Detection of non-specific protein adsorption at artificial surfaces by the use of acoustic plate mode sensors

ELSEVIER

B Sensors and Actuators B 35-36 (1996) 497-505

Detection of non-specific protein adsorption at artificial surfaces by the use of acoustic plate mode sensors R. Dahint ~,*, R. Ros SeigeP, P. Harder a, M. Grunze -~, F. Josse b aAngewandte Physikalische Chemie, UniversitiltHeidelberg, lm Neuenheimer Feld 253, 69120 Heidelberg, Germany bDepartment of Electrical and Computer Engineering and Microsensor Research Laboratory, Marquette University, Milwaukee, W153233, USA

Abstract

The interaction of proteins with artificial surfaces is important to many medical and biochemical applications. Su h examples involve the incorporation of catheters and prostheses as well as the non-specific adsorption of pharmacological proteins at the walls of a container, which may drastically reduce their activity. A fast analytical tool capable of determining the specific adsorption characteristics at these surfaces would, therefore, support technological progress. Contrary to traditional immunoassays, acoustic wave-based sensors allow an on-line and direct detection of label-free proteins, thus saving time and providing the opportunity to monitor the kinetics of the binding process. In this study, Cr/Au-coated acoustic plate mode (APM) sensors have been used to investigate the interaction of immunoglobulin (3 (lgO) and fibrinogen with differently terminated self-assembled monolayers (SAMs) of thiols. By this method, both the low affinity of hexa(ethylene glycol)-terminated (HS-(CH2) II-(O-CH2-CH2)6-OH) alkanethiol SAMs and the high affinity of methyl-terminated (HS-(CH2)IrCH3) surfaces towards protein adsorption were confirmed. It was found that the amount of bound proteins depends on the pH of the solution. At low pH values, protein binding to methyl-terminated surfaces is drastically reduced. The adsorption characteristics of fibrinogen at methyl-terminated surfaces are explained by a kinetic model which involves the initial binding of native proteins and a subsequent unfolding process. Complete regeneration of the sensor element is achieved by the use of sodium dodecylsulfate. Keywords: Protein adsorption: Artificial surfaces: Acoustic plate mode sensors

1. Introduction

A detailed understanding of the interaction of proteins with artificial surfaces is critical for many medical and biochemical applications [1,2]. Such knowledge could, for example, improve the implementation of synthetic coatings for catheters and prostheses which come in contact with biological fluids. Other examples involve the non-specific adsorption of pharmacological proteins at the walls of a container, which may reduce their activity [3], and the adhesion of pathogenic bacteria, which is determined by the affinity of the surface towards proteins [4]. What is ueeded to support technical progress is a rapid, on-line analytical tool capable of detecting proteins and determining the specific adsorption behavior at various surfaces. Moreover, well-defined surface characteristics

* Corresponding author. 0925-4005/96/$15.00 © 1996 ElsevierScience S.A. All rights reserved Pll $0925-4005(96)02072-2

are required in order to create general model systems for studying the effects of surface properties on protein adsorption. Several approaches have been investigated and used for real-time, on-line monitoring of biospecific interactions. They include optical, electrical, electrochemical and acoustic techniques. Among the various technologies being investigated, acoustic wave-based sensors apparently offer the opportunity to establish a low-cost system for a one-step detection of proteins [5]. In particular, the use of acoustic plate modes (APMs) has proven to be a promising concept as sensor electrodes and analyte are strictly separated, thus facilitating the sensor cleaning and preparation processes. Contrary to the traditional timeconsuming multi-stage immunoassay, they allow a direct, rapid and on-line detection of label-free proteins, thus providing the opportunity to study the kinetics of the binding process [6,7]. Moreover, o n l / a small amount of analyte is required for the analysis.

