Interdigitated Array Microelectrode-Based Electrochemical Impedance

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Anal. Chem. 2004, 76, 1107-1113

Interdigitated Array Microelectrode-Based Electrochemical Impedance Immunosensor for Detection of Escherichia coli O157:H7 Liju Yang,† Yanbin Li,*,†,‡ and Gisela F. Erf‡

Department of Biological & Agricultural Engineering, and Center of Excellence for Poultry Science, University of Arkansas, Fayetteville, Arkansas 72701

A label-free electrochemical impedance immunosensor for rapid detection of Escherichia coli O157:H7 was developed by immobilizing anti-E. coli antibodies onto an indium-tin oxide interdigitated array (IDA) microelectrode. Based on the general electronic equivalent model of an electrochemical cell and the behavior of the IDA microelectrode, an equivalent circuit, consisting of an ohmic resistor of the electrolyte between two electrodes and a double layer capacitor, an electron-transfer resistor, and a Warburg impedance around each electrode, was introduced for interpretation of the impedance components of the IDA microelectrode system. The results showed that the immobilization of antibodies and the binding of E. coli cells to the IDA microelectrode surface increased the electron-transfer resistance, which was directly measured with electrochemical impedance spectroscopy in the presence of [Fe(CN)6]3-/4- as a redox probe. The electron-transfer resistance was correlated with the concentration of E. coli cells in a range from 4.36 × 105 to 4.36 × 108 cfu/mL with the detection limit of 106 cfu/mL. The detection of foodborne pathogenic bacteria remains a challenging and important issue for ensuring food safety and security. It is reported that foodborne diseases cause approximately 76 million illnesses, 325 000 hospitalizations, and 5000 deaths each year in the United States.1 Pathogenic bacteria cause over 90% of foodborne illness. Escherichia coli O157:H7 is one of the foodborne pathogenic bacteria that are of most concern today. The infection of E. coli O157:H7 may cause life-threatening complicationsshemolytical uremic syndrome and hemorrhagic colitis in humans.2 According to the Center for Disease Control and Prevention (CDC), an estimated 73 000 cases of infection and 61 deaths associated with E. coli O157:H7 occur in the United States each year.1 * To whom correspondence should be addressed. Phone: 479-575-2424. Fax: 479-575-7139. E-mail: [email protected]. † Department of Biological & Agricultural Engineering. ‡ Center of Excellence for Poultry Science. (1) Mead, P. S.; Slutsker, L.; Dietz, V.; McCaig, L. F.; Bresee, J. S.; Shapiro, C.; Grifrm, P.; Tauxe, R. V. Emerg. Infect. Dis. 1999, 5, 607-625. (2) Jay, J. M. Modern Food Microbiology; Aspen Publishers: Gaithersburg, MD, 2000. 10.1021/ac0352575 CCC: $27.50 Published on Web 01/15/2004

© 2004 American Chemical Society

Conventional methods for detection of pathogenic bacteria usually involve microbiological culturing and isolation of the pathogen, followed by confirmation with biochemical or serological tests. While these methods have low detection limits and can be used in complex food samples, they are typically time-consuming and labor intensive. Particularly, the required time to get confirmed results, which may take 5-7 days, makes this approach unsuitable for use in an industry-based laboratory. As a result, a number of more rapid methods have been developed for detection of pathogens, including miniaturized biochemical tests, immunological assays, and methods based on nucleic acid probes and the polymerase chain reaction. In recent years, biosensors play an increasingly important role in the detection of pathogenic organisms for environment monitoring and food safety. Most of the biosensors for detection of pathogenic bacteria are labeldependent immunosensors, in which labeled secondary antibodies are required to convert the antibody/antigen interaction into detectable optical or electrochemical signals. Label-free immunosensors, in which the immune interaction between antibody and antigen can be directly monitored when none of the reaction partners is labeled, have attractive advantages with respect to speed and simplicity of operation. Several transducing techniques, including quartz crystal microbalance (QCM)3-6 and surface plasmon resonance (SPR)7,8 have been used for developing labelfree immunosensors for detection of different bacteria. Impedance technique is an alternative for developing biosensors for detection of bacteria. Generally, there are two categories in impedance measurements: non-Faradiac and Faradiac impedance. Non-Faradaic impedance is performed in the absence of any redox probe. The growth of bacteria can cause a change in the ionic composition and hence the impedance of the medium either in the regular volume (10-15 mL)9,10 or in a microfluidic biochip (3) Park, I.; Kim, N. Biosen. Bioelectron. 1998, 13, 1091-1097. (4) Park, I.; Kim, W.; Kim, N. Biosens. Bioelectron. 2000, 15, 167-172. (5) Fung, Y. S.; Wong, Y. Y. Anal. Chem. 2001, 73, 5302-5309. (6) Ye, J.; Letcher, S. V.; Rand, A. G. J. Food Sci. 1997, 62, 1067-1086. (7) Koubova, V.; Brynda, E.; Karasova, L.; Skvor, J.; Homola, J.; Dostalek, J.; Tobiska, P.; Rosicky, J. Sens. Actuators, B 2001, 74, 100-105. (8) Fratamico, P. M.; Strobaugh, T. P.; Medina, M. B.; Gehring, A. G. Biotechnol. Tech. 1998, 12, 571-576. (9) Silley, P.; Forsythe, S. J. Appl. Bacteriol. 1996, 80, 233-243. (10) Wawerla, M.; Stolle, A.; Schalch, B.; Eisgruber, H. J. Food Prot. 1999, 62, 1488-1496.

