Electrochemical Protein Chip with Arrayed Immunosensors with

An electrochemical protein chip was microfabricated. A thin-film three-electrode system, including an array of 36 platinum working electrodes, a set o...
0 downloads 0 Views 253KB Size
Download PDF
Anal. Chem. 2003, 75, 1116-1122

Electrochemical Protein Chip with Arrayed Immunosensors with Antibodies Immobilized in a Plasma-Polymerized Film Kenichi Kojima,† Atsunori Hiratsuka,‡ Hiroaki Suzuki,† Kazuyoshi Yano,‡,§ Kazunori Ikebukuro,| and Isao Karube*,‡,§

Institute of Materials Science, University of Tsukuba, 1-1-1 Tennodai, Tsukuba, Ibaraki 305-8573, Japan, Research Center for Advanced Science and Technology, The University of Tokyo, 4-6-1 Komaba, Meguro-ku, Tokyo 153-8904, Japan, and Department of Biotechnology, Tokyo University of Agriculture and Technology, 2-24-16 Naka-cho, Koganei-city, Tokyo 184-8588, Japan

An electrochemical protein chip was microfabricated. A thin-film three-electrode system, including an array of 36 platinum working electrodes, a set of thin-film Ag/AgCl electrodes, and platinum auxiliary electrodes, was integrated on a glass substrate. Capture antibodies were immobilized in a 4.5-nm-thick double layer of a hexamethyldisiloxane plasma-polymerized film. Because of their highly cross-linked network structure, the capture antibodies could be firmly immobilized. No nonspecific adsorption was observed during a series of procedures to detect target proteins, and electrochemical cross talk between neighboring sites was negligible. The sandwich immunoassay was conducted on a single chip using model proteins, r-1-fetoprotein and β2-microglobulin. A distinct current increase following the oxidation of hydrogen peroxide produced by the enzymatic reaction of glucose oxidase was observed, which indicates that the capture proteins could actually bind the target proteins. Two kinds of protein were detected independently on multiple sites with respective capture antibodies. With the end of the human genome project,1,2 research interests are shifting from the detection of specific sequences of DNA or RNA to the functional analysis of respective genes or proteomes expressed from them. Although the DNA sequence contains information about what the future might hold, it does not provide information about the present. For this reason, conventional DNA chips have limited diagnostic applications. Protein chips that are used to check blood or urine for abnormal proteins related to cancer, arthritis, or heart disease will revolutionize clinical analysis. In addition, protein chips will contribute to basic research that will reveal complicated protein-protein * Corresponding author. Tel: +81-426-37-2111. Fax: +81-426-37-3134. Email: [email protected]. † University of Tsukuba. ‡ The University of Tokyo. § Present address: Tokyo University of Technology, 1404-1 Katakura, Hachioji, Tokyo 192-0982, Japan. | Tokyo University of Agriculture and Technology. (1) Craig, J. C.; et al. Science 2001, 291, 1304-1351. (2) International Human Genome Sequencing Consortium. Nature 2001, 409, 860-921.

