Biosensors& Bioekctronics 6 (1991) 595-607
An evanescent fluorescence biosensor using ion-exchanged buried waveguides and the enhancement of peak fluorescence* Y. Zhou, P. J. R. Laybourn, J. V. Magill 81 R. M. De La Rue Department of Electronics and Electrical Engineering, University of Glasgow, Glasgow, G12 SQQ,
UK
(Received 24 May 1990,received in revised form 19 November 1990; accepted 4 December 1990)
Abstract: The principle of an optical molecular sensor using ion-exchanged buried planarwaveguides in glass has been demonstrated. We have shown both theoretically and experimentally that the intensity of the peak evanescent fluorescence can be increased by several orders of magnitude with the use of an index-matching material. The method of differential measurement has been used to improve the differentiation between specific and non-specific binding. We used h-IgG (human immunoglobulin G) as the immobilized antibody on the surface of the waveguide and protein A-PITC (fluorescein isothiocyanate) as the fluorescently labelled antigen or anti-antibody to be detected, and have shown that a concentration of protein A as low as 24 nM can easily be detected. Keywords: biosensor, immunosensor. ion-exchanged waveguide, total internal reflection fluorescence (TURF),tluoroimmunoassay. immobilization of proteins, antibody-antigen binding.
INTRODUCTION
Optical molecular sensors are analytical devices which convert a chemical parameter of a molecular species (e.g. concentration, intra- or inter-molecular structure, or orientation) into an optical signal (e.g. intensity, frequency, phase or polarization). As the number of molecular species which might be present in addition to the one to be detected is usually large, specificity or selectivity is definitely needed for the discrimination of low concentrations of analytes in complex matrices. To employ this property for *Paper presented at Biosensors90, Singapore,2-4 May 1990.
a wide range of assays for both antigens and antibodies, one of the most commonly used methods is to label a known antigen or antibody so that their location and quantity can be established. For example, antibodies can first be immobilized onto a solid support such as a glass microscope slide and a solution of antigens with a fluorescent label can then be applied to the immobilized antibodies. Immobilized antibodies raised against a specific antigen will selectively bind to that antigen, leaving the non-specific antibodies in the solution. After a certain time of incubation, the free antigens can be removed by a washing step. The glass slide is then illuminated by light of a suitable excitation wavelength in an
0956-5663/91/$03.50 (0 1991 Elsevier Science Publishers Ltd.
595
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optical microscope. The excitation light is filtered out before reaching the observer or the detector. Thus, only the fluorescent light from the glass slide is observed or detected and the position as well as the amount of labelled antigen can be established. A number of optical techniques have been used for molecular sensing purposes, including total internal reflection (TIR), ellipsometry, surface plasmon resonance (SPR) and light scattering. Of these techniques, however, only those which involve the use of evanescent interactions of light with molecules seem promising. The reason is that the dimension of the penetration depth of an optical evanescent wave (typically 100 nm) is approximately of the same order as that of the bimolecular complex (about ten to several tens of nanometers, depending on the size of the antigen and the antibody) and consequently the interaction of light with the molecules in the cover medium is mainly restricted to those molecules which are bound to the surface of the optical transducer. As a result, the washing step for the separation of bound molecules from unbound ones can readily be avoided. Although total internal reflection is the most often used method to generate an evanescent wave, integrated optical waveguides and optical fibres are preferred as the sensing element because their sensitivity can be increased as a result of the increase in the number of internal reflections over a reasonable length. Until now, several devices using evanescent interactions have been proposed and developed (Liedberg et al., 1983;Sutherland & Dahne, 1987; Badley et al., 1987; Cullen et al., 1988; Love ef al., 1989; Rogers et al., 1989; Nellen & Lukosz, 1990; Lukosz er al., 1990; Sloper et al., 1990).The idea of using an evanescent wave to study chemical reactions at a surface was put forward more than two decades ago (Harrick, 1967). It was then demonstrated by Rronick & Little (1973, 1975) that a wash-free immunoassay could be performed with a prism using total internal reflection fluorescence. An advance was made when Liedberg et al. (1983)published their results on the use of surface plasmon resonance (SPR) for the detection of gas and for biosensing. Later, the use of a diffraction grating as an alternative to prisms in SPR devices was reported (Cullen et al., 1988;Parry&al., 1989).Although the sensitivity of the two devices was similar, the use of a diffraction grating implied the possibility of cheap devices, as mass production of high5%
Biosensors& Bioekctronics 6 (1991) 595407
quality diffraction gratings has become a reality. The reporting of the fluorescence capillary fill device (FCFD) (Badley er al., 1987) introduced another device which enhances the concept of an inexpensive, disposable and manufacturable product. More recently, the use of diffraction gratings on planar dielectric waveguides for biosensing has been studied, and a sensitivity as high as 10s9 M has been reached (Nellen & Lukosz, 1990).In addition, investigations are also being made into fabrication of cheap planar monomode or low-mode order waveguides and the application of these waveguides to evanescent fluorescence biosensing (Sloper & Flanagan, 1988). A common problem associated with all these devices, especially the direct devices such as the SPR device (Liedberg et al., 1983; Cullen et al., 1988) and the grating coupling device (Nellen & Lukosz, 1990; Lukosz et al., 1990), is their sensitivity to non-specific interactions at the interface. In the FCFD (Badley et al., 1987),the discrimination between surface-bound fluorescence and bulk fluorescence is achieved by comparing the intensity of the fluorescent light emerging at smaller angles to the axis of the slide with that at larger angles. However, because the discrimination against fluorescence from within the fluid bulk is disrupted by the appearance of some bulk fluorescence in the smaller angle zone, the background fluorescence noise is high. A solution to this problem is the use of narrow planar waveguides so that evanescent excitation of fluorescence can be employed. Although an evanescent fluorescence immunosensor based on low-mode order planar waveguides has been reported (Sloper et al., l!WO), problems concerning its sensitivity, background noise and differentiation of specific binding from nonspecific binding still exist. In this paper, we describe the development of a planar guided-wave optical molecular sensor which attempts to overcome these problems. The device consists of a patterned planar waveguide fabricated on a glass microscope slide. A two-step ion-exchange process produces a completely buried monomode or low-mode order planar waveguide on the glass slide. Because of the complete burial of the waveguide, there is very little penetration of the evanescent wave beyond the glass surface. A patterned waveguide is then made by etching two rectangular patterns on the slide, contlning the penetration of the evanescent
Biosensors & Bioelectmnics 6
(1991)595-607
wave beyond the glass surface to within the two etched areas (see Figs l-3 later). Particular advantages of the ion-exchange process include low cost, simplicity and flexibility of fabrication. Although much effort has been spent on the fabrication of different kinds of optical waveguides using the ion-exchange technique, and a report has been made recently on the burial of K+-exchanged waveguides using electric-field assistance during the second step Na+-exchange (Milliou et al., 1989), this technique does not appear to have been applied to the development of integrated optical molecular sensors. The process developed by us for the fabrication of patterned waveguides is simple and the need for electric field assistance is eliminated. There are several advantages associated with patterned waveguides. Firstly, surface scattering of light guided by the waveguide can be greatly reduced, leading to less fluorescent noise produced by the surface-scattered light. Secondly, the penetration depth of the evanescent wave beyond the etched areas can be adjusted by appropriate etching. Thirdly, the complete burial of the waveguide eliminates the need for a patterned buffer layer on the slide to limit evanescent interactions within defined areas. Finally, by immobilizing specific proteins on one etched area and control proteins on the other, a differential measurement can be performed, and this can lead to a great improvement in the differentiation between specific binding and nonspecific binding. Another particular feature of our device is the use of an index-matching material as the cover medium, which leads to an increase of the peak evanescently excited fluorescence by several orders of magnitude.
MATERIALS AND METHODS
An evanescent jluorescence biosensor
weight 2000,75% dehydrolysed) were from Sigma Chemical Co. (St Louis, MO, USA). Fabrication of buried planar waveguides
Glass slides were cleaned by soaking successively in each of the following solutions in an ultrasonic bath for 5 min: (a) trichloroethylene; (b) methanol; (c)acetone. The slides were then rinsed with reverse osmosis (RO) water, blown dry with compressed nitrogen gas and placed over a hot plate (90°C) for 10 min. Potassium ion exchange was carried out at 490°C using a bath of molten potassium nitrate in a tube furnace. Glass slides were preheated in the furnace for 5 min to prevent cracking. The slides were then immersed in the molten potassium nitrate for 30 min, taken out of the bath, cooled to room temperature in open air and rinsed with hot water. For waveguide burying, the process was repeated under the same conditions using sodium nitrate instead of potassium nitrate. Patterning of the buried waveguide
The slides were cleaned as described in the previous section. A resist coating was made on the slide surface by spinning. After curing at 90°C for 1 h, two transparent rectangular patterns were formed on the resist coating near the centre of the slide by W light exposure and development. The etching of the patterns on the glass slide was performed by immersing the glass slides in Si02 etch (4~1 HF) with mild agitation for 2 min, rinsing with RO water and blowing dry with compressed nitrogen gas. One end of the slide (about 10 mm) was also etched by immersing it in the SiOz etch for 4 min. This was intended to facilitate the coupling of light into the waveguide when prism coupling was later used. We refer to slides treated to this stage as patterned waveguides.
Materials Protein immobilization by physical adsorption
Standard soda-lime glass microscope slides (75.5 mm X 26 mm X 1 mm) were purchased from BDH Ltd (Poole, Dorset, UK). Photoresist (145OJ) and Microposit Developer for Photoresist 14505 were supplied by Shipley Europe Ltd (Coventry, UK). H-&G (human immunoglobulin G), protein A-FITC (fluorescein isothiocyanate), BSA (bovine serum albumin), dichlorodimethylsilane (Si(CH&C12) and polyvinyl alcohol (molecular
The immobilization of proteins onto the sensor surface was based on a method described by Elwing & Stenberg (1981). Patterned waveguides were cleaned in 30% hydrogen peroxide in concentrated sulphuric acid for lo-15 min, rinsed with RO water, blown dry with compressed nitrogen gas and dried over a hot plate (90°C) for about 10 min. 597
Y: Zhou et al.
