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. 2026 Jun 22;11(26):39445–39455. doi: 10.1021/acsomega.6c05341

Long-Term Stable Wearable Electrochemical Aptamer Sensor for Cortisol in Artificial Sweat Based on a Gold Nanoparticle–Carboxymethyl Cellulose–Methylene Blue Interface

Suparat Cotchim †,‡,§, Supatinee Kongkaew †,‡,§, Yudtapum Thipwimonmas §,∥, Warakorn Limbut †,‡,§,⊥,*
PMCID: PMC13347639  PMID: 42428912

Abstract

A wearable electrochemical aptamer sensor for cortisol was developed on a screen-printed electrode bearing a gold nanoparticle–carboxymethyl cellulose–methylene blue interface. The redox layer was formed by on-electrode electropolymerization of methylene blue in the presence of gold nanoparticles and carboxymethyl cellulose, which helps stabilize the film and supports Au–S anchoring of the thiolated aptamer. Analytical cortisol measurements were performed in artificial sweat with controlled ionic strength and pH. The sensor provided a limit of detection of 0.09 pg mL–1 and a linear range of 0.0001–1000 ng mL–1 (mean ± SD, n ≥ 3). Interference from common sweat constituents was minimal, and batch-to-batch reproducibility was good. The device retained 93.4% of its initial response after 56 days of storage at 4 °C (RSD = 2.4%). These results demonstrate a stable wearable transducer platform for cortisol sensing in artificial sweat, and validation in real human sweat will be pursued in future work.

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Introduction

Cortisol is a glucocorticoid hormone secreted by the adrenal cortex under regulation of the hypothalamic–pituitary–adrenal (HPA) axis and is widely used as a biomarker of physiological stress. Typically, the cortisol level in biofluids fluctuates in response to the circadian rhythm throughout the day, with the highest levels occurring in the early morning and the lowest at night. In addition to circadian regulation, cortisol secretion is also influenced by various psychological and physiological stressors. Physical stress, such as the intense and prolonged nature of athletic training and competition, can trigger a complex hormonal response involving cortisol. Understanding the role of cortisol in training adaptation is crucial for sports and health professionals, as it helps prevent overtraining and supports the long-term health and performance of athletes.

Conventional cortisol assays, including enzyme-linked immunosorbent assay (ELISA), , chemiluminescent immunoassay (CLIA), , and liquid chromatography-tandem mass spectrometry (LC–MS/MS), , provide accurate quantification in serum and plasma but require laboratory instruments and trained personnel, making them unsuitable for point-of-care (POC) use. Sweat is a noninvasive biofluid that can be collected repeatedly and safely from eccrine glands without pain or infection risk. Reported sweat cortisol levels span ∼7.9–151 ng mL–1, ,− suggesting potential for stress monitoring. In this context, electrochemical biosensing has gained interest because of its low cost, rapid response, high sensitivity, small sample-volume requirements, and compatibility with miniaturized, wearable formats. Recent advances in sweat-based electrochemical biosensors further support prospects for continuous, real-time monitoring. ,−

Advances in electrochemical sensing materials have accelerated the development of intelligent biosensing systems, wearable platforms, and micro- and nanointegrated analytical devices. Recent progress includes integrating flexible sensors with microfluidic systems and miniaturized electronics to enable continuous, real-time monitoring in complex biological environments. , These systems offer high sensitivity, portability, user comfort, and data-driven analysis, supporting personalized health monitoring and point-of-care applications. Wearable biosensors are particularly valuable for noninvasive, dynamic biochemical monitoring. Micro- and nanoscale engineering allows precise control of interfacial properties and enhance signal transduction efficiency. Developing stable, adaptable sensing interfaces remains critical for linking material performance with system-level integration.

DNA aptamers, single-stranded oligonucleotides capable of selectively binding to specific targets with high affinity, have shown great potential as biorecognition elements. They offer high stability, low cost, and ease of synthesis and modification, making them suitable for biomedical sensing applications. Nanomaterials-assisted electrode modification further improves the sensitivity and selectivity of biosensors. Gold nanoparticles (AuNPs) are widely used due to their excellent electrical conductivity, large surface area, and strong affinity to biomolecules such as aptamers. Methylene blue (MB), a redox-active mediator, facilitates electrochemical signal generation because of its stable redox behavior and ability to intercalate into nucleic acids. Carboxymethyl cellulose (CMC), a water-soluble biopolymer, acts as a biocompatible linker and stabilizing matrix that enhances the uniform dispersion of AuNPs and protects MB on the electrode surface. , Its carboxylate groups retain cationic MB/PMB electrostatically, help disperse AuNPs, and form a hydrophilic film that limits fouling and supports ion transport. CMC is also compatible with water-based screen printing on the wearable substrate and, in this work, yielded more reproducible films than neutral polymers such as poly­(vinyl alcohol) (PVA) and polyethylene glycol (PEG).

