Langmuir 2005, 21, 6025-6029
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Amperometric Immunosensor for Direct Detection Based upon Functional Lipid Vesicles Immobilized on Nanowell Array Electrode Ho Sup Jung,† Jong Min Kim,‡ Jong Wan Park,† Hea Yeon Lee,*,† and Tomoji Kawai† Institute for Scientific and Industrial Research, Osaka University, 8-1 Mihogaoka, Ibaraki, Osaka 567-0047, Japan, and Department of Chemical Engineering, Dong-A University, 840 Hadan-dong, Saha-gu, Pusan 604-714, Korea Received November 13, 2004. In Final Form: March 23, 2005 An original electrochemical immunosensor has now been developed that is based upon the spontaneous immobilization of biotinylated, functional lipid vesicles (FLVs) on a polymeric resist layer. An electrode was fabricated utilizing a form of electron-beam (e-beam) that has been used to fabricate 200 nm (nanoscale) wells in the resist layer covering of the gold electrode. The stability of adhered FLVs upon the nanowell (NW) electrode was observed by atomic force microscopy (AFM). From these observations, we were able to determine that the assembled FLVs primarily adhered as individual molecules, that is, without the aggregation or fusion noted in earlier designs. Additionally, these immobilized FLVs demonstrated clearly defined redox activity in electrochemical measurements. Streptavidin, biotinylated capture antibody, and target proteins were consequently injected in order to set up the immunoassay environment. Electrochemical immunoassay experimentation was performed on the NW array electrode with model proteins, such as human serum albumin (HSA) and carbonic anhydrase from bovine (CAB). We observed a notable current decrease, following the redox path, interrupted by the target HSA, indicating the binding of the capture antibody with the target antigen. On the basis of these results, we propose a new type of immunosensor incorporating this mechanism of spontaneous immobilization of FLVs.
Introduction The development of immunoassay designs, featuring high throughput and sensitivity, for the detection of various biochemical substances is required due to the everincreasing numbers of biological analytes in the clinical, environmental, and bioindustrial fields. In recent years, various conceptual approaches have been reported in the hopes of realizing such immunosensor designs, which are capable of specific detection of the intended reaction between a receptor and its ligand.1-5 In realizing such devices, optical-based systems such as the enzyme-linked immunosorbent assay (ELISA) method6 have been used primarily to develop devices that will detect signals from specific sites. Although these methods have proved to be effective for high-density arrays, (numbering in the tens or hundreds of thousands), it is more advantageous to possess the ability to perform electrochemical detection for immunoassays upon much smaller array systems. Electrochemical immunosensors possess significant advantages over optical approaches; they provide comparable instrumental sensitivity yet are much more * Corresponding author: tel +81-6-6879-8446; fax +81-6-68752440; e-mail
[email protected]. † Osaka University. ‡ Dong-A University. (1) Wang, J.; Liu, R.; Hawkins, M.; Barzilai, N.; Rossetti, L. Nature 1998, 393, 684. (2) Memoli, A.; Annesini, M. C.; Mascini, M.; Papale, S.; Petralito, S. J. Pharm. Biomed. Anal. 2002, 29, 1045. (3) Hayes, F. J.; Halsall, H. B.; Heineman, W. R. Anal. Chem. 1994, 66, 1860. (4) Rubtsova, M. Y.; Kovba, G. V.; Egorov, A. M. Biosensors Bioelectron. 1998, 13, 75. (5) Kumada, Y.; Maehara, M.; Tomioka, K.; Katoh, S. Biotechnol. Bioeng. 2002, 80, 414. (6) Ronnmarkoa, J.; Kampfb, C.; Asplundc, A.; Hoiden-Guthenbergb, I.; Westerc, K.; Pontenc, F.; Uhlena, M.; Nygrena, P. A. J. Immunol. Methods 2003, 281, 149.