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with small volume (nanoliter),11,12 which can be detected in the absence of a redox probe. The main application of this technique is for enumeration of bacteria. Another important application field of non-Faradiac impedance technique is for cell-based sensors. This application is based on the measurement of impedance resulting from the bacteria cells that are adherently growing on the electrode surface.13,14 The impedance signal is influenced by the changes in number, growth, and morphological behavior of adherent cells, mainly owing to the insulating effects of the cell membranes. Electrochemical impedance spectroscopy (EIS) is a Faradaic impedance technique that is performed in the presence of a redox probe. It is regarded as an effective way for sensing the formation of antigen-antibody,15 biotin-avidin complexes,16 and oligonucleotide-DNA interaction17 on electrode surfaces by probing the features of the interfacial properties (capacitance, electron-transfer resistance) of electrodes. However, EIS immunosensors for detection of pathogenic bacteria are rarely reported. In our previous study,18 electrochemical impedance spectroscopy was exploited for developing a label-dependent immunosensor for detection of E. coli O157:H7 using a planar indium-tin oxide (ITO) electrode. Alkaline phosphatase-labeled secondary antibodies were used to produce insoluble precipitation for amplifying the change in electron-transfer resistance resulting from the binding of E. coli cells to the immunosensor surface. The immunosensor achieved a detection limit of 103 colony-forming units (cfu)/mL. More attractively, integration of EIS with interdigitated array (IDA) microelectrodes could provide sensitive label-free biosensing systems for detection of small molecules, such as luteinizing hormone and DNA fragments.19,20 Other studies also showed that IDA microelectrodes had great promises in the field of label-free impedimetric immunosensing and biosensing,21-24 due to its capability of monitoring the changes of the electrical properties in the immediate neighborhood of their surface. (11) Gomez, R.; Bashir, R.; Sarikaya, A.; Ladisch, M. R.; Sturgis, J.; Robinson, J. P.; Geng, T.; Bhunia, A. K.; Apple, H. L.; Wereley, S. Biomed. Microdevices 2001, 3, 210-209. (12) Bhunia, A. K.; Jaradat, Z. W.; Naschansky, K.; Shroyer, M.; Morgan, M.; Gomez, R.; Bashir, R.; Ladisch, M. Proc, SPIE-Int. Soc. Opt. Eng. 2001, 4206, 32-39. (13) Ehret, R.; Baumann, W.; Brischwein, M.; Schwinde, A.; Wolf, B. Med. Biol. Eng. Comput. 1998, 36, 365-370. (14) Ehret, R.; Baumann, W.; Brischwein, M.; Schwinde, A.; Stegbauer, K.; Wolf, B. Biosens. Bioelectron. 1997, 12, 29-41. (15) Kharitonov, A. B.; Alfonta, L.; Katz, E., Willner, I. J. Electroanal. Chem. 2000, 487, 133-141. (16) Athey, D.; Ball, M.; McNeil, C. J.; Armstrong, R. D. Electroanalysis 1995, 7, 270-273. (17) Bardea, A.; Patolsky, F.; Dagan, A.; Willner, I. Chem. Commun. 1999, 2122. (18) Ruan, C.; Yang, L.; Li, Y. Anal. Chem. 2002, 76, 3814-4820. (19) Lillie, G.; Payne, P.; Vadgama, P. Sens. Actuators, B 2001, 78, 249-256. (20) Farace, G.; Lillie, G.; Hianik, T.; Payne, P.; Vadgama, P. Bioelectrochemistry 2002, 55, 1-3. (21) Laureyn, W.; Frederix, F.; Van Gerwen, P.; Maes, G. Transducers’99, Digest of Technical Papers, Sendai, Japan, 1999; pp 1884-1885. (22) Laureyn, W.; Nelis, D.; Van Gerwen, P.; Baert, K.; Hermans, L.; Maes, G. Eurosensors XIII, The 13th European conference on solid-state transducers, 1999, September 12-15, the Hague, The Netherland. (23) Laureyn, W.; Nelis, D.; Van Gerwen, P.; Baert, K.; Hermans, L.; Magnee, R.; Pireaux, J. J.; Maes, G. Sens. Actuators, B 2000, 68, 360-370. (24) Van Gerwen, P.; Laureyn, W.; Laureys, W.; Huyberechts, G.; Op De Beek, M.; Baert, K.; Suls, J.; Sansen, W.; Jacobs. P.; Hermans, L.; Mertens, R. Sens. Actuators, B 1998, 49, 73-80.