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R. Dahint et al. /Sensors a,td Actu,~iors B33'-36 (1996) 497-505

One of the primary interaction mechanisms in these devices is mass loading caused by molecules bound or adsorbed at the surface of the device. Such perturbation of the sensor/liquid interface results in a change of the acoustic wave frequency and/or attenuation. This change has been shown to be a function of the amount of molecules bound to the sensing surface. In the present study, APM sensors have been used to study the non-specific adsorption of immunoglobulin G (IgG) and fibrinogen at artificial surfaces. The binding process at these surfaces is investigated on-iin~, and the protein adsorption characteristics are determined. Hexa(ethylene glycol)-terminated and methyl-terminated selfassembled monolayers (SAMs) of alkanethiols served as model coatings. While the former should exhibit very low protein adsorption, the latter tends to non-specifically adsorb a large amount of proteins [8,9], As a clean sensor response is crucial for the detectio, of small quantities of molecules, a new APM device design has been developed. Utilizing different interdigitai transducer (IDT) structures for input and output, the performance of the sensor is significantly improved. A dual delay line configuration is used to compensate for secondary effects and directly compare the adsorption behavior of differently derivatized crystal surfaces.

2. Experimental section 2. !. Improved APM sensor design Harmonic generation of acoustic waves by interdigitai transducers (IDTs) provides device operation at frequencies where the metal strip period is a multiple integer of the acoustic wavelength. Efficient excitation of APMs by spatial harmonics of the IDT will offer the possibility of realizing high frequency devices with relatively high mass sensitivity, However, these harmonic responses of. ten coincide with higher order APMs, which leads to distortion and deterioration of the sensor response. In most cases, external filtering is used to select the desired harmonic response. This yields external circuitry with additional driving sources and microcircuit components. In order to maintain the small and compact size of the APM device, and to reduce mode interference and distortion for improved sensor applications, a design synthesis is used, which does not require any extern~ti circuit filtering. Two different multi-electrode transducers (with different numbers of electrodes per wavelength) serve as the input and output IDTs. Fig. 1 shows a diagram of the new APM design. It utilizes four fingers per period (4F-IDT) and three fingers per period (3F*IDT) in the input and output, respectively. The 4F-IDT is known to operate at the fundamental frequency and odd harmonics, while the 3F-IDT operates at the fundamental frequency and mostly even harmonics, depending on the metallization ratio [10]. The individual frequencies are selected such that

TRANSMITTER 4-Finger lOT

3-Finger lOT RECEIVER

dominant mode i adjacent modes I /r] v +// 150 (3rcl) 250 (5th) 5O

I,f]ll,g, I[Jll P' I

~'~150 (2 ncl)

i i i

f [MHz]

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adjacent modes dominant mode

Fig, 1. The new sensor design utilizes a 4F-IDT for the transmitter and a 3F-IDT for the receiver. The lines represent the various modes excited (or received) by each IDT, While the transmitter is exciting the fundamental frequency and odd harmonies, the receiver is detecting the fundamental frequency and even harmonics. If the ratio of the individual fundamental frequencies is selected such that one or more harmonies of the dominant modes coincide, adjacent modes, however, interfere destructively, spurious sensor respon~ can be significantly reduced.

one or more harmonics coincide or overlap at one selected band of frequencies while other harmonics, adjacent modes and higher order APMs interfere destructively. Using this design, spurious sensor response and distortion should be drastically reduced. Note, that a practical implementation of this design requires an overlapping of the selected band of frequencies. This is complicated for APM devices whose responses depend on the actual plate thickness [ 11]. Thus, for the implementation of this design synthesis, a large number of APM spectra were calculated [12] and an appropriate combination of input and output IDT fundamental frequency and higher order modes was selected. The design parameters were ~. = 92.0,um and ,~ = 61.21tin for the period (wavelength) of the 4F- and 3F-IDT, respectively. On 0.5 mm thick ZX-LiNbO3, the selected crystal thickness and orientation, these values correspond to fundamental frequencies of about 48 Hz and 72 MHz, respectively. The remaining harmonic responses, as well as other higher order modes, which do not overlap, are illustrated in Fig. 1. The device is operated at about 145 MHz where the 2nd harmonic response of the 3F-IDT and the 3rd harmonic response of the 4F-IDT coincide. Note that the upper crystal surface was sputtered with a 20 mn adhesive layer of Cr and 60 nm of Au, which is known to slightly shift the device frequency response as a result of changes in the boundary conditions. The metaUized surface was used to eliminate acoustoelectric interactions with the analyte [13] and to facilitate the formation of thiol SAMs. Fig. 2 compares the response of the conventional 4F!4F-IDT design with that of the new 4F/3F-IDT design. The data in Figs. 2a,c represent non-filtered results while

