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An optical fiber biosensor based on anomalous reflection of gold
Sensors and Actuators B 89 (2003) 126±130
An optical ®ber biosensor based on anomalous re¯ection of gold Mitsuaki Watanabe, Kotaro Kajikawa* Department of Information Processing, Interdisciplinary Graduate School of Science and Engineering, Tokyo Institute of Technology, 4259 Nagatsuta, Midori-ku, Yokohama 226-8502, Japan Received 19 June 2002; received in revised form 19 October 2002; accepted 22 November 2002
Abstract We propose a simple and sensitive optical sensing method for biological applications. Since gold behaves as a dielectric with a large extinction index under blue or violet light, presence of a transparent surface layer on gold produces a large decrease in the re¯ectivity of the gold surface due to multiple re¯ections in the surface layer. We call this phenomenon anomalous re¯ection (AR) of the gold surface. AR is applicable to af®nity biosensors based on ®ber optics, so that inexpensive and disposable micrometer-sized biosensor probes are possible based on this technique. Here, we demonstrate the application of AR to real-time measurements of the adsorption process of octadecanethiol (ODT) on gold and the af®nity of streptavidin to a biotin-labeled monomolecular layer on gold. # 2002 Elsevier Science B.V. All rights reserved. Keywords: SPR biosensor; Optical ®ber sensor; Re¯ection spectroscopy; Self-assembled monolayer; Biotin; Streptavidin
1. Introduction As a result of the development of genetic engineering and bioengineering, highly sensitive tools are needed to measure biomolecular interaction. The surface plasmon resonance (SPR) biosensor has been widely used in the last 10 years to analyze the af®nity between ligand and analyte [1,2]. The SPR condition is usually achieved in the attenuated total re¯ection (ATR) geometry by using prism-coupling optics. This optical setup is usually bulky and is incompatible with micrometer-sized probes. In place of the conventional SPR, interest in localized plasmon resonance (LPR) has grown in recent years, since LPR is quite compatible with such microsensors [3±6]. The LPR of gold appears as an absorption band at about 550 nm in the presence of thin molecular layers on gold. In this condition, the LPR produces a large enhancement of the electric ®eld around small gold particles or roughness of a gold surface, so that it is applicable in a highly sensitive af®nity sensor of size less than microns. Recent reports by Okamoto et al. [3] and Himmelhaus and Takei [4] have demonstrated LPR sensing with small gold particles ®xed on a substrate. However, the LPR condition is sensitive for the microscopic surface structure, reproducible way to control the surface is needed to make quantitative measurements. * Corresponding author. Tel./fax: 81-45-924-5596. E-mail address: [email protected] (K. Kajikawa).
The purpose of this paper is to propose a simple and reproducible method to probe the af®nity of biomolecular interaction with high sensitivity. Gold is metallic for light of wavelength longer than 550 nm, but it behaves as a dielectric rather than a metal for blue or violet light so that the re¯ectivity is less than 50%. In this case, adsorption of a monomolecular layer on a smooth gold surface produces a great reduction in re¯ectivity even when the layer is transparent, while the reduction in re¯ectivity for red light is small. Below, we call this phenomenon anomalous re¯ection (AR) of gold. The phenomenon of AR was reported in 1991 as a change in re¯ectivity of a gold surface in the presence of an inorganic thin ®lm of MgF2 deposited by vacuum evaporation [7]. The report discusses the relation between inhomogeneity of the thin ®lm and an LPR peak appearing at around 520 nm. However, practical application of AR has not yet been proposed to our knowledge. The AR af®nity biosensors proposed in this paper have several advantages over conventional SPR and LPR methods. First, the optical geometry is simple. This allows use in a variety of applications. Since AR measurements require only an optical geometry to probe re¯ectivity, the size of the probe is not limited. Fiber optics can therefore be exploited to make micrometer-sized probes. Secondly, results can be analyzed quantitatively by solving Maxwell equations with appropriate boundary conditions. Thirdly, the gold surface can be prepared by the simple vacuum evaporation method without any additional treatment. The conventional SPR
0925-4005/02/$ ± see front matter # 2002 Elsevier Science B.V. All rights reserved. doi:10.1016/S0925-4005(02)00453-7
M. Watanabe, K. Kajikawa / Sensors and Actuators B 89 (2003) 126±130
method demands that the thickness of the gold layer be in a narrow range for highly sensitive measurements. This simple preparation process enables us to fabricate simple and inexpensive devices. Fourthly, since AR is not a resonance effect and occurs over a large wavelength range of light, from 350 to 500 nm, incoherent light sources with a broad emission band such as a light emitting diode (LED) can be used. Below, we demonstrate AR biosensors based on ®ber optics. A blue LED (l 470 nm) was used as the light source. As expected from a simulation, the broad line width of LED does not cause any serious problem in these measurements. As an example, and to demonstrate the applicability of micrometer-sized af®nity biosensors with ®ber optics, we demonstrate the adsorption process of octadecanethiol (ODT) on gold. The optical thickness evaluated using this method is in good agreement with the value reported previously. We also show a measurement of adsorption process of streptavidin to the biotin-labeled SAMs on gold. 2. Experimental Fig. 1 shows our optical setup. The light source was a blue LED (l 470 nm), which was coupled into a multimode optical ®ber of diameter 62.5 mm. A 50 50 coupler was installed to split the incident light. The endface of the optical ®ber was brought to the sample surface with a stage. The distance between the endface of the ®ber and the gold surface was about 1 mm. The sample surface was irradiated with light from the optical ®ber at normal incidence. The re¯ected light was coupled into the ®ber and detected via a photomultiplier tube. The reference light was monitored with a photodiode to compensate for ¯uctuations of the light source. The ODT compound (Aldrich) used was puri®ed by recrystallization in ethanol. It was dissolved in pure ethanol at a concentration of 1 mM. The substrate was a silica plate coated with a 300 nm thick gold thin ®lm deposited by vacuum vapor
