A molecularly-imprinted-electrochemical-sensor modified with nano-carbon-dots with high sensitivity and selectivity for rapid determination of glucose

Analytical Biochemistry 555 (2018) 42–49

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A molecularly-imprinted-electrochemical-sensor modified with nanocarbon-dots with high sensitivity and selectivity for rapid determination of glucose

T

Wei Zhenga,1, Haiyan Wub,c,d,1, Yan Jiangc, Jicheng Xuc,d, Xin Lia, Wenchi Zhanga, Fengxian Qiua,∗ a

School of Chemistry and Chemical Engineering, Jiangsu University, Zhenjiang 212013, China School of Chemical and Environmental Engineering, Jiangsu University of Technology, Changzhou 213001, China Institute of Medicine & Chemical Engineering, Zhenjiang College, Zhenjiang 212003, China d School of Environment and Safety Engineering, Jiangsu University, Zhenjiang 212013, China b c

A R T I C LE I N FO

A B S T R A C T

Keywords: Carbon dots Molecularly imprinted electrochemical sensor Glucose 3-Aminobenzeneboronic acid

In this work, a novel molecularly imprinted electrochemical sensor (MIECS) based on a glassy carbon electrode (GCE) modified with carbon dots (CDs) and chitosan (CS) for the determination of glucose was proposed for the first time. The use of the environmental-friendly CDs and CS as electrode modifications improved the active area and electron-transport ability substantially, while 3-aminobenzeneboronic acid was used as a functional monomer and glucose as template for the fabrication of molecularly imprinted polymer (MIP) film to detect glucose via differential pulse voltammetry. Transmission electron microscope, Fourier transform infrared spectroscopy, energy dispersive x-ray spectrometry, cyclic voltammetry and electrochemical impedance spectroscopy (EIS) were applied to characterize the fabricated sensor. Experimental conditions such as molar ratio of functional monomer to template, volume ratio of CDs to CS, incubation time and elution time were optimized. By using glucose as a model analyte, the MIECS had two assay ranges of 0.5–40 μM and 50–600 μM, and fairly low limit of detection (LOD) of 0.09 μM (S/N = 3) under the optimized conditions. The MIECS also exhibited excellent selectivity, good reproducibility, and stability. The proposed sensor was successfully applied to a preliminary test for glucose analysis in real human blood serum samples.

Introduction Glucose is not only one of the most important metabolites in human body, but also the source of energy [1,2]. In recent years, just like all the countries in the world, the prevalence of diabetes in China is gradually rising, and the impact of diabetes on people's health is becoming more serious. Although the diabetes prevalence rate in China was low, but the number of patients with diabetes has been ranked second in the world, increasing at an alarming rate [3]. Therefore, rapid and efficient detection of glucose concentration in human blood is particularly important for the diagnosis of diabetes mellitus [4]. Recently, there are many methods to detect glucose, such as spectrophotometry [5,6], colorimetry [7–9], chromatography [10], surface enhanced Raman scattering [11], and electrochemical sensing method [12–14], etc. Among these methods, electrochemical sensor attracts more attention owing to its portability, easy operation, fast response,

∗

1

Corresponding author. E-mail address: [email protected] (F. Qiu). These authors contributed equally to this work and should be considered co-first authors.

https://doi.org/10.1016/j.ab.2018.06.004 Received 10 April 2018; Received in revised form 5 June 2018; Accepted 5 June 2018 Available online 14 June 2018 0003-2697/ © 2018 Elsevier Inc. All rights reserved.

high sensitivity, and easier to be miniaturized [15]. The enzyme glucose sensor is the most commonly utilized electrochemical sensor and has been widely studied. Although the enzyme-based sensors show relatively high selectivity, they suffer from some disadvantages, such as chemical changes and thermal deformation caused by the inherent nature of the enzyme in blood samples and other oxidable species serious interference, which may limit their application [16]. Non-enzymatic glucose sensors based on electro catalysis and oxidation of glucose directly on the electrode surface are not only convenient [17], but also have the advantage of avoiding the shortcomings of enzyme electrodes, so they have received a lot of attention over the years [18–20]. Unlike normal amino acids and lipids, glucose are hydrophilic species with multiple hydroxyl groups. It is difficult to be extracted from water by traditional pre-treatment methods and thus causes the cross-interference of the determination signals. Moreover, the differentiate of the structure similarity carbohydrates mainly depends on the

