Precipitation of an Insoluble Product on Enzyme Monolayer Electrodes

Precipitation of an insoluble, insulating product on monolayer-functionalized electrodes ..... Electrical conductivity measurements of bacterial nanow...
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Anal. Chem. 1999, 71, 3171-3180

Precipitation of an Insoluble Product on Enzyme Monolayer Electrodes for Biosensor Applications: Characterization by Faradaic Impedance Spectroscopy, Cyclic Voltammetry, and Microgravimetric Quartz Crystal Microbalance Analyses Fernando Patolsky, Maya Zayats, Eugenii Katz, and Itamar Willner*

Institute of Chemistry, The Hebrew University of Jerusalem, Jerusalem 91904, Israel

Precipitation of an insoluble, insulating product on monolayer-functionalized electrodes enables the development of new electrochemical biosensors. Faradaic impedance spectroscopy and cyclic voltammetry are used to probe the electron-transfer resistance at the conductive support upon the accumulation of the insoluble product on the electrode surface. Similarly, microgravimetric quartz crystal microbalance, QCM, analyses were used to assay the formation of the precipitate on the electrode. A horseradish peroxidase, HRP, monolayer electrode is used to analyze H2O2 via the biocatalyzed oxidation of 4-chloro1-naphthol (1) and the precipitation of the insoluble product (2). A bienzyme-layered electrode consisting of HRP and glucose oxidase, GOx, is used to sense glucose. Biocatalyzed oxidation of glucose by O2, in the presence of GOx, yields H2O2, and the generated hydrogen peroxide effects the formation of the insoluble product (2) in the presence of HRP. The insoluble product accumulated on the electrode, and the extent of the resulting electrontransfer resistance, correlated with the amounts of H2O2 or glucose, and appropriate calibration curves are extracted. Monolayer and multilayer enzyme electrodes are widely used as amperometric biosensors.1,2 One approach involves the electrochemical detection of a product generated by the biocatalyst associated with the electrode, e.g., electrical detection of hydrogen peroxide.3,4 A second approach involves the electrical contacting * To whom correspondence should be addressed. Telephone: 972-2-6585272. Fax: 972-2-6527715. E-mail: [email protected]. (1) (a) Willner, I.; Katz, E.; Willner, B. Electroanalysis 1997, 9, 965-977. (b) Willner, I.; Katz, E.; Willner, B. In Biosensors and Their Applications; Yang, V. C., Ngo, T. T., Eds.; Plenum Press: New York, in press. (c) Katz, E.; Heleg-Shabtai, V.; Willner, B.; Willner, I.; Bu ¨ ckmann, A. F. Bioelectrochem. Bioenerg. 1997, 42, 95-104. (2) (a) Heller, A. J. Phys. Chem. 1992, 96, 3579-3587. (b) Schmidt, H.-L.; Schuhmann, W. Biosens. Bioelectron. 1996, 11, 127-135. (c) Jin, W.; Bier, F.; Wollenberger, U.; Scheller, F. Biosens. Bioelectron. 1995, 10, 823-829. (d) Kuwabata, S.; Okamoto, T.; Kajiya, J.; Yoneyama, H. Anal. Chem. 1995, 67, 1684-1690. (e) Creager, S. E.; Olsen, K. G. Anal. Chim. Acta 1995, 307, 277-289. (f) Williams, R. A.; Blanch, H. W. Biosens. Bioelectron. 1994, 9, 159-167. 10.1021/ac9901541 CCC: $18.00 Published on Web 06/22/1999

© 1999 American Chemical Society

of redox enzymes and the electrode support and the direct electrical activation of the biocatalyst. Chemical functionalization of the enzyme layer with tethered redox-relay groups provides a general means for establishing electron-transfer communication between redox proteins and electrodes.5 For example, glucose oxidase layers assembled on electrodes were electrically contacted with the conducting support by the covalent linkage of ferrocene relay units to the protein.5d A different method to tailor electrically contacted enzyme electrodes includes the reconstitution of apoflavoenzymes onto relay-FAD (FAD ) flavin adenine dinucleotide) monolayer-functionalized electrodes.6 The alignment of the protein on the electrode surface and the design of a vectorial electron-transfer path in this system led to layered enzyme electrodes of unprecedented efficient electrical communication. Recently, integrated enzyme electrodes were prepared by the cross-linking of layered affinity complexes between cofactor and the enzyme7 or a synthetic electron mediator and the enzyme.8 For example, cross-linking of a layered affinity complex between a pyrroloquinoline quinone/NAD+ (NAD ) nicotinamide adenine dinucleotide) monolayer and lactate dehydrogenase yielded an electrically contacted enzyme electrode for the amperometric detection of lactate.7 Similarly, cross-linking of affinity complexes (3) (a) Bourdillon, C.; Bourgeois, J. P.; Thomas, D. J. Am. Chem. Soc. 1980, 102, 4231-4235. (b) Wieck, H. J.; Shea, C.; Yacynych, A. M. Anal. Chim. Acta 1982, 142, 277-279. (4) (a) Urban, G.; Jobst, G.; Kohl, F.; Jachimowicz, A.; Olcaytug, F.; Tilado, O.; Goiser, P.; Nauer, G.; Pittner, F.; Schalkhammer, T.; Mann-Buxbaum, E. Biosens. Bioelectron. 1991, 6, 555-562. (b) Lowry, J. P.; McAteer, K.; Atrash, S. E.; O’Neill, R. D. J. Chem. Soc., Chem. Commun. 1994, 2483-2484. (5) (a) Willner, I.; Katz, E.; Riklin, A.; Kasher, R. J. Am. Chem. Soc. 1992, 114, 10965-10966. (b) Willner, I.; Lapidot, N.; Riklin, A.; Kasher, R.; Zahavy, E.; Katz, E. J. Am. Chem. Soc. 1994, 116, 1428-1441. (c) Katz, E.; Riklin, A.; Willner, I. J. Electroanal. Chem. 1993, 354, 129-144. (d) Willner, I.; Riklin, A.; Shoham, B.; Rivenzon, D.; Katz, E. Adv. Mater. 1993, 5, 912-915. (6) (a) Willner, I.; Heleg-Shabtai, V.; Blonder, R.; Katz, E.; Tao, G.; Bu ¨ ckmann, A. F.; Heller, A. J. Am. Chem. Soc. 1996, 118, 10321-10322. (b) Katz, E.; Riklin, A.; Heleg-Shabtai, V.; Willner, I.; Bu ¨ ckmann, A. F. Anal. Chim. Acta 1999, 385, 45-58. (7) (a) Bardea, A.; Katz, E.; Bu ¨ ckmann, A. F.; Willner, I. J. Am. Chem. Soc. 1997, 119, 9114-9119. (b) Katz, E.; Heleg-Shabtai, V.; Bardea, A.; Willner, I.; Rau, H. K.; Haehnel, W. Biosens. Bioelectron. 1998, 13, 741-756. (8) (a) Patolsky, F.; Katz, E.; Heleg-Shabtai, V.; Willner, I. Chem. Eur. J. 1998, 4, 1068-1073. (b) Katz, E.; Heleg-Shabtai, V.; Willner, I.; Rau, H. K.; Haehnel, W. Angew. Chem., Int. Ed. Engl. 1998, 37, 3253-3256.

