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Study of acoustic Love wave devices for real time bacteriophage detection
Sensors and Actuators B 91 (2003) 275–284
Study of acoustic Love wave devices for real time bacteriophage detection Ollivier Tamarina, Corinne De´jousa,*, Dominique Rebie`rea, Jacques Pistre´a, Sylvie Comeaub, Daniel Moynetc, Jean Bezianc a
Laboratoire IXL, CNRS UMR5818, ENSEIRB, Univ. Bordeaux 1, 351 Cours de la Libe´ration, F-33405 Talence Cedex, France b Cellule Adera/Bordeaux 2, Bordeaux, France c Laboratoire d’Immunologie Mole´culaire, Univ. Victor Segalen Bordeaux 2, 146 Rue Leo Saignat, F-33076 Bordeaux, France
Abstract Acoustic wave devices have shown their good potentialities for real time monitoring of immunoreactions. Different acoustic wave devices (BAW, SHAPM, Love waves) were described for applications in liquid medium. Love wave delay line structures (ST cut quartz substrate with interdigital transducers, SiO2 guiding layer) present several advantages, in particular, the pure shear horizontal polarisation adapted to liquid medium, and its very high sensitivity related to the wave confining in the thin guiding layer. In this paper an analytical method based on multilayer propagating structure is first presented: it allows us to estimate the Love wave phase velocity and then the mass loading effect sensitivity. A few theoretical results are exposed; they show that this theoretical analysis can allow to optimise physical parameters in order to conceive powerful devices for detection applications in liquid medium. As a model for virus or bacteria detection in liquids (drinking or bathing water, food, etc.), we design a model using M13 bacteriophage. The first step is the anti-M13 antibody binding. By using a Labwindows CVI software, we can monitor in real time the graft of the anti-M13 antibody sensitive coating, as well as the detection of the M13 bacteriophages. Experimental results are exposed, analysed and discussed. Love waves sensors appear to be a powerful approach for immunodetection, as theoretically predicted. # 2003 Elsevier Science B.V. All rights reserved. Keywords: Biosensors; Immunodetection; Love wave; Gravimetric sensitivity
1. Introduction The scope of acoustic (bio)sensors receive a great deal of attention for a few years as they can realise real time and in situ experiments in a large domain (e.g. medical applications, food industries or environment). A lot of studies concern detection in gas medium using mainly SAW devices [1]. Applications in liquid medium need specific acoustic polarisation as pure shear horizontal (SH) waves. Thus detection tests in liquid medium with quartz crystal microbalance (QCM) devices [2], and SH acoustic plate mode (SHAPM) devices [3] were presented, and allowed to validate real time experiments. Our aim is to develop powerful devices in terms of mass loading effect sensitivity (gravimetric sensitivity). Love wave devices, as the wave energy is confined in a thin *
Corresponding author. Tel.: þ33-5-56-84-2848; fax: þ33-5-56-37-1345. E-mail address: [email protected] (C. De´jous).
guiding layer (a few micrometers), present a high sensitivity compared to other acoustic devices [4], moreover the SH polarisation allows to work in liquid medium. The realisation of experimental Love wave delay lines needs a very fine technology, as the interdigital transducers which generate the wave are periodical long fingers (10 mm width, separated by 10 mm), without any electrical short. In order to minimise the technological realisations and conceive best adapted devices before manufacturing, an analytical method allowing to express a real dispersion relation linking the Love wave phase velocity to the other parameters of the device is presented. A second step allows to model the device gravimetric sensitivity. Thus we can estimate the liquid and device physical parameters influence on the Love wave sensor sensitivity, and so define optimised parameters. Then, experimental tests, using one of the theoretically defined devices, aim at validating the feasibility of biodetection using such a Love wave sensor. For that, an immunological model is proposed. An anti-M13 antibody sensitive
0925-4005/03/$ – see front matter # 2003 Elsevier Science B.V. All rights reserved. doi:10.1016/S0925-4005(03)00106-0
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layer is grafted on the Love wave delay line surface. As model analyte, we use the associated M13 bacteriophage. Results are analysed and discussed in order to conclude about the great interest of Love wave delay lines for detection in liquid medium.
2. Theoretical multilayered model Love wave (bio)chemical sensors, consist in the guided-SH acoustic delay line associated with a (bio)chemical layer sensitive and selective to the specie to detect. Previous theoretical works on detection tests in gas medium [4], and on the influence of a Newtonian liquid medium [5] on the propagation characteristics of the Love wave allow to estimate the phase velocity versus mechanical properties of the system. Here, we must take into account other physical properties due to the sensitive coating. This induces other interfaces like guiding layer/sensitive coating, sensitive coating/liquid. Changes due to the interaction between the overlay and the acoustic waves induce wave velocity variations. In many applications for gas detection, the most significant and best controlled effect is the wave velocity perturbation due to mass loading. Our first aim was to develop a model of the Love wave structure allowing an analytical resolution in order to predict its mass sensitivity, when working in a liquid medium. Love waves propagate in layered structures (cf. Fig. 1) consisting of a piezoelectric substrate (e.g. quartz) with an isotropic guiding layer (e.g. SiO2), the isotropic sensitive layer and the isotropic Newtonian liquid in our application. The analytical resolution consists in the determination of the SH acoustic wave propagation in the x1 direction in each volumic material. Then, a dispersion relation linking the physical parameters of the device to the Love wave velocity is deducted from the boundary conditions at each interfaces and at infinities of the multilayered structure. 2.1. Bulk SH waves In this part we will express the transverse displacements in each volumic material of the multilayered structure. As in
Fig. 1, each layer is characterised by the material physical constants (elastic modulus: Cijkl,s,h,b,L, density: rs,h,b,L, viscosity: Z for the liquid). The indices s, h, b, L characterise the physical constants in the substrate, the guiding layer, the sensitive coating and the liquid, respectively. The layers thicknesses are h and b for, respectively, the guiding layer and the sensitive coating. The general expression of the pure SH wave propagating along the acoustic path (between the interdigital transducers) is expressed as U2a ðx1 ; x3 ; tÞ ¼ U2a ðx3 Þ exp½jðot kx1 Þ where U2a ðx1 ; x3 ; tÞ is the acoustic SH particle motion, in the x2 direction, U2a ðx3 Þ the transverse acoustic displacement in the x3 direction (the expression of U2a ðx3 Þ depends on the medium of propagation represented by a), and exp½jðot kx1 Þ represents the wave propagation along the x1 direction, where k ¼ 2p=l (l the wavelength) is the wave number. The pulsation o of the wave is defined as: o ¼ l=Vp , where VP is the Love wave phase velocity. 2.1.1. The semi-infinite quartz substrate The weak electromechanical coupling coefficient ðK 2 0:1%Þ of the quartz, used for the substrate in order to generate pure SH waves, allows to neglect the piezoelectricity of this material in the first approximation. With this assumption, only the mechanical behaviour of the system (transverse acoustic displacement, and normal stress) has to be considered. In order to express the transverse acoustic displacement we use the Hooke’s and Newton’s laws written, respectively, as follows: Tij ¼ Cijkm
@Um ; @xk
i; j; k; m ¼ 1; 2; 3;
@Tij @ 2 Ui ¼r 2 @xj @t
Using these expressions, in the case of transverse horizontal waves, we obtain the equation of motion in elastic nonisotropic solids [4] C44s
@ 2 U2s @U2s 2jkC46s þ ðro2 k2 C66s ÞU2s ¼ 0 @x3 @x23
Fig. 1. Basic configuration of the Love wave delay line with a sensitive coating (sensor configuration).
(1)
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By solving the above equation, we can write the expression of the transverse acoustic displacement in the quartz substrate as follows: kC46s U2s ðx1 ; x3 ; tÞ ¼ U0s exp j þ kws x3 C44s exp½jðot kx1 Þ;
x3 < 0
(2)
where 2.1.2. The sensitive coating and the guiding layer Using Eq. (1) in the case of isotropic solids (C44a ¼ C66a ¼ ma and C46a ¼ 0), we can write the expression of the transverse acoustic displacement in both guiding layers (Eq. (3)) and sensitive coating (Eq. (4)) media: U2h ðx1 ; x3 ; tÞ ¼ U0h cosðkwh x3 þ jÞ exp½jðot kx1 Þ; 0 x3 h
(3)
qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi where wh ¼ rh Vp2 =mh and j is a constant. U2b ðx1 ; x3 ; tÞ ¼ U0b cosðkwb x3 þ j0 Þ exp½jðot kx1 Þ; h x3 h þ b ¼ L qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi where wb ¼ rb Vp2 =mb and j0 is a constant.
