A Plasma-Polymerized Film for Surface Plasmon Resonance

Runa Nakamura,† Hitoshi Muguruma,*,‡,§ Kazunori Ikebukuro,‡ Satoshi Sasaki,‡ Ryohei Nagata,†. Isao Karube,‡ and Henrik Pedersen|. Central...
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Anal. Chem. 1997, 69, 4649-4652

A Plasma-Polymerized Film for Surface Plasmon Resonance Immunosensing Runa Nakamura,† Hitoshi Muguruma,*,‡,§ Kazunori Ikebukuro,‡ Satoshi Sasaki,‡ Ryohei Nagata,† Isao Karube,‡ and Henrik Pedersen|

Central Research Institute, Dai Nippon Printing Company, Ltd., 250-1 Wakashiba, Kashiwa, Chiba 277, Japan, Research Center for Advanced Science and Technology, University of Tokyo, 4-6-1 Komaba, Meguro-ku, Tokyo 153, Japan, and Department of Chemical and Biochemical Engineering, Rutgers, The State University of New Jersey, Piscataway, New Jersey 08855-0909

We propose a novel ethylenediamine plasma-polymerized film matrix that is deposited on gold surfaces for use in surface plasmon resonance immunosensing. The films formed on the gold surfaces are extremely thin (∼100 nm), are homogeneous, demonstrate good adhesion, have a flat profile, and incorporate amino groups to introduce a chemical functionality. A sensor chip made with the film has many advantages when compared with conventionally used designs such as carboxylated dextran hydrogels. For example, a sensor chip made with the film shows a better sensor response than a conventional design partly because antibodies are densely and two-dimensionally immobilized onto the surface of the plasma-polymerized film. Surface plasmon resonance (SPR) detection for monitoring biomolecular interactions such as those occurring during antigenantibody, receptor-ligand, or nucleic acid base pair association has been widely used as a label-free, real-time measurement system by many researchers.1-15 The interactions play an extremely important role in both applied and fundamental studies within the biochemical and medical fields. One system, BIAcore, has been developed by BIAcore AB (Uppsala, Sweden)9-15 into a †

Dai Nippon Printing Co., Ltd. University of Tokyo. § Rutgers, The State University of New Jersey. | Present address: Rutgers, The State University of New Jersey. (1) Liedberg, B.; Nylander, C.; Lungstro ¨m, I. Sens. Actuators 1983, 4, 299304. (2) Flanagan, M. T.; Pantell, R. H. Electron. Lett. 1984, 20, 968-970. (3) Kooyman, R. P. H.; Kolkman, H.; Van Gent, J.; Greve, J. Anal. Chim. Acta 1988, 213, 35-45. (4) Cullen, D. C.; Lowe, C. R. Sens. Actuators B 1990, 1, 576-579. (5) Brigham-Burke, M.; Edwards, J. R.; O’Shannessy, D. J. Anal. Biochem. 1992, 205, 125-131. (6) O’Shannessy, D. J.; Brigham-Burke, M.; Peck, K. Anal. Biochem. 1992, 205, 132-136. (7) Buckle, P. E.; Davis, R. J.; Kinning, T.; Yeung, D.; Edwards, P. R.; PollardKnight, D.; Lowe, C. R. Biosens. Bioelectron. 1993, 8, 355-363. (8) Parsons, I. D.; Persson, B.; Mekhalfia, A.; Blackburn, G. M.; Stockley, P. G. Nucleic Acid Res. 1995, 23, 211-216. (9) Lo¨fa˚s, S.; Johnsson, B. J. Chem. Soc., Chem. Commun. 1990, 1526-1528. (10) Lo¨fa˚s, S.; Magnus, M.; Ro¨nnberg, I.; Stenberg, E.; Liedberg, B.; Lundstro¨m, I. Sens. Actuators B 1991, 5, 79-84. (11) Karlsson, R.; Michaelsson, A.; Mattsson, L. J. Immunol. Methods 1991, 145, 229-240. (12) Johnsson, B.; Lo¨fa˚s, S.; Lindquist, G. Anal. Biochem. 1991, 198, 268-277. (13) Stenberg, E.; Persson, B.; Roos, H.; Urbaniczky, C. J. Colloid Interface Sci. 1991, 143, 513-526. (14) Liedberg, B.; Lundstro ¨m, I.; Stenberg, E. Sens. Actuators B 1993, 11, 6372. (15) Fa¨gerstam, L. G.; Frostell-Karlsson, A° .; Karlsson, R.; Persson, B.; Ro¨nnberg, I. J. Chromatogr. 1992, 597, 397-410. ‡

