A plasma-polymerized film for capacitance immunosensing

Biosensors and Bioelectronics 20 (2004) 841–847

A plasma-polymerized film for capacitance immunosensing Jishan Li, Hua Wang, Ting Deng, Zhaoyang Wu, Guoli Shen∗ , Ruqin Yu State Key Laboratory of Chemo/Biosensing and Chemometrics, College of Chemistry and Chemical Engineering, Hunan University, Changsha 410082, PR China Received 28 October 2003; received in revised form 19 March 2004; accepted 24 March 2004 Available online 24 April 2004

Abstract A capacitance immunosensor based on a plasma-polymerized ethylenediamine film (PPEF) has been developed. The resulting PPEF is studied with scanning electrode micrograph (SEM), IR reflection spectrum and cyclic voltammetry. SEM and IR reflection spectrum showed that the plasma-polymerized film (PPF) formed on the gold electrode surface is quite homogeneous, flat, nonporous and contains plenty of free-reacted –NH2 . Moreover, cyclic voltammetry showed that the hexacyanoferrate redox reactions were blocked well by the formed PPF, that is to say, the formed PPF has excellent insulating characteristics. To investigate its applicability for capacitive immunosensing, goat-anti-human IgG antibody (IgGAb) was coupled to the PPF-coated gold electrode surface via glutaraldehyde (GA) to form an immunoglobulin G (IgG) probe. Alternating current (ac) impedance and capacitance measurement were used in the immunoassay. The experiment results show that the PPEF is applicable to form insulating layer of capacitive immunosensors. © 2004 Elsevier B.V. All rights reserved. Keywords: Plasma-polymerized film; Capacitance immunosensor; Alternating current (ac) impedance

1. Introduction Capacitance measurement could be a useful tool in immunoassay (Christine et al., 2001). As for the construction of an immunosensor based on a capacitive transducer, the first and probably most important step is the immobilization of the recognition elements to provide the sites of antibody–antigen interaction. In order to circumvent the possible interference from electroactive species in the electrolyte solution and reduce the Faradic current, the layer containing the recognition sites should be electrically nonconducting and at the same time not too thick to guarantee a low detection limit. Some earlier works (Thust et al., 1999; Newman et al., 1986; Pierre et al., 1988; Andreas et al., 1992) focused on metal oxide and semiconductor thin layers. However, the infrastructure for production of the sensor chips is not accessible for most of the biosensor laboratories. In recent years, organosulfur-based species self-assembled at noble metal surfaces are quite popular in capacitance immunosensing ∗ Corresponding author. Tel.: +86-731-8821355; fax: +86-731-8822237. E-mail address: [email protected] (G. Shen).

0956-5663/$ – see front matter © 2004 Elsevier B.V. All rights reserved. doi:10.1016/j.bios.2004.03.023

technology (Berggren and Johansson, 1997; Vladimir et al., 1997; Christine et al., 1998; Dijksma et al., 2000, 2001; Bordia et al., 2002; Howard et al., 2003; Anderson et al., 1996). Long-chain alkanethiols containing a second functional group available for covalent attachment of intact immunoreagents have been used to form well-organized structure on gold substrates for the preparation of capacitive immunosensors (Berggren and Johansson, 1997; Anderson et al., 1996). However, the capacitive immunosensors using long-chain alkanethiols bear a problem that the self-assembled layer (SAL) might not be able to completely cover the gold electrode surface to guarantee the electrical isolation (Ameur et al., 2000). A remedy for this is to block any possible defects with alkanethiols, such as n-dodecanethiol. Such a treatment, however, might cause the loss of activity of the capacitance sensor (Berggren and Johansson, 1997). Plasma-polymer deposition techniques offer a promising alternative to the conventional coating. Since Yasuda and Lamaze (1973) firstly synthesized an organic thin film using plasma polymerization method, this technique has been used in reverse osmosis and membrane separation processes. The plasma-polymerization films (PPFs) with high cross-linking degree mainly possess following advantages