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Figure 1. Layout of the interdigitated array microelectrode.

In this study, we describe a label-free electrochemical impedance immunosensor using an IDA microelectrode for rapid detection of E. coli O157:H7. Anti-E. coli O157:H7 antibodies were immobilized on an ITO IDA microelectrode. The immobilization of antibodies and the formation of antibody-cell complexes on the surface of the IDA microelectrode were characterized by cyclic voltammetry and two-electrode Faradaic electrochemical impedance spectroscopy in the presence of [Fe(CN)6]3-/4- as a redox probe. An electrical equivalent circuit was proposed for understanding stepwise changes in impedance components of the immunosensor system. Electron-transfer resistance of the electrode increased upon the immobilization of antibody and the binding of E. coli cells. A correlation between the electron-transfer resistance of the electrode and the concentration of E. coli cells in the sample was found. EXPERIMENTAL SECTION Chemicals and Biochemicals. Affinity purified goat anti-E. coli O157:H7 antibody (1.0 mL, 4.6 mg/mL) was obtained from Biodesign International (Saco, ME). A 1:1 dilution was prepared with 50% glycerin solution in water before use. Phosphate-buffered saline (PBS; 0.01 M, pH 7.4) and Tris(hydroxymethyl)aminomethane-buffered solution (25 mM, pH 7.4) were purchased from Sigma-Aldrich (St. Louis, MO). Potassium ferrocyanide and potassium ferricyanide were purchased from Fisher (Fair Lawn, NJ). All solutions were prepared with deionized water from Millipore (Milli-Q, 18.2 MΩ‚cm, Bedford, MA). Bacteria and Surface Plating Method. E. coli O157:H7 (ATCC 43888) was obtained from American Type Culture Collection (Rockville, MD). The pure culture of E. coli O157:H7 was prepared in brain heart infusion broth (Remel, Lenexa, KS) at 37 °C for 20 h. The culture was serially diluted to 10-7 with physiological saline solution. The viable cell number was determined by surface plating each of 0.1-mL dilutions of 10-5, 10-6, and 10-7 on MacConkey sorbitol agar (Remel) plates. After incubation at 37 °C for 24 h, E. coli O157:H7 colonies on the plates were counted to determine the number of viable cells in cfu/mL. The culture was then heat-killed in a boiling water bath for 15 min for further use. Interdigitated Array Microelectrodes and Immobilization of Anti-E. coli O157:H7 Antibodies. ITO-coated IDA microelectrodes were obtained from ABtech Scientific, Inc. (Richmond, VA). The ITO layer was sputtered on the Schott D263 borosilicate glass substrate metallized with 100-Å Ti in monolithic configuration. The surface resistance of ITO was 10 Ω/0. Each electrode had 25 digital pairs with 15-µm digit width, 15-µm interdigit space, and a digit length of 2985 µm (Figure 1).