S0003-2700(97)00571-4 CCC: $14.00

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widely used commercial instrument.16 Recently, a compact and optoelectronically integrated SPR sensing system has also been evaluated.17 A surface plasmon wave is a nonpropagating evanescent wave formed at a metal-coated surface when light is directed toward the interface at a very specific reflection angle. It extends from the metal surface into the sample solution, decaying exponentially as a function of distance. Refractive index changes, localized near the gold and resulting from the interaction between biomolecules such as antigen-antibody pairs, perturb the evanescent wave and alter the propagation characteristics of the surface plasmon. As a result, the resonance angle changes. In practice, the BIAcore system tracks the change of the resonance angle, which is expressed as resonance units (RU) and is proportional to the mass interacting at the sensor surfaces. Since the evanescent wave decays exponentially as a function of distance, the interaction of biomolecule interactions should be carried out close to the gold surface to obtain a high sensitivity. A typical penetration depth of the evanescent wave from the gold surface is ∼200-300 nm.13,14 When antibodies are directly immobilized on the gold surface, it is often difficult to attach large amounts of antibodies with good adhesion. As a result, antibodies are apt to be denatured and the sensor surface suffers from low surface coverage as well as nonspecific adsorption. Thus, many types of coatings have been considered to facilitate antigen attachment. In the BIAcore system, a carboxylated dextran9,12,13 hydrogel is often used on the gold surface, as shown in Figure 1a, and is designated as the “sensor chip CM5”.18 In our work, an alternative coating is investigated. In particular, a plasma-polymerized film (PPF) is deposited using glow discharge or a plasma of organic vapors. The resulting coating is extremely thin (∼100 nm), is homogeneous, has good adhesion to the gold substrate, and forms a flat surface because of its highly cross-linked structure.19 Previously, our group demonstrated that ethylenediamine PPFs were suitable for use on gold surfacemodified quartz crystal microbalances as immunosensors.20,21 There are many additional reports concerning biomolecular attachment onto PPF surfaces for various applications including glucose sensing,22,23 immunoassays,24 semiconductor immun(16) There are numerous research reports using the BIAcore system as a tool. See: www.biacore.com/products/sensor.html on the Internet, for example. (17) Melendez, J.; Carr, R.; Bartholomew, D. U.; Kukanskis, K.; Elkind, J.; Yee, S.; Furlong, C.; Woodbury, R. Sens. Actuators B 1996, 35, 212-216. (18) There are four kinds of sensor chip commercialized by BIAcore AB, the sensor chip HPA, SA, NTA and CM5. See: www.biacore.com/scientific/ reflist.html.

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Figure 1. Schematic illustration of the SPR sensor chip on which the gold surface was modified by (a) carboxylated dextran (sensor chip CM5) and (b) plasma-polymerized film (sensor chip PPF), where d denotes the thickness of each layer.