Recent research has focused on developing advanced sensitive materials for electrochemical biosensing. Current strategies emphasize integrating nanostructured materials, conductive polymers, and hybrid nanocomposites to improve sensitivity, stability, and antifouling properties. Nanomaterial-based electrodes are frequently used to increase surface area and facilitate electron transfer, which enhances detection limits and response. Additionally, mechanical interfaces and molecular design are important for achieving high selectivity and robustness in complex biological environments. Functional polymer matrices and hybrid sensing platforms are also under investigation to improve film stability, biocompatibility, and resistance to biofouling. , The integration of gold nanoparticles with polymeric matrices and redox-active species is a promising approach for constructing stable and high-performance electrochemical sensing interfaces.

In this work, we developed a wearable DNA aptamer sensor based on an AuNPs–CMC–MB modified wearable screen-printed electrode (WSPE) for noninvasive cortisol sensing via sweat. Our aim is to establish a stable redox and aptamer-anchoring interface and to evaluate transducer-level performance in artificial sweat under controlled composition. On-body studies and clinical validation will be pursued in future work.

Materials and Methods

Materials

The sequence of the aptamer consisted of an 85-mer cortisol-binding ssDNA aptamer (5′ thiol-C6GGAATGGATCCACATCCATGGATGGGCAATGCGGGGTGGAGAATGGTTGC CGCACTTCGGCTTCACTGCAGACTTGACGAAGCTT-3′) and was purchased from Bio Basic Inc. (Markham, Canada). Hydrocortisone (≥98% purity), gold­(III) chloride trihydrate, 6-mercapto-1-hexanol (MCH), ascorbic acid, uric acid, d-glucose, urea, sodium bicarbonate, monosodium dihydrogen orthophosphate, sodium chloride, potassium chloride, sodium citrate, methylene blue were purchased from Sigma-Aldrich. Lactic acid was purchased from LOBA Chemie. The adhesive tape (Fuh Shuen brand) and polypropylene sheets were obtained from a local stationery store (Hatyai, Songkhla, Thailand). Conductive inks were obtained from Gwent Electronic Materials Ltd. (Gwent, UK). All solutions used in the procedures were prepared with deionized water (resistivity ≥18.2 MΩ cm–1) obtained from a Milli-Q water purification system (Millipore, Bedford, USA).

Instruments

The morphology of synthesis AuNPs was evaluated by field emission transmission electron microscope (FE-TEM) (Talos F200i, Thermo Scientific, Czech Republic) and UV–vis spectrophotometer (T60UV, PG Instruments Limited, United Kingdom). The morphology of AuNPs–CMC–MB was evaluated by scanning electron microscopy (SEM) with energy dispersive X-ray spectroscopy (EDX) (Apreo, FEI, Netherlands). All electrochemical investigations, including cyclic voltammetry (CV), electrochemical impedance spectroscopy (EIS), differential pulse voltammetry (DPV), and chronoamperometry, were performed using PalmSens4 potentiostat-galvanostat with PSTrace software (PalmSens BV, Houten, Netherlands).

Preparation of Artificial Sweat

Artificial sweat solution was prepared following a previously reported method. Briefly, sodium bicarbonate (0.0630 g), sodium dihydrogen phosphate (0.0121 g), sodium chloride (0.3587 g), potassium chloride (0.1138 g), and urea (0.24 g) were dissolved in 150 mL of deionized water. Then, 350 μL of lactic acid and 80 μL of ammonia were added to the above solution and transferred to a 250 mL volumetric flask. The prepared artificial sweat solution was used in all subsequent experiments.

Synthesis of AuNPs

Before synthesis, the glassware was washed with a freshly prepared aqua regia acid solution, a mixture of HCl and HNO3 with a molar ratio of 3:1, to prevent contamination. The AuNPs were synthesized using the Turkevich method with slight modification. , In this synthesis, sodium citrate was used to reduce gold chloride trihydrate solution (HAuCl4·3H2O) and as a stabilizing agent for the nanoparticles. First, 2.2 mmol L–1 of sodium citrate was heated at 70 °C. Then, 25.0 mmol L–1 of HAuCl4 was suddenly added and vigorously stirred under constant conditions for 30 min. The solution turned from transparent to a light purple, indicating that the reduction of gold ions had entirely occurred to nanoparticles. The final AuNP dispersion was cooled to room temperature and stored at 4 °C until use.