Immediately prior to immobilization of antibodies, IDA microeletrodes were cleaned with acetone, alcohol, 20% ethanolamine solution in water, and deionized water. They were then immersed into a solution of 1:1:5 (v/v) H2O2/NH4OH/H2O for 30 min. Finally, they were thoroughly rinsed with deionized water and dried with a stream of nitrogen. The diluted antibody solution (5 µL) was spread on the surface of the clean IDA microelectrode. The IDA microelectrode was then kept at 4 °C overnight. After this, the IDA was slowly dipped into a PBS buffer (pH 7.4) for 30 s, allowing the diffusion of unbound antibodies away from the electrode surface. This was then followed by immersion of the electrode in a fresh PBS at room temperature for 30 min. The electrode was then rinsed extensively with PBS (pH 7.4) containing 1% BSA and deionized water, and then it was dried with nitrogen. E. coli cultures with different cell numbers (20 µL) were dropped 3 times onto the surfaces of the sensors, and the sensors were incubated at 37 °C until the solution was evaporated off. To remove nonspecifically bound proteins and cells, the sensors were then rinsed thoroughly with PBS containing 0.05% Tween followed by several rinses with deionized water. Electrochemical Impedance Spectroscopy Measurements. EIS measurements were performed using an IM-6 impedance analyzer (BAS, West Lafeyette, IN) with the IM-6/THALES software. The IDA microelectrode was immersed in 4 mL of 0.01 M PBS solution, pH 7.4, containing 10 mM [Fe(CN)6]3-/4- (1:1). One of the two microband array electrodes was connected to the test and sense probes, and the other was connected to the reference and counter electrodes on the IM-6 impedance analyzer. 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. Bode (impedance and phase vs frequency) and Nyquist (imaginary impedance vs real impedance) diagrams were recorded. The difference in electron-transfer resistance before and after the cell binding was taken as the signal produced by the immune reaction between immobilized antibodies and bacterial cells. Simulation was performed using the SIM program. Fifty points of data from each measured spectrum were automatically selected by the software for input into an equivalent circuit to generate a fitting spectrum. Cyclic Voltammetry (CV) Measurements. All CV measurements were performed using the IM-6 impedance analyzer with the IM-6/THALES software under the CV program. The IDA microelectrode was immersed in 4 mL of 0.01 M PBS solution, pH 7.4, containing 10 mM [Fe(CN)6]3-/4- (1:1). Both sets of microband array electrodes were connected to both the test and sense probes as a working electrode, Ag/AgCl was used as a reference electrode and platinum as a counter electrode. Potential scanned from -0.1 to 0.6 V with a scan rate of 100 mV/s.

Figure 2. Principle of the direct impedance immunosensor contructed by interdigitated array microelectrode: (A) bare electrode; (B) with antibody immobilization; (C) with cell binding. Gray oval, E. coli O157:H7 cell; Y, anti-E. coli antibody.

RESULTS AND DISCUSSION Preparation and Transduction Principle of the Immunosensor. Among various semiconductor electrodes, ITO has been considered as a very promising material for the characterization of biological systems.25-29 ITO surfaces are stable under physio-

logical conditions because of their polarizable properties, maintaining high sensitivity without insulating oxide layers. Biomolecules, including antibodies, antigens, and enzymes can be linked to the ITO surface through the formation of M+COO- covalent linkages (where M ) indium or tin) between their free carboxyl groups and the abundant reactive hydroxyl groups (∼12-13 OH groups/nm2) on the ITO surface.25-27 In this study, anti-E. coli O157:H7 antibodies were immobilized onto the ITO surface through covalent linkages between carboxyl groups on antibodies and the abundant reactive hydroxyl groups on the ITO surface. In addition to the covalent linkage ester bonds, the OH groups on the ITO surface also help to stabilize the immobilization through noncovalent interactions such as hydrogen bonding with carboxyl or amine groups from the antibodies and van der Waals interactions. Carboxylic acid groups bind more effectively than the amino groups to the ITO surface, and the abundance of OH groups ensures thorough interactions at all possible binding sites.29 This one-step immobilization method is much more convenient and time saving than most other methods with various cross-linkers. The transduction principle of the immunosensor is shown in Figure 2. It is based on measurements of electrochemical Faradaic impedance in the presence of [Fe(CN)6]3-/4- as a redox probe. As shown in the Figure 2A, when a bare ITO IDA microelectrode is immersed into an electrolyte solution containing the redox couple, and a small-amplitude ac potential (5 mV) is applied to the electrode, the Faradaic process of oxidation and reduction of the redox couple occur, and then electrons would be transferred between the two sets of array electrodes through the redox couple. When antibodies are immobilized onto the electrode surface

