A molecularly imprinted electrochemical sensor based on conductive polymer hydrogel

Abstract

This invention provides a molecularly imprinted electrochemical sensor based on a conductive polymer hydrogel. The invention combines the antifouling properties of the hydrogel with molecularly imprinted polymers to prepare an antifouling molecularly imprinted hydrogel. This hydrogel possesses a three-dimensional porous structure, increased electrochemical active area, excellent conductivity, and antifouling properties. The molecularly imprinted electrochemical sensor prepared based on this hydrogel has the function of specifically recognizing structurally similar substances. It exhibits outstanding antifouling ability in eliminating non-specific adsorption of single proteins and complex media, and demonstrates strong long-term antifouling capability in complex media. Simultaneously, this sensor can accurately detect trace amounts of cortisol in real samples (saliva and sweat). This invention provides new inspiration for the development of ultrasensitive biosensors for complex media, particularly in practical implantable devices.

Description

Technical Field

[0001] This invention belongs to the field of medical materials technology, specifically relating to a molecularly imprinted electrochemical sensor based on conductive polymer hydrogel. Background Technology

[0002] Cortisol (CT) is a steroid hormone released by the human adrenal glands, which plays a role in regulating adrenal cortex dysfunction and chronic stress. It is crucial in regulating blood pressure, blood glucose levels, metabolism, and immune responses. Abnormal cortisol levels can lead to many diseases, including post-traumatic stress disorder, Cushing's syndrome, and weakened immune responses. Cortisol can serve as a biomarker to assess the severity of diseases caused by abnormal cortisol levels; therefore, rapid and reliable detection of cortisol levels is invaluable for the diagnosis and treatment of various diseases.

[0003] Traditional methods for detecting cortisol include enzyme-linked immunosorbent assay (ELISA) and liquid chromatography. These traditional methods have the following drawbacks:

[0004] 1) The testing process is cumbersome and data analysis is time-consuming, taking approximately 2 hours;

[0005] 2) Low detection sensitivity;

[0006] 3) Poor selectivity for detecting other similar steroid hormones;

[0007] 4) Specific recognition of cortisol depends on expensive and poorly recyclable biological antibodies.

[0008] While there are reports of using biosensors to detect biomolecules, most biosensors, when used in the detection of complex biological fluids (such as blood, serum, saliva, sweat, and urine), experience non-specific adsorption of biomolecules, cells, and other coexisting contaminants. This generates overwhelming signal interference and leads to sensor malfunction. The impact of biocontamination on the electrochemical signals at the biosensor interface severely reduces the sensitivity and selectivity of biosensors in detecting disease biomarkers, interfering with the accuracy of early disease diagnosis. Therefore, the development of anti-contamination materials is crucial for reducing false detections caused by non-specific adsorption. Summary of the Invention

[0009] To address the aforementioned problems, this invention provides a molecularly imprinted electrochemical sensor based on a conductive polymer hydrogel, specifically an antifouling molecularly imprinted electrochemical sensor based on a conductive sodium alginate-doped poly(3,4-ethyldioxothiophene) (PEDOT / SA) hydrogel, for the detection of cortisol in complex human body fluids. The antifouling molecularly imprinted hydrogel used in this invention, due to its three-dimensional pore structure, can specifically recognize target analytes and exhibits high inhibition of nonspecific adsorption.

[0010] This invention provides a molecularly imprinted polymer biosensor, which involves immersing a pretreated electrode in a mixed electrolyte solution to electrochemically reduce the electrode surface and generate a copper coating. After the copper-coated electrode dries, it is oxidized at a constant voltage in a mixed solution to generate a PEDOT / SA / CT hydrogel electrode. A specific solution is then drop-coated onto the surface of the hydrogel electrode, and the resulting sensor is named GCE / PEDOT / SA / CT. The sensor is then immersed in an ethanol solution to elute the cortisol template from the imprinted hydrogel, leaving cortisol-selective binding sites in the PEDOT / SA matrix, thus generating the molecularly imprinted polymer biosensor (GCE / PEDOT / SA-MIP).

