---
title: "Photonic Crystal Biosensor Based on Optical Surface Waves"
authors: ["Valery N. Konopsky", "Tanya Karakouz", "Elena V. Alieva", "Chiara Vicario", "Sergey K. Sekatskii", "Giovanni Dietler"]
affiliation: "% Institute of Spectroscopy, Russian Academy of Sciences, Fizicheskaya, 5, Troitsk, Moscow region, 142190, Russia Laboratoire de Physique de la Mati `ere Vivante, Institut de Physique des Syst `emes Biologiques, Ecole Polytechnique F 'ed 'erale de Lausanne, CH-1015 Lausanne, Switzerland"
journal: "Sensors"
year: 2013
volume: "13"
issue: "2"
article_number: ""
pages: "2566--2578"
doi: "10.3390/s130202566"
type: journal-article
site_group: ""
url_abstract: ""
url_pdf: "https://valery.konopsky.com/kvnlocal/sensors-13-02566.pdf"
language: en
source_tex: "Z:\\ValeryData\\Valery_New\\my_articles\\Sensors2013\\Konopsky2Sensors.tex"
source_pdf: "Z:\\ValeryData\\Valery_New\\my_articles\\Sensors2013\\published\\sensors-13-02566.pdf"
---
## Abstract

A label-free biosensor device based on registration of photonic crystal surface waves is described. Angular interrogation of optical surface wave resonance is used to detect changes in the thickness of an adsorption layer, while an additional simultaneous detection of the critical angle of total internal reflection provides independent data of the liquid refractive index. This device is tested using binding of biotin molecules to a streptavidin monolayer, and by measuring association and dissociation kinetic of immunoglobulin G proteins. The deposition of PSS\,/\,PAN multilayer assembly is also recorded and separate thicknesses of PSS and PAN monolayers are calculated.

# Introduction

Photonic crystals (PCs) are materials that possess a periodic modulation of their refraction indices (RIs) on the scale of the wavelength of light [1]. The multiple reflection from the periodic RI boundaries in such materials can lead to the destructive interference of the optical waves and to the formation of bands where light propagation is forbidden. A simple periodic multilayer stack (dielectric mirror) is an example of one-dimensional (1D) PC structure. Such dielectric mirror can reflect all light within its forbidden band gaps, but also can support propagation of optical surface waves along the external border of the mirror. These photonic crystal surface waves (PC SWs) have important features that distinguish them from other waveguide waves and surface waves propagating along an interface under investigation.

The existence of optical SWs in the forbidden band gap of the PC may be deduced from an analogy between electron waves traveling in the periodic potential of the ordinary crystals and optical waves traveling in PC crystals [2]. In both cases, frequency intervals exist in which wave propagation is forbidden. This analogy may be extended to include surface levels, which can exist in band gaps of electronic crystals (i.e. Tamm states). In PCs, they correspond to optical SWs with dispersion curves located inside the photonic band gap. Sometimes these PC SWs are also called Bloch surface waves [3] or optical Tamm states [4].

Optical surface modes in 1D PCs were studied in the 1970s, both theoretically [5, 6] and experimentally [7]. Twenty years later, the excitation of optical SWs in a Kretschmann-like configuration was demonstrated [8, 9]. In recent years, PC SWs have been used in ever-widening applications in the field of optical sensors [12, 11, 10, 13]. In contrast to surface plasmon-polaritons (SPPs), both *p* -polarized and *s* -polarized optical surface waves can be used in PC SW sensor applications [10].

In the present article we describe a label-free biosensor device based on PC SWs, where *s* -polarization is used for detection of PC SW, which is sensitive both to adlayer thickness and RI of the liquid, while *p* -polarization is used for detection of a critical angle, which is sensitive to RI of the liquid only. The simultaneous registration of these two angles gives possibility to derive both the RI of the liquid and the adlayer thickness. This permitted us to segregate the volume and the surface signals from the analyte and increase the sensitivity of label-free biomolecule detection.