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formed between nitrate reductase and microperoxidase-11,8a or an iron(II)-protoporphyrin IX-reconstituted de novo synthetic protein8b and associated as a monolayer on electrodes, yielded electrically contacted nitrate sensing electrodes. Various enzymes stimulate the biocatalytic precipitation of products applied in histochemical analysis.9 This property was also used to develop microgravimetric quartz crystal microbalance, QCM, sensors.10 For example, an oxidase-based glucose sensor that microgravimetrically analyzes the resulting precipitate was reported.10b Also, a quartz crystal microbalance immunosensor for the detection of the African swine fever virus using oxidase-labeled antibody for the precipitation of an insoluble product was designed.10c The precipitation of products on conductive supports is anticipated to alter the interfacial electron-transfer feature at the electrode surface. Capacitance and electron-transfer resistance at the electrode interface are expected to change upon the biocatalytic precipitation of a product on the electrode surface. The time-dependent buildup of an insulating layer on the electrode alters and enhances the perturbation of the interface electrical features of the electrode, thereby providing a mechanism for amplification of sensing processes. To our knowledge, the art of assembling monolayer enzyme electrodes has not been coupled to precipitating biocatalysts that alter the interfacial electrontransfer features of conductive supports. Here we report on a novel approach for assembling enzyme electrodes on the basis of precipitating biocatalysts organized as monolayers on electrode supports. We apply Faradaic impedance spectroscopy11 and cyclic voltammetry as electrochemical sensing methods. We emphasize that this method could offer a very general approach to design enzyme-based sensors,12 immunosensors,13 and DNA sensors.14 We also apply microgravimetric QCM measurements15 to probe the precipitation of the insoluble product upon sensing of the respective analyses. Other transduction means, such as surface plasmon resonance or ellipsosmetry, could (9) (a) Nineham, A. W. Chem. Rev. 1955, 55, 355-483. (b) Tsou, K.-C.; Cheng, C.-S.; Nachlas, M. M.; Seligman, A. M. J. Am. Chem. Soc. 1956, 78, 61396144. (c) Feinstein, R. N.; Lindahl, R. Anal. Biochem. 1973, 56, 353-360. (d) Feidberg, R. S.; Datta, P. Science 1970, 170, 1414-1416. (e) Horwitz, J. P. Easwaran, C. V.; Wolf, P. L. Biochim. Biophys. Acta 1972, 276, 206214. (f) Wolf, P. L.; Horwitz, J. P.; Freisler, J. V.; Vazquez, J.; von der Muehll, E. Biochim. Biophys. Acta 1968, 159, 212-214. (g) Horwitz, J. P.; Easwaran, C. V.; Wolf, P. L.; Kowalczyk, L. S. Biochim. Biophys. Acta 1969, 185, 143152. (10) (a) Ebersole, R. C.; Ward, M. D. J. Am. Chem. Soc. 1988, 110, 8623-8628. (b) Reddy, S. M.; Jones, J. P.; Lewis, T. J.; Vadgama, P. M. Anal. Chim. Acta 1998, 363, 203-213. (c) Abad, J. M.; Pariente, F.; Ferna´ndez, L.; Lorenzo, E. Anal. Chim. Acta 1998, 368, 183-189. (11) (a) Bard, A. J.; Faulkner, L. R. Electrochemical Methods: Fundamentals and Applications: Wiley: New York, 1980. (b) Stoynov, Z. B.; Grafov, B. M.; Savova-Stoynov, B. S.; Elkin, V. V. Electrochemical Impedance; Nauka: Moscow, 1991. (12) (a) Calvo, E. J.; Etchenique, R.; Danilowicz, C.; Diaz, L. Anal. Chem. 1996, 68, 4186-4193. (b) Mirsky, V. M.; Krause, C.; Heckmann, K. D. Thin Solid Films 1996, 284-285, 930-941. (13) (a) Patolsky, F.; Filanovsky, B.; Katz, E.; Willner, I. J. Phys. Chem. B 1998, 102, 10359-10367. (b) DeSilva, M. S.; Zhang, Y.; Hesketh, P. J.; Maclay, G. J.; Gendel, S. M.; Stetter, J. R. Biosens. Bioelectron. 1995, 10, 675-682. (c) Rickert, J.; Go¨pel, W.; Beck, W. B.; Jung, G.; Heiduschka, P. Biosens. Bioelectron. 1996, 11, 757-768. (d) Maupas, H.; Soldatkin, A. P.; Martelet, C.; Jaffrezic, N.; Mandrand, B. J. Electroanal. Chem. 1997, 421, 165-171. (14) (a) Bardea, A.; Patolsky, F.; Dagan, A.; Willner, I. Chem. Commun. 1999, 21-22. (b) Patolsky, F.; Katz, E.; Bardea, A.; Willner, I. Langmuir 1999, 15, 3703-3706. (15) Buttry, D. A.; Ward, M. D. Chem. Rev. 1992, 92, 1355-1379.