2.1.3. The semi-infinite liquid medium In the liquid medium, we express the SH particle motion U2 ðx1 ; x3 ; tÞ with the Navier–Stokes law. Furthermore, we suppose a perfect horizontal system, then the pressure and the gravity can be neglected: @ 2 U2 @ @ 2 U2 @ 2 U2 rL ¼Z þ (5) @t @x23 @t2 @x21 For the resolution of Eq. (5) we suppose the x1, x3, and t variables independent and we express U2 ðx1 ; x3 ; tÞ as: U2L ðx1 ; x3 ; tÞ ¼ Sðx3 Þ expðjkx1 Þ expðjotÞ, where S(x3) represents the particle motion in the x3 direction. By including the last expression of U2L ðx1 ; x3 ; tÞ in Eq. (5), we obtain 1 @ 2 Sðx3 Þ r o ¼ j L þ k2 2 Sðx3 Þ @x3 Z
(6)
As the sound velocity for a shear mode in a solid is much larger than in a liquid, we can write k2 ! orL =Z [6]. This assumption can be verified by considering typical values of Love wave delay lines physical parameters. For example, with l ¼ 40 mm, f ¼ 100 MHz, rL ¼ 103 kg m3, Z ¼ 10 cP, we obtain k2 2 1010 m2
and
orL 6 1013 m2 Z
Then Eq. (6) can be simplified as 1 @ 2 Sðx3 Þ r o ¼j L Sðx3 Þ @x23 Z
The solution of this simplified Navier–Stokes law gives us the expression of U2L ðx3 ; tÞ. For this expression, as there is no source of energy in the liquid, we suppose a wave propagating in the x3 direction, this induces that the wave amplitude is equal to 0 for x3 ! þ1: U2L ðx1 ; x3 ; tÞ ¼ U0L exp½ð1 þ jÞkL x3 exp½jðot kx1 Þ (7) pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi with kL ¼ orL =2Z . We can estimate the expression of the decay length defined for a diminution of the wave amplitude with a factor 1/e. This will be useful for experimental tests, in order to verify the assumption of a semi-infinite liquid layer, and so neglect the influence of a liquid volume change. Expression (7) allows us to determine an approximate value of the decay length d in the x3 direction 1 d¼ ¼ kL
(4)
277
sffiffiffiffiffiffiffiffiffi 2Z orL
This expression of d is used by several authors in the case of SH waves [6–8]. This is an approximate value of the real decay length. In order to estimate the real decay length, we would have to consider the Navier–Stokes law in two space dimensions (x1, x3), given by Eq. (5). These first steps lead to the SH wave displacement form in each material. In a second step we apply the equations of continuity at each interface of the four layered device. 2.2. Sensor: quartz substrate, SiO2 guiding layer, sensitive coating, liquid In this structure (cf. Fig. 1), we consider three interfaces: on both faces of the sensitive coating, with the liquid ðx3 ¼ LÞ and with the guiding layer ðx3 ¼ hÞ, and one between the quartz substrate and the SiO2 guiding layer. This allows to write a complex dispersion relation. 2.2.1. SiO2/liquid interface: x3 ¼ h By applying the continuity of acoustic displacement velocity between the liquid and the sensitive coating, U2L ðx3 ¼ Lþ Þ ¼ U2b ðx3 ¼ L Þ (Eqs. (4) and (7), respectively), and the continuity of the normal stress T32L ðx3 ¼ Lþ Þ ¼ T32b ðx3 ¼ L Þ (Hooke’s law), we obtain the following set of equations: joU0b cosðkwb L þ j0 Þ ¼ U0L exp½ð1 þ jÞkL L; kwb mb U0b sinðkwb L þ j0 Þ ¼ ð1 þ jÞkL ZU0L exp½ð1 þ jÞkL L The ratio of the equations of the previous system leads to
kL Vp Z 1 j0 arctan ð1 þ jÞ (8) k¼ wb L mb wb wb L
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2.2.2. Sensitive coating/SiO2 interface: x3 ¼ h, determination of j0 In the same way, we consider the continuity of acoustic displacement velocity between the sensitive coating and the guiding layer, U2b ðx3 ¼ hþ Þ ¼ U2h ðx3 ¼ h Þ (Eqs. (3) and (4), respectively), and the continuity of the normal stress T32b ðx3 ¼ hþ Þ ¼ T32h ðx3 ¼ h Þ (Hooke’s law). This leads to the following set of equations: U0h cosðkwh h þ jÞ ¼ U0b cosðkwb h þ j0 Þ; kwh mh U0h sinðk1 wh h þ jÞ ¼ kwb mb U0b sinðk1 wb h þ j0 Þ The ratio of these equations gives us the quantity j0 :
mh wh 0 tanðkwh h þ jÞ kwb h j ¼ arctan mb wb
Eq. (10) allows to estimate the Love wave phase velocity (Vp) in a particular structure with a sensitive layer and in a Newtonian viscous liquid, provided that we know all material parameters. Note that if we consider a system with Z ¼ 0, we obtain dispersion relations defined in gas medium detection applications [4]. Furthermore, compared to a complete numerical resolution, the analytical approach presents two main advantages. First of all, the Love wave phase velocity can so be calculated quite easily in only a fraction of a second, instead of several minutes with a numerical approach. And the analytical expression allows to visualise physically how the materials properties are involved in the wave propagation.
3. Theoretical results
In this expression and the following ones, all ‘‘arctan’’ functions are modulo p. 2.2.3. Quartz/SiO2 interface: x3 ¼ 0, determination of j Continuities of the acoustic displacement U2b ðx3 ¼ 0þ Þ ¼ U2h ðx3 ¼ 0 Þ (Eqs. (2) and (3)) and of the normal stress T32b ðx3 ¼ 0þ Þ ¼ T32h ðx3 ¼ 0 Þ give the following set of equations: U0s ¼ U0h cosðjÞ; U0h kC44s ws cosðjÞ ¼ U0h kC44h wh sinðjÞ Then we can write the quantity j: C44s ws j ¼ arctan C44h wh By including the quantities j and j0 in Eq. (6), after simplifications, we obtain
k L Vp Z kwb b ¼ arctan ð1 þ jÞ C44b wb
C44h wh C44s ws arctan tan kwh h arctan C44b wb C44h wh (9) with b ¼ L h. In order to write the dispersion relation, we identify the real part of the second member of Eq. (7) to kwbb, and we obtain the following dispersion relation:
C44h wh C44s ws kwb b ¼ arctan tan kwh h arctan C44b wb C44h wh 1 k L Vp Z arctan (10) 2 C44b wb The imaginary part of the second member is related to a contribution of the viscous damping of the wave. In order to take into account this part, we should have considered in this case a complex wave number k ¼ k þ jg instead of k; and in this case the analytical approach becomes very difficult or even impossible.
For detection application, among all possible detection mechanisms, one of the main effects induced is the sensitive layer mass increase when target molecules are sorbed on it. So, it is interesting to use previous studies in order to estimate the gravimetric sensitivity (i.e. sensitivity to mass loading effect) of the sensor when it is in contact with a liquid medium as a function of the device parameters, and of the liquid parameters and more particularly its viscosity. First of all, is allows to compare several acoustic propagations in terms of powerfulness for (bio)detection tests. And secondly it allows to improve delay lines conception in a very short time. To model this effect, up to now for gaseous application, we consider a little increase Drb of the material density, which induces a phase velocity shift DVP during the detection. Then the velocity sensitivity to the mass loading effect is defined by the following expression: Smv ¼
DVP 1 VP0 Drb b
(11)
where VP ¼ VP VP0 , VP0 the Love wave velocity with a sensitive layer density r ¼ rb, and VP the Love wave velocity with a sensitive layer density r ¼ rb þ Drb. In order to calculate this gravimetric sensitivity we define a Love wave test device with the following physical characteristics: a quartz substrate, a wavelength l ¼ 52 mm, a SiO2 guiding layer thickness of 6 mm (mh ¼ 3:12 1010 N/ m2, rh ¼ 2203 kg/m3 [9]), a PMMA sensitive coating layer thickness of 100 nm (mb ¼ 0:11 1010 N/m2, rb ¼ 1109 kg/m3) [10]. The PMMA was chosen in a first step as sensitive material because it was already used in previous works in gas medium [4], and so its physical parameters are well known. 3.1. Influence of the delay line parameters on the gravimetric sensitivity In order to define well-adapted devices for detection application, we investigate first the influence of the delay line parameters (wavelength, SiO2 guiding layer thickness) on the gravimetric sensitivity.
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Fig. 2. Influence of h/l on the gravimetric sensitivity in a Newtonian viscous liquid medium (Z ¼ 3 cP, rL ¼ 103 kg/m3).
Fig. 4. Influence of Newtonian liquid viscosity on Love wave phase velocity (test device [5]).
Results reported in Fig. 2 show that it is possible to increase the gravimetric sensitivity up to a factor 3 if we reduce the wavelength from 52 to 20 mm. The maximum value of gravimetric sensitivity is always obtained with an appropriate guiding layer (h=l 0:17 in this case). Similar results were also shown theoretically and experimentally for Love wave devices used for detection application in gas medium [4].
important parameter, provided that it is low enough to allow wave propagation. On the other hand, previous studies using aqueous solution of glycerol [5] showed that the viscosity had a great influence on the wave phase velocity, as represented in Fig. 4. And so, any slight viscosity change during experimental tests will induce a great velocity shift. Finally, this theoretical study shows that it is possible to define Love wave devices offering a particularly high gravimetric sensitivity. Best parameters so defined for detection in a liquid medium are very closed to those obtained in gas applications. This is partially related to the choice of a typical polymeric material as sensitive layer, as the acoustic properties of biological films are unknown yet.
3.2. Influence of the liquids mechanical parameters on the gravimetric sensitivity Using Eq. (11) again, it is also interesting to estimate the gravimetric sensitivity of the Love wave test device in a Newtonian viscous liquid medium as a function of the liquid viscosity Z. We consider only weak values of viscosities in order to avoid excessive loss of acoustic energy due to high viscous damping [7]. In Fig. 3 the results with three viscosities from 1 to 7 cP are shown. These curves show that an increase of the liquid viscosity only slightly increases the device gravimetric sensitivity. So, the liquid test viscosity should not be an
Fig. 3. Influence of the liquid viscosity on the gravimetric sensitivity (rL ¼ 103 kg/m3).