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national drug group chemical reagents Co., Ltd (Shanghai, China). The pledget was used without further purification. Ascorbic acid, D-fructose and dopamine were obtained from Shanghai Aladdin Industrial Corporation (Shanghai, China). All other reagents are analytical reagents. Ultrapure water (18.2 MΩ cm−1) was used throughout all experiments.

physicochemical signal changes (i.e. swelling/shrinking degree, diffraction and conductivity) of the receptor exposed to the sample [21]. Therefore, the identification of glucose in a multi-carbohydrate system is not satisfied because of the nonspecific electro catalytic activity of the active material to the analyte [22]. Metals and its alloys are always utilized to improve the electrocatalytic activity and selectivity toward the direct oxidation of glucose [23]. In addition, the metal-based sensors may sometimes poison the surface of sensor with oxidation products [23]. Therefore, it is of great importance to find an environmental-friendly, convenient method to detect glucose with high selectivity and sensitivity. Molecularly imprinted polymer (MIP) can be used to recognize some particular target analytes with high selectivity [24]. Molecular imprinting technology (MIT) is an interdisciplinary technique that involves polymer chemistry, materials science, chemical engineering and biochemistry [25]. MIT is increasingly favored by researchers with its high specificity, stability and reusability [26]. Molecularly imprinted electrochemical sensor (MIECS) combines MIT and electrochemical sensor [27], which is widely used in analytical chemistry field [28,29]. However, the sensitivity of MIECS is relatively low due to the poor conductivity and electrocatalytical activity in previous report [30]. Thus, many efforts have been made to enhance its conductivity by using conductive nanomaterials, such as carbon-based materials [31]. As a new carbon nanomaterial, carbon dots (CDs) have gained more and more interest amongst researchers to study. Fluorescent CDs are form of spherical particles, which size is generally within 10 nm. The surface of CDs usually contains a large number of oxygen containing groups, but different synthetic methods and surface treatments can be used to make the oxygen content of the surface differently [32]. In addition, due to the high fluorescence stability, suitable conductivity, low particle and toxicity that make the fluorescent CDs have attracted a lot of attention from scientists since its discovery [33–35]. Nevertheless, the research of CDs was used to fabricate the electrochemical sensors to monitor analytes with high selectivity and sensitivity still remains at an early stage [31]. Chitosan (CS) is a kind of natural, nonpolluting natural polysaccharide polymer [36]. It is a natural polymer membrane material exhibits good film forming property and has shown great potential in electrode modification [36,37]. Owing to its relatively poor conductivity, CS was usually combined with conducting materials such as carbon-based materials, redox mediators, metal nanoparticles, and ionic liquidate form conductive film [38,39]. Herein, we report a novel, simple, sensitive and selective MIECS for the quantification of glucose. It is aimed at the development of a CS and CDs composite film modified MIECS for the determination of glucose with high sensitivity. The CDs synthesized by pollution-free green substances. The electrode was modified with CDs-CS solution to improve its electrochemical properties. A MIP film was fabricated on the modified electrode surface via elecrtopolymerization method of 3aminobenzeneboronic acid (APBA) in the presence of glucose template molecules. The proposed new MIECS demonstrated high sensitivity, selectivity, good reproducibility, stability. Furthermore, the proposed MIECS was successfully applied to a preliminary test for glucose analysis in human blood serum samples with satisfactory results, which are verified by the records obtained from hospital's commercial instrument. Importantly, enzyme free, metal free, low cost and no tedious sample pretreatment procedure made it possible to apply the sensor in clinical, biological and other complex matrix, such as biochemistry monitoring.

Instruments and measurements Cyclic voltammetry (CV), electrochemical impedance spectroscopy (EIS) and differential pulse voltammetry (DPV) were carried out on a CHI 660E electrochemical workstation (Shanghai ChenHua Instruments Co., China). A conventional three-electrode system was used for all electrochemical experiments, which consisted of a platinum wire as the auxiliary electrode, an Ag/AgCl/saturated KCl as the reference electrode, and a bare or modified GCE as the working electrode. The pH measurements were carried out on a pHS-3C exact digital pH metre (Shanghai Mettler-Toledo Instruments Co., Ltd), which was calibrated with standard pH buffer solutions. The fluorescence spectra were recorded by a Cary Eclipse fluorescence spectrophotometer (Australia Varian Co., Ltd). The UV–vis absorption was performed on a UV-2450 spectrophotometer (Japan Shimadzu Co., Ltd). Transmission electron microscopy (TEM) was performed on a JEM-2100 electron microscope (JEOL, Ltd., Japan). Energy dispersive x-ray spectrometry (EDX) was performed an energy-dispersive X-ray analyzer (EDX, Thermofisher NSS7). Fourier transform infrared (FT-IR) spectrum was recorded using a Nicolet Nexus 470 FT-IR spectroscope (Nicolet Instrument Co., USA).