Label-free optical biosensors play a key role in the selective recognition of target biomolecules and in biomolecular interaction analysis, providing kinetic data about biological binding events in real time without labeling. The advantages of the label-free concept are the elimination of detrimental effects from labels that may interfere with fundamental interaction and the absence of time-consuming pretreatment [14, 15, 16, 17]. The disadvantages of all label-free techniques including the most mature one, the surface plasmon resonance (SPR) technique [18], are deficient sensitivity to a specific signal and undesirable susceptibilities to non-specific signals, e.g., to the volume effect of refraction index variations. These variations arise from temperature fluctuations and drifts, and they are the limiting factor for many state-of-the-art optical biosensors. To overcome these disadvantages, we describe the design, realization, and testing of the optical biosensor based on detection of PC SWs angle and critical angle.

# Experimental Section

## PC SW biosensor setup

The PC SW biosensor with an independent registration of the critical angle of total internal reflection (TIR) from the liquid is outlined in Fig. 1 (A). In this Figure a sketch of the biosensor and typical signals from the photodetector are shown. A laser beam from fiber-coupled diode laser ($\lambda=658$ nm) is sent to the sensor surface through a polarization-maintaining fiber cable (to improve the quality of a beam profile). The beam is focused by a cylindrical lens so that the excitation angle of one *s* -polarized PC SW (existing in this 1D PC) structure and TIR angle (in *p* -polarization) are contained in the convergence angle of the beam.

![A sketch of the biosensor based on angle interrogation of a PC SW. The typical reflection profile is shown near the CMOS matrix in (A) and is illustrated in (B) at different distances from the 1D PC. The angular resonance curves are shown in red for s -polarization, and in blue for p -polarization.](media/Sensors2013/Figure1.eps)

*A sketch of the biosensor based on angle interrogation of a PC SW. The typical reflection profile is shown near the CMOS matrix in (A) and is illustrated in (B) at different distances from the 1D PC. The angular resonance curves are shown in red for s -polarization, and in blue for p -polarization.*

Such types of PC SW sensors also possess one-dimensional spatial selectivity in a direction perpendicular to the plane of the Figure 1 (A) (i.e., along the focus line of the cylindrical lens). This fact permits recording of several reactions with an analyte simultaneously if different ligands are deposited on the PC in several linear target bands. In this way several tests can be performed at once that increases the throughput of the sensor.

## 1D PC structure

The desirable 1D PC structure may be theoretically deduced, for example, using a previously described impedance approach [19]. The following 1D PC structure was derived by this method and was used in experiments: *substrate* /$(LH)^{3}L'$/ *water*, where $L$ is a $SiO_2$ layer with thickness $d_1=183.2$ nm, $H$ is a $Ta_2O_5$ layer with $d_2=111.2$ nm and $L'$ is a $SiO_2$ layer with $d_3=341.6$ nm. The $SiO_2/Ta_2O_5$ 7-layers structure (started and finished by $SiO_2$ layers) is deposited by magnetron sputtering. The prism and the glass plate substrate are made from BK-7 glass. The RIs of the substrate, $SiO_2$, $Ta_2O_5$ and water at $\lambda=658$ nm, are $n_0=1.514$, $n_1=n_3=1.47$, $n_2=2.1$ and $n_e=1.331$, respectively.

## Angular resonance curves in reflection profile of the 1D PC structure

After reflection from the sensor surface, the reflection profile contains information about the TIR angle (transferred by the *p* -polarized part of the beam) and about the angle of the PC SW excitation (transferred by the *s* -polarized part of the beam). The reflection profile and fringe patterns near the resonance dip and TIR angle are illustrated in Figure 1 (B) for different distances from the prism with PC.

The fringe pattern, observed on the larger-angle side of the resonance dip, is the distinguishing feature of all long-range surface waves and a similar fringe pattern was observed not only with PC SWs [11], but also with ultra long-range SPPs [20, 21]. It may be easily inferred that for very large distances, where wavefront curvatures of both waves (reflected and reradiated from the surface) equalize, the fringe pattern will disappear, and only the resonance dip will be preserved [20]. Indeed on Figure 1 (B) one can see that this fringe pattern is more pronounced on the distance of 5 cm, less pronounced on the distance of 15 cm, and nearly disappears on the distance of 35 cm.