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(Nakanishi et al., 1996; Zhu and Liu, 1992): (i) the thickness of film is controllable, and a nanometer-leveled film can easily be obtained; (ii) they have a strong adhesion to the substrate; (iii) they are homogeneous and hole-free; (iv) they are mechanically and chemically stable; (v) many compounds can be used as precursors to deposit PPFs; (vi) PPFs can be deposited on various substrates, such as metals, glasses, Si plates, and organic materials. Due to their outstanding properties, some efforts (Kurosawa et al., 1997; Yan et al., 1998) have been made on depositing PPFs for chemically sensing application. Particularly, some recent reports (Nakamura et al., 1997; Wu et al., 2000; Muguruma and Karube, 1999) demonstrated the successful applications of PPF in biosensing. Such a polymer film seems very suitable for preparation of capacitive biosensors. In this paper, a new coating method for capacitive immunosensor preparation based on plasma-polymerized ethylenediamine film (PPEF) is reported. The resulting coating, which is extremely thin and homogeneous with a flat surface and highly cross-linked structure, shows good adhesion to the gold substrate. Additionally, there are plenty of amine groups on the surface of PPEF which are useful for glutaraldehyde cross-linking. Using goat-anti-human IgG antibody as the model substrate, a direct detection of IgG by capacitance measurements was demonstrated using a heterostructure of Au/PPEF/covalent-coupled antibody. When such a heterostructure is immersed in a solution containing the specific antigen, the interaction of antibody with the antigen leads to the thickening of the dielectric layer and induces a capacitance decrease which can be directly related to the amount of antigen to be quantified. 2. Materials and methods 2.1. Reagents and materials Goat-anti-human IgG antibody (IgGAb, affinity purification) and normal human reference serum (NHRS, containing 10.9 mg ml−1 of immunoglobulin G (IgG)) were purchased from Sino-America Biotechnology Company (Shanghai, China). Ethylenediamine (AR), glutaraldehyde (GA) (25% solution) and gold wire (1.0 mm diameter, 99.99%) were purchased from Changsha Chemical Reagents (Changsha, China). All other reagents and solvents were of analytical reagent grade and the doubly distilled water was used throughout. Phosphate buffer is a 8 mM Na2 HPO4 –2 mM NaH2 PO4 solution (PBS) of pH 7.4. The 1 mM hexacyanoferrate(III) or 1 mM hexacyanoferrate(II) solutions used in electrochemical system were prepared using PBS of pH 7.4. 2.2. Analytical system For alternating current (ac) impedance and capacitance measurements, a Model 283 electrochemistry system and

frequency response detector with Powersine220 software (EG&G Princeton Applied Research, Princeton, NJ, USA) were used. The volume of capacitance measuring cell is 5.0 ml and the position of working and reference electrode is settled (the distance from each other is 0.5 cm). The three-electrode system used in cyclic voltammetry experiments consisted of working electrode of interest, a saturated calomel reference electrode (SCE) and a platinum foil auxiliary electrode. The capacitance changes of the immunosensor could be obtained directly from the experiments. 2.3. Methods 2.3.1. Preparation of the sensor (a) Pretreatment of the gold electrode. Gold electrode was constructed by sealing a 1.0 mm diameter pure gold wire in glass tube with an exposed surface area of 0.00785 cm2 , and then polished with alumina slurries down to 0.05 ␮m. After mechanical polishing, the Au electrode was immersed in “piranha” solution containing reagent grade 30% H2 O2 and concentrated H2 SO4 (3/7 v/v) at 70 ◦ C for 10 min. Finally, the electrode was cleaned ultrasonically in distilled water and ethanol each for 15 min. Caution: “piranha” solutions react violently with many organic materials and should be handled with extreme care! (b) Manufacture of the PPEF. The PPEFs were prepared as follows: the gold electrode was placed in a stainless reactor with a Microwave Plasma Chemical Vapor Deposition System (MPCDS) (Guowei Hightech Ltd., Hubei, China). Ethylenediamine was used as a precursor, and pure hydrogen as a carrying gas. During the deposition process, the flow rate of carrying gas was kept at 50 ml min−1 , the pressure of the reactor chamber was 160 Pa, and the apparent rf power was 60 W. (c) IR reflection spectrum, scanning electron micrograph and thickness of the PPEF. IR reflection spectrum of PPEF was taken using a WQF-410 FT-IR spectrometer (Beijing Second optical instrument Factory). Scanning electron micrographs of the PPEF surfaces were produced using JSM-5600LV SEM (JEOL Ltd., Japan). In order to evaluate the thickness of the PPE film formed on the gold electrode surface, the PPEF was formed on the surface of quartz crystal microbalance (QCM) under the same conditions as described in part (b) above and the frequency shifts were measured. At the same time, the electric conductivity of the PPEF-coated gold electrode was evaluated using cyclic voltammograms versus SCE with the scan rate of 100 mV/s. 2.3.2. Immobilization of protein A covalent attachment of the antibodies to the surface of PPEF was obtained by GA as described elsewhere (Nakanishi et al., 1996). For coupling with GA, the PPEF-modified electrode was incubated for 60 min in GA (2.5%, w/w in water), rinsed three times with water. The gold electrode with coupled layer was immersed for