4. Experimental setup 4.1. Love wave device For detection tests, Love wave delay lines such as that previously defined in part III were realised. The technological realisation of this delay line is very important in order to obtain good sensing properties. A careful attention must concern, the engraving of the IDTs, and the deposition of the guiding layer. In our application, 50 pairs of aluminium split fingers with a wavelength l ¼ 52 mm are engraved on the quartz substrate. Split fingers allow to attenuate reflection when acoustic wave arrive under the transducers, and so to greatly attenuate the triple transit echo. The aperture, and the centre to centre distance between the IDTs are, respectively, 2.6 and 6.5 mm. This realisation needs a very fine technology in order to obtain a so periodical structure (long fingers, width l/8, separated by l/8, without any electrical short). The SiO2 guiding layer is realised on the substrate and the IDTs using PECVD (plasma enhanced chemical vapour deposition), which allows to obtain up to a few microns of SiO2 (6 mm for the test device) at low tem-
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Fig. 5. Direct frequency response of the Love wave delay line (S21).
peratures (the quartz must be kept under its phase transition temperature). These characteristics allowed to obtain transmission losses near 17 dB at the centre frequency, about 88 MHz (cf. Fig. 5). The very linear phase allows to realise an oscillating configuration with a very high short term stability. The involved technological steps are used in standard planar microelectronics processes, and so constitute an advantage compared to QCM devices. Furthermore, these technologies greatly get from the huge development of delay lines and resonators for telecommunications applications. 4.2. Test bench For operation in liquid medium, a specific test bench is used. Firstly, the Love wave delay line (500 nm thick) is glued with a silicon seal and gold wire-bonded on a printed circuit board (PCB). It becomes then an interchangeable device which can be placed in and removed from an oscillator loop. In order to measure the sensor temperature an AD590 temperature sensor placed between the delay line and the PCB is used. Secondly, a Teflon cell with a rectangular cavity (121 mm2) localised on the acoustic path of the delay line is placed on the sensor with a Teflon seal.
Finally, this system is placed in a thermoregulated dry bath. For all experiments, a single delay line is used. In order to estimate the perturbation of the Love wave characteristics, we used the delay line mounted in an oscillator configuration, and/or connected to a network analyser (Anritsu Scorpion MS4623B). The measurement of the oscillation frequency allows to obtain the highest resolution on any wave phase velocity variation. Moreover, this configuration for a complete system can be portable as demonstrated by other works [11]. Here, real time experiments are realised by recording the frequency and temperature as a function of the time with a Labwindows CVI software. Typically, with our oscillation loop, we can obtain a short term stability of 1 Hz/s. When more parameters are needed, the delay line can be characterised with a network analyser (Fig. 5). Then both wave phase velocity and transmission losses can be measured, as a function of frequency. 4.3. Biological model 4.3.1. Choice of antibody and bacteriophage immunoreaction model The M13 bacteriophage (bacteria infecting virus) was chosen as the target particle to detect. Indeed, it is quite easy to numerate the phages by using their property to infect ‘‘male’’ E. coli and to form plagues on a bacterial lawn.
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Fig. 6. Sensor surface after the detection application.
Several well-characterised anti-M13 monoclonal antibodies directed against the major proteins constituting the M13 are commercially available. And the M13 size is comparable to many pathogen viruses (900 nm long and 9 nm in diameter). These reasons justified the choice of this bacteriophage, as an immunological model for future developments. 4.3.2. Immunoreaction procedure This procedure consists of two main steps: the immobilisation of a biological layer, and of the selected bacteriophage (detection). The first step is the immobilisation of the antibody antiM13. For that, the sensor surface is cleaned: H2SO4 (4 N), three washes with H2O(d), two washes with 95% ethanol, two washes with acetone, and a final wash with PBS buffer (NaCl: 137 mM, KCl: 2.7 mM, Na2HPO4: 8 mM and KH2PO4: 1.45 mM). Each cleaning steps is realised for 10 min. Then the cleaned SiO2 surface is incubated with a 0.4% DTPS (dithiobisuccinimidyl-propionate, Sigma– Aldrich) solution in DMSO (dimethylsulphoxide) for 75 min, which creates a coupling monolayer (bifunctional crosslinker DTPS chemisorbs rapidly to quartz surface as described in [12]). A final wash with PBS buffer solution is realised. An anti-M13-PVIII monoclonal antibody (IgG, Amersham Biosciences) solution at a final concentration of 250 mg/ml is incubated for 3 h at 37 8C. Finally, the sensor is washed three times with PBS and all non-specific binding sites are blocked with a BSA (bovine serum albumin) solution for 1 h. After this first step, the Love wave delay line is in a ‘‘sensor configuration’’ and allows us to realise detection tests with the M13 bacteriophage in the second step. A solution of variable concentration (X to Y) of M13 bacteriophage (Stratagene) is applied into the test cell in a 500 ml PBS start solution (in order to satisfy the semi-infinite layer assumption) for 2 h at 37 8C. The surface of the sensor at the end of the detection is represented in Fig. 6. At the end of the experiment, it is possible to elute and numerate immobilised bacteriophages. The result can be confirmed by numerating phages left in the incubated solution after immobilisation.
After two washes with PBS (10 min each) and 10 washes with a TBS/Tween solution (Tris: 25 mM, NaCl: 136 mM, KCl: 20.7 mM and 0.1% Tween 20), bound phages are eluted with a glycine solution (0.1 M, pH 2.2) for 10 min. Secondly, the numeration of eluted M13 bacteriophages is realised using classical methods: serial phages dilutions (100 ml) are incubated with 200 ml of indicative cells E. coli (Xl1Blue:recA1, endA1, gyrA96, thi-1, hsdR17, supE44, relA1, lac[F0 proAB lacZDM15 Tn10(Tetr)], Stratagene) for 30 min. Then these dilutions are plated in a second layer of soft agar onto a Petri dish on top of 30 ml of hard agar L broth (10 g/l tryptone, 5 g/l yeast extract, 5 g/l NaCl, and 17 g/l agar). After over night incubation at 37 8C the plaques of lyses are counted: the phage concentration in PFU/ml (plaque-forming unit) is so determined. By using Love wave devices with the test bench it is possible to follow the all immunoreaction protocol in real time by a study of the sensor frequency response.
5. Experimental results 5.1. Real time sensor responses 5.1.1. Sensor response to sensitive coating deposition In Fig. 7, we can see a typical real time response of the oscillator during the AM13 binding procedure, after all preliminary surface preparations. A base line is obtained with a PBS buffer as start solution placed in the cell, with a thickness higher than several decay lengths (cf. Section 2.1.3), and so a volume sufficient to assure semi-infinite liquid assumption. When the antibody in solution is added, the frequency decreases first very quickly, this has been attributed to the instantaneous variation of the liquid physical properties [5], as we have seen that the device is very sensitive to some of these parameters, like the viscosity (cf. Fig. 4). Then can be seen the signal associated to the sensitive layer deposition, with a more slow frequency decrease. The frequency shift is near 24 kHz in this case, after 3 h immobilisation. The slight saturation of the curve is due to the fact that all accessible area are engaged. Several similar experiments showed a good reproducibility (cf. Table 1). The differences observed between
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Fig. 7. Typical acoustic signal during the AM13 grafting (sensitive layer). Table 1 Reproducibility of the AM13 deposition on the guiding layer, influence of the temperature Measured frequency shift (kHz)
Variation of temperature (8C)
Frequency shift corrected in temperature (kHz)
24 25.55 22.81
0.4 0.4 0.1
22.5 24.1 22.44
sensitive layer depositions can be attributed mainly to a temperature effect, which appears as a very limitant factor. Indeed, the influence of the temperature on the test device in gas medium has been determined theoretically and experimentally to be as high as 4 kHz/8C. With the thermoregulated dry bath, the maximal frequency shift observed is near 0.4 8C for 3 h, which corresponds to the accuracy of 10% on the frequency shift due to the sensitive layer binding. 5.1.2. Detection of the bacteriophage After the sensitive coating deposition, the device is kept at 4 8C over night. The delay line is then ready to be used in a sensor configuration. The detection of the bacteriophage begins after a stabilisation of the system temperature near 37 8C with a start solution of PBS for the base line. Like in the sensitive coating protocol, we can monitor in real time the detection of bacteriophage. The typical response represented in Fig. 8 is corrected in temperature in order to minimise this parameter influence. This response for the model analyte detection is very similar to that obtained for the sensitive coating deposition (cf. Fig. 7). The frequency is firstly stabilised with the start PBS solution (semi-infinite medium assumption). Then, adding M13 in solution induces an instantaneous frequency decrease followed by a slower decrease all along the M13/AM13 reaction on the sensor surface. The associated acoustic signal after 2 h is about 15 kHz for this slow frequency decrease. At the end of the experiment, the elution and titration of the captured bacteriophages allowed to count 1 1010 pfu.
Fig. 8. Typical acoustic signal of the bacteriophage M13 detection (model analyte) with temperature correction.