amenable to miniaturization.7-10 Additionally, the time required for detection can be significantly reduced due to the electrochemical reactions, which are being measured, occurring on the surface of a metal electrode rather than in bulk solution. Electrochemical immunosensors are generally composed of an electrochemical transducer and a biorecognition substance (typically an antigen or structured antibodies), which is immobilized upon the transducer surface. The formation of an antigen-antibody association and the localization of nonspecific antibodies on this complex (which are usually labeled with enzymes, metal ions, chemiluminescence, or bioluminescence) result in the generation of a differential electrochemical response.1,3,4 The labeling techniques have the ability to improve the detection limits. However, the bioactivity of the complex is adversely affected, and the labor intensity remains quite high in comparison to direct electrochemical detection. Of course, as label-free techniques, they are potentially very versatile. But, they are more susceptible to effects from nonspecific adsorption than established electrochemical methods utilizing, for example, enzymatic labeling.7,11-13 Recently, Peng et al.14 reported on the design of an amperometric biosensor for toxins on a sol-gel thinfilm electrode using lipid bilayer vesicles consisting of ferrocenic diacetylene lipid. In this report, we propose an (7) Sadik, O. A.; Xu, H.; Gheorghiu, E.; Andreescu, D.; Balut, C.; Gheorghiu, M.; Bratu, D. Anal. Chem. 2002, 74, 3142. (8) Kojima, K.; Hiratsuka, A.; Suzuki, H.; Yano, K.; Ikebukuro, K.; Karube, I. Anal. Chem. 2003, 75, 1116. (9) Dequaire, M.; Degrand, C.; Limoges, B. Anal. Chem. 2000, 72, 5521. (10) Xuan, G. S.; Oh, S. W.; Choi, E. Y. Biosensors Bioelectron. 2003, 19, 365. (11) Grant, S.; Davis, F.; Pritchard, J. A.; Law, K. A.; Higson, S. P. J.; Gibson, T. D. Anal. Chim. Acta 2003, 495, 21. (12) Dai, Z.; Yan, F.; Chen, J.; Ju, H. Anal. Chem. 2003, 75, 429. (13) Corry, B.; Uilk, J.; Crawley, C. Anal. Chim. Acta 2003, 496, 103. (14) Peng, T.; Cheng, Q.; Stevens, R. C. Anal. Chem. 2000, 72, 1611.
10.1021/la047212k CCC: $30.25 © 2005 American Chemical Society Published on Web 05/21/2005
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amperometric immunosensor for human serum albumin (HSA) that utilizes redox lipids, which are coanchored to the bilayer vesicles on the nanowell (NW) electrode. The key goals of any immunoassay protocol are first, the immobilization of the biorecognition substance(s) onto suitable substrate(s), and second, preserving sufficient bioactivity of the antibodies’ interaction with the antigens. Therefore, to create the ideal immunosensor, the method of immobilization for receptors on electrode substrates is of utmost importance, since it directly influences reliability and reproducibility. Several methods for accomplishing this have been proposed, such as avidin-biotin interaction,15 covalent coupling,16 a plasma-polymerized film,8 and entrapment in a sol-gel.14 These more conventional methods, however, require multiple steps, are laborintensive, and utilize an increased quantity of reagent. Additionally, direct immobilization on a solid substrate often causes undesirable conformational changes and results in the loss of antibody bioactivity. In this study, the structure of the liposome (or lipid vesicle) that mimics the cell membrane has been carefully considered in an attempt to resolve the aforementioned problems in practical applications by utilizing advanced stabilizers.5,14,17 Identical liposomes are relatively easy to prepare on a large scale and are easily bound to the membrane by utilizing various functional ligands. Additionally, proteins or antibodies immobilized on the liposome membrane retain their structure and bioactivity because liposomes themselves are stable in the presence of various stressing factors such as heating, osmotic pressure, pH, and denaturants.17,18 Moreover, we are able to prevent nonspecific adsorption by taking advantage of the strong hydrophilic properties of the liposome membrane. Although the liposome is useful as a biosensor material, none of the proposed models utilizing liposomes fixed to gold electrodes have allowed the retention of molecular structure or bioactivity. In general, individual liposomes spontaneously form either a flat lipid bilayer structure through fusion processes on their hydrophilic surfaces19 or form an aggregate through similar processes on their hydrophobic surfaces. Indeed, the disruptions of liposomes result in a decreasing electrochemical signal. To resolve these design problems, we have developed a new immobilization method on the gold substrate using electron-beam (e-beam) lithography. This process is based upon preparation of the NW array electrode by utilizing e-beam lithography to spontaneously immobilize the functional lipid vesicles (FLVs) while preserving their stability. In this report, we describe the analytical characteristics of the resultant amperometric immunosensor pertaining to detection of the model proteins human serum albumin (HSA) and carbonic anhydrase from bovine (CAB). Experimental Section Chemicals. 1-Palmitoyl-2-oleoyl-sn-glycero-3-phosphocholine (POPC), 1,2-dimyristoyl-sn-glycero-3-phospho-1-glycerol (DMPG), and 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine N-(Cap biotinyl) (biotinyl Cap-PE) were purchased from Avanti Polar Lipids Inc. (Alabaster, AL). N-(10,12-Pentacosadiynoic)acetylferrocene (Fc-PDA) was synthesized by a previously reported method.14 (15) Dontha, N.; Nowall, W. B.; Kuhr, W. G Anal. Chem. 1997, 69, 2619. (16) Suzuki, M.; Nakashima, Y.; Mori, Y. Sensors Actuators B 1999, 54, 176. (17) Rozema, D.; Gellman, S. H. Biochemistry 1996, 35, 15760. (18) Yoshimoto, M.; Shimanouchi, T.; Umakoshi, H.; Kuboi, R. J. Chromatogr. B 2000, 743, 93. (19) Binder, W. H.; Barragan, V.; Menger, F. M. Angew. Chem., Int. Ed. 2003, 42, 5802.
