: The following sentences were added to the “Discussion” section to provide a clearer understanding of the rationale behind our testing range and the quantitation limits of our device. For your convenience, we added the text here:\n“In vivo results demonstrate accurate glucose prediction within physiological range of 4-30 mM across healthy and diabetic subjects. Future work should validate this capability by extending the tested range.” (Lines 399-401 in the revised Manuscript)\n6. Did the authors make several of these devices and if yes, what was the variability observed across different devices in terms of calibration and sensitivity to skin pH changes? If some discussion on this can be added, it will improve the robustness of the proposed model/device.\nResponse: Thank you for raising this important point regarding the reproducibility and robustness of our model/device. We appreciate your attention to these critical aspects.\nThe consistency testing results of the device showed minimal variation in sensitivity across glucose (218.4±16.4 nA mM-1, CV=7.5%, Supplementary Fig. 9h), Na⁺ (52.6±6.0 mV decade-1, CV=11.3%, Supplementary Fig. 12e), and pH sensors (- 54.1±1.8 mV decade-1, CV=3.4%, Supplementary Fig. 13e). To evaluate individual-level performance, we fabricated 21 devices and conducted OGTT experiments on 6 healthy participants and 15 diabetic patients. After applying skin surface pH calibration, predicted glucose values aligned more closely with reference measurements (Fig. 5b, 5c, and Supplementary Fig. 38). Notably, the MAPE was reduced from 11.2% to 2.9% in vitro, and the MARD improved from 34.44% to 14.78% in vivo, demonstrating that the pH-calibrated model and device maintain strong applicability across individuals. Thank you very much again for your valuable suggestions.\nChanges: The following sentence was added in the “Discussion” section. For your convenience, we added the text here:\n“More refined models of blood-to-ISF biomarkers kinetics may be developed to address physiological delays and achieve personalized calibration.” (Lines 397-399 in the revised Manuscript)\nReviewer #2:\nGeneral Comments: Authors have made detailed responses to my comments but have not, in my opinion, provided compelling evidence that the addition of pH monitoring represents a true step-change in the use of reverse iontophoresis for glucose monitoring.\nResponse: We thank the reviewer for their thoughtful assessment. In this work, we sought to provide both theoretical and experimental evidence that transient pH fluctuations at the epidermal interface during RI are not merely incidental but exert a measurable influence on electroosmotic transport and glucose prediction accuracy. Our in vivo results demonstrate that integrating pH calibration markedly enhances accuracy, increasing the percentage of data points in Clarke’s zone A (healthy: 78.33% vs. 58.34% with Na⁺ calibration alone vs. 56.67% without calibration, and 61.54% with Na⁺ calibration alone reported in Sci. Adv. 7, eabd0199, 2021 1; diabetic: 67.14% vs. 38.57% with Na⁺ calibration alone vs. 25.17% without calibration) and reducing prediction errors (healthy: 10.91% vs. 20.04% with Na⁺ calibration alone vs. 25.49% without calibration; diabetic: 18.10% vs. 33.24% with Na⁺ calibration alone vs. 42.11% without calibration) — well below the FDA’s acceptance threshold (≤20%). We therefore suggest that pH monitoring addresses an important, previously unaccounted biophysical variable that can enhance the robustness and accuracy of RI-based biosensing. Specifically, the pH calibration compensates for fluctuations in glucose extraction flux caused by skin charge, while the Na+ calibration accounts for variations in glucose extraction due to skin impedance 1. These findings establish pH calibration as a critical step toward reliable and clinically meaningful RI-based biosensing.\nReference:\n1. Pu Z, et al. A thermal activated and differential self-calibrated flexible epidermal biomicrofluidic device for wearable accurate blood glucose monitoring. Science Advances 7, eabd0199 (2021).\nChanges: The following sentences were added to the “Discussion” section.\n“Nevertheless, some limitations remain. The mechanisms underlying the observed transient elevation of epidermal microfluidic pH during prolonged RI remain incompletely understood. Potential contributors include the rapid accumulation of H⁺ due to their high mobility, the release of H⁺ from biochemical reactions at the epidermal interface, and other, as-yet unidentified factors. These processes are complex and currently lack consensus in the literature, precluding a detailed mechanistic account at this stage and highlighting the need for further investigation. Additionally, the consequences of such transiently elevated pH on the local skin microenvironment, including potential irritation, need further systematic study.” (Lines 398-405 in the revised Manuscript)\n2\n1. Authors have accepted that they measure a bulk pH in the fluid on the skin surface; they do not measure the flux of H+ ions. Relatedly, the authors now state that the pH measured is the same as that in the \"epidermal electroosmotic microchannel\" but how do they know this, given that they cannot actually measure what is going on in that space?