127
deposition at a pressure of 6 10 4 Pa. The refractive index of the gold was measured with a VASE 32 ellipsometer (J.A. Woollam Co., Inc.) using a Xe lamp source. For measurements in air, the gold surface of the sample plate was half coated with SAMs that had been formed by immersion in the ethanol solution for 2 h. Samples for measurements of the af®nity between streptavidin and the biotin-labeled monomolecular layer were prepared as follows [4]. The substrate was a silica plate coated with a 300 nm thick gold ®lm. The 1-amino-11undecanethiol (AUT) compound was purchased from Dojindo, Japan, and was used as received. The gold surface was half coated with an AUT SAM by exposure to the ethanol solution at a concentration of 0.1 mM for 2 h. Then the substrate with the AUT SAM was dipped in a borate buffer solution of sulfosuccinimidyl-N-[N0 -(D-biotinyl)-6-aminohexanoyl]-60 -aminohexanoate (biotin-(AC5)2 Sulfo-Osu) (Dojindo Laboratories, Japan) solution at a concentration of 0.1 mM for 2 h. Formation of the biotin-labeled AUT SAM had been checked beforehand by a real-time SPR measurement [8]. The streptavidin Mw 66 000) (Wako Chemical, Japan) was dissolved in borate buffer at a concentration of 1.5 mM. The biotin-labeled SAM was immersed in the streptavidin solution for 2 h and rinsed with the buffer solution. The strong af®nity between the streptavidin and the biotin-labeled AUT SAM was also con®rmed beforehand by a real-time SPR measurement [8]. 3. Results and discussion 3.1. Principle of the AR method Let us assume the three-layer model as depicted in the inset of Fig. 2(b). The surrounding medium 1 is air, water or ethanol with a refractive index n1; medium 2 is an adsorbed layer having a refractive index n2 and a thickness d2; medium 3 is gold used as a substrate. A re¯ectivity in the
Fig. 1. The optical setup used in the present study.
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M. Watanabe, K. Kajikawa / Sensors and Actuators B 89 (2003) 126±130
Fig. 2. The reflective index of gold measured by spectroscopic ellipsometry (a) and the reflectivity difference DR in the presence of the 1 nm thick transparent layer (medium 2) on gold (medium 3) simulated with the refractive index (b) as a function of the wavelength of the incident light in various surroundings (medium 1: air, water and ethanol). The three-layer model is also depicted in the inset.
absence of medium 2 corresponds to a re¯ectivity of the gold surface (medium 3). The re¯ectivity change upon adsorption of the adsorbed layer (medium 2) on the gold surface, DR, is de®ned as 1 R=Rgold , where R denotes a re¯ectivity of the gold surface covered with the layer and Rgold does that of a bare gold surface. DR can be obtained quantitatively with the model using the transfer matrix technique [9]. The refractive indexes used in the calculation were n1 1:0 for air, 1.33 for water and 1.36 for ethanol, n2 1:5 for the adsorbed layer and a complex refractive index of n3(l) for medium 3 of gold. Here we regarded only medium 3 of gold as a dispersible medium. Thus we measured n3(l) by spectroscopic ellipsometry as plotted in Fig. 2(a) against the wavelength l. It was in quite good agreement with those reported in the literature [10]. Fig. 2(b) shows the calculation of the re¯ectivity difference DR in air, water and ethanol at normal incidence in the presence of a 1 nm thick transparent dielectric layer 2 n2 1:5 on gold as a function of the wavelength of light l. This DR pro®le corresponds to re¯ection absorption spectrum of the transparent dielectric ®lm on gold. One may ®nd a decrease in re¯ectivity, i.e. absorption, occurs in spite of no absorption in the surface layer for blue and purple light, l < 480 nm, called AR. It is also observed that a small absorption peak due to LPR exists at about l 490 nm. Since intensity of the LPR peak is sensitive for the surface structure, we must measure the complex refractive index of gold to simulate this peak at each time by spectroscopic ellipsometry. On the other hand, DR of AR can be simulated with appropriate optical parameters because it is originated from the optical property of bulk gold. We calculated absorption at l 470 nm as a function of the thickness of the dielectric layer as shown in Fig. 3. The absorption increases almost proportionally with the thickness of the surface layer over the thickness of 30 nm. The principle of AR is as follows. For l > 550 nm, gold behaves as a metal because the real part of n3(l) is much
smaller than unity. On the other hand, for blue and purple light of l < 500 nm, the real part of the refractive index of gold is more than unity so that gold behaves as a dielectric rather than a metal with a large imaginary component (extinction index). In this case, multiple re¯ections in the dielectric layer produces a large decrease in the overall re¯ectivity because of the large imaginary component of n3. Hence, the substrate that shows AR should have a dielectric character with a large extinction index as is gold for blue and purple light. Thus AR can be observed in a copper surface, although it is less sensitive. 3.2. Demonstration of the AR method with SAMs in the air Table 1 summarizes DR values obtained in air by the re¯ectivity measurements on adsorption of various molecules on the gold surface with the optical setup depicted in Fig. 1. Formation of the ODT SAM yields DR 2:5%, and a value of DR 7:7% was found in the avidin/biotin-labeled AUT SAM on gold. These values are rather small compared
Fig. 3. The reflectivity difference DR in the presence of transparent layer on gold as a function of the layer thickness.