Jung et al. Biotinlayed anti-human serum albumin (anti-HSA) was purchased from American Qualex International Inc.d (San Clemente, CA). HSA and CAB were obtained from Sigma (St. Louis, MO). All other chemicals were purchased from Wako Pure Chemical Industries, Ltd. (Osaka, Japan). FLVs and NWs Electrode. The FLVs were prepared by an extrusion method (LipoFast; Avestin Inc.).20 Briefly, the FLV solution was composed on POPC, DMPG, biotinyl Cap-PE, FcPDA, and cholesterol at a molar ratio of 7:3:1:1:3. The lipid solution was dehydrated by a 100 mM phosphate-buffered saline solution (PBS, pH 7.4). Subsequently, five freeze-thaw cycles were applied and the lipid vesicles were repeatedly extruded through a polycarbonate film with 100 nm pores by use of an extruder device in order to produce FLVs of uniform size. The lipid vesicle size was confirmed in all cases, prior to immobilization, by the dynamic light scattering method (DLS-700 Ar; Otsuka Electronics Co., Ltd., Japan). Also, we specially fabricated multiarray electrodes on which to immobilize the FLVs. An e-beam lithography technique was used for the fabrication of a substrate by use of a standard positive polymer resist (ZEP520; Zeon Corp.). ZEP is a copolymer of chloromethacrylate and methylstyrene. It is generally used for high-resolution e-beam lithography instead of poly(methyl methacrylate) (PMMA). The e-beam exposed area was approximately 100 × 100 µm2. Arrays of 200-nm holes were formed in the resist layer by e-beam lithography. The total fabrication time was about 30-60 min depending on the structure of the 100 × 100 µm2 area. The dose amount utilized was adjusted to 150 C/cm2 by use of a 75 kV scanning electron microscope (ELC-2; Elionix Co. Ltd.). The immobilization of the FLVs was performed by incubating a 10 µL portion of the liposome solution on the e-beam exposed substrate for 30 min. The FLV-modified electrode was then carefully rinsed with 100 mM PBS solution to remove any unbound FLVs. To immobilize the capture antibody (anti-HSA), 2 µL of streptavidin (10 µg/mL) was imposed upon the FLVmodified electrode for at least 30 min and was then washed with Millipore Milli-Q (18 MΩ‚cm) water. A 10 µg/mL portion of biotinylated anti-HSA was dropped on the modified electrode for 30 min, and target proteins (10 µg/mL) were injected successively for 15 min at 37 °C. All surface treatment processes were carried out at room temperature in a high humidity environment. Surface Plasmon Resonance and Electrochemical Measurements. SPR is measured at 25 °C by real-time biomolecular interaction analysis in a Biacore 3000 instrument (Biacore AB, Uppsala, Sweden). The NW sensor chip consisting of a gold electrode, which was modified with e-beam lithography, was obtained from Biacore (SIA Kit Au). The FLV solution was injected into the flow cells with 100 mM PBS (pH 7.4) as carrier buffer solution. Streptavidin (10 µg/mL), anti-HSA (10 µg/mL), 0.1 mg/ mL BSA (blocking solution), and HSA were successively injected at a flow rate (5 µL/min). For all other washing processes the flow rate was set to 10 µL/min. Electrochemical measurements were performed on a BAS 100 B/W potentiostat (Bioanalytical Systems, Inc.) at room temperature. Electrodes used were gold working NW electrodes, an Ag/ AgCl reference electrode in 3 M KCl, and a platinum-wire counterelectrode (1 mm). Square-wave voltammetry (SWV) measurements were performed in a solution containing 100 mM PBS solution (pH 7.4) at a scan rate of 100 mV/s. SWV was performed with the following parameters: -0.2 V initial potential, 0.8 V end potential, 25 mV amplitude, 4 mV step potentials, and 50 Hz