R. Dahint et aL /Sensors and Actuators B35-36 (! 996) 497-505 i

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Fig. 2. Comparison of the conventional 4F/4F design (a,b) and the new 4F/3F design (c,d). For almost the same frequency of operation the non-filtered data of the 4F/3F structure yield a drastically improved sensor response (a,e). Only if spurious signals are numerically removed by gating (b,d) the same ,gnal quality is obtained for the two responses.

the spectra in Figs. 2b,d are obtained by numerically removing electromagnetic crosstalk and some spurious adjacent modes. This latter technique is known as 'gating'. It can be seen that the non-filtered data of the 4F/3F-IDT design indicate a much better sensor response, It is also shown that only after removing spurious signals by gating the quality of the two responses is the same. Based on these results, the new 4FI3F design was selected for almost all of the on-line adsorption studies described below. Only few experiments were performed with sensors of the conventional 4F/4F design. These devices have a IDT period of ;t = 88.0/~m and were fabricated on 0.5 mm thick ZX-LiNbO3 as well. They were operated at about 150 MHz, which is the 3rd harmonic response of the IDT structure. In any case, a dual delay line configuration was used in order to compensate for secondary effects and directly compare the adsorption behavior of differently derivatized crystal surfaces. The single propagation paths will be addressed as sensing and reference line in what follows.

2.2. Chemicals

Ethanol and H 2 0 2 (30%) were purchased from Baker (GroB-Gerau, Germany), H2SO4 (95-97%) ff~m Rie<~elde Hahn (Seelze, Germany), sodium dodecylsulfate (>99.9%) from Serva (Heidelberg, Germany) and ldodecanethiol from Sigma Chemie (Deisenhofen, Germany). Hexa(ethylene glycol)-terminated alkane-thiol (HS_(CH2)n-(OCH2)CH2)~)H) was synthesized by the group of Professor Whitesides, Harvard University [14]. Albumin bovine Fraction V (BSA, assay (CAF)> 99%, 11924) was bought from Serva, while phosphate buffered saline (PBS, P-4417), fibrinogen (F-4883, -50% protein, 20% sodium citrate and ~30% NaCI, -94% clottability) and ribonuclease A (RNA, R-5125) were obtained from Sigma. Catalog numbers are indicated in parentheses. Both affinity chromatography purified goat anti-rabbit IgG (lot no. 30081) and affinity chromatography purified rabbit anti-goat IgG (lot no. 30081) were purchased as solution (2.3 mg m1-1 PBS) from Jackson Immuno Re-

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search, Inc. (West Grove, PA 19390). Throughout the experiments, PBS was used at pH 7.4. As high concentrations of fibrinogen dissolved only slowly in PBS, the solutions were allowed to stand overnight. At the beginning of each experiment, they were filtered through 0.22,um hydrophilic cellulose acetate membrane filter units (Model (=450.1) obtained from Carl Roth GmbH (Karlsruhe, Germany). The actual concennations of the dissolved fibrinogen were determined by comparing the UV absorption at 214 nm with the absorption ofO.l%o RNA and 0.1%o BSA solutions. It was found [ 15] that at this wavelength, the light adsorption is almost independent of pH in the range from 4 to 8. Furthermore, the variation in amino acid composition of different proteins has relatively little effect on the measured adsorption values. This is due to the fact that at the low wavelength used, the dominant interaction mechanism is the absorption by peptide bonds. 2.3. Sensor preparation