(25) Fang, A.; Ng, H. T.; Li, S. F. Y. Langmuir 2001, 17, 4360-4366. (26) Fang, A.; Ng, H. T.; Su, X.; Li, S. F. Y. Langmuir 2000, 16, 5221-5226. (27) Ng, H. T.; Fang, A.; Huang, L.; Li, S. F. Y. Langmuir 2002, 18, 6324-6329.

(28) Berlin, A.; Zotti, G.; Schiavon, G.; Zecchin, S. J. Am. Chem. Soc. 1998, 120, 13453-13460. (29) Zotti, G.; Schiavon, G.; Zecchin, S. Langmuir 1998, 14, 1728-1733.

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Scheme 1. Equivalent Circuit of Electrochemical Impedance Measurement System with the Interdigitated Array Microelectrodea

Figure 3. Nyquist diagram of electrochemical impedance spectra of the bare interdigitated ITO electrode (a), after antibody immobilization (b), and after E. coli cells binding (c) in the frequency range from 1 Hz to 100 kHz. Data points from left to right correpond to decreasing frequency. Amplitude voltage, 5 mV; electrolyte, 10 mM [Fe(CN)6]3-/4(1:1) in 0.01 M PBS, pH7.4; E. coli O157:H7, 2.6 × 107cells.

(Figure 2B), they form a layer that would inhibit the electron transfer between the electrodes. An increase in the electrontransfer resistance would be expected. It is reported that the membranes of natural biological cells (thickness 5-10 nm) show a capacitance of 0.5-1.3 µF/cm2 and a resistance of 102-105 Ω‚ cm2.31 If bacterial cells attach to the electrode surface (Figure 2C), the formation of antibody-antigen complexes could create a further barrier for the electrochemical process, thereby hindering the access of the redox probe to the electrode surface, resulting in an further increase in the electron-transfer resistance. The magnitude of the increase in electron-transfer resistance is related to the number of bacterial cells captured by the immobilized antibodies. Electrochemical Impedance Spectroscopy Characterization. A Nyquist diagram of the electrochemical impedance spectrum is an effective way to measure the electron-transfer resistance. Figure 3 shows representative Nyquist diagrams of the electrochemical impedance spectra of a bare IDA microelectrode (curve a), after antibody immobilization (curve b), and after an E. coli O157:H7 (2.6 × 107 cells) cell binding (curve c) in the presence of [Fe(CN)6]3-/4- as a redox probe. As shown in Figure 3, each of the three impedance spectra includes a semicircle portion and a linear line portion, which correspond to the electrontransfer process and diffusion process, respectively. The diameter of the semicircle represents the electron-transfer resistance at the electrode surface. The electron-transfer resistances (the diameters of the semicircles) of the bare electrode, after antibody immobilization, and after cell binding were 563, 850, and 2120 Ω, respectively. This result indicated that electrochemical impedance spectroscopy was capable of monitoring the change in electrontransfer resistance resulting from the immobilization of antibodies and the binding of E. coli O157:H7 cells. The result also demonstrated that the IDA microelectrode was suitable for the detection of bacterial cells bound to its surface using electrochemical impedance spectroscopy without any amplification step. Equivalent Circuit of the Immunosensor System. The data of the electrochemical impedance spectra can be simulated with (30) Patolsky, F; Zayats, M.; Katz, E.; Willner, I. Anal. Chem. 1999, 71, 31713180. (31) Pethig, R. Dielectric and Electronic Properties of Biological Materials; J. Wiley: New York, 1979.

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a C represents double layer capacitance, R is the electrondl et transfer resistance, ZW is the Warburg impedance, Rs represents the resistance of the electrolyte solution. i, ic, and if represent total current, double layer current, and Faradaic current, respectively.