frequency. The normalized signal response (S), which is defined as (St/Sab) × 100, was determined under the following experimental conditions: Sab is the peak current in immobilized capture anti-has, and St is the peak current after binding to the target proteins for 15 min. Tapping-Mode AFM. The FLVs were immobilized by dropping a 10 µL portion of the FLV solution on the sample substrate for 30 min. The substrate was then carefully rinsed five times with a PBS solution (pH 7.4) in preparation for AFM measurement. AFM measurements were performed promptly after immobilization. All tapping-mode AFM images were obtained on a commercially available instrument (Dimension 3100; Veeco (20) MacDonald, R. C.; MacDonald, R. I.; Menco, B. P.; Takeshita, K.; Subbarao, N. K.; Hu, L. Biochim. Biophys. Acta 1991, 1061, 297.
Nanowell Electrode-Based Amperometric Immunosensor
Figure 1. Schematic illustration of an FLV immunosensor system: (1) N-(10,12-pentacosadiynoic)acetylferrocene (FcPDA); (2) 1,2-dioleoylphosphatidylethanolamine-N-caproylamine (Cap-PE); (3) streptavidin; (4) biotinylated capture antibody; (5) target protein. Instruments) without any environmental control. A normal tapping-mode silicon cantilever with an oscillation frequency of 350 kHz and a spring constant of 40 N/m (OTESP7; Olympus Optical Co. Ltd., Japan) was used for all AFM imaging.
Results and Discussion The principles of direct detection with FLVs as the amperometric immunosensor are detailed in Figure 1. Initally, the FLVs are firmly immobilized around the NWs of the arrayed electrode. The FLVs themselves are composed on supramolecular assemblies containing biotinyl Cap-PE, cholesterol, and a facile redox probe (FcPDA), which is important for electrochemical detection. Second, streptavidin, which is used to immobilize the capture antibody, was injected onto the FLVs to bind with biotinyl Cap-PE. Finally, the antigen was captured by the antibody and then the immunosensor was evaluated by observation and measurement of the current change before and after injection of the antigens. It is imperative that the FLVs are immobilized in a stable form onto the NW electrode without resulting rupture or formation of planar bilayers. Although the redox probes are coanchored into the lipid bilayer vesicles (to provide electrochemical signaling since the antigen is electrochemically inactive), the aggregation or fusion of lipid vesicles allows us to decrease the sensing signal because this design mandates direct investigation of the dependency of electron transport on the amount of biomolecular recognition that has taken place on the vesicles. We have observed the stability of adhered FLVs on the NW electrodes by AFM. Figure 2 shows photography (a) and typical AFM images (b, c) of the e-beam exposed NW array electrode for immobilized FLVs. The size of the whole electrode substrate is 2.5 × 2.5 cm2, the experimental area is 500×500 µm2, and the e-beam developed area is 100 × 100 µm2 (Figure 2a). The electrode was fabricated with a 200 × 200 nm2 well size at 500-nm
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Figure 2. Picture of the nanowell arrayed electrode (a) and AFM topographic images (b, c) for the sample substrate used. The size of the whole substrate is 2.5 × 2.5 cm2, the experimental area is 500 × 500 µm2, and the e-beam developed area is 100 × 100 µm2 (a). Panel b displays an AFM image of e-beam developed area on the scale of 15 × 15 µm2, and panel c displays an AFM topographic image of a typical FLV immobilized pattern on the e-beam exposed surfaces in the scan areas of 2.5 × 2.5 µm2 (white circle indicate the 200 nm nanowell).