In order to facilitate the formation of thiol SAMs, the non-electroded surface of the APM devices was sputtered with a 20 nm adhesive layer of Cr and 60 nm of Au. A single chamber teflon cell was mounted on top of the devices covering both sensing and reference line. After cleaning the gold surface for 15 rain in a freshly prepared piranha solution (!:3 mixture of H202 (30%), and H2SO4 (96%)) cooled down to 500C, the cell was rinsed ten times with ultraclean water (Miilipore) and blown dry with nitrogen. Caution. Piranha solution reacts violently, even explosively, with organic materials! [16] Both sensing and reference line were coated with Idodecanethiol by incubating the gold surface with 2 ml of a 5 mM thiol solution in ethanol for 12 h. After this solution had been removed, the cell was rinsed several times with ethanol and water and dried with nitrogen. Next, crystal and cell were tilted and 50°C hot piranha solution was added dropwise on the reference line. As determined by X-ray photoelectron spectroscopy (XPS), an incubation time of 15 rain is sufficient to remove the organic film completely. After extensively rinsing the reference line with Millipore water and a nitrogen blow dry, the bare gold region was exposed to a 5 mM hexa(ethylene glycol)-terminated aikanethiol (HS(CH2),I(OCH2CH2)6OH) [ 14] solution in ethanol for 1 h. Finally, the whole teflon cell was rinsed several times with ethanol and Millipore water. The film quality was checked by XPS, FT-IR spectroscopy and contact angle measurements on some larger (1.5 × 2.5 cm2) gold substrates prepared according to the same procedure. 2.4. A P M sensor measurements

In order to protect the sensor electrodes from corrosion and to facilitate cleaning and preparation procedures, the

Peristaltic Pump

"Three-Way-;al've; #2

Three-Way.Valve

/

#1

Plexlglass Cell

APM - Sensor

1 PBS Buffer 2 Protein Solution 3 Waste

Temperature Chamber

Fig 3. Block diagram of the continuous flow system.

APM device is mounted with the IDTs on the bottom surface. By this technique, the non-electroded crystal surface can be used for the sensing process. For each device, a liquid cell was attached to the top surface of the crystal plate to provide a well-defined interaction region between the analyte and the propagating acoustic wave. A fiat rubber gasket is used to guarantee a water-tight seal maintaining a constant liquid volume on the APM device. The whole arrangement is inserted into a temperature chamber with an integrated flow system as depicted in Fig. 3. The flow system consists of two liquid tanks containing the solutions used in the experiments and a container for waste. A peristaltic pump ensures continuous sample delivery at a controlled flow rate of about 0.6 ml min -I. By the use of two 3-way-valves it can be selected whether liquid is pumped from tank I or 2 and whether the solution is directed into the waste or back to tank 2. All components are mounted on a copper plate to provide good thermal contact. The measurements were performed at 25°C with temperature variations of less than :e0.2°C. Two techniques were used to expose the sensing surface to the analyte: a static injection system and a continuous flow system. In the static case, the liquid cell was filled with about 1.5 ml of PBS. After the system had stabilized, small amounts of protein solution were added to the buffer through a small opening between the reference and the sensing line. For the flow system, the liquid cell volume was about 0.5 ml. First, PBS was pumped through the cell, next replaced by a well-defined concentration of the protein solution. The amount of analyte used i~ reduced by switching 3-way-valve #2 in such position that the solution is pumped back into tank 2 after 5 or 10 rain for the experiments shown in Figs. 8 and 9, respectively. After the sensor response had reached a plateau, valve #2 was switched back to its initial position and the protein solution was substituted with PBS. Regeneration of the sensing surface is achieved by pumping a solution of sodium dodecylsuifate (SDS, 1% per weight in PBS at pH 7.4) through the liquid cell using again tank 2.

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R. Dahint et al. / Sensors and Actuators B35-36 (1996) 497-505

...... Relay

D'~10

0

Relay

described in detail [ 13,17] A P t 100 thermistor mounted at the bottom of the device monitors any temperature variations. The entire measurement is controlled by a personal computer. 3. Results and discussion

t

For the on-line detection of protein adsorption, the sensing line of the APM device was coated with hydrophobic, methyl-terminated SAMs (HS-(CH2) t1-CH3) and the reference line with hydrophilic, hexa(ethylene glycoi)-terminated SAMs (HS-(CH2) I r(O-CHz.-CH2) 6OH). If not indicated otherwise, the studies were performed with a network analyzer and sensors of the new 4F/3F design.