These first tests were done with highly concentrated bacteriophages solutions (the number of bacteriophages injected in the cell is very higher than the antibodies sensitive layer accessible sites), in order to validate the sensor feasibility. Then, several experiments with different concentrations of bacteriophages solutions were done. The final aim is to estimate the ability of the sensor to detect lower quantities of bacteriophages, and to improve our knowledge on the biosensing mechanisms, by comparing these results to each other. To study the influence of the model analyte concentration, the third parts of the sensor responses, corresponding to the slower frequency decrease, have been reported in Fig. 9. By doing so, the differences of physical properties (such as viscosity) of various concentrated liquid samples can be corrected in the first approximation, and we can compare the responses directly related to the bacteriophage immobilisation on the sensitive layer. Fig. 9 shows that the amplitude of the sensor response, represented here during the first minute of detection,
Fig. 9. Short term response of the sensor during bacteriophage immobilisation for two concentrations of bacteriophages injected.
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increases with the quantity of bacteriophages injected in the cell. Such results are promising for future studies: species could be detected and even quantified in a few minutes only, without waiting for the steady state. 5.2. Discussion 5.2.1. Estimation of the sensor sensitivity and detection level For the estimation of the sensor sensitivity and detection level, we consider the long-term response of the device shown in Fig. 8. The numerated bacteriophages 1 1010 pfu were captured by the antibody layer in all the accessible cell surface (121 mm2). Assuming that 1 pfu is equal to one bacteriophage (mass: 3 1015 g), we can estimate the mass accumulation near 250 ng mm2 (total mass: 30 mg). If we consider the frequency shift during deposition (Df ¼ 15 kHz), we can estimate the device sensitivity near 60 Hz/(ng mm2). We can note that an estimation of the device sensitivity such as defined in Eq. (11) would not be useful here. Indeed, it could not be compared directly to theoretical results, due to differences between polymeric and biological layers behaviours. In future more systematic experiments compared to each other and to theoretical calculations, should allow to obtain parameters for these biological layers, provided that they can be represented, at least in the first approximation, as continuous and isotropic layers, as assumed in the model. Undergoing tests with lower bacteriophages concentration aim at estimating the effective detection limit of the actual device. But the test bench has to be improved in order to reach the real detection limit of the acoustic sensor only (cf. Section 5.2.2). Indeed, placed in the oscillation loop the device offers a short term stability around 1 Hz/s. So, it should be possible to detect a variation of frequency due to bacteriophage deposition near 100 Hz or even less. If we consider linear the sensor response with the phage concentration it would induce a detection limit with the test device near 1.7 ng/mm2, corresponding to a mass deposition about 200 ng in all the cell. Additional experiments are now undertaken in order to validate these results. 5.2.2. Setup improvement Several ways can be investigated in order to improve the sensor sensitivity and its detection limit. A first way consists in improving the acoustic device. First of all the device based on the ST cut quartz is very sensitive to temperature (around 40 ppm/8C). This is a restrictive factor in term of sensor resolution. By using an AT cut quartz, it is possible to realise delay lines having a very low temperature coefficient [13]. Secondly, as presented in Fig. 2, the gravimetric sensitivity could be improved up to a factor 3 by reducing the wavelength, with the appropriate guiding layer thickness. This value of 20 mm for the wavelength was already realised by other authors, and can be adapted for biological detection tests [14].
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The second way of improvement concerns the setup used for experiments. For example the use of a dual channel configuration (reference and sensing lines) could allow to increase the device resolution. With this configuration, the antibody specific to the specie to detect is placed on the sensing line, and the reference line is equipped with a nonspecific (blocking) biomolecule. Then the reference line should allow to compensate non-specific immunoreactions as well as some variations of parameters like the temperature and physical properties of the liquid test (viscosity). Furthermore, in order to automate the protocol, a next step in the near future will be to realise a flow injection cell, in which the sample is prepared ahead the test. This system avoids instabilities due to liquid injection with a micropipette directly on the sensor, which also greatly limits the sensor resolution.
6. Conclusion In this paper, an analytical resolution of the Love wave propagation in a multilayered model is presented. It allows to estimate the Love wave phase velocity of a quartz substrate, SiO2 guiding layer, and sensitive coating device. Thus it is possible to study theoretically the gravimetric sensitivity in liquid medium in order to conceive powerful Love wave sensors. For applications in liquid medium, an AM13 antibody sensitive layer, and a M13 bacteriophage model analyte were used. The feasibility of detection applications in liquid medium has been demonstrated. Significant responses are obtained versus the quantity of bacteriophages injected in the cell. The oscillator loop associated with automate control allows to realise in situ real time experiments. Which is very interesting for immunologists, and can be applied for a great number of applications, in many fields like health, but also environment or food industry. Other works with the presented Love wave delay line are undertaken in order to validate the predicted detection level, and compare it, as well as the sensitivity, with those of other acoustic waves (QCM). These more systematic experiments should also allow to improve our knowledge of biological films and biodetection mechanisms involved in such sensors. Furthermore, other lines and test cells will be realised in order to improve the gravimetric sensitivity of the sensor and reduce the effect of the environmental perturbations on the Love waves propagation’s characteristics.
Acknowledgements We sincerely thank Dr. M.C. Bezian for her expert help in the Salmonella titration and Mrs. Nicole Lavigne and Mr. Serge Destor, members of the IXL technical team for their contribution on this work. This work was supported by a grant from the Conseil Re´ gional d’Aquitaine.
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References [1] S.J. Martin, G.C. Frye, A.J. Ricco, Multiple-frequency SAW devices for chemical sensing and material characterization, Transducers, San Francisco, 1991, pp. 385–388. [2] E. Uttenthaler, M. Schra¨ ml, J. Mandel, S. Drost, Ultrasensitive quartz crystal microbalance sensors for detection of M13-phages in liquids, Biosens. Bioelectron. 16 (2001) 735–743. [3] R. Dahint, R. Ros Seigel, P. Harder, F. Josse, Detection of nonspecific protein at artificial surfaces by the use of acoustic plate mode sensors, Sens. Actuator B 35–36 (1996) 497–505. [4] C. Zimmermann, D. Rebie`re, C. De´ jous, J. Pistre´ , E. Chastain, R. Planade, A Love wave gas sensor coated with functionalized polysiloxane for sensing organophosphorus compounds, Sens. Actuator B 76 (1–3) (2001) 88–94. [5] O. Tamarin, C. De´ jous, D. Rebie`re, J. Pistre´ , Simple analytical method to estimate the influence of liquids viscosity on Love wave chemical sensors, in: Proceedings of the Symposium on IEEE International Ultrasonics, Atlanta, GA, October 7–10, 2001. [6] S.J. Martin, A.J. Ricco, T.M. Niemczyk, G.C. Frye, Characterization of SH acoustic plate mode sensors, Sens. Actuator B 20 (1989) 253–268. [7] B. Jakoby, M.J. Vellekoop, Viscosity sensing using a Love-wave device, Sens. Actuator A 68 (1998) 275–281. [8] C. Mc Mullan, H. Mehta, E. Gizeli, C.R. Lowe, Modelling of the mass sensitivity of the Love wave device in the presence of a viscous liquid, J. Phys. D 33 (2000) 3053–3059. [9] Slobodnik, Acoustic Wave Handbook, Masson, Paris, 1973. [10] D.W. Van krevelen, Properties of Polymers, Elsevier, Amsterdam, 1990, p. 374. [11] C. Cazaubon, H. Le´vi, C. Bordieu, D. Rebie`re, J. Pistre´ , Syste`me multicapteurs de de´ tection, portable, utilisant la technique du feneˆ trage temporel, Revue d’E´ lectricite´ et d’E´ lectronique 3 (1999) 55–58. [12] J. Decker, K. Weinberger, E. Prohaska, S. Hauck, C. Ko¨ ßlinger, H. Wolf, A. Hengerer, Characterisation of a human pancreatic secretary trypsin inhibitor mutant binding to Legionella pneumophila as determined by a quartz crystal microbalance, J. Immunol. Meth. 233 (2000) 159–165. [13] B. Jakoby, J. Beistemeijer, M.J. Vellekoop, Temperature-compensated Love-wave sensors on quartz substrates, Sens. Actuator 82 (2000) 83–88. [14] G. Kovacs, G.W. Lubking, M.J. Vellekoop, A. Venema, Love waves for (bio)chemical sensing in liquids, in: Proceedings of the IEEE Ultrasonic Symposium, 1992, pp. 281–285.