After cleaning, the slides were immersed for 30 s in a 2%solution of dichlorodimethylsilane in trichloroethylene, rinsed in trichloroethylene and blown dry with compressed nitrogen gas. This treatment generated a hydrophobic surface which would spontaneously adsorb BAS molecules and antibodies. Physical adsorption was achieved by leaving with a pipette a small amount (approximately 25 ~1) of 0.1 mg ml-’ h-IgG and O-1mg ml-’ BSA solution in 0.01 M PBS (phosphate-buffered saline) (pH 7.2) on the two etched areas respectively for 3-18 h at room temperature. To maintain a high humidity to prevent the evaporation of water, an airtight plastic wet box was used to contain the slides.
Biosensors t Bioekctronics 6 (1991) 595-607 Dimension: etched areas 1Omm x Smm separation of the two areas 5mm width of slide 26mm length of slide 76mm
mirror
chopper /
Application of the index-matching material and FlTC-labelled proteins
The slides with physically adsorbed proteins were rinsed with RO water to remove any unadsorbed proteins and blown dry with compressed nitrogen gas. Immediately after this, a small amount (about 025 ml) of 10% (w/v) polyvinyl alcohol solution was applied to the etched patterns, giving a uniform distribution of the index-matching material on the etched surfaces. The slides were always kept level so that a uniform film of the index-matching material (about 0.5 mm thick) was formed on the central area of the slides. A small equal amount of FITC-labelled protein A solution of various concentrations in 0.01 M PBS (pH 7.2) was then applied to the index-matching film of the slides, and the slides were left in open air until the water in the polyvinyl alcohol had completely evaporated (i.e. the material had dried). Alternatively, equal amounts of protein A-FITC solution and polyvinyl alcohol solution could be mixed together and then applied to the etched areas. Apparatus and evanescent fluorescence measurements
Figure 1 shows our experimental set-up for the measurement of the evanescent fluorescence. A slide was mounted on a rotation table. The 488 nm light beam from an Ar+ laser was passed through a chopper, reflected by a mirror and coupled into the waveguide through an input coupling prism. As the waveguide was completely buried, the prism was placed on the etched end of the slide and the rotation table was turned until 598
Fig I. Experimental set-up for the measurement of light intensities,
good input coupling of the excitation light was reached. Two detectors were placed at a fixed distance from the edges of the slide and polyvinyl alcohol solution was painted on the edges to reduce diffused reflection of light caused by the roughness of the edges. In front of detector 1, a bandpass filter (centre wavelength ;1, = 520 nm, bandwidth = 8.2 nm) and a long-pass filter (OG5 15, i.e. cut-off wavelength & = 515 nm) were placed to pass the fluorescent light and eliminate the light at the excitation wavelength. This detector moved along the edge of the slide and gave two maximum readouts, Z,and Z,,,where Z,is the fluorescent intensity from the specific area and Z, is the fluorescent intensity from the nonspecific area. Detector 2 was placed at the front end of the waveguide and measured the intensity of the reference guided exciting light, Z,. The chopping frequency for the excitation light was about 600 Hz and a lock-in amplifier was used to give the readouts of the measured intensities. PRINCIPLE OF OPERATION AND THEORETICAL ANALYSIS OF THE ENHANCEMENT OF EVANESCENT FLUORESCENCE Evaaescent excitation of fluorescence
When a beam of light is incident on the interface of two optically different media, it will in general
Biosensors t Bioelectronics 6 (1991) 595-607
be partially reflected and partially transmitted. However, if the light beam is incident from an optically denser medium onto an optically rarer medium and the angle of incidence 8i is greater than the critical angle defined as
e, = sin-‘(n&r,)
(1) where n, and nP are the refractive indices of the optically rarer and denser medium respectively, total internal reflection takes place. As a result, an evanescent wave is generated on the interface in the optically rarer medium. Typically, for a glass prism (n,, = 1.5) and an aqueous cover solution (4 = 1*33),0, Z 62”. A particular feature of the evanescent wave is that its amplitude and intensity decay exponentially from the interface. A planar waveguide can be used for the generation of an evanescent wave because the light guided by the waveguide undergoes total internal reflection. As shown in Fig. 2, a planar waveguide is a thin film of optically denser medium sandwiched between two optically rarer media. If n, is the refractive index of the cover medium, nB is the refractive index of the guiding layer, n, is the refractive index of the substrate, d is the thickness of the guiding layer, and ng > n,, n,, the electromagnetic lield beyond the interface in the cover medium can be expressed as E = A
exp(-6x)
forx>O
(2) where E is the amplitude of the electric field at x, A is the amplitude at the interface x = 0, & = 2n(n,fl* - nC2)“*/& h is the wavelength of the electromagnetic wave in free space, n,m is the effective refractive in&x of a mode guided by the planar waveguide (Garmire, 1985), and n, is the refractive index of the cover medium. The above expressions apply to all the guided modes, provided that different values of nemare used for different modes. The penetration depth of the evanescent wave is defined as
An evanescmt jluorescence biosensor
and typically, dp is about 100 nm for potassium ion-exchanged waveguides and aqueous cover solution, as will be shown later. In the case of total internal reflection, when the cover medium is a fluorescent dye solution, fluorescence can be excited by the evanescent wave. Furthermore, this fluorescent light emitted in an optically rarer medium can be coupled back into the optically denser medium and observed below the interface beyond the critical angle. A simple explanation of this is that as the exciting wave is evanescent, the excited fluorescent wave is also evanescent, and consequently it can be coupled back in the same way as an evanescent wave is generated. Principle of operation of the device
Figure 3 is an illustration of our device. Antibodies and control proteins such as BSA molecules are immobilized on the two etched areas of the patterned waveguide respectively and the surfaces are then covered with a film of indexmatching material (such as polyvinyl alcohol). Because of the gelatinous property of the indexmatching material, fluorescently labelled antigens or anti-antibodies in aqueous solution are able, when applied, to diffuse through the film and bind to the immobilized antibodies, ifthere is a significant degree of specificity between them. Because of the specific binding, more fluorescent molecules are localized close to the specific surface than to the non-specific one (i.e. the BSAimmobilized surface). When light of the excitation wavelength is guided in the waveguide, evanescent excitation of fluorescence takes place. As the optical power used for the excitation of
substrate
Fig. 2. Evanescent wave on a planar waveguide.
Fig. 3. Schematic illustration of the sensor device: n,. substrate or superstme tejkctive index; nF waveguide index; n,, re$uctive index of index-matching material. 599
Y: Zhou et
al.
evanescent fluorescence is only a small fraction of the total guided optical power, the evanescent exciting intensities for the two etched areas are roughly the same. However, because of the difference in the number of fluorescent molecules in the two etched areas, more evanescent fluorescence is expected from the specific area than from the non-specific area. If 1, is the fluorescence intensity from the specific area, 1, is the fluorescence intensity from the non-specific area and 1, is the reference intensity of the guided exciting light, the ratio of the evanescently excited fluorescence intensity difference to the intensity of the guided exciting light (i.e. (I, - I&) is proportional to the fractional surface coverage of the applied antigen or anti-antibody which is, in turn, dependent upon the concentration as well as the afftnity constant of the antigen or antiantibody to the immobilized antibody. Therefore, by measuringl& and I,, the concentration of the applied specific antigen or anti-antibody can be determined. The differential measurement method has the merit that when there is no specificity between the applied antigen or anti-antibody and the immobilized antibody, ideally& and& will be the same, regardless of the applied concentration, if the propagation loss of light is assumed to be negligible. As (I, - I,,)/& is equal to zero, the influence on optical output of the non-specific binding is eliminated. In practice, there is always some optical propagation loss; however, as long as this loss is taken into consideration, the differential measurement method can still be used. Fluorescence enhancement though index-matching
The angular distribution of fluorescence excited by an evanescent wave has been investigated both theoretically and experimentally by Lee et al. (1979). As shown in Fig. 4, the following assumptions are made: n, is the refractive index of the cover or sample solution, n,, is the refractive index of the hemicylindrical prism, Oi is the angle of incidence of the exciting beam of light, Oci is the critical angle pertaining to the wavelength of the exciting light, O,r is the critical angle pertaining to the wavelength of the fluorescent light, and 0s is the angle of fluorescence observation. Iflr(Oi, 0,) represents the fluorescence intensity excited at angle Oiand observed at angle 0s per 600
BiosensorsL Bioelectmnics6 (1991) 595-607
dye sample
reflected
light
1
incident
light
Fig. 4. Evanescentjluorescence.