3. !. Non.specific and specific lgG adsorption Fig, 4. Block diagram of the electronics measurement set-up, For signal analysis, the sensor element can either be connected to a network analyzer or a combination of signal generatorand vector voltmeter. Two

relays allow to monitor sensing aml referenceline separately.Temperatur¢ is measuredwith a Pti00 and read out by the use of a digital mul° timeter.

The measurements consist of monitoring the phase and signal amplitude of the electrical output of the APM device relative to the input of the sensor as binding occurs. Two miniature relays allow to switch between the two propagation paths of the dual delay line configuration. Thus, both the individual responses of the two lines and the difference between the two responses can be monitored. The latter will be denoted as 'difference signal' in the following. In order to determine the relative signal changes due to protein adsorption, the response of the PBS covered sensor is taken as the reference. The measured shift in phase was used to calculate the shift in frequency as it would be observed in a phase-locked measurement. For this purpose, the phase-frequency characteristics of the two lines were determined at the beginning of each experiment. Single line responses are temperaturecorrected using the experimentally determined frequencytemperature characteristics (FTC) of 76 ppm/°C [ 13]. Two electronic configurations can be used for signal generation and analysis: a n:..twork analyzer and a combination of signal generator and vector voltmeter (Fig. 4). As the network analyzer allows to remove signal distortion from electromagnetic crosstalk and spurious acoustic modes by means of digital filtering, baseline stability and reproducibility of the measurements is significantly enhanced. However, this type of signal processing is rather time-consuming. Therefore, about 25 s are required to measure sensing and reference line separately with good resolution. In contrast, the combination of signal generator and vector voltmeter does not provide any filtering of the sensor response but accelerates signal analysis by a factor of 5. The two electronic systems have already been

FoUowing former experiments on the detection of antigen/antibody reactions [6,7], the static injection system was used for the first investigations. Fig. 5 shows the affinity of the two lines towards protein adsorption. As no devices with improved IDT design were available at that time, a sensor with conventional 4F/4F design was selected. After filling the liquid cell with about 1.5 ml PBS and monitoring signal stability for 30rain, l l5/~g goat anti-rabbit IgG dissolved in 50/zl PBS were added to the buffer solution. While no signal change is observed for the reference line, the sensing line shows a frequency decrease of about 350 Hz. Thus, it is confirmed in an online experiment that hexa(ethylene glycol)-terminated SAMs exhibit very low protein adsorption, methylterminated SAMs on the other hand tend to nonspecifically adsorb large amounts of protein. This is in good agreement with previous studies using ex situ spectroscopy [8] or in situ surface plasmon resonance (SPR) i

0.2 oO

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Time [min] Fig. 5. The sensing line of the APM device was coated with Idod,eeaneth~.ol, the reference line with a hexa(¢lhylene glycol)terminated SAM. After30 min I 15/~g goat anti-rabbit igG dissolvedin 50/JI PBS were injected into the liquid cell. While no binding processes are observed for the referenceline, the sensing line shows a frequency decreaseof about 350 Hz due to non-specificlgG adsorption.

502

R. Dahint et al. I Sensors and Actuators B35-36 (1996) 497-505

coo°

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•

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Fig. 6. Conlinuationof the experiment hi Fig. 5. After rinsing the liquid cell with PBS, 23/~g rabbit anti-goat IgO dissolved in 10pl PBS were injected. As expected, a strong reduction in frequency is noticed due to the specific rcacqion of tim proteins wilh Ihe preadsorbed goat antirabbit igG,

analysis [9]. It can, therefore, be concluded that an hexa(ethylene glycol)-coated sensor line is an excellent reference for all further experiments as protein adsorption is effectively suppressed. Having verified that hexa(ethylene glycol)-terminated SAMs are protein resistant, the following studies were focused on the adsorption characteristic.,~ of the methylterminated surfaces. In order to effectively eliminate secondary effects, such as temperature changes, from the sensor response without any numerical computation, the glycol-terminated line was taken as a reference and the difference signal, defined as sensing line minus reference line response, was considered, In continuation of the experiment described above, the liquid cell was rinsed with PBS to remove any lgG which had not adsorbed onto the sensor surface. Afterwards the cell was refilled with about 1.5 ml buffer solution and

