Materials Science and Engineering C 58 (2016) 953–959
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Label-free electrochemical immunosensor based on cerium oxide nanowires for Vibrio cholerae O1 detection Phuong Dinh Tam ⁎, Cao Xuan Thang Advanced Institute for Science and Technology (AIST), Hanoi University of Science and Technology, Viet Nam
a r t i c l e
i n f o
Article history: Received 5 February 2015 Received in revised form 27 July 2015 Accepted 7 September 2015 Available online 12 September 2015 Keywords: Antibody Antigen Immunosensor Nanowire CeO2
a b s t r a c t This paper developed a label-free immunosensor based on cerium oxide nanowire for Vibrio cholerae O1 detection application. The CeO2 nanowires were synthesized by hydrothermal reaction. The immobilization of Anti-V. cholerae O1 onto CeO2 nanowire-deposited sensor was performed via an amino ester, which was created by using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide, and sulfo-N-hydroxysuccinimide. The electrochemical responses of the immunosensor were studied by electrochemical impedance spectroscopy with [Fe (CN) 6] 3−/4− as redox probe. A linear response in electron transfer resistance for cell of V. cholerae O1 concentration was found in the range of 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL. The detection limit of the immunosensor was 1.0 × 102 CFU/mL. The immunosensor sensitivity was 56.82 Ω/CFU·mL−1. Furthermore, the parameters affecting immunosensor response were also investigated, as follows: pH value, immunoreaction time, incubation temperature, and anti-V. cholerae O1 concentration. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Vibrio cholerae O1, a member of the family Vibrionaceae, is a Gram-negative with about 1.4–2.6 μm long. It is well determined by biochemical tests and DNA homology investigations [1]. To date, a number of approaches for detection of Vibrio cholerae O1 have been developed such as polymerase chain reaction (PCR) [2], Immunofluorescent-Aggregation Assay [3], DNA sensor [4] and immunosensor [5]. Immunosensors are most commonly used as analytical tools for clinical diagnosis [6–8], food security [9–12], and environmental pollutants [13–15] because they are simple, highly sensitive, and easy to use. Several materials have been studied to fabricate immunosensors, such as conducting polymers [16–19], carbon nanotubes [20,21], graphene [22,23], and metal nanoparticles [21,24–26]. Recent studies show that nanostructured semiconductor metal oxides, such as zinc oxide, titanium oxide, tin oxide, tungsten oxide, and cerium oxide, have been studied for immunosensor fabrication [27–33]. Kyu et al. [27] studied a titanium dioxide nanotube array-based immunosensor. They used protein A capture to immobilize antibodies on the inner pore walls of the nanotube by electrostatic adsorption. The fundamental response of material to liquid infiltration was determined. The aqueous stabilities of porous TiO2 and SiO2 were compared in the pH range of 2 to 12. The response signals of immunosensor were observed by reflectivity spectra measurement. Ronghui Wang et al. [28] studied a TiO2 nanowire bundle-based ⁎ Corresponding author at: Advanced Institute for Science and Technology (AIST), Hanoi University of Science and Technology, No.1 Dai Co Viet, Hanoi, Viet Nam. E-mail addresses: [email protected] (P.D. Tam), [email protected] (C.X. Thang).
http://dx.doi.org/10.1016/j.msec.2015.09.027 0928-4931/© 2015 Elsevier B.V. All rights reserved.
immunosensor for rapid and sensitive detection of Listeria monocytogenes concentration. TiO2 nanowire bundle was prepared by a hydrothermal reaction of alkali with TiO2 powder. Monoclonal antibodies were immobilized on the surface of TiO2 nanowire bundle to specifically capture L. monocytogenes. The TiO2 nanowire bundle-based immunosensor could detect L. monocytogenes at a concentration as low as 4.7 × 102 CFU/mL and at a response time of 50 min. Chi-Chang Lin et al. [29] constructed an immunosensor of antibodies/conducting polymer/TiO2 nanowire composite film. TiO2 nanowires were synthesized by hydrothermal method and spin-coated on a gold microelectrode. Conducting polymer and antibodies were electrochemically polymerized on patterned nanowire. The immunosensor responses were characterized by measuring changes in current-voltage. As a result, immunosensors could detect anti-rabbit IgG within a linear range of 11.2 μg/mL to 112 μg/mL, the immunosensor sensitivity was − 0.64 A/(g/mL). Pavel and co-worker [30] reported a zinc oxide thin film transistor-based immunosensor. Primary monoclonal antibodies were attached to the ZnO channel surface. Detection of antibody and antigen interactions was performed by channel carrier modulation via pseudo double gating field effect, which was caused by the biochemical reaction. The immunosensor sensitivity was 10 fM. Anees et al. [31] developed a nanostructured nano zinc oxide film immunosensor for mycotoxin detection. The antibodies and bovine serum albumin were co-immobilized on zinc oxide film. Fourier transform infrared spectroscopy, scanning electron microscopy, and electrochemical impedance spectroscopy were used to analyze the immobilization characterizations. The immunosensor response was characterized by electrochemical method with a detection limit of 0.006 nM/dm3, response time of 25 s, and sensitivity of 189 Ω/nM. Michael et al. [32] developed an immunosensor based
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on iridium oxide thin film matrices. Antibodies were attached to iridium oxide by physical entrapment in the 3D matrix. The immunosensor displayed a linear range for IgG concentrations (10 and 200 ng/mL) and a low detection limit of 8 ng/mL. An immunosensor with cerium oxide as medium material for food-borne mycotoxin detection was studied by Ajeet et al. [33]. Here, the co-immobilization of r-IgG and bovine serum albumin (BSA) onto nano cerium oxide film was prepared. Electrochemical studies confirmed that the immunosensor exhibited a detection limit of 0.25 ng/dl and a response time of 25 s. Pratima et al. [34] developed a cerium oxide film based label – free capacitive immunosensor for detection of human chorionic gonadotropin hormone (hCG). The nano CeO2 film was fabricated onto indium tin oxide (ITO), which was used for the immobilization of anti-hCG antibody (Abs). According to Pratima group, the nano CeO2 film has a role in higher loading antibodies which led to the improvement of the immunosensing response. The sensitivity immunosensor obtained of 0.838 pF/mIU/mL in the detection range of 0–500 mIU/mL. The storage stability of immunosensor exhibits 95% response after about 5 week with relative standard deviation (RSD = 3.4%). An electrochemical immunoassay for the prostate specific antigen (PSA) using ceria mesoporous nanospheres was investigated by Juan Peng et al. [35]. A glassy carbon electrode was coated by