Novel PMI Marker Electrochemical Biosensor Based on Quantum Dot Dissolution Using a Dual Labeling Strategy
SEM, EDS, confocal microscopy, XPS, immunosensor probe CV
Figure 2 shows the SEM images of the modified Au surfaces GO-, Cys-GO- and Cys-GO/QD. The GO electrode exhibited a thick and robust morphology in the SEM image. For the Cys-GO electrode, a large spot was observed on the smooth SAM surface, confirming that GO was incorporated into Cys. The Cys-GO/QD-modified Au surface showed some ramifications due to QD binding after QD binding. However, QDs are small nanoparticles of 5 nm in size; thus, they are not visible in the images.
EDS was performed to characterize the surface of the Au/Cys-GO/QD electrode (Fig. 2). Survey spectra of Au/GO and Au/Cys-GO indicated the presence of C, O, S, and K and C, N, O, and S, respectively. However, the Au/Cys-GO/QD electrode was found to possess C, N, O, S and Cd. The presence of Cd on the surface indicated the successful immobilization of QDs on the immunosensor.
XPS analysis was performed to characterize the covalent bond of QD with Au/Cys-GO, as shown in Fig. S1. The C1S peak in Au/Cys-GO/QD and the N1S peak in Au/Cys-GO shifted to higher and lower energies, respectively, clearly indicate the formation of the covalent bond between the carboxylic acid of QD and the amine groups cysteamine in Au/Cys-GO.
Confocal microscopy was performed to further characterize the surface of the GAPDH immunosensor. Figure S2 shows the fluorescence images of Au/Cys and Au/Cys-GO/QD. Fluorescence was evident from the surface of the Au/Cys-GO/QD electrode due to QD emission, while the Au/Cys electrode showed no fluorescence. These results confirm that the QDs were successfully immobilized on the Au/Cys-GO surface.
The electrochemical behavior of the Au/Cys-GO/QD electrode was analyzed by CV. Figure S3 shows the CV profiles of Au/Cys-, Au/Cys-GO- and Au/Cys-GO/QD modified electrodes in PBS. A very small redox peak was observed for the modified Au/Cys electrode due to the absence of electroactive materials such as nanoparticles for the redox reaction. Moreover, for the modified Au/Cys-GO electrode, a significantly increased redox peak between −0.1 V and 0.3 V compared to Ag/AgCl was observed, suggesting that the introduction of GO improved the conductivity and increased the maximum current. After Au/Cys-GO/QD modification, the redox peak was located at the same position as observed for the Au/Cys-GO electrode, with an increase in peak current. The separation between the reduction and oxidation peaks has been calculated at 0.4 V, which implies a quasi-reversible electron transfer reaction. Table S1 shows the rate of increase of the redox peak, indicating a gradual increase in current following further steps. These analysis results confirmed that Cys-GO/QD was successfully immobilized on the Au electrode surface with an electroactive characteristic and can be used for the fabrication of the GAPDH immunosensor.
UV/Vis characteristics of QD dissolution
QD dissolution was analyzed by UV/Vis spectroscopy. Fig. S4 shows the UV/Vis spectra obtained for 10 µp/ml QD, 10 µg/ml QDs reacting with 5% H2O2 for 5 minutes, and 10 µg/ml QDs reacting with 1% H2O2 for 5 mins. As expected, the QDs showed a peak between 440 and 460 nm. After the QDs react with H2O2, the absorbance was remarkably reduced; the high concentration of H2O2 offered lower absorbance at the low concentration, confirming that QD dissolution was related to the presence of H2O2.
To obtain the best response in the analysis of real samples, the dilution rate of the antibody has been optimized. Figure S5 shows the anti-GAPDH antibody response to detection probes diluted in various ratios from 1:2000 to 1:40 under 1 ng/mL free GAPDH conditions. The current response increased significantly when the dilution factor Ab ranged from 1:2000 to 1:200 and did not increase further when the dilution factor was 1:200 or more. Thus, the optimal ratio of anti-GAPDH antibodies was finalized at 1:200. Optimizations of pH and incubation time are important for maximum sensitivity in the immunosensor based on antibody-antigen interaction. However, we did not try to optimize pH and incubation time in this study because our goal was to use this immunosensor at physiological pH (pH = 7.0 ~ 7.4). Regarding the incubation time of antibody-antigen interaction, previous results on antibody-antigen binding have shown incubations of 30 min to 2 h appropriate for maximal responses.25,26,27.. Considering that most antibody-antigen interactions should have a similar binding time, we did not attempt to optimize it. However, for maximum antibody-antigen binding, we used an incubation time of 4 h because the incubation step was performed at 4℃.