Fabrication of Wearable Screen-Printed Electrode (WSPE)

The WSPE was fabricated using a cutting machine (Silhouette CAMEO, version 3.2). The WSPE pattern was designed using the Inkscape program and transferred onto an adhesive tape, which was affixed to a polypropylene (PP) plastic sheet with two sides as a mask for electrode printing. The substance was punched at the connecting area to ensure the two sides were connected. The graphene ink was then applied to one side of the substrate using a squeegee and heated at 80 °C for 20 min to evaporate the solvent. This process was repeated for the other side. Finally, the mask was removed from two sides, and a pseudoreference electrode was created by painting Ag/AgCl ink onto the reference electrode area and drying it at 80 °C for 5 min (Figure S1).

Fabrication of AuNPs-CMC–MB Modified on WSPE

A wearable DNA biosensor for cortisol in artificial sweat was fabricated via stepwise surface modification on WSPE coated with AuNPs–CMC–MB particles (Figure ). First, a WSPE was cleaned to remove impurities using water and then cycled in an artificial sweat solution using CV, with a range of −0.5 to 0.3 V at a scan rate of 0.1 V s–1 for 50 cycles. Afterward, a mixture of AuNPs, CMC, and MB was drop-cast onto the electrode surface and electropolymerized by cyclic voltammetry in the potential range of +1.2 to −0.9 V at a scan rate of 0.05 V s–1. This process formed a poly­(methylene blue) (PMB) film in the presence of AuNPs–CMC, yielding a stable and conductive polymer matrix. The concentration of CMC, MB, and the number of electropolymerization cycles were optimized. The modified WSPE was rinsed with artificial sweat to remove any unreacted monomer.

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Illustration shows the stepwise fabrication of the AuNPs–CMC–MB modified WSPE and the immobilization of the DNA aptamer for cortisol detection.

It should be noted that the AuNPs were uniformly distributed in the CMC/MB monomer solution. During polymerization, the AuNPs were embedded in the growing PMB matrix, which served as a binder and stabilizer to prevent aggregation and improve film stability.

Subsequently, 1.0 μL of cortisol-binding aptamer (optimized concentration) was immobilized on the surface of AuNPs–CMC–MB/WSPE and incubated at 4 °C for 12 h. The aptamer molecules were covalently attached to AuNPs via thiol–gold bonds. The electrode was rinsed with deionized water and incubated in 2.0 mmol L–1 MCH for 30 min at room temperature to block nonspecific binding sites. The resulting wearable DNA aptamer sensor was stored at 4 °C before use.

Results and Discussion

Design and Fabrication of a Wearable Screen-Printed Electrode

A wearable screen-printed electrode (WSPE) and the ports of each electrode were designed to be compatible with the Sensit Wearable system (PalmSens, Netherlands). This work produced five electrode patterns, namely P1, P2, P3, P4, and P5, as shown in Figure A. The physical structure, including the working electrode diameter of each pattern, was fixed. The electrochemical behavior of these five electrode patterns was investigated using cyclic voltammetry (CV) in a 5.0 mmol L–1 [Fe­(CN)6]3–/4– containing 0.10 mol L–1 of KCl (Figure B). The performance of the fabricated WSPEs was compared with that of a commercial screen-printed electrode (SPE) under identical conditions.

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(A) Template pattern of wearable sensing. (B) CV of five electrode template patterns (P1, P2, P3, P4, and P5) and commercial SPE. (C) The histogram of oxidation current and the reduction current in 5.0 mmol L–1 [Fe­(CN)6]3–/4– containing 0.10 mol L–1 of KCl solution. (D) The reproducibility of the preparation of wearable electrode.

In the P1 design, the electrode consisted of only two electrodes, namely a working electrode (WE) and a reference electrode (RE). The redox peak current obtained from P1 was lower than those from the other electrode patterns employing a three-electrode configuration that included a counter electrode (CE) in addition to the WE and RE. In a two-electrode system, the RE must simultaneously maintain the electrode potential and carry the current generated during the redox process. This dual function can lead to potential instability and reduced electrochemical performance, resulting in lower peak currents. In contrast, in a three-electrode configuration, the CE is responsible for carrying the current generated during the electrochemical reaction, while the RE maintains a stable potential at the WE. Consequently, the electron-transfer process becomes more stable and produces higher redox peak currents.