M. Watanabe, K. Kajikawa / Sensors and Actuators B 89 (2003) 126±130 Table 1 The reflectivity change DR and the corresponding thickness and mass of the adsorbed molecules in the region of the sensing area, 62.5 mm in diameter (the data were obtained in the measurement in air) Layer
DR (%)
Thickness of layer (nm)
Mass of adsorbed molecules (pg)
ODT AUT Biotin±AUT Avidin/biotin±AUT
2.50 2.36 5.07 7.68
1.9 1.8 3.7 5.4
5.8 5.5 11.4 16.6
with the change in re¯ectivity obtained by a conventional SPR measurement using an ATR geometry, but are suf®cient for the practical biosensing. The DR values give the layer thickness and the amount of adsorbed molecules on the sensing surface observed with the ®ber probe. DR 2:5% for the ODT SAM corresponds to a thickness of the monomolecular layer of 1.9 nm assuming n2 1:5 for the adsorbed molecular layer. The refractive index n2 1:5 is a typical value for organic monomolecular layers without optical absorption. The estimated thickness, d2 1:9 nm, is in agreement with the value of 2.2 nm reported previously [11]. The results for the AUT SAM, the biotin-labeled AUT SAM, and avidin/biotin-labeled
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AUT SAM are also in close agreement with values predicted from the molecular models. Suppose that the density of the layer is 1.0, we estimate the mass of the adsorbed molecules in the region of the sensing surface, 62.5 mm in diameter, which is the size of the optical ®ber core. These results show the high sensitivity of the AR method allowing the measurements of the order of magnitude of 10 12 g. 3.3. Demonstration of the AR method for real-time measurements Here we demonstrate the adsorption process or af®nity to the surface layer using AR in real time. Fig. 4(a) shows the real-time measurement of an adsorption process of ODT molecules on gold, normalized by the re¯ectivity of the bare gold surface, as a function of time t. A 1.0 mM ethanol solution of ODT was injected at t 0 into a cuvette of volume 200 ml. This procedure yielded a total concentration of 0.1 mM after complete mixture of the injected solution. The ODT SAM starts to form on gold at 500 s, with the delay due to slow mixture of the injected solution. The change in re¯ectivity stopped at t 900 s, and stayed constant. Since the surrounding medium has a refractive index n1 1:36 higher than that of air, the change in re¯ectivity DR is smaller than that in air (Table 2), as simulated results in
Fig. 4. Real-time in situ measurements of reflectivity normalized by the reflectivity at t 0. The adsorption process of ODT molecules on gold (a), and that of AUT on gold (b) were measured in the corresponding solution of ethanol. The reaction process of biotin-(AC5)2 Sulfo-Osu to AUT-labeled SAM (c) and that of streptavidin to biotin-labeled AUT SAM (d) was carried out in borate buffer solution.
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Table 2 The rate of decrease in reflectivity and the corresponding thickness and mass of the adsorbed molecules in the region of the sensing area, 62.5 mm in diametera Layer
Concentration (mM)
Rate of decrease (%)
Thickness of layer (nm)
Mass of adsorbed molecules (pg)
ODT AUT Biotin Avidin
0.1 0.01 0.01 0.00015
1.10 0.96 1.12 0.85
2.3 2.0 1.9 (3.9) 1.5 (5.4)
7.1 6.1 5.8 (11.9) 4.6 (16.5)
a
The rate of decrease was measured in ethanol for ODT and AUT and in borate buffer for biotin and avidin. The thickness and mass are the values of each layer. Total thickness and mass are addressed in the parenthesis to compare the results obtained in air (Table 1).
