Anal. Chem. 1994, 66, 2888-2892
Acoustic Plate Mode Sensor for Immunochemical Reactions R. Dahlnt,t M. Grunze,*pt F. Jesse,'** and J. Renkent Lehrstuhl fur Ange wandte Physikalische Chemie, Universitat Heidelberg, Im Neuenheimer Feld 253, 69 120 Heidelberg, Germany, and Department of Electrical and Computer Engineering, Marquette University, Milwaukee, Wisconsin 53233
A practical immunosensor, based on a 150-MHz acoustic plate mode (APM) device on ZX-LiNbO3 has been built and is employed to monitor directly, in real time, antigen-antibody reactions via mass loading of the crystal surface. The results show frequencyshifts 1 kHz above the noise level upon antigen adsorption. The mass density detection limit of our present sensor is -200 pg/mm2, corresponding to 8.108 molecules/ mm2 utilizing a typical IgG of 150 000 Da. Antigen binding capacitiesof aminosilane films are measured independently by a modified ELISA test and are related to the APM sensor response.
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In the areas of medical diagnostics and biological research, there is a strong interest in direct and rapid diagnostic methods which are capable of separating specific molecules from a mixture. One of the most important techniques in the field of immunochemistry is the detection of antigens by the use of antibodies or vice versa. Recent developments in sensor technology focus on rapid on-line monitoring of antigenantibody reactions. Biosensors based on the surface plasmon resonance (SPR) have been developed1 utilizing changes in the refractive index at the sensor surface due to changes in adsorbed mass. However, a relatively expensive optical analytical system is required. In contrast, acoustic wave sensors offer the opportunity to establish a low-cost system for a one-step detection of immunochemical reactions. Moreover, only a small amount of liquid is required for the analysis.24 The most common acoustic wave sensor configurations utilize bulk acoustic wave resonators and surface acoustic wave (SAW) devices. In both configurations, the liquid analyte is in contact with (or near) both the crystal surface and the electrodes. Corrosion problems on the electrodes may occur, resulting in instabilities and deterioration of the sensor response. Acoustic plate modes (APMs) excited by interdigital transducers (IDTs) present no such problems since APM devices are mounted such that the biological system is in contact with the nonelectroded crystal surface. There exist various interaction mechanisms between the acoustic fields and the medium being analyzed that result in the sensor response^.^-^ One of the primary interaction
mechanisms in acoustic wave sensors is mass loading caused by the amount of particles bound to the device surface. The added mass, typically a few nanograms, results in changes of the properties of the acoustic wave. In a previous study,9 the mass sensitivity of APMs on ZXLiNb03 was investigated as a function of frequency. The purpose of the study was to determine whether APM devices on ZX-LiNb03 can be used as mass-sensitive practical immunosensors. The mass loading study was conducted in liquid environments to accurately reflect the sensitivity of an actual biosensor. It was found that the frequency mass sensitivity df/dms = -0.0366$, where df (Hz) denotes the change in the device operating frequency,f (MHz), due to a change in surface mass density, dm, (ng/mm2). Figure 1 shows the frequency-mass sensitivity characteristics of APMs on ZX-LiNb039 and other acoustic waves on quartz. It is seen that the sensitivity of APMs on ZX-LiNbO3 is less than that for SAWs. However, the shear horizontally (SH) polarized APMs on ZX-LiNb03 appear more suitable for biosensing applications. This is because the biological system is in contact with the nonelectroded surface of the APM device and SAWs are characterized by a large propagation loss in liquid environments due to mode conversion caused by the SAW compressional component. Quartz bulk acoustic wave resonators, which are the most mass-sensitive detectors, are limited by the highest practical operating frequency (-30 MHz) that can be achieved. It can then be seen that the S H polarized APMs on ZX-LiNb03 appear to be a good trade-off with a relatively high mass sensitivity. An estimation of the area required by one antibody molecule varies between 30 and 100 nm2. Thus, the mass density of a close-packed monolayer of IgG-used as the antigen component-ranges from 2.5 to 8.3 ng/mm2. Because this added mass is small, it may not be possible to develop a practical high-sensitivity APM biosensor. However, the above frequency dependence of the sensitivity indicates that mass sensitivity can be enhanced by using APM devices with higher frequencies of operation. In the present study, a practical mass-sensitive APM biosensor is demonstrated. A 150-MHz dual delay line APM biosensor on ZX-LiNbO3 was designed and fabricated, and ( 5 ) Martin, S. J.; Ricco, A. J.; Niemczyk, T. M.; Frye, G. C. Sens. Acruators
Universitat Heidelberg. 1 Marquette University. (1) Malmqvist, M. Nature 1993, 361, 186. (2) Thompson, M.; Dhaliwal, G. K.; Arthur, C. L.; Calebrese, G. S. I€€€ Trans. Ultrason.. Ferroelec.. Freq. Conrr. 1987, UFFC-34,127. (3) Andle, J. C.; Vetelino, J. F.; Lec, R.; McAllister, D. Proc. I € € E Ultrason. Symp. (IEEE, New York) 1989, 579. (4) Gizeli, E.; Stevenson, A. C.; Goddard, N. J.; Lowe, C. R. Transducers ’91 (I€€€, New York) 1991, 690.