intervals by e-beam lithography on the resist layer surface. Figure 2 panels b and c show immobilized FLVs as a whole (15 × 15 µm2) and on an enlarged scale (2.5 × 2.5 µm2). These AFM images allow us to quantify the FLVs’ size and morphological distribution on the e-beam-exposed NW arrayed electrode. Sizes of the individual FLVs were shown to be 90-150 nm full width at half-maximum (fwhm), which are comparable to the size of the originally designed FLVs. Moreover, we find that the FLVs are mainly adhered as individual entities without the aggregation or fusion common to previously outlined results. This mechanism for FLV immobilization on a resist surface was first suggested by us.21 Briefly, we infer that vesicle immobilization is engendered by the electrostatic charge interaction between the polymeric resist surface and FLVs. The immobilized FLVs around and on the interface between the e-beam exposed area and the undeveloped resist area offer confirmation that highly dense liposome immobilization was achieved on the e-beam exposed side. [See Supporting Information Figure S1, which shows (a) the bare NW electrode and (b) two of the immobilization regions. Enlarged images can be seen in Figures S1c,d.] Streptavidin plays a key role in the binding process between the receptor and the FLV’s sensor ligand. The biotinylated capture antibody (anti-HSA) was firmly immobilized as a receptor onto the FLV membrane by streptavidin-biotin interaction. We used SPR to con(21) Kim, J. M.; Jung, H. S.; Park, J. W.; Lee, H. Y.; Kawai, T. J. Am. Chem. Soc. 2005, 127, 2358.
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Figure 3. SPR sensorgrams for the immunoassay step: injection of FLV solution (30 µL, 10 mM) over sensor chip flow cell. (a) Sensor chip treated with streptavidin (20 µL, 10 µg/ mL) at time 150 s; (b) without streptavidin; (c) sensor chip treated with streptavidin but not with blocking solution (0.1 mg/mL BSA). After the anti-HSA (20 µL, 10 µg/mL) was injected over all sensor chips for 300 s at 25 °C, the HSA was injected for 300 s. For all washing processes, the flow rate was set to 10 µL/min with 100 mM PBS (pH 7.4).
firm whether the anti-HSA was indeed firmly immobilized onto the FLVs, which were then treated both with and without streptavidin. After immobilization with streptavidin, the FLVs were treated with blocking solution (0.1 mg/mL BSA) to avoid nonspecific adsorption. The result of the immunoassay is shown in Figure 3. As expected, a sensor chip surface treated with streptavidin (a) is exposed to anti-HSA and shows a strong response (450 ( 10 RU). Although a sensor chip that was not treated with streptavidin (b) also showed a response (60 ( 5 RU), the increase observed is not nearly as significant as that of the sensor chip treated with streptavidin. These results suggest that biotin-streptavidin interaction may be useful for immobilization of the antibody. Additionally, the effect of immobilized FLVs on the sensor chip has been examined by SPR with a blocking solution. This nonspecific interaction is of major concern in immunosensor design, since electrochemical discrimination between nonspecific and specific adsorption in electrochemical sensing of labelfree antigens cannot be distinguished. The results obtained (sensor chip c) display only a slight difference (∼5 RU) from those for sensor chip a. It is proposed that nonspecific binding processes could be avoided and background interaction reduced to a minimum by immobilization of FLVs, thereby permitting the attainment of highly reproducible sensor signals.22 We evaluated cyclic voltammograms, shown in Figure 4 (inset), of the liposome layer both in the presence and absence of Fc-PDA liposome membrane probes in 100 mM PBS solution. Well-defined current responses were obtained for the Fc-PDA-containing liposome electrode. However, no redox response was observed for the immobilized liposome in the absence of Fc-PDA. For that reason, it is supposed that the liposome modified with Fc-PDA can be used to detect the target antigen by redox current changes on the immobilized electrode. Although FLVs are generally immobilized on the resist surface, the e-beam-exposed NWs assist electron transport via FLVs immobilized around the NW’s edge areas (Figure 2c). The electron transport can flow through the NW and the resist layer cracks similar to sol-gel film entrapped liposome (22) Webster, M. A. C. C.; Packman, L. C.; Williams, D. H.; Gray, J. C. Nucleic Acids Res. 2000, 1618, 28.
Jung et al.
