---
title: "Photonic crystal surface waves for optical biosensors"
authors: ["Valery N. Konopsky", "Elena V. Alieva"]
affiliation: "Institute of Spectroscopy, Russian Academy of Sciences, Troitsk, Moscow region, Russia."
journal: "Analytical Chemistry"
year: 2007
volume: "79"
issue: "12"
article_number: ""
pages: "4729--4735"
doi: "10.1021/ac070275y"
type: journal-article
site_group: "Biosensors based on optical surface modes"
url_abstract: "https://valery.konopsky.com/paper/AC2007/AC2007.htm"
url_pdf: "https://valery.konopsky.com/kvnlocal/ac070275y.pdf"
language: en
source_tex: "Z:\\ValeryData\\Valery_New\\my_articles\\Analytical Chemistry2007\\Analytical Chemistry\\Sent\\Konop6.tex"
source_pdf: "Z:\\ValeryData\\Valery_New\\my_articles\\Analytical Chemistry2007\\Analytical Chemistry\\published\\ac070275y.pdf"
---
## Abstract
We present a new optical biosensor technique based on registration of dual optical s-polarized modes on a photonic crystal surface. The simultaneous registration of two optical surface waves with different evanescent depths from the same surface spot permits the segregation of the volume and the surface contributions from an analyte, while the absence of metal damping permits an increase in the propagation length of the optical surface waves and the sensitivity of the biosensor. Our technique was tested with the binding of biotin molecules to a streptavidin monolayer that has been detected with signal/noise ratio of about 15 at 1 second signal accumulation time.
Institute of Spectroscopy, Russian Academy of Sciences, Troitsk, Moscow region, 142190, Russia.
Optical biosensors have played a key role in the selective recognition of target biomolecules and in biomolecular interaction analysis (BIA), providing kinetics data of biological binding events in real time without labelling. Advantages of the label-free concept are the elimination of undue detrimental effects from labels that may interfere with fundamental interaction and the absence of a time consuming pretreatment. [1] Disadvantages of all label-free techniques, including the most mature one – surface plasmon resonance (SPR) technique, [2] are a deficient sensitivity to a specific signal and undesirable susceptibilities to non-specific signals, e.g., to the volume effect of refraction index variations. Our goal was to overcome these variations caused by temperature fluctuations and drifts that are a problem for many state-of-the-art optical biosensors.
Registration of optical waves propagating along the surface under investigation is the much used method in the label-free optical biosensors. [3] In the SPR technique these waves are surface plasmon-polaritons [4] propagating along a gold or silver surface, while in the resonant mirror (RM) technique [5] the waves are waveguade modes excited in a high refractive index dielectric layer via the frustrated total internal reflection (TIR) from a low refractive index spacer. In both cases an evanescent field of the optical wave (with penetration depth in water $\sim 100$ nm) is sensitive not only to biomolecular interactions at the surface, but also to changes in the volume refraction index (RI) of the liquid due to variations of the liquid temperature, composition and so on. For example, a water temperature change of 0.1$^0$C gives a water RI change of about 10$^{-5}$.
Therefore, a need exists for a biosensor technique that would be able to segregate the volume and the surface contributions from an analyte in detected signals. To obtain these two parameters one needs to detect at least two optical waves with different characteristics (e.g., with different penetration depths) simultaneously. In the work [6] we exploited a bulk optical wave, propagating above the sensing surface as a reference of the volume RI fluctuations. The weakness of this method is the impossibility of decreasing the flow cell height (and therefore the flow cell volume) because of using the bulk optical wave. In a dual-waveguide interferometric technique [8] the measurement of propagation constants of two modes with s- and p-polarizations is used to seek an adsorption layer thickness and its RI. It is worth noting here that the exploitation of the modes with the ortogonal polarizations may be stated as a weakness of the method, because of an implicit assumption that the adalayer is an isotropic substance, while the adalayer is almost always anizotropic (and birefringent to some extent) in consequence of its binding to the surface.