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also be used to follow and characterize the insoluble precipitate on the respective transducers. EXPERIMENTAL SECTION Chemicals. Peroxidase (EC 1.11.1.7, from horseradish, HRP), glucose oxidase (EC 1.1.3.4, from Aspergilus niger, GOx), 4-chloro1-naphthol (1), and all other chemicals (Aldrich or Sigma) were used as supplied without further purification. Ultrapure water from Elgastat (UHQ) source was used throughout this work. Electrode Modification. Gold wire electrodes (0.5 mm diameter; ca. 0.2 cm2 geometrical area; roughness coefficient, ca. 1.2-1.5) were used for electrochemical measurements. Prior to the modification and measurements they were cleaned according to the published technique.16 Cleanness of the electrodes and roughness coefficients of their surface were determined by cyclic voltammetry in 0.1 M H2SO4.17 Primary modification of the Au electrode surface with active ester groups18 was performed by the electrode incubation in 10 mM dithiobis(succinimidylpropionate) in dimethyl sulfoxide (DMSO) for 30 min, followed by rinsing with pure DMSO and then water. The active ester-functionalized electrode was reacted with HRP (6 mg‚mL-1 in 0.1 M phosphate buffer, pH 7.0) for 1.5 h and then rinsed with phosphate buffer to yield the HRP-functionalized electrode. The bienzyme HRP-GOx electrode was prepared by initially treating the HRP-monolayermodified Au electrode with 15% (v/v) glutaric dialdehyde for 20 min and then rinsing with the phosphate buffer solution. The HRPmodified electrode after its activation with glutaric dialdehyde was reacted with GOx (2 mg‚mL-1 in 0.1 M phosphate buffer, pH 7.0) for 30 min and then rinsed with phosphate buffer. The bienzymemodified (HRP-GOx) Au electrodes were stored in a refrigerator at ca. 4 °C before use in the electrochemical measurements. Electrochemical Measurements. A conventional threeelectrode cell, consisting of the enzyme-modified Au electrode, a glassy carbon auxiliary electrode isolated by a glass frit, and a saturated calomel electrode (SCE) connected to the working volume with a Luggin capillary, was used for the electrochemical measurements. The cell was enclosed in a grounded Faraday cage. Impedance and cyclic voltammetry measurements were performed using an electrochemical impedance analyzer and potentiostat (EG&G, model 6310) connected to a computer (EG&G software no. 398 and no. 270/250 for impedance and cyclic voltammetry, respectively). All electrochemical measurements were performed in 0.1 M phosphate buffer, pH 7.0, as a background electrolyte. Faradaic impedance measurements were performed in the presence of a 10 mM K3[Fe(CN)6]/K4[Fe(CN)6] (1:1) mixture as a redox probe at the formal potential of the system, E° ) 170 mV, using alternating voltage, 10 mV. Impedance measurements were performed in the frequency range from 100 mHz to 50 kHz. The impedance spectra were plotted in the form of complex plane diagrams (Nyquist plots). The experimental impedance data were analyzed by the Kramers-Kronig procedure to confirm true frequency-dependence impedance. For this purpose commercial software (Equivalent Circuit, EQUIVCRT.PAS, version 4.5, formulated by B. A. Boukamp, Twente University) was employed, (16) Katz, E.; Schlereth, D. D.; Schmidt, H.-L. J. Electroanal. Chem. 1994, 367, 59-70. (17) Woods, R. In Electroanalytical Chemistry; Bard, A. J., Ed.; Dekker: New York, 1978; Vol. 9, pp 1-162. (18) Katz, E. J. Electroanal. Chem. 1990, 291, 257-260.

Scheme 1. Precipitation of the Insoluble Insulating Material (2) on the Enzyme-Modified Electrode Surface as a Result of Biochemical Reactions: (A) Au Electrode Modified with HRP in the Presence of 1 and H2O2; (B) Au Electrode Modified with HRP-GOx in the Presence of 1 and Glucose

and the calculated impedance data were found to fit the experimental results. Microgravimetric Measurements. A QCM analyzer (EG&G model QCA 917) linked to a personal computer was employed for the microgravimetric analysis. Quartz crystals (AT-cut, EG&G) sandwiched between two Au electrodes (area, 0.196 cm2; roughness factor, ca. 3.5) were used. The fundamental frequency of the crystals was ca. 9 × 106 Hz. The frequency changes of the crystals were used to evaluate the surface coverage with the primary and secondary enzymes and for the quantitative analyses of the insoluble product precipitated onto the electrodes by the respective enzymes. Analytical Procedure. 4-Chloro-1-naphthol (1) was initially dissolved in ethanol, and then the ethanolic stock solution was diluted with 0.1 M phosphate buffer, pH 7.0, to compose the analyte solution that includes 1 mM 1 and 2% (v/v) ethanol. The HRP-modified Au electrode was incubated in an analyte solution consisting of 1 mM 1 and H2O2 (variable concentration) for 10 min, unless stated otherwise. The bienzyme-modified (HRP-GOx) electrode was incubated in an analyte solution consisting of 1 mM 1 and glucose (variable concentration) for 10 min unless stated otherwise. After incubation in the respective analyte solution, the electrode was rinsed with 0.1 M phosphate buffer, pH 7.0, and applied for the respective electrochemical measurements. In control experiments, HRP- and HRP-GOx-modified electrodes were incubated in the solution of 1 mM 1 in the absence of H2O2 and glucose, respectively, for 10 min. RESULTS AND DISCUSSION Among the many enzymes that stimulate the precipitation of an insoluble substrate, HRP induces10c the H2O2 biocatalyzed oxidation of 1 to an insoluble product, benzo-4-chlorocyclohexa-