goat antl-rabblt IgG

0,2

,

1

HCO

23/tg rabbit anti-goat IgG dissolved in 10/~! PBS were injected (Fig. 6). This protein can specifically react with goat anti-rabbit IgG adsorbed at the surface of the sensor in the first stage of the experiment. As expected, a strong reduction in frequency is observed. A comparison of Figs. 5 and 6 suggests that almost twice the amount of protein is bound in the second step. This is not surprising as for the specific antigen/antibody system used, both proteins can bind one another via their antigen binding sites. Thus, a 2:1 binding can easily be achieved even if not all antigen binding sites of the preadsorbed antibody are accessible for the added protein. As indicated above, all experiments described in the following have been performed with sensors utilizing the new 4F/3F IDT-design. The fact that the amount of protein adsorbed at the sensing line is not only determined by the hydrophobicity of the coating but also a function of the pH of the solution is demonstrated in Fig. 7. Lowering the pH of the buffer by successive injection of 10/tl of 2 N HCI releases all preadsorbed proteins. However, a complete removal of the protein layer could not be achieved in all experiments. The reasons for this behavior are still unclear. The above studies show that valuable information on protein adsorption can be obtained by the use of a static injection system. A major advantage of this technique is that only small amounts of analyte are required for analysis. However, it is hard to extract kinetic data from the curves, as the concentration profile of the proteins in the buffer solution is not well-defined during the adsorption process. As the proteins are injected directly across the sensor lines, the sensing surface is exposed to a comparatively high protein concentration right after injection. Afterwards the concentration is reduced by diffusion processes into the bulk of the liquid until equilibrium is obtained. Thus, for quantitative kinetic studies the use of a continuous flow system was preferred.

=Ibrinogen

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0,0 . die w

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°

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-0,6

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30

60

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90

Time [rain] Fig, 7, lh'oleins axe adsorbed onto the methyl-terminated sensing line at pH 7,4 by adding twice 23,ug rabbit anti-goat lgG dissolved in lOpl PBS Io ~ buffer solution. Note, that the protein adsorption sites at the stnso¢ surface have already been saturated after the first injection. Lowering the pH of the buffer by adding HCI results in a complete removal oftbe lxotein layer.

o

'

2'0

'

20

'

6'0

Time [min] Fig. 8. After monitoring stability of the s~nsor for 20 rain, the sensor is exposed to a 0.8 mg ml-I fibdnogen solution. Subsequent flushing with PBS only removes a small portion of the adsod)ed proteins. However, by the use of SD$ and PBS the sensor element can be completely re-

generated.

503

R. Dahint et al. / Sensors and Actuators B35-36 (1996) 497-505 •

i

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Time [min] Fig. 9, Adsorption of fibrinogen is monitored for different concentrations of protein. It is observed that frequency shifts at equilibrium strongly depend on protein concentration. The solid lines show a fit to the experimental data according to Eqs. (6) and (9).

3.2. Non-spec(~c adsorption of fibrinogen

After proving that the glycol-terminated SAMs also behave protein resistant with respect to fibrinogen exposure, the adsorption characteristics of fibrinogen at methyl-terminated surfaces were investigated by the use of a continuous flow system. Again, a glycol-terminated line served as the reference. Reversibihty of protein adsi ,rption was checked by pumping a 0.8 mg ml -I fibrinogen solution through the liquid cell. Due to the adsorption process, a fast frequency decrease of about 700 Hz is obtained (Fig. 8). As flushing the cell with PBS reduces the sensor response only by about 120 Hz, it has to be concluded that the main portion of fibrinogen is irreversibly attached to the methyl-terminated line. However, complete regeneration of the sensor element can be achieved by the use of SDS and a subsequent flow of PBS. All proteins are removed from the surface and the sensor returns to its initial state. Again, these results are in good agreement with SPR analysis [9]. Moreover, they indicate that one coated sensor element can be used for a complete set of measurements, thus effectively reducing possible systematic errors due to slightly different mass sensitivities of the devices or minor differences in the quality of the coatings. Combining the above observation with the results on antigen/antibody reactions (Fig. 6), it can also be concluded that thiol coated sensors allow the implementation of a new type of immunosensors, which are easily prepared by non-specific adsorption of antibodies and easily reusable with different antigen/antibody systems by completely removing the antigen/antibody complexes from the surface. This may be advantageous compared to conventional biosensors where the receptors are irreversibly linked to the sensing surface and which, therefore, only allow the detection of a single species of antigen. A complete set of adsorption experiments is presented