For the P2, P3, and P4 designs, the area of the RE was progressively increased, whereas the area of the CE was reduced. This modification was intended to investigate the influence of electrode geometry on electrochemical performance. In particular, P3 possessed a smaller CE area relative to P2, resulting in increased current density at the CE surface. The reduced CE area generated higher overpotential and increased electrolyte resistance, thereby decreasing electron-transfer efficiency and overall electrochemical stability. Furthermore, when the CE area became too small, the CE could not effectively support the required current during the redox process, which may have caused partial current leakage through the RE and reduced the reproducibility of the electrochemical response.

In contrast, the P5 design employed larger connection tracks and an optimized three-electrode layout, which reduced the overall electrical resistance of the system and minimized the ohmic drop between the electrode and the potentiostat. As a result, P5 exhibited the highest oxidation and reduction peak currents together with the smallest peak-to-peak separation (ΔE), indicating superior electron-transfer efficiency and electrochemical reversibility (Figure C). In addition, the electrochemical performance of the SPE pattern was comparable to that of the optimized P5 design. The SPE exhibited high oxidation and reduction peak currents with a relatively small ΔE value, confirming efficient electron-transfer behavior. However, the P5 configuration produced slightly higher peak currents and lower ΔE values than the SPE pattern, indicating that the optimized wearable electrode geometry provided improved electrochemical performance while maintaining a design compatible with wearable sensing applications. Consequently, the P5 electrode pattern was selected as the optimized template for developing the wearable cortisol biosensor in this study.

The reproducibility of the fabricated WSPE was also examined under the same redox conditions (5.0 mmol L–1 [Fe­(CN)6]3–/4– in 0.10 mol L–1 KCl) using 24 electrodes. The WSPE demonstrated excellent reproducibility, with relative standard deviations (RSDs) of 4.8% for the oxidation peak current and 4.3% for the reduction peak current (Figure D). These results confirm that the fabricated WSPEs possess high reproducibility, stable signal response, and strong potential for wearable biosensor applications, including DNA-based cortisol detection.

Characterization of Synthesized AuNPs

The colloidal AuNPs were synthesized by chemical reduction of HAuCl4 using sodium citrate. The light-purple color of the final product indicates the formation of AuNPs. The synthesis of AuNPs was characterized using UV–vis spectrophotometry and field emission transmission electron microscopy (FETEM). The UV–vis spectrum of the AuNPs exhibited a maximum absorbance peak at 529 nm, confirming the presence of well-dispersed nanoparticles with a symmetric spherical shape, as shown in Figure A. Based on the Beer–Lambert law (eq ), the final AuNP concentration was estimated to be 1.08 nM.

A=εbC 1

where A = absorbance, ε = molar absorptivity, b = length of light path, and C = concentration.

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Characterization of synthesized AuNPs. (A) UV–vis spectrum and photograph of the AuNP solution. (B) Particle size distribution histogram from TEM analysis. (C) TEM image showing uniform spherical AuNPs. (D) SAED diffraction pattern, and (E) the HRTEM image of AuNP.

Moreover, the AuNP suspension exhibited a pinkish-red color observable to the naked eye (inset of Figure A), consistent with typical optical characteristics of gold nanoparticles. The FETEM images revealed that the AuNPs possessed a spherical morphology with uniform distribution (Figure C). The particle size histogram indicated an average diameter of 12 ± 1 nm (Figure B). To further confirm the crystalline structure, high-resolution TEM (HRTEM) and selected area electron diffraction (SAED) analyses were conducted. The SAED pattern displayed distinct ring patterns corresponding to the (111), (200), (220), and (311) planes of face-centered cubic (fcc) gold, as shown in Figure D. The HRTEM image (Figure E) showed well-defined lattice fringes with a spacing of 0.23 nm, corresponding to the (111) plane of gold. These results verify that the synthesized AuNPs possess high crystallinity, uniform size distribution, and structural stability, confirming their suitability for use as a conductive nanomaterial in the biosensor fabrication process.

Characterization of WSPE and AuNPs–CMC–MB Modified on WSPE

The morphology of the wearable screen-printed electrode (WSPE) and the AuNPs–CMC–MB modified WSPE was analyzed using scanning electron microscopy (SEM) coupled with energy-dispersive X-ray spectroscopy (EDX). Figure S2A shows the surface morphology of the bare WSPE, which exhibits a relatively flat structure with slight roughness. After modification with AuNPs–CMC–MB, the surface became more textured and displayed granular-like nanoparticle features, as shown in Figure S2B. It was observed that the AuNPs were either uniformly distributed or partially embedded within the CMC–MB polymer matrix. The flat graphene regions were noticeably reduced, indicating that a polymeric coating layer had formed on the electrode surface.