osensing,25 and alteration of biomaterials.26 PPFs also have potential usage for gold surface modification of SPR immunosensing chips. In this context, we propose a novel sensor chip using ethylenediamine PPFs as a coating matrix on gold surfaces for SPR immunosensing and show that this method provides significant advantages over current strategies. EXPERIMENTAL METHODS The sputtering apparatus for metal thin-film preparation was obtained from Shibaura Engineering Works Co., Ltd. (Model CFS4ES-231, Tokyo, Japan) and the deposition was performed under normal conditions. The plasma-polymerization apparatus and the detailed experimental conditions of film preparation have been described in previous reports.20,21 In brief, chromium and gold were first sputtered on a cleaned glass slide (18 × 18 × 0.15 mm) to a thickness of 4 and 50 nm, respectively. Chromium was deposited to obtain good adhesion between the gold and the glass slide. The gold film thickness was determined from a surface profiler (Dektak3 ST, Veeco Instruments Inc., Tokyo, Japan). Subsequently, an ethylenediamine PPF was deposited. Its thick(19) There are many reports concerned with properties of plasma-polymerized films. See these books and references therein: Plasma Chemistry of Polymers; Shen, M. Ed. Marcel Dekker: New York, 1976. Yasuda, H. Plasma Polymerization; Academic Press: New York, 1985. Plasma Polymerization and Plasma Interactions with Polymeric Materials; Yasuda, H. K., Ed.; Journal of Applied Polymer Science, Applied Polymer Symposia 46; Wiley: New York, 1990. Plasma Deposition of Polymeric Thin Films; Danilich, M. J., Marchant, R. E., Eds.; Journal of Applied Polymer Science, Applied Polymer Symposia 54; Wiley: New York, 1994. (20) Nakanishi, K.; Muguruma, H.; Karube, I. Anal. Chem. 1996, 68, 16951700. (21) Nakanishi, K.; Adachi, M.; Sako, Y.; Ishida, Y.; Muguruma, H.; Karube, I. Anal. Lett. 1996, 29, 1247-1258. (22) Kampfrath, G.; Hintsche, R. Anal. Lett. 1989, 22, 2423-2431. (23) Danilich, M. J.; Gervasio, D.; Marchant, R. E. In Biosensors and Chemical Sensors; Edelman, P. G., Wang, J., Eds.; ACS Symposium Series 487; American Chemical Society: Washington, DC, 1992; pp 84-98. (24) Muratsugu, M.; Kurosawa, S.; Kamo, N. J. Colloid Interface Sci. 1991, 147, 378-386. (25) Jimbo, Y.; Saito, M. J. Mol. Electron. 1988, 4, 111-118. (26) Hoffman, A. S. In Plasma Polymerization and Plasma Interactions with Polymeric Materials; Yasuda, H. K., Ed.; Journal of Applied Polymer Science, Applied Polymer Symposia 46; Wiley: New York, 1990; pp 341-359.

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ness and refractive index were measured by ellipsometry and found to be 100 nm and 1.63 (at 760 nm), respectively. The amino groups on the film surface were investigated with X-ray photoelectron spectroscopy (XPS).20 This chip is designated “sensor chip PPF” and was set into a BIAcore 2000 instrument. Subsequent processes, e.g., immobilization of antibodies and measurement of immunoreaction kinetics, were carried out with automated equipment associated with the instrument. The light source equipped with the BIAcore is a 760 nm light-emitting diode. Antibodies were rabbit anti-human albumin IgG obtained from Sigma Chemical Co. (St. Louis, MO). Human serum albumin (HSA) as a target molecule (antigen) and bovine serum albumin (BSA) were purchased from Wako Co. (Osaka, Japan). Typical procedures are as follows: 0.15 M HEPES buffer (pH 7.4) was pumped at 5 µL/min to prepare the surface; 100 µL of 5% glutaraldehyde solution, 750 µL of 50 µg/mL antibody, 35 µL of ethanolamine, and 5 µL of 0.1 M HCl were sequentially pumped at 5 µL/min except for the antibody solution, which was pumped at 1 µL/min. The subsequent immunoreactions with various concentration of HSA using both the sensor chip PPF and CM5 modules were carried out at 25 °C in 0.15 M HEPES buffer (pH 7.4). The reaction time was 7 min (flow rate 5 µL/min), and the experiments were carried out in triplicate. The procedure for the sensor chip CM5 was obtained from BIAcore, and activation of the dextran matrix was carried out according to reports in the literature12,13 except for the immunoreactions just described. A change of resonant angle equal to 10-4° was defined as 1 RU. RESULTS AND DISCUSSION The plasma-polymerized technique enables various kinds of monomers to be polymerized even if the monomer has no double bond in its structure.19 However, important reasons should be noted as to why we selected ethylenediamine as a monomer. One is that the surface of ethylenediamine PPF is hydrophilic and has many amino groups, which are the active groups for further modification.20 Another reason is that the polymerization rates of aliphatic amines (e.g., ethylenediamine20 and n-propylamine27) are slow compared to those of other well-known nitrogencontaining PPFs27,28 such as allylamine,27 acrylonitrile,29 and pyridine.30 This allows for easy control of the film thickness with good reproducibility. In our experiment, the deposition rate of ethylenediamine PPF was 100 nm/min. Figure 1 is a schematic illustration of the SPR sensor chip on which the gold surface is modified by (a) carboxylated dextran (sensor chip CM5) and (b) PPF (sensor chip PPF). In normal use,9,10,12 good adhesion of the matrix to the gold surface is achieved by adsorption of long-chain 1,ω-hydroxyalkanethiols to generate a self-assembled monolayer31 also called the “linker layer”.9 This layer serves partly as a barrier to prevent proteins and other ligands from coming into contact with the metal and partly as a functionalized structure for further modification of the surface.9 In particular, a dextran matrix is attached to the linker layer. In our case, the PPF serves not only as the linker layer but also as the matrix; the PPF was deposited on the gold surface (27) Bell, A. T.; Wydeven, T.; Johnson, C. C. J. Appl. Polym. Sci. 1975, 19, 1911. (28) Yasuda, H.; Lamaze, J. Appl. Polym. Sci. 1973, 17, 201. (29) Hozumi, K.; Kitamura, K.; Hashimoto, H.; Hamoka, T.; Fujisawa, H.; Ishizawa, T. J. Appl. Polym. Sci. 1983, 28, 1651. (30) Tajima, I.; Suda, T.; Yamamoto, M.; Satta, K.; Morimoto, H. Polym. J. 1988, 20, 919. (31) Bain, C. D.; Troughton, E. B.; Tao, Y. -T.; Evall, J.; Whitesides, G. M.; Nuzzo, R. G. J. Am. Chem. Soc. 1989, 111, 321-335.