[0011] The pretreatment described herein is polishing and cleaning using a cleaning solution.

[0012] The mixed electrolyte solution contains 0.01 g mL of electrolyte. -1 A 10% acetic acid solution of CuSO4.

[0013] The electrochemical reduction was carried out at a constant potential of -0.4V for 300s.

[0014] The mixed solution is a mixture containing SA (0.01g), KOH (0.028g), EDOT (10μL) and CT (0.8mg).

[0015] The constant voltage oxidation is oxidation at a constant voltage of 0.5V for 350s.

[0016] The specially prepared solution is a 5% perfluorosulfonic acid polymer solution (nafion solution).

[0017] The molecularly imprinted polymer biosensor prepared in this invention is used to detect cortisol;

[0018] The detection of cortisol described herein refers to the detection of cortisol in a liquid.

[0019] The molecularly imprinted polymer biosensor prepared in this invention can also be used to detect cortisol products.

[0020] The conductive PEDOT / SA hydrogel prepared in this invention is biocompatible. Its conductivity promotes electron transfer, enabling rapid and stable transmission of sensing responses. The specific binding of cortisol molecules inhibits electron transfer, increasing impedance (EIS mode) and decreasing the sensor's active area (CA mode). This hydrogel exhibits excellent conductivity, and by surrounding numerous imprinted cavities, it can more sensitively detect changes in the sensor's electrochemical performance, thereby improving sensing performance. The hydrogel's inherent superhydrophilic properties give the sensor excellent anti-fouling properties. Attached Figure Description

[0021] Figure 1 It curve of Cu coating deposited at a constant potential of -0.4V.

[0022] Figure 2 The it curve of Cu oxide coating under a constant potential of 0.5V.

[0023] Figure 3 Figure 1 shows the effect of preparation conditions on sensor impedance. Figure 2 shows the effect of cortisol mass in 5 mL of polymerization solution on sensor impedance; Figure 3 shows the effect of conductive polymer hydrogel polymerization time on sensor impedance; Figure 4 shows the effect of cortisol elution time on sensor impedance; and Figure 5 shows the effect of cortisol recombination time on sensor impedance.

[0024] Figure 4 SA / CT composite material (A), PEDOT / SA / CT hydrogel (B and C), SEM image of PEDOT / SA hydrogel after CT extraction (D); EDS map of PEDOT / SA / CT hydrogel (E).

[0025] Figure 5 Figure (A) shows the FTIR spectra of three hydrogels, where a is SA / CT hydrogel, b is PEDOT / CT hydrogel, and c is PEDOT / SA / CT; Figure (B) shows the contact angles of three electrodes, where B1 is the contact angle of bare GCE, B2 is the contact angle of GCE / PEDOT / CT, and B3 is the contact angle of GCE / PEDOT / SA / CT.

[0026] Figure 6 Three sensors at 50.0 mM [Fe(CN)6] 3- / 4- The curves are shown in the 0.1M KCl solution, where (A) is the CV sensor curve, (B) is the EIS sensor curve, and (C) is the CC sensor curve. Figure (D) shows the Q and t values. 1 / 2The relationships between them are as follows: a (black line) is bare GCE, b (red line) is GCE / PEDOT / SA / CT, c (blue line) is GCE / PEDOT / SA-MIP (CT washout), d (green line) is GCE / PEDOT / SA / CT (CT recapture), and e (purple line) is GCE / SA / CT.