Figure 4. Square-wave voltammograms obtained upon detection of target protein (HSA) following the modification where streptavidin (a) and biotinylated capture-antibody (anti-HSA) bound onto the FLVs immobilized electrode. Cyclic voltammograms of the liposome layer in both the presence and absence of Fc-PDA probes in membranes are shown in the inset.
electrodes proposed by Stevens et al.14 That study proposed that ferrocene species, which are covalently adhered to the bulky vesicles, on the gel surface near and far from the crack sites participate in the electron transfer. Similarity, in our experiments, the current response is a typical diffusion-control process, indicating that FLVs on the NW electrode near and far from the wells sites participate in the electrode’s signal. Therefore, our NW electrode plays a key role in transporting electrons through the path and immobilization onto the substrate which is flexible in its design. Figure 4 shows square-wave voltammograms (SWV) results obtained immediately after the immobilization of streptavidin (a), anti-HSA (b), and HSA (c) on the FLVs. Current responses of the FLV electrodes decreased successively following immobilization of anti-HSA and HSA. Current signals of the FLV electrode decreased noticeably when the HSA interacted with the FLV sensors. This result suggests that the target antigen binding to anti-HSA on the FLV’s membrane subsequently blocks the electron transport path into the NW electrode. Therefore, the stacking of target antigen onto the FLV leads to the suppression of further electron-transfer currents. To further examine the specific binding interaction, the detection capabilities for HSA and CAB were tested with the multichannel NW arrayed electrode design shown in Figure 5. The subsequent results also demonstrate the normalized signal generated on respective working electrodes for each case. A distinct normalized signal increase was observed at the binding sites for HSA that were treated with streptavidin. In the case of the injection of CAB, no apparent signal was observed regardless of whether streptavidin was applied to the FLVs. On the other hand, the signals from sites without specific binding or streptavidin were actually the same as the background levels (Sab) noted. Regarding channels 2 and 4, which were not treated with streptavidin, the current changes show a similar tendency as noted in the SPR sensorgrams. These results indicate that if capture antibody is strictly immobilized on the FLV membrane, specific detection of the target protein can be achieved with proper differentiation from other proteins. Furthermore, these results also suggest that the FLVs successfully function as stabilized immunosensor substrates for suppression of nonspecific protein adsorption on the FLVs. Our major concern in this study was to detect a target antigen electrochemically using lipid vesicles immobilized onto the NW electrode.
Nanowell Electrode-Based Amperometric Immunosensor
Figure 5. Detection of HSA and CAB at the multichannel nanowell arrayed electrodes. The antibody for HSA was immobilized on all electrodes with or without streptavidin (SAv). Normal current increase was observed when HSA and CAB were injected separately onto the electrodes.
However, we have yet to perform specifically designed experimentation relative to the immobilization of different antibodies for multiple protein sensing. This is an area where further work is planned. Conclusions A newly designed immunosensor for amperometric detection of a target antigen (HSA) has been developed that utilizes FLVs on a NW electrode. This sensor has so
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far facilitated the fabrication of a planar immunosensor chip using integrated-circuit technology, allowing rapid access of the target antigens to the FLVs for differential detection. Compared to other designs involving lipid bilayer vesicles for the immunosensor, the process has remarkably reduced the steps necessary for detection (similar to use of the sandwich method) as well as significantly reducing nonspecific interactions. One of the more important conclusions resulting from this work is that the shape and functional structure of lipid vesicles are maintained during fabrication of the electrode. They can be immobilized individually onto the resist layer around the NW, facilitating lateral electron transport with the FLVs. These factors can be combined for molecular recognition and utilized for signal detection at the moment the target antigen binds with the capture antibody by blocking the electron transport path via near and far FLVs. Therefore, the capture antibody (anti-HSA) can be immobilized firmly and apparently made to interact specifically with the target antigen, which retains its orientation and bioactivity. This is difficult when the liposome layer aggregates or forms a fusion layer on the electrode surface. Acknowledgment. Support for this work was generously provided by the New Energy and Industrial Technology Development Organization (NEDO), Japan, and Matsushita Electric Industrial Co., Ltd., and is gratefully acknowledged. Supporting Information Available: AFM images of the bare NW electrode and two of the immobilization regions. This information is available free of charge via the Internet at http://pubs.acs.org. LA047212K