Here we present a technique based on the simultaneous registration of two s-polarized optical surface waves on a one-dimensional photonic crystal surface. Photonic crystals (PCs) are materials that possess a periodic modulation of their refraction index on the scale of the wavelength of light. [9] Such materials can exhibit photonic band gaps that are very much like the electronic band gaps for electron waves travelling in the periodic potential of the crystal. In both cases, frequency intervals exist where the wave propagation is forbidden. This analogy may be extended [10] to include surface levels, which can exist in band gaps of electronic crystals. In PCs, they correspond to optical surface waves with dispersion curves located inside the photonic band gap.
The one-dimensional photonic crystal (1D PC) is a simple periodic multilayer stack. Optical surface modes in 1D PCs were studied in the 1970s, both theoretically [11] and experimentally. [12] Twenty years later, the excitation of optical surface waves in a Kretschmann-like configuration was demonstrated. [13] Despite several theoretical proposals [13, 14] that suggested that the photonic crystal surface waves (PC SWs) have the potential to be superior alternatives in sensor applications to surface plasmons (due to low damping of PC SWs), there are no experimental demonstrations of such applications to date. In our opinion, the reason is the above mentioned point that the limiting factor for the SPR technique is not the instrumental sensitivity but the temperature fluctuations and drifts. From this point of view the increase of a propagation length of surface waves itself is ineffective without a concurrent compensation of the fluctuations of the liquid.
We show that in addition to the low loss propagation (which is not unique among other all-dielectric biosensors), the presented technique based on *dual* optical surface waves in 1D PCs has some additional advantages over all the above mentioned biosensor techniques. Unique tunable properties of 1D PCs permit the design of a 1D PC structure that can support two long-range surface modes at the same wavelength (this is impossible in the SPR technique), with one mode exited very close to the angle of TIR from the water (this is unfeasible in any other waveguide techniques). The mode, in which the exited angle is infinitesimally close to the angle of TIR from the external medium, has a very large penetration depth in this medium (e.g., water) and may be used as a reference of the water RI fluctuations. Indeed, the weak localization of this mode reduces its sensitivity to overlayers and increases its sensitivity to changes in the RI of the water. Simultaneous detection of two modes, with one of them being more sensitive to changes of the RI of the liquid then the other, permit us to derive both the RI of the liquid, $n_e=n_e(\rho_1,\rho_2)$, and the adalayer thickness, $d_a=d_a(\rho_1,\rho_2)$, as functions of the detected angular parameters $\rho_1$ and $\rho_2$ of two PC SWs.
**EXPERIMENTAL SECTION**
**Photonic crystal structure.** The following 1D PC structure was used in experiments:\
substrate/$(LH)^{3}L'$/water, where $L$ is a $SiO_2$ layer with $d_1=154.0$ nm, $H$ is a $Ta_2O_5$ layer with thickness $d_2=89.4$ nm and $L'$ is a $SiO_2$ layer with $d_3=638.5$ nm. The $SiO_2/Ta_2O_5$ 7-layers structure (started and finished by $SiO_2$ layers) was deposited by ion sputtering. The prism and substrate were made from BK-7 glass. The RIs of the substrate, $SiO_2$, $Ta_2O_5$ and water at $\lambda=532$ nm, were $n_0=1.52$, $n_1=n_3=1.49$, $n_2=2.12$ and $n_e=1.335$, correspondingly. The RIs at other wavelengths were derived using dispersion data presented by Palik. [15]
**Absolute angle measurements.** The excitation angles of the optical surface waves, indicated as black diamonds in Fig. 1, were experimentally measured with an angular accuracy of $\pm 1'$ by parallel laser beam at $\lambda=532$ nm (2nd harmonic of Nd-YAG laser) and $\lambda=442$ nm (He-Cd laser).
**Materials.** All biochemicals (except streptavidin) were purchased from Sigma-Aldrich (Germany) and were used immediately after preparation. A dialkoxy aminosilane 3-(2-Aminoethylamino) propyl-dimethoxymethylsilane [molecular weight – 206.36 ] was used to convert $OH$-terminated $SiO_2$ surface to $NH_2$-terminated one. [16] Biotin-XX, SSE [Sulfosuccinimidyl Ester sodium salt or Sulfo-NHS-LC-LC-Biotin; molecular weight – 669.74; mass added to target – 452.6 ] was used for the biotinylation of the amino-terminated surface. The streptavidin from Amersham (UK) [molecular weight $\sim 60\,000$] was deposited on the biotinylated surface. The free biotin [vitamin H; molecular weight $244.31$ ] was used as a test small molecule binding with streptavidin monolayer. All experiments were done in the PBS [phosphate-buffered saline; pH=7.2] except absolute angle measurements of PC SWs excitation, which were done in pure water.