J. Li et al. / Biosensors and Bioelectronics 20 (2004) 841–847

60 min in 0.6 ␮g ml−1 antibody in phosphate buffer solution (PBS, 10 mM, pH 7.03) at room temperature (30 ◦ C). The antibody-embedded gold electrode formed was soaked in 5% (v/v) ethaneamine solution to react with any unreacted carbonyl groups on the surface, then washed with abundant amount of doubly distilled water. 2.3.3. Measurement procedure The immunosensor was constructed in the cell of 5.0 ml with the antibody-embedded gold electrode as the working electrode, and the SCE as the reference. Then 5.0 ml of PBS (pH 7.4) containing appropriate amount of IgG was added into the measuring cell following the detection of the capacitance changes.

3. Results and discussion 3.1. Characterization of the PPEF The sensitivity of the sensor increases with decreasing thickness of the insulating layer. However, a homogeneous, nonporous and insulating coating is very important for capacitive immunosensor. Fig. 1 shows a typical result obtained by the formation of PPEF on gold substrate. One can predict from Fig. 1 that an optimal thickness of the dielectric layer could be obtained when the formation time is between ca. 20 and 25 min. The thickness of the PPEF was evaluated by forming PPF on the surface of quartz crystal microbalance under the same formation conditions. In this work, PPEF corresponding to ca. 1100 Hz of frequency decrease was deposited on the crystal in 20 min. According to the Sauerbery’s (1959) equation, the thickness (t) (Å) of

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coating on the gold electrode surface or the crystal can be estimated from the frequency shift (F),
 F t = 0.27 (1) ρ where ρ is the density of the film. Here, the density of ethylenediamine is 0.898 g cm−3 , so the thickness of PPEF is estimated to be about 33 nm. To show the presence of amine groups in the PPEF, IR reflection spectrum of the film was recorded. The adsorption peak at 3100–3400 cm−1 indicates the stretch vibration of N–H bonding. The peaks at 1167 and 1075 cm−1 show the C–N stretch vibration. The adsorption peaks in the range of 1630–1543 cm−1 are the deformation vibrations of N–H bonding. Here the adsorption peak at 1630 cm−1 shows the presence of normal amine. The IR spectra indicate that amine groups of different forms are present in the PPEF. At the same time, Scanning electron micrograph of the PPEF surfaces shows that the PPF surface is quite homogeneous, flat and nonporous (Fig. 2a). For comparison, the scanning electron micrograph of the bare gold electrode surface was also taken (Fig. 2b). 3.2. Optimization of the excitation frequency Information about the optimal excitation frequency was gathered by the following experiment. Goat-anti-human IgG antibody was immobilized on the PPEF-coated electrode surface via GA. Frequencies from 1 to 100,000 Hz were applied for the detection of IgG (Fig. 3). For this analyte, an excitation frequency of 24.42 kHz resulted in a maximum change in capacitance; hence the monitoring was performed at this frequency (f = 24.42 kHz).

