Non-Invasive Detection of Interferon-Gamma in Sweat Using a Wearable DNA Hydrogel-Based Electrochemical Sensor

Abstract

Monitoring of immune factors, including interferon-gamma (IFN-γ), holds great importance for understanding immune responses and disease diagnosis. Wearable sensors enable continuous and non-invasive detection of immune markers in sweat, drawing significant attention to their potential in real-time health monitoring and personalized medicine. Among these, electrochemical sensors are particularly advantageous, due to their excellent signal responsiveness, cost-effectiveness, miniaturization, and broad applicability, making them ideal for constructing wearable sweat sensors. In this study, we present a flexible and sensitive wearable platform for the detection of IFN-γ, utilizing a DNA hydrogel with favorable loading performance and sample collection capability, and the application of mobile software achieves immediate data analysis and processing. This platform integrates three-dimensional DNA hydrogel functionalized with IFN-γ-specific aptamers for precise target recognition and efficient sweat collection. Signal amplification is achieved through target-triggered catalytic hairpin assembly (CHA), with DNA hairpins remarkably enhancing sensitivity. Ferrocene-labeled reporting strands immobilized on a screen-printed carbon electrode are displayed via CHA-mediated strand displacement, leading to a measurable reduction in electrical signals. These changes are transmitted to a custom-developed mobile application via a portable electrochemical workstation for real-time data analysis and recording. This wearable sensor platform combines the specificity of DNA aptamers, advanced signal amplification, and the convenience of mobile data processing. It offers a high-sensitivity approach to detecting low-abundance targets in sweat, paving the way for new applications in point-of-care diagnostics and wearable health monitoring.

1. Introduction

Wearable sensors offer a convenient approach to monitoring biomarkers, enabling continuous, real-time analysis in a non-invasive or minimally invasive manner. These features facilitate long-term monitoring and the collection of valuable data for clinical analysis. Among the various sensing technologies, electrochemical sensors stand out as the preferred choice for wearable sensors due to their exceptional performance, cost-effectiveness, miniaturization capabilities, and wide applicability. Wearable electrochemical biosensors are used in a wide range of applications. For example, graphene-based smart wireless health-monitoring systems have been developed to detect cortisol in sweat [1,2], while other electrochemical biosensors monitor nitric oxide levels [3]. Additionally, certain sensors target Pseudomonascan aeruginosa by detecting its phenazine metabolites as biomarkers [4]. Electrochemical biosensors have also been used to measure zinc ion concentrations, enabling the dynamic monitoring of insulin release [5]. Furthermore, protein-imprinted sensors have been electrochemically fabricated on screen-printed carbon electrodes (SPCEs) for the detection of the cytokine interleukin-1β (IL-1β) [6]. Biofluids such as sweat, tissue fluids, and tears present unique advantages as analytes due to their ease of sampling and compatibility with real-time monitoring. By providing continuous physiological information and insights into the deeper biomolecular state of patients, these biofluids hold immense potential for advancing personalized healthcare.

Diverse kinds of materials have been employed to create wearable sensors, such as polymer materials and plastics. Among them, DNA nanomaterials stand out as promising building blocks for the construction of wearable sensors due to their flexible programmability, superior biocompatibility, and target responsiveness. The specific recognition of DNA aptamers for biomarkers like interferon-gamma (IFN-γ) ensures accurate and reliable detection of physiological indicators in sweat. Furthermore, the integration of signal amplification strategies, including DNAzyme reactions, catalytic hairpin assembly (CHA), and hybridization chain reaction, extremely enhance the sensitivity of target detection, thereby enabling the detection of biomarkers in low abundance. This capability is particularly crucial in immunotherapy, where early detection and prompt intervention are particularly important for effective disease management and improved treatment outcomes. For instance, in pancreatic cancer studies, biomarkers have been highlighted to help in detecting and identifying unclear lesions early [7].

To achieve the effective collection of sweat, a hydrogel with a three-dimensional network configuration offers a powerful tool for sample accumulation and analysis. The incorporation of functional DNA modules further provides the desired physical properties and functionalities, and the high water content and biocompatibility provide a soft and flexible interface with human skin [8]. This ensures comfortable, long-term wear and efficient sweat collection. More importantly, the integration of a DNA hydrogel into wearable sensors also benefits from its electrical conductivity, which can be further enhanced by incorporating conductive materials or designing DNA modifications to establish conductive pathways. This capability facilitates real-time monitoring and data transmission, making them suitable for advanced bioelectronic applications.