The highest curve in Figure 1 (B) is the reflection curve from the prism without multilayer coating (no PC). It is provided to illustrate that the sharpness of the curve near the $\mathrm{TIR}$ angle is much higher for our PC structure, than the one on the bare surface. Therefore, this setup also may be used as a critical-angle refractometer with the enhanced measurement precision, if only $\theta_\mathrm{TIR}$ angle is measured.

From Figure 1 (B) one can see that the fringe pattern appears not only near the SW resonance ($\theta_\mathrm{PC\,SW}=62.63$ grad), but also near the critical angle $\theta_\mathrm{TIR}=61.54$ grad. This fine interference pattern in the reflection profile of a focused laser beam is the result of enhanced Goos-Hänchen shift [22]. Both interference patterns, appearing near the critical angle of TIR and near the resonance dip of long-range surface waves are very important and useful for sensing applications, because the increase in number of points with large values of the first derivative near the dip and near the critical angle leads to an increase in sensitivity to changes in their position. In particular, subpixel registration of a fringe pattern shift becomes possible through detection of lots of these very sharp fringes, which are distributed on many CMOS matrix pixels.

The RI of the liquid may be derived as for classical critical-angle Abbe refractometers through the angle of total internal reflection $\theta_\mathrm{TIR}$. The liquid RI is then given by
``` math
\begin{equation}
   n_e=n_0\sin(\theta_\mathrm{TIR})\, ,
   
\end{equation}
```
where $n_0$ is the RI of the prism in which the critical angle $\theta_\mathrm{TIR}$ is measured. To derive changes in the adlayer thickness from changes in the resonance angle $\Delta\theta_\mathrm{SW}$ and $\Delta n_e$ (known from (1)), one can use the dispersion relation valid for both polarizations, which is derived in [19].

## Reagents

Hydrochloric acid (37%, HCl, Fluka); ethanol (ACS, Fluka); acetone; ethanolamine (ACS, Fluka); glycine (98%, Aldrich); toluene (99%, Fluka); toluene anhydrous (99.8%, AlfaAesar); 3-triethoxysylilpropylsuccinic anhydride (94%, TESPSA, ABCR); 3-aminopropyl triethoxysilane (APTES, 221.37 Da Sigma); N-hydroxysuccinimide (98%, NHS, Aldrich); Biotinamidohexanoyl-6-aminohexanoic acid N-hydro­xysuc­cini­mide ester (Biotin-X-X-NHS, 567.7 Da, Sigma); N,N-dimethyl­forma­mide (DMF, Sigma); free biotine (244 Da, Sigma-Aldrich); streptavidin (ca. 60 kDa, Sigma); N-(3-dimethylamnopropyl)-N’-ethylcarbodiimide (EDC, Fluka); phosphate buffered saline tablets (PBS, Sigma); polyallylamine hydrochloride (PAH, 58 kDa, Sigma-Aldrich); polystyrene sulfonic acid, sodium salt (PSS, 77 kDa, Fluka); sodium chloride (NaCl, Sigma); mouse IgG (Millipore); rabbit IgG (Sigma); goat anti-rabbit (Millipore) and goat anti-mouse IgG (Millipore) were used as received. Citric acid – disodium phosphate buffer solution (CAP, pH = 5.6) was prepared by mixing 42 ml of 0.1 M citric acid (Sigma) with 58 ml of 0.2 M disodium phosphatedodecanhydrate (Sigma).

## PC SW substrate cleaning and functionalization

Before functionalization PC SW substrates were sonicated one time in acetone and twice in ethanol (5 min each), followed each time by drying under a stream of nitrogen. After washing PC SW substrates were treated in plasma cleaner (HARRICK PLASMA PDC-32G) 1 min in medium power, air pressure ca. 400 mTorr. Then, these precleaned slides (with expected $OH$ bonds on the ultra-hydrophilic $SiO_2$ surface) were modified to obtain either biotinylated surface for streptavidin/biotin experiments or carboxylate-derivatized surface for experiments with IgG.

In the first case APTES was used to convert the $OH$-terminated $SiO_2$ surface to an $NH_2$-terminated one, and then Biotin-X-X-NHS dissolved in DMF was used to biotinylate the amino-terminated surface. Afterwards, the slides with a biotinylated surface were sequentially sonicated and thoroughly rinsed with DMF and PBS to remove any excess of Biotin-X-X-NHS.