Fig. 1. Interdependence of capacitance, the frequency shift of QCM, and oxidative peak current in 1 mM hexacyanoferrate(II)/hexacyanoferrate(III) PBS of pH 7.4 during the PPEF forming. The capacitance measurement was at E = +0.0 V, Eac = 10 mV and f = 24.42 kHz, and the scan rate of cyclic voltammograms was 100 mV/s. All potentials were given vs. SCE.

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3.3. Properties of the grafted IgGAb layer

Fig. 2. Scanning electron micrographs of the bare gold electrode surface (b) and that coated PPEF (a).

Fig. 3. Dependence of capacitance change on the excitation frequency for the detection of IgG, measured in PBS (pH 7.4) at E = +0.0 V, Eac = 10 mV and f = 100 kHz–1 Hz. The IgG concentration is 43.6 ng ml−1 and the incubation time is 60 min.

Cyclic voltammetry of electroactive species in a conducting aqueous solution is a valuable mean of probing the insulation of a layer on gold (Porter et al., 1987; Chidsey and Loiacono, 1990). Because electron transfer between a chemical species in the solution and the electrode surface must occur either by tunneling through the layer or by approaching the electrode at a “pinhole” or defect in the layer, the extent of surface passivation to electron transfer is useful to detect defects in the layer (Finklea et al., 1993; Creager et al., 1992). The defect site in the layer can result in partially short-circuiting, which increases its capacitance. It is crucial to obtain a well-ordered surface structure of polymer without such defects. The degree of insulation can be examined using cyclic voltammetry with permeable redox couple, such as K3 [Fe(CN)6 ], in the solution. A layer of ethylenediamine plasma polymer with a capacitance of 2531.2 nF cm−2 seems less defective than other reported self-assembled monolayer (SAM) did. The permeability of ions through the layer is so small that the conducting ions in the solution can seldom penetrate it (Fig. 4b). Coupling with GA and covalent-coupling of IgGAb further reduce the penetration of the conducting ions in the solution (Fig. 4c) and electrically insulates the surface, showing a capacitance of 1351.3 nF cm−2 . Impurities in the gold surface could not be covered by the self-assembled monolayer of long-chain alkanethiols and thus might be accessible to ions from the solution. After the gold surface were covered thoroughly by PPFs, the layer of plasma polymer on the gold surface is the ethylenediamine derivative of a highly cross-linking polymer, and contains free unreacted –NH2

Fig. 4. Cyclic voltammograms recorded in 1 mM hexacyanoferrate(II)/hexacyanoferrate(III) in PBS of pH 7.4 using: (a) bare gold electrode, (b) gold electrode modified with PPEF and coupling with glutaraldehyde and (c) as (b) but with IgGAb immobilized. The scan rate was 100 mV/s. All potentials were given vs. SCE.

J. Li et al. / Biosensors and Bioelectronics 20 (2004) 841–847

Fig. 5. Capacitance change of the PPEF-modified sensor with different concentrations of IgGAb incubated via GA for 60 min, measured in PBS (pH 7.4) at E = +0.0 V, Eac = 10 mV and f = 24.42 kHz.

groups among the backbone, which can be applied for the covalent-coupling of GA. 3.4. Immobilizing IgGAb on gold The PPEF-modified gold electrode treated by glutaraldehyde was dipped into the PBS (pH 7.4) solution containing IgGAb. The capacitance measurement showed that the capacitance change increases with the concentration of IgGAb increasing and reaches saturation at 600 ng ml−1 (Fig. 5). After covalent-coupling IgGAb, ethaneamine was used as a sealing agent to react with any unreacted surface carbonyl groups. 3.5. Investigation of nonspecific binding To investigate nonspecific binding phenomena, ac impedance and capacitance measurements were used. The antibody(IgGAb)-embedded immunosensors were incubated in NHRS and BSA at the same concentration level, respectively. Both the real and the imaginary part of the impedance increase when the sensor was incubated with NHRS, which is assumed to result from an increase in the mean thickness of the layer, caused by the specific interaction of antigen with antibody on the electrode surface to form an additional layer on the surface. Moreover, a capacitance shift of ca. 340 nF cm−2 (f = 24.42 kHz) was recorded as compared to no significant shift for the case of incubating in BSA. This suggests that the observed capacitance change was a specific one and not being caused by nonspecific absorption of protein at the sensor surface. The physical basis for the capacitance response is considered to be associated with the displacement of the polar water molecules close to the electrode surface with much less polar species. Nonspecific interaction seems not be able to drive the polar water molecules far away from the electrode