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1989, 20, 253. (6) Martin, B. A,; Wenzel, S. W.; White, R. M. Sens. Actuators 1990, A21-A23, 704. - .. (7) Andle, J. C.; Vetelino, J . F.; Lade, M.; McAllister, D. Proc. IEEE Ultrason. Symp. ( I € € € , New York) 1990, 291. (8) Josse, F.;Shana, Z. A,; Haworth, D. T.; Liew,S.;Grunze, M.Sens. Actuators B 1992, 9, 97. (9) Bender, F.; Dahint, R.; Josse, F.; Grunze. M.;v. Schickfus, M. J. Acoust. Soc. Am. 1994, 95 (3), 1386.
0003-2700/94/0366-2888$04.50/0
0 1994 Amerlcan Chemical Society
Output IDT
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Figure 1. Frequency-mass sensitivity characterlstlcs of dlfferent acoustlc wave sensors. The APM (ZX-LiNb03) curve is obtalned by fling measured data. The curves representingthe quartz bulk acoustlc wave resonators and SAWS are theoretlcal curves shown here for comparison.
the ability of the device to perform as a mass-sensitivebiosensor is tested by monitoring antigen-antibody reactions at the crystal surface. Antibodies are covalently linked to the sensor surface by the use of aminosilane films.1° Antigen binding capacities of the films are independently determined by standard immunological tests and related to the frequency response of the APM sensor.
EXPERIMENTAL SECTION Immobilization of Antibodies. Before the actual (nonelectroded) surface derivatization with aminosilane, the crystals were cleaned by rinsing with 2-propanol, followed by a treatment with a 1:1 mixture of methanol and concentrated hydrochloric acid for 30 min and subsequently with concentrated sulfuric acid for another 30 min. After neutralization, the surface was activated with 5 N NaOH for 10 min in order to hydroxylate the surface. The well-dried sensor was then treated with a solution of 2% (aminopropy1)trimethoxysilane in dry ethanol. After 20 min the solution was removed and the crystal carefully rinsed with dry ethanol. Before adding a solution of 2.5% glutaraldehyde in phosphate buffer saline (PBS) for 30 min, the film was allowed to cure at 105 "C for 1h. After carefully washing the crystal with PBS, the sensor surface was incubated with 25 pg of AC-purified goat anti-rabbit IgG in 1 mL of PBS. Excess aldehyde functions were deactivated with 50 pL of ethanolamine in 1 mL of PBS for 15 min. The last step consisted of washing the substrate carefully with PBS-Tween20. All immunochemicals used in the experiments were purchased from Sigma Chemie, Deisenhofen, Germany. Throughout all experiments, PBS was used at pH 7.4. A solution of PBS containing 0.1% of Tween-20 (Merck, Darmstadt, Germany) is denoted as PBS-Tween-20. The abbreviations AC and IEC refer to affinity chromatography
and ion exchange chromatography, respectively. All ACpurified antibodies are bought from Dianova, Hamburg, Germany. ELISA Measurements. The binding capacities for antigen were determined by a modified enzyme-linked immunosorbent assay (ELISA) on glass substrates prepared by following the procedures used for the LiNb03 sensor element. The glass substrates were incubated with 5 pL of rabbit antihorse IgG peroxidase (POD) conjugate in 1 mL of PBS for 1 h. After the substrates were washed with PBS-Tween-20, the amount of adsorbed peroxidase conjugate was measured photometrically using a 2,2'-azinobis(3-ethylbenzthiazoline-6-sulfonic acid) (ABTS) as enzyme substrate and H202 as oxidizing agent. The aminosilane films prepared with IEC-purified goat anti-rabbit IgG show a binding capacity of 0.18 f 0.03 ng/ mm2 for rabbit anti-horse IgG POD conjugate, whereas ACpurified antibodies show a higher binding capacity of 0.82 f 0.1 1 ng/mm2. This is expected due to the elimination of nonanti-rabbit IgG from the IgG fraction of goat serum by affinity chromatography. A nonspecific adsorption (goat anti-guinea pig IgG POD conjugate) of 15% (-0.12 ng/mm2) was observed. Antigen Detection Using APMs. The sensor experiments on the LiNbO3 crystal surface were performed using labelfree or gold-labeled goat anti-rabbit IgG as antigen component instead of the POD conjugate which was necessary for the ELISA. Thus, the actual values of the binding capacity for the sensor experiment must be corrected by the ratio of molecular weight of the actual antigen and the POD