The EDX spectra and elemental mapping confirmed the successful surface modification (Figure S2C,D). The distinct Au peaks in the EDX spectra, together with the bright spots corresponding to Au in the elemental mapping, provide clear evidence of the successful deposition of AuNPs onto the electrode surface. These results collectively verify the effective coating and uniform distribution of AuNPs–CMC–MB on the WSPE, which contributes to improved surface roughness and enhanced electrochemical activity.

Electrochemical Characterization

The electrochemical properties of the Bare WSPE, AuNPs–CMC–MB, Ap/AuNPs–CMC–MB, MCH/Ap/AuNPs–CMC–MB, and cortisol/MCH/Ap/AuNPs–CMC–MB modified on WSPE were first evaluated using cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS) to verify each electrode modification step. For CV, the electrochemical response of methylene blue (MB) during each modification step was recorded in an artificial sweat solution within the potential range of −0.5 to 0.3 V at a scan rate of 0.05 V s–1. Figure S3A,B present the cyclic voltammograms of the different electrode configurations. The bare WSPE exhibited no apparent redox peaks, confirming the absence of MB on the electrode surface. After modification with AuNPs–CMC–MB, a well-defined redox couple appeared, indicating that MB was successfully deposited and that the AuNPs–CMC–MB layer enhanced the electrode conductivity. Upon immobilization of the aptamer on the AuNPs–CMC–MB-modified electrode, the redox peak current of MB decreased, suggesting a hindered electron transfer process due to the formation of the DNA layer on the surface. These results confirmed that the aptamer molecules were successfully immobilized onto the AuNPs–CMC–MB surface. Subsequent treatment with MCH further reduced the peak current, consistent with the blocking of nonspecific binding sites. Finally, after cortisol binding, the current response of MB decreased further, confirming that cortisol molecules were effectively captured on the electrode surface, thereby increasing the interfacial resistance and restricting electron transfer at the modified WSPE. These electrochemical changes collectively confirm the successful stepwise surface modification and construction of the sensing interface.

The electron-transfer behavior was further investigated by EIS measurements performed in a 5.0 mmol L–1 [Fe­(CN)6]3–/4– containing 0.10 mol L–1 KCl solution. The experiments were carried out over a frequency range of 100 kHz to 0.1 Hz with an amplitude of 0.02 V. The resulting Nyquist plots (Figure S3C,D) consist of a semicircular region, representing the electron-transfer-limited process, followed by a linear diffusion-controlled region. The inset in Figure S3C,D shows the equivalent circuit model used to interpret the impedance behavior of the modified electrodes. The diameter of the semicircle corresponds to the charge-transfer resistance (R ct) at the electrode–electrolyte interface. The bare WSPE displayed the largest semicircle with an R ct value of 4180 ± 3 Ω. After modification with AuNPs–CMC–MB, the R ct value decreased drastically to 544.6 ± 0.5 Ω, indicating that the nanocomposite layer significantly improved the electron-transfer efficiency at the electrode surface. When the aptamer was introduced, the R ct value increased to 769.4 ± 0.7 Ω, confirming that the aptamer molecules were successfully immobilized on the electrode surface, leading to partial obstruction of the electron pathway. After subsequent modifications with MCH and cortisol, the R ct values further increased to 2072 ± 29 Ω and 2194 ± 3 Ω, respectively. These findings confirm that both MCH and cortisol were successfully attached to the electrode surface, each contributing to additional resistance, thereby verifying the successful stepwise fabrication of the aptamer sensor. These EIS results are consistent with the corresponding CV observations and collectively confirm the successful stepwise construction of the sensing interface.

Electrochemical Behavior and Effective Surface Area of MCH/Ap/AuNPs–CMC–MB Modified on WSPE

The electrochemical behavior of MCH/Ap/AuNPs–CMC–MB modified on WSPE was evaluated using CV in [Fe­(CN6)3–/4–] at different scan rates. The redox peak current of [Fe­(CN6)3–/4–] gradually increased as the scan rate increased from 20 to 200 mV s–1, as shown in Figure A.

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(A) Cyclic voltammograms of the MCH/Ap/AuNPs–CMC–MB/WSPE recorded at scan rates of 20–200 mV s–1. (B) Dependence of anodic (I pa) and cathodic (I pc) peak currents on scan rate. (C) Plot of I p versus the square root of scan rate. (D) Linear relationship between log I p and log ν, confirming a diffusion-controlled process.

Both the anodic and cathodic peak currents displayed a linear relationship with the square root of the scan rate (Figure C), compared to their direct dependence on scan rate (Figure B). This observation indicates that the electrochemical reaction on the MCH/Ap/AuNPs–CMC–MB/WSPE surface is diffusion-controlled. The kinetic behavior was further analyzed by plotting log I p versus log ν, which yielded a linear regression equation of log I p = 0.374 ± 0.004 log ν + 1.1884 ± 0.007 (R 2 = 0.9992, n = 3) (Figure D). The slope of approximately 0.37 is close to the theoretical value of 0.5, confirming that the redox process on the modified WSPE follows a diffusion-controlled mechanism.