in Fig. 9. Here, fibrinogen adsorption is monitored for protein concentrations ranging from 3 to 280/~g ml -I. After each measurement, the sensor is regenerated by flushing the liquid cell with SDS and PBS. It is clearly seen that sensor response is enhanced if protein concentration increases. The observed frequency shifts at equilibrium are shown in Fig. 10. Concentration dependent equilibrium states usually occur for reversible adsorption/desorption processes as in classical Langmuir kinetics. However, the observation that the main portion of fibrinogen is irreversibly bound to the methyl-terminated surface refutes this model. Another explanation for the above behavior is based on the fact that adsorbed proteins can change their threedimensional structure by interaction with artificial surfaces [18]. In this so-called unfolded state they may occupy larger surface areas than in their original native configuration. At high protein concentrations, however, the adsorption process is accelerated. Thus, the surface will be completely covered with proteins before a considerable amount of them can unfold. As a result, a higher number of proteins will be able to bind to the surface than in the low concentration regime. A relatively simple mathematical description of the latter approach is obtained by the following kinetic model: As illustrated by Fig. l l, native proteins may adsorb from solution onto an artificial surface. The rate of adsorption is assumed to be proportional to the concentration of proteins in solution, c, and to the free surface area, ( l - 0): dO

--_..L = kac( 1 _ O) dt

(I)

In this equation, 0 = 01 + 02 denotes the percentage of surface area covered with proteins. It consists of two contributions: 01, which is the area that would be occupied, if all proteins were still in their initial state (black circles in Fig. 11), and 02, which is the increase in coy-

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(9).

504

R. Dahint et aLI Sensors and Actuators B35-36 (1996) 497-505

ered area due to the unfolding process (grey rings in Fig. 1|). Both 0t and 02 are normalized to the total surface area. ka is the time constant for the adsorption process. It should be emphasized that 01 is proportional to the number of adsorbed proteins and thus determines the stre,tgth of the sensor response in case of mass loading. After adsorption, the proteins will unfold at a rate that is proportional to the number of adsorbed molecules and to the non-covered surface area:

dO-.--Z.2 = kuO 1(I - O)

(2)

dt

where k, denotes the time constant at which unfolding occurs. To reduce the complexity of the mathematical solution, it is assumed, that the unfolding process is only limited by the available surface area and not by any restrictions due to the actual size of the molecules. This simplification is regarded to be acceptable in case of not too low surface coverages. Considering Eqs. (1) and (2) we find that

Table 1 Computed time constants for the experiments shown in Fig. 9 c (rag ml-I)

ku (I)

ka (ml mg -I)

3 14 28 140 280

0.246 ± 0.039 0.374 ± 0,032 0.266 ± 0.017 0.275 ± 0.017 0.146 ± 0.009

(2.41 ± 0.38) x (3.66 ± 0.34) x (2.60 ± 0.17) x (2.69 :l: 0.18) x (1.43 ± 0.10) x

10-3 10-3 10-3 10-3 10-3

The fit routine used is based on the Levenberg Marquardt algorithm. It is seen that the values of ku and ka fall within the same range, although the variations are higher than the computed errors. 2kac

I

01 (r) =

(6) with

dO

d0l +d02

d01 + k u 0i dOI

k:

(3)

"77

I 2kac l 2 8k~c

(7)

't'=t-77f) + which is solved by

O=Ol + ku 02 2kac

(4)

if the boundary conditions O(t = 0) = 0 and Ol(t = O) = 0 are taken into account. Using again Eq. (1), we finally obtain

I

dOI : k a ¢ 1 - 0 I

dt

2k=c

(5)

J

The physical solution of the above differential equation is

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Unfolding " " " Protein

Native Protein

Fig, ! 1. Schematic view of the adsorption process: Proteins binding to an artificial surface will cover a relatively small area (black circles) as long as they maintain their native state, Due to the unfolding process additional surface area is occupied (grey rings).