Figure 2. Correlation between resonance unit changes and human serum albmin concentration. The gold surfaces of the sensor chip were modified by carboxylated dextran (b; sensor chip CM5) and plasma-polymerized film (O; sensor chip PPF). Fitted data. sensor chip CM5: RU ) 44.24 + 33.67 log[HSA]. Correlation coefficient R ) 0.999. The intercept at the x axis is 0.049 mg/L. Sensor chip PPF: RU ) 55.19 + 21.32 log[HSA]. R ) 0.988. The intercept at the x axis is 0.0026 mg/L.

with good adhesion, and PPF layers have a highly cross-linked and three-dimensional polymer matrix. In related reports, PPFs have also been applied to serve as passivation films for metals32 and semiconductors.33 Antibodies are immobilized only on the PPF surface and do not penetrate in the film (vide infra). Antibodies can be immobilized, for example, with glutaraldehyde as a cross-linking agent between the amino groups on the film surface and the antibodies’ amino groups (See Figure 1b). Figure 2 shows a comparison of response profiles for the sensor chip CM5 and PPF modules. The calibration curves demonstrate that the sensor chip PPF has somewhat better sensitivity than the sensor chip CM5 under the same experimental conditions; the linear regions of sensor chips CM5 and PPF are 0.1-50 and 0.01-50 mg/mL, respectively. Antibodies are likely to be immobilized onto the PPF surface at high density whereas immobilization onto the dextran matrix occurs at a lower density but throughout the three-dimensional structure (See Figure 1).34 For instance, it was reported that antibodies are densely adsorbed or immobilized onto PPF surfaces when these surfaces were used in alternative sensing schemes not involving SPR.24,25 Therefore, the PPF surface in our case is also likely to be populated by a continuous, uniformly accessible surface of antibody ligands. This suggests that antigen-antibody binding to the PPF surface can give a more efficient sensor response than the same reaction on a sensor chip CM5 partly because mass transport11 to the surface (32) Wertheimer, M. R.; Klemberg-Sapieha, J. E.; Schreiber, H. P. Thin Solid Films 1984, 115, 109-124. (33) Brosset, D.; Ai, B.; Segui, Y. Appl. Phys. Lett. 1978, 33, 87-89. (34) Surface concentration of antibodies immobilized to the matrix is an important value characterizing the sensor response.13 These values were investigated with a radiolabeled assay. The values for the sensor chip CM5 and PPF (allylamine monomer) were 1.3-25.3 13 and 26.4 pmol/cm2,24 respectively. Total amount of attached antibodies onto sensor chip CM5 and PPF are similar. However, the features of surface coverage are different; the antibodies immobilized onto sensor chip PPF are two-dimensional whereas those on the sensor chip CM5 are three-dimensional (see Figure 1).