\nResponse: We respectfully clarify that while we indeed measure the bulk pH in the thin fluid layer on the skin surface, this measurement is widely accepted as a faithful proxy of the electroosmotic environment in the epidermal microchannels under an applied electric field. During RI, electroosmotic flux is restricted to these microchannels (such as pores and sweat glands) within the skin's porous structure, which serve as the transport pathways for glucose and other neutral molecules. Therefore, the pH in this interfacial layer directly reflects the environment governing analyte transport.\nThis assumption is not unique to our work but is foundational to the field of epidermal biosensing. Leading groups, including Richard H. Guy 2, John A. Rogers 3, 4, Joseph Wang 5-7, and Wei Gao 8-12, have all employed this approach to obtain physiologically relevant signals. The validity of this coupling between the skin-surface interfacial layer and the underlying microchannels is thus well established and underpins decades of peer-reviewed studies.\nReferences:\n2. Lipani L, et al. Non-invasive, transdermal, path-selective and specific glucose monitoring via a graphene-based platform. Nature Nanotechnology 13, 504-511 (2018).\n3. Cho S, et al. A skin-interfaced microfluidic platform supports dynamic sweat biochemical analysis during human exercise. Science Translational Medicine 16, eado5366 (2024).\n4. Koh A, et al. A soft, wearable microfluidic device for the capture, storage, and colorimetric sensing of sweat. Science Translational Medicine 8, 366ra165-366ra165 (2016).\n5. Ding S, et al. A fingertip-wearable microgrid system for autonomous energy management and metabolic monitoring. Nature Electronics 7, 788-799 (2024).\n6. Xu Y, et al. In-ear integrated sensor array for the continuous monitoring of brain activity and of lactate in sweat. Nature Biomedical Engineering 7, 1307-1320 (2023).\n7. Sempionatto JR, et al. An epidermal patch for the simultaneous monitoring of haemodynamic and metabolic biomarkers. Nature Biomedical Engineering 5, 737-748 (2021).\n8. Ye C, et al. A wearable aptamer nanobiosensor for non-invasive female hormone monitoring. Nature Nanotechnology 19, 330-337 (2024).\n9. Xu C, et al. A physicochemical-sensing electronic skin for stress response monitoring. Nature Electronics 7, 168-179 (2024).\n10. Tu J, et al. A wireless patch for the monitoring of C-reactive protein in sweat. Nature Biomedical Engineering 7, 1293-1306 (2023).\n11. Wang M, et al. A wearable electrochemical biosensor for the monitoring of metabolites and nutrients. Nature Biomedical Engineering 6, 1225-1235 (2022).\n12. Gao W, et al. Fully integrated wearable sensor arrays for multiplexed in situ perspiration analysis. Nature 529, 509-514 (2016).\n2. Authors also accept that they do not know the quantitative contribution that H+ makes to carrying the applied current across the skin. They acknowledge that it must small (given the very low concentration of H+ ions available to do so) and this - by extrapolation - challenges their premise that correcting for the amount of H+ around is needed.\nResponse: We agree that the quantitative contribution of H⁺ to carrying the applied current is small, given their low bulk concentration in ISF. However, our premise is not about the proportion of H⁺ flux in current transfer, but rather that the accumulation and local redistribution of H⁺ at the epidermal interface significantly alter the pH values of the microenvironment. We have replaced “H⁺ flux” with “H⁺ movement” throughout the text. Specifically, these pH fluctuations modulate the zeta potential of the epidermal electric double layer (EDL), which in turn governs the electroosmotic flux of neutral analytes such as glucose.\nThe origin of pH fluctuations during RI is likely multifactorial, involving: (i) rapid accumulation of H⁺ due to their high mobility, (ii) H⁺ release/consumption from biochemical reactions at the epidermal interface, as the reviewer also noted, and (iii) additional factors yet to be fully elucidated. While the precise contributions remain a matter of ongoing investigation in the literature, the experimental evidence in our study (Fig. 5) clearly demonstrates that these pH fluctuations are not negligible but rather exert a measurable influence on glucose extraction flux.