By incorporating the functional DNA hydrogel with a screen-printed carbon electrode (SPCE), we designed a sensitive and convenient sensor for portable sweat collection and IFN-γ detection (Scheme 1). This was achieved by coupling an IFN-γ aptamer into a DNA hydrogel and immobilizing the report strand on an SPCE. Initially, the trigger strand (TS) was hybridized with an IFN-γ aptamer anchored onto the DNA hydrogel strand, while hairpin 1 (H1) and hairpin 2 (H2) were loaded into the hydrogel for signal amplification via a circular reaction. In the presence of the target IFN-γ, the trigger strand gained liberation under the combination of the aptamer with the IFN-γ, initiating the CHA of H1 and H2. The produced double-stranded complex (DS) could activate the strand displacement reaction of ferrocene, labeled L2 on the surface of the electrode, which made the ferrocene molecules drift away from the electrode and, thus, resulted in the reduction of the electrical signal. Leveraging the high loading capacity and excellent sweat collection capabilities of DNA hydrogel, this signal amplification strategy enables highly sensitive detection of the target biomarker IFN-γ in sweat. The use of SPCE, combined with a portable electrochemical workstation, significantly enhances the portability and convenience of this detection platform. Furthermore, we developed a mobile application for real-time data analysis and processing, facilitating instant target acquisition and analysis. This platform not only enables portable detection and real-time data analysis but also holds great significance for on-site diagnostics and disease monitoring.

2. Experimental

2.1. Materials and Reagents

HPLC-purified oligonucleotides were purchased from Sangon Co. Ltd. (Shanghai, China) and the detailed sequences are listed in Table S1. IFN-γ, IFN-α, and IFN-β were obtained from Yeasen Biotechnology Co., Ltd. (Shanghai, China). The 1 × PBS solution, 30% Acrylamide, 5 × TBE (tris-borate buffer), 10% ammonium persulfate, TEMED (tetramethyl ethylenediamine), TBS buffer, agarose, and 1 × TAE were provided by Solarbio Life Science (Beijing, China). TCEP and MCH (2-mercaptoethanol) were purchased from Macklin Biochemical Co., Ltd. (Shanghai, China). Chloroauric acid trihydrate (HAuCl₄·4H₂O) was purchased from Adamas (Shanghai, China). The screen-printed carbon electrode (SPCE) was obtained from Shenzhen Haoyang Technology Co., Ltd. (Shenzhen, China). Artificial sweat was purchased from Shanghai Yuanye Bio-Technology Co., Ltd. (Shanghai, China). NaCl was purchased from Shanghai Aladdin Biochemical Technology Co., Ltd. (Shanghai, China).

2.2. Apparatus

The polyacrylamide gel electrophoresis (PAGE) was conducted with a Mini-Protean Tetra Cell and a PowerPacTM Basic Power Supply (Bio-Rad, Hercules, CA, USA), and the gel imaging was performed on the ChemiDoc Imaging Systems (Bio-Rad, Hercules, CA, USA). Fluorescence spectra were recorded on an F-4700 spectrofluorophotometer (Hitachi, Tokyo, Japan). Differential pulse voltammetry (DPV) testing was obtained through the HY-E100X portable electrochemical workstation (Shenzhen Haoyang Technology Co., Ltd., Shenzhen, China). The app was exploited based on the Android studio. PBSs with different pH values were prepared with a pH meter FE28-Standard (Mettler Toledo, Zurich, Switzerland).

2.3. Optimization of the Binding Bases Between Aptamer and Trigger

One µM of the aptamer was heated with the trigger containing different numbers of binding bases (Lock-8, Lock-10, Lock-12, and Lock-14, respectively) in a thermostatic water bath at 95 °C for 5 min. After the reagents cooled slowly to room temperature, the cleaved Lock-8, Lock-10, Lock-12, and Lock-14 were each mixed 1:1 with the cleaved aptamer, and incubated at 37 °C for 2 h. Ten uL of the aptamer, Lock-8, Lock-10, Lock-12, Lock-14, Apt-lock-8, Apt-lock-10, Apt-lock-12, and Apt-lock-14 were injected into a 15% polyacrylamide gel. The gel electrophoresis was run in 1 × TBE buffer at 130 V for 70 min. Then, they were labeled with diluted SYBR Gold solution for 15 min, and finally, the results were analyzed by the Bio-Rad gel imaging system. The aptamer variants and aptamers were screened by polyacrylamide gel electrophoresis experiments (PAGE) to identify the antisense strands Lock-12 and Lock-14 with higher affinity to the aptamers.