In the second case, PC SW substrates were modified to obtain carboxylate functionalities [23, 24] following further procedure: overnight immersion in a 0.5% (v:v) TESPSA solution in anhydrous toluene in the glove box environment, followed by rinsing and sonication in toluene, acidic water (0.1 M HCl) and drying first under a stream of nitrogen and finally in low vacuum for ca. 30 min.

## Immobilization of the streptavidin

Bounding of the streptavidin to the biotinylated surface was monitored in situ by the PC SW biosensor. After a streptavidin monolayer formation on the the biotinylated surface (see Figure 2), the free biotin was used as a test to detect small molecule binding with the streptavidin monolayer (see Figure 3). All experiments were carried out in PBS (pH = 7.2).

## Immobilization of the IgG

All further procedures were performed while the carboxylate-derivatized PC SW substrate was mounted to the flow cell of the PC SW instrument allowing in situ monitoring of the adsorbed layer thickness and bulk RI. The flow rate during all the experiments was ca. 0.5 ml/min. Bioactivation of the PC SW substrates was obtained by covalent linking of biological ligands using amine coupling via reactive esters [25]. The standard amine coupling includes a three step reaction with EDC/ NHS chemistry: activation, ligand linking and deactivation. The activation of the surface was obtained by running of 0.1M EDC/ NHC solution in water (freshly prepared) for 5 min following by washing with a coupling buffer (CAP). This procedure was repeated 3 times to increase the activation yield. After the activation, the ligand (rabbit IgG or mouse IgG) was linked to the surface by running a 40 g/ml solution of it in the CAP buffer for ca. 15 min following by washing with the same buffer. After ligand immobilization the unreacted esters were neutralized by deactivation with 1M ethanolamine and washed with a coupling buffer for 5 min. Finally the bioactive interface was stabilized by alternate runs of 0.05M HCl and a coupling buffer for 5 min each, repeated 3 times.

## Polyelectrolyte multilayers

The polyelectrolyte assembly was carried out using the positive polyelectrolyte PAH and the negative polyelectrolyte PSS, 1 mM solutions (concentration calculated with respect to the monomer) in an aqueous solution of 0.1 M NaCl. The PC SW substrate was mounted to the flow cell of the PC SW instrument immediately after plasma treatment and amine functionalization was obtained by flowing 1% (v:v) APTES in water for 5 min to get positively charged surface. Following silanization the substrate was alternatingly exposed to PSS or PAH solutions for 10 min, starting with PSS. After each adsorption step the substrate was rinsed with water for 100 s, and with an aqueous solution of 0.1 M NaCl for 200 s, and exposed to next electrolyte solution.

# Results and Discussion

## Detection of free biotin binding to the streptavidin monolayer

To illustrate sensitivity of the PC SW biosensor device, we present the unsmoothed experimental data of free biotin binding on the streptavidin monolayer. As described in the experimental section and shown in Fig. 2, we initially present the buildup of the streptavidin monolayer on the biotinylated $SiO_2$ surface. Then, the flow cell was rinsed by fresh PBS solution (a region between dashed lines in Fig. 2) and biotin was injected into PBS, running through the flow cell (Fig. 3). In the color insets corresponding processes are illustrated.

![Immobilization of streptavidin on a biotinylated surface.](media/Sensors2013/Figure_strept_ins.eps)

*Immobilization of streptavidin on a biotinylated surface.*

Streptavidin (diluted in PBS to a concentration of $c_\mathrm{str}=12~\mu$g/ml) was run through the flow cell with a volumetric flow rate of $v_\mathrm{str}=0.3$ ml/min. Fig. 2 illustrates that the increase of the adlayer thickness due to immobilization of streptavidin on a biotinylated surface (top) occurs with kinetics different from those of the RI change of buffer during injection (bottom). This fact indicates that the volume and surface contributions from an analyte are indeed separated into different registration channels.