\nMost importantly, our in-vivo and clinical results validate the necessity of correcting for pH fluctuations. Specifically, integrating pH calibration markedly enhances blood glucose prediction accuracy, increasing the percentage of data points in Clarke’s zone A (healthy: 78.33% vs. 58.34% with Na⁺ calibration alone vs. 56.67% without calibration, and 61.54% with Na⁺ calibration alone reported in Sci. Adv. 7, eabd0199, 2021 1; diabetic: 67.14% vs. 38.57% with Na⁺ calibration alone vs. 25.17% without calibration) and reducing prediction errors (healthy: 10.91% vs. 20.04% with Na⁺ calibration alone vs. 25.49% without calibration; diabetic: 18.10% vs. 33.24% with Na⁺ calibration alone vs. 42.11% without calibration) — well below the FDA’s acceptance threshold (≤20%).\nReference:\n1. Pu Z, et al. A thermal activated and differential self-calibrated flexible epidermal biomicrofluidic device for wearable accurate blood glucose monitoring. Science Advances 7, eabd0199 (2021).\n3. Authors' definition of 'skin recovery' remains vague. Imposing an electric field across a membrane drives in ions and causes membrane impedance to decrease; when the field is removed, ions will diffuse out down the created concentration gradient and impedance returns to its original value... is this recovery or just physics?\nResponse: We respectfully note that “skin recovery” is a well-established concept in transdermal research, referring not only to passive ion diffusion but also to active biological processes by which the skin restores barrier integrity and impedance after electrical or chemical perturbation. Prior studies have clearly demonstrated this phenomenon 13-15. In our work, we further describe this as “skin self-recovery,” encompassing the skin’s intrinsic ability to buffer pH fluctuations and regulate fluid loss via structural and physiological barriers such as the stratum corneum and sweat glands.\nAccordingly, our use of the term reflects both passive and active mechanisms by which the skin resists and recovers from RI-induced ISF extraction. It is therefore consistent with the literature and extends beyond a purely physical description, representing a recognized biophysical process.\nReferences:\n13. Curdy C, Kalia YN, Guy RH. Post-iontophoresis recovery of human skin impedance in vivo. European Journal of Pharmaceutics and Biopharmaceutics 53, 15-21 (2002).\n14. Manda P, Angamuthu M, Hiremath SR, Raman V, Murthy SN. Iontophoretic Drug Delivery for the Treatment of Scars. Journal of Pharmaceutical Sciences 103, 1638-1642 (2014).\n15. Abe Y, Nishizawa M. Electrical aspects of skin as a pathway to engineering skin devices. APL Bioengineering 5 (2021).\n4. Authors' responses refer repeatedly to H+ changes being due to \"extracted H+\" (which cannot be very much from a physiological fluid like the ISF that has a pH close to neutral) and \"released H+\" from 'biochemical reactions' at the epidermal interface. What are these reactions and how much H+ do they release? Can they really be sufficient to cause the local pH to increase to 10?\nResponse: We acknowledge that the precise biochemical reactions at the epidermal interface leading to H⁺ release remain incompletely understood and are the subject of ongoing debate in the literature. As noted by you, possible interfacial reactions may contribute, but their exact contributions are not yet quantitatively defined. The high mobility of H⁺ under an applied electric field further exacerbates local accumulation, which can amplify pH fluctuations.\nImportantly, our in-vivo experiments directly demonstrate that the epidermal microfluidic pH can transiently rise to ~10 during prolonged RI. These are measured facts and form the empirical basis of our study. We acknowledge two main limitations of this work. First, although we have demonstrated that skin-surface pH fluctuations significantly affect ISF extraction, the precise mechanisms underlying these fluctuations remain incompletely understood and may involve multiple biochemical and electrochemical processes, warranting further investigation. Second, while no irritation was observed in our in vivo studies, the transient rise of skin-surface pH to ~10 underscores the need for systematic evaluation of long-term biocompatibility and user safety. We have explicitly acknowledged this limitation in the revised discussion and emphasize that further mechanistic studies are warranted.\nChanges: The following sentences were added to the “Discussion” section.\n“Nevertheless, some limitations remain. The mechanisms underlying the observed transient elevation of epidermal microfluidic pH during prolonged RI remain incompletely understood. Potential contributors include the rapid accumulation of H⁺ due to their high mobility, the release of H⁺ from biochemical reactions at the epidermal interface, and other, as-yet unidentified factors. These processes are complex and currently lack consensus in the literature, precluding a detailed mechanistic account at this stage and highlighting the need for further investigation. Additionally, the consequences of such transiently elevated pH on the local skin microenvironment, including potential irritation, need further systematic study.” (Lines 398-405 in the revised Manuscript)\n5. The response about the hydrolipid film is obscure and quickly digresses to discuss the mysterious \"epidermal electroosmotic microchannel\"...