Fabrication of MIP/CDs-CS/GCE sensor The CDs were prepared based on our previous work [35]. The CDs were prepared via hydrothermal process as follows: 1.5 g pledget and 65 mL distilled water were transferred into autoclave. The mixture was heated at 200 °C for 12 h and then cooled to room temperature naturally. Lastly, a yellow CDs solution was obtained. Nano solid CDs were obtained by filtration, centrifugation, dialysis and freeze drying. The GCE was polished with a 1.0 μm, 0.3 μm, and 0.05 μm α-alumina slurry respectively, and then ultrasonic cleaned by distilled water, ethanol and distilled water, and dried in air before use. 1 mL of 1.0% CS solution was added to 3 mL of CDs solution with vigorous ultrasonication. Then 9.0 μL of the mixture solution was carefully dropped onto the surface of the GCE, and dried in an oven at 60 °C for 30 min. Then the CDs-CS/GCE was obtained. The MIP film was prepared on CDs-CS/GCE surface by electropolymerization in the mixed solution containing 1.29 mM glucose, 1.29 mM APBA in PBS (pH = 9.0). Electropolymerization was carried out by CV in the potential range of −0.2-0.6 V for 20 cycles with scanning rate of 50 mV s−1. After electropolymerization, the template molecules (glucose) were extracted from the MIP film by CV scanning in the range of 0–1.5 V in 0.1 M HCl for 250 cycles with a scanning rate of 50 mV s−1 until a pair of stable peaks emerged in probe solution (5 mM [Fe(CN)6]3-/4- solution containing 0.1 M KCl (pH 7.0)). Thus, MIP/CDs-CS/GCE was obtained. The schematic representation for preparing MIP/CDs-CS/GCE is illustrated in Scheme 1. In the same way, the non-imprinted polymer (NIP) modified electrode was also prepared without addition of glucose template molecules.

Experimental Electrochemical measurement Reagents The differently modified electrodes were incubated in the analyte solution at certain concentration for 10 min, and then ready for the CV, EIS and DPV analyses. After detecting the target, a washing step was taken to remove the adsorbed compounds by CV in 0.1 M HCl at 0–1.5 V until stable peaks reappeared in probe solution.

Glucose and chitosan (CS) were purchased from Jinhua Institute of science and technology (Shanghai, China), other commercially available chemicals such as hydrochloric acid, sodium hydroxide and 3aminobenzeneboronic acid (APBA) were purchased from Shanghai 43

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Scheme 1. Schematic illustration of the preparation of and MIP/CDs-CS/GCE.

gradually decline to the visible region. Several humps over the wavelength range between 250 and 700 nm are observed. The fluorescence excitation and emission wavelengths appear at 355 and 440 nm.

Detection of glucose in real samples To evaluate the practical applications of the proposed MIECS in complex matrix analysis, human blood serum obtained from the local hospital were divided into two parts for analysis. One is assayed by standard curve method. The sample without treatment was diluted with double distilled water and the calibration curves were obtained under the optimal experiment conditions. Another is investigated via standard addition method. Sample preparation was as follows:briefly, before electrochemical detection, 5.0 mL of the human blood serum was pipetted into a 10 mL centrifugal tube and centrifuged for 5 min at 8000 rpm for 3 times. The clear supernatant was diluted with PBS (pH = 7.0)and added different known concentrations of glucose. All experiments were repeated 3 times.