1116 Analytical Chemistry, Vol. 75, No. 5, March 1, 2003

interactions in different cell types and complex chains of chemical intracellular communication. A final objective of this study is to develop a device that could detect numerous target proteins. As a natural consequence of moving into the postgenome era, the technologies developed for the massive parallel detection of DNA are beginning to be applied to the analysis of proteins. In realizing such devices, optical approaches have been used primarily to develop devices that will detect signals from specific sites.3-5 Although the methods are effective for high-density arrays of tens or hundreds of thousands, electrochemical detection for the above-mentioned clinical applications will be more advantageous for a smaller number of arrays (e.g., fewer than 100). A peculiar feature of electrochemical techniques is that the entire system, including parts for signal processing, will be made very compact. In addition, the time required for detection will be significantly reduced because electrochemical reactions proceeding on the surface of a metal electrode are used rather than a reaction that proceeds in the bulk of the solution. Furthermore, various electrochemical techniques can be applied to achieve a very low detection limit.6 In this study, an electrochemical device containing an array of 36 working electrodes to detect proteins was fabricated. Because thinfilm Ag/AgCl electrodes and auxiliary electrodes were integrated, all measurements could be conducted on a chip. The detection was based on the protocol of enzyme immunoassay. In realizing such a device, the reliability and reproducibility will be influenced by the method of immobilization of receptors on the sensitive area of transducers. As possible candidates, biotin-avidin interaction,7 a plasma-treated nitrocellulose membrane,8 physical adsorption,9 covalent coupling,10 or entrapment (3) Arenkov, P.; Kukhtin, A.; Gemmell, A.; Voloshchuk, S.; Chupeeva, V.; Mirzabekov, A. Anal. Biochem. 2000, 278, 123-131. (4) MacBeath, G.; Schreiber, S. L. Science 2000, 289, 1760-1763. (5) Uetz, P.; Giot, L.; Cagney, G.; Mansfield, T. A.; Judson, R. S.; Knight, J. R.; Lockshon, D.; Narayan, V.; Srinivasan, M.; Pochart, P.; Qureshi-Emili, A.; Li, Y.; Godwin, B.; Conover, D.; Kalbfleisch, T.; Vijayadamodar, G.; Yang, M.; Johonston, M.; Fields, S.; Rothberg, J. M. Nature 2000, 403, 623-627. (6) Heineman, W. R.; Halsall, H. B. Anal. Chem. 1985, 57, 1321A-1331A. (7) Dontha, N.; Nowall, W. B.; Kuhr, W. G. Anal. Chem. 1997, 69, 2619-2625. (8) Rejeb, S. B.; Tatoulian, M.; Khonsari, F. A.; Durand, N. F.; Martel, A.; Lawrence, J. F.; Amouroux, J.; Goffic, F. L. Anal. Chim. Acta 1998, 376, 133-138. (9) Brizzolara, R. A. Biosens. Bioelectron. 2000, 15, 63-68. 10.1021/ac0257391 CCC: $25.00

© 2003 American Chemical Society Published on Web 01/28/2003

in a gel3,11 have been proposed. However, conventional methods require a peculiar time-consuming process and, in general, an increase in the quantity of reagents. In addition, direct immobilization of biomolecules often results in conformational changes and the loss of activity. To solve the problem in the light of its practical applications, biomolecules have been immobilized using plasmapolymerized films (PPFs).12 A very uniform pinhole-free film can be obtained by a completely dry process, which is compatible with other microfabrication processes. In other words, the formation of the membrane can be strictly controlled, which is a marked contrast to the conventional wet processing methods mentioned above. In addition to this, PPFs are chemically and mechanically stable, and their adhesion to substrates is strong, which is critical in obtaining a reproducible sensing performance. Furthermore, because of the high energy of plasma, the selection of monomers is versatile, which affords freedom in the design of a film with preferred functions. In this study, a PPF formed from hexamethyldisiloxane (HMDS) was used to immobilize antibodies on the working electrodes. One aspect in which our device was different from others was that we attempted to embed antibodies with two layers of PPF. The approach has never been tried and is a promising method to immobilize many antibodies mildly using a dry process. Furthermore, the HMDS-PPF has an excellent property with respect to unspecific adhesion. Capture antibodies could actually be immobilized firmly while retaining their activity. Because the method of immobilization is simple and suitable to make arrays of biomolecules, with the electrochemical method of detection, an array of immunosensors could easily be realized, and clear signals could actually be obtained. The fabrication and basic characterization of the device will be presented in this report. EXPERIMENTAL SECTION Materials and Reagents. Glass wafers (7740, 3-in. diameter, 500 µm thick) were purchased from Corning Japan (Tokyo, Japan). A photosensitive polyimide prepolymer (Photoneece 3100) was purchased from Toray (Tokyo, Japan). A positive photoresist (S1400-31) used in fabrication was purchased from Shipley Far East (Tokyo, Japan). HMDS used for monomers of PPF and an RTV-silicone rubber, KE347T, were purchased from Shin-Etsu Chemical (Tokyo, Japan). Mouse monoclonal anti-human IgG was purchased from Calbiochem-Novabiochem. Goat anti-mouse IgG and goat anti-human IgG, both labeled with glucose oxidase (GOx), were purchased from American Qualex. These two were used for preliminary experiments to check the performance of the electrodes with the PPF. A blocking agent (Block Ace) was purchased from Dainippon Pharmaceutical (Osaka, Japan) and diluted with water four times. Human AFP (90890 units/mL) was purchased from Dako Japan (Kyoto, Japan). Human β2-microglobulin (β2MG) was purchased from Sigma. The concentrations of R-1-fetoprotein (AFP) and β2MG were adjusted to 0.5 and 20 µg/ mL, respectively, with a 0.1 M phosphate buffer solution containing 0.1 M NaCl (pH 7.0). Rabbit anti-human AFP IgG was purchased from Dako Japan. Rabbit anti-human β2MG IgG was purchased from Sigma. These two were immobilized in the PPF as capture antibodies. Their concentration was adjusted to 440 µg/mL and (10) Suzuki, M.; Nakashima, Y.; Mori, Y. Sens. Actuators 1999, B54, 176-181. (11) Gerry, N. P.; Witowski, N. E.; Day, J.; Hammer, R. P.; Barany, G.; Barany, F. J. Mol. Biol. 1999, 292, 251-262. (12) Hiratsuka, A.; Karube, I. Electroanalysis 2000, 12, 695-702.