\nResponse: We appreciate the reviewer’s comment and would like to clarify our intent. The hydrolipid film (HF) is a thin physiological layer on the skin surface, but it plays only a minor role in RI-driven extraction and detection. Consistent with prior studies, RI transport primarily occurs through epidermal electroosmotic microchannels, which are the established pathways for neutral molecule extraction under an applied electric field. Therefore, while the HF is a component of the skin barrier system, it is peripheral to the mechanisms investigated in this work. Our study focuses on how transient pH fluctuations at the epidermal interface modulate electroosmotic transport and extraction robustness, which we believe is the key biophysical factor influencing RI performance.\n6. Authors state without any 'back-up' or verification that any contribution to H+ presence resulting from the enzymatic reaction between glucose and glucose oxidase is small compared to H+ extraction in the EOF. Is the Gox reaction the \"biochemical reactions\" referred to in previous and other answers, or are there others (and, if yes, then what are they)? Either way, the authors need to be up-front here.\nResponse: Thanks again for your comments. As noted in Comment 2, the substantial pH fluctuations observed during prolonged RI likely arise from a combination of factors: the rapid accumulation of H⁺ due to their high mobility, H⁺ release from epidermal interfacial biochemical reactions (as the reviewer mentioned), and other yet unidentified mechanisms. The “biochemical reactions” referenced include buffering and metabolic processes at the epidermal interface (e.g., amino acids, sweat constituents, keratinocyte activity), which remain incompletely characterized. These processes are complex and remain without consensus in the literature; thus, further study is required. We have clarified this point in the revised Discussion. Importantly, our work does not depend on resolving the exact biochemical sources of H⁺. Instead, it demonstrates both theoretically and experimentally, that pH fluctuations occur, significantly impact RI extraction, and necessitate calibration.\n7. It's quite difficult to come up with an explanation for pH increasing to 10 during a short period of reverse iontophoresis without some undesirable electrochemistry occurring that results in water electrolysis...\nResponse: We acknowledge the reviewer’s concern that extreme pH shifts (up to 10) may suggest contributions from electrochemical side reactions, such as local water electrolysis, during RI. Indeed, undesirable electrochemical processes at the electrode-skin interface cannot be fully excluded. However, our in-vivo measurements consistently show transient alkaline excursions under prolonged RI, confirming that such fluctuations are a real physiological/electrochemical phenomenon rather than experimental artifact. While the precise mechanisms, whether from water electrolysis, H+ accumulation due to their high mobility, or epidermal interfacial reactions, remain incompletely resolved in the literature, our focus is on the downstream consequence: these pH deviations markedly alter the zeta potential of the epidermal EDL and thereby affect glucose extraction. We have clarified in the revised Discussion that further mechanistic work is needed, but the need for pH calibration is supported directly by experimental evidence.\n8. This reviewer remains unconvinced about the relevance of the the artificial membrane experiments.\nResponse: We appreciate the reviewer’s concern regarding the use of artificial membranes. While these membranes are not full biological equivalents to mammalian skin, they are widely used in epidermal research and transdermal diagnostics because of their structural and physicochemical stability 16-20. This stability provides a reproducible environment that enables controlled experimental validation of the theoretical framework. In contrast, in vivo and ex vivo membranes exhibit substantial variability in temperature, impedance, hydration, and ISF flow, which makes it challenging to isolate the specific impact of pH. Using the artificial membrane, we were able to experimentally validate the effect of skin surface pH on RI-driven ISF extraction, thereby reinforcing the mechanistic basis and practical rationale for incorporating pH calibration in our approach.\nReferences:\n16. Zheng L, Zhu D, Xiao Y, Zheng X, Chen P. Microneedle coupled epidermal sensor for multiplexed electrochemical detection of kidney disease biomarkers. Biosensors and Bioelectronics 237, 115506 (2023).\n17. Corsi M, et al. Bioresorbable Nanostructured Chemical Sensor for Monitoring of pH Level In Vivo. Advanced Science 9, 2202062 (2022).\n18. Stamatialis DF, Rolevink HHM, Koops GH. Controlled transport of timolol maleate through artificial membranes under passive and iontophoretic conditions. Journal of Controlled Release 81, 335-345 (2002).