2.4. The Signal Cyclic Amplification Verification

To verify the catalytic hairpin assembly reaction of H1 and H2, 10 μM H1 and 10 μM H2 were annealed, respectively, at 95 °C for 5 min and gradually cooled to room temperature. Then 1 μM trigger strand was mixed with 1 μM prepared H1 and 1 μM prepared H2 and incubated for 30 min at 37 °C. Production 1 was characterized by the polyacrylamide gel electrophoresis (PAGE) analysis, and the gel electrophoresis was conducted at 110 V for 65 min. Then, 10 μM L1 and 10 μM L2 were mixed and annealed at 95 °C for 5 min, followed by gradually cooling to room temperature to form the L1–L2 hybrid assembly. The signal cyclic amplification was validated by incubating production 1 with the L1–L2 assembly for 1 h at 37 °C, and the reaction was characterized by a PAGE analysis at 110 V for 70 min.

The fluorescence monitoring was conducted with the same approach, except that Cy3-labeled H1, BHQ2-labeled H2, FAM-labeled L1, and BHQ1-labeled L2 were used as substitutions respectively.

2.5. Assembly of L1 on the Electrode

Referencing the conventional carbon electrode modification of gold [9,10], a gold solution was configured by diluting 40 mg of HAuCl₄·4H₂O with 10 mL of 0.5 M H₂SO₄ solution. The gold solution was placed onto the carbon electrode and scanned by cyclic voltammetry over a potential range of 50 mVs⁻¹ until the electrode surface was uniformly covered by gold. The electrode was removed and rinsed with ethanol and ultrapure water and finally washed and blown dry. The sulfhydryl-labeled L1 was anchored on the gold-covered electrode according to the previous report [11,12]. The method is as follows. Before modifying the gold-covered electrode, L1 was diluted in TBE buffer to achieve the desired L1 concentration of 0.2 µM. The diluted L1 was water-bathed at 95 °C for 5 min to form a single strand and left to set at 4 °C for 10 min. The 0.2 µM thiosulfate solution L1-HS solution was mixed 1:5 with 15 mM TCEP solution and water-bathed at 25 °C for 1 h to form the HS structure. To make L1 fixed on the gold electrode, the thiosulfate solution was incubated on the surface of the gold-covered electrode for 1 h. After incubation, the gold-covered electrode was rinsed with TBE buffer and immersed in an aqueous solution of MCH for 1 h to displace non-specifically adsorbed L1 molecules and to passivate the gold-covered electrode surface. Finally, a TBE buffer was used to wash the gold-covered electrode again.

2.6. Electrochemical Characterization

To achieve an electrochemical signal response, Ferrocene-labeled L2 was hybridized with L1 by incubating at 95 °C for 5 min and then cooling to room temperature slowly. Then, the hybridization product was anchored on the gold-covered SPCE with the same approach as described above. The aptamer–trigger coalition was prepared through incubation at 95 °C for 5 min and then cooled-down to room temperature for over 4 h. Then, different concentrations of IFN-γ were mixed with 1 μM of the aptamer–trigger coalition. H1 and H2 were dropped on the gold-covered SPCE to incubate for 2 h at 37 °C. Afterward, the electrode was washed with TBE buffer 3 times to detect the electrochemical signal response with the DPV mode on the portable electrochemical workstation. The impact of salt on the DPV signals was evaluated according to similar procedures, except that the IFN-γ was incubated in PBS with 100, 300, or 500 nM NaCl. The impact of pH environments on the DPV signals was evaluated according to similar procedures, except that the IFN-γ was incubated in PBS with different pH values (6.0, 7.4, and 8.0). The specificity experiments were conducted according to similar procedures, except that the IFN-γ was substituted with IFN-α or IFN-β, and the control group was also handled using a strand with random sequences as substitution for the aptamer.