**Sample preparation.** Samples (i.e, their top silicon oxide layers with thickness 638.5 nm) were cleaned as follows: first, they were sonicated in ethanol and acetone for 5 min each and then immersed into a Piranha solution ($H_2SO_4:H_2O_2=3:1$) for 15 min (caution, piranha solution reacts violently with organic solvents). The glass slides were then exposed to UV-ozone (185 nm and 254 nm) for 45 min and finally thoroughly rinsed with DI water. The precleaned glass slides (with expected $OH$ bonds on $SiO_2$ surface) were immersed in 1% aminosilane solution in 95% acetone/water for 5 min. The slides were then rinsed with acetone and baked for 30 min at $120^0$C. Then the sample was mounted in the flow cell, and further sample treatment was made *in situ*. For the biotinylation of the $NH_2$-terminated surface of the slides, Sulfo-NHS-LC-LC-Biotin (2mg/mL in PBS) was flowed over the flow cell for several minutes and then the fluid flow was stopped for several hours or even overnight and the biotinylation of the surface was monitored in real time. Then the flow cell system was thoroughly rinsed by PBS.
**Liquid handling.** The flow cell consists of a glass slide with two holes in which two glass tubes are fitted, serving as inlet and outlet, respectively. The height of the cell is determined by a thickness of a Teflon film, which serves as a sealing gasket and as a spacer between the sample and the glass slide. We have used the Teflon films with thicknesses 35 $\mu$m or 100 $\mu$m. The flow cell volume was 3.5 $\mu$L or 10 $\mu$L, correspondingly. The dead volume of the flow cell system was approximately 25 $\mu$L. Gravity flows of streptavidin or biotin solutions and pure PBS buffer were used with volumetric flow rate up to 1 mL/min.
**Angular resonance curves measurements.** The angular resonance curves in Fig. 2b were measured by focusing both parts of the splitted laser beam (with diameter $D\simeq 3$ mm) in the same spot on the structure surface with the objective of a focal length of $f=60$ mm, and detecting the intensity distribution of reflected light with a 512-pixels Hamamatsu photodiode array placed 385 mm$|$ 442 mm ($\rho_1|\,\rho_2$) apart from the structure.
**Data handling.** Data acquisition from the diode array, data processing and presentation were done with software we wrote on a personal computer running under Windows.
**RESULTS AND DISCUSSION**
We prepared the test photonic crystal structure and measured absolute angles of the PC SWs excitation at $\lambda=532$ nm and $\lambda=442$ nm wavelengths. In Fig. 1 a calculated dispersion of our 1D PC structure in water is presented as the logarithm of optical field enhancement (i.e., as $\lg[(E^{\phantom *}_eE^*_e)/(E^{\phantom *}_0E^*_0)]$) in the external medium near the structure. Good correspondence is seen between experimental points (black diamonds) and the calculated dispersion curves of the surface modes. The dispersion is presented in coordinate $\lambda(\rho)$, where $\lambda$ is an optical wavelength and $\rho$ is a numerical aperture $\rho=n_0\sin(\theta_0)$. Hereafter we use the numerical aperture $\rho$ as an angle variable instead of angles $\theta_j$ in different layers. This is a unified angle variable for all layers since, according to Snell’s law, $\rho=n_0\sin(\theta_0)=n_j\sin(\theta_j)$, for any layer $j$.
