Microneedle-assisted interstitial fluid extraction combined with a 2D tellurene-based glucose sensor for personal diabetes management

Abstract

The objective of this work is to develop a selective glucose sensor for real-time in vivo monitoring of interstitial fluid (ISF) glucose levels, addressing challenges of interference from other biomolecules. A microneedle (MN) patch was employed for ISF extraction, while a 2D tellurene (Te)-based glucose oxidase (GOX) GOX/Te/carbon cloth (CC) electrode served as the sensing element for electrochemical measurements. The sensor exhibited a low detection limit (LOD) of 0.357 μM, and a sensitivity of 0.10378 mA μM⁻¹ cm⁻². In vivo studies in an animal model demonstrated an average glucose concentration of 5.05 mM with a rapid response time of ∼95 s. Importantly, the electrochemical measurements showed a strong correlation with commercial glucometer readings, confirming accuracy and reliability. This study highlights the potential of the proposed sensor for real-time in vivo glucose monitoring, paving the way for future home care diabetes management devices.

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1. Introduction

Diabetes mellitus is a prevalent chronic condition resulting from insulin dysfunction, which can cause severe complications such as blindness, kidney failure, stroke, numbness, and heart attacks.¹ Continuous glucose monitoring (CGM) in interstitial fluid (ISF) offers real-time glucose tracking, reducing the need for frequent finger pricks and improving diabetes management by detecting glucose fluctuations and optimizing treatments.² This approach lowers healthcare costs and enhances patient outcomes.³ Microneedle (MN)-based glucose sensors are attracting interest for their minimally invasive, nearly painless design, which penetrates the skin to detect biomarkers present in ISF, improving CGM sensor performance.

Frequent glucose monitoring is crucial in diabetes management. While subcutaneous CGM devices like Medtronic MiniMed 670G,⁴ Dexcom G6,⁵ and Abbott FreeStyle Libre⁶ are effective, they have drawbacks such as skin pain, acupuncture anxiety, and allergies. Non-invasive methods using sweat, saliva, and tears have emerged as promising but lack accuracy due to inconsistent glucose correlations,^7,8 CGM sensors also face issues with long-term sensitivity and biocompatibility. Microneedles offer the potential for real-time monitoring but currently lack the required accuracy for diabetes management.

Since graphene's discovery in 2004,⁹ various 2D materials with remarkable properties, such as transition metal dichalcogenides (TMDCs), h-BN, and phosphorene, have been developed, enabling applications in electronics, photovoltaics, thermoelectrics, and sensors.^10–14 Recently, monolayer tellurene, the first 2D material from group VI, has shown promising features like a moderate band gap, high environmental stability, excellent carrier mobility, and a high on/off ratio.^15–22 Despite these advantages, tellurene's potential as a biosensor material remains underexplored. 2D materials, in general, are promising for electrochemical sensing due to their high surface area and tunability.²³ Tellurene, with its semiconducting and photoconductive properties, is ideal for sensors and energy devices,^24–27 and its strong interactions with metals improve electrocatalytic performance, especially with noble metals.²⁸ However, challenges like nanostructure synthesis and stability persist. While 1D tellurium has been studied for nonenzymatic H₂O₂ detection, research on glucose-sensing with 2D tellurene is limited.^29–32 The high surface area and efficient electron transport of 2D tellurene make it a promising candidate for glucose sensors, offering efficient enzyme immobilization and electron transfer.³³ Their scalable, cost-effective synthesis of 2D tellurene further enhances its potential for next-generation biosensors, highlighting the need for new 2D tellurene materials.

This study involves three distinct experimental phases: (i) in vitro electrode characterization of the GOX/Te/CC electrode in buffer solution, (ii) ex vivo ISF analysis where ISF extracted via the MN patch was analysed separately using the electrode, and (iii) in vivo real time sensing in which the MN patch remained in place for continuous ISF access while electrochemical measurements were performed using a standard three electrode setup. The MN patch served solely for ISF access and was not physically integrated with the sensing electrode. This study demonstrated a microneedle (MN)-assisted ISF extraction combined with a 2D Te-based glucose sensor for determining glucose in an ISF sample and in vivo real-time monitoring, and an animal model is used to study the pathogenesis of diabetes disease in Sprague Dawley (SD) rats. MN patches were fabricated by using a 3D printing technique with a height of 1.5 mm, base diameter of 0.4 mm, and tip diameter of 0.02 mm. In vivo tests were performed in a 7–10-week-old male SD rat model and an MN patch, consisting of 100 needles was inserted into the dorsal skin by finger press. Then patches were extracted to collect glucose containing ISF. The GOX/Te/CC modified electrode was fabricated by the direct drop casting method and carbon cloth (CC) was used as an electrode base material that allowed for the electrocatalytic sensing of glucose biomolecules. The glucose oxidase (GOX) enzyme is immobilized on the working electrode surface. The GOX/Te/CC electrode served as a working electrode and showed the selectivity of 0.10378 mA μM⁻¹ cm⁻² with a correlation coefficient of R² = 0.97 and the limit of detection (LOD) was found to be 0.357 μM. Using the GOX/Te/CC electrode the glucose concentration in the ISF samples was found to be 0.99 μM and 1.13 μM, and by commercial glucometer (Apollo Healthco Limited, India), it was 1.09 μM and 1.21 μM, respectively. The data obtained from the GOX/Te/CC electrode were correlated with the commercial glucometer data. The GOX/Te/CC electrode exhibited favorable behavior towards glucose oxidation, high sensitivity to different glucose concentrations, reproducibility, and exceptional selectivity towards multiple interfering species. Due to these obvious advantages, the modified electrode has promising prospects for practical applications.