2.7. Development of Sensor Based on DNA Hydrogel

The agarose hydrogel was chosen to serve as the carrier of DNA probes. The shape fixation was performed using microscope slides, circular hole patches, etc., to form circular hydrogel patches. We prepare 2% (weigh 180 mg of agarose on an electronic analytical balance and add 8820 μL 1 × TAE) and 3% agarose gels (weigh 270 mg of agarose on an electronic analytical balance and add 8820 μL 1 × TAE), respectively. Place in a microwave oven and heat until the solution boils for 10 sec. Remove the conical flask, shake, and mix with the aptamer–trigger coalition, H1, and H2, and repeat four times until the gel is clear and particle-free. Quickly pipette 60 μL of solution onto the circular patch wells of the slide and cool naturally for 10 min to show the circular hydrogel.

To evaluate the detection performance of this DNA hydrogel platform, different concentrations of IFN-γ were dropped onto the bottoms of hydrogel patches, respectively, and incubated for different times at 37 °C. Then, the surfaces of the circular hydrogel patches were rinsed with 1× TBS buffer, and the electrochemical signal was recorded through the DPV mode of the portable electrochemical workstation. The sweat detection was conducted according to similar procedures, except that the artificial sweat was added with 100 nM INF-γ and was dropped and incubated on the bottom of hydrogel patches.

3. Results and Discussion

3.1. Target Triggered Cyclic Reaction

To achieve the signal response to IFN-γ, we employed IFN-γ aptamer for specific recognition and subsequent activation of DNA reaction. First, the polyacrylamide gel electrophoresis (PAGE) experiments were performed to verify the hybridization conformation of the aptamer with the trigger strand, as well as its release upon target binding. As shown in Figure S1, the aptamer strand and the trigger strand formed a band with much lower mobility, while in the presence of IFN-γ, a band with the same mobility as the trigger strand could be observed, indicating the release of the trigger strand after the treatment with the target. This demonstrated the target-induced release of the trigger strand, which is a key mechanism for achieving the targeted response and activating the signal generation.

The signal amplification strategy is necessary for the detection of targets with low abundance. Thus, a catalytic hairpin assembly reaction was introduced following the strand displacement of the trigger strand in the presence of IFN-γ. The CHA reaction in the solution was first verified through PAGE characterization. As shown in Figure 1A, after incubation with the trigger strand, a distinct bright band with a much slower migration could be observed in the mixture of H1 and H2, and the bands of the H1, H2, and trigger strand became dim, indicating the successful generation of their assembly complex (HAC). Additionally, Cy3-labeled H1 and BHQ2-labeled H2 were employed to further certify the trigger strand activated assembly, which displayed and obviously reduced fluorescence signal of Cy3 after incubation with the trigger strand because of the adjacencies of Cy3 in H1 and BHQ2 in H2 (Figure 1B).

Afterward, the signal had cyclic amplification through the strand displacement of L1 with HAC assembly. As shown in Figure 1C, after treatment with the produced assembly complex of H1, H2, and trigger (HAC), the L1–L2 hybrid assembly generated a band with faster migration, which displayed the same position as the band of L1, demonstrating the successful strand displacement and their cyclic reaction. In addition, fluorescence monitoring was further employed to evidence the trigger strand-activated signal cyclic amplification. FAM and BHQ1 were labeled on L1 and L2, respectively, to indicate the hybridization of L1 and L2 through the attenuation of FAM fluorescence (Figure 1D). In the presence of the trigger strand, the fluorescent intensity of FAM exhibited significant enhancement, suggesting the assembly of H1 and H2, as well as the subsequent strand separation of L1 and L2. The fluorescence variation was also used to indicate the target-triggered cyclic reaction. After the addition of IFN-γ, obvious fluorescence recovery could be observed because of the division of FAM-labeled L1 with BHQ1-labeled L2 (Figure 1D), indicating the successful assembly of H1 and H2 and the strand displacement of L1 for the signal cyclic amplification. These results demonstrated the feasibility of this CHA-based signal amplification strategy for the sensitive detection of IFN-γ.