1.2 mg /mL, respectively, with deionized water. Mouse anti-human R-1-fetoprotein IgG and mouse anti-human β2-microglobulin IgG used as secondary antibodies were purchased from the Institute of Immunology (Tokyo, Japan). Their concentration was adjusted to 50 µg/mL with a 0.1 M phosphate buffer solution containing 0.1 M NaCl (pH 7.0). The tertiary antibody, goat anti-mouse IgG labeled with glucose oxidase, was purchased from American Qualex. The concentration was adjusted to 50 µg/mL with a 0.1 M phosphate buffer solution containing 0.1 M NaCl (pH 7.0). All reagents used for fabrication were semiconductor grade and purchased from Kanto Chemicals (Tokyo, Japan) and Wako Pure Chemicals Industries (Osaka, Japan). All reagents used to examine the device performance were purchased from Wako Pure Chemicals Industries. Deionized water was used throughout the experiments. Device Structure and Fabrication. Construction of the device is shown in Figure 1a. The dimensions of the chip are 17 mm × 17 mm. The electrode patterns were formed on a glass substrate. All metal layers were sputter-deposited, and a liftoff process was used to pattern the metal thin films. After a 200-nmthick platinum layer with a 40-nm-thick chromium adhesive layer was patterned, a 400-nm-thick silver layer was formed only on the reference electrode area of the platinum layer. The active areas of the electrodes were delineated, and leads were insulated with a 3.0-µm-thick polyimide layer. The dimensions of the working electrode were 600 µm × 600 µm. The working electrodes were placed in an array of 6 × 6 with the same interelectrode distance of 700 µm. In addition, two pseudo Ag/AgCl reference electrodes and two platinum auxiliary electrodes were formed at the corners. The details of the structure of the Ag/AgCl electrode are described elsewhere.13 The AgCl layer was grown from square pinholes of 30 µm × 30 µm formed in a polyimide protecting layer into the silver layer. A constant current of 500 nA was impressed for 60 min using a galvanostat in an unstirred 1.0 M KCl-HCl buffer solution (pH 2.2). Approximately 30% of the silver layer was converted into AgCl. Twelve chips were batch-fabricated on a 3-in. glass wafer and cut out using a dicing saw (Figure 1b). An epoxy reservoir to store a sample solution was adhered to the periphery of the chip with RTV-silicone rubber (Figure 1a).The internal volume of the container was 300 µL. The chip was set in a specially designed chip carrier when conducting measurements, and all signal lines were connected to the chip through the carrier (Figure 1c). Immobilization of Antibodies in the PPF. Antibodies were embedded in a double layer of PPF. Before growing the PPF, the areas for the reference electrodes and the auxiliary electrodes were protected with a paraffin film, and only the area of the working electrodes was exposed. After the chips with the detecting electrode patterns were rinsed well with deionized water and dried, they were placed in the plasma chamber of a Samco BP-1 Basic Plasma Kit. Then, the chamber was evacuated, and a carrier gas of HMDS was added to the chamber at a constant flow rate of 8.5 sccm. The background pressure was lower than 13.3 Pa. Plasma was discharged at an rf power of 200 W to polymerize and deposit the PPF on the entire active area on the chip. The thickness of the first PPF layer was 1.5 nm, which was measured using an Ulvac ESM-1AT ellipsometer. Then, 0.3 µL of a capture IgG (13) Suzuki, H.; Taura, T. J. Electrochem. Soc. 2001, 148, E468-E474.