dienone (2). Thus, the immobilization of an HRP monolayer on a Au electrode could stimulate the H2O2-induced precipitation of 2 on the electrode, Scheme 1A. The amount of precipitate is controlled by the concentration of H2O2 and the time duration over which precipitation occurs. Thus, by the design of a transduction signal that monitors the extent of precipitation, a H2O2 biosensor can be tailored. Furthermore, numerous oxidases induce the O2-mediated oxidation of the respective substrates with the concomitant generation of H2O2, e.g., glucose oxidase, lactate oxidase, choline oxidase, etc. Thus, by the organization of bienzyme-layered electrodes, Scheme 1B, consisting of the oxidase as a H2O2-generating biocatalyst and HRP as a product-precipitating enzyme, many biosensor devices can be envisaged. That is, the amount of generated H2O2 and the resulting precipitate (2) are controlled by the analyte substrate oxidized by the oxidase. In fact, many biosensor electrodes based on the bioelectrocatalytic detection of a product generated by a biocatalytic transformation of the analyte substrate with a primary enzyme were reported.19 Specifically, the HRP bioelectrocatalytic detection of H2O2 generated by different oxidases is the principle of many bienzyme sensor electrodes.20 Organized multilayered enzyme electrodes have (19) Ruzgas, T.; Cso ¨regi, E.; Emne´us, J.; Gorton, L.; Marko-Varga, G. Anal. Chim. Acta 1996, 330, 123-138. (20) (a) Kulys, J. J.; Schmid, R. D. Bioelectrochem. Bioenerg. 1990, 24, 305311. (b) Wollenberger, U.; Bogdanovskaya, V.; Bobrin, S.; Scheller, F.; Tarasevich, M. Anal. Lett. 1990, 23, 1795-1808. (c) Gorton, L.; Bremle, G.; Cso ¨regi, E.; Jo ¨nsson-Pettersson, G.; Persson, B. Anal. Chim. Acta 1991, 249, 43-54. (d) Tatsuma, T.; Watanabe, T.; Watanabe, T. J. Electroanal. Chem. 1993, 356, 245-253. (e) Buttler, T. A.; Johansson, K.; Gorton, L.; Marko-Varga, G. A. Anal. Chem. 1993, 65, 2628-2636. (f) Kacaniklic, V.; Johansson, K.; Marko-Varga, G.; Gorton, L.; Jo¨nsson-Pettersson, G.; Cso¨regi, E. Electroanalysis 1994, 6, 381-390.

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Scheme 2. General Equivalent Circuit for the Faradaic Impedance Spectroscopy Measurements Performed at the Au Electrode Modified with Enzymes and upon the Precipitation of the Insoluble Layer of 2 in the Presence of [Fe(CN)6]3-/4- as a Redox Probe

originating from an electrode surface modifier, Cmod. The doublelayer capacitance is expressed by eq 1. Any electrode modifier of

1 1 1 ) + Cdl CAu Cmod

(1)

insulating features decreases the double-layer capacitance as compared to the pure metal electrode. Ret controls the electrontransfer kinetics of the redox probe at the electrode interface. Thus, the insulating modifier on the electrode is expected to retard the interfacial electron-transfer kinetics and to increase the electron-transfer resistance. The electron-transfer resistance at the electrode will be given by eq 2, where RAu and Rmod are the

Ret ) RAu + Rmod

already been reported;21 e.g., a layered electrode consisting of choline oxidase and choline esterase was reported22 as an electrochemical biosensor for acetylcholine. Thus, the combination of two enzymes (or more) in an organized, layered array on an electrode could lead to the biocatalyzed precipitation of an insoluble product and thereby establish a new principle for the development of biosensor systems. Impedance spectroscopy is an effective method for probing the features of surface-modified electrodes.23 The complex impedance can be presented as the sum of the real, Zre(ω), and imaginary, Zim(ω) components that originate mainly from the resistance and capacitance of the cell, respectively. The general electronic equivalent scheme (Randles and Ershler model11), Scheme 2, includes the ohmic resistance of the electrolyte solution, Rs, the Warburg impedance, ZW, resulting from the diffusion of ions from the bulk electrolyte to the electrode interface, the double layer capacitance, Cdl, and electron-transfer resistance, Ret, that exists if a redox probe is present in the electrolyte solution. The two components of the electronic scheme, Rs and ZW, represent bulk properties of the electrolyte solution and diffusion features of the redox probe in solution. Therefore, these parameters are not affected by chemical transformations occurring at the electrode surface. The other two components in the scheme, Cdl and Ret, depend on the dielectric and insulating features at the electrode/electrolyte interface. The double-layer capacitance consists of the constant capacitance of an unmodified electrode (e.g. for an Au electrode, CAu ≈ 40-60 µF‚cm-2, depending on the applied potential24) and a variable capacitance (21) (a) Shoham, B.; Migron, Y.; Riklin, A.; Willner, I.; Tartakovsky, B. Biosens. Bioelectron. 1995, 10, 341-352. (b) Jin, W.; Bier, F.; Wollenberger, U.; Scheller, F. Biosens. Bioelectron. 1995, 10, 823-829. (22) Riklin, A.; Willner, I. Anal. Chem. 1995, 67, 4118-4126. (23) (a) Ren, X.; Pickup, P. G. J. Electroanal. Chem. 1997, 420, 251-257. (b) Hall, E. A. H.; Skinner, N. G.; Jung, C.; Szunerits, S. Electroanalysis 1995, 7, 830-836. (c) Brillas, E.; Cabot, P. L.; Garrido, A.; Montilla, M.; Rodriguez, R. M.; Carrasco, J. J. Electroanal. Chem. 1997, 430, 133-140. (d) Ehret, R.; Baumann, W.; Brischwein, M.; Schwinde, A.; Stegbauer, K.; Wolf, B. Biosens. Bioelectron. 1997, 12, 29-41.

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(2)

electron-transfer resistance of the nonmodified electrode and the variable electron-transfer resistance introduced by the modifier, in the presence of the solubilized redox probe, respectively. A typical shape of a Faradaic impedance spectrum (presented in the form of a Nyquist plot) includes a semicircle region lying on the Zre-axis followed by a straight line. The semicircle portion, observed at higher frequencies, corresponds to the electrontransfer-limited process, whereas the linear part is characteristic of the lower frequencies range and represents the diffusionallimited electron-transfer process. In the case of very fast electrontransfer processes, the impedance spectrum could include only the linear part, whereas a very slow electron-transfer step results in a big semicircle region that is not accompanied by a straight line. The electron-transfer kinetics and diffusional characteristics can be extracted from the spectra. The semicircle diameter equals Ret. The intercept of the semicircle with the Zre-axis at high frequencies (ω f ∞) is equal to the Rs. Extrapolation of the circle to lower frequencies yields an intercept corresponding to Rs + Ret. The characteristic frequency, ωo, given by eq 3 has the meaning of the reciprocal of the time constant of the equivalent circuit. The maximum value of the imaginary impedance in the semicircle part corresponds to Zim ) Ret/2 and is achieved at the characteristic frequency, ωo.