\n19. Haq A, Dorrani M, Goodyear B, Joshi V, Michniak-Kohn B. Membrane properties for permeability testing: Skin versus synthetic membranes. International Journal of Pharmaceutics 539, 58-64 (2018).\n20. Haq A, Goodyear B, Ameen D, Joshi V, Michniak-Kohn B. Strat-M® synthetic membrane: Permeability comparison to human cadaver skin. International Journal of Pharmaceutics 547, 432-437 (2018).\n9. It remains unclearas to whether the addition of a pH correction, in addition to that achievable with the Na+ correction, provides any real-world added-value in terms of accuracy/prediction given that - except perhaps in very extreme cases as the Authors acknowledge - the pH variation possible is really quite small and unlikely to affect EOF. New Figures R3/R4 support this point of view.\nResponse: We thank the reviewer for this comment. The Na⁺-based calibration, first introduced by Richard’s group 21, was originally aimed at reducing finger-prick frequency or enabling blood-free calibration. Subsequent work revealed that Na⁺ calibration also helps compensate for fluctuations in ISF extraction caused by changes in skin impedance 1. Building on this foundation, this work introduces an additional pH calibration step to account for variations in skin surface charge that directly affect electroosmotic transport and, consequently, glucose extraction flux. As shown in Figure R4, incorporating pH calibration on top of Na⁺ calibration yields substantial improvements in predictive accuracy. Specifically, Our in vivo results demonstrate that integrating pH calibration markedly enhances accuracy, increasing the percentage of data points in Clarke’s zone A (healthy: 78.33% vs. 58.34% with Na⁺ calibration alone vs. 56.67% without calibration, and 61.54% with Na⁺ calibration alone reported in Sci. Adv. 7, eabd0199, 2021 1; diabetic: 67.14% vs. 38.57% with Na⁺ calibration alone vs. 25.17% without calibration) and reducing prediction errors (healthy: 10.91% vs. 20.04% with Na⁺ calibration alone vs. 25.49% without calibration; diabetic: 18.10% vs. 33.24% with Na⁺ calibration alone vs. 42.11% without calibration) — well below the FDA’s acceptance threshold. These findings establish pH calibration as a critical step toward reliable and clinically meaningful RI-based biosensing.\nReferences:\n21. Sieg A, Guy RH, Delgado-Charro MB. Reverse iontophoresis for noninvasive glucose monitoring: the internal standard concept. J Pharm Sci 92, 2295-2302 (2003).\n1. Pu Z, et al. A thermal activated and differential self-calibrated flexible epidermal biomicrofluidic device for wearable accurate blood glucose monitoring. Science Advances 7, eabd0199 (2021).\n10. This reviewer's concerns about Figure 5 are not assuaged by the Authors' reply which is highly speculative and not supported by any hard facts.\nResponse: We apologize for not having stated this point clearly in our earlier response. In our in vivo study, transient elevations of skin-surface pH up to ~10 were observed during prolonged RI; importantly, no visible irritation was detected in either 6 healthy or 15 diabetic participants. We attribute this to the localized and short-lived nature of the pH fluctuations, combined with the intrinsic buffering capacity of the skin's microfluidic environment (self-recovery). Nevertheless, we acknowledge that further investigation is warranted to fully assess long-term safety. We have added this point to the Discussion section, highlighting the need for future systematic biocompatibility studies during RI.\nRegarding Figure 5, we emphasize that the conclusions are based on direct experimental measurements rather than speculation. The data from both healthy volunteers and diabetic participants consistently show that cooperative Na⁺ and pH calibration yields significant improvements in prediction accuracy.\n11. The term \"reverse ions\" is unusual. I think the Authors mean 'counterions', i.e., ions with a charge opposite to that of the skin.\nResponse: Thank you for pointing this out. Indeed, the term \"reverse ions\" is synonymous with \"counterions,\" which refers to ions with a charge opposite to that of the skin. We have replaced \"reverse ions\" with \"counterions\" throughout the manuscript. Thank you for your attention to this detail.\n12. I think that maintaining a pH of 10 on the skin for very long will typically cause visible irritation.\nResponse: We acknowledge that sustained high skin pH may cause irritation. In our study, while the epidermal microfluidic pH transiently reached ~10 during RI, no adverse reactions were observed in either 6 healthy or 15 diabetic participants. We attribute this to the transient nature of the pH change, its confinement within a localized microenvironment, and the skin’s intrinsic buffering capacity. Safety and comfort remain priorities, and we will continue to monitor these parameters in future studies. We also acknowledge this part in the “Discussion” section."}]

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