2. Experimental section

2.1. Materials

All chemicals used were of analytical grade. Sodium tellurite (Na₂TeO₃), polyvinylpyrrolidone (PVP), hydrazine hydrate (N₂H₄), ammonium hydroxide solution (NH₄OH), D-⊕-glucose, glucose oxidase (GOX) (from Aspergillus Niger type VII, ≥100 000 units per g solid), dopamine hydrochloride (DA), L-ascorbic acid (AA), uric acid (UA), and Nafion (5 wt%) were purchased from Sigma-Aldrich-USA. Sodium phosphate dibasic dehydrate (Na₂HPO₄·2H₂O) and sodium phosphate monobasic dehydrate (NaH₂PO₄·2H₂O) were purchased from Sisco Research Laboratories Pvt. Ltd (SRL)-India. Carbon cloth (CC) was obtained from Boffin Butler Pvt. Ltd, India. Stock solutions of glucose, ascorbic acid, dopamine, and uric acid were prepared using DI water.

2.2. 2D tellurene (Te) synthesis

2D Te was synthesized by using the hydrothermal method. Fig. S1a shows a schematic diagram of the 2D Te synthesis process. Na₂TeO₃ (99.71 mg) and PVP (343.45 mg) were dissolved in 33 mL of deionized water using a magnetic stirrer for 20 min. A mixture of ammonia solution (1.66 mL) and hydrazine hydrate (1 mL) was prepared and placed in a Teflon-lined stainless-steel autoclave for hydrothermal treatment. The autoclave was then heated in an oven at 180 °C for 20 hours. Once cooled down to room temperature, the resulting precipitates were separated by centrifugation at 5000 rpm for 15 min. The final product was thoroughly washed with deionized water several times and dried at 100 °C for 12 hours.

2.3. Fabrication of the working electrode

The electrode fabrication process of the GOX/Te/CC electrode is shown in Fig. S1b. Before the fabrication of the electrode, carbon cloth (CC, cut into 3 × 1 cm² dimensions) was ultra-sonicated using ethanol and deionized water for 10 min and rinsed thoroughly. Te (20 mg) was dispersed in 100 μL of deionized water and the solution was ultrasonically stirred to make a homogeneous solution for 15 min, and 10 μL of the as prepared nanocomposite was deposited by a drop-casting method on a cleaned CC surface area of 1 × 1 cm² and stored at room temperature for drying. 6 mg of GOX was dispersed in 100 μL PBS (0.1 M) solution and the mixture was ultrasonically stirred to achieve homogeneous dispersion. 5 μL of GOX solution was immobilized on the Te/CC surface. To avoid leaching of the GOX enzyme a final modification of the Nafion solution was deposited. Electrochemical measurements were performed using a K-Lyte potentiostat workstation (Kanopy Techno Solutions Pvt. Ltd, India) for all in vitro, ex vivo, and in vivo GOX/Te/CC electrode studies.

2.4. Fabrication of 3D printed microneedles for accessing ISF in skin

Microneedles were fabricated using a fast and efficient masked stereolithography apparatus (mSLA) 3D printing technique, as shown in Fig. 1a. The complete fabrication of the MN patch was possible within 3 hours with minimal error. A 10 × 10 array of pyramidal microneedles with a height of 1.5 mm, base diameter of 0.4 mm and tip diameter of 0.02 mm was 3D printed by mSLA technology (Sonic mini 8K, Phrozen Technology, Taiwan). It is equipped with a 405 nm UV light. Poly (ethylene glycol) diacrylate (PEGDA) with an average M_n of 700 was used as the photo-responsive material of construction. PEGDA is a biocompatible material, which is categorized as a Generally Regarded As Safe (GRAS) material by the USFDA. Hydrogel monoliths made from PEGDA have been recently used for drug delivery. Lithium phenyl-2,4,6-trimethylbenzoylphosphinate (LAP), a water soluble free radical generator, was added as the photoinitiator. Various combinations of PEGDA and LAP were tested to obtain the desired level of mechanical strength. Finally, in an amber glass flask, 0.3 g of LAP was dissolved in 17.7 mL of deionized water. To this, 70 mL of PEGDA was added and mixed until a homogenous solution was formed. 12 mL of glycerol was added as a plasticizer. The as prepared ink was poured into the vat of the printer.

Fig. 1 (a) Schematic illustration of 3D printed PEGDA MN patch, and (b) schematic representation of an enzymatic reaction of glucose.

The MN patch was designed using computer-aided design CAD software (AutoCAD 2023, Autodesk, Inc., SF, California, USA). The design was then exported as an .stl file and used for slicing. The .ctb file was configured in the slicer software (CHITUBOX basic, V1.9.1). The design was adjusted at an x-axis angle of 35°. Conical supports with a contact depth of 0.02 mm were strategically placed on the overall base of the patch. The layer exposure time was set at 3 s with the bottom exposure at 5 s. Other parameters like lifting distance and lifting speed were set at 3 mm and 10 mm min⁻¹, respectively. Specific parameters like rest time before lift and retract were optimized. The printing time was 2 hours 19 min.

After the printing was finished, the printed structure was carefully removed from the build plate and washed gently by repeatedly dipping in IPA for 5 min. The supports were carefully clipped using scissors. They were then dried and subjected to a post-curing process for 15 min each using the Phrozen Cure station (Phrozen Technology, Taiwan).

The MN patch was used only to access ISF and contained no sensing element. In vitro experiments used only the GOX/Te/CC electrode in the three electrode setup. For ex vivo experiments, ISF collected via the MN patch was analyzed separately. During in vivo real-time sensing, the MN patch remained attached for ISF access while the GOX/Te/CC, reference, and counter electrodes were connected to the potentiostat.

2.4.1. Aspiration of ISF from animals. In vivo experiments were performed in a 7–10-week-old male Sprague Dawley rat model. All animal experiments were conducted following the guidelines set by the Committee for the Purpose of Control and Supervision Experimental Animals (CPCSEA), India. All experimental protocols were approved by the Institutional Animal Ethics Committee (IAEC) of BITS Pilani Hyderabad Campus (Approved No. BITS-HYD/IAEC/2020/05). The MN patch of 100 needles was inserted into the dorsal skin by finger press. Adult rats (7–10 week-old) were entrained to a 12-hour light–dark cycle for 7 days with unrestricted access to food and water. Following this period, the hair on their dorsal side was gently trimmed away with a hair clipper and a depilatory cream. The animals were kept under mild anesthesia with an isoflurane-oxygen pump. The patch was held in place for 10 min using occlusive tape. The patch was then withdrawn and extracted for ISF. At the conclusion of the study, the rats were euthanized through CO₂ asphyxiation, a humane and efficient method that ensures a quick and painless process. CO₂ was introduced at a controlled rate, causing unconsciousness before vital functions ceased. This procedure adhered to CPCSEA guidelines to minimize distress.