Figure 4. Bode digram of electrochemical impedance spectra of the IDA microelectrode immunosensor together with the fitting spectra. The EIS was obtained in 0.01 M PBS solution containing 10 mM [Fe(CN)6]3-/4-. Amplitude voltage, 5 mV; dc potential, equilibrium potential of [Fe(CN)6]3-/4-; frequency range, 1 Hz to 100 kHz; solid line, measurement data; dot line, fitting data.

an equivalent circuit of the system. Based on the general electronic equivalent model of an electrochemical cell30 and the behavior of the IDA microelectrode,24 an equivalent circuit, which consists of ohmic resistance (Rs) of the electrolyte between two electrodes and double layer capacitance (Cdl), electron-transfer resistance (Ret), and Warburg impedance (Zw) around each electrode, was proposed for interpretation of the impedance measurement of this IDA system (Scheme 1). Since all the current must pass through the uncompensated resistance of the electrolyte and each set of the array electrodes, Rs and the two branch circuits are connected in series. Each branch circuit represents the behavior of each set of microband array electrodes. Since the total current through the electrode surface is the sum of Faradaic current (if) and double layer current (ic), in the branch circuit, the Warburg impedance is considered to be in series with the electron-transfer resistance and both of them are connected in parallel with the double layer capacitance. The two elements of the scheme, Rs and Zw, represent the properties of the bulk solution and the diffusion of the redox probe; thus, they are not affected by the reaction occurring at the electrode surface. The other two elements, Cdl and Ret, depend on the dielectric and insulating features at the electrode/ electrolyte interface. Figure 4 (solid line) shows the Bode diagram impedance spectra of the bare interdigitated electrode measured in PBS buffer containing 10 mM [Fe(CN)6]3-/4- (1:1). This is a typical Bode

Table 1. Simulated Values of All Elements in the Equivalent Circuit for the Bare Electrode, after Antibody Immobilization, and after E. coli Cell Binding, as Well as the Percentage of Their Changesa

bare electrode antibody immobilized change antibody immobilized E. coli binding change a

Cdl (nF)

Zw (Ω/s1/2)

Ret (Ω)

Rs (Ω)

382.2 ( 15.7 401.1 ( 16.8 4.9% 401.1 ( 16.8 407.4 ( 17.4 1.6%

790.1 ( 52.9 789.5 ( 41.6 -0.0075% 789.5 ( 41.6 805.7 ( 38.7 2.1%

315.6 ( 3.8 423.8 ( 5.5 34.3% 423.8 ( 5.5 541.8 ( 6.5 27.8%

319.3 ( 18.0 320.3 ( 20.5 0.31% 320.3 ( 20.5 378.9 ( 20.4 18.3%

Number of data for simulation, 50; E. coli O157:H7, 2.6 × 106 cells.

plot for a system where polarization is due to a combination of kinetic and diffusion processes. To validate the equivalent circuit, 50 points of the measured data on the impedance spectrum were automatically selected by the software and used as input to the equivalent circuit, generating a fitting impedance spectrum (Figure 4 dash line). Using this simulation, the values of Cdl, Zw, Ret, and Rs were 382.2 nF, 790.1 Ω/s1/2, 315.6 Ω, and 319.3 Ω, respectively. The mean modulus impedance error and the phase angle error were 0.4% and 0.2°, and the maximum error of the modulus impedance and the phase angle were 4.7% and 3.1°, respectively. The agreement between the measured data and the fitting spectra indicated that this equivalent circuit provided a feasible, if not unique, model to describe the performance of the IDA electrode system. By fitting the electrochemical impedance spectra to the equivalent circuit, the value of each electrical element in the equivalent circuit was obtained (Table 1). The results showed that the immobilization of antibodies onto the IDA electrode surface changed in the double layer capacitance, the Warburg impedance, the electron-transfer resistance, and the solution resistance by 4.9, -0.0075, 34.3, and 0.31%, respectively. There was almost no change in Rs and Zw, which demonstrated that they were not affected by any modifications on the electrode surface. The most significant change was found in the electron-transfer resistance, indicating that assembly of the antibody protein layer on the electrode surface introduced an electron-transfer barrier. Similarly, the electron-transfer resistance changed from 423.8 to 541.8 Ω with an increase of 27.8% upon the binding of E. coli cells on the antibody-immobilized IDA electrode surface, which is also the most significant change among all the elements. Compared with those values obtained from Figure 3, the electron-transfer resistances of the bare electrode and the antibody-immobilized electrode are approximately half of the values obtained from Figure 3, indicating that the Nyquist diagram measured the electron-transfer resistance of the whole system, which consisted of two electron-transfer resistances of the two sets of array electrodes (two identical Ret in the equivalent circuit). This result demonstrated that bound bacterial cells on the immunosensor surface created a barrier for the electrochemical process, and the access of the redox probe was hindered to some degree, resulting in an increase in the Ret. As shown in Table 1, there was almost no change in double layer capacitance upon the bacterial cell binding. In contrast, an increase of 18% in Rs was observed upon the bacterial cell binding. When bacterial cells attached to the electrode surface, they were usually separated by a gap of 10-20 nm (up to several hundred nanometers).32 The aqueous gap