3.2. Characterization and Optimization of Electrical Signal Response

To achieve a target-triggered electrical signal response, L2 was labeled with ferrocene (FC) and anchored on the electrode through hybridization with L1. The sulfhydryl (HS)-modified L1 was immobilized on a gold-coated screen-printed carbon electrode (SPCE) by a sulfur–gold (Au-S) bond (Figure S2). The electrical signal response was generated from the liberation of FC-modified L2 and characterized by differential pulse voltammetry. Differential pulse voltammetry (DPV) reduces the effect of currents generated by non-analytes, increases the detection sensitivity, and has become the electrochemical analysis method of choice for the sensitive detection of redox reactions on electrode surfaces in the laboratory. Therefore, we monitored the DPV signal variation after treatment with the target.

The strand displacement sensing mechanism for the electrical signal response was first verified through incubation with an HAC assembly complex that was synthesized from the H1, H2, and trigger strand. L1 was modified with a sulfhydryl group (HS) on the 3′ end to anchor on the gold-coated SPCE through Au-S binding [13,14], and L2 labeled with ferrocene was attached through hybridization with L1 to generate an electrical signal. As shown in Figure S3, the modification of FC-labeled DNA strands on the gold-coated SPCE generated an obvious peak in the DPV curve, as well as the cyclic voltammetry (CV) curve (Figure S3A–C). And in the presence of the HAC assembly, its substitution for L2 triggered cyclic amplification, resulting in the liberation of L2 and, thus, reducing the electrical signal (Figure S3D), which suggested the feasibility of this strand displacement mechanism for electrical sensing.

The IFN-γ triggered electrical signal variation was then monitored by incubating the target with probes for cyclic amplification in a solution. Figure 2A shows a typical surface electrochemical reaction with a reduction peak at around +0.2 V, indicating the presence of Lock-Fc on the surface. In the absence of IFN-γ, the mixture containing the trigger–aptamer complex, H1, and H2 did not have an evident influence on the electrical signal. After incubation with the target IFN-γ, an obvious decrease of the DPV peak could be observed, and the peak values displayed a falling off as the concentration of IFN-γ increased. The stability of the designed DNA sensor was investigated by monitoring the DPV signal under different concentrations of NaCl or pH conditions to simulate diverse physiological environments that may affect sensor performance. As shown in Figure S4, the sensor maintains consistent signal responses across a range of salt concentrations and pH levels. This adaptability suggested that the sensor can effectively operate in different detection environments, making it a reliable tool for real-world applications in sweat analysis and other biological assessments. To evaluate the specificity of this platform for IFN-γ detection, bull serum albumin (BSA) was employed as a control to incubate with the probe for DPV examination. As shown in Figure 2B, the BSA-treated group did not exhibit a distinct variation in the DPV signal in comparison with the blank group, while the IFN-γ incubated probe generated a significant reduction in the electrical signal. In addition, other types of cytokine interference (including IFN-α and IFN-β) did not cause a notable diminution of the DPV signals (Figure S5), further indicating the selectivity and specificity of this sensor in distinguishing IFN-γ from the complex physiological environment. These results suggested the specific concentration-dependent responsiveness of this platform to IFN-γ and the feasibility of IFN-γ detection. Additionally, a DNA strand with random sequences was used to substitute the aptamer strand for the control, which induced a slight variation in the DPV signal after incubation with INF-γ (Figure S5), indicating the specific binding of the aptamer and INF-γ for target recognition and electrochemistry signal response.

To enhance the sensitivity and signal-to-noise ratio (SNR) of this electrochemical sensor, we optimized the concentration of the FC-labeled L2 on the electrode surface. The L1-L2 assemblies were prepared at different concentrations for modification onto the electrode surface, and their performances in the presence of target molecules were examined through DPV signal variation. Following the application of these assemblies, the introduction of specific probes allowed for effective binding with the target, leading to a measurable electrochemical response. As shown in Figure S6, the choice of 0.2 μM presented as the optimal concentration based on a quantitative analysis of the SNR, which is defined as the value of (initial peak current–final peak current)/initial peak current. The significant increase in SNR at this concentration suggests a favorable interaction between the target and the recognition elements, potentially leading to improved detection capabilities in practical applications. Therefore, this concentration would be employed in subsequent experiments, with the aim of further signal enhancement and facilitating the sensor’s applicability to IFN-γ detection.