2.4.2. Extraction of ISF from the MN patches. The patches were extracted for the collected ISF containing glucose. The patches were cut into small pieces using a scalpel blade. They were transferred into a 5 mL centrifuge tube and 1.5 mL of deionized water was added and left to equilibrate for 15 min. This was followed by homogenization at 12 000 rpm for 3 min and centrifugation for 15 min at 7000 rpm at 4 °C. The supernatant was collected and stored at −80 °C until further use. The MN patch served solely to extract ISF and remained attached during in vivo monitoring to allow continuous ISF access; electrochemical sensing was performed separately using the GOX/Te/CC electrode in a standard three-electrode configuration. The glucose concentration reported for the ISF samples in Table 1 reflects values measured from the diluted extracts obtained after MN patch collection and processing. These do not represent undiluted physiological ISF concentrations, which are higher in vivo. All glucose concentrations in this study are reported in μM, mM, or mg dL⁻¹ depending on experimental context, with clear indication of the units in each case. Where relevant, values have been converted between units for clarity and consistency across tables and figures.

Table 1 Real sample analysis of glucose in rat ISF

S. No.	Sample	Glucose concentration (μM) detected by GOX/Te/CC sensor (μM)	Glucose concentration (μM) detected by glucometer (μM)
1	Rat ISF sample 1	0.99	1.09
2	Rat ISF sample 2	1.13	1.21

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2.5. Redox activity of glucose enzyme

The sensing mechanism of glucose is based on the redox reactions between analyte (glucose) and enzyme (GOX), described in the following equations.³⁴

H₂O₂ → O₂ + 2H⁺ + 2e⁻ (2)

GOX catalyzes the oxidation of glucose to gluconic acid and the reaction involves the transfer of electrons from glucose to GOX and generates H₂O₂, which is electrochemically detected at the electrode surface. The glucose molecules adsorb and diffuse through the electrode (GOX/Te/CC) surface, which accelerates both the adsorption and diffusion processes due to the large surface area and surface active sites of Te. The oxidizing H₂O₂ environment inhibits the reduction of TeO₂, resulting in a reduced signal at the negative potential of CV at elevated concentrations, along with the electron transfer process shown in Fig. 1b.

2.6. Characterization

The crystal structure of the Te/CC was examined by using an X-ray diffractometer (XRD, Rigaku Ultima IV). A field emission scanning electron microscope (FE-SEM, FEI-Apreo S) and transmission electron microscopy (TEM, Philips CM 200) were used to study the surface morphology of the 2D tellurene, bare CC, and Te/CC electrodes. Furthermore, the electronic and chemical states of Te/CC were analyzed by using X-ray photoelectron spectroscopy (XPS). Raman spectroscopy (Sunterra via Opus (Brucker) Spectrometer) was employed to further study the structure of the as synthesized 2d tellurene nano flakes.

2.7 Electrochemical measurements

Electrochemical experiments were carried out using a K-Lyte electrochemical workstation (Kanopy Techno Solutions Pvt. Ltd, India) in a 0.1 M phosphate buffered saline (PBS, pH 7.4) solution at room temperature. PBS is essential in enzymatic glucose detection as it maintains a stable pH, and ensures optimal enzyme activity. Its isotonic nature and buffering capacity prevent enzyme denaturization, making it a reliable and cost effective solution for accurate glucose measurements. The analysis was performed using a three-electrode cell setup, with a carbon cloth substrate as the working electrode, a platinum mesh as the counter electrode and Ag/AgCl as the reference electrode.

In the CV measurement, the potential between the working electrode and reference electrode was swept in a linear manner, while the current between the working electrode and the counter electrode was measured. In CA measurement, the potential applied between the working electrode and reference electrode was set based on the anodic peak voltage (i.e. potential difference where maximum current is obtained) from the CV curve, and the current flowing from the working electrode to the counter electrode was measured as a function of time. The CA measurements were conducted to quantitatively detect the glucose at the GOX/Te/CC electrode. Glucose was detected at the following concentrations: 1 μM to 5 μM and 0.5 mM to 7 mM, respectively.

3. Results and discussion

An overview of the experimental stages and device configuration is presented in Fig. 2. The workflow consisted of four stages: (i) in vitro electrochemical characterization of the GOX/Te/CC electrode; (ii) ex vivo ISF analysis, where ISF was extracted using MN patch and analysed separately; (iii) in vivo real-time sensing, where the MN patch remained inserted for continuous ISF access and glucose detection; and (iv) animal model studies comparing ISF and serum glucose levels under fasting, fed, and insulin-treated conditions.

Fig. 2 Schematic diagram illustrating the experimental flow and device configurations used in this study: (i) in vitro GOX/Te/CC electrode characterization in buffer; (ii) ex vivo ISF analysis where ISF extracted using an MN patch was analysed using the electrode; (iii) in vivo real-time sensing with the MN patch in situ and the GOX/Te/CC electrode connected in a standard three-electrode setup; and (iv) the animal model study under fasting, fed, and insulin-treated conditions.