[0027] Figure 7 The sensor's resistance to nonspecific adsorption was characterized. Figure (A) shows the nonspecific adsorption of different concentrations of single proteins by four different sensors. Sensors a, b, c, and d are sensor a: naked GCE; sensor b: GCE / PEDOT / CT; sensor c: GCE / PEDOT / SA / CT; sensor d: GCE / PEDOT / SA-MIP (after elution of CT), respectively. (B) shows the adsorption of naked GCE and GCE / PEDOT / SA / CT in tears, saliva, and sweat (red bars represent naked GCE, green bars represent GCE / PEDOT / SA-MIP). (C) shows the long-term adsorption of naked GCE and GCE / PEDOT / SA-MIP in tears, saliva, and sweat (V / V, 10%) over 30 days.

[0028] Figure 8 (A1-3) Laser confocal microscopy images of naked GCE and GCE / PEDOT / SA-MIP (B1-3); (A4) Demonstration of nonspecific adsorption of proteins on naked GCE; (B4) Schematic diagram of the resistance of the hydrophilic interface hydration layer to protein adsorption.

[0029] Figure 9 Figure (A) shows the EIS response of the electrochemical MIP sensor to different concentrations of cortisol; Figures (B) / (C) show the calibration linearity curves of the EIS signal against cortisol in PBS (B) and 10% (V / V) FBS (C); (D) shows the CC response of the sensor to cortisol; Figures (E) / (F) show the calibration linearity curves of the electrochemical active area against cortisol in PBS (E) and 10% (V / V) FBS (F). Curve ai represents the concentrations of cortisol at 0 and 10%, respectively. -12 -10 -6 M. (G) 50-cycle cyclic voltammetry curves of the MIP sensor; the sensor measures progesterone (a), estriol (b), estrone (c), glucose (d), urea (e), ascorbic acid (f), albumin (g), globulin (H), lysozyme (I), dopamine (j), lactate (k), Na3PO4 (l), K2CO3 (m), cortisol (n), and blank (o) using EIS (H) and CC (I).

[0030] Figure 10Intra-electrode and inter-electrode reproducibility of biosensors, where (A) is the intra-electrode reproducibility of detecting 1.0 nM cortisol seven times with a single electrode, and (B) is the inter-electrode reproducibility of detecting 1.0 nM cortisol using seven different electrodes. Detailed Implementation

[0031] The PEDOT / SA hydrogel-modified sensor provided by this invention first involves electrodepositing copper onto a pretreated glassy carbon electrode (GCE) surface, then oxidizing it to copper ions that diffuse into the electrolyte, resulting in the gelation of a PEDOT / SA / CT hydrogel. The conductive polymer micro-hydrogel particles repel each other and flow due to the negatively charged carboxyl groups (from SA) on their surface. A sufficiently high concentration of metal ions increases the ionic strength of these dispersions, shielding electrostatic repulsion and causing the particles to aggregate, forming a physically cross-linked and stable hydrogel.

[0032] The pretreatment described herein is polishing and cleaning using a cleaning solution.

[0033] The mixed electrolyte solution has a concentration of 0.01 g / mL. -1 A mixed solution of CuSO4 and acetic acid at a ratio of 1:10 v / v.

[0034] The electrodeposition was performed using IT technology at a constant potential of -0.4V for 300 seconds.

[0035] The constant voltage oxidation is oxidation at a constant voltage of 0.5V for 350s.

[0036] Glassy carbon electrode: GCE; Poly(3,4-ethyldioxothiophene): PEDOT; Cortisol: CT; Sodium alginate: SA; Fluorescently labeled bovine serum albumin: FITC-BSA; Lysozyme: LYS; Bovine hemoglobin: Hg; Human serum albumin: HAS; Cyclic voltammetry: CV; Chronocoulometric method: CC; Electrochemical impedance spectroscopy: EIS; Fourier transform infrared spectroscopy: FTIR; Molecularly imprinted polymer: MIP.