unit incident light intensity, and zero dispersion in the relative refractive index nCP (=n,/n,) between the two wavelengths is assumed, the peak fluorescence intensity per unit incident light intensity for 0, > &i, and 0s = &r can be expressed as Mi, %d = a I W4 1 I * 9 I Wb)
(4) where a and n are the dye absorption coefftcient and the dye fluorescence efficiency respectively and are both assumed to be isotropic. T(&) and T(&) are the transmission terms for the incident intensity and for the fluorescent intensity, respectively. They are defined as the transmitted energy flux (or power) divided by the incident energy flux (or power). An important conclusion arising from Lee’s investigation is that, with evanescent excitation, the overall fluorescence intensity of a thick cover film (i.e. one satisfying the condition t > (2/dpi + 2/d,f)-‘, where t is the thickness of the film and dpi and dpf are the penetration depths of the exciting and fluorescent light respectively) of dye solution radiating into the optically denser medium is largest at an incident angle obeying 8i = &i, and, for a fixed Oi> B,i, the relative fluorescence intensity peaks at e, = ecf. The transmission term for the incident beam polarized vertically with respect to the plane of incidence can be related to the Fresnel transmission factor t, as follows (Born & Wolf, 1959): I *dpi /2
IT’(&)l* = nCPIt,(ei) I */cos ei The Fresnel transmission factor is
Itl(ei) I* =
I
ices
ei + ~~pY~sin2e.)l/2] I
(5)
I* t6)
Where the fluorescent light is concerned, its transmission term I Tl (0,) I is also given by eqn (5), with Oireplaced by 0, and nC,,replaced byits
Biosensors& Bioelectnmics6 (1991) 595407
An evanescentfluorescencebiosensor
value at the fluorescence wavelength. Corresponding equations for the horizontally polarized light have been published (Lee er al., 1979). The reason for choosing vertically polarized light here is simply that the expressions are simpler. In the case of a planar waveguide (Fig. 2), nc = n, and n,, = nB, where ns is the refractive index of the guiding layer. For a guided mode, if n,w is the effective refractive index of the guided mode (Garmire, 1985) 6, = sin-’ (n&t& and 8i > 8,-i are always valid. Taking into consideration the fact that (ncs2- sin28i)“2 is now imaginary, where nCg= nC/ng, substitution of n,, nP and 8i into eqns (6) and (5) results in 1T” (e,)
12 =
“$
1
:e$“2 c
If we restrict our observation of fluorescence to the peak value for vertical polarization, then e, = 0, = sin-’ (n&r,), where 0, is the critical angle for the fluorescent light. Replacement of 8i by 19, in eqns (5) and (6) leads to 1Tl (0,) I* = 4n,/(ni
- nz)“2
(8)
By substituting eqns (3), (7) and (8) into eqn (4), the peak fluorescence intensity per unit incident light intensity can be re-expressed in terms of n,, nB and nen as 4atjAi
Mnc,n,,neff) = y-
certainly true for our K+-exchanged waveguides (Adams, 1981). To use eqn (9) directly, the K+exchanged waveguide is treated as a step-index planar waveguide. Referring to Fig. 2, the following assumptions are made: the refractive index of the substrate, n, = l-510, the refractive index of the guiding layer, nB = 1.519, the thickness of the guiding layer, d = 1.4pm, and the refractive index of the cover medium n, lies between 1.33 and 1.51. To compare the peak evanescent fluorescence of an index-matching medium with that of an aqueous solution, we have chosen the refractive index of aqueous solutions, 1.33, as the starting point of the range of possible values for n,. As an ordinary glass microscope slide has a refractive index of 1.510 at the wavelength of sodium light (589 nm), this value is taken as the refractive index of the substrate. The reason for choosing ng = 1.519is that K+-exchange gives a maximum increase of 0009 in the refractive index of glass. Additionally, d = 1.4pm ensures that the waveguide supports a single mode within the range 1.33 < n, < l-510. Figure 5 is the calculated fluorescence enhancement as a function of n, for the TE mode. The fluorescence enhancement Et is defined as the peak fluorescence intensity at n, with respect to that at n, = 1.33, i.e. Et =
WC, ng9 nefd I&
= 1.33, nB, neft)
(10)
nC2(ng2- n,fFZ)“’ ’ (net? - n,2)‘R(nB2 - nC2)3/2
(9) As eqn (4) requires that the intensity of the light incident at the interface is normalized, the assumption is made that the optical power guided by the waveguide per unit width is constant. However, this does not mean that there is no change in the amplitude of the electromagnetic field at the interface, but rather that this change is accounted for by the transmission term T(&). As will be shown below, the increase in the fluorescence intensity is mainly caused by the increase in this field, although there is also an increase in the penetration depth of the evanescent wave. Typically, for a potassium ionexchanged planar waveguide and aqueous cover solution, ng = l-519, n, = 1.33, ner z 1.514 and consequently, If z 0.97 o&Ii. Generally speaking, a planar waveguide made by diffusion has a graded index profile, and this is
“C
Fig. 5. Enhancement of peak evanescent fluorescence. 601
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Biosensors& Bioelectmnics6 (1991) 595407
et al.
In Fig 5, the starting point is one, i.e. Er( 1.33) = 1.With increase in the refractive index of the cover sample, Er increases. Its value and rate of increase can be very high as n, approaches R,. What this means is that, when every quantity except n, is held the same, the use of a cover medium with a higher refractive index, instead of an aqueous solution as in the case of most molecular sensors, can result in a substantial increase in the intensity of the evanescently excited fluorescence. However, as will be discussed below, the increase in n, also means an increase in the penetration depth, dr. This means that more fluorescent molecules not bound to the surface will be excited. Meanwhile, there exists a value for n, beyond which the waveguide will not be able to support any mode, and therefore choosing a value for n, very close to ns is not necessarily beneficial. A problem with index-matching is that the increase of n, leads to an increase in the penetration depth of the evanescent wave, and this will reduce the differentiation of fluorescence emitted by surface-bound molecules from that emitted by those in the bulk As a comparison, the increase of the penetration depth dp is plotted as a function of n, (eqn (3)) in Fig 6. As can be seen from Figs 5 and 6, the peak evanescent fluorescence increases much faster than the penetration depth, as n, approaches the indexd,
(urn)
A2 /A’0
1.33 1.35 1.37 1.39 1.41 1.43 1.4s I.47 1.49 1.51 no
Fig. 7. Enhancement of intensi~ (AZ/A&versus%.
matching point. In the case of polyvinyl alcohol (n, = 1.509) the enhancement is about 500 whereas the penetration depth is only five times that of the value at n, = 1.33.This actually means that the enhancement is mainly brought about by the increase in the electromagnetic intensity at the interface. To show this, we include here Fig 7, which gives the change in the intensity of the electromagnetic wave at the interface (i.e. A*, eqn (2)) as a function of n, for TE modes. The curve is derived from the following expression (Marcuse, 1974):
‘j A* =
0.1 1.33 1.36 1.37 1.39 141 I
h
#
I
4
I
I
nC
Fig. 6. Penetmtbn depth dp vemus q. 602
I
I
1.43 1.46 1.47 I.49 1.51
4&&P I/3I (d + l/y + l/S)&* + S*)
(11)
where A is the electric field amplitude at the interface, P is the total optical power per unit width, K = (2n/;l)(ns* - n,m2)ln is the transverse propagation constant, y = (2n/A)(n,fi* - n,2)“* and l/y = dpe, where dpl is the penetration depth into the substrate, 6 = (2n/~)(n,r,* - nC2)lDand l/S = d,,, where d,, is the penetration depth into the cover medium, d is the thickness of the guiding layer, 1j3 I= (2n/L)n,n is the longitudinal propagation constant, o = 2nf is the angular frequency of the light, and PO is the magnetic permeability. The enhancement of the electromagnetic intensity is defined as A* at n, divided by A2 at n, = 1.33, i.e. A*/A*(n, = l-33), with P assumed
Biosensors & Bioel~nics
An evanescentjluonscence bier
6 (1991) 595-601
unchanged. It is worth pointing out that lo~~~rnic coordinates have been used in Figs 5-7. When high sensitivity is required, a high value for the ratio of evanescent fluorescence intensity to total guided optical power is preferred. In spite of the fact that index-matching achieves this at the cost of an increase in the ~netration depth, the influence of the fluorescent molecules not bound to the sensor surface can be greatly reduced by the method of differential measurement.
RESULTS AND DISCUSSION
b“’ \
Laser beam
II
rubrtmte
-I
I
Fig. 8. ~~‘rn~t~ set-up Jbr the measuremmt of the output coupling elfiency of a completely buried pianur wavqukk
Ion-exchange condltion: K’-exchange and Na*-exchange; T-490%, t6Om
‘l%eformation of completely buried potassium ion-excsurged waveguides Our first goal was to bury the waveguide completely. Experiments showed that although the K+-exchange process could produce a surface planar waveguide with sufficient depth by increasing either the temperature or the time duration of the fab~~~on process, the optimum conditions for the complete burial of the waveguide through the Na+-exchange process were stringent Although lower temperature and shorter time would give a planar waveguide not completely buried, high temperature and long time would always lead to the cracking of the guiding layer. The plausible temperature and time ranges for both steps were found to be 480-495”C and 25-30 min, respectively. To confirm that the waveguide was completely buried and that a controllable evanescent penetration depth could be achieved by proper etching, the output coupling effkiency was measured using the experimental set-up shown in Fig. 8. The front end of the buried waveguide was etched step by step in 41 HF and a prism was placed on this area after each step. To obtain maximum coupling of the guided light out from the gradually etched area, the position of the prism and the pressure exerted by it on the waveguide were properly adjusted each time. Figure 9 shows the measured coupling eficiency, n = f,/(&, + le), as a function of the etching time, where 1, is the light intensity coupled out by the prism and& is the light intensity emitted from the end of the waveguide. As can be seen from Fig 9, when the waveguide is not etched, the coupling efticiency is virtually
0.0
0.6
1.0
l.6
2.0
2.3
3.0
3.3
4.0
4.6
etching time (minutes) Fig. 9. The output coupling ~~~ of a ~~~ buried waveguide as afinction of the et&kg time.
zero and it increases with the etching time until a maximum value is reached. This confirms that the waveguide is completely buried and that the evanescent penetration depth can be controlled by appropriate etching. Furtheretching leads to a decrease in the thickness of the guiding layer and finally the disappearance of the waveguiding effect. Index-mat~~mgmd thnwenee
enhcemeat
As pointed out above, index-matching can lead to a substantial increase in the evanescent fluorescence. However, the index-matching material must have the right refractive index, 603
Y: Zhou et al.
good optical quality to reduce scattering and a gelatinous property so that diffusion of proteins in aqueous solution can take place. In the search for such a material, we found that the readily available suitable materials were very limited. Materials that have been tested include gelatin, agar, agarose, polyacrylamide and polyvinyl alcohol. Of these materials, we found that only polyvinyl alcohol meets adequately the above requirements. Gelatin, agar and polyacrylamide were found to have a higher refractive index than that of the glass slide. The use of these materials would mean complete leakage of the guided light. Agarose does have a slightly lower refractive index than that of the glass slide but its optical quality is so poor that after the evaporation of water in the agarose, a lot of scattering can be readily observed under illumination. The evaporation of water is needed because the refractive index of water-tilled gelatinous materials is usually very. close to that of pure water and evaporation significantly increases the index. To confirm the effect of evanescent fluorescence enhancement through indexmatching, human IgG was physically adsorbed onto a patterned waveguide as described in the section ‘Protein immobilization by physical adsorption’. Then 1Opg ml-’ (w/v) protein A-FITC solution in 0.01 M PBS (pH 7.2) was applied to the patterned area and incubated at 37°C for 1 h. The slide was rinsed in RO water and blown dry, and 50% (w/v) polyvinyl alcohol solution in RO water was then applied to the etched areas of the patterned waveguide. Using the experimental set-up as shown in Fig. 1, measurement of the intensities of the evanescent fluorescence before and after the evaporation of water in polyvinyl alcohol was carried out, and the result clearly indicated an enhancement of at least two orders of magnitude.