ωo ) (CdlRet)-1

(3)

The electron-transfer resistance can be translated into the exchange current under equilibrium, Io, eq 4, and then the heterogeneous electron-transfer rate constant, ket, can be evaluated, eq 5, where R is the gas constant, T is the temperature (K), A is the electrode area (cm2), [S] corresponds to the bulk concentration of the redox probe (mol‚cm-3), and n is the number of transferred electrons per molecule of the redox probe.

Ret ) RT(nFIo)-1

(4)

Io ) nFAket[S]

(5)

Previous studies have employed Faradaic impedance measurements to characterize the formation of modified thin films on the (24) Champagne, G. Y.; Belanger, D.; Fortier, G. Bioelectrochem. Bioenerg. 1989, 22, 159-166.

Scheme 3. Assembly of the Bienzyme HRP-GOx Layer on the Surface of the Au Electrode

electrode surface. For example, the self-assembly aromatic disulfides on gold electrodes,25 or the electrodeposition of polymer films on electrodes,26 were characterized by Faradaic impedance spectroscopy. Recently, the affinity interactions of the dinitrophenyl antibody, DNP-Ab, with an antigen-monolayer electrode, and specifically the photoswitchable interactions of DNP-Ab with a photoisomerizable antigen monolayer associated with a Au electrode, were reported.13a Scheme 3 outlines the stepwise assembly of the layered enzyme electrodes consisting of the HRP-monolayer electrode and the HRP-GOx-layered electrode. Microgravimetric quartz crystal microbalance experiments revealed that the surface coverage of the Au surfaces with the HRP and the GOx enzymes corresponds to 2.6 × 10-12 and 1.5 × 10-12 mol‚cm-2, respectively. The surface coverage that we found for GOx is ca. 60% of the theoretical limit for the densely packed GOx monolayer.6a Figure 1 shows the Faradaic impedance spectra, presented as Nyquist plots (Zim vs Zre), of the bare Au electrode (A), the electrode after modification with the HRP monolayer (B, curve a), the HRP-monolayermodified Au electrode after its treatment with 1, in the absence (25) (a) Bandyopadhyay, K.; Vijayamohanan, K.; Shekhawat, G. S.; Gupta, R. P. J. Electroanal. Chem. 1998, 447, 11-16. (b) Bandyopadhyay, K.; Patil, V.; Sastry, M.; Vijayamohanan, K. Langmuir 1998, 14, 3808-3814. (26) (a) Rossberg, K.; Paasch, G.; Dunsch, L.; Ludwig, S. J. Electroanal. Chem. 1998, 443, 49-62. (b) Komaba, S.; Osaka, T. J. Electroanal. Chem. 1998, 453, 19-23.

of H2O2 (B, curve b), and the HRP-monolayer-modified Au electrode after its treatment with an analyte solution consisting of 1 and H2O2 (B, curve c). (Note different scales in the two parts of Figure 1.) It can be seen from Figure 1A that the bare Au electrode exhibits an almost straight line that is characteristic of a diffusional limiting step of the electrochemical process. Assembly of HRP on the electrode surface generates a hydrophobic insulating layer on the electrode that introduces a barrier to the interfacial electron transfer. This is reflected by the appearance of the semicircle part of the spectrum (Figure 1B, curve a) corresponding to the interfacial electron-transfer resistance, Ret ) 4.2 kΩ. The treatment of the HRP-monolayer-modified electrode with 1 in the absence of H2O2 results in a very small increase of the semicircle diameter, reflecting insignificant adsorption of 1 on the electrode surface (Figure 1B, curve b). On the other hand, incubation of the HRP-monolayer-modified Au electrode in the analyte solution composed of 1 and H2O2 results in a very significant change in the Faradaic impedance spectrum (Figure 1B, curve c). The semicircle region reveals a very large diameter, implying high electron-transfer resistance, and the Ret value corresponds to 117.0 kΩ. This is due to the precipitation of the insoluble ketone (2) onto the electrode, resulting in an insulating layer on the conducting support, Scheme 1A. The insulating layer introduces a barrier for electron transfer between the redox probe in the electrolyte solution and the electrode. Note that upon the Analytical Chemistry, Vol. 71, No. 15, August 1, 1999

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Figure 1. Nyquist diagram (Zim vs Zre) for the Faradaic impedance measurements in the presence of 10 mM [Fe(CN)6]3-/4- at (A) a bare Au electrode; (B) a HRP-monolayer-modified Au electrode before any further treatment (a), after its incubation in 1 mM 1 solution for 10 min (b), and after its incubation in 1 mM 1 solution containing 5 × 10-3 M H2O2 for 10 min (c). Inset: Enlarged impedance spectra corresponding to curves a and b.

precipitation of 2 the interfacial electron-transfer resistance increases from 4.2 kΩ to 117.0 kΩ. The precipitate (2) accumulates on the electrode support by the nonelectrochemical biocatalyzed process. Therefore, the insulating layer is anticipated to become thicker as time proceeds. Curves b-e of Figure 2A show the Faradaic impedance spectra of the HRP-monolayer electrode in the presence of H2O2, 5 × 10-3 M, and 1 at different time intervals. The semicircle diameters of the impedance plots increase as the time intervals for precipitation of 2 are longer, Figure 2B. Table 1 summarizes the electrontransfer resistances, the corresponding electron-transfer rate constants, and double-layer capacitance values at the electrode surface at different time intervals of precipitation of 2. It is evident that upon the accumulation of the insoluble product on the electrode surface, the interfacial electron transfer is retarded and the capacitance is decreased. Cyclic voltammetry experiments further confirm that accumulation of the insulating film on the electrode perturbs the electrical communication between the redox probe in the electrolyte solution and the electrode. Figure 3 shows the cyclic voltammograms of the Fe(CN)63-/Fe(CN)64- couple in the presence of the bare Au electrode, curve a, upon the assembly of the HRP monolayer, curve b, and after exposure of the HRPmonolayer electrode to H2O2 and 1, curve c. While the cyclic voltammogram of the redox probe reveals quasi-reversible features at the bare electrode, the assembly of the HRP monolayer results 3176 Analytical Chemistry, Vol. 71, No. 15, August 1, 1999