[0037] The relevant information regarding the chemicals and instruments used in the embodiments of this invention is as follows:

[0038] 3,4-Ethiothiophene (EDOT), cortisol, sodium dihydrogen phosphate dihydrate (NaH2PO4.2H2O), disodium hydrogen phosphate dodecahydrate (Na2HPO4.12H2O), potassium ferrocyanide (K4Fe(CN)6) and potassium ferricyanide (K3Fe(CN)6) were purchased from Aladdin Reagent Co., Ltd. (Shanghai, China). Sodium alginate (SA) was purchased from Tianjin Bailuns Biotechnology Co., Ltd., copper sulfate pentahydrate (CuSO45H2O) was purchased from Tianjin BASF Chemical Co., Ltd., and human serum albumin (HSA), human hemoglobin (Hg), lysozyme (LYZ), and FITC-labeled bovine serum albumin (BSA) were purchased from Sangon Biotech Co., Ltd. (Shanghai, China).

[0039] The sensor was analyzed using cyclic voltammetry (CV), chronocoulometric spectroscopy (CC), and electrochemical impedance spectroscopy (EIS) in the presence of a redox probe. The analysis was performed using a CHI660E electrochemical workstation (Shanghai Chenhua) equipped with a three-electrode system, including a platinum counter electrode (auxiliary electrode), an Ag / AgCl (saturated KCl) reference electrode, and a modified glassy carbon electrode (GCE, working electrode). The material properties of the MIP were characterized using a JSM-IT500 (Japan) scanning electron microscope, energy dispersive spectroscopy (EDS), a contact angle meter (DSA 25S, Germany Kruss), and Fourier transform infrared spectroscopy (FTIR) (Thermo Fisher Nicolet IS10).

[0040] Electrochemical CV measurements were performed between 0 and 0.5 V at a scan rate of 100 mV·s. -1 EIS measurements were performed by applying a 10 mV pulse (frequency range from 0.2 Hz to 100 kHz). Chronocoulometric (CC) tests were conducted directly within the range of -0.5 to 0.8 V with a pulse width of 0.25 s. All electrochemical tests were performed at 50 mM [Fe(CN)6]. 4- / 3- Recorded in 0.1M KCl solution.

[0041] The present invention will now be described in detail with reference to the embodiments and accompanying drawings.

[0042] Example 1: Fabrication of a molecularly imprinted electrochemical sensor based on conductive polymer hydrogel

[0043] First, the glassy carbon electrode (GCE) was polished, cleaned, and pretreated. Then, it was subjected to a process using copper sulfate (0.01 g·mL⁻¹). -1 A mixed electrolyte consisting of acetic acid (1:10 v / v) and acetic acid was used to electrochemically reduce copper on a GCE surface. A purplish-red copper coating was obtained using IT techniques at a constant potential of -0.4 V for 300 s. Figure 1 ).

[0044] After drying, the copper-coated electrode was immersed in a 5 mL solution containing SA (0.01 g), KOH (0.028 g), EDOT (10 μL), and CT (0.8 mg), and oxidized for 350 s under a constant voltage of 0.5 V to generate a PEDOT / SA / CT hydrogel. Figure 2 ).

[0045] 10 μL (5%) of Nafion was dropped onto the deposited electrode surface and stored in a humid environment for 30 min to form a hydrogel.

[0046] A distinct transparent gel layer was observed on the electrode surface, forming a well-defined hydrogel. The sensor was then immersed in an ethanol solution for 20 min to extract the cortisol template from the imprinted hydrogel, leaving cortisol-selective binding sites in the PEDOT / SA matrix for cortisol detection. The resulting molecularly imprinted sensor was defined as GCE / PEDOT / SA-MIP.

[0047] To obtain the best sensing performance of the molecularly imprinted electrochemical sensor, the mass of cortisol in 5 mL of polymerization solution, the polymerization time of the conductive polymer hydrogel, and the cortisol elution and re-recognition time were optimized. Figure 3 As can be seen from Figure A, the mass of cortisol in the 5 mL polymerization solution played a role in the sensor construction process; therefore, the mass of cortisol in the polymerization solution was optimized. With increasing cortisol mass, the sensor impedance gradually increased until it remained constant. Consequently, a hierarchical porous nanostructure was obtained using 0.8 mg of cortisol. Doping the hydrogel with more non-conductive cortisol led to an increase in the hydrogel's impedance. The sensor impedance remained unchanged only when the cortisol doping level reached its maximum.