Evanescent fluorescence sensing of specific binding
The excitation wavelength from the Ar+ laser is 488 nm, which is very close to one of the absorption peaks of the FITC molecule at 490nm. As the fluorescent emission of FITC molecules lies within the wavelength range between 515 and 525 nm and this is not very far from the excitation wavlength, a bandpass optical A, = 520 nm, wavelength filter (central bandwidth M = 8.2 nm) and a longpass tilter 604
Biosensors & Bioelectmnics 6 (1991) 595-607
(cut-off wavelength, & = 5 15 nm) were combined to make sure that the exciting light and the fluorescent light could be well separated. Experiments showed that this combination could reduce the 488 nm light intensity by a factor of 10-3-10-4, while allowing a reasonably high percentage of the fluorescent light to pass through. As the edges of our microscope glass slide were not well polished, it was found that fluorescent light would be diffused and scattered irregularly by the edges. To overcome this problem and extract more fluorescent light from the slide, polyvinyl alcohol solution was painted on the edges to suppress scattering from surface roughness. Figure 10 shows the variation of the fluorescence difference relative to the intensity of the guided exciting light, (I, - 1,)/I,, as a function of protein A concentration. As can be seen from the curve, for protein A, a concentration as low as 24 nM could easily be detected. After a number of attempts to test the reproducibility of the curve, it was found that difference in each waveguide caused during fabrication and fluorescence fading are two major reasons which lead to the large variation as shown. However, the curye is Key: intensities Is: specific
area,
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-
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(arbitrary
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-
t /
30
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/’
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50
100
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An evanescentfluorescence biosensor
Biosensors& Bioehronics 6 (1991) 595-607
similar in shape to Eddowes’s (1987-1988) theoretical curve, which takes the form of %
=
L’-kJWlec,+ K)
(12)
where 8, is the fractional surface coverage at equilibrium, [Aleg is the concentration of antigen or anti-antibody in solution at equilibrium and K is the dissociation constant. Attempts have also been made to detect fluorescence intensities at the level of 2 nM of protein A concentration, because at this level a difference in evanescent fluorescence from the two areas was still observable with the naked eye in darkness. Unfortunately, the fluorescence signal at this level was overwhelmed by the noise coming from background fluorescence, the detector or the lock-in amplitier. As a result, the sensitivity of our present device is only of the order of 10 nM. A higher sensitivity is expected when photomultipliers and/or APDs (avalanche photodiodes) are used and further improvements are made to reduce noise. Also shown in Fig. 10 is the curve when water is added to the polyvinyl alcohol on the patterned waveguide, resulting in the decrease of the medium. index refractive of the cover Comparison of the two curves confirms that an enhancement can be achieved through indexmatching. The fact that the enhancement decreases with decreasing concentrations of protein A-FITC is attributable to the detection limit caused by noise. Another problem possibly associated with the detection of evanescent fluorescence is that, as the two etched areas were separated by only 5 mm, when detection of the evanescent fluorescence from one etched area is carried out, the evanescent fluorescence from the other etched area might enter the detector and interfere. Although placing a slit in front of the detector should reduce interference, a better way would be to use lenses to focus the fluorescent light onto the detector. In addition, the two etched areas can be further separated. Discussion
As described in this paper, particular advantages of the optical waveguide biosensor are the potentially low cost of fabrication, and the simplicity and flexibility of the device. This combination enhances the concept of a disposable device. Although efforts have already
been made to develop an inexpensive process for the deposition of hard glassy metal phosphate films which act as optical waveguides (Sloper & Flanagan, 1988) the ion-exchange technique is a well-characterized and relatively cheap alternative for waveguide fabrication. In addition, because of the specific properties of the ion-exchange process, the buried waveguide has a continuous refractive index profile and, compared with film-deposited waveguides, which usually have a rough surface, this can lead to a substantial reduction in the surface scattering of guided light. The patterning of our waveguide offers other advantages over deposited planar waveguides, including the elimination of an additional buffer layer on top of the waveguide surface to restrict evanescent interaction to defined areas and the possibility of adjusting the evanescent penetration depth by appropriate etching. Although the patterning of our waveguides was done using standard lithography, which may sound a little complicated, this is not absolutely necessary, as the two etched wells do not need to have very well-defined walls, and simpler and less costly etching or grinding technique can be employed. A disadvantage of using a gelatinous material for index-matching is that the evaporation of water in the material takes time (usually 20-60 min depending on the volume) and the glass slide must be kept level during the evaporation to avoid extremely non-uniform distribution of the index-matching film. Solutions to the problem include the speeding-up of the evaporation process by either a little heating or the use of a dehydration agent, the use of other index-matching material and the fabrication of waveguides on a low refractive index substrate which matches the refractive index of aqueous solutions. The device has not been used for a real immunoassay yet, as the analyte to be sensed must be fluorescently labelled, and the measurements made so far are all controlled ones. However, this does not limit its adaptation to real immunoassays such as competitive immunoassays, as described by Badley et al. (1987) and sandwich immunoassays. A simple way of performing a real immunoassay is to construct a configuration similar to that of the FCFD. A plain glass slide may be placed at a fixed reasonable distance above the patterned waveguide immobilized with antibodies. Before 605
Y. Zhou et al.
this, a polyvinyl alcohol film can be formed on the patterned waveguide surface and the upper plate can be coated with dissolvable fluorescently labelled antigens specific for the immobilized antibodies. When the sample solution with unlabelled antigens fills the gap through, say, the capillary effect, competition for the limited number of immobilized antibodies between the labelled and unlabelled antigens then occurs.
CONCLUSION We have demonstrated for, we believe, the first time the principle of operation of an integrated optical molecular sensor using ion-exchanged waveguides. This prototype sensor enhances the concept of a disposable device as it has potential advantages such as low cost, simplicity and flexibility in waveguide fabrication, the possibility of a simple one-step operation, high sensitivity and selectivity, as well as possibilities for future device miniaturization. The evanescent fluorescence has been increased by several orders of magnitude through the use of an indexmatching material. The method of differential measurement, which improves the differentiation between specific and non-specific binding, has been proposed and performed. Although further investigations are required before the device is proved to be viable with a real analyte and a teal sample, the first attempt with ‘model’ binding showed that, for protein A-FITC, a concentration down to the order of 20 nM could be detected readily.
ACKNOWLEDGEMENTS We would like to thank Dr C. Marshman of Unilever, Professor J. Lamb, Dr P. Connolly, Dr S. Britland, Dr W. Cushley and Professor I. A. Shanks for helpful discussions and assistance. Financial support from the British Council, the SERC under grant GR/E 18704 and Unilever is gratefully acknowledged.
REFERENCES Adams, M. J. (1981). An Introduction to Optical Waveguides. Wiley, New York, pp. 90-134.
Biosensors & Bioeiectronics 6 (1991) 595-601
Drake,R.A L.,Shanks, I. A, Smith, A M. & Stephenson, P. R (1987).Optical biosensors for immunoassay - the fluorescence capillary fill device.Philos. Trans. R Sot. London, 316,143-&J.
Badley, R A,
Born, M. & Wolf, E. (1959). Principles of Optics. Pergamon, Oxford pp. 39-50. Cullen, D. C., Brown, R G. W. & Lowe, C. R (1988). Detection of immune-complex formation via surface plasmon resonance on gold-coated diffraction gratings. Biosensors, 3, 211-25. Eddowes, M. J. (1987-1988). Direct immunochemical basic chemical principles and sensing: fundamental limitations. Biosensots, 3, l-15. Elwing, H. & Stenberg, M. (1981). Biospecitic binding reactions - a bimolecular new method for their detection, ellipsometric quantification and characterization. J Immunol. Meth., 44,343-9. Garmire, E. (1985). Fundamentals of waveguides. Electromagnetic Surface Excitations, I+weedings of an International Summer School at the Ettore Majorana Centre, Eke, Italy. Springer-Verlag, Heidelberg, pp. 188-201. Harrick, N. J. (1967). Internal Reflection Sptwmscopy. Interscience. New York Kronick, M. N. &Little, W. A (1973). A new fluorescent immunoassay. Bull. Am Phys. Sot.. 18,782. Kronick M. N. & Little, W. A (1975). A new immunoassay based on fluorescence excitation by internal reflection spectrosc0py.J Immunol. Meth., 8, 235-42. Lee, El-H., Benner, R. E., Fenn, J. B. & Chang, R K. (1979). Angular distribution of fluorescence from liquids and monodispersed spheres by evanescent wave excitation. Appi. Optics. 18(6), 862-8. Liedberg, B., Nylander, C. & Lundstrom, I. (1983). Surface plasmon resonance for gas detection and biosensing. Proc. Int. Conf Solid State Transducers, IV; De@ Love. W. F., Walczak, I. M. & Slovacek, R E. (1989). High sensitivity fiber optic evanescent wave sensing for fluoroimmunoassay. Optical Fiber Sensors, Springer Proc. Phys., 44,431-5. Lukosz, W.. Nellen, Ph. M., Stamm, C. H. & Weiss, P. (1990). Output grating couplers on planar waveguides as integrated optical chemical sensors. Sens. Actuators, Bl, 585-8. Marcuse, D. (1974). Theory of Dielecttic Optical Waveguides. Academic Press, New York, pp. 714. Milliou, A, Zhenguang, H., Cheng, H. C., Srivastava, R. & Ramaswamy, R V. (1989). Fiber-compatible K+-Na+ ion-exchanged channel waveguides: fabrication and characterization. IEEE .I. Quantum. Electron., 25(8), 1889-97. Nellen, Ph. M. & Lukosz, W. (1990). Integrated optical couplers as chemoand input grating ‘immunosensors. Sens. Actuators. Bl, 592-6.
Biosensors & Bioelectrwnics 6 (1991) 595-601 Parry, R., Deacon, J. K., Robinson, G. A, Skehel, J. J. & Forrest, G. C. (1989). Surface plasmon resonance immunosensom. Pmt. F@h Znt. Symp. on Rapid Method and Automation in Microbiology and Immunology, Florence, November 1987, ed. k Turano, A Balows & R. C. Tilton. Brixia Academic Press, Brescia, pp. 641-8. Rogers, K R, Valdes, J. J. & Eldefrawi, M. E. (1989). Acetylcholine receptor fiber-optic evanescent fluorosensor. Anal. B&hem., 182,353-9.
An evanescent fluorescence biosensor Sloper, A. N. & Flanagan, M. T. (1988). Novel iron phosphate optical waveguides fabricated by a low temperature process. Electron. L&t., 24,353-5. Sloper, A. N., Deacon, J. IL & Flanagan, M. T. (1990). A planar indium phosphate monomode waveguide evanescent field immunosensor. Sens. Actuators, Bl, 589-91. Sutherland, R & Dahne, C. (1987). IRS devices for optical immunoassays. Biosensors, Oxford Science Publications, Oxford, pp. 655-78.