Fig. 2(b). The measured value DR 1:1% corresponds to a layer thickness of 2.3 nm, which is by 0.4 nm larger than that obtained in air. This is because of the presence of the additional physisorbed layer on the ODT SAM in solution, as reported previously [11,12]. Then we move to the real-time measurements of an adsorption process of streptavidin to the biotin-labeled AUT SAM. To prepare the biotin-labeled AUT SAM, we ®rst deposited the AUT molecules on gold in a 0.01 mM ethanol solution of AUT, and then deposit biotin molecule on it in the 0.01 mM borate buffer solution of biotin-(AC5)2 Sulfo-Osu. The deposition process was monitored by the AR method in real time as shown in Fig. 4(b) and (c) in order. The re¯ectivity is normalized by the average re¯ectivity before injection of the solution from t 500 to 0, where the cuvette was ®lled with the solvent, ethanol for AUT and borate buffer for biotin-(AC5)2 Sulfo-Osu. Injection of the solution t 0 yields a drastic decrease of the re¯ectivity, suggesting the successful adsorption of the molecules. The rates of decrease are summarized in Table 2. The thicknesses calculated with the rates are in agreement with the molecular length predicted from the molecular model. Then we carefully rinsed the substrate surface and put the substrate covered with the biotin-labeled AUT SAM in the cuvette. At t 0, the cuvette was ®lled with borate buffer. Upon injection of the streptavidin solution, the re¯ectivity decreases drastically to 99.15% compared with the re¯ectivity of the biotin-labeled AUT SAM. This value corresponds to a mass of the adsorbed streptavidin layer of a thickness of 1.5 nm (4:6 10 12 g), which is in good agreement with the value obtained in the air (Table 2), a thickness of 1.7 nm (5:2 10 12 g). This implies that excess streptavidin molecules are scarcely physisorbed on the surface of the avidin/biotin-labeled AUT SAM. The total thickness and mass for the biotin-labeled AUT SAM and the avidin/biotin-labeled AUT SAM are given in parenthesis in Table 2 for comparison with the values in Table 1.
4. Conclusion We have proposed and demonstrated that the AR method is applicable for af®nity optical biosensors. This method is compatible with ®ber optics, allowing inexpensive and disposable micrometer-sized probes. Optimization of the system to increase the sensitivity is in progress and will be reported shortly. Acknowledgements We thank Prof. K. Seki and Prof. Y. Ouchi of Nagoya University for kindly allowing us to use the ellipsometer. We gratefully acknowledge support from Nakatani Electronic Measuring Technology Association of Japan and the CASIO Foundation. References [1] K.R. Rogers, A. Mulchandani (Eds.), Affinity Biosensors Techniques and Protocols, Humanae Press, New Jersey, 1998. [2] J. Homola, S.S. Yee, G. Gaugliz, Sensors Actuators B 54 (1999) 3± 15. [3] T. Okamoto, I. Yamaguchi, T. Kobayashi, Opt. Lett. 25 (2000) 372± 374. [4] M. Himmelhaus, H. Takei, Sensors Actuators B 63 (2000) 24±30. [5] L.A. Lyon, M.D. Musick, P.C. Smith, B.D. Reiss, D.J. Pena, M.J. Natan, Sensors Actuators B 54 (1999) 118±124. [6] F. Meriaudeau, T. Downey, A. Wig, A. Passian, M. Buncick, T.L. Ferrell, Sensors Actuators B 54 (1999) 106±117. [7] V.-V. Truong, P.V. Ashrit, G. Bader, P. Courteau, F.E. Girouard, T. Yamaguchi, Can. J. Phys. 69 (1991) 107±113. [8] K. Kajikawa, M. Hara, H. Sasabe, W. Knoll, Jpn. J. Appl. Phys. 36 (1997) L1116±L1119. [9] D.S. Bethune, J. Opt. Soc. Am. B 6 (1989) 910±916. [10] E.D. Palik (Ed.), Handbook of Optical Constants of Solids, Academic Press, San Diego, 1985, pp. 286±295. [11] K.A. Peterlinz, R. Georgiadis, Langmuir 12 (1996) 4731±4740. [12] K. Tamada, M. Hara, H. Sasabe, W. Knoll, Langmuir 13 (1997) 1558±1566.
Biographies Mitsuaki Watanabe received BSc (Eng) in electronics in 2002 from Tokyo Institute of Technology. He is now a graduate student of Tokyo Institute of Technology. His work is focused on biosensors. Kotaro Kajikawa received the Doctoral Degree of Engineering from Tokyo Institute of Technology in 1992. He was with the Frontier Research Program, The Institute of Physical and Chemical Research (RIKEN) from 1993 to 1996. He was with the Department of Chemistry, Nagoya University from 1996 to 1999. He is now an associate professor of Tokyo Institute of Technology. His research group investigates biological sensors based on surface plasmon resonance and nonlinear optics in nanophotonics.