Characterization of different modified electrode Elemental analysis was investigated by EDX to provide the composition of MIP/CDs-CS/GCE before (A) and after extraction of glucose (B) in Fig. S3. The results display the presence of C, O, N, B can be observed with these two electrodes, while after removal of template molecule (glucose), C and O signals are decreased. It can be explained by the fact that a successful synthesis of MIECS. FT-IR spectra also provide a direct evidence for the polymer film preparation, which are illustrated in Fig. S4. Compared to the FT-IR spectrum of MIP after extraction of glucose (Fig. S4b), the disappearance of vibration at ∼1017 and ∼3300 cm−1 (with a possible component of N-H stretching) vibrations are observed in the spectrum of MIP before extraction of glucose (Fig. S4a). It is corresponding to BOH bending mode and stretching mode, respectively [40,41]. In addition, an increase in the intensity of ∼1459 cm−1 vibration attributed to asymmetric B-O stretching mode is also observed [40,41]. This is consistent with the loss of the free B-OH group which occurs with an increase in asymmetric B-O bond formation. The adsorption band at 1460–1650 cm−1 are derived from the vibration of -NH2 stretching mode, and the two sharp peaks at 2850–2910 cm−1 were attributed to probably due to the stretching vibration modes in the alkane chain of CS (Figs. S4b–c). It is predominately due to the exposure of the CDs-CS film in the MIP after extraction of glucose (Fig. S4b) and NIP (Fig. S4c). These results suggest that the MIP film has been prepared onto the modified electrode successfully. In addition, the electrochemical properties were investigated to characterize the stepwise fabrication procedure of the different modified electrode. As shown in Fig. 2A, the electrochemical responses of different modified electrodes were investigated in a 5 mM [Fe(CN)6]3-/ 4solution containing 0.1 M KCl (pH 7.0) at 50 mV s−1 with CV in the potential range of −0.2-0.6 V. The CV curve of bare GCE displays a couple of well-defined reversible redox peaks (curve a). Benefited from the affinity of positively charged surface of CS/GCE to negatively charged [Fe(CN)6]3-/4-, the diffusion of [Fe(CN)6]3-/4- to the electrode surface was encouraged. After modification on the GCE surface with CDs-CS film, the redox peak current significantly increased (curve b), which was owing to the surface structure of CDs can spur the electron transfer rate of the redox system. While both peak currents decrease remarkably after MIP modification (curve c), which can be explained by

Results and discussion Characterization of the nano CDs The surface morphologies of CDs were characterized by TEM. From Fig. 1, the image CDs showed well dispersion and spherical shape. The particle size is mainly between 2 and 5 nm (Fig. S1). Typical UV–vis absorption spectra and fluorescence spectra of the CDs are shown in Fig. S2. CDs solution has an absorption in the ultraviolet region and then

Fig. 1. TEM image of CDs. 44

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Fig. 2. CV (A) and EIS Nyquist plots (B) for bare GCE (a), CDs-CS/GCE (b), MIP/CDs-CS/GCE before (c) and after extraction of glucose (d), and MIP/CDs-CS/GCE after rebinding of glucose (e). Supporting electrolyte was 5 mM [Fe(CN)6]3-/4- solution containing 0.1 M KCl (pH 7.0). Scan rate of CV was 50 mV s−1. Frequency range of EIS was 0.1 Hz–100 KHZ.

functional monomer is relatively small, it is difficult to fully polymerize. In addition, if the amount of functional monomer excesses, template molecule may be embedded and difficult to be eluted. Moreover, the peak current and sensitivity are decreased [44]. Current peak values of CV of different ratios of APBA to glucose on different electrodes are shown in Fig. 3B (The CV curves were presented in Fig. S5). The effect of the molar ratio shown was in the range from 2:3 to 2:1. When the mole ratio of APBA to glucose was 1:1, the gap of peak current between MIP/CDs-CS/GCE and NIP/CDs-CS/GCE was the most obvious among the four proportions in the investigation. Hence, considering the imprinting effects and identification performance for glucose, the molar ratio of 1:1 was chose for the following analytical experiments.