607
An evanescent fluorescence biosensor using ion-exchanged buried waveguides and the enhancement of peak fluorescence* Y. Zhou, P. J. R. Laybourn, J. V. Magill 81 R. M. De La Rue Department of Electronics and Electrical Engineering, University of Glasgow, Glasgow, G12 SQQ,
UK
(Received 24 May 1990,received in revised form 19 November 1990; accepted 4 December 1990)
Abstract: The principle of an optical molecular sensor using ion-exchanged buried planarwaveguides in glass has been demonstrated. We have shown both theoretically and experimentally that the intensity of the peak evanescent fluorescence can be increased by several orders of magnitude with the use of an index-matching material. The method of differential measurement has been used to improve the differentiation between specific and non-specific binding. We used h-IgG (human immunoglobulin G) as the immobilized antibody on the surface of the waveguide and protein A-PITC (fluorescein isothiocyanate) as the fluorescently labelled antigen or anti-antibody to be detected, and have shown that a concentration of protein A as low as 24 nM can easily be detected. Keywords: biosensor, immunosensor. ion-exchanged waveguide, total internal reflection fluorescence (TURF),tluoroimmunoassay. immobilization of proteins, antibody-antigen binding.
INTRODUCTION
Optical molecular sensors are analytical devices which convert a chemical parameter of a molecular species (e.g. concentration, intra- or inter-molecular structure, or orientation) into an optical signal (e.g. intensity, frequency, phase or polarization). As the number of molecular species which might be present in addition to the one to be detected is usually large, specificity or selectivity is definitely needed for the discrimination of low concentrations of analytes in complex matrices. To employ this property for *Paper presented at Biosensors90, Singapore,2-4 May 1990.
a wide range of assays for both antigens and antibodies, one of the most commonly used methods is to label a known antigen or antibody so that their location and quantity can be established. For example, antibodies can first be immobilized onto a solid support such as a glass microscope slide and a solution of antigens with a fluorescent label can then be applied to the immobilized antibodies. Immobilized antibodies raised against a specific antigen will selectively bind to that antigen, leaving the non-specific antibodies in the solution. After a certain time of incubation, the free antigens can be removed by a washing step. The glass slide is then illuminated by light of a suitable excitation wavelength in an
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optical microscope. The excitation light is filtered out before reaching the observer or the detector. Thus, only the fluorescent light from the glass slide is observed or detected and the position as well as the amount of labelled antigen can be established. A number of optical techniques have been used for molecular sensing purposes, including total internal reflection (TIR), ellipsometry, surface plasmon resonance (SPR) and light scattering. Of these techniques, however, only those which involve the use of evanescent interactions of light with molecules seem promising. The reason is that the dimension of the penetration depth of an optical evanescent wave (typically 100 nm) is approximately of the same order as that of the bimolecular complex (about ten to several tens of nanometers, depending on the size of the antigen and the antibody) and consequently the interaction of light with the molecules in the cover medium is mainly restricted to those molecules which are bound to the surface of the optical transducer. As a result, the washing step for the separation of bound molecules from unbound ones can readily be avoided. Although total internal reflection is the most often used method to generate an evanescent wave, integrated optical waveguides and optical fibres are preferred as the sensing element because their sensitivity can be increased as a result of the increase in the number of internal reflections over a reasonable length. Until now, several devices using evanescent interactions have been proposed and developed (Liedberg et al., 1983;Sutherland & Dahne, 1987; Badley et al., 1987; Cullen et al., 1988; Love ef al., 1989; Rogers et al., 1989; Nellen & Lukosz, 1990; Lukosz er al., 1990; Sloper et al., 1990).The idea of using an evanescent wave to study chemical reactions at a surface was put forward more than two decades ago (Harrick, 1967). It was then demonstrated by Rronick & Little (1973, 1975) that a wash-free immunoassay could be performed with a prism using total internal reflection fluorescence. An advance was made when Liedberg et al. (1983)published their results on the use of surface plasmon resonance (SPR) for the detection of gas and for biosensing. Later, the use of a diffraction grating as an alternative to prisms in SPR devices was reported (Cullen et al., 1988;Parry&al., 1989).Although the sensitivity of the two devices was similar, the use of a diffraction grating implied the possibility of cheap devices, as mass production of high5%
Biosensors& Bioekctronics 6 (1991) 595407
quality diffraction gratings has become a reality. The reporting of the fluorescence capillary fill device (FCFD) (Badley er al., 1987) introduced another device which enhances the concept of an inexpensive, disposable and manufacturable product. More recently, the use of diffraction gratings on planar dielectric waveguides for biosensing has been studied, and a sensitivity as high as 10s9 M has been reached (Nellen & Lukosz, 1990).In addition, investigations are also being made into fabrication of cheap planar monomode or low-mode order waveguides and the application of these waveguides to evanescent fluorescence biosensing (Sloper & Flanagan, 1988). A common problem associated with all these devices, especially the direct devices such as the SPR device (Liedberg et al., 1983; Cullen et al., 1988) and the grating coupling device (Nellen & Lukosz, 1990; Lukosz et al., 1990), is their sensitivity to non-specific interactions at the interface. In the FCFD (Badley et al., 1987),the discrimination between surface-bound fluorescence and bulk fluorescence is achieved by comparing the intensity of the fluorescent light emerging at smaller angles to the axis of the slide with that at larger angles. However, because the discrimination against fluorescence from within the fluid bulk is disrupted by the appearance of some bulk fluorescence in the smaller angle zone, the background fluorescence noise is high. A solution to this problem is the use of narrow planar waveguides so that evanescent excitation of fluorescence can be employed. Although an evanescent fluorescence immunosensor based on low-mode order planar waveguides has been reported (Sloper et al., l!WO), problems concerning its sensitivity, background noise and differentiation of specific binding from nonspecific binding still exist. In this paper, we describe the development of a planar guided-wave optical molecular sensor which attempts to overcome these problems. The device consists of a patterned planar waveguide fabricated on a glass microscope slide. A two-step ion-exchange process produces a completely buried monomode or low-mode order planar waveguide on the glass slide. Because of the complete burial of the waveguide, there is very little penetration of the evanescent wave beyond the glass surface. A patterned waveguide is then made by etching two rectangular patterns on the slide, contlning the penetration of the evanescent
Biosensors & Bioelectmnics 6
(1991)595-607
wave beyond the glass surface to within the two etched areas (see Figs l-3 later). Particular advantages of the ion-exchange process include low cost, simplicity and flexibility of fabrication. Although much effort has been spent on the fabrication of different kinds of optical waveguides using the ion-exchange technique, and a report has been made recently on the burial of K+-exchanged waveguides using electric-field assistance during the second step Na+-exchange (Milliou et al., 1989), this technique does not appear to have been applied to the development of integrated optical molecular sensors. The process developed by us for the fabrication of patterned waveguides is simple and the need for electric field assistance is eliminated. There are several advantages associated with patterned waveguides. Firstly, surface scattering of light guided by the waveguide can be greatly reduced, leading to less fluorescent noise produced by the surface-scattered light. Secondly, the penetration depth of the evanescent wave beyond the etched areas can be adjusted by appropriate etching. Thirdly, the complete burial of the waveguide eliminates the need for a patterned buffer layer on the slide to limit evanescent interactions within defined areas. Finally, by immobilizing specific proteins on one etched area and control proteins on the other, a differential measurement can be performed, and this can lead to a great improvement in the differentiation between specific binding and nonspecific binding. Another particular feature of our device is the use of an index-matching material as the cover medium, which leads to an increase of the peak evanescently excited fluorescence by several orders of magnitude.
MATERIALS AND METHODS
An evanescent jluorescence biosensor
weight 2000,75% dehydrolysed) were from Sigma Chemical Co. (St Louis, MO, USA). Fabrication of buried planar waveguides
Glass slides were cleaned by soaking successively in each of the following solutions in an ultrasonic bath for 5 min: (a) trichloroethylene; (b) methanol; (c)acetone. The slides were then rinsed with reverse osmosis (RO) water, blown dry with compressed nitrogen gas and placed over a hot plate (90°C) for 10 min. Potassium ion exchange was carried out at 490°C using a bath of molten potassium nitrate in a tube furnace. Glass slides were preheated in the furnace for 5 min to prevent cracking. The slides were then immersed in the molten potassium nitrate for 30 min, taken out of the bath, cooled to room temperature in open air and rinsed with hot water. For waveguide burying, the process was repeated under the same conditions using sodium nitrate instead of potassium nitrate. Patterning of the buried waveguide
The slides were cleaned as described in the previous section. A resist coating was made on the slide surface by spinning. After curing at 90°C for 1 h, two transparent rectangular patterns were formed on the resist coating near the centre of the slide by W light exposure and development. The etching of the patterns on the glass slide was performed by immersing the glass slides in Si02 etch (4~1 HF) with mild agitation for 2 min, rinsing with RO water and blowing dry with compressed nitrogen gas. One end of the slide (about 10 mm) was also etched by immersing it in the SiOz etch for 4 min. This was intended to facilitate the coupling of light into the waveguide when prism coupling was later used. We refer to slides treated to this stage as patterned waveguides.
Materials Protein immobilization by physical adsorption
Standard soda-lime glass microscope slides (75.5 mm X 26 mm X 1 mm) were purchased from BDH Ltd (Poole, Dorset, UK). Photoresist (145OJ) and Microposit Developer for Photoresist 14505 were supplied by Shipley Europe Ltd (Coventry, UK). H-&G (human immunoglobulin G), protein A-FITC (fluorescein isothiocyanate), BSA (bovine serum albumin), dichlorodimethylsilane (Si(CH&C12) and polyvinyl alcohol (molecular
The immobilization of proteins onto the sensor surface was based on a method described by Elwing & Stenberg (1981). Patterned waveguides were cleaned in 30% hydrogen peroxide in concentrated sulphuric acid for lo-15 min, rinsed with RO water, blown dry with compressed nitrogen gas and dried over a hot plate (90°C) for about 10 min. 597
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After cleaning, the slides were immersed for 30 s in a 2%solution of dichlorodimethylsilane in trichloroethylene, rinsed in trichloroethylene and blown dry with compressed nitrogen gas. This treatment generated a hydrophobic surface which would spontaneously adsorb BAS molecules and antibodies. Physical adsorption was achieved by leaving with a pipette a small amount (approximately 25 ~1) of 0.1 mg ml-’ h-IgG and O-1mg ml-’ BSA solution in 0.01 M PBS (phosphate-buffered saline) (pH 7.2) on the two etched areas respectively for 3-18 h at room temperature. To maintain a high humidity to prevent the evaporation of water, an airtight plastic wet box was used to contain the slides.