An optical ®ber biosensor based on anomalous re¯ection of gold Mitsuaki Watanabe, Kotaro Kajikawa* Department of Information Processing, Interdisciplinary Graduate School of Science and Engineering, Tokyo Institute of Technology, 4259 Nagatsuta, Midori-ku, Yokohama 226-8502, Japan Received 19 June 2002; received in revised form 19 October 2002; accepted 22 November 2002
Abstract We propose a simple and sensitive optical sensing method for biological applications. Since gold behaves as a dielectric with a large extinction index under blue or violet light, presence of a transparent surface layer on gold produces a large decrease in the re¯ectivity of the gold surface due to multiple re¯ections in the surface layer. We call this phenomenon anomalous re¯ection (AR) of the gold surface. AR is applicable to af®nity biosensors based on ®ber optics, so that inexpensive and disposable micrometer-sized biosensor probes are possible based on this technique. Here, we demonstrate the application of AR to real-time measurements of the adsorption process of octadecanethiol (ODT) on gold and the af®nity of streptavidin to a biotin-labeled monomolecular layer on gold. # 2002 Elsevier Science B.V. All rights reserved. Keywords: SPR biosensor; Optical ®ber sensor; Re¯ection spectroscopy; Self-assembled monolayer; Biotin; Streptavidin
1. Introduction As a result of the development of genetic engineering and bioengineering, highly sensitive tools are needed to measure biomolecular interaction. The surface plasmon resonance (SPR) biosensor has been widely used in the last 10 years to analyze the af®nity between ligand and analyte [1,2]. The SPR condition is usually achieved in the attenuated total re¯ection (ATR) geometry by using prism-coupling optics. This optical setup is usually bulky and is incompatible with micrometer-sized probes. In place of the conventional SPR, interest in localized plasmon resonance (LPR) has grown in recent years, since LPR is quite compatible with such microsensors [3±6]. The LPR of gold appears as an absorption band at about 550 nm in the presence of thin molecular layers on gold. In this condition, the LPR produces a large enhancement of the electric ®eld around small gold particles or roughness of a gold surface, so that it is applicable in a highly sensitive af®nity sensor of size less than microns. Recent reports by Okamoto et al. [3] and Himmelhaus and Takei [4] have demonstrated LPR sensing with small gold particles ®xed on a substrate. However, the LPR condition is sensitive for the microscopic surface structure, reproducible way to control the surface is needed to make quantitative measurements. * Corresponding author. Tel./fax: 81-45-924-5596. E-mail address: [email protected] (K. Kajikawa).
The purpose of this paper is to propose a simple and reproducible method to probe the af®nity of biomolecular interaction with high sensitivity. Gold is metallic for light of wavelength longer than 550 nm, but it behaves as a dielectric rather than a metal for blue or violet light so that the re¯ectivity is less than 50%. In this case, adsorption of a monomolecular layer on a smooth gold surface produces a great reduction in re¯ectivity even when the layer is transparent, while the reduction in re¯ectivity for red light is small. Below, we call this phenomenon anomalous re¯ection (AR) of gold. The phenomenon of AR was reported in 1991 as a change in re¯ectivity of a gold surface in the presence of an inorganic thin ®lm of MgF2 deposited by vacuum evaporation [7]. The report discusses the relation between inhomogeneity of the thin ®lm and an LPR peak appearing at around 520 nm. However, practical application of AR has not yet been proposed to our knowledge. The AR af®nity biosensors proposed in this paper have several advantages over conventional SPR and LPR methods. First, the optical geometry is simple. This allows use in a variety of applications. Since AR measurements require only an optical geometry to probe re¯ectivity, the size of the probe is not limited. Fiber optics can therefore be exploited to make micrometer-sized probes. Secondly, results can be analyzed quantitatively by solving Maxwell equations with appropriate boundary conditions. Thirdly, the gold surface can be prepared by the simple vacuum evaporation method without any additional treatment. The conventional SPR
0925-4005/02/$ ± see front matter # 2002 Elsevier Science B.V. All rights reserved. doi:10.1016/S0925-4005(02)00453-7
M. Watanabe, K. Kajikawa / Sensors and Actuators B 89 (2003) 126±130
method demands that the thickness of the gold layer be in a narrow range for highly sensitive measurements. This simple preparation process enables us to fabricate simple and inexpensive devices. Fourthly, since AR is not a resonance effect and occurs over a large wavelength range of light, from 350 to 500 nm, incoherent light sources with a broad emission band such as a light emitting diode (LED) can be used. Below, we demonstrate AR biosensors based on ®ber optics. A blue LED (l 470 nm) was used as the light source. As expected from a simulation, the broad line width of LED does not cause any serious problem in these measurements. As an example, and to demonstrate the applicability of micrometer-sized af®nity biosensors with ®ber optics, we demonstrate the adsorption process of octadecanethiol (ODT) on gold. The optical thickness evaluated using this method is in good agreement with the value reported previously. We also show a measurement of adsorption process of streptavidin to the biotin-labeled SAMs on gold. 2. Experimental Fig. 1 shows our optical setup. The light source was a blue LED (l 470 nm), which was coupled into a multimode optical ®ber of diameter 62.5 mm. A 50 50 coupler was installed to split the incident light. The endface of the optical ®ber was brought to the sample surface with a stage. The distance between the endface of the ®ber and the gold surface was about 1 mm. The sample surface was irradiated with light from the optical ®ber at normal incidence. The re¯ected light was coupled into the ®ber and detected via a photomultiplier tube. The reference light was monitored with a photodiode to compensate for ¯uctuations of the light source. The ODT compound (Aldrich) used was puri®ed by recrystallization in ethanol. It was dissolved in pure ethanol at a concentration of 1 mM. The substrate was a silica plate coated with a 300 nm thick gold thin ®lm deposited by vacuum vapor