multiwalled carbon nanotube, poly(dimethyldiallylammonium chloride), CeO2 and PSA using layer by layer method for immunosensor application. A linear relationship between the decrease in current and concentration of PSA was found in the range from 0.01 pg/mL to 1.000 pg/mL. The detection limit was 4 fg/mL. Thus, many researchers have studied semiconductor metal oxide-based immunosensors. However, published information is lacking on immunosensors that use CeO2 nanowires for label-free detection of Vibrio cholerae O1 bacteria. In this paper, we reported a CeO2 nanowire-based immunosensor for label-free detection of V. cholerae O1 bacteria. The CeO2 nanowires are fabricated by hydrothermal method using Teflon autoclave. The covalent method was performed to immobilize anti-V. cholerae O1 on CeO2 nanowire-modified sensor surface. Electrochemical impedance spectroscopy was used to detect V. cholerae O1 cell concentration with [Fe (CN) 6] 3−/4− as redox probe. Electron transfer resistance (Ret) increased linearly in the range of 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL after interaction with V. cholerae O1 cells. The CeO2 nanowire-based immunosensor exhibited low detection limit, highly sensitivity and easy to use. 2. Experimental 2.1. Chemical reagents CeO(NO3)3·6H2O, H2O2, toluene, anti-V. cholerae O1 were provide by Invitrogen Co. Phosphate-buffered saline (PBS 1×, pH 7.4), 1-ethyl-3(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC), sulfo-Nhydroxysuccinimide (NHS), bovine serum albumin BSA, H2SO4 98%, KCr2O7, , and 3-aminopropyl triethoxy-silane (APTES) were purchased from Sigma-Aldrich. Potassium ferrocyanide and potassium derricyanide were purchased from Beijing Chemical Reagent (China). All solutions were prepared with de-ionized (DI) water. 2.2. CeO2 nanowires synthesis We transferred 10 mL of 1 mol/L of CeO(NO3)3 6H2O, H2O2, and toluene into a 50 mL Teflon lined stainless steel autoclave that was placed into a furnace. The temperature was controlled to react at 160 °C for 72 h. The obtained product was directly precipitated on the silicon substrate, which was placed in the autoclave. Subsequently, the nanomaterial could be dispersed in ethanol after the silicon substrate was removed. The products were dried in an oven for 12 h at 80 °C before antibody immobilization.
2.3. Immunosensor fabrication In this work, the microelectrode was utilized as a sensor for the electrochemical impedance spectroscopy measurement. The sensor was fabricated by sputtering 10 nm Cr and 200 nm Pt on a ~ 100 nm thick silicon dioxide (SiO2) layer thermally grown on top of a silicon wafer. Then, the surface of the sensor was initially cleaned with KCr2O7 in 98% H2SO4, followed by cyclic voltammograms (swept potential from − 1 V to + 2 V; scan rate of 50 mV/s) in 0.5 M H2SO4 to activate the sensor surface. Subsequently, 10 μL of silanized-CeO2 nanowires were drop-coated on the sensor surface using APTES and were dried in a desiccator. The sensor was then immersed in the mixture of 100 mM EDC and 50 mM NHS in H2O for 60 min at room temperature to shift the terminal carboxylic group to activate NHS ester. The modified sensor was rinsed with DI water to remove EDC and NHS molecules, which were not covalently bound to the surface of the sensor and then dried in nitrogen flow. Subsequently, 2 μg/mL of anti-V. cholerae O1 was passed on the sensor surface. The NHS moiety reacted spontaneously with a primary amine group in anti-V. cholerae O1. Covalent bonding between anti-V. cholerae O1 and matrix was formed. Afterward, the immunosensor surface was rinsed with double-distilled water and dried in nitrogen flow. Finally, 1 mg/mL BSA was added to the modified immunosensor's surface to block nonspecific sites. The immunosensor was rinsed with DI water and dried in nitrogen flow. When not in use, the immunosensors were kept at 4 °C in the refrigerator. 2.4. Bacterial binding measurement IM6-impedance analyzer with IM6-THALES software was used to detect concentration of cell of V. cholerae O1. In this work, antiV. cholerae O1 modified sensor was immersed in a measuring cells and was filled with 5 mL of 1 mM PBS solution (pH 7.3) containing defined concentration of cells of V. cholerae O1 for 90 min at room temperature to form an antibody-antigen complex. The immunosensor was rinsed thrice with buffer solution to remove the non-specifically adsorbed cells. The immunosensor responses were monitored by dipping the modified sensor in 2 mL of 1 mM PBS solution containing 20 mM [Fe(CN)6]3−/4− as an indicator probe. The detected immunosensor was connected to the test and sense probe, and Pt electrode was connected to the counter electrode on the IM6 impedance analyzer. Ag/AgCl electrode was used as a reference electrode. All tests were conducted in an open circuit. The tested frequency range was from 1 Hz to 100 kHz with an amplitude of ±5 mV. The Nyquist was recorded. The differences in electron transfer resistance (Ret) were considered the signal produced by the interaction reaction between antibodies and cells. 3. Results and discussion 3.1. Electrochemical impedance spectroscopy characterizations of the immunosensor The schematic diagram of the immunosensor fabrication and cells of V. cholerae O1 binding is displayed in Fig. 1. As mentioned in [36], the impedance spectra included a semicircle portion that corresponds to the electron transfer process and a linear portion that corresponds to the diffusion process. The semicircle diameter is the electron transfer resistance. This resistance restrains the electron transfer kinetics of redoxprobe at the interface of the sensor. In this work, impedance spectra used to test interaction between antibodies and cells of V. cholerae O1. The anti-V. cholerae O1 modified sensor was dipped in 1 mM PBS solution with 20 mM [Fe (CN) 6] 3−/4− as an indicating probe. The impedance measurements were carried out at the open circuit voltage. The tested frequency range was from 1 Hz to 100 kHz with an amplitude of ±5 mV. The Nyquist frequency was recorded. The difference between the electron transfer resistance (Ret) before and after immunoreaction was considered as the signal produced. A Nyquist plots for bare sensor,
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Fig. 1. Schematic illustration of CeO2 nanowire-based immunosensor fabrication for V. cholera O1 detection.
CeO2 nanowire modified sensor, anti-V. cholerae O1/CeO2 nanowire modified sensor, control sample, and detection of V. cholerae O1 cells are presented in Fig. 2. From the Nyquist plots, we can observe that when the bare sensor was immersed in an electrolyte solution containing the redox probe, the reduction process of the redox probe occurred, and electrons were transferred between the two electrodes through the redox probe [Fe (CN) 6] 3−/4−. The electron transfer was not blocked by any monolayer on the sensor's surface. Ret was determined to be 437 Ω.