Biosensors t Bioekctronics 6 (1991) 595-607 Dimension: etched areas 1Omm x Smm separation of the two areas 5mm width of slide 26mm length of slide 76mm
mirror
chopper /
Application of the index-matching material and FlTC-labelled proteins
The slides with physically adsorbed proteins were rinsed with RO water to remove any unadsorbed proteins and blown dry with compressed nitrogen gas. Immediately after this, a small amount (about 025 ml) of 10% (w/v) polyvinyl alcohol solution was applied to the etched patterns, giving a uniform distribution of the index-matching material on the etched surfaces. The slides were always kept level so that a uniform film of the index-matching material (about 0.5 mm thick) was formed on the central area of the slides. A small equal amount of FITC-labelled protein A solution of various concentrations in 0.01 M PBS (pH 7.2) was then applied to the index-matching film of the slides, and the slides were left in open air until the water in the polyvinyl alcohol had completely evaporated (i.e. the material had dried). Alternatively, equal amounts of protein A-FITC solution and polyvinyl alcohol solution could be mixed together and then applied to the etched areas. Apparatus and evanescent fluorescence measurements
Figure 1 shows our experimental set-up for the measurement of the evanescent fluorescence. A slide was mounted on a rotation table. The 488 nm light beam from an Ar+ laser was passed through a chopper, reflected by a mirror and coupled into the waveguide through an input coupling prism. As the waveguide was completely buried, the prism was placed on the etched end of the slide and the rotation table was turned until 598
Fig I. Experimental set-up for the measurement of light intensities,
good input coupling of the excitation light was reached. Two detectors were placed at a fixed distance from the edges of the slide and polyvinyl alcohol solution was painted on the edges to reduce diffused reflection of light caused by the roughness of the edges. In front of detector 1, a bandpass filter (centre wavelength ;1, = 520 nm, bandwidth = 8.2 nm) and a long-pass filter (OG5 15, i.e. cut-off wavelength & = 515 nm) were placed to pass the fluorescent light and eliminate the light at the excitation wavelength. This detector moved along the edge of the slide and gave two maximum readouts, Z,and Z,,,where Z,is the fluorescent intensity from the specific area and Z, is the fluorescent intensity from the nonspecific area. Detector 2 was placed at the front end of the waveguide and measured the intensity of the reference guided exciting light, Z,. The chopping frequency for the excitation light was about 600 Hz and a lock-in amplifier was used to give the readouts of the measured intensities. PRINCIPLE OF OPERATION AND THEORETICAL ANALYSIS OF THE ENHANCEMENT OF EVANESCENT FLUORESCENCE Evaaescent excitation of fluorescence
When a beam of light is incident on the interface of two optically different media, it will in general
Biosensors t Bioelectronics 6 (1991) 595-607
be partially reflected and partially transmitted. However, if the light beam is incident from an optically denser medium onto an optically rarer medium and the angle of incidence 8i is greater than the critical angle defined as
e, = sin-‘(n&r,)
(1) where n, and nP are the refractive indices of the optically rarer and denser medium respectively, total internal reflection takes place. As a result, an evanescent wave is generated on the interface in the optically rarer medium. Typically, for a glass prism (n,, = 1.5) and an aqueous cover solution (4 = 1*33),0, Z 62”. A particular feature of the evanescent wave is that its amplitude and intensity decay exponentially from the interface. A planar waveguide can be used for the generation of an evanescent wave because the light guided by the waveguide undergoes total internal reflection. As shown in Fig. 2, a planar waveguide is a thin film of optically denser medium sandwiched between two optically rarer media. If n, is the refractive index of the cover medium, nB is the refractive index of the guiding layer, n, is the refractive index of the substrate, d is the thickness of the guiding layer, and ng > n,, n,, the electromagnetic lield beyond the interface in the cover medium can be expressed as E = A
exp(-6x)
forx>O
(2) where E is the amplitude of the electric field at x, A is the amplitude at the interface x = 0, & = 2n(n,fl* - nC2)“*/& h is the wavelength of the electromagnetic wave in free space, n,m is the effective refractive in&x of a mode guided by the planar waveguide (Garmire, 1985), and n, is the refractive index of the cover medium. The above expressions apply to all the guided modes, provided that different values of nemare used for different modes. The penetration depth of the evanescent wave is defined as
An evanescmt jluorescence biosensor
and typically, dp is about 100 nm for potassium ion-exchanged waveguides and aqueous cover solution, as will be shown later. In the case of total internal reflection, when the cover medium is a fluorescent dye solution, fluorescence can be excited by the evanescent wave. Furthermore, this fluorescent light emitted in an optically rarer medium can be coupled back into the optically denser medium and observed below the interface beyond the critical angle. A simple explanation of this is that as the exciting wave is evanescent, the excited fluorescent wave is also evanescent, and consequently it can be coupled back in the same way as an evanescent wave is generated. Principle of operation of the device
Figure 3 is an illustration of our device. Antibodies and control proteins such as BSA molecules are immobilized on the two etched areas of the patterned waveguide respectively and the surfaces are then covered with a film of indexmatching material (such as polyvinyl alcohol). Because of the gelatinous property of the indexmatching material, fluorescently labelled antigens or anti-antibodies in aqueous solution are able, when applied, to diffuse through the film and bind to the immobilized antibodies, ifthere is a significant degree of specificity between them. Because of the specific binding, more fluorescent molecules are localized close to the specific surface than to the non-specific one (i.e. the BSAimmobilized surface). When light of the excitation wavelength is guided in the waveguide, evanescent excitation of fluorescence takes place. As the optical power used for the excitation of
substrate
Fig. 2. Evanescent wave on a planar waveguide.
Fig. 3. Schematic illustration of the sensor device: n,. substrate or superstme tejkctive index; nF waveguide index; n,, re$uctive index of index-matching material. 599
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al.
evanescent fluorescence is only a small fraction of the total guided optical power, the evanescent exciting intensities for the two etched areas are roughly the same. However, because of the difference in the number of fluorescent molecules in the two etched areas, more evanescent fluorescence is expected from the specific area than from the non-specific area. If 1, is the fluorescence intensity from the specific area, 1, is the fluorescence intensity from the non-specific area and 1, is the reference intensity of the guided exciting light, the ratio of the evanescently excited fluorescence intensity difference to the intensity of the guided exciting light (i.e. (I, - I&) is proportional to the fractional surface coverage of the applied antigen or anti-antibody which is, in turn, dependent upon the concentration as well as the afftnity constant of the antigen or antiantibody to the immobilized antibody. Therefore, by measuringl& and I,, the concentration of the applied specific antigen or anti-antibody can be determined. The differential measurement method has the merit that when there is no specificity between the applied antigen or anti-antibody and the immobilized antibody, ideally& and& will be the same, regardless of the applied concentration, if the propagation loss of light is assumed to be negligible. As (I, - I,,)/& is equal to zero, the influence on optical output of the non-specific binding is eliminated. In practice, there is always some optical propagation loss; however, as long as this loss is taken into consideration, the differential measurement method can still be used. Fluorescence enhancement though index-matching
The angular distribution of fluorescence excited by an evanescent wave has been investigated both theoretically and experimentally by Lee et al. (1979). As shown in Fig. 4, the following assumptions are made: n, is the refractive index of the cover or sample solution, n,, is the refractive index of the hemicylindrical prism, Oi is the angle of incidence of the exciting beam of light, Oci is the critical angle pertaining to the wavelength of the exciting light, O,r is the critical angle pertaining to the wavelength of the fluorescent light, and 0s is the angle of fluorescence observation. Iflr(Oi, 0,) represents the fluorescence intensity excited at angle Oiand observed at angle 0s per 600
BiosensorsL Bioelectmnics6 (1991) 595-607
dye sample
reflected
light
1
incident
light
Fig. 4. Evanescentjluorescence.