Biographies Ollivier Tamarin received his Licence de Physique from Antilles-Guyane University, FWI, in 1997, then the Maıˆtrise de Physique et Applications,
and the Diploˆ me d’Etudes Approfondies in instrumentation and measurements from Bordeaux 1 University, respectively, in 1998 and 1999. He is now a PhD student at the IXL Microelectronics Laboratory, where he studies Love wave acoustic sensors for applications in liquid medium. Corinne De´ jous studied electronics engineering and received her Diploˆ me d’Inge´ nieur from the E´ cole Nationale Supe´ rieure d’E´ lectronique, Informatique et Radiocommunications de Bordeaux (ENSEIRB) in 1991. She received her PhD in 1994, and is now Associate Professor at Bordeaux 1 University of Science and Technology, France. Since 1991, she has been working at the IXL Microelectronics Laboratory, where her main interest is acoustic waves used in chemical sensors. Dominique Rebie`re received his Maıˆtrise d’E´ lectronique, E´ lectrotechnique, Automatique, the Diploˆ me d’E´ tudes Approfondies in electronics and his PhD from Bordeaux 1 University, France, in 1987, 1988 and 1992, respectively. He has been involved in research on surface acoustic wave sensors since 1989 at Bordeaux 1 University, IXL Microelectronics Laboratory, and is Associate Professor at Bordeaux 1 University in electronic engineering. Jacques Pistre´ received an MEng degree in electronics from Bordeaux University, France, in 1968, and a The`se d’E´ tat degree in ionic conductivity of thin films in 1979. Since joining the IXL Laboratory (formerly Laboratoire d’E´ lectronique Applique´ e), he has worked in several areas of thick-film microelectronics, including applications to microwave circuits and sensors. He was sent on secondment for 18 months to the French company Thomson (now Thales), working in the division of Radars, Countermeasures, Missiles. In 1987, he returned to the IXL Laboratory where he is in charge of the sensors and microsystems group. This group is mainly involved in chemical sensors, using either elastic waves in solids, resonant microstructures or microwave devices. He is Professor at ENSEIRB, a French ‘‘Grande E´ cole’’, where he teaches electronic systems. Sylvie Comeau received his PhD in 1992 from Bordeaux 2 University. She is presently working as a research engineer in Pr. Bezian group on the Plantibody project. Daniel Moynet received his The`se d’E´ tat degree in genetics in 1984 from Toulouse University and is now an Assistant Professor at Bordeaux 2 University. After working for 15 years on the HTLV-I retrovirus variability he joined the immunology laboratory and is working on recombinant antibodies and on phage display. Jean Bezian received his MD from Bordeaux University in 1960. He is Professor of immunology since 1982 and was in charge of the Immunological Laboratory at the Bordeaux hospital. His main interest was on monoclonal antibodies and recombinant antibodies. He was one of the founder of Sorebio (now part of Ares-Serono group). His group is now involved in expressing human antibodies in plants.
Study of acoustic Love wave devices for real time bacteriophage detection Ollivier Tamarina, Corinne De´jousa,*, Dominique Rebie`rea, Jacques Pistre´a, Sylvie Comeaub, Daniel Moynetc, Jean Bezianc a
Laboratoire IXL, CNRS UMR5818, ENSEIRB, Univ. Bordeaux 1, 351 Cours de la Libe´ration, F-33405 Talence Cedex, France b Cellule Adera/Bordeaux 2, Bordeaux, France c Laboratoire d’Immunologie Mole´culaire, Univ. Victor Segalen Bordeaux 2, 146 Rue Leo Saignat, F-33076 Bordeaux, France
Abstract Acoustic wave devices have shown their good potentialities for real time monitoring of immunoreactions. Different acoustic wave devices (BAW, SHAPM, Love waves) were described for applications in liquid medium. Love wave delay line structures (ST cut quartz substrate with interdigital transducers, SiO2 guiding layer) present several advantages, in particular, the pure shear horizontal polarisation adapted to liquid medium, and its very high sensitivity related to the wave confining in the thin guiding layer. In this paper an analytical method based on multilayer propagating structure is first presented: it allows us to estimate the Love wave phase velocity and then the mass loading effect sensitivity. A few theoretical results are exposed; they show that this theoretical analysis can allow to optimise physical parameters in order to conceive powerful devices for detection applications in liquid medium. As a model for virus or bacteria detection in liquids (drinking or bathing water, food, etc.), we design a model using M13 bacteriophage. The first step is the anti-M13 antibody binding. By using a Labwindows CVI software, we can monitor in real time the graft of the anti-M13 antibody sensitive coating, as well as the detection of the M13 bacteriophages. Experimental results are exposed, analysed and discussed. Love waves sensors appear to be a powerful approach for immunodetection, as theoretically predicted. # 2003 Elsevier Science B.V. All rights reserved. Keywords: Biosensors; Immunodetection; Love wave; Gravimetric sensitivity
1. Introduction The scope of acoustic (bio)sensors receive a great deal of attention for a few years as they can realise real time and in situ experiments in a large domain (e.g. medical applications, food industries or environment). A lot of studies concern detection in gas medium using mainly SAW devices [1]. Applications in liquid medium need specific acoustic polarisation as pure shear horizontal (SH) waves. Thus detection tests in liquid medium with quartz crystal microbalance (QCM) devices [2], and SH acoustic plate mode (SHAPM) devices [3] were presented, and allowed to validate real time experiments. Our aim is to develop powerful devices in terms of mass loading effect sensitivity (gravimetric sensitivity). Love wave devices, as the wave energy is confined in a thin *
Corresponding author. Tel.: þ33-5-56-84-2848; fax: þ33-5-56-37-1345. E-mail address: [email protected] (C. De´jous).
guiding layer (a few micrometers), present a high sensitivity compared to other acoustic devices [4], moreover the SH polarisation allows to work in liquid medium. The realisation of experimental Love wave delay lines needs a very fine technology, as the interdigital transducers which generate the wave are periodical long fingers (10 mm width, separated by 10 mm), without any electrical short. In order to minimise the technological realisations and conceive best adapted devices before manufacturing, an analytical method allowing to express a real dispersion relation linking the Love wave phase velocity to the other parameters of the device is presented. A second step allows to model the device gravimetric sensitivity. Thus we can estimate the liquid and device physical parameters influence on the Love wave sensor sensitivity, and so define optimised parameters. Then, experimental tests, using one of the theoretically defined devices, aim at validating the feasibility of biodetection using such a Love wave sensor. For that, an immunological model is proposed. An anti-M13 antibody sensitive
0925-4005/03/$ – see front matter # 2003 Elsevier Science B.V. All rights reserved. doi:10.1016/S0925-4005(03)00106-0
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layer is grafted on the Love wave delay line surface. As model analyte, we use the associated M13 bacteriophage. Results are analysed and discussed in order to conclude about the great interest of Love wave delay lines for detection in liquid medium.
2. Theoretical multilayered model Love wave (bio)chemical sensors, consist in the guided-SH acoustic delay line associated with a (bio)chemical layer sensitive and selective to the specie to detect. Previous theoretical works on detection tests in gas medium [4], and on the influence of a Newtonian liquid medium [5] on the propagation characteristics of the Love wave allow to estimate the phase velocity versus mechanical properties of the system. Here, we must take into account other physical properties due to the sensitive coating. This induces other interfaces like guiding layer/sensitive coating, sensitive coating/liquid. Changes due to the interaction between the overlay and the acoustic waves induce wave velocity variations. In many applications for gas detection, the most significant and best controlled effect is the wave velocity perturbation due to mass loading. Our first aim was to develop a model of the Love wave structure allowing an analytical resolution in order to predict its mass sensitivity, when working in a liquid medium. Love waves propagate in layered structures (cf. Fig. 1) consisting of a piezoelectric substrate (e.g. quartz) with an isotropic guiding layer (e.g. SiO2), the isotropic sensitive layer and the isotropic Newtonian liquid in our application. The analytical resolution consists in the determination of the SH acoustic wave propagation in the x1 direction in each volumic material. Then, a dispersion relation linking the physical parameters of the device to the Love wave velocity is deducted from the boundary conditions at each interfaces and at infinities of the multilayered structure. 2.1. Bulk SH waves In this part we will express the transverse displacements in each volumic material of the multilayered structure. As in
Fig. 1, each layer is characterised by the material physical constants (elastic modulus: Cijkl,s,h,b,L, density: rs,h,b,L, viscosity: Z for the liquid). The indices s, h, b, L characterise the physical constants in the substrate, the guiding layer, the sensitive coating and the liquid, respectively. The layers thicknesses are h and b for, respectively, the guiding layer and the sensitive coating. The general expression of the pure SH wave propagating along the acoustic path (between the interdigital transducers) is expressed as U2a ðx1 ; x3 ; tÞ ¼ U2a ðx3 Þ exp½jðot kx1 Þ where U2a ðx1 ; x3 ; tÞ is the acoustic SH particle motion, in the x2 direction, U2a ðx3 Þ the transverse acoustic displacement in the x3 direction (the expression of U2a ðx3 Þ depends on the medium of propagation represented by a), and exp½jðot kx1 Þ represents the wave propagation along the x1 direction, where k ¼ 2p=l (l the wavelength) is the wave number. The pulsation o of the wave is defined as: o ¼ l=Vp , where VP is the Love wave phase velocity. 2.1.1. The semi-infinite quartz substrate The weak electromechanical coupling coefficient ðK 2 0:1%Þ of the quartz, used for the substrate in order to generate pure SH waves, allows to neglect the piezoelectricity of this material in the first approximation. With this assumption, only the mechanical behaviour of the system (transverse acoustic displacement, and normal stress) has to be considered. In order to express the transverse acoustic displacement we use the Hooke’s and Newton’s laws written, respectively, as follows: Tij ¼ Cijkm
@Um ; @xk
i; j; k; m ¼ 1; 2; 3;
@Tij @ 2 Ui ¼r 2 @xj @t
Using these expressions, in the case of transverse horizontal waves, we obtain the equation of motion in elastic nonisotropic solids [4] C44s
@ 2 U2s @U2s 2jkC46s þ ðro2 k2 C66s ÞU2s ¼ 0 @x3 @x23
Fig. 1. Basic configuration of the Love wave delay line with a sensitive coating (sensor configuration).