Fig. 2. Nyquist diagram for impedance measurement of sensor with the present 20 mM [Fe(CN)6]3−/4− in 1 mM PBS solution. The impedance spectra was recorded within the frequency range from 1 kHz to 100 kHz at the formal potential of the [Fe(CN)6]3−/4− redox couple. The amplitude of the alternate voltage 5 mV, anti-V. cholerae O1 concentration of 2 μg/mL, cell of V. cholerae O1 concentration of 1.0 × 102 CFU/mL, room temperature. Inset: Equivalent circuit of electrochemical impedance measurement system, Cdl represents double layer capacitance, Ret is electron transfer resistance, Zw is the Warburg impedance, and Rs represents the resistance of the electrolyte solution.
On the surface of CeO2 nanowire-modified sensor, a thin film was formed, and this film could hinder electron transfer of hexacyanoferrates into the conductive sensor surface. A small semicircle domain was formed, which corresponded to Ret = 479 Ω. On anti-V. cholerae O1/CeO2 nanowiremodified sensor surface, a thinner film was formed, and the electron transfer of hexacyanoferrates was continuously inhibited. Ret was determined to be 632 Ω. Thus, the impedance change showed that antibodies were attached to CeO2 nanowire-modified sensor surface. The antibody immobilization was continuously confirmed by interaction tests between antibodies and cells, as indicated in Fig. 2. The diameter of the semicircle increased continuously. Ret value was 764 kΩ when antibodies/cells reaction occurred because of the creation of a thick barrier layer, which blocked the access of the redox probe to the sensor surface. By contrast, no significant signal change was observed for the immunosensor exposed to the control sample. Based on these results, an equivalent circuit of the system based on the models of Randles [37] was simulated and presented in the inset of Fig. 2. This equivalent circuit includes the following: ohmic resistance of the electrolyte solution (Rs), depending on the ionic concentration, type of ions, temperature, and the geometry of the area in which current is carried; and the Warburg impedance (Zw), which is the impedance caused by the diffusion of the redox probe to the interface from the electrolyte bulk. Two elements were unaffected by the reaction on the sensor surface. The double layer capacitance (Cdl) represented the electrical double layer at the electrode/solution interface that was formed as ions from the solution attached to the electrode surface. The value of the double layer capacitance depends on many variables, including electrode potential, temperature, ionic concentrations, type of ionic, and electrode roughness. The Ret is the electron transfer resistance that shows electron transfer kinetics of the redox probe at the electrode diffusion layer. The Cdl and Ret represent the interface properties of the sensor. By fitting the electrochemical impedance spectra to the
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Table 1 Simulated value of all elements in the Randles equivalent circuit model for base sensor, CeO2 NWs modified sensor, after antibodies/NWs modified sensor and after V. cholera O1 cell binding.
Bare sensor CeO2 NWs modified sensor Antibodies/NWs modified sensor Cell of V. cholera O1 binding 1
Zw (Ω/s1/2)
Ret (Ω)
Cdl (nF)
Rs (Ω)1
512 518 611 759
470 510 630 765
430 496 478 576
500 500 500 500
Table 2 Comparison of analytical parameters of the developed sensors for the V. cholera O1 detection. Transducer
Detection limit
Linearity range
Reference
Cantilever Electrochemical SPR Electrochemical Electrochemical
103 CFU/mL 105 Cell/mL – 80 CFU/mL 102 CFU/mL
103–107 CFU/mL – 105–109 cell/mL – 102–104 CFU/mL
[38] [39] [5] [40] This work
Rs fixed 500 Ω.
equivalent circuit, the value of each electrical element in the equivalent circuit was obtained, as shown in Table 1. As mentioned above, the double layer capacitance and electron transfer resistance described the interface properties of the sensor/electrolyte and changed because of the modified sensor surface. As shown in Table 1, the change of double layer capacitance was not as high as the electron transfer resistance. The change in electron transfer resistance before and after modification of sensor surface was dramatically increased to 765 Ω. Therefore, the electron transfer resistance change was selected as a parameter to indicate the detection of V. cholerae O1 using the immunosensor based on the CeO2 nanowire. The variation in the electron transfer resistance value induced by binding with cells of V. cholerae O1 at different concentrations is shown in Fig. 3A. When the concentration of cells increased from 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL, the electron transfer resistance increased linearly, as presented in Fig. 3B. The cell concentration continuously increased, and the Ret response was saturated (1.0 × 107 CFU/mL). The relationship of Ret and cells concentration was presented as Y = 56.82X + 815.53, R = 0.97. The detection limit of the sensor was 1.0 × 102 CFU/mL, and the sensor's sensitivity was 56.82 Ω/CFUmL−1. In natural infections, the concentration of V. cholerae O1 in feces is 105 CFU/mL. Thus, CeO2 nanowire-based immunosensor could be used for V. cholerae O1 detection in real samples. The comparison of analytical parameters of the developed sensors for the V. cholera O1 detection was presented in Table 2. 3.2. Optimization of experimental conditions 3.2.1. pH value As mentioned in previously lectures [41,42], the pH value of solution is one of the parameters influencing the immunosensor response because of its effect on stability and biological activity of antibodies/ antigens. Thus, the influence of pH value has been studied in range of pH 5 to pH 8. Fig. 4a shows that the Ret value increases with increasing pH value from 5 to 7.3. The maximum Ret value was obtained at pH 7.3, which subsequently decreased with continuous pH level increase to pH 8. The reason for this phenomenon could be the denaturation of the biological activity of antibodies/antigens in acid or alkaline solutions.