Figure 2. (A) Nyquist diagram (Zim vs Zre) for the Faradaic impedance measurements in the presence of 10 mM [Fe(CN)6]3-/4at the HRP-modified Au electrode after its incubation in the solution composed of 1 mM 1 and 5 × 10-3 M H2O2 for (a) 0, (b) 2, (c) 4, (d) 7, and (e) 15 min. Inset shows curve a with a higher resolution scale. (B) Electron-transfer resistance, Ret, as a function of the HRP-modified Au electrode incubation time in 1 mM 1 and 5 × 10-3 M H2O2. Table 1. Electron-Transfer Resistances, Interfacial Electron-Transfer Rate Constants, and Double-Layer Capacitance Values at the HRP-Modified Au Electrode at Different Time Intervals of Precipitation of 2a time (min)

Retb (kΩ)

ketc (cm‚s-1)

Cdld (µF‚cm-2)

2 4 7 15

19.1 52.6 72.9 148.0

6.8 × 10-6 2.5 × 10-6 1.8 × 10-6 0.9 × 10-6

1.5 1.3 0.94 0.83

a Precipitation induced in the presence of 1 × 10-3 M 1 and 5 × 10-3 M H2O2. b Extracted from the diameter of the respective impedance spectra. c Calculated from the respective Ret values using eq 4 and eq 5. d Calculated using eq 3 and taking into account the electrode area.

in larger peak-to-peak separation, Figure 3. This originates from the assembly of a hydrophobic protein layer on the electrode that retards the electron transfer between the redox probe and the electrode.27 Similarly, blocking of the interfacial electron transfer between a redox probe and the electrode upon the binding of a protein to the electrode surface was previously reported,28 e.g., (27) Willner, I.; Rubin, S.; Cohen, Y. J. Am. Chem. Soc. 1993, 115, 4937-4938. (28) (a) Willner, I.; Blonder, R.; Dagan, A. J. Am. Chem. Soc. 1994, 116, 93659366. (b) Blonder, R.; Levi, S.; Tao, G.; Ben-Dov, I.; Willner, I. J. Am. Chem. Soc. 1997, 119, 10467-10478.

Figure 3. Cyclic voltammograms of a Au electrode measured in the presence of 10 mM [Fe(CN)6]3-/4- after different steps of modification: (a) bare Au electrode; (b) HRP-modified Au electrode; (c) HRP-modified Au electrode after incubation in a solution consisting of 1 mM 1 and 5 × 10-3 M H2O2 for 10 min. All data were recorded under argon in 0.1 M phosphate buffer solution, pH 7.0; potential scan rate, 100 mV‚s-1.

upon the association of an antibody to an antigen monolayer. The precipitation of the insoluble product (2) on the electrode blocks entirely the electrical contact between the redox probe and the conducting surface, and the electrical response of Fe(CN)63- is completely depleted, Figure 3. The time-dependent blocking of the electrical communication between the redox label and the electrode, upon precipitation of 2, can be followed by cyclic voltammetry. Exposure of the HRP-monolayer electrode to H2O2 and 1 for longer time intervals results in an increase in the peakto-peak separation of the redox probe in solution that ultimately leads to the complete insulation of the electrode upon the accumulation of the insoluble product on the electrode surface. Faradaic impedance spectroscopy, and, specifically, the characterization of the electron-transfer resistance at the electrode, is a sensitive method for probing the extent of deposition of the insoluble product on the electrode surface. The extent of insulation of the electrode by the insoluble precipitate is controlled by two parameters: (i) the concentration of H2O2 that induces the biocatalyzed oxidation of 1 and the formation of the insoluble product (2); (ii) the time interval over which the monolayerenzyme electrode is subjected to the precipitation process. In fact, it should be noted that the time-dependent formation of the insoluble product provides a route for amplifying the analysis of the substrate, e.g., H2O2. That is, even low concentrations of the analyte can be sensed by allowing the system to generate and accumulate the insoluble product on the electrode. Thus, to use the HRP-monolayer electrode to sense H2O2, the Faradaic impedance spectra should be recorded at different concentrations of the analyte and upon the exposure of the electrode to the precipitation process for a fixed time. The extent of insulation of the electrode surface, and the resulting electron-transfer resistance at the electrode, will be controlled by the concentration of H2O2. Figure 4 shows the derived calibration curve for the HRPmonolayer electrode, corresponding to the electron-transfer resistance at different concentrations of H2O2. This calibration curve was extracted when the insoluble product was allowed to accumulate on the electrode surface for a time interval of 10 min. By permitting the formation of the precipitate for longer time intervals, very low concentrations of H2O2 could be determined.

Figure 4. Electron-transfer resistance, Ret, extracted from the Faradaic impedance spectra registered with the HRP-modified Au electrode after its incubation for 10 min in the analyte solution composed of 1 × 10-3 M 1 and variable concentrations of H2O2.

For example, by allowing the accumulation of 2 for 2 h, H2O2 was determined at a sensitivity level corresponding to 1 × 10-8 M. To our knowledge, this is the highest sensitivity level achieved to date for the detection of H2O2 by an HRP-based biosensor electrode.19 The ultrasensitive detection of H2O2 is very interesting for clinical diagnosis. Reactive oxygen intermediates may be generated in the lungs during various pathological processes. These intermediates, as well as hydrogen peroxide, are toxic to cells through their oxidizing effects on proteins, membranes, and DNA. Detection of submicromolar concentrations of H2O2 is vital because these peroxide levels can damage mammalian cells.29 Detection of very low concentrations of hydrogen peroxide is also important for environmental protection; e.g., detection of 0.0318 µM H2O2 is necessary when analyzing photochemical smog.19 Microgravimetric, QCM experiments provide a further means for following the precipitation of 2 on the enzyme-functionalized surface. The accumulation of the precipitate on the piezoelectric crystal is anticipated to cause a mass change, ∆m, that results in a frequency change, ∆f, of the crystal, given by the Sauerbrey equation, eq 6, where ∆f is the measured frequency shift, fo is