Analytical performance of the GAPDH immunosensor
DPV is a more sensitive technique than CV; thus, it was used for the quantification of free GAPDH. Figure 3 shows the DPV responses measured after dissolution of the QDs at different concentrations of free GAPDH between 10 fg/mL and 100 ng/mL. The response was measured after incubation of the immunosensor in PBS containing conjugated GAPDH and various concentrations of GAPDH for 10 h. As shown in Figure 3, an etching peak at −0.43 V relative to Ag/AgCl was observed due to the dissolution of Cd into Cd+. Then the immunosensor was placed with conjugated GAPDH and free GAPDH in PBS. Free GAPDH competed for binding to active sites in the anti-GAPDH antibody; thus, the amount of conjugated GAPDH bound to the active sites of the antibody was reduced and hence the peak Cd stripping reduction current decreased due to the presence of free GAPDH. Because the bounded GOx interacting with β-glucose in PBS solutions generated H2O2 30enzymatically generated H2O2 could dissolve metal Cd from QDs. Therefore, the Cd stripping peak gradually decreased and its intensity was proportional to the amount of free GAPDH.
In order to obtain a sensitive GAPDH detection, the analytical performances were carried out under optimized experimental conditions. When the concentration of free GAPDH was increased and the amount of conjugated GAPDH was fixed, a small amount of GOx coupled to antibody binding sites via the competitive binding strategy31generating a small amount of H2O2 for QD dissolution. As a result, the intensity of the Cd stripping peak did not decrease significantly. The stripping current response increased upon decreasing free GAPDH concentration in PBS and a direct linear relationship was observed between current response and free GAPDH concentration, allowing quantification of free GAPDH by this strategy. of dissolution.
Figure 4 shows the calibration plot using the intensity of the etching current obtained from Figure 3. The linear detection range of GAPDH was determined to be 10 fg/mL to 100 ng/mL. The linear dependence between intensity and concentration of free GAPDH produced a regression equation of I(y) = 0.0969 log x + 0.7769 with a correlation coefficient of 0.9802. The relative standard deviation was approximately 5.16% (n=5) at a GAPDH concentration of 1 ng/mL. Based on three measurements of the standard deviation of the white noise (95% confidence level, k=3, n=5), the detection limit was determined to be 2 fg/mL. All analytical parameters are summarized in Table S2. The analytical parameters obtained in this study and those of previously reported QD-based biosensors are compared in Table 1.32,33,34,35,36,37,38demonstrating the high sensitivity of the GAPDH sensor developed in this work.
Selectivity, stability and analysis of real samples
To investigate selectivity, competitive immunosensor binding was investigated using conjugated GAPDH and similar biomarkers such as PSA, human immunoglobulin G (hIgG), carcinoembryonic antigen (CEA) CRP, horseradish peroxidase (HRP) and thrombin (TB). Figure 5 shows the DPV responses obtained before and after a competitive reaction. Except for GAPDH, all proteins tested showed no significant current changes after competition, suggesting that the proposed immunosensor is highly selective and that the aforementioned proteins did not interfere with the detection of GAPDH. .
The stability of the fabricated sensor was determined by measuring the response to 100 pg/mL GAPDH for one month. After each measurement, the immunosensor was washed with a PBS solution and immersed in a 0.2 M Gly-HCl buffer solution (pH=2.8) for 5 min. After three washes with PBS solution, the immunosensor was stored in a dry state at 4°C until used for competitive binding of free and conjugated GAPDH. The current response did not decrease significantly (0.516, 0.523, 0.488, and 0.482) for one month, as shown in Figure 6. The immunosensor retained nearly 93.4% of its initial response for a period of ‘a month. These results indicated that the GAPDH immunosensor exhibited not only high selectivity but also long-term stability.
The applicability of this GAPDH immunosensor for the analysis of real samples was investigated using human blood serum. For this purpose, blood serum samples were prepared by adding 0.1 ng/mL ± 0.03 and 0.2 ng/mL of GAPDH. To calculate the GAPDH concentration, the standard addition method was used; the plot obtained is shown in Fig. 7. GAPDH concentration was detected at 0.1 ± 0.03 and 0.19 ± 0.03 ng/mL. The recovery was nearly 100% for the first sample and 95% for the second sample, indicating acceptable precision. This result confirms that the proposed GAPDH immunosensor can be applied for the detection of GAPDH in real human serum samples.