that MIP layer was non-conductive and nearly covered the entire electrode surface. Removal of glucose creates some cavities in polymeric network so that the probe ions could pass through to the electrode surface, which is illustrate by the amplification of the peak currents (curve d). After rebinding glucose, the electrode exhibits reduced peak currents (curve e), which could be caused by the blockage of some cavities prohibiting electron transfer of [Fe(CN)6]3-/4-. The different modification processes of the sensor were also obtained by electrochemical impedance spectroscopy (EIS). EIS results were in consistence with the above results from CV. For better comparison, we also present the corresponding resistance fitted with the software named ZSimDemo software package (version 3.3). In the modified equivalent circuit (inset), Rs, Rct, C and W represent the solution resistance, charge transfer resistance, and the double-layer capacitance and the Warburg impedance, respectively. The value of the Rct depends on the dielectric and insulating features at the electrode–electrolyte interface [42,43]. As shown in Fig. 2B, there is a small semicircle domain present at CDs-CS/GCE, implying a low electron transfer resistance (259.9 Ω) to the redox probes in the electrolyte solution, which was much smaller than the Ret value of bare GCE (638 Ω). This was strongly proved that CDs-CS composite film had an excellent conductivity and could be a promising electrochemical platform for sensing. With MIP coating, the interfacial resistance increased substantially (262 KΩ), suggesting formation of hindered pathways for electron transfer. Extraction of glucose decreased resistance (103.9 KΩ) by opening the channels while rebinding of glucose increased resistance (136 KΩ) again.

The optimization of elution time for template The appropriate elution time should be selected in the preparation of MIP/CDs-CS/GCE. Current peak values of CV of different elution time for template are shown in Fig. 3C. The CV curves are presented in Fig. S6. It can be seen that the current response increased with the increase of elution time for 50–250 cycles but decreased distinctly when the elution time reached 300 cycles. This phenomenon may due to the elution equilibrium of glucose was reached for 250 cycles. On the other hand, too long elution time may also lead to the destruction of the MIP film. Therefore, the optimum elution time was 250 cycles. The optimization of adsorption time for target The adsorption time was also an important parameter for the evaluating of the efficiency of MIECS. In this work, the current response decreased rapidly between 3 and 9 min, the amount of glucose adsorbed onto the MIP sensor reached the minimum current value at 10 min, as shown in Fig. 3D. Then almost underwent not change, indicating a termination of the adsorption. As a result, the optimum adsorption time for glucose was 10 min.

Condition optimization The effect of the ratio of CDs-CS To obtain a satisfactory electrochemical sensor, it was very important to investigate the ratio of the CDs to CS on electrodes. Considering the CS has the ability of film-forming, good adhesion and other advantages, which can be used as a dispersant to form a stable film on the electrode surface, different ratio of CDs to CS (2:1, 3:1, 4:1, 5:1) were used to evaluate the most suitable ratio. As shown in Fig. 3A, the GCE modified with CDs-CS at a volume ratio of 3:1 exhibited the strongest current response. Thus, the optimal volume ratio of CDs to CS was set as 3:1.

The effect of scan rates The effect of the scan rate on the redox of glucose was also investigated. Fig. 4 reveals the CV of 20 μM glucose on MIP/GCE with different scan rates. With the increase in scan rate, the redox peak current increased gradually. As shown in Fig. 4A, the redox peak current of glucose increased linearly with the square root of the scan rate in the range of 15–300 mV s−1, which indicated that the electro-redox of glucose on the MIP/CDs-CS/GCE was a typical diffusion-controlled process [45]. In addition, with the increased scan rate, the redox peak current of glucose shifted positively. Fig. 4B displays the dependence of peak current (Ipa, Ipc, μA) on the square root of scan rate (v1/2, (v s−101/

The effect of the ratio of monomer to template In the process of preparation of molecularly imprinted sensors, it is essential to select an appropriate proportion of functional monomer to template molecule. When the amount of template molecule to 45

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Fig. 3. Effects of the volume ratio CDs to CS (A); the molar ratio functional monomers to template molecules (B), elution time (C) on the CV response at the MIP/CDsCS/GCE and NIP/CDs-CS/GCE; and incubation time on the DPV response of 20 μM glucose in 0.05 M PBS at the MIP/CDs-CS/GCE (D), other conditions were the same. 2

Electrochemical detection of glucose

).

Ipa = 4.65367 + 68.71088 v Ipc = −10.06037–72.1278 v

1/2

, R=

1/2

, R=

0.999

Under the optimal conditions, the electrochemical response of the prepared MIECS in glucose solution with different standard concentration was determined by DPV. The peak current responses of the MIP/CDs-CS/GCE to the glucose solution with different concentrations are shown in Fig. 5. With the increase of glucose concentration, the peak current decreased obviously. Besides, it was linear with the concentration of glucose in the range from 0.5 to 40 μM and from 50 to 600 μM with a low limit of detection of 0.09 μM (S/N = 3). The corresponding linear regression equations were Ip (μA) = 52.19–0.2223c

0.998

The obtained results specified that the charge-transport through the MIP/CDs-CS/GCE is a diffusion-controlled process too. This limitingdiffusion process can be due to the charge neutralization of the modified electrode surface during the redox process [46].