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Fig. 6. Interaction between IgGAb and IgG as monitored by capacitance with time. The antibody-embedded immunosensor dipped into a PBS (pH 7.4) with IgG concentration of 43.6 ng ml−1 , measured at E = +0.0 V, Eac = 10 mV and f = 24.42 kHz. The standard deviations obtained by five repeated measurements were shown as the error bars.

surface. The capacitance immunosensor seemed to be selective enough without much interference from nonspecific absorption. 3.6. Detection of human IgG The interaction of IgG with IgGAb can be directly monitored in the measuring cell by determining capacitance versus time. Fig. 6 shows an example of such a measurement. The saturation was reached after an incubation time of 40 min. Samples of 2.18–327.0 ng ml−1 IgG were added to the measuring cell separately and the capacitance changes were measured. A nearly linear relationship between the capacitance change and the logarithm of the IgG concentration was obtained in the concentration range of 2.18–109.0 ng ml−1 and reached a saturation at ca. 174.4 ng ml−1 (Fig. 7). The linear regression equation was C (nF cm−2 ) = −142.45 log[T ] − 104.24, with a correlation coefficient of 0.9987. The detection limit was ca. 0.19 ng ml−1 as calculated from the smallest measurable value of the capacitance variation under optimum experimental conditions, which is about 210 pF cm−2 . 3.7. Activity restoration after washing by acidic solution To restore the immunosensor activity, the IgG and IgGAb conjugation on the sensor surface was washed with an acidic solution (glycine-HCl, pH 2.1). The analyzed immunoelectrode was immersed in the stirred acid solution for 15 min and then IgG was released from the sensor surface. The sensor was then put in the detection cell for capacitance measurement with PBS (pH 7.4) to check whether the

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tance change for the same amount of IgG, and no noticeable hysteresis was observed. 3.8. Analysis of human serum samples Finally, in order to investigate the possibility of using the prepared capacitive immunosensor for clinical analysis, some patient human serum samples were examined by the developed capacitive immunosensor and the results obtained were compared with that by enzyme-linked immunosorbent assay (ELISA). The comparison of results was shown in Table 2. The analytical results show that the developed capacitive immunosensor based on PPEF is suitable for the applications in clinical area.

Fig. 7. Capacitance changes vs. the logarithm of the IgG concentration. For each experiment, the incubation time was 50 min, and measured at E = 0.0 V, Eac = 10 mV and f = 24.42 kHz. Table 1 Regeneration of the prepared capacitive immunosensor 1 Ci Ca C

1347.6 1004.7 −342.9

2

3

1341.8 1003.3 −338.5

1343.5 1012.4 −331.1

4 1343.3 1006.3 −337.4

5

6

1335.2 1008.6 −326.6

1332.7 1015.9 −316.7

Ci is the initial capacitance of the immunosensor regenerated by using glycine-HCl solution, Ca the capacitance after the binding of antigen, C the capacitance change. For each experiment, the antibody(IgGAb)-embedded immunosensor was incubated in 43.6 ng ml−1 of IgG solution for 40 min.

blank capacitance of the sensor was restored. The IgG-free sensor was then ready for the next IgG assay. Six restoring experiments were shown in Table 1. As can be seen from the result, that the sensor could provide nearly the same capaci-