Otherwise, the antigen/antibodies binding could easily be dissociated in acid or alkaline environmental. Thus, more acidic and alkaline medium is less favorable for immunosensor response. The pH value of solution strongly affected the stability of the immunosensor. Therefore, the pH 7.3 was selected as the optimal pH value in further experiments. 3.2.2. Anti-V. cholerae O1 concentration The concentration of antibodies immobilized on the sensor's surface is a very important factor in the performance of the immunosensor because of the effects on sensitivity and response time. To obtain the optimal concentration of antibodies, concentrations of anti-V. cholerae O1 varied from 0.01 μg/mL to 2 μg/mL. As shown in Fig. 4b, the changes in Ret value increased from 670 Ω to 2100 Ω with increasing anti-V. cholerae O1 concentration from 0.01 μg/mL to 2 μg/mL, respectively. The Ret response obtained saturate value when anti-V. cholerae O1 concentration increased to higher than 2 μg /mL (data not shown), thereby indicating that the anti-V. cholerae O1 amount was immobilized fully on sensor surface and they could not further on the sensor surface. Thus, the concentration of antibodies attached on sensor surface played a key role in the response of the immunosensor. In this work, 2 μg/mL of anti-V. cholerae O1 was chosen in subsequent experiments. 3.2.3. Immuno-reaction time The reaction time is another parameter that affects the performance of the immunosensor. The full reaction did not occur over a short time. A reaction that persists for a long time could cause dissociation of the antibody/cell complex. As displayed in Fig. 4c, when reaction time was less than 1 h and 30 min, the Ret value increased along with increasing reaction time. When the reaction time increased continuously, the Ret value tended to reach a saturation value because the interaction between cells and antibodies was unstable or the cells completely interacted with antibody-immobilized sensor surface because of saturation in the system. Thus, a long reaction time did not improve the response signal of the immunosensor. It was suggested that 1 h and 30 min was the optimal reaction time for the immunosensor. 3.2.4. Incubation temperature The influence of incubation temperature on the immunosensor response was also investigated in this work. As shown in Fig. 4d, the
Fig. 3. (A) Impedance spectra of the immunosensor with different cell of V. cholerae O1 concentration in the presence of [Fe(CN)6]3−/4− as redox probe, (a) 1.0 × 102, (b) 5.0 × 102, (c) 1.0 × 103, (d) 5.0 × 103, (e) 1.0 × 104, (f) 1.0 × 105, (g) 1.0 × 106, (h) 1.0 × 107 CFU/mL, (B) the relationship between cell of V.cholerae O1 concentrations and the electron transfer resistance.
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Fig. 4. Optimal of experimental parameters, influence of (a) pH value, (b) reaction time, (c) anti-V. cholerae O1 concentration, (d) incubation temperature on immunosensor response.
response signal increased with increasing temperature until it reached a maximum value at 36 °C. The response signal decreased when temperature increased continuously over 36 °C, thereby indicating that incubation temperature is very important in the activity of biomolecules. The incubation temperature for obtaining the best response signal in this work was 36 °C. However, as presented in lectures [43–46], incubation temperature was reportedly in the range of 26 °C to 37 °C. High incubation temperature could damage the activity of biomolecules (denaturization of antibodies/antigens), thereby affecting the sensitivity and lifetime of immunosensors. Thus, incubation temperature at room
temperature (about 26 °C) was recommended as the optimal temperature for this immunosensor.
3.3. Reproducibility, stability, specificity, and regeneration of immunosensor Fig. 5a presents result of reproducibility of immunosensor. To evaluate the reproducibility of immunosensor, ten immunosensors were prepared for detect 2.103 CFU/ml cells of V. cholerae O1. The experimental results showed that the relative standard deviation (RSD) of the parallel
Fig. 5. (a) The reproducibility of the immunosensor, (b) The lifetime of the immunosensor, (c) The specificity of the proposed immunosensor, (d) The regeneration of the immunosensor.
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measurements for 10 sensors was 8.4%, thereby confirming that the immunosensor has good reproducibility. Furthermore, the stability of immunosensor was also evaluated. The 10 immunosensors were stored in 0.1 M PBS (pH 7.4) at 4 °C for 100 d and were analyzed at different times (10 d/time). A repeatable signal occurred up to 40 d, after which the response signal of stored immunosensor decreased by approximately 5.4%, 6.1%, 11.6%, 15%, and 17.8% corresponding to 50, 60, 70, 80, and 90 days, respectively. The response signal was not found when the immunosensor was stored for 100 days. This indicated that biological activity of anti-V. cholerae O1 could be denatured, and no binding with cells of V. cholerae O was observed. With these results, we can conclude that the immunosensor has acceptable stability (Fig. 5b). Specificity of the immunosensor is a crucial factor in the development of microorganism detection tools. To investigate the specificity of the immunosensor, we used herpes simplex virus, Salmonella, and E. coli O157: H7 bacterium as control sample. In Fig. 5c, the response signal of the immunosensor was not detected for herpes simplex, Salmonella, and E. coli O157: H7. The shift in Ret value was ~ 1200 Ω for the detection of V. cholerae O1, thereby indicating that the specificity of immunosensor was quite good. Regeneration is a significant factor in the development of immunosensor for in-field/on-site detection. To study the regeneration of the immunosensor, anti-V. cholerae O1-immobilized sensor was immersed in a buffer solution containing V. cholerae O1 for 90 min. Washing was performed by using a PBS buffer solution and DI water, and drying was performed by using nitrogen gas. The V. cholerae O1 cell concentration was determined by the change in measurement of Ret, as presented in Fig. 5d. After detection of cells of V. cholerae O1, the immunosensor was dipped into the glycine-HCl buffer (pH 2.8) for about 10 min to remove cells. Subsequently, the sensor was washed with PBS buffer solution and DI water, and drying was conducted by using nitrogen gas. The immunosensor was again measured with cells of V. cholerae O1 under the same conditions. The obtained results in Fig. 5 d indicated that the response signal decreased by approximately 5% on average because biomolecules (antibodies) could be denatured or destroyed by using glycine-HCl buffer. 4. Conclusion In summary, a CeO2 nanowire-based electrochemical immunosensor for label-free detection of V. cholerae O1 cell was developed. Electrochemical impedance spectroscopy was used to detect V. cholerae O1 cell concentration with [Fe (CN) 6] 3−/4− as redox probe. A linear relationship between electron transfer resistance and V. cholerae O1 cell concentration was found in the range of 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL. The detection limit of immunosensor was 1.0 × 102 CFU/mL. The immunosensor sensitivity was 58.82 Ω/CFU·mL−1. This immunosensor benefits from the use of a label-free detection method, which is easy to use and simple. The immunosensor could be useful for saving time and reducing costs in clinical testing and could be applied to portable devices used for home tests, as well as for in-field/outside detection. Thus, future research would focus on the reduction of the size and improving sensitivity of the immunosensor. Acknowledgments The work was supported by the Ministry of Education and Training under research Project code: B2014.01.78 References [1] P. Baumann, A.L. Furniss, J.V. Lee, I. Genus, Vibrio Pacini 1854, in: N.R. Krieg, J.G. Holt (Eds.), Bergey's Manual of Systematic Bacteriology, 1, Williams and Wilkins, Baltimore 1984, pp. 518–538.