3.1. Characterization of the 3D printed MN patch

The 3DP MN patches were observed under the stereomicroscope and images were captured. The observed needles were distinctly pyramidal in shape. The length, base width and tip diameter were found to be 780 ± 30 μM, 350 ± 10 μM, and 40 ± 5 μM, respectively Fig. 3(a) and (b) show stereomicrographs of the 3DP MNs at different magnifications. Field emission scanning electron microscopy (FESEM) was also performed to confirm the shape geometry and dimensions. Fig. 3c depicts representative images of the micrographs captured. Fig. 3d shows the force vs. displacement graph of the 3DP MNs captured during the compression testing using the texture analyzer. The compression strength was found to be 0.46 ± 0.09 N per needle. This strength is more than the reported minimum strength required for easy penetration into the skin.^35,36

Fig. 3 (a) and (b) Stereomicroscopic images; (c) representative SEM image and (d) force vs. displacement graph of the 3D printed MN patch.

3.2. Characterization of the Te/CC electrode

Fig. 4a displays the X-ray diffraction (XRD) pattern of Te/CC. The XRD pattern of a Te/CC has peaks at different diffraction angles (2θ) of 25° (1 0 0), 27.53° (0 0 2*), 29.53° (1 0 1), 40° (1 0 2), 42.5° (1 1 0), 45.3° (1 1 1, 1 0 1*), 51.6° (2 0 1), 59.1° (2 0 2), 64.9° (1 1 3), 68° (1 0 4), 74.15° (2 1 2), and 77.7° (3 0 1) with JCPDS no. 36-1452. The CC exhibited a broad diffraction peak at 27.52° and a small peak at 45.2° corresponding to the planes (0 0 2) and (1 0 1), respectively, coherent with that report in the literature.³⁷ The distinct peak at 29.52° corresponds to the (1 0 1) plane of hexagonal Te, which belongs to the P3₁21 space group. The diffraction peaks observed in the XRD pattern of Te/CC closely align with all the peaks listed in JCPDS no. 36-1452.³⁸ X-ray photoelectron spectroscopy (XPS) study was performed to determine the oxidation state and chemical composition of the synthesized Te, and Fig. 4b displays the survey spectra of the Te/CC electrode, where Te 4d, C 1s, O 1s, Te 3d_5/2 and Te 3d_3/2 are assigned 43.09 eV, 284.07 eV, 532.07 eV, 576.06 eV and 586.07 eV binding energies, respectively, with a peak separation energy of 10.55 eV.^39,40Fig. 4c shows the core level fitting spectra of Te 3d and Te O, displaying two prominent peaks in the Te 3d core-level spectrum. These peaks correspond to the 3d_3/2 and 3d_5/2 states of Te²⁺ at binding energies of 586.58 and 576.28 eV, respectively. The two adjacent peaks at 583.58 and 573.15 eV are attributed to the binding energy of the Te⁴⁺ ions with oxygen,⁴¹ resulting from the surface oxidation of Te occurring throughout the transfer of the sample for analysis. Fig. S2(a) and (b) display the surface morphology of the bare carbon cloth (CC) and Te/CC observed by FE-SEM at high magnification. Fig. S2a shows that the CC is composed of a large number of carbon microfibers having a diameter of ∼10 μM to create a porous structure. Fig. S2b shows the dimensions of hexagonal Te nanorods with diameters of ∼100 nm and lengths of 500 nm, which have an atomic weight percentage of Te 63.87%, carbon 34.34%, and oxygen 1.78%, which demonstrates that Te is successfully deposited on the CC surface. Transmission electron microscopy (TEM) was performed to study the small surface structures and confirm the occurrence of a sheet-like structure with adjacent dimensions ranging from 30 μM to 6 μM as shown in Fig. 4(d) and (e). To further confirm how the structural vibrational modes change depending on the morphology of tellurene, Raman spectroscopy measurements were also performed as shown in Fig. 4f. The 2D tellurene exhibited three major peaks at 92.6, 120.86, and 141.17 cm⁻¹, corresponding to E₁, A₁, and E₂ modes. This denotes that all the vibration modes are consistent with those in the literature,^42,43 the A₁ mode is owing to atoms moving in the basal plane, and the E₁ and E₂ modes correspond to bond bending (a-axis) and bond stretching (c-axis).

Fig. 4 (a) XRD spectra of the Te/CC electrode; (b) and (c) XPS spectra of the Te/CC electrode; (d) and (e) TEM images of 2D tellurene, and (f) Raman spectra of 2D tellurene.

3.3. In vitro electrode characterization

The electrochemical performance of the GOX/Te/CC electrode was first evaluated in vitro in buffer solution to establish baseline sensitivity, selectivity, and stability prior to ISF testing. Fig. 5a shows the electrochemical performance of different electrodes (bare CC, Te/CC and GOX/Te/CC) in phosphate buffer saline (PBS, pH 7.4) solution with a voltage range from −1.3 to 1 V. From the voltammetry response it is noted that the bare CC electrode exhibited no noticeable oxidation current, whereas the Te/CC electrode showed a peak current density of ∼0.97 mA cm⁻² at 0.4 V. After the modification with GOX (GOX/Te/CC electrode), the anodic peak current density increased to 1.768 mA cm⁻² and the cathodic peak current density decreased to −1.85 mA cm⁻². The oxidation peak current increases more with the GOX/Te/CC electrode compared to the Te/CC electrode. This result might be assigned to the GOX enzyme, in which GOX enhances the electrochemical reactions of glucose. The electron produced during the glucose oxidation reaction is transported to the Te surface, which, as a conductive material, efficiently conducts electrons away from the electrode, generating a quantifiable electric current that can be detected and measured. The amplification of the redox peak with increasing concentration confirms the effectiveness of the GOX/Te/CC sensor.

Fig. 5 (a) CV curve of bare CC, Te/CC and GOX/Te/CC electrode; (b) CV curve of the GOX/Te/CC electrode at different scan rates; (c) CV curve of the GOX/Te/CC electrode at different concentrations; (d) and (f) chronoamperometry response of glucose at different concentrations; (e) calibration curve of (d); (g) selectivity study of the GOX/Te/CC electrode for the detection of glucose in the presence of various interferents, such as urea, dopamine, uric acid, and ascorbic acid; (h) relative response curve of the GOX/Te/CC electrode in the presence of different interferents; (i) glucose current response of the GOX/Te/CC electrode detected for 30 days; (j) repeatability of three different glucose sensors; (k) and (l) chronoamperometry response of ISF sample 1 & 2.