127
deposition at a pressure of 6 10 4 Pa. The refractive index of the gold was measured with a VASE 32 ellipsometer (J.A. Woollam Co., Inc.) using a Xe lamp source. For measurements in air, the gold surface of the sample plate was half coated with SAMs that had been formed by immersion in the ethanol solution for 2 h. Samples for measurements of the af®nity between streptavidin and the biotin-labeled monomolecular layer were prepared as follows [4]. The substrate was a silica plate coated with a 300 nm thick gold ®lm. The 1-amino-11undecanethiol (AUT) compound was purchased from Dojindo, Japan, and was used as received. The gold surface was half coated with an AUT SAM by exposure to the ethanol solution at a concentration of 0.1 mM for 2 h. Then the substrate with the AUT SAM was dipped in a borate buffer solution of sulfosuccinimidyl-N-[N0 -(D-biotinyl)-6-aminohexanoyl]-60 -aminohexanoate (biotin-(AC5)2 Sulfo-Osu) (Dojindo Laboratories, Japan) solution at a concentration of 0.1 mM for 2 h. Formation of the biotin-labeled AUT SAM had been checked beforehand by a real-time SPR measurement [8]. The streptavidin Mw 66 000) (Wako Chemical, Japan) was dissolved in borate buffer at a concentration of 1.5 mM. The biotin-labeled SAM was immersed in the streptavidin solution for 2 h and rinsed with the buffer solution. The strong af®nity between the streptavidin and the biotin-labeled AUT SAM was also con®rmed beforehand by a real-time SPR measurement [8]. 3. Results and discussion 3.1. Principle of the AR method Let us assume the three-layer model as depicted in the inset of Fig. 2(b). The surrounding medium 1 is air, water or ethanol with a refractive index n1; medium 2 is an adsorbed layer having a refractive index n2 and a thickness d2; medium 3 is gold used as a substrate. A re¯ectivity in the
Fig. 1. The optical setup used in the present study.
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M. Watanabe, K. Kajikawa / Sensors and Actuators B 89 (2003) 126±130
Fig. 2. The reflective index of gold measured by spectroscopic ellipsometry (a) and the reflectivity difference DR in the presence of the 1 nm thick transparent layer (medium 2) on gold (medium 3) simulated with the refractive index (b) as a function of the wavelength of the incident light in various surroundings (medium 1: air, water and ethanol). The three-layer model is also depicted in the inset.
absence of medium 2 corresponds to a re¯ectivity of the gold surface (medium 3). The re¯ectivity change upon adsorption of the adsorbed layer (medium 2) on the gold surface, DR, is de®ned as 1 R=Rgold , where R denotes a re¯ectivity of the gold surface covered with the layer and Rgold does that of a bare gold surface. DR can be obtained quantitatively with the model using the transfer matrix technique [9]. The refractive indexes used in the calculation were n1 1:0 for air, 1.33 for water and 1.36 for ethanol, n2 1:5 for the adsorbed layer and a complex refractive index of n3(l) for medium 3 of gold. Here we regarded only medium 3 of gold as a dispersible medium. Thus we measured n3(l) by spectroscopic ellipsometry as plotted in Fig. 2(a) against the wavelength l. It was in quite good agreement with those reported in the literature [10]. Fig. 2(b) shows the calculation of the re¯ectivity difference DR in air, water and ethanol at normal incidence in the presence of a 1 nm thick transparent dielectric layer 2 n2 1:5 on gold as a function of the wavelength of light l. This DR pro®le corresponds to re¯ection absorption spectrum of the transparent dielectric ®lm on gold. One may ®nd a decrease in re¯ectivity, i.e. absorption, occurs in spite of no absorption in the surface layer for blue and purple light, l < 480 nm, called AR. It is also observed that a small absorption peak due to LPR exists at about l 490 nm. Since intensity of the LPR peak is sensitive for the surface structure, we must measure the complex refractive index of gold to simulate this peak at each time by spectroscopic ellipsometry. On the other hand, DR of AR can be simulated with appropriate optical parameters because it is originated from the optical property of bulk gold. We calculated absorption at l 470 nm as a function of the thickness of the dielectric layer as shown in Fig. 3. The absorption increases almost proportionally with the thickness of the surface layer over the thickness of 30 nm. The principle of AR is as follows. For l > 550 nm, gold behaves as a metal because the real part of n3(l) is much
smaller than unity. On the other hand, for blue and purple light of l < 500 nm, the real part of the refractive index of gold is more than unity so that gold behaves as a dielectric rather than a metal with a large imaginary component (extinction index). In this case, multiple re¯ections in the dielectric layer produces a large decrease in the overall re¯ectivity because of the large imaginary component of n3. Hence, the substrate that shows AR should have a dielectric character with a large extinction index as is gold for blue and purple light. Thus AR can be observed in a copper surface, although it is less sensitive. 3.2. Demonstration of the AR method with SAMs in the air Table 1 summarizes DR values obtained in air by the re¯ectivity measurements on adsorption of various molecules on the gold surface with the optical setup depicted in Fig. 1. Formation of the ODT SAM yields DR 2:5%, and a value of DR 7:7% was found in the avidin/biotin-labeled AUT SAM on gold. These values are rather small compared
Fig. 3. The reflectivity difference DR in the presence of transparent layer on gold as a function of the layer thickness.