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
Label-free electrochemical immunosensor based on cerium oxide nanowires for Vibrio cholerae O1 detection Phuong Dinh Tam ⁎, Cao Xuan Thang Advanced Institute for Science and Technology (AIST), Hanoi University of Science and Technology, Viet Nam
a r t i c l e
i n f o
Article history: Received 5 February 2015 Received in revised form 27 July 2015 Accepted 7 September 2015 Available online 12 September 2015 Keywords: Antibody Antigen Immunosensor Nanowire CeO2
a b s t r a c t This paper developed a label-free immunosensor based on cerium oxide nanowire for Vibrio cholerae O1 detection application. The CeO2 nanowires were synthesized by hydrothermal reaction. The immobilization of Anti-V. cholerae O1 onto CeO2 nanowire-deposited sensor was performed via an amino ester, which was created by using 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide, and sulfo-N-hydroxysuccinimide. The electrochemical responses of the immunosensor were studied by electrochemical impedance spectroscopy with [Fe (CN) 6] 3−/4− as redox probe. A linear response in electron transfer resistance for cell of V. cholerae O1 concentration was found in the range of 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL. The detection limit of the immunosensor was 1.0 × 102 CFU/mL. The immunosensor sensitivity was 56.82 Ω/CFU·mL−1. Furthermore, the parameters affecting immunosensor response were also investigated, as follows: pH value, immunoreaction time, incubation temperature, and anti-V. cholerae O1 concentration. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Vibrio cholerae O1, a member of the family Vibrionaceae, is a Gram-negative with about 1.4–2.6 μm long. It is well determined by biochemical tests and DNA homology investigations [1]. To date, a number of approaches for detection of Vibrio cholerae O1 have been developed such as polymerase chain reaction (PCR) [2], Immunofluorescent-Aggregation Assay [3], DNA sensor [4] and immunosensor [5]. Immunosensors are most commonly used as analytical tools for clinical diagnosis [6–8], food security [9–12], and environmental pollutants [13–15] because they are simple, highly sensitive, and easy to use. Several materials have been studied to fabricate immunosensors, such as conducting polymers [16–19], carbon nanotubes [20,21], graphene [22,23], and metal nanoparticles [21,24–26]. Recent studies show that nanostructured semiconductor metal oxides, such as zinc oxide, titanium oxide, tin oxide, tungsten oxide, and cerium oxide, have been studied for immunosensor fabrication [27–33]. Kyu et al. [27] studied a titanium dioxide nanotube array-based immunosensor. They used protein A capture to immobilize antibodies on the inner pore walls of the nanotube by electrostatic adsorption. The fundamental response of material to liquid infiltration was determined. The aqueous stabilities of porous TiO2 and SiO2 were compared in the pH range of 2 to 12. The response signals of immunosensor were observed by reflectivity spectra measurement. Ronghui Wang et al. [28] studied a TiO2 nanowire bundle-based ⁎ Corresponding author at: Advanced Institute for Science and Technology (AIST), Hanoi University of Science and Technology, No.1 Dai Co Viet, Hanoi, Viet Nam. E-mail addresses: [email protected] (P.D. Tam), [email protected] (C.X. Thang).
http://dx.doi.org/10.1016/j.msec.2015.09.027 0928-4931/© 2015 Elsevier B.V. All rights reserved.
immunosensor for rapid and sensitive detection of Listeria monocytogenes concentration. TiO2 nanowire bundle was prepared by a hydrothermal reaction of alkali with TiO2 powder. Monoclonal antibodies were immobilized on the surface of TiO2 nanowire bundle to specifically capture L. monocytogenes. The TiO2 nanowire bundle-based immunosensor could detect L. monocytogenes at a concentration as low as 4.7 × 102 CFU/mL and at a response time of 50 min. Chi-Chang Lin et al. [29] constructed an immunosensor of antibodies/conducting polymer/TiO2 nanowire composite film. TiO2 nanowires were synthesized by hydrothermal method and spin-coated on a gold microelectrode. Conducting polymer and antibodies were electrochemically polymerized on patterned nanowire. The immunosensor responses were characterized by measuring changes in current-voltage. As a result, immunosensors could detect anti-rabbit IgG within a linear range of 11.2 μg/mL to 112 μg/mL, the immunosensor sensitivity was − 0.64 A/(g/mL). Pavel and co-worker [30] reported a zinc oxide thin film transistor-based immunosensor. Primary monoclonal antibodies were attached to the ZnO channel surface. Detection of antibody and antigen interactions was performed by channel carrier modulation via pseudo double gating field effect, which was caused by the biochemical reaction. The immunosensor sensitivity was 10 fM. Anees et al. [31] developed a nanostructured nano zinc oxide film immunosensor for mycotoxin detection. The antibodies and bovine serum albumin were co-immobilized on zinc oxide film. Fourier transform infrared spectroscopy, scanning electron microscopy, and electrochemical impedance spectroscopy were used to analyze the immobilization characterizations. The immunosensor response was characterized by electrochemical method with a detection limit of 0.006 nM/dm3, response time of 25 s, and sensitivity of 189 Ω/nM. Michael et al. [32] developed an immunosensor based
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on iridium oxide thin film matrices. Antibodies were attached to iridium oxide by physical entrapment in the 3D matrix. The immunosensor displayed a linear range for IgG concentrations (10 and 200 ng/mL) and a low detection limit of 8 ng/mL. An immunosensor with cerium oxide as medium material for food-borne mycotoxin detection was studied by Ajeet et al. [33]. Here, the co-immobilization of r-IgG and bovine serum albumin (BSA) onto nano cerium oxide film was prepared. Electrochemical studies confirmed that the immunosensor exhibited a detection limit of 0.25 ng/dl and a response time of 25 s. Pratima et al. [34] developed a cerium oxide film based label – free capacitive immunosensor for detection of human chorionic gonadotropin hormone (hCG). The nano CeO2 film was fabricated onto indium tin oxide (ITO), which was used for the immobilization of anti-hCG antibody (Abs). According to Pratima group, the nano CeO2 film has a role in higher loading antibodies which led to the improvement of the immunosensing response. The sensitivity immunosensor obtained of 0.838 pF/mIU/mL in the detection range of 0–500 mIU/mL. The storage stability of immunosensor exhibits 95% response after about 5 week with relative standard deviation (RSD = 3.4%). An electrochemical immunoassay for the prostate specific antigen (PSA) using ceria mesoporous nanospheres was investigated by Juan Peng et al. [35]. A glassy carbon electrode was coated by multiwalled carbon