From Fig. 1, one can see that it is possible to excite one of the PC SWs in close proximity to the TIR angle from the water by appropriately choosing the laser wavelength and/or by appropriately choosing the PC structure. The penetration length of the evanescent wave intensity (i.e., $E^{\phantom *}_eE^*_e$) in the external medium, which is
``` math
\begin{equation}
l_e=\frac{\lambda}{4\pi \sqrt{\rho_1^2-\rho_{\mathsf{TIR}}^2}} \; ,
\end{equation}
```
may be very large for this mode if the difference $(\rho_1-\rho_{\mathsf{TIR}})=(\rho_1-n_e)$ is small. This is a unique property of PC SWs, because in any standard waveguide techniques [5, 8] the numeric aperture or (in other words) the effective RI of the waveguide mode $\rho_{\mathsf{mode}}$ is always more then a RI of the low refractive index spacer $n_{\mathsf{spacer}}$. Therefore the difference $(\rho_{\mathsf{mode}}-\rho_{\mathsf{TIR}}) \ge (n_{\mathsf{spacer}}-n_e)$ cannot be made small in the standard waveguides, taking into account the RI of the water ($n_e \simeq 1.33$) and the RI of the spacer (usually made from $SiO_2$, $n_{\mathsf{spacer}}\simeq 1.49$).
Another unique property of PC SWs is the possibility to excite them in the structure, where the final dielectric layer (the silicon oxide layer in our case) may have a low RI , while the standard waveguide has a high RI layer on a low RI spacer. This simplifies the procedures of biochemical modification of the external surface, which is now the standard $SiO_2$ surface.
In Fig. 2 the biosensor setup scheme and a typical raw experimental signal from the setup are shown. The Hamamatsu photodiode array was used to record the experimental signals – angles of two PC SWs ($\rho_1$ and $\rho_2$) simultaneously. The interference near resonance curves is the distinguishing feature of long-range PC SWs propagation. We observed similar interference in our recent work, [17] dealing with long-range surface plasmon-polaritons propagation. The appearance of such interference means that the propagation distance of the PC SW becomes much more than the waist of an incident Gaussian beam at the surface (this also may be easily seen on the sample surface by a naked eye). Note that the resonance peaks are very sharp (due to the long-range PC SWs propagation) and this allows measurement of the resonance peaks position change with high precision.
To verify the sensitivity of the biosensor and to compare it with existing label-free methods we present the unsmoothed experimental data of free biotin binding on the streptavidin monolayer. Initially (Fig. 3a), we present the build-up of the streptavidin monolayer on the biotinylated surface. Streptavidin (diluted in PBS to a concentration of 16 $\mu$g/mL) was run through the flow cell with volumetric flow rate 0.4 mL/min. Then the flow cell was rinsed by PBS. In Fig. 3a one can see that the adalayer thickness increases on 6.2 nm during streptavidin binding to biotinylated surface. The adalayer RI was chosen as $n_a=1.43$ (it is a free parameter in our program).
Fig. 3b presents the change of $d_a$ during free biotin binding to the streptavidin monolayer, while Fig. 3c shows RI changes of the analyte during these biotin injections. Biotin (diluted in PBS to a concentration of 3 $\mu$g/mL) was injected into PBS running through the flow cell with volumetric flow rate 0.6 mL/min. Fig. 3b shows that the streptavidin monolayer at first increases its thickness, but then contracts to a value slightly less than the initial one. At the same time, Fig. 3c shows that the external medium RI is not changed until the second biotin injection (from 1500 sec until 1600 sec $n_e\simeq$ const). So, in Fig. 3b, during this time period, we observe the act of streptavidin conformation while biotin molecules penetrate into streptavidin molecules. The second biotin injection did not result in the same streptavidin conformation, because most streptavidin subunits are already occupied by biotin molecules.
In compliance with the work [22] we suppose that the process of the free biotin-streptavidin binding is a good candidate for comparison of the signal/noise (S/N) ratio of different label-free techniques. We believe that for comparison of the S/N ratio it is also very important to point out the measurement time and the fact of posterior data averaging and/or smoothing (which increase the effective time of the measurement). In other words, the noise should be reduced to $1/\sqrt\mathrm{Hz}$ value. In our experiments the signal accumulation time was 1 second per point and no posterior data averaging or smoothing was done. The noise (i.e., standard deviation – std) of the thickness measurement was equal $\delta d =\mathrm{std}(d_a)\simeq 1.3~\mathrm{pm}/\sqrt\mathrm{Hz}$. The noise of the measurement of the external medium RI was $\delta n=\mathrm{std}(n_e) \simeq 5\cdot 10^{-7}~/\sqrt\mathrm{Hz}$. In Fig. 3b one can see that we detected the streptavidin conformation process during free biotin binding with an S/N ratio of about 15. We believe that the noise of the presented technique could be further decreased by improving the quality of the dielectric multilayer coating and by decreasing laser noise.