Figure 5. Cyclic voltammetry of the bare interdigitated ITO electrode (a), after antibody immobilization (b), and after E. coli cells binding (c) in the presence of 10 mM [Fe(CN)6]3-/4- (1:1) in 0.01 M PBS, pH7.4. Scan rate, 100 mV/s; Ag/AgCl as a reference electrode; platinum as a counter electrode; E. coli O157:H7, 2.6 × 107cells.

between the cell membrane and the electrode surface prevented a direct influence of the cell membrane capacitance on the impedance of the electrode. However, the cell membrane resistance of these attached bacterial cells affected the interface resistance in the IDA microelectrode system.13,14 These attached cells acted as resistors connected in series with the solution resistance in the equivalent circuit. Therefore, the Rs value (after cell binding) shown in Table 1 indeed included the original solution resistance and the additional resistance due to the attached bacterial cells. Hence, the increase in Rs was due to the bound cells’ membrane resistance. Cyclic Voltammetry Characterization. Cyclic voltammetry was used for further confirmation of the stepwise changes of the IDA microelectrode-based immunosensor. It was demonstrated by Hicks et al.33 and Horswell et al.34 that cyclic voltammetry and impedance spectroscopy are comparable for determination of electron-transfer rate on modified electrodes. Theoretically, the changes in peak current and the separation of peak potentials in voltammograms at different electrode surfaces are related to the electron-transfer rate constant and, thus, the electron-transfer resistance. Figure 5 shows the cyclic voltammograms obtained at a bare IDA (curve a), after antibody immobilization (curve b), and after the cell binding (2.6 × 107 cells) (curve c) in PBS solution (32) Gingell, D. Cell contact with solid surfaces. In Biophysics of the Cell Surface; Glaser, R., Gingell, D., Eds.; Springer, New York, 1990; pp 263-286. (33) Hicks, J. F.; Zamborini, F. P., Murray, R. W. J. Phys. Chem. B 2002, 106, 7751-7757. (34) Horswell, S. L.; O’Neil, I. A.; Schiffrin, D. J. J. Phys. Chem. B 2003, 107, 4844-4854.

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Figure 6. Impedance spectra of the interdigitated array microelectrodes based immunosensor with different numbers of E. coli O157: H7 cells on their surface: (a) antibodies and (b) 4.36 × 105, (c) 4.36 × 106, (d) 4.36 × 107, and (e) 4.36 × 108 cfu/mL. Other conditions were the same as those in Figure 3.

containing 10 mM [Fe(CN)6]3-/4- (1:1 in moles). The voltammetry shows that the voltammetric behavior of the redox probe is influenced by the electrode surface modification. When the electrode was immobilized with antibodies, a decrease in peak current (from 211.1 to 153.7 µA) and an increase in the separation of peak potentials (from 188 to 195 mV) were observed. Especially, an increase of ∼30% in the separation of peak potentials (from 195 to 251 mV) was clearly observed upon the binding of cells to the electrode surface. The cell binding also resulted in a decrease in peak current (from 153.7 to 133.4 µA). The results of this single scan rate voltammetry were consistent with the increases in the electron-transfer resistance observed in the impedance spectra. The increase in the separation of the peak potentials and the decrease in peak current implied that the immobilization of antibody and the binding of cells perturb the electron transfer right above the IDA microelectrode surface. In Figure 4, an increase of 150% in electron-transfer resistance between the antibodies immobilized electrode (850 Ω) and the cell-bound electrode (2120 Ω) was observed, whereas in Figure 5, when the same electrodes were measured by cyclic voltammetry, only a 13.2% change (from 153.7 to 133.4 µA) in the anodic current was observed. This result indicated that the impedance spectroscopic measurement of electron-transfer resistance was more sensitive than the cyclic voltammetric measurement of current for sensing the change on the electrode surface upon the binding of bacterial cells. Detection of E. coli O157:H7 Cells. Figure 6 presents one group of the impedance spectroscopic responses of an IDA microelectrode based immunosensor to different cell numbers of E. coli O157:H7. Spectra a-e were obtained using the electrodes with antibodies (a) and 4.36 × 105 (b), 4.36 × 106 (c), 4.36 × 107 (d), and 4.36 × 108 cfu/mL (e) E. coli cells on their surfaces. While there is almost no difference in the diameter between spectra a and b, a significant difference in the diameter of the semicircle can be seen between spectra a and c, indicating that the increase in electron-transfer resistance between (a) and (c) is due to the sufficient number of bacterial cells bound on the immunosensor surface. It also can be seen that there are increases in the diameters from spectrum c to spectrum d and from spectrum d to spectrum e, indicating the electron-transfer resistances in1112 Analytical Chemistry, Vol. 76, No. 4, February 15, 2004