unit incident light intensity, and zero dispersion in the relative refractive index nCP (=n,/n,) between the two wavelengths is assumed, the peak fluorescence intensity per unit incident light intensity for 0, > &i, and 0s = &r can be expressed as Mi, %d = a I W4 1 I * 9 I Wb)
(4) where a and n are the dye absorption coefftcient and the dye fluorescence efficiency respectively and are both assumed to be isotropic. T(&) and T(&) are the transmission terms for the incident intensity and for the fluorescent intensity, respectively. They are defined as the transmitted energy flux (or power) divided by the incident energy flux (or power). An important conclusion arising from Lee’s investigation is that, with evanescent excitation, the overall fluorescence intensity of a thick cover film (i.e. one satisfying the condition t > (2/dpi + 2/d,f)-‘, where t is the thickness of the film and dpi and dpf are the penetration depths of the exciting and fluorescent light respectively) of dye solution radiating into the optically denser medium is largest at an incident angle obeying 8i = &i, and, for a fixed Oi> B,i, the relative fluorescence intensity peaks at e, = ecf. The transmission term for the incident beam polarized vertically with respect to the plane of incidence can be related to the Fresnel transmission factor t, as follows (Born & Wolf, 1959): I *dpi /2
IT’(&)l* = nCPIt,(ei) I */cos ei The Fresnel transmission factor is
Itl(ei) I* =
I
ices
ei + ~~pY~sin2e.)l/2] I
(5)
I* t6)
Where the fluorescent light is concerned, its transmission term I Tl (0,) I is also given by eqn (5), with Oireplaced by 0, and nC,,replaced byits
Biosensors& Bioelectnmics6 (1991) 595407
An evanescentfluorescencebiosensor
value at the fluorescence wavelength. Corresponding equations for the horizontally polarized light have been published (Lee er al., 1979). The reason for choosing vertically polarized light here is simply that the expressions are simpler. In the case of a planar waveguide (Fig. 2), nc = n, and n,, = nB, where ns is the refractive index of the guiding layer. For a guided mode, if n,w is the effective refractive index of the guided mode (Garmire, 1985) 6, = sin-’ (n&t& and 8i > 8,-i are always valid. Taking into consideration the fact that (ncs2- sin28i)“2 is now imaginary, where nCg= nC/ng, substitution of n,, nP and 8i into eqns (6) and (5) results in 1T” (e,)
12 =
“$
1
:e$“2 c
If we restrict our observation of fluorescence to the peak value for vertical polarization, then e, = 0, = sin-’ (n&r,), where 0, is the critical angle for the fluorescent light. Replacement of 8i by 19, in eqns (5) and (6) leads to 1Tl (0,) I* = 4n,/(ni
- nz)“2
(8)
By substituting eqns (3), (7) and (8) into eqn (4), the peak fluorescence intensity per unit incident light intensity can be re-expressed in terms of n,, nB and nen as 4atjAi
Mnc,n,,neff) = y-
certainly true for our K+-exchanged waveguides (Adams, 1981). To use eqn (9) directly, the K+exchanged waveguide is treated as a step-index planar waveguide. Referring to Fig. 2, the following assumptions are made: the refractive index of the substrate, n, = l-510, the refractive index of the guiding layer, nB = 1.519, the thickness of the guiding layer, d = 1.4pm, and the refractive index of the cover medium n, lies between 1.33 and 1.51. To compare the peak evanescent fluorescence of an index-matching medium with that of an aqueous solution, we have chosen the refractive index of aqueous solutions, 1.33, as the starting point of the range of possible values for n,. As an ordinary glass microscope slide has a refractive index of 1.510 at the wavelength of sodium light (589 nm), this value is taken as the refractive index of the substrate. The reason for choosing ng = 1.519is that K+-exchange gives a maximum increase of 0009 in the refractive index of glass. Additionally, d = 1.4pm ensures that the waveguide supports a single mode within the range 1.33 < n, < l-510. Figure 5 is the calculated fluorescence enhancement as a function of n, for the TE mode. The fluorescence enhancement Et is defined as the peak fluorescence intensity at n, with respect to that at n, = 1.33, i.e. Et =
WC, ng9 nefd I&
= 1.33, nB, neft)
(10)
nC2(ng2- n,fFZ)“’ ’ (net? - n,2)‘R(nB2 - nC2)3/2
(9) As eqn (4) requires that the intensity of the light incident at the interface is normalized, the assumption is made that the optical power guided by the waveguide per unit width is constant. However, this does not mean that there is no change in the amplitude of the electromagnetic field at the interface, but rather that this change is accounted for by the transmission term T(&). As will be shown below, the increase in the fluorescence intensity is mainly caused by the increase in this field, although there is also an increase in the penetration depth of the evanescent wave. Typically, for a potassium ionexchanged planar waveguide and aqueous cover solution, ng = l-519, n, = 1.33, ner z 1.514 and consequently, If z 0.97 o&Ii. Generally speaking, a planar waveguide made by diffusion has a graded index profile, and this is
“C
Fig. 5. Enhancement of peak evanescent fluorescence. 601
Y: Zhou
Biosensors& Bioelectmnics6 (1991) 595407
et al.
In Fig 5, the starting point is one, i.e. Er( 1.33) = 1.With increase in the refractive index of the cover sample, Er increases. Its value and rate of increase can be very high as n, approaches R,. What this means is that, when every quantity except n, is held the same, the use of a cover medium with a higher refractive index, instead of an aqueous solution as in the case of most molecular sensors, can result in a substantial increase in the intensity of the evanescently excited fluorescence. However, as will be discussed below, the increase in n, also means an increase in the penetration depth, dr. This means that more fluorescent molecules not bound to the surface will be excited. Meanwhile, there exists a value for n, beyond which the waveguide will not be able to support any mode, and therefore choosing a value for n, very close to ns is not necessarily beneficial. A problem with index-matching is that the increase of n, leads to an increase in the penetration depth of the evanescent wave, and this will reduce the differentiation of fluorescence emitted by surface-bound molecules from that emitted by those in the bulk As a comparison, the increase of the penetration depth dp is plotted as a function of n, (eqn (3)) in Fig 6. As can be seen from Figs 5 and 6, the peak evanescent fluorescence increases much faster than the penetration depth, as n, approaches the indexd,
(urn)
A2 /A’0
1.33 1.35 1.37 1.39 1.41 1.43 1.4s I.47 1.49 1.51 no
Fig. 7. Enhancement of intensi~ (AZ/A&versus%.
matching point. In the case of polyvinyl alcohol (n, = 1.509) the enhancement is about 500 whereas the penetration depth is only five times that of the value at n, = 1.33.This actually means that the enhancement is mainly brought about by the increase in the electromagnetic intensity at the interface. To show this, we include here Fig 7, which gives the change in the intensity of the electromagnetic wave at the interface (i.e. A*, eqn (2)) as a function of n, for TE modes. The curve is derived from the following expression (Marcuse, 1974):
‘j A* =
0.1 1.33 1.36 1.37 1.39 141 I
h
#
I
4
I
I
nC
Fig. 6. Penetmtbn depth dp vemus q. 602
I
I
1.43 1.46 1.47 I.49 1.51
4&&P I/3I (d + l/y + l/S)&* + S*)
(11)
where A is the electric field amplitude at the interface, P is the total optical power per unit width, K = (2n/;l)(ns* - n,m2)ln is the transverse propagation constant, y = (2n/A)(n,fi* - n,2)“* and l/y = dpe, where dpl is the penetration depth into the substrate, 6 = (2n/~)(n,r,* - nC2)lDand l/S = d,,, where d,, is the penetration depth into the cover medium, d is the thickness of the guiding layer, 1j3 I= (2n/L)n,n is the longitudinal propagation constant, o = 2nf is the angular frequency of the light, and PO is the magnetic permeability. The enhancement of the electromagnetic intensity is defined as A* at n, divided by A2 at n, = 1.33, i.e. A*/A*(n, = l-33), with P assumed
Biosensors & Bioel~nics
An evanescentjluonscence bier
6 (1991) 595-601
unchanged. It is worth pointing out that lo~~~rnic coordinates have been used in Figs 5-7. When high sensitivity is required, a high value for the ratio of evanescent fluorescence intensity to total guided optical power is preferred. In spite of the fact that index-matching achieves this at the cost of an increase in the ~netration depth, the influence of the fluorescent molecules not bound to the sensor surface can be greatly reduced by the method of differential measurement.
RESULTS AND DISCUSSION
b“’ \
Laser beam
II
rubrtmte
-I
I
Fig. 8. ~~‘rn~t~ set-up Jbr the measuremmt of the output coupling elfiency of a completely buried pianur wavqukk
Ion-exchange condltion: K’-exchange and Na*-exchange; T-490%, t6Om
‘l%eformation of completely buried potassium ion-excsurged waveguides Our first goal was to bury the waveguide completely. Experiments showed that although the K+-exchange process could produce a surface planar waveguide with sufficient depth by increasing either the temperature or the time duration of the fab~~~on process, the optimum conditions for the complete burial of the waveguide through the Na+-exchange process were stringent Although lower temperature and shorter time would give a planar waveguide not completely buried, high temperature and long time would always lead to the cracking of the guiding layer. The plausible temperature and time ranges for both steps were found to be 480-495”C and 25-30 min, respectively. To confirm that the waveguide was completely buried and that a controllable evanescent penetration depth could be achieved by proper etching, the output coupling effkiency was measured using the experimental set-up shown in Fig. 8. The front end of the buried waveguide was etched step by step in 41 HF and a prism was placed on this area after each step. To obtain maximum coupling of the guided light out from the gradually etched area, the position of the prism and the pressure exerted by it on the waveguide were properly adjusted each time. Figure 9 shows the measured coupling eficiency, n = f,/(&, + le), as a function of the etching time, where 1, is the light intensity coupled out by the prism and& is the light intensity emitted from the end of the waveguide. As can be seen from Fig 9, when the waveguide is not etched, the coupling efticiency is virtually
0.0
0.6
1.0
l.6
2.0
2.3
3.0
3.3
4.0
4.6
etching time (minutes) Fig. 9. The output coupling ~~~ of a ~~~ buried waveguide as afinction of the et&kg time.
zero and it increases with the etching time until a maximum value is reached. This confirms that the waveguide is completely buried and that the evanescent penetration depth can be controlled by appropriate etching. Furtheretching leads to a decrease in the thickness of the guiding layer and finally the disappearance of the waveguiding effect. Index-mat~~mgmd thnwenee
enhcemeat
As pointed out above, index-matching can lead to a substantial increase in the evanescent fluorescence. However, the index-matching material must have the right refractive index, 603
Y: Zhou et al.
good optical quality to reduce scattering and a gelatinous property so that diffusion of proteins in aqueous solution can take place. In the search for such a material, we found that the readily available suitable materials were very limited. Materials that have been tested include gelatin, agar, agarose, polyacrylamide and polyvinyl alcohol. Of these materials, we found that only polyvinyl alcohol meets adequately the above requirements. Gelatin, agar and polyacrylamide were found to have a higher refractive index than that of the glass slide. The use of these materials would mean complete leakage of the guided light. Agarose does have a slightly lower refractive index than that of the glass slide but its optical quality is so poor that after the evaporation of water in the agarose, a lot of scattering can be readily observed under illumination. The evaporation of water is needed because the refractive index of water-tilled gelatinous materials is usually very. close to that of pure water and evaporation significantly increases the index. To confirm the effect of evanescent fluorescence enhancement through indexmatching, human IgG was physically adsorbed onto a patterned waveguide as described in the section ‘Protein immobilization by physical adsorption’. Then 1Opg ml-’ (w/v) protein A-FITC solution in 0.01 M PBS (pH 7.2) was applied to the patterned area and incubated at 37°C for 1 h. The slide was rinsed in RO water and blown dry, and 50% (w/v) polyvinyl alcohol solution in RO water was then applied to the etched areas of the patterned waveguide. Using the experimental set-up as shown in Fig. 1, measurement of the intensities of the evanescent fluorescence before and after the evaporation of water in polyvinyl alcohol was carried out, and the result clearly indicated an enhancement of at least two orders of magnitude.