conjugate. As rabbit anti-goat IgG was coupled to the sensor surface, goat anti-rabbit IgG may also react via its antigen binding sites with the immobilized antibodies." ELISA tests show that antigen binding capacity is enhanced by a factor of -2.5 due to the additional binding mechanism. In the dual delay line configuration as shown in Figure 2, one line is used as a reference line to compensate for temperature effects, changes in the liquid properties, and nonspecific adsorption. Therefore, it has to be guaranteed that specific binding of antigens occurs only on one line, the sensing line. Two different techniques were used to achieve this situation. The first technique utilizes the fact that the
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(IO) Jdnsson, U.; Malmquist, M.; Olofsson, G.; Rdnnberg, I. Methods Enzymol. 1988, 137, 381.
(1 1) Davey, B. Immunologic; Birkhauser: Berlin, 1991.
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binding sites of antibodies can be deactivated by UV light. Thus, the reference line is exposed to 0.254-pm UV radiation for several hours. The deactivation process is performed under PBS. In order to avoid a degeneration of the antibodies on the sensing line due to heat diffusion, the deactivation process was conducted in a temperature chamber at 10 “C. The second technique that was also used consists of designing a compartimental liquid cell such that only the sensing line is exposed to bind antibodies during the immobilization process. Remaining aldehyde groups on both lines are deactivated by treatment with 50 pL of ethanolamine in 1 mL of PBS. The liquid cell is next filled with 1.6 mL of PBS, and the sensor is inserted into a temperature chamber to provide a constant temperature of -27 OC for device operation. The APM devices used in the experiments were designed to operate at a fundamental frequency of -51 MHz on a 0.5-mm-thick crystal plate of ZX-LiNbO3 with h/X = 5.7 (where h/X is the ratio of plate thickness h to the interdigital transducer period A). The design and characteristics of this type of device have been detailed elsewhere.* The optically polished LiNbO3 wafers were purchased from Crystal Technology Inc. As indicated earlier, a dual delay line configuration was used with each IDT utilizing a double-electrode pattern with four electrodes per wavelength. This electrode pattern is known to operate efficiently at higher harmonics if the electrode metalization ratio is properly selected.* To achieve a high sensitivity, the device was operated at a third harmonic corresponding to an excitation frequency of 150 MHz. It should be noted that a similar device utilizing the same acoustic mode and operating at an equivalent fundamental frequency would require an IDT with period X = 30 pm on a plate -0.17 mm thick. While such IDTs could be easily fabricated, handling, preparation for biosensing, and encapsulation of such devices would have been a problem. For all experiments, the derivatized APM devices were mounted with the IDTs on the bottom surface to allow the nonelectroded surface of the crystal to be exposed to the biochemical fluid. A liquid cell was attached to the top surface of the crystal to provide a well-defined interaction region between the acoustic fields and the biological film-fluid composite. A rubber gasket was used to provide a liquidtight seal. The measurements consist of monitoring the phase (frequency) shifts and signal amplitude (attenuation) of the output of both lines relative to the input as the antigen-antibody complexes are formed, i.e., as mass loading occurs. A Network analyzer (Hewlett-Packard 8752A) was used initially to electrically characterize the APM devices with and without liquid loading and to determine their suitability as sensors. A constant-amplitude fixed-frequency output signal of a frequency synthesizer (Rhode & Schwarz) was split into two equal signals. One signal is used as the reference and the other is fed through the delay lines. A vector voltmeter (HP 8508A) was used to measure the output of the APM device relative to the reference signal. Two miniature relay switches allow switching between the active line and the nonactive line. Thus, both the individual responses and the difference response between the two lines can be monitored. In what follows, the term “difference signal” will denote the sensor response as the response of the sensing line minus the response of the reference