Fig. 4. (A) CV of 20 μM glucose on MIP/CDs-CS/GCE at different scan rates (15, 30, 50, 100, 150, 200, 250, 300 mV s−1) (from a to h). (B) The plot for the dependence of peak current on the square root of scan rate. 46

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Fig. 5. (A) DPV responses of 0.5, 5, 10, 20, 30, 40, 50, 100, 200, 400, 600 μM glucose (from a to k) on MIP/CDs-CS/GCE. (B) The relationship between the peak current and glucose concentration.

Table 1 Comparison of different electrochemical sensors based on molecularly imprinted technique for glucose. Electrode

Linear range (μM)

Detection limit (μM)

References

MAAa-based MIP pTBAb/AuNPs/ SPCEc GIPd/CuCo/SPCE MIP@Ni foam MIPMs/Gold MIP/CDs-CS/GCE

20–5.0 × 103 0.32–1.0 × 103

— 0.19

[47] [23]

1.0–2.5 × 104 1.0 × 104–5.5 × 104 2.0 × 102–8.0 × 103 0.5–40/50–6.0 × 102

0.65 — 1.0 × 104 0.09

[48] [49] [50] This work

MAAa: methacrylic acid; pTBAb: benzoic acid-functionalized poly(terthiophene); SPCEc: screen printed carbon electrode; GIPd: glucose-imprinted polymer.

(μM), R = 0.996 and Ip (μA) = 41.42–0.02366c (μM), R = 0.997, respectively. The slope and intercept standard deviation were provided in Table S1. The comparison of our method with other reported methods for the determination of glucose is listed in Table 1, which demonstrates the proposed MIECS has a lower limit for glucose analysis.

Fig. 6. Current variation of DPV responses of 20 μM glucose, AA, DA, D-fructose on MIP sensor and NIP sensor respectively.

MIP/CDs-CS/GCE in 20 μM glucose solution every 3 days. The results indicated that MIECS could maintain 87.8% of its initial response after 18 days. In addition, the relative standard deviation (RSD) was 2.75% by assaying 20 μM glucose solution with the same electrode for 11 continuous times test. These results indicate that the MIECS has a satisfactory reproducibility and stability.

Selectivity of the sensor For the sake of evaluating the selectivity of the prepared MIECS, structurally similar analog of glucose (D-fructose) and the common interfering substances in blood (ascorbic acid (AA), dopamine (DA)) were investigated. The normal physiological level of glucose in the human blood is about 30 times greater than DA and AA, 10 times greater than D-fructose [51]. Therefore, we investigated the electrochemical responses of glucose, ascorbic acid (AA), dopamine (DA) and D-fructose at the same concentration (20 μM) by DPV, respectively. As shown in Fig. 6, it demonstrated that the proposed sensor shows much higher response towards glucose than that induced by other interferes. The main reason was that when the template molecules were eluted, imprinted membrane formed with selective recognition of cavities, its size and arrangement of functional groups matched with the space structure of glucose molecules. Thus, glucose made diffusion in the imprinted membrane more easily than other chemicals. Due to the lack of imprinted cavities and interaction between functional groups, NIP sensor could only rely on non-specific adsorption, so the responses were relatively low [52]. In conclusion, the presence of AA, DA and D-fructose did not interfere in glucose determination. Hence, the proposed MIECS exhibits outstanding recognition ability for glucose detection.

Real sample analysis The reliability and accuracy of the proposed sensor were investigated in the measurement of real human blood serum samples. The standard curve of blood sample exhibited the DPV response was linear with the concentration of glucose in the range from 0.5 to 40 μM and from 50 to 661.25 μM with a low LOD of 0.11 μM (Fig. S7). The corresponding linear regression equations were Ip (μA) = 52.12–0.21609c (μM), R = 0.998 and Ip (μA) = 41.58–0.02369c (μM), R = 0.999, respectively (Table S2). Compared to the records of hospital commercial instrument, T-test shows no significant difference in the response at MIP/CDs-CS/GCE towards the glucose with confidence coefficient 95%. As shown in Table 2, the values of recovery vary from 96.9% to 100.9% with RSD varying from 3.2% to 4.3% in the spiked recovery experiment. The results suggested that the proposed MIECS exhibits a strong validity and practicability.