The effective electroactive surface area of the MCH/Ap/AuNPs–CMC–MB/WSPE was estimated using the Randles–Ševčík eq (eq ).

Ip=2.69×105n3/2AD1/2Cν1/2 2

where I p is the oxidation current peak current, n is the number of electrons transferred in the redox reaction (n = 1), A is the electrode’s effective surface area (cm2), D is the diffusion coefficient of [Fe­(CN6)3–/4–] solution, C is the concentration of the ferricyanide solution and ν is the scan rate (V s–1). The calculated effective surface area of the MCH/Ap/AuNPs–CMC–MB/WSPE was 0.075 cm2, which is significantly larger than the geometric area (0.035 cm2), demonstrating the enhanced electroactive surface and improved electron transfer kinetics resulting from the AuNPs–CMC–MB modification.

Optimization

Concentration of Carboxymethyl Cellulose (CMC)

In this work, CMC was employed as a biocompatible linker with MB and as a stabilizing and dispersing agent for AuNPs. Therefore, the effect of CMC concentration was investigated at 0, 0.50, 0.75, 1.00, and 1.50% (w/v) in artificial sweat (Figure A). The peak current of MB increased as the CMC concentration increased from 0 to 1.0%. This enhancement is attributed to the electrostatic interaction between the anionic polymer (CMC) and cationic MB molecules, which facilitates the formation of a uniform and stable protective layer on the electrode surface, thereby improving MB stability and electron transfer. However, when the CMC concentration exceeded 1.0%, the MB peak current decreased because the polymer film became thicker, hindering electron transfer between MB and the electrode surface. Therefore, 1.0% CMC was selected as the optimum concentration for subsequent experiments.

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(A) Effect of CMC concentration, (B) effect of MB concentration, (C) effect of electropolymerization cycles, and (D) effect of aptamer concentration on the electrochemical signal in sweat solution.

Concentration of Methylene Blue (MB)

The electrochemical signal of the WSPE originated from the redox activity of MB, which acts as a redox mediator. The concentration of MB was optimized at 0.10, 0.25, 0.50, and 1.00 mmol L–1 in sweat solution (Figure B). The peak current increased with increasing MB concentration up to 0.50 mmol L–1, indicating that more MB monomers were available for electropolymerization and electron transfer. At concentrations higher than 0.50 mmol L–1, the MB film became too dense, reducing electron mobility and partially blocking active sites on the electrode surface. Therefore, 0.50 mmol L–1 of MB was chosen as the optimal concentration for subsequent experiments.

Electropolymerization Cycle of MB

The number of electropolymerization cycles during MB deposition on the electrode was optimized at 5, 10, 15, and 20 cycles in sweat solution (Figure C). The MB peak current increased progressively from 5 to 15 cycles, indicating that each cycle deposited additional MB film and enhanced the electrochemical activity of the electrode. Beyond 15 cycles, however, the MB peak current declined because the excessively thick polymer film caused electron-transfer hindrance, as many active sites became buried within the matrix. Thus, 15 electropolymerization cycles were selected as the optimum condition to ensure a stable and highly conductive MB film on the WSPE surface.

Concentration of Aptamer

The concentration of the cortisol-specific aptamer was optimized at 2.0, 4.0, 6.0, and 8.0 μmol L–1 in sweat solution (Figure D). When the aptamer concentration during immobilization increased from 2.0 to 4.0 μmol L–1, the ΔI (difference in current response) increased significantly. This result suggests that more aptamer molecules were successfully conjugated to the AuNPs surface via thiol–Au covalent interactions, providing a higher density of recognition elements for cortisol binding. When the aptamer concentration exceeded 4.0 μmol L–1, the ΔI value decreased. At high aptamer densities, the immobilized strands became electrostatically repulsive and sterically hindered, adopting stretched or unfolded conformations that reduced their ability to bind cortisol efficiently. Therefore, 4.0 μmol L–1 of aptamer was chosen as the optimal concentration for subsequent experiments.