(1)
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By solving the above equation, we can write the expression of the transverse acoustic displacement in the quartz substrate as follows: kC46s U2s ðx1 ; x3 ; tÞ ¼ U0s exp j þ kws x3 C44s exp½jðot kx1 Þ;
x3 < 0
(2)
where 2.1.2. The sensitive coating and the guiding layer Using Eq. (1) in the case of isotropic solids (C44a ¼ C66a ¼ ma and C46a ¼ 0), we can write the expression of the transverse acoustic displacement in both guiding layers (Eq. (3)) and sensitive coating (Eq. (4)) media: U2h ðx1 ; x3 ; tÞ ¼ U0h cosðkwh x3 þ jÞ exp½jðot kx1 Þ; 0 x3 h
(3)
qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi where wh ¼ rh Vp2 =mh and j is a constant. U2b ðx1 ; x3 ; tÞ ¼ U0b cosðkwb x3 þ j0 Þ exp½jðot kx1 Þ; h x3 h þ b ¼ L qffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi where wb ¼ rb Vp2 =mb and j0 is a constant.
2.1.3. The semi-infinite liquid medium In the liquid medium, we express the SH particle motion U2 ðx1 ; x3 ; tÞ with the Navier–Stokes law. Furthermore, we suppose a perfect horizontal system, then the pressure and the gravity can be neglected: @ 2 U2 @ @ 2 U2 @ 2 U2 rL ¼Z þ (5) @t @x23 @t2 @x21 For the resolution of Eq. (5) we suppose the x1, x3, and t variables independent and we express U2 ðx1 ; x3 ; tÞ as: U2L ðx1 ; x3 ; tÞ ¼ Sðx3 Þ expðjkx1 Þ expðjotÞ, where S(x3) represents the particle motion in the x3 direction. By including the last expression of U2L ðx1 ; x3 ; tÞ in Eq. (5), we obtain 1 @ 2 Sðx3 Þ r o ¼ j L þ k2 2 Sðx3 Þ @x3 Z
(6)
As the sound velocity for a shear mode in a solid is much larger than in a liquid, we can write k2 ! orL =Z [6]. This assumption can be verified by considering typical values of Love wave delay lines physical parameters. For example, with l ¼ 40 mm, f ¼ 100 MHz, rL ¼ 103 kg m3, Z ¼ 10 cP, we obtain k2 2 1010 m2
and
orL 6 1013 m2 Z
Then Eq. (6) can be simplified as 1 @ 2 Sðx3 Þ r o ¼j L Sðx3 Þ @x23 Z
The solution of this simplified Navier–Stokes law gives us the expression of U2L ðx3 ; tÞ. For this expression, as there is no source of energy in the liquid, we suppose a wave propagating in the x3 direction, this induces that the wave amplitude is equal to 0 for x3 ! þ1: U2L ðx1 ; x3 ; tÞ ¼ U0L exp½ð1 þ jÞkL x3 exp½jðot kx1 Þ (7) pffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi with kL ¼ orL =2Z . We can estimate the expression of the decay length defined for a diminution of the wave amplitude with a factor 1/e. This will be useful for experimental tests, in order to verify the assumption of a semi-infinite liquid layer, and so neglect the influence of a liquid volume change. Expression (7) allows us to determine an approximate value of the decay length d in the x3 direction 1 d¼ ¼ kL
(4)
277
sffiffiffiffiffiffiffiffiffi 2Z orL
This expression of d is used by several authors in the case of SH waves [6–8]. This is an approximate value of the real decay length. In order to estimate the real decay length, we would have to consider the Navier–Stokes law in two space dimensions (x1, x3), given by Eq. (5). These first steps lead to the SH wave displacement form in each material. In a second step we apply the equations of continuity at each interface of the four layered device. 2.2. Sensor: quartz substrate, SiO2 guiding layer, sensitive coating, liquid In this structure (cf. Fig. 1), we consider three interfaces: on both faces of the sensitive coating, with the liquid ðx3 ¼ LÞ and with the guiding layer ðx3 ¼ hÞ, and one between the quartz substrate and the SiO2 guiding layer. This allows to write a complex dispersion relation. 2.2.1. SiO2/liquid interface: x3 ¼ h By applying the continuity of acoustic displacement velocity between the liquid and the sensitive coating, U2L ðx3 ¼ Lþ Þ ¼ U2b ðx3 ¼ L Þ (Eqs. (4) and (7), respectively), and the continuity of the normal stress T32L ðx3 ¼ Lþ Þ ¼ T32b ðx3 ¼ L Þ (Hooke’s law), we obtain the following set of equations: joU0b cosðkwb L þ j0 Þ ¼ U0L exp½ð1 þ jÞkL L; kwb mb U0b sinðkwb L þ j0 Þ ¼ ð1 þ jÞkL ZU0L exp½ð1 þ jÞkL L The ratio of the equations of the previous system leads to
kL Vp Z 1 j0 arctan ð1 þ jÞ (8) k¼ wb L mb wb wb L
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2.2.2. Sensitive coating/SiO2 interface: x3 ¼ h, determination of j0 In the same way, we consider the continuity of acoustic displacement velocity between the sensitive coating and the guiding layer, U2b ðx3 ¼ hþ Þ ¼ U2h ðx3 ¼ h Þ (Eqs. (3) and (4), respectively), and the continuity of the normal stress T32b ðx3 ¼ hþ Þ ¼ T32h ðx3 ¼ h Þ (Hooke’s law). This leads to the following set of equations: U0h cosðkwh h þ jÞ ¼ U0b cosðkwb h þ j0 Þ; kwh mh U0h sinðk1 wh h þ jÞ ¼ kwb mb U0b sinðk1 wb h þ j0 Þ The ratio of these equations gives us the quantity j0 :
mh wh 0 tanðkwh h þ jÞ kwb h j ¼ arctan mb wb
Eq. (10) allows to estimate the Love wave phase velocity (Vp) in a particular structure with a sensitive layer and in a Newtonian viscous liquid, provided that we know all material parameters. Note that if we consider a system with Z ¼ 0, we obtain dispersion relations defined in gas medium detection applications [4]. Furthermore, compared to a complete numerical resolution, the analytical approach presents two main advantages. First of all, the Love wave phase velocity can so be calculated quite easily in only a fraction of a second, instead of several minutes with a numerical approach. And the analytical expression allows to visualise physically how the materials properties are involved in the wave propagation.
3. Theoretical results
In this expression and the following ones, all ‘‘arctan’’ functions are modulo p. 2.2.3. Quartz/SiO2 interface: x3 ¼ 0, determination of j Continuities of the acoustic displacement U2b ðx3 ¼ 0þ Þ ¼ U2h ðx3 ¼ 0 Þ (Eqs. (2) and (3)) and of the normal stress T32b ðx3 ¼ 0þ Þ ¼ T32h ðx3 ¼ 0 Þ give the following set of equations: U0s ¼ U0h cosðjÞ; U0h kC44s ws cosðjÞ ¼ U0h kC44h wh sinðjÞ Then we can write the quantity j: C44s ws j ¼ arctan C44h wh By including the quantities j and j0 in Eq. (6), after simplifications, we obtain
k L Vp Z kwb b ¼ arctan ð1 þ jÞ C44b wb
C44h wh C44s ws arctan tan kwh h arctan C44b wb C44h wh (9) with b ¼ L h. In order to write the dispersion relation, we identify the real part of the second member of Eq. (7) to kwbb, and we obtain the following dispersion relation:
C44h wh C44s ws kwb b ¼ arctan tan kwh h arctan C44b wb C44h wh 1 k L Vp Z arctan (10) 2 C44b wb The imaginary part of the second member is related to a contribution of the viscous damping of the wave. In order to take into account this part, we should have considered in this case a complex wave number k ¼ k þ jg instead of k; and in this case the analytical approach becomes very difficult or even impossible.
For detection application, among all possible detection mechanisms, one of the main effects induced is the sensitive layer mass increase when target molecules are sorbed on it. So, it is interesting to use previous studies in order to estimate the gravimetric sensitivity (i.e. sensitivity to mass loading effect) of the sensor when it is in contact with a liquid medium as a function of the device parameters, and of the liquid parameters and more particularly its viscosity. First of all, is allows to compare several acoustic propagations in terms of powerfulness for (bio)detection tests. And secondly it allows to improve delay lines conception in a very short time. To model this effect, up to now for gaseous application, we consider a little increase Drb of the material density, which induces a phase velocity shift DVP during the detection. Then the velocity sensitivity to the mass loading effect is defined by the following expression: Smv ¼
DVP 1 VP0 Drb b
(11)
where VP ¼ VP VP0 , VP0 the Love wave velocity with a sensitive layer density r ¼ rb, and VP the Love wave velocity with a sensitive layer density r ¼ rb þ Drb. In order to calculate this gravimetric sensitivity we define a Love wave test device with the following physical characteristics: a quartz substrate, a wavelength l ¼ 52 mm, a SiO2 guiding layer thickness of 6 mm (mh ¼ 3:12 1010 N/ m2, rh ¼ 2203 kg/m3 [9]), a PMMA sensitive coating layer thickness of 100 nm (mb ¼ 0:11 1010 N/m2, rb ¼ 1109 kg/m3) [10]. The PMMA was chosen in a first step as sensitive material because it was already used in previous works in gas medium [4], and so its physical parameters are well known. 3.1. Influence of the delay line parameters on the gravimetric sensitivity In order to define well-adapted devices for detection application, we investigate first the influence of the delay line parameters (wavelength, SiO2 guiding layer thickness) on the gravimetric sensitivity.
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Fig. 2. Influence of h/l on the gravimetric sensitivity in a Newtonian viscous liquid medium (Z ¼ 3 cP, rL ¼ 103 kg/m3).
Fig. 4. Influence of Newtonian liquid viscosity on Love wave phase velocity (test device [5]).