3.3.1. Effect of scan rate. The effect of the scan rates on the cyclic voltammetry measurement of the GOX/Te/CC electrode was investigated as shown in Fig. 5b. The cyclic voltammogram reproduces the same curve shapes with varying scanning speeds of 50 mV s⁻¹, 60 mV s⁻¹, 70 mV s⁻¹, 80 mV s⁻¹, and 90 mV s⁻¹ in 0.1 M PBS solution containing 10 μM glucose concentration. A linear relationship was observed between peak current and square root of the scan rate (Fig. S3), with a correlation coefficient of R² = 0.98. With increasing scan rate, the reaction occurs faster and leads to a decrease in the size of the diffusion layer; as a consequence, higher currents were observed.⁴⁴ Both the anodic current and cathodic current were enhanced with increasing scan rate.

3.3.2. Effect of glucose concentration. Fig. 5c displays the CV analysis of the GOX/Te/CC electrode observed in the absence (0 μM) and presence (3 μM, 7 μM, 10 μM) of glucose in PBS solution (pH 7.4) with a scan rate of 50 mV s⁻¹ to decrease the background current and to obtain high-density sensitivity. As shown in Fig. 5c, in the absence of glucose, an anodic peak exists at ∼0.4 V with a current density of 1.640 mA cm⁻² and a cathodic peak at −0.8 V with a current density of −1.603 mA cm⁻². When 3 μM of glucose was added to the electrolyte, the anodic peak current density was increased to 1.74 mA cm⁻², and the cathodic peak current density was decreased to −1.886 mA cm⁻², respectively. Furthermore, with increasing concentration of glucose (7 μM, 10 μM), the oxidation peak currents were also increased gradually. Thus, GOX/Te/CC electrodes can be used in glucose sensing materials due to the direct relationship between the current response and glucose concentration at a potential of ∼0.4 V.

To further validate the electrocatalytic performance of the GOX/Te/CC electrode, chronoamperometric measurements were conducted. Fig. 5(d) and (f) show the chronoamperometric response recorded at an optimal potential of +0.45 V in 0.1 M PBS solution under stirring conditions to minimize the thickness of the diffusion layer around the working electrode, with sequential addition of glucose. Two separate experiments were performed: one covering glucose concentration from 1 μM to 5 μM (Fig. 5d) and another in a wider range from 0.5 mM to 7 mM (Fig. 5f). In both cases, a stepwise increase in current was observed with each successive addition of glucose. Furthermore, the inset of Fig. 5(d) and (e) provides a magnified view of the lower concentration range, showing a strong linear correlation (R² = 0.97). Notably, a distinct current response was detected even at 1 μM glucose, demonstrating the sensors’ high sensitivity of 0.1037 mA μM⁻¹ cm⁻² and the LOD was found to be 0.357 μM. These results clarify that the GOX/Te/CC electrode demonstrates remarkable specificity towards the detection of glucose. The limit of detection (LOD) of the GOX/Te/CC electrode was determined by the following equation:

LOD = 3.3σ/S

where σ is the standard deviation of the noise while S is the slope of the linear calibration curve measured by the GOX/Te/CC glucose sensor. Fig. 5f shows a linear range from 0.5 mM to 7 mM with the sensitivity and correlation coefficient of 0.0786 mM mA⁻¹ cm⁻² and R² = 0.97.

3.3.3. Selectivity, stability, and reproducibility of the GOX/Te/CC electrode. To assess the operation of stability, the GOX/Te/CC was used for sensing in 0.1 M PBS solution (pH = 7.4) for 50 cycles as shown in Fig. S4. Additionally, the anti-interference ability and selectivity of the GOX/Te/CC sensor were studied by introducing electroactive species like urea, L-ascorbic acid, dopamine and uric acid as displayed in Fig. 5g. The GOX/Te/CC electrode retains approximately 70% of its electrocatalytic activity in the electro-oxidation of glucose in the presence of various electroactive molecules, such as ascorbic acid (3.5%), urea (2.3%), dopamine (2.5%), and uric acid (2%), which showed no major interface with the GOX/Te/CC electrode as shown in Fig. 5h because the fabricated electrode was immobilized with a special enzyme (GOX) that reacts specifically with glucose. As depicted in Fig. 5i, the current response to 10 μM glucose decreases by approximately 7% within the first 15 days due to a reduction in enzyme activity, after which it stabilizes for the following 15 days. Over the entire 30 days storage and testing period, no significant structural changes or loss of electrode conductivity were observed, and the electrode maintained >90% of its initial sensitivity. For the duration of individual in vivo or ex vivo experiments, the sensitivity remained stable with no observable drift, indicating that the electrode can be reliably used for repeated measurements within this timescale. The reproducibility of glucose sensing was assessed using three different sensors, as shown in Fig. 5j. The relative standard deviation (RSD) value of the three sensors was 2.25% in response to 10 μM glucose, indicating excellent stability.

3.4. Ex vivo ISF analysis

In this phase, ISF was extracted from the rat skin using the MN patch and subsequently analyzed using the GOX/Te/CC electrode in a standard three-electrode configuration. These measurements were performed ex vivo, with the MN patch serving solely for ISF collection and not containing any sensing element. After the glucose sensor was fully characterized electrochemically, the glucose concentration was determined in the rat ISF sample using a GOX/Te/CC electrode. Fig. 5(k) and (l) shows the chronoamperometry response curves and the corresponding glucose concentration levels in the rat ISF samples. The unknown glucose concentration detected by the GOX/Te/CC electrode was calculated using the linear regression equation of Fig. 5e. For rat sample 1, the glucose concentration was found at 0.99 μM and sample 2 at 1.13 μM, respectively. Meanwhile, the glucose level of the rat ISF samples was also measured using a commercially available glucometer. The data obtained from the GOX/Te/CC sensor correlated well with the commercial glucometer data as shown in Table 1; demonstrating the reliability of the results. Therefore, the GOX/Te/CC glucose sensor is very reliable as an ISF glucose detector.