nanotube, poly(dimethyldiallylammonium chloride), CeO2 and PSA using layer by layer method for immunosensor application. A linear relationship between the decrease in current and concentration of PSA was found in the range from 0.01 pg/mL to 1.000 pg/mL. The detection limit was 4 fg/mL. Thus, many researchers have studied semiconductor metal oxide-based immunosensors. However, published information is lacking on immunosensors that use CeO2 nanowires for label-free detection of Vibrio cholerae O1 bacteria. In this paper, we reported a CeO2 nanowire-based immunosensor for label-free detection of V. cholerae O1 bacteria. The CeO2 nanowires are fabricated by hydrothermal method using Teflon autoclave. The covalent method was performed to immobilize anti-V. cholerae O1 on CeO2 nanowire-modified sensor surface. Electrochemical impedance spectroscopy was used to detect V. cholerae O1 cell concentration with [Fe (CN) 6] 3−/4− as redox probe. Electron transfer resistance (Ret) increased linearly in the range of 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL after interaction with V. cholerae O1 cells. The CeO2 nanowire-based immunosensor exhibited low detection limit, highly sensitivity and easy to use. 2. Experimental 2.1. Chemical reagents CeO(NO3)3·6H2O, H2O2, toluene, anti-V. cholerae O1 were provide by Invitrogen Co. Phosphate-buffered saline (PBS 1×, pH 7.4), 1-ethyl-3(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC), sulfo-Nhydroxysuccinimide (NHS), bovine serum albumin BSA, H2SO4 98%, KCr2O7, , and 3-aminopropyl triethoxy-silane (APTES) were purchased from Sigma-Aldrich. Potassium ferrocyanide and potassium derricyanide were purchased from Beijing Chemical Reagent (China). All solutions were prepared with de-ionized (DI) water. 2.2. CeO2 nanowires synthesis We transferred 10 mL of 1 mol/L of CeO(NO3)3 6H2O, H2O2, and toluene into a 50 mL Teflon lined stainless steel autoclave that was placed into a furnace. The temperature was controlled to react at 160 °C for 72 h. The obtained product was directly precipitated on the silicon substrate, which was placed in the autoclave. Subsequently, the nanomaterial could be dispersed in ethanol after the silicon substrate was removed. The products were dried in an oven for 12 h at 80 °C before antibody immobilization.
2.3. Immunosensor fabrication In this work, the microelectrode was utilized as a sensor for the electrochemical impedance spectroscopy measurement. The sensor was fabricated by sputtering 10 nm Cr and 200 nm Pt on a ~ 100 nm thick silicon dioxide (SiO2) layer thermally grown on top of a silicon wafer. Then, the surface of the sensor was initially cleaned with KCr2O7 in 98% H2SO4, followed by cyclic voltammograms (swept potential from − 1 V to + 2 V; scan rate of 50 mV/s) in 0.5 M H2SO4 to activate the sensor surface. Subsequently, 10 μL of silanized-CeO2 nanowires were drop-coated on the sensor surface using APTES and were dried in a desiccator. The sensor was then immersed in the mixture of 100 mM EDC and 50 mM NHS in H2O for 60 min at room temperature to shift the terminal carboxylic group to activate NHS ester. The modified sensor was rinsed with DI water to remove EDC and NHS molecules, which were not covalently bound to the surface of the sensor and then dried in nitrogen flow. Subsequently, 2 μg/mL of anti-V. cholerae O1 was passed on the sensor surface. The NHS moiety reacted spontaneously with a primary amine group in anti-V. cholerae O1. Covalent bonding between anti-V. cholerae O1 and matrix was formed. Afterward, the immunosensor surface was rinsed with double-distilled water and dried in nitrogen flow. Finally, 1 mg/mL BSA was added to the modified immunosensor's surface to block nonspecific sites. The immunosensor was rinsed with DI water and dried in nitrogen flow. When not in use, the immunosensors were kept at 4 °C in the refrigerator. 2.4. Bacterial binding measurement IM6-impedance analyzer with IM6-THALES software was used to detect concentration of cell of V. cholerae O1. In this work, antiV. cholerae O1 modified sensor was immersed in a measuring cells and was filled with 5 mL of 1 mM PBS solution (pH 7.3) containing defined concentration of cells of V. cholerae O1 for 90 min at room temperature to form an antibody-antigen complex. The immunosensor was rinsed thrice with buffer solution to remove the non-specifically adsorbed cells. The immunosensor responses were monitored by dipping the modified sensor in 2 mL of 1 mM PBS solution containing 20 mM [Fe(CN)6]3−/4− as an indicator probe. The detected immunosensor was connected to the test and sense probe, and Pt electrode was connected to the counter electrode on the IM6 impedance analyzer. Ag/AgCl electrode was used as a reference electrode. All tests were conducted in an open circuit. The tested frequency range was from 1 Hz to 100 kHz with an amplitude of ±5 mV. The Nyquist was recorded. The differences in electron transfer resistance (Ret) were considered the signal produced by the interaction reaction between antibodies and cells. 3. Results and discussion 3.1. Electrochemical impedance spectroscopy characterizations of the immunosensor The schematic diagram of the immunosensor fabrication and cells of V. cholerae O1 binding is displayed in Fig. 1. As mentioned in [36], the impedance spectra included a semicircle portion that corresponds to the electron transfer process and a linear portion that corresponds to the diffusion process. The semicircle diameter is the electron transfer resistance. This resistance restrains the electron transfer kinetics of redoxprobe at the interface of the sensor. In this work, impedance spectra used to test interaction between antibodies and cells of V. cholerae O1. The anti-V. cholerae O1 modified sensor was dipped in 1 mM PBS solution with 20 mM [Fe (CN) 6] 3−/4− as an indicating probe. The impedance measurements were carried out at the open circuit voltage. The tested frequency range was from 1 Hz to 100 kHz with an amplitude of ±5 mV. The Nyquist frequency was recorded. The difference between the electron transfer resistance (Ret) before and after immunoreaction was considered as the signal produced. A Nyquist plots for bare sensor,
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Fig. 1. Schematic illustration of CeO2 nanowire-based immunosensor fabrication for V. cholera O1 detection.
CeO2 nanowire modified sensor, anti-V. cholerae O1/CeO2 nanowire modified sensor, control sample, and detection of V. cholerae O1 cells are presented in Fig. 2. From the Nyquist plots, we can observe that when the bare sensor was immersed in an electrolyte solution containing the redox probe, the reduction process of the redox probe occurred, and electrons were transferred between the two electrodes through the redox probe [Fe (CN) 6] 3−/4−. The electron transfer was not blocked by any monolayer on the sensor's surface. Ret was determined to be 437 Ω.