Analytical Chemistry, Vol. 75, No. 5, March 1, 2003

1117

Figure 1. Construction of the electrochemical protein chip: (a) decomposed structure shown layer by layer; (b) picture of the chip (R.E., reference electrode; A.E., auxillary electrode; W.E., working electrode); (c) the chip mounted on a chip carrier. Picture seen from the reverse side.

solution (100 µg/mL) was dropped onto the respective working electrode using a micropipet. In this study, we dropped the capture IgG solution manually. However, for the batch fabrication of this device, innovating technologies, including the use of an arrayer and microcontact printing,14,15 and modified techniques from photolithography14 will be of great help. After the chip was dried, another 3.0-nm-thick layer of PPF was deposited on the first PPF layer with the adsorbed antibodies. In this study, another film structure without the upper PPF was also tried and compared with the above-mentioned double-layered structure. Procedures in Optimizing the Formation of the PPF. The HMDS-PPF was characterized by taking cyclic voltammograms using a Bio Analytical Systems (BAS) CV-50W voltammetric analyzer unless otherwise noted. Although our final goal will be to operate the respective immunosensors amperometrically at a constant potential, cyclic voltammetry was used to characterize the basic performance of the system more clearly. One of the working electrodes on the chip was used along with a commercial Ag/AgCl reference electrode (Cypress Systems EE008) and a platinum wire auxiliary electrode. The solution used was a 0.1 M phosphate buffer solution (pH 7.0) containing 0.1 M NaCl. When necessary, 20 mM hydrogen peroxide was added. The potential was scanned between 0 and +1.0 V (vs Ag/AgCl), and the scan rate was set at 50 mV s-1. All electrochemical data, including those for the following experiments using antigens and antibodies, were taken at 25 °C. In a preliminary study to characterize the PPF as a supporting layer to immobilize antibodies, enzyme-labeled antibodies (goat anti-mouse IgG) were reacted with the antibodies (mouse antihuman IgG) immobilized on the PPF. Unlike ordinary cases, the latter became the antigens for the former antibodies. The unusual (14) Blawas, A. S.; Reichert, W. M. Biomaterials 1998, 19, 595-609. (15) Kane, R. S.; Takayama, S.; Ostuni, E.; Ingber, D. E.; Whitesides, G. M. Biomaterials 1999, 20, 2363-2376.

1118

Analytical Chemistry, Vol. 75, No. 5, March 1, 2003

combination was used to simulate the interaction between the secondary and tertiary antibodies explained later in the detection of general proteins. Basic procedures followed the protocol of ELISA. After the surface of the active area of the chip was rinsed with deionized water, the conjugates (0.3 µL, 100 µg/mL) were dropped there, incubated for 10 min at 25 °C, and rinsed well with deionized water. Finally, 200 µL of a 0.1 M phosphate buffer solution (pH 7.0) containing 0.1 M NaCl and 3 M glucose was added to the container. The highest concentration of glucose was used to extract the highest activity of the enzyme (Km, 3.3 × 10-2 M) irrespective of the variation in the concentration of glucose, even when glucose was locally depleted. Because the model analyte was modified with glucose oxidase enzymes, the enzymatic oxidation of glucose proceeded. The hydrogen peroxide produced in the reaction was oxidized electrochemically on the platinum working electrode, and the generated current was detected using the analyzer mentioned above. The accompanying increase in current suggests that the antigens and antibodies have actually bound on the working electrode site. The currents at 700 mV (vs on-chip Ag/AgCl) were read and used to analyze the data. In the experiment to check the function of the blocking agent, 0.3 µL of the solution was dropped to cover the working electrode area with the immobilized antibodies and adhered well to the PPF layer by maintaining it at 4 °C for 10 min. The same procedure was followed thereafter. Procedures Used for the Detection of Proteins. The procedures for the use of immunosensors are as follows. The experiment was conducted at room temperature. After the surface of the PPF with immobilized antibodies (rabbit anti-human AFP IgG or rabbit anti-human β2MG IgG) was rinsed with deionized water; 200 µL of a standard solution containing AFP or β2MG was poured into the reservoir and incubated for 10 min. After the reservoir had been rinsed with deionized water well, 200 µL of a solution containing secondary antibodies (mouse anti-human AFP