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line. Using the measured phase-frequency characteristics of the device, the experimentally determined phase change can be related to the corresponding change in the device frequency of operation. Note that theentire measurement was automated using a personal computer.
RESULTS AND DISCUSSION Figure 3 shows the sensor response due to specific antigen binding at the sensor surface. AC-purified antibodies (rabbit anti-goat IgG) are immobilized at the surface of the sensing line using an aminosilane film. No antibodies are linked to the reference line. Stability of the system was monitored for 1 h, after which 48 pg of IEC-purified antigen (goat antirabbit IgG) dissolved in 20 pL of PBS was injected into the corner of the liquid cell and no stirring was done. As shown in Figure 3, the difference signal was observed to decrease by 1.4 kHz. The injection (same amount) was repeated 1.5 and 2.5 h later at the same location. A small change of 100 Hz was observed for the second injection and none for the third. This is to be expected since independent ELISA tests show that an antigen concentration of 30 pg/mL is sufficient to saturate almost all binding sites of the immobilized antibodies. The detection of gold-labeled antigens (AC-purified goat anti-rabbit IgG-gold conjugate) is demonstrated in Figure 4. AC-purified antibodies (rabbit anti-goat IgG) were covalently linked to the sensor crystal surface using aminosilane. Afterward, the reference line was deactivated by UV radiation. 1.6 mL of PBS was added to the liquid cell, and stability was observed. Note from the device response that stability was already reached after -20 min, but the device response was monitored up to 2 h to study the stability of the silanePBS solution system. After 2 h, 1.1 pg of gold-labeled antigen (-24 ng of antigen attached to 1080 ng of gold label) dissolved in 10 pL of a tris buffer solution was added to the liquid cell and the difference signal was monitored. As shown in Figure 4, a sensor response of 0.76 kHz was observed. The injection (same amount) was repeated 2 and 4 h later at the same location. A change of -0.27 kHz was measured
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for the second injection and none for the third. Similar behavior has been observed for the experiment with the detection of nonlabeled antigens. After the second injection, almost all binding sites of the immobilized antibodies are occupied and a large amount of surplus antigen exists in the solution. This statement is confirmed by a simple experiment in which 26 p g of AC-purified antibodies (rabbit anti-goat IgG) dissolved in 10 pL of PBS was added to the system at the same location after the eighth h. This type of antibody is capable of reacting with goat anti-rabbit IgG gold conjugate both in the solution and at the sensor surface. As the amount of added antibody is large compared to the amount of surplus antigen in the solution, only small antigen-antibody complexes will be formed.” These complexes are also able to bind to the crystal surface via specific and nonspecific adsorption. In either case, a sensor response is to be expected due to added masson both the sensing line and the reference line. As shown in Figure 4,the difference signal indicates a frequency decrease of 0.92 kHz at the injection of rabbit anti-goat IgG. Using the observed frequency change together with the frequencymass sensitivity characteristics of APMs on ZX-LiNbO3, it is concluded that 1.1 ng/mm2 antibody and antigen-antibody complexes are added to the sensor surface due to specific adsorption. In order to quantify the results shown above, the added mass resulting from specific and nonspecific adsorption of antigens was estimated from the ELISA data on the glass substrates as described above. The corresponding predicted frequency changes as well as the measured frequency change are listed in Table 1. The predicted response was calculated using the frequency-mass sensitivity characteristics of APMs on ZX-LiNbO3 according to Figure 1. The actual measured sensor response that accounts for the reference line response is also shown, i.e., nonspecific adsorption of antigens as well as temperature variation effects and other secondary effects. These results indicate a measurable frequency shift of more than 1 kHz, well above the noise level. Experiments with all