![Free biotin binding to the streptavidin monolayer.](media/Sensors2013/Figure_biot_ins.eps)

*Free biotin binding to the streptavidin monolayer.*

Figure 3 presents the adlayer thickness changes observed during free biotin ($M_\mathrm{biot}\simeq244$ Da) binding to the streptavidin monolayer and the RI changes of the analyte (bottom) during this biotin solution injection. Biotin (in a concentration of $c_\mathrm{biot}=0.9~\mu$g/ml) was injected into PBS running through the flow cell with volumetric flow rate of $v_\mathrm{biot}=0.4$ ml/min. It is clear that the biosensor can reliably detect the increase in streptavidin monolayer thickness upon free biotin binding.

## Polyelectrolyte multilayers deposition

The 22 layers of polyelectrolyte assembly were deposited, started with PSS and ended by PAN. After three initial bilayers, the total adlayer thickness increases linearly with each adsorbed layer and no sign of saturation is observed. Fragment of the deposition is presented in Figure 4(A) as the thickness growth of the polyelectrolyte assembly. The total layer thickness registered upon polyelectrolyte assembly is presented in Figure 4(B).

![A time slice of the thickness growth of the polyelectrolyte assembly (A) and the total layer thickness registered upon polyelectrolyte assembly for different RI of the adsorption layer (B).](media/Sensors2013/PSS_PAN2.eps)

*A time slice of the thickness growth of the polyelectrolyte assembly (A) and the total layer thickness registered upon polyelectrolyte assembly for different RI of the adsorption layer (B).*

At small adlayer thickness, all optical methods can measure the changes in optical thickness only (i.e., physical thickness of the adlayer multiplied by the adlayer RI: $d\!\times\!n_\mathrm{a}$). Therefore, some assumption about the adlayer RI (immersed in the liquid) is needed (if independent non-optical measurement of $d$ do not possible). In our measurement we used $n_\mathrm{a}=1.43$ and measure the the total thickness of the 22 layers equals 54.28 nm. Recalculation of the thickness for other $n_\mathrm{a}$ is also possible. As example, we recalculate the adlayer thickness, assuming that adlayer RI = 1.56, and receive the total thickness approximately as twice as small (24.27 nm). This thickness is in good correspondence with results reported in the work [26], where the thickness of the PSS / PAH bilayer was measured to be $2.09\pm0.03$ nm with assumption that $n_\mathrm{a}=1.56$.

To derive the PSS / PAH bilayer thickness more precisely, we should mention that initial PSS and PAH layers have a smaller thickness, probably due to incomplete assembling of the initial monolayers. Taking into account only complete monolayers (i.e., excluding the first three bilayers), we find that PSS monolayer thickness is $d^\mathrm{1.56}_\mathrm{PSS}=1.48\pm0.04$ nm and PAN monolayer thickness is $d^\mathrm{1.56}_\mathrm{PAN}=0.98\pm0.07$ nm, in assumption that $n_\mathrm{a}=1.56$, while in assumption that $n_\mathrm{a}=1.43$, we get: $d^\mathrm{1.43}_\mathrm{PSS}=3.41\pm0.07$ nm and $d^\mathrm{1.43}_\mathrm{PAN}=2.09\pm0.17$ nm.

That is, for $n_\mathrm{a}=1.43$, we measure the PSS / PAH bilayer thickness to be $d^\mathrm{1.43}_\mathrm{PSS+PAN}\simeq5.5$ nm. In the earlier works [27], the thickness of PSS / PAH bilayer was measured to be $5.1\pm0.2$ nm according to x-ray reflectometry, and $6.1\pm0.7$ nm (with adlayer RI $n_\mathrm{a}=1.50\pm0.05$), according to ellipsometry. Taken into account that the precise measurement of the adlayer RI for very thin films is hardly possible, it may stated that our thickness measurements are in good correspondence with previous ones. In addition to the previously published results, the separate determination of the thickness of each monolayer in the bilayer is possible.

## Ligand-receptor interactions

In this study the applicability of the PC SW optical sensor to characterizing ligand – receptor interactions is demonstrated by measuring association and dissociation kinetic of well-studied system of immunoglobulin G (IgG) proteins: rabbit and mouse IgG (ligands) and goat anti-rabbit and anti-mouse IgG proteins (receptors). The PC SW substrates were modified to get a biological recognition interface for IgG protein binding, as schematically shown in Figure 5 and described in details in the Experimental Section.