The measured ISF glucose concentrations (Table 1) correspond to the diluted samples due to the extraction process (Section 2.4.2) and are not direct physiological values. Unit consistency has been maintained throughout this work, with μM values used for low concentration detection range studies and mM or mg dL⁻¹ used for physiological and in vivo data.

3.5. In vivo real-time sensing

In this study, continuous refers to uninterrupted real-time chronoamperometric measurements over the experimental monitoring period rather than extended wearable use. To conduct the real-time glucose monitoring, an MN patch was attached to a small tube (Fig. 6a). Male Sprague Dawley rats were used for the study, and the hair on the dorsal skin was removed prior to MN patch insertion (Fig. 6b). Previously weighed MN patches were pressed into the dorsal side of the rat and held in place for 10 min with the help of an occlusive tape (Fig. 5c), and then three electrode cells were placed as shown in Fig. 6d. In this configuration (Fig. 6d), the small tube surrounding the MN patch contained a PBS solution. It is important to note that the MN patch was not removed after puncture; instead it remained inserted in the rat skin throughout the measurement to ensure continuous ISF access for electrochemical sensing. The solid PEGDA MNs create microchannels in the skin, allowing ISF to diffuse into this PBS layer. This hydrated interface ensures good ionic conductivity and facilitates mixing of ISF with the buffer, enabling real-time glucose measurement by the GOX/Te/CC working electrode positioned within the tube. The tube also serves as a mechanical holder to stabilize the electrode MN assembly and minimize motion artifacts during in vivo monitoring. Fig. 6e shows the complete experimental setup for real-time glucose monitoring using a K-Lyte electrochemical workstation (Kanopy Techno Solutions Pvt. Ltd, India). During real-time monitoring, the MN patch remained attached to the rat skin for continuous ISF access, while the GOX/Te/CC working electrode, reference electrode and counter electrode were connected to the potentiostat for chronoamperometric monitoring. After the measurement, the insertion impression from the MN patch is seen in Fig. 6f, which disappeared within 30 min. There was no erythema or edema observed at the site of MN application. ISF volume was quantified gravimetrically by weighing MN patches before and after insertion. For the present MN patch (10 × 10 array, 100 needles), an average of 15 ± 1 μL of ISF were extracted within 10 min of insertion, which was sufficient for subsequent electrochemical analysis. This yield is comparable to previous reports for PEGDA MNs of similar geometry and was adequate for both ex vivo and in vivo glucose measurements.^45,46 The ISF was transferred to the sensor following the extraction process described in Section 2.4.2. ISF was collected from the MN patch following the extraction procedure mentioned earlier before being measured with the GOX/Te/CC. The glucose level fluctuations were detected via the chronoamperometry technique as shown in Fig. 6g. Corresponding to the linear regression equation in Fig. 4f, the average glucose concentration of 5.05 mM was detected and the response time was observed to be ∼95 s. At the same time, 5.17 mM glucose concentration was detected using the glucometer. These results suggest that the combined MN-assisted ISF extraction and GOX/Te/CC sensing approach has a favorable feature for continuous glucose monitoring, and glucose levels were effectively detected using the GOX/Te/CC electrode in rats.

Fig. 6 (a) MN patch adhered to the small tube container; (b) removal of rat's hair; (c) insertion of MNs into the shaved backs of healthy rats; (d) three-electrode cell inserted into the SD rat skin; (e) complete setup for real-time monitoring; (f) impressions seen on the back immediately after removal of the MN patch; (g) in vivo performance response of real-time monitoring.

3.6. Animal model study

The study utilized adult male Sprague Dawley rats weighing between 200–250 g. Fasting and fed state glucose levels were analyzed in ISF using MN-assisted sampling with a GOX/Te/CC sensing electrode (Fig. 7(a) and (b)), and compared with those in serum through blood withdrawal from the retro-orbital route as shown in Table 2. Briefly, the rats were given access to a normal pellet diet and water ad libitum for one week. The animals were then kept fasting for 12 hours with continued access to water. The MN patch-sensor assembly was then used for ISF collection and on-line detection and measurement of glucose (fasting state). The animals were then provided with a normal pellet diet and the change in serum and ISF glucose levels were analyzed after 2 hours (fed-state). A type II diabetes model was established by administering a high-fat diet along with a low-dose of streptozotocin (STZ). The animals were then provided with a commercially available high-fat diet (HFD) for three weeks. The diet was composed of 35.15% fat and 18.20% protein. 3120 kcal kg⁻¹ (60%) of energy was provided from the fat present in the HFD. Body weight, food intake, and serum glucose levels were monitored during this period. After three weeks of dietary manipulation, the animals were administered 35 mg kg⁻¹ of STZ in citrate buffer intraperitoneally. Animals with serum glucose values of ≥300 mg dL⁻¹ after 7 days of STZ administration were taken forward for the study. To further assess the efficacy of the MN integrated sensor device for analysis of ISF glucose gradient, fasting, and fed-state glucose levels were measured. The animals were fasted overnight. Serum and ISF glucose levels were then analyzed using blood collection (through retro-orbital route) and the sensor-integrated MN device as shown in Fig. 6(a)–(d), respectively. The animals were then provided with HFD for two hours, and glucose values were again analyzed similarly as shown in Table 3. It is worth noting that for Rat3, the fasting glucose value was slightly higher than the high fed state value (Table 3), which deviates from the expected physiological trend. This variation reflects normal inter-animal biological variability in glucose metabolism and feeding response, but does not affect the overall conclusion, as the sensor reliably captured physiologically consistent trends across the majority of animals, including the expected decrease following insulin administration.

Fig. 7 In vivo glucose monitoring performance for the basal state of animal model induction: (a) fed state & (b) fasting state.