Fig. 2. Nyquist diagram for impedance measurement of sensor with the present 20 mM [Fe(CN)6]3−/4− in 1 mM PBS solution. The impedance spectra was recorded within the frequency range from 1 kHz to 100 kHz at the formal potential of the [Fe(CN)6]3−/4− redox couple. The amplitude of the alternate voltage 5 mV, anti-V. cholerae O1 concentration of 2 μg/mL, cell of V. cholerae O1 concentration of 1.0 × 102 CFU/mL, room temperature. Inset: Equivalent circuit of electrochemical impedance measurement system, Cdl represents double layer capacitance, Ret is electron transfer resistance, Zw is the Warburg impedance, and Rs represents the resistance of the electrolyte solution.
On the surface of CeO2 nanowire-modified sensor, a thin film was formed, and this film could hinder electron transfer of hexacyanoferrates into the conductive sensor surface. A small semicircle domain was formed, which corresponded to Ret = 479 Ω. On anti-V. cholerae O1/CeO2 nanowiremodified sensor surface, a thinner film was formed, and the electron transfer of hexacyanoferrates was continuously inhibited. Ret was determined to be 632 Ω. Thus, the impedance change showed that antibodies were attached to CeO2 nanowire-modified sensor surface. The antibody immobilization was continuously confirmed by interaction tests between antibodies and cells, as indicated in Fig. 2. The diameter of the semicircle increased continuously. Ret value was 764 kΩ when antibodies/cells reaction occurred because of the creation of a thick barrier layer, which blocked the access of the redox probe to the sensor surface. By contrast, no significant signal change was observed for the immunosensor exposed to the control sample. Based on these results, an equivalent circuit of the system based on the models of Randles [37] was simulated and presented in the inset of Fig. 2. This equivalent circuit includes the following: ohmic resistance of the electrolyte solution (Rs), depending on the ionic concentration, type of ions, temperature, and the geometry of the area in which current is carried; and the Warburg impedance (Zw), which is the impedance caused by the diffusion of the redox probe to the interface from the electrolyte bulk. Two elements were unaffected by the reaction on the sensor surface. The double layer capacitance (Cdl) represented the electrical double layer at the electrode/solution interface that was formed as ions from the solution attached to the electrode surface. The value of the double layer capacitance depends on many variables, including electrode potential, temperature, ionic concentrations, type of ionic, and electrode roughness. The Ret is the electron transfer resistance that shows electron transfer kinetics of the redox probe at the electrode diffusion layer. The Cdl and Ret represent the interface properties of the sensor. By fitting the electrochemical impedance spectra to the
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Table 1 Simulated value of all elements in the Randles equivalent circuit model for base sensor, CeO2 NWs modified sensor, after antibodies/NWs modified sensor and after V. cholera O1 cell binding.
Bare sensor CeO2 NWs modified sensor Antibodies/NWs modified sensor Cell of V. cholera O1 binding 1
Zw (Ω/s1/2)
Ret (Ω)
Cdl (nF)
Rs (Ω)1
512 518 611 759
470 510 630 765
430 496 478 576
500 500 500 500
Table 2 Comparison of analytical parameters of the developed sensors for the V. cholera O1 detection. Transducer
Detection limit
Linearity range
Reference
Cantilever Electrochemical SPR Electrochemical Electrochemical
103 CFU/mL 105 Cell/mL – 80 CFU/mL 102 CFU/mL
103–107 CFU/mL – 105–109 cell/mL – 102–104 CFU/mL
[38] [39] [5] [40] This work
Rs fixed 500 Ω.
equivalent circuit, the value of each electrical element in the equivalent circuit was obtained, as shown in Table 1. As mentioned above, the double layer capacitance and electron transfer resistance described the interface properties of the sensor/electrolyte and changed because of the modified sensor surface. As shown in Table 1, the change of double layer capacitance was not as high as the electron transfer resistance. The change in electron transfer resistance before and after modification of sensor surface was dramatically increased to 765 Ω. Therefore, the electron transfer resistance change was selected as a parameter to indicate the detection of V. cholerae O1 using the immunosensor based on the CeO2 nanowire. The variation in the electron transfer resistance value induced by binding with cells of V. cholerae O1 at different concentrations is shown in Fig. 3A. When the concentration of cells increased from 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL, the electron transfer resistance increased linearly, as presented in Fig. 3B. The cell concentration continuously increased, and the Ret response was saturated (1.0 × 107 CFU/mL). The relationship of Ret and cells concentration was presented as Y = 56.82X + 815.53, R = 0.97. The detection limit of the sensor was 1.0 × 102 CFU/mL, and the sensor's sensitivity was 56.82 Ω/CFUmL−1. In natural infections, the concentration of V. cholerae O1 in feces is 105 CFU/mL. Thus, CeO2 nanowire-based immunosensor could be used for V. cholerae O1 detection in real samples. The comparison of analytical parameters of the developed sensors for the V. cholera O1 detection was presented in Table 2. 3.2. Optimization of experimental conditions 3.2.1. pH value As mentioned in previously lectures [41,42], the pH value of solution is one of the parameters influencing the immunosensor response because of its effect on stability and biological activity of antibodies/ antigens. Thus, the influence of pH value has been studied in range of pH 5 to pH 8. Fig. 4a shows that the Ret value increases with increasing pH value from 5 to 7.3. The maximum Ret value was obtained at pH 7.3, which subsequently decreased with continuous pH level increase to pH 8. The reason for this phenomenon could be the denaturation of the biological activity of antibodies/antigens in acid or alkaline solutions.