Analytical Performance

Linear Range, Limit of Detection, and Limit of Quantification

The analytical performance of the wearable electrochemical biosensor for cortisol detection was evaluated by chronoamperometry in artificial sweat. The measurements were conducted at 100 s intervals with varying cortisol concentrations (Figure A). As the cortisol concentration increased, the cathodic current became less negative because cortisol binding hindered electron transfer. Consequently, the baseline-referenced current change, ΔI, increased with cortisol concentration. The calibration curve between ΔI and cortisol concentration from 0.0001 to 1000 ng mL–1 displayed an excellent linear relationship with a correlation coefficient (R 2 = 0.9994) (Figure B). The limit of detection (LOD) and limit of quantification (LOQ) were calculated using the formulas 3 × (Sa/b) for LOD and 10 × (Sa/b) for LOQ, where Sa is the standard deviation of the blank signal and b is the slope of the calibration curve. The LOD and LOQ of the wearable aptamer sensor were determined to be 0.09 pg mL–1 and 0.3 pg mL–1, respectively. These outstanding analytical figures confirm the high sensitivity of the proposed wearable aptamer sensor for cortisol detection.

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(A) Chronoamperometric responses at different cortisol concentrations (0–1000 ng mL–1) in artificial sweat at a fixed applied potential of –0.071 V and pH 5.0. (B) Calibration plot of ΔI versus cortisol concentration. (C) Reproducibility using six WSPEs at 10.0 ng mL–1 cortisol. (D) Repeatability of a single WSPE over 12 measurements. (E) Storage stability of WSPEs over 56 days. (F) Selectivity of the biosensor in the presence of possible interfering substances.

The AuNPs–CMC–MB-modified wearable aptamer sensor compares favorably with recent electrochemical and wearable cortisol sensors. Prior reports, including CNT-based flexible electrodes, Bio-inspired MIP sensor, Ni–Co–MOF, Microfluidic Patch, MOF-assisted MIP systems, Wearable MIP Sensor, and both NFC-enabled devices and aptamer-FET (Table S1), typically achieve LODs in the pg mL–1 to ng mL–1 range and often involve complex fabrication. Our device offers pg mL–1-level LOD, wide linear range, and 56 day storage stability, and good batch-to-batch reproducibility, while using a wearable screen-printed electrode and on-electrode electropolymerization to form the redox layer. These analytical characteristics are attributed to the synergistic combination of AuNPs for Au–S aptamer immobilization, CMC as a stabilizing biopolymer matrix, and MB as a redox mediator within a printable and reproducible sensing interface.

Reproducibility, Repeatability, Storage Stability, and Selectivity

The reproducibility of the biosensor was examined using six independently prepared WSPEs tested with 10.0 ng mL–1 cortisol in artificial sweat. The relative standard deviation (RSD) of the responses was 0.30%, which is significantly lower than the acceptable threshold according to AOAC guidelines, demonstrating the excellent reproducibility of the fabricated electrodes (Figure C). The repeatability of a single WSPE was assessed by performing 12 consecutive measurements of 10.0 ng mL–1 cortisol under identical conditions. The RSD value was 0.35%, also well within AOAC limits, indicating high repeatability of the proposed sensor (Figure D).

The long-term stability of the WSPEs was tested by storing the electrodes at 4 °C and periodically evaluating their responses over 56 days. The electrodes retained 93.4% of their initial response, with an RSD of 2.42% after 56 days, confirming excellent storage stability and suitability for long-term applications (Figure E). This stability likely arises from the CMC–PMB/AuNP film, which provides a hydrophilic, carboxylate-rich matrix that retains MB/PMB and disperses AuNPs into a continuous layer. The AuNPs also furnish Au–S anchoring sites for the thiolated aptamer. These results suggest that the observed stability mainly arises from film retention and stable aptamer anchoring rather than loss of redox species or film delamination. The slight decrease in current response observed during long-term storage may be associated with gradual conformational changes of the immobilized aptamer and minor loss of electrochemical activity within the MB/PMB redox layer during long-term storage. Nevertheless, the decrease remained limited throughout the storage period, indicating good long-term stability of the sensing interface. The selectivity of the biosensor was investigated by testing potential interferents, including ascorbic acid (100.0 μmol L–1), uric acid (50.0 μmol L–1), glucose (1.0 mmol L–1), lactate (25.0 mmol L–1), and urea (50.0 mmol L–1), both individually and in mixtures with 10.0 ng mL–1 cortisol. The observed current variations were less than 5%, and no statistically significant difference (p > 0.05) was observed between the cortisol-only and mixed samples (Figure F). These results clearly demonstrate the high selectivity of the developed wearable electrochemical aptamer sensor toward cortisol detection.