Table 2 Basal glucose level values before animal model induction

S. no.	ISF glucose value (GOX/Te/CC sensor)		Serum glucose values (Trinder's method)
	Fed state (mg dL^−1)	Fasting state (mg dL^−1)	Fed state (mg dL^−1)	Fasting state (mg dL^−1)
1. Rat1	133.62	101.20	128.28	98.81
2. Rat2	117.86	98.50	125.46	93.86
3. Rat3	153.13	119.13	141.50	109.91

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Table 3 Glucose level values for animal model development

S. no.	ISF glucose value (GOX/Te/CC sensor)			Serum glucose values (Trinder's method)
	High fed state (mg dL^−1)	Fasting state (mg dL^−1)	Insulin administration (mg dL^−1)	High fed state (mg dL^−1)	Fasting state (mg dL^−1)	Insulin administration (mg dL^−1)
1. Rat1	380.58	269.75	227.52	359.67	286.47	217.73
2. Rat2	278.23	227.45	199.98	298.02	218.67	205.81
3. Rat3	185.09	194.41	164.03	169.81	184.79	156.83

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The electrochemical data shown in Fig. 7 and 8 represent the raw chronoamperometric current responses of the GOX/Te/CC sensor during in vivo monitoring under fasting, fed, and insulin administration states. These current signals were not directly reported as glucose concentrations; instead, they were converted using the calibration curve established in Fig. 5f. For each physiological condition, the current responses recorded over the monitoring window were averaged to obtain a representative value, which was then translated into absolute glucose concentration through the calibration equation and expressed in mg dL⁻¹ for physiological relevance. The final values derived from this conversion are summarized in Tables 2 and 3. The following day, 0.75 IU kg⁻¹ of insulin was administered intraperitoneally. Serum and ISF glucose levels were analyzed 1 h after insulin administration. Trinder's method was used for serum glucose level analysis. Briefly, 10 μL of the sample was added to 1.0 mL of enzyme reagent provided with the commercial kit. The enzyme reagent is composed of glucose oxidase, peroxidase, phenol, and phosphate buffer. The sample and reagent were mixed well and incubated at 37 °C for 10 min. The absorbance was then read at 505 nm in a UV spectrophotometer. An analysis of Tables 2 and 3 reveals that the glucose levels detected in ISF using the developed glucose sensor exhibit a similar trend observed in serum glucose levels measured using Trinder's method. The ISF glucose measurement (GOX/Te/CC sensor) using the microneedle-assisted sensor showed strong correlation with serum glucose measured by Trinder's method, across basal, diabetic, and insulin-treated states. This correlation confirms that ISF reliably reflects serum glucose dynamics. Although ISF values may exhibit a short physiological lag of 5–15 min relative to serum,⁴⁷ such delays did not significantly affect the accuracy of trend-based monitoring of this study. This correlation highlights the sensor's ability to reliably track changes in glucose concentration, demonstrating its potential for MN-assisted ISF sampling combined with electrochemical sensing for glucose monitoring. It is important to note that the sensor was specifically developed and validated for glucose detection in ISF. While serum glucose values were measured using Trinder's method to provide a physiological reference, direct glucose measurements from serum using the GOX/Te/CC sensor were not conducted in this study. Future work will explore extending the sensor application to serum and other biological fluids. Fig. 7, 8 and Tables 2, 3 follow the same trend of higher glucose levels in the fed state compared to the fasting state for all rats, with reductions following insulin administration. In all cases, the trends in ISF glucose closely mirror those in serum glucose, demonstrating the sensor's capability to track physiological changes. This close agreement supports existing evidence that ISF glucose levels provide a reliable substitute for serum glucose monitoring, with a typical physiological lag of 5–15 min due to glucose diffusion from capillaries to interstitial compartments. In our study, correlation between ISF and serum glucose was high in both normoglycemic and diabetic states, across fasting, fed, and insulin-treated conditions (Tables 2 and 3). This confirms that ISF can be performed repeatedly without significant discomfort or infection risk. Furthermore, this method eliminates the need for trained phlebotomists, making it highly suitable for wearable and home-based continuous glucose monitoring systems.

Fig. 8 In vivo glucose monitoring performance for the animal model development. (a) High fed state; (b) fasting state; and (c) insulin administration.

4. State of the art

In the past decade, significant research has been conducted on non-invasive glucose monitoring sensors, which have prominent applications in continuous glucose monitoring. In 2024, Hu et al. developed a fluorescent MN sensor patch designed for real-time, on-site glucose monitoring in ISF.⁴⁸ The sensor exhibited a LOD of 0.193 mM in artificial ISF, with a sensitivity of 0.029 Mm⁻¹ across the glucose range in human ISF. It also demonstrated a response time of 7.7 min in human serum. The sensor was successfully tested for continuous glucose monitoring during simulated meal and resting periods using ex vivo porcine skin. Further in vivo and clinical trials are needed to assess the MN sensor's performance. Bao et al. (2024) designed wearable MNPs utilizing double cross-linked hydrogel for the simultaneous detection of various type 2 diabetes (T2D)-related biomarkers, offering colorimetric signal readouts for both in vitro and in vivo analysis.⁴⁹ The present study introduces a novel approach, being the first to develop a 2D Te-based enzymatic glucose sensor for real-time glucose monitoring through an electrochemical method (in vivo). The material's high electrical conductivity facilitates rapid electron transfer, improving detection efficiency and sensitivity. Its large surface area provides more active sites for glucose interactions, enhancing sensitivity. Additionally, strong electrochemical activity arises from 2D Te's high catalytic properties and efficient charge carrier mobility, which accelerates electron exchange and promotes glucose oxidation. These properties enable accurate, reliable glucose monitoring, with a sensitivity of 0.10378 mA μM⁻¹ cm⁻² and a limit of detection of 0.357 μM. The analytical parameters obtained for the GOX/Te/CC electrode are compared to those in the previously reported literature as shown in Table 4. Overall, the sensor developed in this study demonstrates a remarkable LOD of 0.357 μM, significantly surpassing previous reports, and its exceptional rapid response time of ∼95 s enables real-time, highly sensitive glucose detection for in vivo monitoring.