IgG or mouse anti-human β2MG IgG) was poured into the reservoir and incubated for 10 min. After the reservoir had been rinsed well with deionized water again, 200 µL of a solution containing tertiary antibodies (goat anti-mouse IgG labeled with GOD) was poured into the reservoir and incubated for 10 min. For the electrochemical detection of signals, 200 µL of a 0.1 M phosphate buffer solution (pH 7.0) containing 0.1 M NaCl or the same solution containing 0.3 M glucose was poured into the reservoir. The same procedure using cyclic voltammetry was followed as mentioned before. In the ELISA, it usually takes a couple of hours for the antigen-antibody reaction to proceed. In the present device, which uses antibodies immobilized on the electrode, 10 min was enough for the reaction to complete (data not shown). RESULTS AND DISCUSSION Characterization of the HMDS-PPF. Prior to the characterization of the HMDS-PPF, the detection of hydrogen peroxide was checked using only the three-electrode system on the chip. The potential of the on-chip Ag/AgCl electrodes settled at a value corresponding to the activity of Cl- ions in the electrolyte, and no problems with respect to durability were observed when a series of measurements was conducted on the chip. The calibration curve drawn using cyclic voltamograms under different hydrogen peroxide concentrations showed linearity up to 4 mM with a correlation coefficient of 0.999. The structure of the HMDS-PPF is similar to that of poly(dimethylsiloxane) (PDMS).16 Unlike conventional organic polymers, PPFs do not generally have chains with a regular unit of repetition but, rather, have an irregular three-dimensional crosslinked network. Furthermore, the ratio of constituent atoms is different between the surface region and the interior of the film.17 Like PDMS, which has an excellent anticoagulative nature,18 HMDS-PPF has been demonstrated to exhibit excellent biocompatibility.19 The adhesion of the film to the underlying substrate was very strong. To detect the hydrogen peroxide produced by the enzymatic reaction, the film must strongly hold the immobilized antibodies and, at the same time, be sufficiently permeable to hydrogen peroxide. As shown later, it has been confirmed that a pinhole-free homogeneous film could actually be obtained even with a thickness of several nanometers and that the permeability of compounds could be reproducibly controlled. Some parameters influence the nature of PPFs.20 These include power for polymerization, pressure, and flow rates. To examine these effects, a 100-nm-thick PPF was formed under various conditions. Concerning power, the permeability of hydrogen peroxide was checked at 50, 100, and 200 W under a pressure of 100 mTorr and a flow rate of 7.2 sccm. With all the PPF formed under the conditions, a clear wave originating from the oxidation of hydrogen peroxide was observed. The observed current decreased as the power increased. This coincides with a general tendency for the film to become denser at higher powers. Because (16) Tajima, I.; Yamamoto, M. J. Polym. Sci. Polym. Chem. 1985, 23, 615-622. (17) Akovali, G.; Rzaev, Z. M. O.; Mamedov, D. G. Eur. Polym. J. 1996, 32, 375-383. (18) Kyriakides, T. R.; Leach, K. J.; Hoffman, A. S.; Ratner, B. D.; Bornstein, P. Proc. Natl. Acad. Sci. U.S.A. 1999, 96, 4449-4454. (19) Kibarer, G. D.; Akovali, G. Biomaterials 1996, 17, 1473-1479. (20) Morosoff, N. In Plasma deposition, treatment, and etching of polymers; d’Agostino, R., Ed.; Academic Press: San Diego, 1997; pp 9-56.