Evanescent fluorescence sensing of specific binding
The excitation wavelength from the Ar+ laser is 488 nm, which is very close to one of the absorption peaks of the FITC molecule at 490nm. As the fluorescent emission of FITC molecules lies within the wavelength range between 515 and 525 nm and this is not very far from the excitation wavlength, a bandpass optical A, = 520 nm, wavelength filter (central bandwidth M = 8.2 nm) and a longpass tilter 604
Biosensors & Bioelectmnics 6 (1991) 595-607
(cut-off wavelength, & = 5 15 nm) were combined to make sure that the exciting light and the fluorescent light could be well separated. Experiments showed that this combination could reduce the 488 nm light intensity by a factor of 10-3-10-4, while allowing a reasonably high percentage of the fluorescent light to pass through. As the edges of our microscope glass slide were not well polished, it was found that fluorescent light would be diffused and scattered irregularly by the edges. To overcome this problem and extract more fluorescent light from the slide, polyvinyl alcohol solution was painted on the edges to suppress scattering from surface roughness. Figure 10 shows the variation of the fluorescence difference relative to the intensity of the guided exciting light, (I, - 1,)/I,, as a function of protein A concentration. As can be seen from the curve, for protein A, a concentration as low as 24 nM could easily be detected. After a number of attempts to test the reproducibility of the curve, it was found that difference in each waveguide caused during fabrication and fluorescence fading are two major reasons which lead to the large variation as shown. However, the curye is Key: intensities Is: specific
area,
Ir: reference
-
In: non-specific
(guided)
-8-
index matched
(Is-ln)/lr
(arbitrary
area
light
index match removed
unit)
5oi----
40
-
t /
30
,/ ,I--
20 -
10 -
/’
‘i
/
Y x /
_,OO’ 0
50
100
150
concentration Fig. 10. Fluorescence
200
2: i0
hM)
d$wnce versus concentmtion.
protein
A
An evanescentfluorescence biosensor
Biosensors& Bioehronics 6 (1991) 595-607
similar in shape to Eddowes’s (1987-1988) theoretical curve, which takes the form of %
=
L’-kJWlec,+ K)
(12)
where 8, is the fractional surface coverage at equilibrium, [Aleg is the concentration of antigen or anti-antibody in solution at equilibrium and K is the dissociation constant. Attempts have also been made to detect fluorescence intensities at the level of 2 nM of protein A concentration, because at this level a difference in evanescent fluorescence from the two areas was still observable with the naked eye in darkness. Unfortunately, the fluorescence signal at this level was overwhelmed by the noise coming from background fluorescence, the detector or the lock-in amplitier. As a result, the sensitivity of our present device is only of the order of 10 nM. A higher sensitivity is expected when photomultipliers and/or APDs (avalanche photodiodes) are used and further improvements are made to reduce noise. Also shown in Fig. 10 is the curve when water is added to the polyvinyl alcohol on the patterned waveguide, resulting in the decrease of the medium. index refractive of the cover Comparison of the two curves confirms that an enhancement can be achieved through indexmatching. The fact that the enhancement decreases with decreasing concentrations of protein A-FITC is attributable to the detection limit caused by noise. Another problem possibly associated with the detection of evanescent fluorescence is that, as the two etched areas were separated by only 5 mm, when detection of the evanescent fluorescence from one etched area is carried out, the evanescent fluorescence from the other etched area might enter the detector and interfere. Although placing a slit in front of the detector should reduce interference, a better way would be to use lenses to focus the fluorescent light onto the detector. In addition, the two etched areas can be further separated. Discussion
As described in this paper, particular advantages of the optical waveguide biosensor are the potentially low cost of fabrication, and the simplicity and flexibility of the device. This combination enhances the concept of a disposable device. Although efforts have already
been made to develop an inexpensive process for the deposition of hard glassy metal phosphate films which act as optical waveguides (Sloper & Flanagan, 1988) the ion-exchange technique is a well-characterized and relatively cheap alternative for waveguide fabrication. In addition, because of the specific properties of the ion-exchange process, the buried waveguide has a continuous refractive index profile and, compared with film-deposited waveguides, which usually have a rough surface, this can lead to a substantial reduction in the surface scattering of guided light. The patterning of our waveguide offers other advantages over deposited planar waveguides, including the elimination of an additional buffer layer on top of the waveguide surface to restrict evanescent interaction to defined areas and the possibility of adjusting the evanescent penetration depth by appropriate etching. Although the patterning of our waveguides was done using standard lithography, which may sound a little complicated, this is not absolutely necessary, as the two etched wells do not need to have very well-defined walls, and simpler and less costly etching or grinding technique can be employed. A disadvantage of using a gelatinous material for index-matching is that the evaporation of water in the material takes time (usually 20-60 min depending on the volume) and the glass slide must be kept level during the evaporation to avoid extremely non-uniform distribution of the index-matching film. Solutions to the problem include the speeding-up of the evaporation process by either a little heating or the use of a dehydration agent, the use of other index-matching material and the fabrication of waveguides on a low refractive index substrate which matches the refractive index of aqueous solutions. The device has not been used for a real immunoassay yet, as the analyte to be sensed must be fluorescently labelled, and the measurements made so far are all controlled ones. However, this does not limit its adaptation to real immunoassays such as competitive immunoassays, as described by Badley et al. (1987) and sandwich immunoassays. A simple way of performing a real immunoassay is to construct a configuration similar to that of the FCFD. A plain glass slide may be placed at a fixed reasonable distance above the patterned waveguide immobilized with antibodies. Before 605
Y. Zhou et al.
this, a polyvinyl alcohol film can be formed on the patterned waveguide surface and the upper plate can be coated with dissolvable fluorescently labelled antigens specific for the immobilized antibodies. When the sample solution with unlabelled antigens fills the gap through, say, the capillary effect, competition for the limited number of immobilized antibodies between the labelled and unlabelled antigens then occurs.
CONCLUSION We have demonstrated for, we believe, the first time the principle of operation of an integrated optical molecular sensor using ion-exchanged waveguides. This prototype sensor enhances the concept of a disposable device as it has potential advantages such as low cost, simplicity and flexibility in waveguide fabrication, the possibility of a simple one-step operation, high sensitivity and selectivity, as well as possibilities for future device miniaturization. The evanescent fluorescence has been increased by several orders of magnitude through the use of an indexmatching material. The method of differential measurement, which improves the differentiation between specific and non-specific binding, has been proposed and performed. Although further investigations are required before the device is proved to be viable with a real analyte and a teal sample, the first attempt with ‘model’ binding showed that, for protein A-FITC, a concentration down to the order of 20 nM could be detected readily.
ACKNOWLEDGEMENTS We would like to thank Dr C. Marshman of Unilever, Professor J. Lamb, Dr P. Connolly, Dr S. Britland, Dr W. Cushley and Professor I. A. Shanks for helpful discussions and assistance. Financial support from the British Council, the SERC under grant GR/E 18704 and Unilever is gratefully acknowledged.
REFERENCES Adams, M. J. (1981). An Introduction to Optical Waveguides. Wiley, New York, pp. 90-134.
Biosensors & Bioeiectronics 6 (1991) 595-601
Drake,R.A L.,Shanks, I. A, Smith, A M. & Stephenson, P. R (1987).Optical biosensors for immunoassay - the fluorescence capillary fill device.Philos. Trans. R Sot. London, 316,143-&J.
Badley, R A,
Born, M. & Wolf, E. (1959). Principles of Optics. Pergamon, Oxford pp. 39-50. Cullen, D. C., Brown, R G. W. & Lowe, C. R (1988). Detection of immune-complex formation via surface plasmon resonance on gold-coated diffraction gratings. Biosensors, 3, 211-25. Eddowes, M. J. (1987-1988). Direct immunochemical basic chemical principles and sensing: fundamental limitations. Biosensots, 3, l-15. Elwing, H. & Stenberg, M. (1981). Biospecitic binding reactions - a bimolecular new method for their detection, ellipsometric quantification and characterization. J Immunol. Meth., 44,343-9. Garmire, E. (1985). Fundamentals of waveguides. Electromagnetic Surface Excitations, I+weedings of an International Summer School at the Ettore Majorana Centre, Eke, Italy. Springer-Verlag, Heidelberg, pp. 188-201. Harrick, N. J. (1967). Internal Reflection Sptwmscopy. Interscience. New York Kronick, M. N. &Little, W. A (1973). A new fluorescent immunoassay. Bull. Am Phys. Sot.. 18,782. Kronick M. N. & Little, W. A (1975). A new immunoassay based on fluorescence excitation by internal reflection spectrosc0py.J Immunol. Meth., 8, 235-42. Lee, El-H., Benner, R. E., Fenn, J. B. & Chang, R K. (1979). Angular distribution of fluorescence from liquids and monodispersed spheres by evanescent wave excitation. Appi. Optics. 18(6), 862-8. Liedberg, B., Nylander, C. & Lundstrom, I. (1983). Surface plasmon resonance for gas detection and biosensing. Proc. Int. Conf Solid State Transducers, IV; De@ Love. W. F., Walczak, I. M. & Slovacek, R E. (1989). High sensitivity fiber optic evanescent wave sensing for fluoroimmunoassay. Optical Fiber Sensors, Springer Proc. Phys., 44,431-5. Lukosz, W.. Nellen, Ph. M., Stamm, C. H. & Weiss, P. (1990). Output grating couplers on planar waveguides as integrated optical chemical sensors. Sens. Actuators, Bl, 585-8. Marcuse, D. (1974). Theory of Dielecttic Optical Waveguides. Academic Press, New York, pp. 714. Milliou, A, Zhenguang, H., Cheng, H. C., Srivastava, R. & Ramaswamy, R V. (1989). Fiber-compatible K+-Na+ ion-exchanged channel waveguides: fabrication and characterization. IEEE .I. Quantum. Electron., 25(8), 1889-97. Nellen, Ph. M. & Lukosz, W. (1990). Integrated optical couplers as chemoand input grating ‘immunosensors. Sens. Actuators. Bl, 592-6.
Biosensors & Bioelectrwnics 6 (1991) 595-601 Parry, R., Deacon, J. K., Robinson, G. A, Skehel, J. J. & Forrest, G. C. (1989). Surface plasmon resonance immunosensom. Pmt. F@h Znt. Symp. on Rapid Method and Automation in Microbiology and Immunology, Florence, November 1987, ed. k Turano, A Balows & R. C. Tilton. Brixia Academic Press, Brescia, pp. 641-8. Rogers, K R, Valdes, J. J. & Eldefrawi, M. E. (1989). Acetylcholine receptor fiber-optic evanescent fluorosensor. Anal. B&hem., 182,353-9.
An evanescent fluorescence biosensor Sloper, A. N. & Flanagan, M. T. (1988). Novel iron phosphate optical waveguides fabricated by a low temperature process. Electron. L&t., 24,353-5. Sloper, A. N., Deacon, J. IL & Flanagan, M. T. (1990). A planar indium phosphate monomode waveguide evanescent field immunosensor. Sens. Actuators, Bl, 589-91. Sutherland, R & Dahne, C. (1987). IRS devices for optical immunoassays. Biosensors, Oxford Science Publications, Oxford, pp. 655-78.
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