3.3. Characterization of Hydrogel Sensing Signals

Currently, hydrogels have been explored and used in biomedical applications based on their variable physicochemical, structural, and biological properties. Moreover, when combined with functional units, hydrogels can be used in biosensors, bioimaging, and wearable devices [15,16]. In addition, we refer to related texts and found that agarose is a natural polysaccharide hydrogel that is very biocompatible with skin [17,18,19]. To achieve portable detection of targets in sweat, we combined DNA probes with the agarose hydrogel materials to enhance the sample collection capability, while simultaneously facilitating the loading of nucleic acid probes and enabling signal amplification reactions. The DNA-loaded hydrogel was synthesized by mixing the coalition of the aptamer and trigger, H1 and H2 into the agarose hydrogel during the gumming process. The concentration of agarose was first optimized for better morphology and sample collection. Figure S7A shows that the agarose hydrogel with a concentration of 1% is slightly watery, which cannot be firmly adhered to the nanogold screen-printed electrode for sample collection. An agarose hydrogel of 2% and 3% can firmly adhere to the nanogold screen-printed electrode patch, which proves the feasibility of the agarose hydrogel being grown on the surface of the electrode patch. The flow properties of the agarose hydrogels with different concentrations were further evaluated by monitoring the steady-state viscosity as a function of the shear rate at 20 °C, which indicated increased viscosity as the content of agarose raised (Figure S7B). Then, the DPV responsiveness of the aptamer sensor was modified with different concentrations of agarose hydrogel to IFN-γ. After the addition of the analyte protein IFN-γ, the peak current decreased significantly and the 2% agarose hydrogel suggested higher suppression in comparison with the control group without target treatment (Figure S7C), verifying that 2% is the better concentration for the detection of IFN-γ through the DNA cascade reactions.

To achieve a stronger signal response, we optimized the incubation time of the nucleic acids on the electrode surface. As illustrated in Figure S8, the electric signal progressively diminishes with an increase in incubation time. This reduction in the electric signal can be quantitatively assessed by analyzing the rate of decrease over time. According to the relevant literature, we normalized the signal to the final peak current and converted it to signal suppression by following a simple metric (initial peak current–final peak current)/initial peak current [20,21]. It is evident that, as the incubation duration increases, the rate of signal reduction also escalates, ultimately stabilizing around the 30 min mark. This observation suggests that extending the incubation time beyond 30 min does not yield significant improvements in signal intensity, indicating that the optimal binding of nucleic acids to the electrode surface has been reached. The stabilization of the electric signal at approximately 30 min implies that this duration is enough for the nucleic acids to effectively interact with the electrode, leading to a maximized signal response. Therefore, based on these findings, we have selected a 30 min incubation time as the optimal condition for our experiments.

Then the detection performance of this platform was examined by monitoring the variation of the electrical signal after incubation with IFN-γ. As shown in Figure 3, as the concentration of IFN-γ increased, the DPV signal demonstrated a gradual decline, and the signal suppression rate showed a linear correlation between the peak value and target concentration. The limit of detection (LOD) and limit of quantification (LOQ) are calculated to be 0.049 nM and 0.097 nM. The signal-to-noise ratios (SNRs) under different concentrations of IFN-γ data were also measured by the ratio of the signal suppression value under target treatment and the standard deviation of the blank (Figure S9). The observed decrease in signal intensity with rising concentrations suggested that the target effectively triggered the cascade reaction between DNA probes, leading to a measurable response that can be reliably correlated to the concentration levels. This linear relationship suggested the application potential of this platform for the real-time monitoring and portable detection of IFN-γ in sweat.

To validate the application of this designed sensor for the detection of IFN-γ in sweat, the artificial sweat was selected according to the standard ISO 3160-2 with IFN-γ added to simulate the real sample detection [22]. As shown in Figure S10, the DPV signal displayed an obvious decrease after treatment with the artificial sweat containing INF-γ, which indicates the sensor’s efficacy in responding to the target in the artificial sweat. These results corroborate the sensor’s ability to identify and analyze INF-γ within sweat, highlighting its potential for practical applications in the detection of biomarkers in sweat that correlate with physiological states or health conditions.