Figure 2. Influence of pressure of HMDS on the output current during continuous operation of the electrode. The pressure was changed between 100 and 400 mTorr under a power of 200 W and a flow rate of 7.2 sccm. The thickness of the PPF was 100 nm. The concentration of hydrogen peroxide was 20 mM.

a denser film is preferable to immobilize antibodies effectively and block intereferences, the power was set at 200 W in the following formation of the HMDS-PPF, taking into account the possibility that HMDS monomers themselves are damaged at higher powers. To obtain an optimum flow rate, the same experiment was conducted at 2.0, 5.2, 7.2, and 8.5 sccm under a pressure of 100 mTorr and a power of 200 W. As the flow rate increased, the oxidation current of hydrogen peroxide increased. However, the durability of the film was markedly lost as the flow rate increased. The observed results could be explained by the decrease of cross-links and the resulting increase in the permeability of the film. Taking the generated current into consideration, the flow rate of 7.2 sccm was considered to be the best. The influence of pressure of HMDS was checked from the viewpoint of the stability of the output current during continuous operation of the electrode (Figure 2). The pressure was changed between 100 and 400 mTorr under a power of 200 W and a flow rate of 7.2 sccm. As the pressure increased, the output current increased. However, the variation of the current level during the measurement also became significant. During this experiment, the pressure was set at 100 mTorr. Because the HMDS-PPF is hydrophobic, the permeation of hydrogen peroxide should decrease as the thickness of the PPF increases. When immobilizing antibodies, their binding portions must be exposed out of the film, and a very small amount of hydrogen peroxide produced by glucose oxidase must be measured in the constructed immunosensor. Considering this, the 100nm-thick thin film used in the characterization here was rather thick. Therefore, the total thickness of the PPF on the device was reduced to 4.5 nm. PPF as a Matrix for the Immobilization of Antibodies. Cyclic voltammograms after the binding of enzyme-labeled antibodies (goat anti-mouse IgG) to the immobilized antibodies (mouse anti-human IgG) are shown in Figure 3 in the presence and absence of glucose. A distinct current increase was observed in cyclic voltammograms in the presence of glucose. In the absence of glucose, however, the oxidization current of hydrogen peroxide was small, and the difference was distinct. The results showed that the binding of antigens and antibodies could be detected without any problems. To check the specificity of binding, goat anti-human IgG was reacted with the immobilized antibodies. Analytical Chemistry, Vol. 75, No. 5, March 1, 2003

1119

Figure 3. Cyclic voltammograms obtained when detecting hydrogen peroxide following the antigen-antibody reaction: (a) in the absence of glucose; (b) in the presence of glucose.

Figure 4. Dependence of the output on the method of immobilization. (a) Antibodies were embedded between two layers of PPF; (b) antibodies were adsorbed physically on a single layer of PPF; (c) no antibodies were immobilized in the PPF. The above figures illustrate the situation for the respective cases. Enzyme-labeled goat antimouse IgG antibodies were bound to the immobilized mouse antihuman IgG antibodies for (a) and (b).

However, no noticeable change in current was observed compared with the case without corresponding antigens. Because the amount of hydrogen peroxide detected on the working electrode is determined by the amount of bound enzymes, effective immobilization of the capture antibodies at a high density was the first requirement. The dependence of the response on the method of immobilization was examined (Figure 4). In one experiment, the antibodies were embedded in the two layers of PPF, while in the other experiment, they were only adsorbed physically on a single layer of PPF. As already shown in Figure 3, a distinct current increase has been observed in the presence of glucose with antibodies immobilized in the double layer. The obtained output current was reproducible with a relative standard deviation within 10%. On the other hand, with the latter membrane, the current increase was very small, even in the presence of glucose, and was comparable with the charging current originating from the method of measurement. A major cause for the latter will be that a substantial portion of antibodies had been lost during the series of procedures and that the amount of antibodies left on the latter membrane was much smaller than that immobilized in the former double-layered membrane. The result indicates that the second layer of PPF effectively anchors the antibodies 1120