**CONCLUSIONS**
We have employed the two different optical modes on the photonic crystal surface for the optical sensing of biomolecular interactions. Unique properties of photonic crystals was used for the excitation of optical waves along the photonic crystal surface so that the evanescent field of one wave penetrates much deeper into the liquid volume. This wave is used as a reference for the RI of the liquid. The simultaneous registration of the two modes gives possibility to derive both the RI of the liquid and the adalayer thickness. This permited us to segregate the volume and the surface signals from the analyte, increase the sensitivity of biomolecule detection and record the act of streptavidin conformation during binding of biotin molecules.
**ACKNOWLEDGEMENT**
Authors thank S. Grachev for the kind donation of some biochemicals and for helpful advises about surface preparation. This work was partly supported by the European Network of Excellence, NMP3-CT- 2005-515703-2.
**References**
10 url urlprefix
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The calculated dispersion of the 7-layers PC structure in water and measured experimental points (black diamonds) at λ = 532 nm and λ = 442 nm laser wavelengths. The two optical surface modes are clearly seen as red curves (with an enhancement about 1000) inside the band gap (with an enhancement much less than 1). The biosensor scheme (a) and a typical raw experimental signal from the diode array (b). Immobilization of streptavidin on a biotinylated surface (a) and free biotin binding to the streptavidin monolayer: changes of the layer thickness (b) & RI of the buffer (c). The measurement time is 1 second per point (no posterior data averaging and smoothing). In color inserts the corresponding processes are illustrated.
## References
1. Cooper, M. A. Label-free screening of bio-molecular interactions . newblock *Anal. Bioanal. Chem. **377, 834�842 (2003).
2. Homola, J., Yee, S. S. & Gauglitz, G. Surface plasmon resonance sensors: review. newblock *Sensors and Actuators B **54, 3--15 (1999).
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4. Raether, H. newblock *Surface Plasmons (Springer, Berlin, 1988).
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6. Alieva, E. V. & Konopsky, V. N. Biosensor based on surface plasmon interferometry independent on variations of liquid's refraction index. newblock *Sensors and Actuators B **99, 90--97 (2004).
7. Cross, G. *et al. The metrics of surface adsorbed small molecules on the Young's fringe dual-slab waveguide interferometer. newblock *Journal of Physics D Applied Physics **37, 74--80 (2004).
8. Yablonovitch, E. Photonic band-gap structures. newblock *J. Opt. Soc. Am. B **10, 283--295 (1993).
9. Kossel, D. Analogies between thin-film optics and electron band theory of solids. newblock *J. Opt. Soc. Am. **56, 1434--1434 (1966).
10. Yeh, P., Yariv, A. & Hong, C.-S. Electromagnetic propagation in periodic stratified media. I. General theory. newblock *J. Opt. Soc. Am. **67, 423--438 (1977).
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13. Villa, F., Regalado, L., Ramos-Mendieta, F., Gaspar-Armenta, J. & Lopez-Rios, T. Photonic crystal sensor based on surface waves for thin-film characterization. newblock *Opt. Lett. **27, 646--648 (2002).
14. Palik, E. D. newblock *Handbook of Optical Constants of Solids (Academic, London, 1985).
15. Li, J. *et al. Assembly method fabricating linkers for covalently bonding DNA on glass surface. newblock *Sensors **1, 53--59 (2001).
16. Konopsky, V. N. & Alieva, E. V. Long-range propagation of plasmon polaritons in a thin metal film on a one-dimensional photonic crystal surface. newblock *Phys. Rev. Lett. **97, 253904 (2006).
17. Zybin, A. *et al. Double-Wavelength Technique for Surface Plasmon Resonance Measurements: Basic Concept and Applications for Single Sensors and Two-Dimensional Sensor Arrays. newblock *Anal. Chem. **77, 2393--2399 (2005). thebibliography.
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