![Schematic representation of PC SW surface bioactivation and further receptor recognition.](media/Sensors2013/fig3_scheme.eps)

*Schematic representation of PC SW surface bioactivation and further receptor recognition.*

In Figure 6 a typical sensogram registered in situ by the PC SW optical sensor during ligand immobilization and further kinetic assay is shown (steps are in correspondence with a scheme represented in Figure 5). The changes in adsorbed layer thickness (upper panel) and bulk RI of the medium (bottom panel) are determined simultaneously during the experiment. In the first step, the free carboxyl groups on the surface was chemically activated using EDC/ NHS mixture as described in the Experimental Section. Then the ligand (Rabbit IgG) is immobilized following by ethanolamine deactivation of the remaining activated carboxyl groups and detachment of not covalently bound proteins. Finally a receptor binding to the surface modified with a ligand is observed following by IgG complex dissociation in PBS buffer and regeneration procedure. For simple biomolecular interaction analysis several receptor concentrations were bound to the immobilized ligand.

![Typical signal obtained by the PC SW optical sensor upon binding of the ligand (40 μg/ml Rabbit IgG) and the analyte (15 μg/ml anti-Rabbit IgG ).](media/Sensors2013/IgG.eps)

*Typical signal obtained by the PC SW optical sensor upon binding of the ligand (40 μg/ml Rabbit IgG) and the analyte (15 μg/ml anti-Rabbit IgG ).*

Figure 7 depicts normalized sensograms (change in the adlayer thickness as a function of time) for the binding affinity interaction between anti-rabbit IgG at various concentrations and immobilized rabbit IgG. The receptor was injected after baseline stabilization (shown on the graph). The buildup of the signal correlates with rabbit/anti-rabbit IgG complexes formed over time. The PC SW response increases with an increasing antibody concentration ($0.5$–$40\,\mu$g/ml). Further the weak dissociation of the complex is seen upon flowing of PBS buffer in the fluid cell.

![Normalized sensograms of anti-rabbit IgG binding to immobilized rabbit IgG.](media/Sensors2013/sensogram.eps)

*Normalized sensograms of anti-rabbit IgG binding to immobilized rabbit IgG.*

# Conclusions

The development of PC-based biochemical sensors is a fast growing area in recent years. The design flexibility of PC structures permits to devise appropriate sensors for any optical wavelength. The optical surface waves excited on the 1D PC interface is an effective means to guide and concentrate optical waves in the field of interaction between light and sensing material at the external side of the 1D PC. In the presented article we described the label-free biosensor device with 1D PC chips, where the bulk RI of analyte and the adlayer thickness are measured independently. The described here PC SW-based biosensor is commercially available as “EVA 2.0” device [28].

The exploitation of the 1D PCs as substrates supporting the long-range surface wave propagation permits researchers to:\
(1) increase the sensitivity of PC SW biosensors to the level $\delta d_a\simeq3\!\times\!10^{-13}~\mathrm{m}/\mathrm{Hz}^{1/2}$ (that corresponds to mass sensitivity $\delta m_a\simeq 0.3\,\mathrm{pg}/\mathrm{mm}^2$),\
(2) segregate surface and volume events in biosensing (that may be an important advantage in applications where temperature and composition of the liquid under study vary over a wide range),\
(3) enhance the detection of RI variation in the Abbe-like refractometer to the level $n_e\simeq10^{-7}~\mathrm{RIU}/\mathrm{Hz}^{1/2}$,\
(4) work with thick target ligands, such as living cells, with thickness up to 1$\mu m$\
(5) obtain one-dimensional spatial selectivity that makes multichannel registration possible and increases throughput of the sensor,\
(6) use the same PC chip many times, since a thick final $SiO_2$ layer may be effectively cleaned by some active treatment (e.g., in a plasma cleaner).

# Acknowledgements

TK, SKS and GD acknowledge the support of Swiss National Science Foundation, grant No 200021-137711. This work was also partially financially supported by the Science and Technology Cooperation Programme Switzerland–Russia.

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www.pcbiosensors.com

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