∆m ∆f ) -2fo2 A(µqFq)1/2

(6)

the frequency of the quartz crystal prior to a mass change, ∆m is the mass change, A is the piezoelectrically active area, Fq is the density of quartz (2.648 g‚cm-3), and µq is the shear modulus (2.947 × 1011 dyn‚cm-2 for AT-cut quartz).15 Accordingly, the HRP monolayer was assembled onto a Au-quartz crystal (9 MHz), as outlined in Scheme 3. Figure 5 shows the time-dependent frequency change of the crystal modified with the HRP monolayer in the presence of 1 and different concentrations of H2O2. In the absence of H2O2 no significant frequency changes are observed, Figure 5, curve a, indicating that 1 by itself does not bind or precipitate onto the crystal. In the presence of H2O2 the crystal frequency decreases, implying that a mass increase on the crystal occurs. The decrease in the crystal frequency is faster and more pronounced as the bulk concentration of H2O2 increases. The increase of the mass associated with the crystal is attributed to (29) (a) Tatsuma, T.; Gondaira, M.; Watanabe, T. Anal. Chem. 1992, 64, 11831186. (b) Test, S. T.; Weiss, S. J. J. Biol. Chem. 1984, 259, 399-405.

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Figure 5. Time-dependent frequency changes of a Au-quartz crystal functionalized by a HRP monolayer in the presence of (a) 1 × 10-4 M 1without addition of H2O2, (b) 1 × 10-4 M 1 and 1 × 10-3 M H2O2, and (c) 1 × 10-4 M 1 and 1 × 10-2 M H2O2.

the precipitation and accumulation of 2 on the crystal. For example, we observe a frequency decrease of ca. -500 and -1600 Hz after 1000 s of interaction of the Au-quartz crystal with 1 and H2O2 concentrations corresponding to 1 × 10-3 and 1 × 10-2 M, respectively. These changes in the crystal frequency (∆f) correspond to the precipitation of 2.7 and 8.7 µg‚cm-2 of 2 on the crystal surface after 1000 s of interaction with H2O2 at bulk concentrations of 1 × 10-3 and 1 × 10-2 M, respectively. We see that the crystal frequency continues to decrease without levelingoff to a constant value, but the slope of the changes in frequency tends toward a saturation value. These results are consistent with the fact that the biocatalytic process continuously precipitates 2 onto the crystal surface, resulting in a mass increase on the crystal. The slow-down in the frequency changes as time proceeds is attributed to the coverage of the biocatalyst by the precipitate. This partially inhibits the diffusion of 1 to the enzyme active site, and consequently the precipitation of 2 is slowed. The sensing of H2O2 through the precipitation of the insoluble product (2) enables, in principle, the development of many other sensing electrodes by the coupling of HRP to a second oxidase,19 as outlined in Scheme 1B. The oxidase effects the oxidation of its specific substrate by oxygen with the concomitant generation of H2O2. The amount of H2O2 formed depends on the concentration of the substrate, and it controls the extent of precipitated insoluble product (2). Thus, the precipitate accumulated on the electrode is controlled by the concentration of the substrate analyte in the system. We have used glucose oxidase as a model enzyme to assemble the bienzyme electrode, as shown in Scheme 3. Figure 6 shows the Faradaic impedance spectra observed upon the stepwise construction of the bienzyme-layered electrode, curves a-c, and after insulation of the electrode as a result of the oxidation of glucose and precipitation of 2, curve d. The bare Au electrode reveals a straight line, curve a, implying that the electrontransfer process is not a limiting step of the electrochemical process. Assembly of the HRP monolayer results in an increase in the semicircle diameter, curve b, indicating a higher electrontransfer resistance, Ret ) 4.2 kΩ, at the electrode interface. The insulating hydrophobic protein layer generates a barrier for electron transfer. This protein effect on the interfacial electrontransfer resistance is further emphasized upon the organization of the second enzyme layer consisting of GOx, resulting in an enhanced electron-transfer resistance, Ret ) 4.7 kΩ, curve c. Interaction of the bienzyme-layered electrode with glucose or 1 3178 Analytical Chemistry, Vol. 71, No. 15, August 1, 1999

Figure 6. Nyquist diagram (Zim vs Zre) for the Faradaic impedance measurements in the presence of 10 mM [Fe(CN)6]3-/4- at (a) bare Au electrode, (b) HRP-modified Au electrode, (c) bienzyme-HRPGOx-modified Au electrode, and (d) HRP-GOx-modified Au electrode after incubation in the solution composed of 1 mM 1 and 2 × 10-2 M glucose for 10 min.

does not affect the impedance spectrum of the electrode. Yet, exposure of the functionalized electrode to glucose and 1 results in a very high increase in the semicircle diameter of the impedance spectrum, Figure 6, curve d, indicating an enhanced electrontransfer resistance, Ret ) 12.3 kΩ. Control experiments reveal that the HRP and the GOx monolayers are essential for inducing the change in the electron-transfer resistance in the presence of glucose and 1. That is, modification of the Au electrode by either HRP or GOx does not yield any increase in the electron-transfer resistance upon exposure to glucose and 1. The increase in the electron-transfer resistance upon interaction of the bienzyme-layer electrode is attributed to a biocatalytic cascade leading to the precipitation of 2, Scheme 1B. Biocatalyzed oxidation of glucose by molecular oxygen generates H2O2. The HRP layer mediates the oxidation of 4-chloro-1-naphthol by H2O2 and the formation of the insoluble product (2). The generated insulating film increases the interfacial electron-transfer resistance. The accumulation of the insulating film of 2 on the electrode is time-dependent, and the electron-transfer resistance at the electrode increases as the time of exposure of the electrode to the precipitation process is prolonged, Figure 7. In the specific experiment, the electron-transfer resistance at the electrode, Ret, corresponds to 9.5, 15.3, and 24.2 kΩ, upon the exposure of the bienzyme electrode to glucose and 1, for 4, 10, and 20 min, respectively. Cdl, at the respective electrodes, corresponds to 0.61, 0.52, and 0.45 µF‚cm-2, respectively. Cyclic voltammetry experiments further support the stepwise insulation of the conductive surface upon the assembly of the two enzyme layers and upon the precipitation of the insoluble product (2) on the electrode, Figure 8. The redox-label Fe(CN)63-/Fe(CN)64- reveals a reversible cyclic voltammogram at the bare Au electrode (curve a). Immobilization of HRP (curve b) and then the GOx layers (curve c) on the electrode surface causes an increase in the peak-to-peak separation of the anodic and cathodic waves of the redox probe and a decrease in the amperometric responses. These results indicate slow electron-transfer kinetics at the electrode surface originating from the insulating protein