M. Watanabe, K. Kajikawa / Sensors and Actuators B 89 (2003) 126±130 Table 1 The reflectivity change DR and the corresponding thickness and mass of the adsorbed molecules in the region of the sensing area, 62.5 mm in diameter (the data were obtained in the measurement in air) Layer
DR (%)
Thickness of layer (nm)
Mass of adsorbed molecules (pg)
ODT AUT Biotin±AUT Avidin/biotin±AUT
2.50 2.36 5.07 7.68
1.9 1.8 3.7 5.4
5.8 5.5 11.4 16.6
with the change in re¯ectivity obtained by a conventional SPR measurement using an ATR geometry, but are suf®cient for the practical biosensing. The DR values give the layer thickness and the amount of adsorbed molecules on the sensing surface observed with the ®ber probe. DR 2:5% for the ODT SAM corresponds to a thickness of the monomolecular layer of 1.9 nm assuming n2 1:5 for the adsorbed molecular layer. The refractive index n2 1:5 is a typical value for organic monomolecular layers without optical absorption. The estimated thickness, d2 1:9 nm, is in agreement with the value of 2.2 nm reported previously [11]. The results for the AUT SAM, the biotin-labeled AUT SAM, and avidin/biotin-labeled
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AUT SAM are also in close agreement with values predicted from the molecular models. Suppose that the density of the layer is 1.0, we estimate the mass of the adsorbed molecules in the region of the sensing surface, 62.5 mm in diameter, which is the size of the optical ®ber core. These results show the high sensitivity of the AR method allowing the measurements of the order of magnitude of 10 12 g. 3.3. Demonstration of the AR method for real-time measurements Here we demonstrate the adsorption process or af®nity to the surface layer using AR in real time. Fig. 4(a) shows the real-time measurement of an adsorption process of ODT molecules on gold, normalized by the re¯ectivity of the bare gold surface, as a function of time t. A 1.0 mM ethanol solution of ODT was injected at t 0 into a cuvette of volume 200 ml. This procedure yielded a total concentration of 0.1 mM after complete mixture of the injected solution. The ODT SAM starts to form on gold at 500 s, with the delay due to slow mixture of the injected solution. The change in re¯ectivity stopped at t 900 s, and stayed constant. Since the surrounding medium has a refractive index n1 1:36 higher than that of air, the change in re¯ectivity DR is smaller than that in air (Table 2), as simulated results in
Fig. 4. Real-time in situ measurements of reflectivity normalized by the reflectivity at t 0. The adsorption process of ODT molecules on gold (a), and that of AUT on gold (b) were measured in the corresponding solution of ethanol. The reaction process of biotin-(AC5)2 Sulfo-Osu to AUT-labeled SAM (c) and that of streptavidin to biotin-labeled AUT SAM (d) was carried out in borate buffer solution.
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Table 2 The rate of decrease in reflectivity and the corresponding thickness and mass of the adsorbed molecules in the region of the sensing area, 62.5 mm in diametera Layer
Concentration (mM)
Rate of decrease (%)
Thickness of layer (nm)
Mass of adsorbed molecules (pg)
ODT AUT Biotin Avidin
0.1 0.01 0.01 0.00015
1.10 0.96 1.12 0.85
2.3 2.0 1.9 (3.9) 1.5 (5.4)
7.1 6.1 5.8 (11.9) 4.6 (16.5)
a
The rate of decrease was measured in ethanol for ODT and AUT and in borate buffer for biotin and avidin. The thickness and mass are the values of each layer. Total thickness and mass are addressed in the parenthesis to compare the results obtained in air (Table 1).