3.4. Architecture of Intelligent Biosensor

Nowadays, wearable sensors are developing and becoming a hot topic for researchers. The development of wearable biochemical sensors for lactic acid monitoring, diabetes monitoring, and tuberculosis monitoring shows that the electronic skin sweat sensor has broad application prospects and can provide important help for future human health monitoring. One of the wearable biochemical sensors for lactic acid monitoring is a sensor patch that integrates signal acquisition and transmission circuit components as a way to obtain real-time monitoring of the biomarkers in sweat [23,24,25]. Based on this, we use the existing technology to transmit the measurement data from the micro-electrochemical workstation to the computer terminal through a wired connection, use the relevant software to read the data, and then import the data into our self-designed app—Foresight. The maximum value of the standard content range in the human body is compared and analyzed to determine whether the INF-γ in the human body is within the normal range. The system uses electrochemical biosensors to detect the content of inflammatory factor INF-γ and uses artificial intelligence to train and learn in depth. When the detected data are within the standard range, the Foresight app will display the risk-free green text. When the monitored data is out of the standard range, the Foresight app displays a red print of the detected risk. The self-developed Foresight app can judge the test results and remind itself whether there is a storm of inflammatory factors, etc., and detect its abnormalities as soon as possible and treat them. Since hydrogels are soft materials with a three-dimensional network structure, these three-dimensional porous networks can often be used to store, release, or collect materials, or to mimic the natural tissue microenvironment [26,27]. From the above experiments, we can clearly see that the hydrogel prepared with agarose is more suitable for our sensor devices. The agarose hydrogel has a high affinity for the skin, which can effectively collect sweat and enable the analyte to react with the sensor.

The overall operation principle of the system and its detailed equipment component workflow are as follows. Before it all starts, it is first necessary to ensure that all devices are correctly connected. This includes electrochemical workstations, electrode adapters, and related signal transmission lines. Then, through the professional software interface, according to the specific needs of the experiment, accurately set the parameters, such as the target voltage range, data acquisition frequency, etc., to fully prepare for the formal start of the system. When the system receives the start command, the entire workflow will start immediately. First, the device will send the target voltage value to the electrochemical workstation according to the preset procedure. The electrochemical workstation receives this command and quickly adjusts its internal circuit to generate and output the corresponding target voltage. The voltage is then accurately transferred to the electrode through the electrode adapter. At the electrode pad, the voltage forms a closed loop with the solution in contact with the electrode pad. In this circuit, a series of electrochemical reactions occur, which are the key components of the system that need to be monitored and analyzed. As the electrochemical reaction proceeds, the electrode adapter captures the generated target current in real time and transmits it back to the electrochemical workstation again. The electrochemical workstation accurately measures and records the received current and converts it into the corresponding target current value.

Finally, the current value is sent back to the device to complete a complete data acquisition process. The device will automatically change the target voltage value according to the preset program, restart the entire workflow, and perform the next round of data acquisition. This cycle is repeated until all the preset experimental steps and data acquisition tasks are completed. In this process, the various components of the system play a vital role. As shown in Figure 4, there are also L1 aptamer-modified nanogold screen-printed electrode sheets. After dropping INF-γ, it is fully reacted and placed in the electrode placement device. Connect to the electrochemical workstation, perform electrochemical measurements, transmit the data to the mobile phone, and pass the data to the Foresight app. The Foresight app compares the peak of the input data with the peak of the maximum value of the normal range of INF-γ, suggesting that the user’s own health is at risk. They work together to ensure the accuracy and reliability of the experiment. At the same time, the precise control of the software and the real-time recording of the data also provide strong support for the subsequent data analysis and research.

4. Conclusions

As an ideal biological detector, sweat contains various biomarkers, including interferon-gamma (IFN-γ), that indicate the physiological state of the human body and play an important role in health monitoring and disease detection. In this paper, we developed a wearable platform for the sensitive detection of IFN-γ in sweat through DNA cascade reaction-based signal amplification technology. DPV data showed that this platform had good detection specificity for the target and a favorable linear correlation with the target concentration. In addition, the portable electrochemical workstation enables the use of a smartphone for on-site signal detection, and we further developed a mobile software application to facilitate data processing and analysis of the detection results, allowing for the real-time reporting of IFN-γ in sweat. This detection platform provides a convenient and accessible solution for field-deployable detection, expanding the capabilities of DNA nanotechnology in the application of point-of-care testing.