For long adsorption times (t ~ 00), this solution reduces to the concentration dependent equilibrium states

k°c + 0, =---~-

~

2

kac

2kac + k'-'~

(8)

In order to compare the predictions of the above model to the measured adsorption kinetics of fibrinogen, a relationship between sensor response and 01 has to be established. Assuming that mass loading of the crystal surface is the dominant interaction mechanism, the observed frequency shifts, Af, will be proportional to the number of bound molecules. Maximum response, Afmax, is observed if the sensing surface is completely covered with native proteins. Therefore, Af = &fmax0j

(9)

Taking Eqs. (8) and (9) to fit the data of Fig. 10 results in k = kalku = (9.79 _ 0.32) × 10-3 mg /d -] and Afmax= 0.688 :t: 0.089 kHz. Good agreement is achieved between theoretical and experimental data. Next, ku is obtained for each concentration value by fitting the data set of Fig. 9 with Eqs. (6) and (9). During this procedure, k and Afmax are restricted to the above determined values and only varied within experimental error. All fit routines were based on the Levenberg Marquardt algorithm [19]. A summary of the results is presented in Table 1.

R. Dahint et al. / Sensors and Actuators B35-36 (1996) 497-505

505

Although the variations of k, and k a at different protein concentrations are higher than the numerically computed errors, all values fall within the same range. This indicates that the above model is suitable to explain the experimentally determined adsorption kinetics in a self-consistent way. A larger set of experiments could probably further reduce inconsistencies. The observed deviations may, however, also be attributed to small differences in the surface properties after each regeneration step or to slightly varying flow conditions in the liquid ceil. An optimization of the present flow system should overcome the latter problem.

The authors wish to thank G.M. Whitesides and M. Mrksich for their advise regarding the preparation of the thiol films and for providing them with hexa(ethylene glycol)-terminated alkanethiol solutions, which are synthesized at Harvard University and not commercially available. They would also like to acknowledge R. Kohring and A. Lampert for their help in implementing the fitting routines used for data analysis. This work was funded by the Deutsche Forschungsgemeinschaft and the Fond der Chemischen Industrie.

4. Conclusion

References

In the present study, APM sensors with improved sensor characteristics have been used to investigate the interaction of proteins with artificial surfaces in situ. It was found that hexa(ethylene glycol)-terminated alkanethiol SAMs behave protein resistant with respect to both IgG and fibrinogen adsorption, whereas methyl-terminated ( H S - ( C H 2 ) 1 1 - C H 3 ) films adsorbed high amounts of protein at appropriate pH values. Uniform quality of the SAMs was controlled by IR-, XPS- and contact angle measurements. The sensor element was effectively regenerated by flushing the surface with SDS and PBS. Based on experiments on antigen/antibody reactions a new type of immunosensor is proposed which is easily reusable with different antigen/antibody systems. Studying the adsorption kinetics of fibrinogen at methyl-terminated surfaces, it was found that the most part of adsorbed molecules is irreversibly bound. A rise in protein concentration increases steady state sensor response. The observed adsorption behavior is selfconsistently explained by a two-step process which involves the initial binding of native proteins followed by an unfolding process. However, additional biochemical and spectroscopic experiments will be necessary to determine whether the above self-consistent model accurately reflects the real situation. It should, furthermore, be emphasized that the calculated adsorption constants are closely related to the specific experimental situation. Nevertheless, the study clearly reveals the advantages of on-line detection systems: In combipation with mathematical models of the adsorption process, valuable information about the interaction of proteins withartificial surfaces can be obtained that are not directly accessible with other experimental methods. In an attempt to further improve the performance of this technique, new sensor designs with enhanced signal-to-noise ratio are presently developed. An optimization of both the flow system and the quality of the coatings should effectively reduce experimental error.

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Acknowledgements