4. Conclusions The plasma-polymerized ethylenediamine film has been proved to be capable to completely cover the surface of gold electrode. The feasibility of capacitance measurements of antibody–antigen interaction based on the polymer prepared has been studied. The antibodies covalently coupled with the polymer layer showed normal behavior without obvious loss of activity. The nonspecific adsorption of protein species does not show much interference. The restoration of the activity of immunosensor and its reuse are also possible. Moreover, the system should also be useful for the investigation of other ligand–analyte systems (lectins–carbohydrates, receptors–peptides) or DNA–DNA hybridization studies. There are, however, some points which have to be improved. The thickness of PPEFs is limited. When the PPEF is too thin, the faradaic current of the sensor will increases due to some defects in the PPEFs. To improve the sensitivity of the capacitive biosensors based on PPFs, the super-thin and nonporous insulating layer must be obtained. A work in this direction is already in progress.

Table 2 Analysis results of IgG for human serum specimen Samples

1 2 3 4 5 6 7 a

IgG concentration by ELISAa (g/l)

IgG concentration by CIAb (g/l)

6.34 8.90 11.10 12.10 12.30 13.00 13.70

6.65 9.34 11.54 11.44 12.68 12.52 13.36

± ± ± ± ± ± ±

0.32 0.46 0.66 0.64 0.58 0.52 0.72

Enzyme-linked immunosorbent assay method. Capacitance immunosensing assay method, measured at f = 24.42 kHz. As the IgG concentration level of human serum is very high, the human serum specimens have to be diluted to the detection range of the capacitive immunosensors by using PBS according to the IgG concentration by ELISA. The measured value obtained by development capacitive immunosensor multiplied by the dilution ratio was used as the IgG concentration of human serum specimen. The standard deviations were calculated by five repeated measurements.

Acknowledgements We thank Prof. J.-F. Tian for carrying out the SEM and H. Cui for IR measurements. We also wish to acknowledge X. Zeng from the second hospital of Central South University (Changsha, China) who offered us human serum specimen. This work was financially supported by the National Natural Science Foundation of China (grant nos. 20075006, 20205004, 20205005), the Foundation of Ministry of Education for Doctor (no. 20010532008).

b

References Ameur, S., Martelet, C., Jaffrezic-Renault, N., Chovelon, J.M., 2000. Sensitive immunodetection through impedance measurements onto gold functionalized electrodes. Appl. Biochem. Biotechnol. 89 (2–3), 161–170.

J. Li et al. / Biosensors and Bioelectronics 20 (2004) 841–847 Anderson, J.L., Bowden, E.F., Pickup, P.G., 1996. Dynamic electrochemistry: methodology and application. Anal. Chem. 68 (12), 397–444. Andreas, G., Manuel, A.I., Walter, S., Rolf, D.S., 1992. Real-time monitoring of immunochemical interactions with a tantalum capacitance flow-through cell. Anal. Chem. 64, 997–1003. Berggren, C., Johansson, G., 1997. Capacitance measurements of antibody–antigen interactions in a flow system. Anal. Chem. 69, 3651– 3657. Bordia, F., Cametti, C., Gliozzic, A., 2002. Impedance measurements of self-assembled lipid bilayer membranes on the tan electrode. Bioelectrochemistry 57 (1), 39–46. Chidsey, C.E.D., Loiacono, D.N., 1990. Chemical functionality in self-assembled monolayers: structural and electrochemical properties. Langmuir 6, 682–691. Christine, B., Bjarni, B., Gillis, J., 2001. Capacitive biosensors. Electroanalysis 13 (3), 173–180. Christine, B., Bjarni, B., Gillis, J., 1998. An immunological interleukine-6 capacitive biosensor using perturbation wit potentiostatic step. Biosens. Bioelectron. 13 (10), 1061–1068. Creager, S.E.J., Hockett, L.A., Rowe, G.K., 1992. Consequences of microscopic surface roughness for molecular self-assembly. Langmuir 8, 854–861. Dijksma, M., Kamp, B., Hoogvliet, J.C., Bennekom, W.P., 2000. Formation and electrochemical characterization of self-assembled monolayers of thioctic acid on polycrystalline gold electrodes in phosphate buffer pH 7.4. Langmuir 16, 3852–3857. Dijksma, M., Kamp, B., Hoogvliet, J.C., Bennekom, W.P., 2001. Development of an electrochemical immunosensor for direct detection of interferon-␥ at the attomolar level. Anal. Chem. 73, 901–907. Finklea, H.Q., Snider, D.A., Fedyk, J., Sabatani, E., Gafni, Y., Rubinstein, I., 1993. Characterization of octadecanethiol-coated gold electrodes as microarray electrodes by cyclic voltammetry and ac impedance spectroscopy. Langmuir 9, 3660–3667. Howard, A.C., Daniel, H.Z., Marc, O., 2003. In vivo CH3 (CH2 )11 SAu SAM electrodes in the beating heart: in situ analytic studies relevant to pacemakers and interstitial biosensors. Biosens. Bioelectron. 18 (1), 11–21. Kurosawa, S., Tawara-Kondo, E., Minoura, N., Kamo, N., 1997. Detection of polycyclic compounds as mutagens using piezoelectric quartz crystal coated with plasma-polymerized phthalocyanine derivatives. Sens. Actuators B 43 (1–3), 175–179.