Cortisol Detection in Artificial Sweat (Spike-Recovery)

Cortisol standards were spiked into artificial sweat at low (1.0, 5.0, and 10.0 ng mL–1), medium (50 and 100 ng mL–1), and high (200 ng mL–1) levels (samples 1–6) and analyzed under fixed ionic strength, pH, and volume. Currents were processed as baseline-referenced signals (ΔI = I – I blank) using electrode-matched replicates. As summarized in Table , the sensor gave recoveries of 91 ± 1% to 98.4 ± 0.5% with RSD = 0.39–1.36% (mean ± SD, n ≥ 3), indicating good accuracy and precision in this sweat matrix. Apparent concentrations tracked the spikes with minimal bias across the range. These results support reliable cortisol quantification in artificial sweat using the CMC-assisted PMB/AuNP wearable electrode and provide transducer-level evidence for subsequent on-body validation. Ratio-type indices and multianalyte panels for physiological interpretation will be explored in future work when on-body testing becomes available.

1. Cortisol Detection in Artificial Sweat with Recovery and RSD Values for Each Sample.
  concentration (ng mL–1)
 accuracy and precision
samples spiked detected recovery (%) RSD (%)
artificial sweat (#1) 1.0 0.91 ± 0.01 91 ± 1 1.26
artificial sweat (#2) 5.0 4.67 ± 0.04 93.4 ± 0.9 0.92
artificial sweat (#3) 10.0 9.81 ± 0.06 98.1 ± 0.6 0.59
artificial sweat (#4) 50.0 48.9 ± 0.7 98 ± 1 1.36
artificial sweat (#5) 100.0 92.7 ± 0.4 92.7 ± 0.4 0.39
artificial sweat (#6) 200.0 197 ± 1 98.4 ± 0.5 0.55
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Conclusions

A wearable electrochemical aptamer sensor for cortisol was developed on a gold nanoparticle–carboxymethyl cellulose–methylene blue (AuNPs–CMC–MB)–modified screen-printed electrode. The AuNPs–CMC–MB interface provided a stable redox layer and supports Au–S anchoring of the thiolated aptamer. Under controlled artificial-sweat conditions, the sensor achieved an LOD of 0.09 pg mL–1 with a linear range of 0.0001–1000 ng mL–1, along with good repeatability, reproducibility, and minimal interference from common sweat constituents. The device retained 93.4% of its initial response after 56 days of storage at 4 °C, indicating good storage stability of the sensing interface. These results support the AuNPs–CMC–MB platform as a stable wearable transducer for cortisol sensing in artificial sweat. Mechanical stability under dynamic strain conditions, together with real human sweat analysis and on-body evaluation, will be investigated in future studies.

Supplementary Material

ao6c05341_si_001.pdf (639.2KB, pdf)

Acknowledgments

This work was financially supported by a Postdoctoral Fellowship from Prince of Songkla University. The authors gratefully acknowledge the Center of Excellence for Trace Analysis and Biosensors (TAB-CoE), the Talent Management Project, the Forensic Science Innovation and Service Center, Center of Excellence for Innovation in Chemistry (PERCH-CIC), the Division of Health and Applied Sciences, the Division of Physical Science, Faculty of Science, Prince of Songkla University, Hat Yai, Songkhla, Thailand for instruments, apparatus, and financial support. This research and innovation activity is funded by the National Research Council of Thailand (NRCT).

The Supporting Information is available free of charge at https://pubs.acs.org/doi/10.1021/acsomega.6c05341.

  • Schematic illustration of the fabrication process of WSPE (Figure S1), SEM images of (A) bare wearable screen-printed electrode, (B) AuNPs–CMC–MB modified on screen-printed electrode. The EDS spectrum and EDS mapping of (C) bare wearable screen-printed electrode, and (D) AuNPs–CMC–MB modified on screen-printed electrode (Figure S2), electrochemical characterization of the modified electrodes. (A,B) cyclic voltammograms of different modification stages recorded in artificial sweat. (C,D) EIS Nyquist plots for each electrode stage recorded in 5.0 mmol L–1 [Fe­(CN)6]3–/4– containing 0.10 mol L–1 KCl. The inset shows the equivalent circuit model used for interpretation of the impedance behavior (Figure S3), and overview of analytical performance of selected cortisol sensors (Table S1) (PDF)

The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. S.C.: conceptualization, methodology, investigation, formal analysis, validation, data curation, writingoriginal draft, writingreview and editing. S.K.: methodology, formal analysis, validation, writingreview and editing. Y.T.: methodology, formal analysis, validation, writingreview and editing. W.L.: conceptualization, methodology, investigation, formal analysis, resources, data curation, writing-review and editing, supervision, project administration, and funding acquisition.

The authors declare no competing financial interest.

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Supplementary Materials

ao6c05341_si_001.pdf (639.2KB, pdf)

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