Results reported in Fig. 2 show that it is possible to increase the gravimetric sensitivity up to a factor 3 if we reduce the wavelength from 52 to 20 mm. The maximum value of gravimetric sensitivity is always obtained with an appropriate guiding layer (h=l 0:17 in this case). Similar results were also shown theoretically and experimentally for Love wave devices used for detection application in gas medium [4].
important parameter, provided that it is low enough to allow wave propagation. On the other hand, previous studies using aqueous solution of glycerol [5] showed that the viscosity had a great influence on the wave phase velocity, as represented in Fig. 4. And so, any slight viscosity change during experimental tests will induce a great velocity shift. Finally, this theoretical study shows that it is possible to define Love wave devices offering a particularly high gravimetric sensitivity. Best parameters so defined for detection in a liquid medium are very closed to those obtained in gas applications. This is partially related to the choice of a typical polymeric material as sensitive layer, as the acoustic properties of biological films are unknown yet.
3.2. Influence of the liquids mechanical parameters on the gravimetric sensitivity Using Eq. (11) again, it is also interesting to estimate the gravimetric sensitivity of the Love wave test device in a Newtonian viscous liquid medium as a function of the liquid viscosity Z. We consider only weak values of viscosities in order to avoid excessive loss of acoustic energy due to high viscous damping [7]. In Fig. 3 the results with three viscosities from 1 to 7 cP are shown. These curves show that an increase of the liquid viscosity only slightly increases the device gravimetric sensitivity. So, the liquid test viscosity should not be an
Fig. 3. Influence of the liquid viscosity on the gravimetric sensitivity (rL ¼ 103 kg/m3).
4. Experimental setup 4.1. Love wave device For detection tests, Love wave delay lines such as that previously defined in part III were realised. The technological realisation of this delay line is very important in order to obtain good sensing properties. A careful attention must concern, the engraving of the IDTs, and the deposition of the guiding layer. In our application, 50 pairs of aluminium split fingers with a wavelength l ¼ 52 mm are engraved on the quartz substrate. Split fingers allow to attenuate reflection when acoustic wave arrive under the transducers, and so to greatly attenuate the triple transit echo. The aperture, and the centre to centre distance between the IDTs are, respectively, 2.6 and 6.5 mm. This realisation needs a very fine technology in order to obtain a so periodical structure (long fingers, width l/8, separated by l/8, without any electrical short). The SiO2 guiding layer is realised on the substrate and the IDTs using PECVD (plasma enhanced chemical vapour deposition), which allows to obtain up to a few microns of SiO2 (6 mm for the test device) at low tem-
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Fig. 5. Direct frequency response of the Love wave delay line (S21).
peratures (the quartz must be kept under its phase transition temperature). These characteristics allowed to obtain transmission losses near 17 dB at the centre frequency, about 88 MHz (cf. Fig. 5). The very linear phase allows to realise an oscillating configuration with a very high short term stability. The involved technological steps are used in standard planar microelectronics processes, and so constitute an advantage compared to QCM devices. Furthermore, these technologies greatly get from the huge development of delay lines and resonators for telecommunications applications. 4.2. Test bench For operation in liquid medium, a specific test bench is used. Firstly, the Love wave delay line (500 nm thick) is glued with a silicon seal and gold wire-bonded on a printed circuit board (PCB). It becomes then an interchangeable device which can be placed in and removed from an oscillator loop. In order to measure the sensor temperature an AD590 temperature sensor placed between the delay line and the PCB is used. Secondly, a Teflon cell with a rectangular cavity (121 mm2) localised on the acoustic path of the delay line is placed on the sensor with a Teflon seal.
Finally, this system is placed in a thermoregulated dry bath. For all experiments, a single delay line is used. In order to estimate the perturbation of the Love wave characteristics, we used the delay line mounted in an oscillator configuration, and/or connected to a network analyser (Anritsu Scorpion MS4623B). The measurement of the oscillation frequency allows to obtain the highest resolution on any wave phase velocity variation. Moreover, this configuration for a complete system can be portable as demonstrated by other works [11]. Here, real time experiments are realised by recording the frequency and temperature as a function of the time with a Labwindows CVI software. Typically, with our oscillation loop, we can obtain a short term stability of 1 Hz/s. When more parameters are needed, the delay line can be characterised with a network analyser (Fig. 5). Then both wave phase velocity and transmission losses can be measured, as a function of frequency. 4.3. Biological model 4.3.1. Choice of antibody and bacteriophage immunoreaction model The M13 bacteriophage (bacteria infecting virus) was chosen as the target particle to detect. Indeed, it is quite easy to numerate the phages by using their property to infect ‘‘male’’ E. coli and to form plagues on a bacterial lawn.
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Fig. 6. Sensor surface after the detection application.
Several well-characterised anti-M13 monoclonal antibodies directed against the major proteins constituting the M13 are commercially available. And the M13 size is comparable to many pathogen viruses (900 nm long and 9 nm in diameter). These reasons justified the choice of this bacteriophage, as an immunological model for future developments. 4.3.2. Immunoreaction procedure This procedure consists of two main steps: the immobilisation of a biological layer, and of the selected bacteriophage (detection). The first step is the immobilisation of the antibody antiM13. For that, the sensor surface is cleaned: H2SO4 (4 N), three washes with H2O(d), two washes with 95% ethanol, two washes with acetone, and a final wash with PBS buffer (NaCl: 137 mM, KCl: 2.7 mM, Na2HPO4: 8 mM and KH2PO4: 1.45 mM). Each cleaning steps is realised for 10 min. Then the cleaned SiO2 surface is incubated with a 0.4% DTPS (dithiobisuccinimidyl-propionate, Sigma– Aldrich) solution in DMSO (dimethylsulphoxide) for 75 min, which creates a coupling monolayer (bifunctional crosslinker DTPS chemisorbs rapidly to quartz surface as described in [12]). A final wash with PBS buffer solution is realised. An anti-M13-PVIII monoclonal antibody (IgG, Amersham Biosciences) solution at a final concentration of 250 mg/ml is incubated for 3 h at 37 8C. Finally, the sensor is washed three times with PBS and all non-specific binding sites are blocked with a BSA (bovine serum albumin) solution for 1 h. After this first step, the Love wave delay line is in a ‘‘sensor configuration’’ and allows us to realise detection tests with the M13 bacteriophage in the second step. A solution of variable concentration (X to Y) of M13 bacteriophage (Stratagene) is applied into the test cell in a 500 ml PBS start solution (in order to satisfy the semi-infinite layer assumption) for 2 h at 37 8C. The surface of the sensor at the end of the detection is represented in Fig. 6. At the end of the experiment, it is possible to elute and numerate immobilised bacteriophages. The result can be confirmed by numerating phages left in the incubated solution after immobilisation.
After two washes with PBS (10 min each) and 10 washes with a TBS/Tween solution (Tris: 25 mM, NaCl: 136 mM, KCl: 20.7 mM and 0.1% Tween 20), bound phages are eluted with a glycine solution (0.1 M, pH 2.2) for 10 min. Secondly, the numeration of eluted M13 bacteriophages is realised using classical methods: serial phages dilutions (100 ml) are incubated with 200 ml of indicative cells E. coli (Xl1Blue:recA1, endA1, gyrA96, thi-1, hsdR17, supE44, relA1, lac[F0 proAB lacZDM15 Tn10(Tetr)], Stratagene) for 30 min. Then these dilutions are plated in a second layer of soft agar onto a Petri dish on top of 30 ml of hard agar L broth (10 g/l tryptone, 5 g/l yeast extract, 5 g/l NaCl, and 17 g/l agar). After over night incubation at 37 8C the plaques of lyses are counted: the phage concentration in PFU/ml (plaque-forming unit) is so determined. By using Love wave devices with the test bench it is possible to follow the all immunoreaction protocol in real time by a study of the sensor frequency response.
5. Experimental results 5.1. Real time sensor responses 5.1.1. Sensor response to sensitive coating deposition In Fig. 7, we can see a typical real time response of the oscillator during the AM13 binding procedure, after all preliminary surface preparations. A base line is obtained with a PBS buffer as start solution placed in the cell, with a thickness higher than several decay lengths (cf. Section 2.1.3), and so a volume sufficient to assure semi-infinite liquid assumption. When the antibody in solution is added, the frequency decreases first very quickly, this has been attributed to the instantaneous variation of the liquid physical properties [5], as we have seen that the device is very sensitive to some of these parameters, like the viscosity (cf. Fig. 4). Then can be seen the signal associated to the sensitive layer deposition, with a more slow frequency decrease. The frequency shift is near 24 kHz in this case, after 3 h immobilisation. The slight saturation of the curve is due to the fact that all accessible area are engaged. Several similar experiments showed a good reproducibility (cf. Table 1). The differences observed between
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Fig. 7. Typical acoustic signal during the AM13 grafting (sensitive layer). Table 1 Reproducibility of the AM13 deposition on the guiding layer, influence of the temperature Measured frequency shift (kHz)
Variation of temperature (8C)
Frequency shift corrected in temperature (kHz)
24 25.55 22.81
0.4 0.4 0.1
22.5 24.1 22.44
sensitive layer depositions can be attributed mainly to a temperature effect, which appears as a very limitant factor. Indeed, the influence of the temperature on the test device in gas medium has been determined theoretically and experimentally to be as high as 4 kHz/8C. With the thermoregulated dry bath, the maximal frequency shift observed is near 0.4 8C for 3 h, which corresponds to the accuracy of 10% on the frequency shift due to the sensitive layer binding. 5.1.2. Detection of the bacteriophage After the sensitive coating deposition, the device is kept at 4 8C over night. The delay line is then ready to be used in a sensor configuration. The detection of the bacteriophage begins after a stabilisation of the system temperature near 37 8C with a start solution of PBS for the base line. Like in the sensitive coating protocol, we can monitor in real time the detection of bacteriophage. The typical response represented in Fig. 8 is corrected in temperature in order to minimise this parameter influence. This response for the model analyte detection is very similar to that obtained for the sensitive coating deposition (cf. Fig. 7). The frequency is firstly stabilised with the start PBS solution (semi-infinite medium assumption). Then, adding M13 in solution induces an instantaneous frequency decrease followed by a slower decrease all along the M13/AM13 reaction on the sensor surface. The associated acoustic signal after 2 h is about 15 kHz for this slow frequency decrease. At the end of the experiment, the elution and titration of the captured bacteriophages allowed to count 1 1010 pfu.