[0048] The polymerization time of CPHs was optimized to achieve high sensitivity. Figure 3 As shown in Figure B, the present invention selects 300 seconds as the optimal time for the polymerization of conductive polymer hydrogels. Figure 3 The C-axis indicates that the optimal CT elution time is 20 minutes, and the optimized CT rebinding time is 10 minutes. Within the selected time, the impedance value of the modified electrode does not change, indicating that the modified electrode has reached its final stable state.

[0049] Example 2: Physicochemical properties of the prepared electrochemical sensor

[0050] Scanning electron microscopy (SEM) was used to examine porous conductive polymer hydrogels, such as... Figure 2As shown, the SA / CT composite material (without PEDOT) exhibits a planar film with almost no ripples. The preparation method of the SA / CT composite material is the same as that of the PEDOT / SA / CT preparation electrochemical method, the only difference being that 3,4-ethyldioxothiophene (EDOT) is not added to the deposition solution. In contrast, the PEDOT / SA / CT hydrogel of this invention consists of many particles that stack and cross-link with each other at a lower magnification, forming a non-uniform microstructure. Figure 4 B). When the magnification is large ( Figure 4 C), the PEDOT / SA / CT hydrogel exhibits a three-dimensional porous microstructure. For the PEDOT / SA hydrogel ( Figure 4 D, without CT), its morphology is almost identical to that of the cortisol-imprinted hydrogel (PEDOT / SA / CT), with no significant changes observed. The porous structure ensures high water content and expanded surface area at the interface of the hydrogel, thereby promoting the transport of the imprinted cortisol molecule at the interface.

[0051] This invention also recorded energy-dispersive spectroscopy (EDS) experiments to compare the compositional distribution of various elements (C, O, Na, Cu, and S) in PEDOT / SA / CT and PEDOT / SA (after CT elution) hydrogels. The results show that the MIP sensor was successfully fabricated. Figure 4 E, Table 1).

[0052] Table 1: Elemental Content in Molecularly Imprinted Electrochemical Sensors

[0053]

[0054] The prepared MIP sensor was analyzed and characterized by FTIR spectroscopy to demonstrate the successful synthesis of the conductive polymer hydrogel. Figure 5 A). The infrared spectrum of SA / CT (curve a) shows the stretching vibration peak of -OH at 3452 cm⁻¹. -1 The tensile vibration peak of -CH2 is at 2867 cm⁻¹. -1 and 2932cm -1 The tensile vibration peak of COC is at 1245 cm⁻¹. -1 and 1386cm -1 The tensile vibration peak of C-OH is at 1108 cm⁻¹. -1 A 1624 cm⁻¹ was found in the PEDOT / CT infrared spectrum (curve b). -1 The peak value for tensile vibration of C=C is 679 cm⁻¹. -1 The peaks represent the tensile vibrations of CS and are characteristic peaks of PEDOT. The infrared spectrum of PEDOT / SA / CT (curve C) shows COC, CS, C-OH, and other characteristic peaks, confirming the successful preparation of PEDOT / SA / CT.

[0055] from Figure 6 As shown in curve A, the peak current of the redox peak of GCE / PEDOT / SA / CT (curve b) is significantly greater than that of bare GCE (curve a). This is because the SA / CT-doped conductive polymer PEDOT hydrogel has good conductivity. After eluting the template CT from the hydrogel, the peak current of GCE / PEDOT / SA (curve c) further increases, while the peak potential difference of the redox peak decreases. This is because removing the non-conductive cortisol yields more holes and a larger active area. As expected, since the non-conductive cortisol molecules severely block interfacial electron migration, the subsequent rebinding of the sensor with cortisol significantly reduces the peak current (curve d), indicating that the MIP sensor effectively captures cortisol.