Analytical Chemistry, Vol. 75, No. 5, March 1, 2003

adsorbed on the first layer of PPF. In our previous study, proteins were directly placed on a substrate, and a layer of PPF was deposited.21 However, even in the presence of the second PPF layer, the adsorption of the proteins on the substrate was significantly poor even when compared with the above case, in which antibodies were adsorbed on a PPF underlayer. The status of the antibodies embedded between the two layers of PPF is a great concern. In general, PPF is not deposited on the entire substrate but grows only on the already-grown solid phase of PPF. Therefore, it is considered that the antibodies are not covered with the newly grown PPF. In other words, they are anchored with the second PPF layer, and most parts of the antibody molecules are exposed out of the PPF. Because the PPF is very uniform, with very few pinholes, a significant reduction in current should have been observed if the PPF covered the antibodies completely. The result shown in Figure 4 indicates that the assumption is not wrong. As for the orientation of the antibodies in the PPF, there is no reason to think that they are oriented. Unless the interactive portions in the antibodies are buried in the PPF, the antigen-antibody reaction will proceed without any problems. Surface treatment with a blocking agent is required in the ordinary procedure of ELISA to avoid unspecific adsorption of proteins. The process is cumbersome and time-consuming. If the PPF used to immobilize antibodies could minimize unspecific adsorption, it would be more advantageous. As mentioned, the HMDS-PPF has also been shown to have a nature similar to that of PDMS. Therefore, the film is expected to suppress unspecific adsorption effectively. To see the effect, two types of sensors were prepared with antibodies immobilized in the double layer of PPF. In one of them, the blocking agent was applied, while no blocking agent was applied to the other. Surprisingly, the existence of the blocking agent showed no noticeable influence on the output current, as seen in the current levels between the two cases. A similar result has been obtained with the other PPF.21 As a possibility, it is considered that the blocking agent had not adhered to the surface of the PPF and actually had no effect. On the other hand, it is considered that the PPF itself functioned as a blocking agent. In addition, it is possible that proteins that could have bound unspecifically were flushed due to the very short time required for an antigen-antibody reaction. The result suggests that unspecific adsorption is effectively minimized on the surface of the HMDS-PPF. The poor efficiency of immobilization of antibodies when they are directly adsorbed on the HMDS-PPF is considered to reflect the same effect. The result also shows that the procedure with the blocking agent can be omitted in our device. Influence of Cross Talk between Electrodes. Because the device uses an array of 36 working electrodes, electrochemical cross talk might generate residual current, depending upon the distance between electrodes and the time required for the measurements. The cross talk will be caused by hydrogen peroxide produced by the enzymatic reaction on the working electrodes, which will then flow out of the site and affect the output current of the surrounding working electrodes. To examine this effect, a group of nine electrodes was chosen, and the conjugates (21) Miyachi, H.; Hiratsuka, A.; Ikebukuro, K.; Yano, K.; Muguruma, H.; Karube, I. Biotechnol. Bioeng. 2000, 69, 323-329.

Figure 5. Influence of electrochemical cross talk between electrodes. The conjugates were bound to the antibodies immobilized only on the electrode at the center: (a) in the absence of glucose; (b) in the presence of glucose.

were dropped only on the electrode at the center with the immobilized antibodies. The output current generated by the oxidation of hydrogen peroxide was measured in the absence and presence of glucose. The response was measured by cyclic voltammetry in each case, and the time required to obtain the data for all sites was ∼10 min. The result is shown in Figure 5. A significant increase in current was observed only at the central electrode, and the currents on the surrounding electrodes were approximately at the background level. In practical situations, the potential of the working electrodes will be set at a constant potential, and the current on the respective working electrodes will be measured in each case at a constant interval. The required time for all the measurements will be much shorter than that of the present experiment. Therefore, the influence of the cross talk will be disregarded in real situations. Detection of Two Kinds of Proteins. Although the device can analyze 36 proteins, preparing good antibodies specific to target proteins is a common issue in the development of protein chips.22,23 Here, the applicability of the device to multiprotein sensing was tested using two different proteins, AFP and β2MG.