Fig. 8. Typical acoustic signal of the bacteriophage M13 detection (model analyte) with temperature correction.
These first tests were done with highly concentrated bacteriophages solutions (the number of bacteriophages injected in the cell is very higher than the antibodies sensitive layer accessible sites), in order to validate the sensor feasibility. Then, several experiments with different concentrations of bacteriophages solutions were done. The final aim is to estimate the ability of the sensor to detect lower quantities of bacteriophages, and to improve our knowledge on the biosensing mechanisms, by comparing these results to each other. To study the influence of the model analyte concentration, the third parts of the sensor responses, corresponding to the slower frequency decrease, have been reported in Fig. 9. By doing so, the differences of physical properties (such as viscosity) of various concentrated liquid samples can be corrected in the first approximation, and we can compare the responses directly related to the bacteriophage immobilisation on the sensitive layer. Fig. 9 shows that the amplitude of the sensor response, represented here during the first minute of detection,
Fig. 9. Short term response of the sensor during bacteriophage immobilisation for two concentrations of bacteriophages injected.
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increases with the quantity of bacteriophages injected in the cell. Such results are promising for future studies: species could be detected and even quantified in a few minutes only, without waiting for the steady state. 5.2. Discussion 5.2.1. Estimation of the sensor sensitivity and detection level For the estimation of the sensor sensitivity and detection level, we consider the long-term response of the device shown in Fig. 8. The numerated bacteriophages 1 1010 pfu were captured by the antibody layer in all the accessible cell surface (121 mm2). Assuming that 1 pfu is equal to one bacteriophage (mass: 3 1015 g), we can estimate the mass accumulation near 250 ng mm2 (total mass: 30 mg). If we consider the frequency shift during deposition (Df ¼ 15 kHz), we can estimate the device sensitivity near 60 Hz/(ng mm2). We can note that an estimation of the device sensitivity such as defined in Eq. (11) would not be useful here. Indeed, it could not be compared directly to theoretical results, due to differences between polymeric and biological layers behaviours. In future more systematic experiments compared to each other and to theoretical calculations, should allow to obtain parameters for these biological layers, provided that they can be represented, at least in the first approximation, as continuous and isotropic layers, as assumed in the model. Undergoing tests with lower bacteriophages concentration aim at estimating the effective detection limit of the actual device. But the test bench has to be improved in order to reach the real detection limit of the acoustic sensor only (cf. Section 5.2.2). Indeed, placed in the oscillation loop the device offers a short term stability around 1 Hz/s. So, it should be possible to detect a variation of frequency due to bacteriophage deposition near 100 Hz or even less. If we consider linear the sensor response with the phage concentration it would induce a detection limit with the test device near 1.7 ng/mm2, corresponding to a mass deposition about 200 ng in all the cell. Additional experiments are now undertaken in order to validate these results. 5.2.2. Setup improvement Several ways can be investigated in order to improve the sensor sensitivity and its detection limit. A first way consists in improving the acoustic device. First of all the device based on the ST cut quartz is very sensitive to temperature (around 40 ppm/8C). This is a restrictive factor in term of sensor resolution. By using an AT cut quartz, it is possible to realise delay lines having a very low temperature coefficient [13]. Secondly, as presented in Fig. 2, the gravimetric sensitivity could be improved up to a factor 3 by reducing the wavelength, with the appropriate guiding layer thickness. This value of 20 mm for the wavelength was already realised by other authors, and can be adapted for biological detection tests [14].
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The second way of improvement concerns the setup used for experiments. For example the use of a dual channel configuration (reference and sensing lines) could allow to increase the device resolution. With this configuration, the antibody specific to the specie to detect is placed on the sensing line, and the reference line is equipped with a nonspecific (blocking) biomolecule. Then the reference line should allow to compensate non-specific immunoreactions as well as some variations of parameters like the temperature and physical properties of the liquid test (viscosity). Furthermore, in order to automate the protocol, a next step in the near future will be to realise a flow injection cell, in which the sample is prepared ahead the test. This system avoids instabilities due to liquid injection with a micropipette directly on the sensor, which also greatly limits the sensor resolution.
6. Conclusion In this paper, an analytical resolution of the Love wave propagation in a multilayered model is presented. It allows to estimate the Love wave phase velocity of a quartz substrate, SiO2 guiding layer, and sensitive coating device. Thus it is possible to study theoretically the gravimetric sensitivity in liquid medium in order to conceive powerful Love wave sensors. For applications in liquid medium, an AM13 antibody sensitive layer, and a M13 bacteriophage model analyte were used. The feasibility of detection applications in liquid medium has been demonstrated. Significant responses are obtained versus the quantity of bacteriophages injected in the cell. The oscillator loop associated with automate control allows to realise in situ real time experiments. Which is very interesting for immunologists, and can be applied for a great number of applications, in many fields like health, but also environment or food industry. Other works with the presented Love wave delay line are undertaken in order to validate the predicted detection level, and compare it, as well as the sensitivity, with those of other acoustic waves (QCM). These more systematic experiments should also allow to improve our knowledge of biological films and biodetection mechanisms involved in such sensors. Furthermore, other lines and test cells will be realised in order to improve the gravimetric sensitivity of the sensor and reduce the effect of the environmental perturbations on the Love waves propagation’s characteristics.
Acknowledgements We sincerely thank Dr. M.C. Bezian for her expert help in the Salmonella titration and Mrs. Nicole Lavigne and Mr. Serge Destor, members of the IXL technical team for their contribution on this work. This work was supported by a grant from the Conseil Re´ gional d’Aquitaine.
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Biographies Ollivier Tamarin received his Licence de Physique from Antilles-Guyane University, FWI, in 1997, then the Maıˆtrise de Physique et Applications,
and the Diploˆ me d’Etudes Approfondies in instrumentation and measurements from Bordeaux 1 University, respectively, in 1998 and 1999. He is now a PhD student at the IXL Microelectronics Laboratory, where he studies Love wave acoustic sensors for applications in liquid medium. Corinne De´ jous studied electronics engineering and received her Diploˆ me d’Inge´ nieur from the E´ cole Nationale Supe´ rieure d’E´ lectronique, Informatique et Radiocommunications de Bordeaux (ENSEIRB) in 1991. She received her PhD in 1994, and is now Associate Professor at Bordeaux 1 University of Science and Technology, France. Since 1991, she has been working at the IXL Microelectronics Laboratory, where her main interest is acoustic waves used in chemical sensors. Dominique Rebie`re received his Maıˆtrise d’E´ lectronique, E´ lectrotechnique, Automatique, the Diploˆ me d’E´ tudes Approfondies in electronics and his PhD from Bordeaux 1 University, France, in 1987, 1988 and 1992, respectively. He has been involved in research on surface acoustic wave sensors since 1989 at Bordeaux 1 University, IXL Microelectronics Laboratory, and is Associate Professor at Bordeaux 1 University in electronic engineering. Jacques Pistre´ received an MEng degree in electronics from Bordeaux University, France, in 1968, and a The`se d’E´ tat degree in ionic conductivity of thin films in 1979. Since joining the IXL Laboratory (formerly Laboratoire d’E´ lectronique Applique´ e), he has worked in several areas of thick-film microelectronics, including applications to microwave circuits and sensors. He was sent on secondment for 18 months to the French company Thomson (now Thales), working in the division of Radars, Countermeasures, Missiles. In 1987, he returned to the IXL Laboratory where he is in charge of the sensors and microsystems group. This group is mainly involved in chemical sensors, using either elastic waves in solids, resonant microstructures or microwave devices. He is Professor at ENSEIRB, a French ‘‘Grande E´ cole’’, where he teaches electronic systems. Sylvie Comeau received his PhD in 1992 from Bordeaux 2 University. She is presently working as a research engineer in Pr. Bezian group on the Plantibody project. Daniel Moynet received his The`se d’E´ tat degree in genetics in 1984 from Toulouse University and is now an Assistant Professor at Bordeaux 2 University. After working for 15 years on the HTLV-I retrovirus variability he joined the immunology laboratory and is working on recombinant antibodies and on phage display. Jean Bezian received his MD from Bordeaux University in 1960. He is Professor of immunology since 1982 and was in charge of the Immunological Laboratory at the Bordeaux hospital. His main interest was on monoclonal antibodies and recombinant antibodies. He was one of the founder of Sorebio (now part of Ares-Serono group). His group is now involved in expressing human antibodies in plants.