847

Muguruma, H., Karube, I., 1999. Plasma-polymerized films for biosensors. Trends Anal. Chem. 18 (1), 62–68. Nakamura, R., Muguruma, H., Ikebukuro, K., Sasaki, S., Nagata, R., Karube, I., Pedersen, H., 1997. A plasma-polymerized film for surface plasmon resonance immunosensing. Anal. Chem. 69, 4649– 4652. Nakanishi, K., Muguruma, H., Karube, I., 1996. A novel method of immobilizing antibodies on quartz crystal microbalance using plasma-polymerized films for immunosensors. Anal. Chem. 68, 1695– 1700. Newman, A.L., Hunter, K.W., Stanbro, W.D., 1986. Proceedings of the International Meeting on Chemical Sensors, 2nd ed., pp. 596–598. Pierre, B., Gardies, F., Jaffrezic-Renault, N., Martelet, C., 1988. Detection of immunospecies by capacitance measurements. Anal. Chem. 60 (21), 2374–2379. Porter, M.D., Bright, T.B., Allara, D.L., Chidsey, C.E.D., 1987. Spontaneously organized molecular assemblies. 4. Structural characterization of n-alkylthiol monolayers on gold by optical ellipsometry, infrared spectroscopy, and electrochemistry. J. Am. Chem. Soc. 109, 3559– 3568. Sauerbery, G., 1959. Use of a quartz vibrator for weighing thin layers on a microbalance. Z. Phys. 155, 206–222. Thust, M., Schöning, M.J., Schroth, P., Malkoc, Ü., Dicker, C.I., Steffen, A., Kordos, P., Lüth, H., 1999. Enzyme immobilization on planar and porous silicon substrates for biosensor applications. J. Mol. Catal. B: Enzym. 7 (1–4), 77–83. Vladimir, M.M., Michael, R., Otto, S.W., 1997. Capacitive monitoring of protein immobilization and antigen–antibody reaction monomolecular alkylthiol films on gold electrodes. Biosens. Bioelectron. 12 (9–10), 977–988. Wu, Z.Y., Yan, Y.H., Shen, G.L., Yu, R.Q., 2000. A novel approach of antibody immobilization based on n-butyl amine plasma-polymerized films for immunosensors. Anal. Chim. Acta 412 (1), 29–35. Yan, Y., Zeng, Y., Xiang, J., Yin, X., Jin, J., Che, Z., 1998. A novel diamond-like structural film with characterization of recognizing for formic acid vapor. Chin. Sci. Bull. 43 (15), 1672–1675. Yasuda, H., Lamaze, C.E., 1973. Preparation of reverse osmosis membranes by plasma polymerization of organic compounds. J. Appl. Polym. Sci. 17, 201–222. Zhu, C., Liu, M., 1992. Membrane Science and Technology. Zhejiang University Press, Hangzhou, p. 19.