Higher concentrations of AFP can be seen in patients with acute and chronic liver diseases, including hepatitis. β2MG is a marker for certain kinds of cancer relating to white blood cells or kidney disease. Detection of the proteins was based on the sandwich immunoassay. Here, three types of antibodies were used. The first one is capture antibodies immobilized in the PPF. After the target proteins (antigens) are bound to the immobilized antibodies, the secondary antibodies were bound to the proteins. Although the secondary antibodies are usually modified with an enzyme or a fluorescent molecule, we used tertiary antibodies additionally modified with glucose oxidase. The tertiary antibodies recognize the Fc region of the secondary antibodies. If the process of detection requires secondary antibodies modified with the enzyme, their preparation becomes very laborious and time-consuming. However, if the derivation of all the second antibodies is the same and the above-mentioned procedure is used, only one kind of antibody modified with the enzyme is enough. The detection of AFP and β2MG was tested independently. Cyclic voltammograms obtained in the presence of glucose clearly showed that the binding of the proteins was actually detected (Figure 6). To test the applicability to multianalyte testing, the two kinds of capture antibodies were immobilized on four and five working electrodes, as shown in Figure 7, and binding of one of the two kinds of antibodies was checked. The figure also shows the current generated on the respective working electrodes for the respective cases. A distinct current increase was observed at sites where the target proteins bound. On the other hand, the current from sites without specific binding was virtually the same as the background level. The results indicate that different analyte proteins can be detected and differentiated using the immunosensor array. Because our major concern was to detect many proteins electrochemically on a chip, we did not conduct experiments to check the detection limit for the two model proteins. Once target proteins for specific applications are decided and obtained, qualitative analysis, in conjunction with two-dimensional detection, will be more beneficial. As mentioned, 10 min was enough to complete an antigenantibody reaction in the device. Therefore, the minimum time required in performing three antigen-antibody reactions was 30 min. Because it takes 1-2 h for an antigen-antibody reaction to complete in the ELISA, the reduction of time is significant. Furthermore, the time required for the measurement will be on the order of several minutes or tens of minutes, as long as the number of detection sites is limited to fewer than 100. Because cyclic voltammetry was used to characterize the behavior of the

Figure 6. Cyclic voltammograms obtained when detecting hydrogen peroxide in the sandwich immunoassay: (a) in the absence of glucose; (b) in the presence of glucose.

Analytical Chemistry, Vol. 75, No. 5, March 1, 2003

1121

CONCLUSIONS Our on-chip three-electrode system enables electrochemical analysis on a single chip. Antibodies can be immobilized in a double layer of PPF retaining the activity of capture antibodies. Although the PPF is very thin, it can firmly fix the antibodies and, at the same time, control the diffusion of hydrogen peroxide. The process in forming the PPF is compatible with the microfabrication processes. The sandwich immunoassay can be successfully conducted on the chip. A distinct current increase is observed following the enzymatic reaction of glucose oxidase, which clearly shows that target proteins have bound to the capture antibodies. Surface treatment with a blocking agent is not necessary with the PPF. As long as the interelectrode distance is 700 µm, no noticeable cross talk is observed during the measurement. An issue common to the development of this kind of protein chip is to design specific capture molecules that bind to a desired target protein. With a set of capture molecules for clinically or biologically important proteins, our device will play a critical role in clinical diagnosis.

ACKNOWLEDGMENT

Figure 7. Detection of AFP and β2MG at the multiple electrodes. The antibodies for AFP and β2MG were immobilized at the sites in red and blue, respectively. Current increases were measured when AFP (a) and β2MG (b) were added separately onto the chip.

system, charging currents were superimposed as background currents. In practical situations, they will be reduced significantly if the measurement is conducted at a constant potential. (22) Abbott, A. Nature 2002, 415, 112-114. (23) Service, R. F. Science 2001, 294, 2080-2083.

1122 Analytical Chemistry, Vol. 75, No. 5, March 1, 2003

This work was supported by the University of Tsukuba Research Projects. This study was carried out as part of The Project for Technological Development of Biological Resources in Bioconsortia on R&D of New Industrial Science and Technology Frontiers, which was performed by Industrial Science, Ministry of Economy, Trade, and Industry, and entrusted by the New Energy Development Organization (NEDO).

Received for review April 27, 2002. Accepted December 10, 2002. AC0257391