Reproducibility and stability of the sensor The stability of MIECS was measured by testing the response of 47

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Table 2 Determination of glucose in the human blood serum. Sample No. 1 2 3

Hospital method (μM) 3

5.29 × 10 5.29 × 103 5.29 × 103

Determined (μM)

Added (μM)

Found (μM)

RSD (%)

Recovery (%)

5.29 5.29 5.29

5 10 15

10.17 15.42 19.66

3.6 4.3 3.2

98.8 100.9 96.9

Conclusion

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In this paper, a novel sensitive and selective sensor by coupling MIT with environmental-friendly CDs for the rapid determination of glucose has been developed. The excellent sensitivity could stem from the unique properties of CDs, which improved the effective electroactive surface area and electron transportation. Combined with molecular imprinting technology, the as-prepared sensor shows outstanding selectivity. Under the optimal conditions, the obtained MIECS exhibited linear range of response from 0.5 to 40 μM and from 50 to 600 μM with a low limit of detection of 0.09 μM. In addition, the proposed MIECS also has the advantages of low cost, good reproducibility and perfect selectivity to glucose in the presence of common interferents. The successful application of the MIECS was applied to a preliminary determination of glucose in human blood serum, which may contribute to an innovation in the routine glucose monitoring in real samples. Acknowledgements This project was supported by the National Natural Science Foundation of China (31601549 and U1507115), the Natural Science of Jiangsu Province (BK20161362 and 20160500), the Natural Science of Jiangsu Education (16KJB150045) and High-Level Personnel Training Project of Jiangsu Province (BRA2016142) and the Scientific Research Foundation for Advanced Talents. This research was also supported in part by China Postdoctoral Science Foundation funded project (No. 2016M601747), the Qing Lan Project of the Higher Education Institutions of Jiangsu Province, the Senior Talent Start-up Funds of Jiangsu University of Technology and Training Program of Jiangsu Excellent Talents in Higher Vocational College (2017GRFX066). Appendix A. Supplementary data Supplementary data related to this article can be found at http://dx. doi.org/10.1016/j.ab.2018.06.004. References [1] O. Fedrigo, A.D. Pfefferle, C.C. Babbitt, R. Haygood, C.E. Wall, G.A. Wray, A potential role for glucose transporters in the evolution of human brain size, Brain Behav. Evolution 78 (2011) 315–326. [2] C.C. Wang, D.O. Laboratory, Detection of Blood Glucose, Glycosylated Hemoglobin and Insulin in Patients with Type 2 Diabetes Mellitus, (2016) Diabetes New World. [3] K. Han, J. Yao, X. Yin, M. Zhao, Q. Sun, Review on the prevalence of diabetes and risk factors and situation of disease management in floating population in China, Global Health Res. Policy 2 (2017) 33. [4] J.T. Baca, D.N. Finegold, S.A. Asher, Tear glucose analysis for the noninvasive detection and monitoring of diabetes mellitus, Ocul. Surf. 5 (2007) 280. [5] K. Yamakoshi, Y. Yamakoshi, Pulse glucometry: a new approach for noninvasive blood glucose measurement using instantaneous differential near-infrared spectrophotometry, J. Biomed. Optic. 11 (2006) 054028. [6] W.I. Kanchana, T. Sakai, N. Teshima, S. Katoh, K. Grudpan, Successive determination of urinary protein and glucose using spectrophotometric sequential injection method, Anal. Chim. Acta 604 (2007) 139–146. [7] X. Chen, J. Chen, F. Wang, X. Xiang, M. Luo, X. Ji, Z. He, Determination of glucose and uric acid with bienzyme colorimetry on microfluidic paper-based analysis devices, Biosens. Bioelectron. 35 (2012) 363–368. [8] F. Liu, J. He, M. Zeng, J. Hao, Q. Guo, Y. Song, L. Wang, Cu–hemin metal-organic frameworks with peroxidase-like activity as peroxidase mimics for colorimetric sensing of glucose, J. Nanoparticle Res. 18 (2016) 1–9. [9] Y. Xiong, Y. Zhang, P. Rong, J. Yang, W. Wang, D. Liu, A high-throughput colorimetric assay for glucose detection based on glucose oxidase-catalyzed enlargement of gold nanoparticles, Nanoscale 7 (2015) 15584–15588.

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