[0056] The non-conductive properties of traditional hydrogels greatly limit their application in sensors. However, by introducing conductive polymers into the hydrogel structure, excellent conductivity has been achieved. For example... Figure 6 As shown in Figure B, the impedance (R) generated by the formation of the conductive polymer hydrogel (427.2 Ω) is much smaller than that of previously reported hydrogels (approximately 7000 Ω) because the conductive polymer hydrogel accelerates electron transfer and increases interfacial conductivity. With successful removal of the template molecule, the impedance of the GCE / PEDOT / SA-MIP sensor significantly decreases (119.8 Ω). Subsequently, the MIP sensor again identifies and captures cortisol, leading to a gradual increase in impedance (186.8 Ω), indicating that the effective capture of cortisol by holes at the sensing interface blocks interfacial electron exchange.

[0057] The electroactive specific surface area of ​​the sensor GCE / PEDOT / SA-MIP of the present invention is 9.81 times and 2.66 times that of bare GCE and GCE / SA / CT, respectively. Figure 6 (C and D). The results show that the SA / CT-doped macroporous PEDOT hydrogel significantly expands the active surface area. More importantly, cortisol removal and recapture lead to an increase and a decrease in the effective area of ​​the sensor, respectively. These results confirm the successful construction of the molecularly imprinted sensor, consistent with CV and EIS analyses. Furthermore, changes in EIS and CC signals associated with cortisol binding can be used for the specific detection of cortisol.

[0058] Example 3: Anti-pollution performance evaluation

[0059] Hydrophilic interfaces can inhibit the non-specific adsorption of proteins by water-mediated hydrophobicity and hydration. Therefore, the hydrophilicity of an interface is considered one of the important indicators for evaluating its antifouling performance. The water contact angles of bare GCE, GCE / PEDOT / CT, and GCE / PEDOT / SA / CT were tested. Figure 5As shown in Figure B, the contact angle of GCE / PEDOT / CT (56.84°) is much smaller than that of bare GCE (84.21°). The contact angle of GCE / PEDOT / SA / CT is further reduced to 16.92°, highlighting the superhydrophilicity and strong antifouling ability of the interface.

[0060] Example 4: Protein Adsorption Capacity of the Sensor

[0061] The antifouling performance of conductive polymer hydrogel sensors was evaluated by measuring the change in EIS (ΔR / R0) before and after culture in single proteins and complex human fluids. HSA, LYS, and Hg were selected as representative proteins due to their different molecular weights and isoelectric points. Figure 7 As shown in Figure A, in these protein solutions of different concentrations, GCE / PEDOT / SA / CT (sensor c) and GCE / PEDOT / SA-MIP (sensor d, after removal of the template molecule CT) retained most of their initial impedance values ​​with very small rates of change. Impressively, GCE / PEDOT / SA / CT and GCE / PEDOT / SA-MIP exhibited almost identical resistance to non-specific protein adsorption, demonstrating that template molecule removal does not affect the excellent anti-fouling ability of the MIP sensor. However, the significantly higher ΔR / R0 values ​​of bare GCE (sensor a) and GCE / PEDOT / CT (sensor b) indicate that non-specific protein adsorption or biofouling occurred at the interface of these sensors.

[0062] To intuitively perceive the sensor's anti-fouling performance, the adsorption of FITC-labeled BSA on bare GCE and GCE / PEDOT / SA-MIP was studied by recording laser confocal images. Figure 8 The test electrode was incubated in BSA solution for 30 min. A large amount of bovine serum albumin was adsorbed on the bare GCE surface, while no significant protein adhesion was observed on the hydrogel-modified sensor, indicating that the functionalized conductive hydrogel-modified surface can effectively inhibit / eliminate protein adsorption.