Figure 9. Electron-transfer resistance, Ret, extracted from the Faradaic impedance spectra registered with the HRP-GOx-modified Au electrode after its incubation for 10 min in the analyte solution composed of 1 mM 1 and glucose in different concentrations. Figure 7. Nyquist diagram (Zim vs Zre) for the Faradaic impedance measurements in the presence of 10 mM [Fe(CN)6]3-/4- at the HRPGOx-modified Au electrode after its incubation in the solution composed of 1 mM 1 and 2 × 10-3 M glucose for (a) 0, (b) 4, (c) 10, and (d) 20 min.

Figure 10. Time-dependent frequency changes of a Au-quartz crystal functionalized with an HRP-GOx bienzyme-layered assembly in the presence of (a) 1 × 10-4 M 1 and without glucose and (b) 1 × 10-4 M 1 and 2 × 10-3 M glucose.

Figure 8. Cyclic voltammograms of a Au electrode measured in the presence of 10 mM [Fe(CN)6]3-/4- after different steps of modification: (a) bare Au electrode; (b) HRP-modified Au electrode; (c) HRP-GOx-modified Au electrode; (d) HRP-GOx-modified Au electrode after incubation in the solution consisting of 1 mM 1 and 2 × 10-2 M glucose. All data were recorded under argon in 0.1 M phosphate buffer solution, pH 7.0; potential scan rate, 100 mV‚s-1.

layers associated with the electrode. Exposure of the bienzymeHRP-GOx-layered electrode to glucose and 1 blocks the electrical response of the electrode, Figure 8, curve d. The insulating film of the insoluble product (2) inhibits the electrical contact between Fe(CN)63-/Fe(CN)64- and the electrode surface. The sensitivity of the Faradaic impedance spectra to the accumulation of the insoluble product (2) on the electrode paves the way for development of very sensitive glucose sensors. The precipitation of 2 by the HRP-mediated process amplifies the analysis of the primary substrate, glucose. Figure 9 shows the calibration curve corresponding to Ret of the electrode at different bulk concentrations of glucose and accumulation of the insoluble product for a fixed time of 10 min. The electron-transfer resistance increases as the bulk concentration of glucose is elevated. These results are consistent with the fact that higher glucose concentrations yield more H2O2 and enhance the precipitation of the insoluble product (2). The time interval for accumulation of the insulating product (2) can control the sensitivity of the electrode, and by allowing the system to generate 2 for long time intervals, high sensitivities can be reached. With the HRP-GOx electrode

we were able to sense glucose levels as low as 1 × 10-8 M upon exposure of the electrode to the precipitation process for 2 h. QCM measurements further confirm the precipitation of 2 by the HRP-GOx bienzyme assembly and in the presence of 1 and glucose. The HRP-GOx-layered assembly was linked to the Auquartz crystal as outlined in Scheme 3. Figure 10 shows timedependent frequency changes of the functionalized crystal upon interaction with 1 in the absence of glucose (curve a) and in the presence of glucose (curve b). In the presence of glucose, a constant decrease in the crystal frequency is observed, indicating that a mass increase takes place on the crystal. For example, after 1500 s the crystal frequency decreases by ca. 700 Hz. Using eq 6, this frequency decrease translates to a mass increase on the crystal that corresponds to 3.8 µg‚cm-2, as a result of the biocatalyzed precipitation of 2. CONCLUSIONS We have demonstrated a novel method for constructing electrochemical biosensor electrodes based on the nonelectrochemical precipitation of an insoluble product on the conductive support and the examination of the extent of electrode insulation by Faradaic impedance spectroscopy and cyclic voltammetry. Similarly, the precipitation of the insoluble product on the transducer surface was probed using microgravimetric quartz crystal microbalance measurements. This method was exemplified by the assembly of a horseradish peroxidase, HRP, monolayer electrode. Hydrogen peroxide, H2O2, was analyzed with the HRP electrode by the biocatalyzed oxidation of 4-chloro-1-naphthol to Analytical Chemistry, Vol. 71, No. 15, August 1, 1999

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the insoluble product (2) that precipitates on the electrode. The extent of electrode insulation was probed by Faradaic impedance spectroscopy, and the interfacial electron-transfer resistance was correlated with the amount of H2O2. Bienzyme-layered electrodes consisting of HRP and an oxidase enable the use of this method to develop various other biosensors. We exemplified this approach by the tailoring of a bienzyme electrode consisting of HRP and glucose oxidase, GOx, and the electrochemical sensing of glucose by the precipitation of an insoluble product on the electrode surface. Biocatalyzed oxidation of glucose by molecular oxygen generates H2O2 that is sensed by the electrode through the HRPmediated oxidation and precipitation of 2. The precipitation of the insoluble product (2) allows the accumulation of the insulating film on the electrode surface. Thus, the precipitation process provides a method for amplifying the analysis process. The biocatalyzed precipitation of insoluble products is often used to probe enzyme functions in ELISA assays. The electronic transduction of the precipitate formation on the

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electronic transducers enable not only the qualitative determination of the analyte but also the quantitative correlation between the amount of precipitate and the analyte concentration. The electrochemical (Faradaic impedance) transduction of the biocatalyzed formation of the insoluble product on the electrode has some advantages over the microgravimetric QCM analysis of the precipitation process. The possibilities for miniaturizing electrodes and for generating electrode arrays pave the way for future inexpensive enzyme-chip devices. ACKNOWLEDGMENT The research was supported by The Israel Ministry of Science within the infrastructure project on Biomicroelectronics.

Received for review February 10, 1999. Accepted April 27, 1999. AC9901541