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a Specific and nonspecific adsorption are taken into account. The factor of 42 in the table accounts for the wei ht of the gold labels. The surfacemassdensityisestimatedfromtheE~~SAdataonglasssubtrates. c The sensing line responaeincludes both specific and nonspecificadsorption of antigens. Measured data have been temperature corrected. The predicted response is obtained from the frquenc -mass sensitivity characteristics of APMs on ZX-LiNbO,. Actual tiosensor response representing only the frquency change from specific adsorption of antigens.
devices show long-term frequency variations less than f65 Hz if the films are stable and temperature is properly controlled. Comparing the sensor response and the sensing line response in Table 1 for the case of aminosilane films, it is seen that nonspecific adsorption response is 22%and 16%(of the total adsorptionvalue) for the nonlabeled and gold-labeled antigens, respectively. This is about what is expected from the ELISA experiments. It is also observed that for all experiments a discrepancy exists between measured and predicted results. This may be due to calibration problems in the ELISA test and the fact that viscoelastic properties of the film-fluid composite have not been taken into account. Calibration problems may cause systematic errors in the evaluation of the antigen binding capacity, Le., the added mass. Second, the predicted sensor response assumes that only mass loading causes the response. However, viscoelastic effects known to oppose mass loading effect may be important,12 as suggested by the results for the detection of gold-labeled antigens in which most (98%) of the added mass is due to the gold labels and the measured response is more than 1 order of magnitude lower than predicted. Finally, it should be pointed out that the relatively long experiment time was due to both the slow diffusion process and the long observation time, which is not needed in actual sensor operation. The antigens were injected from the corner of the liquid cell and no stirring was done to avoid disturbing the response. A direct injection across the lines in combination with a careful stirring will drastically reduce the diffusion time, hence the reaction time.
CONCLUSION Due to the mass sensitivity of acoustic plate modes on ZXLiNbO3 having an? frequency dependence, an immunosensor with high sensitivity can be designed by using APM devices with higher frequencies of operation. In the present work, the APM biosensor was used to monitor antigen-antibody reactions via mass loading. The results, which show a frequency shift well above the noise level, indicate that practical and inexpensive immunosensors can be designed for real-time detection of antigen-antibody reactions. Present devices in our laboratory can monitor mass density changes down to 200 pg/mm2. Taking into account an active sensor area of 45 (12) Bartley, D. L.:Domingucz, D.D.Awl. Chrm. 1990,62, 1649.
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mm2, a detection limit of 9 ng is obtained, corresponding to -4 X 1O'O antibody molecules of a typical IgG of 150 000 Da . However, several problems remain to be solved before the practical implementation of the devices. These problems range from the reproducibility of film quality and thus antigen binding capacities to accounting for viscoelastic effectswhich tend to oppose the sensor response due to mass loading. This is especially important at high frequency, where viscoelasticity
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cannot be ignored. Thus, a study of these effects needs to be conducted in order to completely quantify these sensors.
ACKNOWLEDGMENT This workwas funded in part by the Max Buchner Stiftung. Received for review February 24, 1994. 1994."
Accepted M Y 17,
*Abstract published in Aduance ACS Abstracts, July 1, 1994.