[0063] Example 5: Non-specific adsorption detection by the sensor

[0064] The specific trapping of cortisol by the MIP sensor hinders the transfer of electrons to the sensor surface, leading to an increase in the sensor's electrochemical impedance and a decrease in the electroactive surface area. Based on the changes in the electrochemical signal of the GCE / PEDOT / SA-MIP, EIS and CC techniques were used to determine cortisol concentration with high selectivity and high sensitivity.

[0065] Under optimal conditions, the sensor's EIS response to different concentrations of cortisol is as follows: Figure 9As shown in Figure A, the impedance increases with increasing cortisol concentration until it almost stabilizes. Since the cavities formed by the molecularly imprinted polymer correspond to electron transport channels, as more cortisol molecules recombine with the imprinted molecular cavities, they hinder the electrochemical probe from reaching the sensor surface. Linear regression equation ( Figure 9 B) is R = 33.14lgC + 552.6(R) 2 =0.992), with a range of 10. -12- 10 -8 mol L -1 The limit of detection (LOD) was as low as 0.314 pM. This high sensing performance is attributed to the three-dimensional porous microstructure of the conductive polymer hydrogel (high surface area, excellent conductivity and antifouling ability) and the advantages of molecular imprinting technology (high specificity and high selectivity). To evaluate the suitability of the prepared MIP sensor, the same experimental tests were performed in 10% (v / v) FBS. Figure 9 C). The linear equation is obtained as (R = 31.20lgC + 534.2, R). 2 =0.996). Clearly, a similar calibration curve ( Figure 9 As shown in B and C, the MIP sensor demonstrates anti-fouling performance and excellent practical potential in complex media.

[0066] Figure 9 The D-display shows that the CC signal of the MIP sensor decreases with increasing cortisol concentration. The electroactive surface area (A) of the sensor exhibits a good linear relationship with the logarithm of cortisol concentration. The linear regression equation is A = -0.4349lgC + 0.8767(R²). 2 =0.994), with a range of 10. -12 -10 -7 The LOD was as low as 0.131 pM (S / N = 3). The significant improvement in sensor recognition efficiency is attributed to the use of conductive polymer hydrogels and molecularly imprinted polymers, which effectively bind target molecules by providing a large surface area and more accessible imprinted cavities. The same experimental scenario was performed using CC technology in 10% (V / V) FBS to further validate the sensor's anti-fouling capability. A linear regression equation was obtained (A = -0.4421lgC + 0.8937, R0). 2 =0.995). It is worth noting that the sensing performance of the MIP sensor remains almost unchanged in complex media. Figure 9 The calibration curves in E and F are similar, and the linear range is the same. This is due to the specificity of the molecularly imprinted sensor enhanced by the anti-sewage gel.

[0067] Real bodily fluid samples (sweat and saliva) were provided by three healthy 25-year-old volunteers (Shandong, China) at the laboratory. Each subject's sample was centrifuged at 5000 rpm for 2 minutes, and the supernatant was collected for electrochemical analysis. Samples were stored at -28°C before analysis. These electrochemical results were compared with those obtained using the ELISA method.

[0068] The antifouling performance of the MIP sensor was further investigated in a series of complex bodily fluids, including sweat, tears, and saliva, with different dilution ratios (V / V). Figure 7 As can be seen from Figure B, even after incubation for 30 minutes in 100% sweat, tears, and saliva, only very small changes in EIS were observed. When incubated in the 1% composite solution, the signal of the bare glassy carbon electrode increased more than twofold (2.57, 5.11, and 3.57, respectively). These results clearly demonstrate that the MIP sensor based on conductive polymer hydrogel exhibits excellent resistance to nonspecific adsorption in real, complex samples and high-concentration protein samples.

[0069] We further compared the long-term antifouling ability of the sensor after immersion in 10% (v / v) human body fluid for more than 30 days. Figure 7 C). Obviously, after continuous immersion in 10% human body fluid, the non-specific adsorption of the bare electrode is very serious.