Fig. 2(b). The measured value DR 1:1% corresponds to a layer thickness of 2.3 nm, which is by 0.4 nm larger than that obtained in air. This is because of the presence of the additional physisorbed layer on the ODT SAM in solution, as reported previously [11,12]. Then we move to the real-time measurements of an adsorption process of streptavidin to the biotin-labeled AUT SAM. To prepare the biotin-labeled AUT SAM, we ®rst deposited the AUT molecules on gold in a 0.01 mM ethanol solution of AUT, and then deposit biotin molecule on it in the 0.01 mM borate buffer solution of biotin-(AC5)2 Sulfo-Osu. The deposition process was monitored by the AR method in real time as shown in Fig. 4(b) and (c) in order. The re¯ectivity is normalized by the average re¯ectivity before injection of the solution from t 500 to 0, where the cuvette was ®lled with the solvent, ethanol for AUT and borate buffer for biotin-(AC5)2 Sulfo-Osu. Injection of the solution t 0 yields a drastic decrease of the re¯ectivity, suggesting the successful adsorption of the molecules. The rates of decrease are summarized in Table 2. The thicknesses calculated with the rates are in agreement with the molecular length predicted from the molecular model. Then we carefully rinsed the substrate surface and put the substrate covered with the biotin-labeled AUT SAM in the cuvette. At t 0, the cuvette was ®lled with borate buffer. Upon injection of the streptavidin solution, the re¯ectivity decreases drastically to 99.15% compared with the re¯ectivity of the biotin-labeled AUT SAM. This value corresponds to a mass of the adsorbed streptavidin layer of a thickness of 1.5 nm (4:6 10 12 g), which is in good agreement with the value obtained in the air (Table 2), a thickness of 1.7 nm (5:2 10 12 g). This implies that excess streptavidin molecules are scarcely physisorbed on the surface of the avidin/biotin-labeled AUT SAM. The total thickness and mass for the biotin-labeled AUT SAM and the avidin/biotin-labeled AUT SAM are given in parenthesis in Table 2 for comparison with the values in Table 1.
4. Conclusion We have proposed and demonstrated that the AR method is applicable for af®nity optical biosensors. This method is compatible with ®ber optics, allowing inexpensive and disposable micrometer-sized probes. Optimization of the system to increase the sensitivity is in progress and will be reported shortly. Acknowledgements We thank Prof. K. Seki and Prof. Y. Ouchi of Nagoya University for kindly allowing us to use the ellipsometer. We gratefully acknowledge support from Nakatani Electronic Measuring Technology Association of Japan and the CASIO Foundation. References [1] K.R. Rogers, A. Mulchandani (Eds.), Affinity Biosensors Techniques and Protocols, Humanae Press, New Jersey, 1998. [2] J. Homola, S.S. Yee, G. Gaugliz, Sensors Actuators B 54 (1999) 3± 15. [3] T. Okamoto, I. Yamaguchi, T. Kobayashi, Opt. Lett. 25 (2000) 372± 374. [4] M. Himmelhaus, H. Takei, Sensors Actuators B 63 (2000) 24±30. [5] L.A. Lyon, M.D. Musick, P.C. Smith, B.D. Reiss, D.J. Pena, M.J. Natan, Sensors Actuators B 54 (1999) 118±124. [6] F. Meriaudeau, T. Downey, A. Wig, A. Passian, M. Buncick, T.L. Ferrell, Sensors Actuators B 54 (1999) 106±117. [7] V.-V. Truong, P.V. Ashrit, G. Bader, P. Courteau, F.E. Girouard, T. Yamaguchi, Can. J. Phys. 69 (1991) 107±113. [8] K. Kajikawa, M. Hara, H. Sasabe, W. Knoll, Jpn. J. Appl. Phys. 36 (1997) L1116±L1119. [9] D.S. Bethune, J. Opt. Soc. Am. B 6 (1989) 910±916. [10] E.D. Palik (Ed.), Handbook of Optical Constants of Solids, Academic Press, San Diego, 1985, pp. 286±295. [11] K.A. Peterlinz, R. Georgiadis, Langmuir 12 (1996) 4731±4740. [12] K. Tamada, M. Hara, H. Sasabe, W. Knoll, Langmuir 13 (1997) 1558±1566.
Biographies Mitsuaki Watanabe received BSc (Eng) in electronics in 2002 from Tokyo Institute of Technology. He is now a graduate student of Tokyo Institute of Technology. His work is focused on biosensors. Kotaro Kajikawa received the Doctoral Degree of Engineering from Tokyo Institute of Technology in 1992. He was with the Frontier Research Program, The Institute of Physical and Chemical Research (RIKEN) from 1993 to 1996. He was with the Department of Chemistry, Nagoya University from 1996 to 1999. He is now an associate professor of Tokyo Institute of Technology. His research group investigates biological sensors based on surface plasmon resonance and nonlinear optics in nanophotonics.