of the sensor chip PPF occurs more rapidly than for sensor chip CM5. Similar results were reported by Nakanishi et al.20 and Lu et al.35 where so-called “orientation-controlled” immobilization generated a better response. The good sensitivity can also be attributed to the homogeneity of the PPF’s optical properties. Christensen and Fowers reported that roughness and inhomogeneity of the metal and sensing region forced the reflectivity curve to broaden and shift.36 For this reason, PPFs have been applied to optical waveguides37 and ultrathin coatings on the surface of contact lenses.38 Although we adopted here a 100 nm thick PPF film, an even thinner layer would further improve the sensitivity. However, it is also important to accurately control the thickness of the film to achieve reproducibility of the sensor response from sample to sample. For this reason, a compromise between sensitivity and reproducibility must be achieved. The response to 50 mg/L BSA was also investigated as a test for nonspecific adsorption. The RU change found for sensor chips CM5 and PPF were 1.6 and 6.6, respectively. Although the signal caused by nonspecific reaction in sensor chip PPF was 4 times larger than that of sensor chip CM5, it is not anticipated to be a problem since the absolute magnitude of the RU change corresponds to equivalent HSA concentrations 1000 times smaller (see Figure 2). Some other practical merits lie in sensor chip PPF. For instance, sensor chip PPF fabrication is carried out in a onestep process whereas sensor chip CM5 fabrication typically needs more steps. Furthermore, it appears possible to fabricate the new sensor chip starting from metal layer to PPF layer in situ (i.e., in a one-chamber fabrication). That should translate to much better uniformity among a series of chips fabricated in the same batch. CONCLUSION We demonstrate here the performance of sensor chip PPF for SPR sensing. We stress the merits of sensor chip PPF for monitoring very specific macromolecular interaction. Moreover, our group has shown39 that these merits increase in an analysis of small (low molecular weight) molecule-antibody interaction, e.g., pesticides-antibodies.40,41 More recently, a sensor chip containing a hydrophobic surface has been developed which can be used to attach lipid monolayers for an analysis of cell adhesion molecule interaction.18,42 In our case, a similar functionality can be realized by selection of a suitable monomer that can generate a hydrophobic surface. For example, a hexamethyldisiloxane PPF has a hydrophobic surface.43 Therefore, the sensor chip PPF can be used not only in antigen-antibody sensing but also in various kinds of other molecular interaction studies. (35) Lu, B.; Xie, J.; Lu, C.; Wu, C.; Wei, Y. Anal. Chem. 1995, 67, 83-87. (36) Christensen, D.; Fowers, D. Biosens. Bioelectron. 1996, 11, 677-684. (37) Tien, P. K.; Smolinsky, G.; Martin, R. J. Appl. Opt. 1972, 11, 637-640. (38) Ho, C. -P.; Yasuda, H. J. Biomed. Mater. Res. 1988, 22, 919-937. (39) Sasaki, S.; Kai, E.; Miyachi, H.; Muguruma, H.; Ikebukuro, K.; Ohkawa, H.; Karube, I. To be published in Anal. Chim. Acta. (40) Sasaki, S.; Tokitsu, Y.; Ikebukuro, K.; Yokoyama, K.; Masuda, Y.; Karube, I. Anal. Lett. 1997, 30, 429-443. (41) Minunni, M.; Mascini, M. Anal. Lett. 1993, 26, 1441-1460 and references therein. (42) van der Merwe, P. A.; Barclay, A. N. Curr. Opin. Immunol. 1996, 8, 257261. (43) Tajima, I.; Yamamoto, M. J. Polym. Sci., Polym. Chem. Ed. 1985, 23, 615622.

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ACKNOWLEDGMENT This work was supported by The Ministry of Education, Science, Sports and Culture: Large Scale Research Projects under the New Program in Grants-in-Aid for Scientific Research. H.M. thanks Research Fellowships of the Japan Society for the Promotion of Science for Young Scientists.

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Received for review June 2, 1997. Accepted September 12, 1997.X AC970571I

X

Abstract published in Advance ACS Abstracts, October 15, 1997.