[0070] The sensor based on conductive polymer hydrogel developed in this invention still exhibits significant anti-fouling capabilities; even after immersion for up to 30 days, the ΔR / R0 value remains almost unchanged.

[0071] The sensor of this invention was used to determine cortisol in complex human fluids (saliva and sweat), and the cortisol content in actual samples was studied using electrochemical techniques (EIS and CC). The results were compared with ELISA results (Table 2). The detection results of EIS and CC methods corroborated each other, with a concordance rate (EIS / CC) between 97.83% and 107.1% (Table 2). This indicates that the detection results of this sensor are in good agreement with the ELISA results, demonstrating that the electrochemical MIP sensor is practically feasible for the detection of cortisol in biological samples.

[0072] Table 2: Cortisol data detected in sweat and saliva using ELISA and the sensor of this invention (n=3)

[0073]

[0074]

[0075] Example 6: Sensor stability, selectivity, and reproducibility

[0076] The stability of the MIP sensor has been proven through long-term anti-fouling tests. Figure 7 C). Simultaneously, CV measurements were used to verify the sensor's stability, with 50 cycles of cyclic voltammetry curves superimposed (…). Figure 9 G) indicates that the sensor based on conductive polymer hydrogel has good stability.

[0077] Within-group and between-group analyses were performed to test the reproducibility of the constructed sensor. For cortisol (10... -9 M) determination, the relative standard deviation (RSD) within and between test groups (M) Figure 10 The percentages were 0.28% and 3.12% (n=7), respectively. These results demonstrate that the CPH-based sensor exhibits good reproducibility.

[0078] The selectivity of the MIP sensor was evaluated by comparing its electrochemical responses to structurally similar 1.0 nM steroid hormones (including cortisol, progesterone, estriol, and estrone) as well as 100 nM glucose, urea, ascorbic acid, albumin, globulin, lysozyme, dopamine, lactate, Na3PO4, and K2CO3. Figure 9 The sensor exhibits a maximum response to the target cortisol, while its response to other potential interfering substances is negligible. These results demonstrate that the MIP sensor possesses high specificity for cortisol identification from other analogues and coexisting substances in real samples. This is due to the high selectivity and accuracy of the MIP, where the cavity and target have a "key-lock" specific recognition relationship.

[0079] In summary, the sensor of the present invention has high specificity and outstanding anti-contamination ability, ensuring accurate detection of trace amounts of cortisol in actual samples (saliva and sweat).

Claims

1. A molecularly imprinted polymer biosensor, characterized in that, The molecularly imprinted polymer biosensor involves immersing a pretreated electrode in a mixed electrolyte solution to electrochemically reduce the electrode surface and generate a copper coating. The coated electrode was obtained, and after the copper coated electrode dried, it was placed in a mixed solution for constant voltage oxidation to generate an imprinted hydrogel. A hydrogel electrode was obtained, and Nafion solution was drop-coated onto the surface of the hydrogel electrode. Then, it was immersed in an ethanol solution to elute the cortisol template from the imprinted hydrogel, thus obtaining the molecularly imprinted polymer biosensor. The mixed electrolyte solution is a 10% acetic acid solution containing CuSO4 at a concentration of 0.01 g / mL; The mixed solution is a mixture containing sodium alginate, KOH, EDOT and cortisol; The constant voltage oxidation refers to oxidation at a constant voltage of 0.5V for 350 seconds.

2. The molecularly imprinted polymer biosensor as described in claim 1, characterized in that, The electrochemical reduction was carried out at a constant potential of -0.4V for 300s.

3. The application of the molecularly imprinted polymer biosensor of claim 1 in the detection of cortisol.

4. The use of the molecularly imprinted polymer biosensor of claim 1 in the preparation of articles for detecting cortisol.

5. The application as described in claim 4, characterized in that, The product described is used to detect cortisol in human body fluids.

6. A product for detecting cortisol, characterized in that, The article comprises the molecularly imprinted polymer biosensor of claim 1.

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