Table 4 Comparison of MNP-based glucose sensors

Materials	MN patch configuration	Sensing material	Real application	LOD	Sensitivity	Response time	Ref.
PEGDA-PAM based hydrogel	11 × 11 array, shape: HMN, height: 446 ± 7 μm, radius: 281 ± 3 μm	—	Ex vivo (porcine skin)	0.193 mM	0.029 mM^−1	7.7 min	45
GelMA + PVA hydrogel	Shape: conical, height: 325.6 ± 1.1 μm, length of bottom circle: 152.3 ± 1.2 μm	GOX + HRP	In vivo (rat)	—	—	10 min	49
GelMA/nano (CMC-pHEA) hydrogel	11 × 11 array, shape: conical, base: 400 μm, height: 600 μm, tip: 600 μm	GOX	In vivo (mice)	—	—	—	50
Silk fibroin	Base width: ∼900 μm, height: ∼500 μm, tip diameter: ∼10 μm	Pt and GOD	In vivo (mice)	—	21.21 nA mM^−1	—	51
Ag–Pt composite hydrogel	Shape: HMN, base width: 200 μm, height: 800 μm, pitch: 500 μm	PEDOT:PSS and Ag–Pt nanoparticles	In vivo (rat)	0.9 mM	—	—	52
Alginate	11 × 11 array, shape: conical, height: 500 μm, base radius: 200 μm, tip radius: 5 μm	mEx4 and GOX	In vivo (mice)	—	—	—	53
Alg-ABA and CS	Shape: pyramidal width: 400 μm, height: 600 μm tip: 35 μm	Ex4 and GOX	In vivo (rat)	160.3 mg dL^−1	—	—	54
PVA/CMCS hydrogel	10 × 10 array shape: pyramid base width: 500 μm, tip length: 1000 μm	GOX + PVA-CS-HRP	In vivo (rat)			10 min	55
Silk	10 × 10 array, height: 650 μm, base: 200 μm, pitch: 500 μm	GF-monomer	In vivo (mice)	—	—	—	56
MeHA hydrogel	15 × 15 array, shape: square pyramid, length: 600 μm, base width: 300 μm, tip-to-tip distance: 500 μm	GOX/Prussian blue (PB)	In vivo (mouse)	—	0.020 ± 0.001 μA mM^−1	—	46
Polymeric	Shape: hexagonal diameter: 1.4 mm, height: 1 mm	GOX–rGO/PEDOT:PSS	In vivo (mice)	—	—	—	57
PEGDA	10 × 10 array shape: pyramid, length: 780 ± 10 μm, width: 350 ± 10 μm, tip: 40 ± 5 μm	GOX/Te/CC	In vivo (rat)	0.357 μM	0.10378 mA μM^−1 cm^−2	95 s	This work

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5. Conclusions

Minimally invasive MN-assisted sampling for CGM provides a practical solution for diagnosing and managing diabetes, by enabling painless and repeated access to ISF for subsequent electrochemical sensing. Although CGM devices are primarily used by individuals with type I diabetes (5–10% of the diabetic population),⁴⁵ painless, affordable, MNPs could expand their accessibility for people that are prediabetic, or have gestational diabetes, or type II diabetes. In this study, we have successfully developed a 2D tellurene-based enzymatic glucose sensor, using a combined MN-assisted ISF sampling and electrochemical sensing approach for glucose detection. The modified GOX/Te/CC electrode demonstrated 0.357 μM LOD, with a 0.10378 mA μM⁻¹ cm⁻² sensitivity. ISF-based glucose monitoring offers several practical advantages over serum sampling, including minimal invasiveness, reduced infection risk, smaller sample volume (∼15 μL), and suitability for continuous and wearable monitoring,⁴⁸ and these benefits make ISF a more efficient medium for real-time glucose monitoring compared to venepuncture-based serum analysis. While absolute ISF values may occasionally be lower or slightly delayed due to transport kinetics,⁴⁷ these limitations are minor compared to the clinical value of continuous, patient – friendly monitoring. To validate this reliability, in vivo studies with a microneedle-assisted sensor showed an average glucose concentration of 5.05 mM, with a rapid response time of ∼95 s, which is notably fast. For in vivo studies, ISF glucose readings obtained from the GOX/Te/CC electrode were compared with those from a commercial glucometer (Apollo Healthco Limited, India). The results highlight the prospects of GOX/Te/CC electrodes for glucose detection in body fluids, making them valuable for detection and monitoring purposes. The animal model study confirmed the sensor's ability to resolve physiologically relevant glucose fluctuations between fasting, fed, and insulin-treated states, with ISF glucose trends matching those observed in serum. The demonstrated equivalence of ISF and serum glucose trends underscores the clinical potential of ISF as a substitute for blood-based glucose monitoring in continuous and non clinical settings. While small temporal lags and slight differences in absolute concentration can occur due to tissue transport kinetics, these do not significantly impact the detection of rapid glucose excursions or long term trends. The combination of high patient comfort, low invasiveness, and strong correlation with serum values positions ISF monitoring as an efficient and reliable alternative for real-time diabetes management. While the present system demonstrates strong proof-of-concept performance for in vivo glucose monitoring, translation toward personal diabetes management would require miniaturization, wireless data transmission, and integration into a wearable platform. Future work will focus on developing a fully integrated MN sensor device with onboard electronics for power management and wireless communication. In conclusion, the successful modification of the GOX/Te/CC electrode, along with its excellent electrochemical performance, suggests its potential for future development into a fully integrated wearable device, particularly in the field of biosensing. The versatility showcased in this study expands the possibilities and opens new avenues for innovative developments in biomedical and biosensor technologies.

Conflicts of interest

The authors confirm that there are no known competing financial interests or personal relationships that could have influenced the outcome of this study.

Data availability

The data supporting this study's findings cannot be accessed publicly due to restrictions set by the data provider.

All experimental supporting data and procedures are available in the SI. See DOI: https://doi.org/10.1039/d5tb00804b.

Acknowledgements

This research was financially supported by the Indian Council of Medical Research (ICMR) with grant number ITR/2021/003371.

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