Figure 7. Linear relationship between the logarithmic value of the concentration of E. coli O157:H7 and the electron-transfer resistance.

creased with the increasing cell number of E. coli from 4.36 × 105 to 4.36 × 108 cfu/mL on the sensor surface. This result demonstrated that the more cells bound on the immunosensor surface, the higher was the electron-transfer resistance of the immunosensors. A linear relationship between the electron-transfer resistance and logarithmic value of E. coli concentrations was found in a range of bacterial concentration from 4.36 × 105 to 4.36 × 108 cfu/mL with a correlation coefficient of 0.98 (Figure 7). This is the basis for enumeration of bacteria cell number using the immunosensor. As shown in Figure 7, when the electron-transfer resistance of the antibody-immobilized IDA microelectrode was taken as the threshold of the signal, the detection limit of this immunosensor was 106 cfu/mL, which is comparable with other label-free immunosensors for detection of pathogenic bacteria using different transducer techniques. For example, Park et al.3,4 reported two QCM immunosensors for detection of Salmonella, achieving detection limits of 3.2 × 106 and 9.9 × 105 cfu/mL; Koubova et al.7 developed SPR immonusensors for detection of Salmonella enteritidis and Listeria monocytogens, which had a detection limit of 106 cfu/mL; Fratamico et al.8 reported a SPR sensor for detection of E. coli O157:H7 with a detection limit of 107 cfu/mL. Moreover, the IDA microelectrode is one of the effective approaches for miniaturization of biosensors, enabling detector device or sample volume to be miniaturized while sustaining its sensitivity and selectivity. The IDA microelectrodebased biosensors may be integrated with microfluidics or microdetectors for biomedical and biotechnological applications. CONCLUSIONS We have demonstrated a label-free immunosensor for detection of E. coli O157:H7 cells using Faradaic impedance spectroscopy with an IDA microelectrode. A proper equivalent circuit, including ohmic resistance of the electrolyte between two electrodes and double layer capacitance, electron-transfer resistance, and Warburg impedance around each electrode, was introduced for modeling the performance of the immunosensor. Among these impedance components, the greatest change was found in electron-transfer resistance due to the binding of E. coli cells. With the IDA microelectrode, the change in the electron-transfer resistance due to the antibody immobilization and binding of E. coli cells was monitored using two-electrode electrochemical

impedance spectroscopy in the presence of [Fe(CN)6]3-/41 without any enzymatic amplification. A linear relationship between the electron-transfer resistance and logarithmic value of the cell concentration was found in a range of bacterial cell concentrations from 105 to 108 cfu/mL. The detection limit of the sensor was 106 cfu/mL, which is comparable to other label-free immunosensors for detection of pathogens using SPR12,13 or QCM8,9 techniques. This technique has potential for detection of other pathogenic bacteria by immobilizing specific antibodies onto the IDA microelectrodes.

ACKNOWLEDGMENT The authors gratefully acknowledge Dr. Ingrid Fritsch, professor in the Chemistry and Biochemistry Department at the University of Arkansas, for helpful discussion. This research was supported in part by the Food Safety Consortium and Arkansas Bioscience Institute.

Received for review December 8, 2003.

October

24,

2003.

Accepted

AC0352575

Analytical Chemistry, Vol. 76, No. 4, February 15, 2004

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