Otherwise, the antigen/antibodies binding could easily be dissociated in acid or alkaline environmental. Thus, more acidic and alkaline medium is less favorable for immunosensor response. The pH value of solution strongly affected the stability of the immunosensor. Therefore, the pH 7.3 was selected as the optimal pH value in further experiments. 3.2.2. Anti-V. cholerae O1 concentration The concentration of antibodies immobilized on the sensor's surface is a very important factor in the performance of the immunosensor because of the effects on sensitivity and response time. To obtain the optimal concentration of antibodies, concentrations of anti-V. cholerae O1 varied from 0.01 μg/mL to 2 μg/mL. As shown in Fig. 4b, the changes in Ret value increased from 670 Ω to 2100 Ω with increasing anti-V. cholerae O1 concentration from 0.01 μg/mL to 2 μg/mL, respectively. The Ret response obtained saturate value when anti-V. cholerae O1 concentration increased to higher than 2 μg /mL (data not shown), thereby indicating that the anti-V. cholerae O1 amount was immobilized fully on sensor surface and they could not further on the sensor surface. Thus, the concentration of antibodies attached on sensor surface played a key role in the response of the immunosensor. In this work, 2 μg/mL of anti-V. cholerae O1 was chosen in subsequent experiments. 3.2.3. Immuno-reaction time The reaction time is another parameter that affects the performance of the immunosensor. The full reaction did not occur over a short time. A reaction that persists for a long time could cause dissociation of the antibody/cell complex. As displayed in Fig. 4c, when reaction time was less than 1 h and 30 min, the Ret value increased along with increasing reaction time. When the reaction time increased continuously, the Ret value tended to reach a saturation value because the interaction between cells and antibodies was unstable or the cells completely interacted with antibody-immobilized sensor surface because of saturation in the system. Thus, a long reaction time did not improve the response signal of the immunosensor. It was suggested that 1 h and 30 min was the optimal reaction time for the immunosensor. 3.2.4. Incubation temperature The influence of incubation temperature on the immunosensor response was also investigated in this work. As shown in Fig. 4d, the
Fig. 3. (A) Impedance spectra of the immunosensor with different cell of V. cholerae O1 concentration in the presence of [Fe(CN)6]3−/4− as redox probe, (a) 1.0 × 102, (b) 5.0 × 102, (c) 1.0 × 103, (d) 5.0 × 103, (e) 1.0 × 104, (f) 1.0 × 105, (g) 1.0 × 106, (h) 1.0 × 107 CFU/mL, (B) the relationship between cell of V.cholerae O1 concentrations and the electron transfer resistance.
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Fig. 4. Optimal of experimental parameters, influence of (a) pH value, (b) reaction time, (c) anti-V. cholerae O1 concentration, (d) incubation temperature on immunosensor response.
response signal increased with increasing temperature until it reached a maximum value at 36 °C. The response signal decreased when temperature increased continuously over 36 °C, thereby indicating that incubation temperature is very important in the activity of biomolecules. The incubation temperature for obtaining the best response signal in this work was 36 °C. However, as presented in lectures [43–46], incubation temperature was reportedly in the range of 26 °C to 37 °C. High incubation temperature could damage the activity of biomolecules (denaturization of antibodies/antigens), thereby affecting the sensitivity and lifetime of immunosensors. Thus, incubation temperature at room
temperature (about 26 °C) was recommended as the optimal temperature for this immunosensor.
3.3. Reproducibility, stability, specificity, and regeneration of immunosensor Fig. 5a presents result of reproducibility of immunosensor. To evaluate the reproducibility of immunosensor, ten immunosensors were prepared for detect 2.103 CFU/ml cells of V. cholerae O1. The experimental results showed that the relative standard deviation (RSD) of the parallel
Fig. 5. (a) The reproducibility of the immunosensor, (b) The lifetime of the immunosensor, (c) The specificity of the proposed immunosensor, (d) The regeneration of the immunosensor.
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measurements for 10 sensors was 8.4%, thereby confirming that the immunosensor has good reproducibility. Furthermore, the stability of immunosensor was also evaluated. The 10 immunosensors were stored in 0.1 M PBS (pH 7.4) at 4 °C for 100 d and were analyzed at different times (10 d/time). A repeatable signal occurred up to 40 d, after which the response signal of stored immunosensor decreased by approximately 5.4%, 6.1%, 11.6%, 15%, and 17.8% corresponding to 50, 60, 70, 80, and 90 days, respectively. The response signal was not found when the immunosensor was stored for 100 days. This indicated that biological activity of anti-V. cholerae O1 could be denatured, and no binding with cells of V. cholerae O was observed. With these results, we can conclude that the immunosensor has acceptable stability (Fig. 5b). Specificity of the immunosensor is a crucial factor in the development of microorganism detection tools. To investigate the specificity of the immunosensor, we used herpes simplex virus, Salmonella, and E. coli O157: H7 bacterium as control sample. In Fig. 5c, the response signal of the immunosensor was not detected for herpes simplex, Salmonella, and E. coli O157: H7. The shift in Ret value was ~ 1200 Ω for the detection of V. cholerae O1, thereby indicating that the specificity of immunosensor was quite good. Regeneration is a significant factor in the development of immunosensor for in-field/on-site detection. To study the regeneration of the immunosensor, anti-V. cholerae O1-immobilized sensor was immersed in a buffer solution containing V. cholerae O1 for 90 min. Washing was performed by using a PBS buffer solution and DI water, and drying was performed by using nitrogen gas. The V. cholerae O1 cell concentration was determined by the change in measurement of Ret, as presented in Fig. 5d. After detection of cells of V. cholerae O1, the immunosensor was dipped into the glycine-HCl buffer (pH 2.8) for about 10 min to remove cells. Subsequently, the sensor was washed with PBS buffer solution and DI water, and drying was conducted by using nitrogen gas. The immunosensor was again measured with cells of V. cholerae O1 under the same conditions. The obtained results in Fig. 5 d indicated that the response signal decreased by approximately 5% on average because biomolecules (antibodies) could be denatured or destroyed by using glycine-HCl buffer. 4. Conclusion In summary, a CeO2 nanowire-based electrochemical immunosensor for label-free detection of V. cholerae O1 cell was developed. Electrochemical impedance spectroscopy was used to detect V. cholerae O1 cell concentration with [Fe (CN) 6] 3−/4− as redox probe. A linear relationship between electron transfer resistance and V. cholerae O1 cell concentration was found in the range of 1.0 × 102 CFU/mL to 1.0 × 104 CFU/mL. The detection limit of immunosensor was 1.0 × 102 CFU/mL. The immunosensor sensitivity was 58.82 Ω/CFU·mL−1. This immunosensor benefits from the use of a label-free detection method, which is easy to use and simple. The immunosensor could be useful for saving time and reducing costs in clinical testing and could be applied to portable devices used for home tests, as well as for in-field/outside detection. Thus, future research would focus on the reduction of the size and improving sensitivity of the immunosensor. Acknowledgments The work was supported by the Ministry of Education and Training under research Project code: B2014.01.78 References [1] P. Baumann, A.L. Furniss, J.V. Lee, I. Genus, Vibrio Pacini 1854, in: N.R. Krieg, J.G. Holt (Eds.), Bergey's Manual of Systematic Bacteriology, 1, Williams and Wilkins, Baltimore 1984, pp. 518–538.
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