---
title: "A biosensor based on photonic crystal surface waves with an independent registration of the liquid refractive index"
authors: ["Valery N. Konopsky", "Elena V. Alieva"]
affiliation: "Institute of Spectroscopy, Russian Academy of Sciences, Troitsk, Moscow region, 142190, RUSSIA."
journal: "Biosens_refract"
year: 2010
volume: "25"
issue: "5"
article_number: ""
pages: "1212--1216"
doi: "10.1016/j.bios.2009.09.011"
type: journal-article
site_group: "Biosensors based on optical surface modes"
url_abstract: "https://valery.konopsky.com/paper/BB2010/BB2010.htm"
url_pdf: "https://valery.konopsky.com/kvnlocal/BIOS3466.pdf"
language: en
source_tex: "Z:\\ValeryData\\Valery_New\\my_articles\\Biosens_refract\\Manuscript\\Konopsky2BB.tex"
source_pdf: "Z:\\ValeryData\\Valery_New\\my_articles\\Biosens_refract\\Published\\BIOS3466.pdf"
---
## Abstract

A high-precision optical biosensor technique capable of independently determining the refractive index (RI) of liquids is presented. Photonic crystal surface waves were used to detect surface binding events, while an independent registration of the critical angle was used for accurate determination of the liquid RI. This technique was tested using binding of biotin molecules to a streptavidin monolayer at various biotin concentrations. The attained baseline noise is $5\!\times\!10^{-13}~\mathrm{m}/\mathrm{Hz}^{1/2}$ for adlayer thickness changes and $9\!\times\!10^{-8}~\mathrm{RIU}/\mathrm{Hz}^{1/2}$ for RI changes.

Label-free optical biosensors ,Photonic crystal surface waves ,Critical-angle refractometry ,Biotin–-streptavidin binding

# Introduction

Registration of a bounded optical wave propagating along a bioactive surface is the most popular method used in state-of-the-art direct optical biosensor techniques [2]. In the surface plasmon resonance (SPR) technique [4], this surface-bound wave is a surface plasmon-polariton propagating along a gold or silver surface, while in optical waveguide techniques this wave is a waveguide mode excited in a high refractive index dielectric layer either via the frustrated total internal reflection (TIR) from a low refractive index spacer [1] or via a grating coupler, like it takes place in the integrated-optical waveguide technique [13, 11]. For each of these techniques, an evanescent field produced by the optical wave (with penetration depth in water $\sim\!100$ nm) is sensitive not only to surface-bound biomolecular interactions but also to changes in the volume refractive index (RI) caused by variations in liquid temperature and composition. Therefore, a need exists for a biosensor technique that would be able to segregate the volume and the surface contributions from analytes in the 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 [8] we used two photonic crystal surface waves (PC SWs) with large differences in penetration depths in order to simultaneously determine the two desired parameters: liquid RI – $n_e$ and the adlayer thickness – $d_a$. It should be noted here that it is impossible to obtain a large penetration mode depth in standard waveguides, where a RI of the medium under the waveguide film is larger than the RI of the liquid above the film [8]. Therefore, standard waveguides are limited in their ability to reliably segregate volume and surface contributions from an analyte, because both modes have a similar (small) penetration depth in a liquid ($\le\!100$ nm). A large penetration depth in a liquid ($>\!1\,\mu$m) can only be obtained using so-called “reverse” waveguides, where RI of the medium under the waveguide film is smaller than the RI of the liquid above the film [5, 6]. One-dimensional photonic crystal (1D PC) structures are particularly advantageous as they possess the similar feature (in their band gap regions) as substrates in the “reverse” waveguide, despite the fact that a PC structure consists of media with RIs larger than the RI of the liquid (e.g., seven alternative layers of $SiO_2$ and $Ta_2O_5$). The large difference among penetration depth of PC SWs in liquids is the main advantage of biosensors based on dual PC SWs [8, 9]. This attribute allows for (reliable) segregation of the volume and the surface contributions from an analyte and increases sensitivity of molecule detection. Nevertheless, some doubts about the validity of target values ($n_e$, $d_a$) derived and about their mutual interdependence may still arise, since each $n_e$ and $d_a$ values depend on both observed PC SW excitation angles and the both calculated target values $n_e$ and $d_a$ are model-based.

In the study presented here we use a direct measurement of the critical TIR angle to immediately obtain the RI $n_e$ of the liquid under investigation. This approach allows one unknown target value – $n_e$ to become model-independent, ultimately increasing the reliability and integrity of results.

# Materials and methods

## PC SW biosensor

![A sketch of the biosensor. The typical reflection profiles are shown near the photodiode array.](media/BB2010/Fig1.eps)

*A sketch of the biosensor. The typical reflection profiles are shown near the photodiode array.*

A novel biosensor with an independent registration of the critical angle of the liquid is employed in this study. In Fig. 1 a sketch of the biosensor and typical signals from the photodiode array are shown. A circular-polarized laser beam from He-Ne laser ($\lambda=632.8$ nm) is sent to the sensor surface through a polarization-maintaining fiber cable (to improve the quality of a beam profile). The beam is split to excite two *s* -polarized PC SWs existing in this 1D PC structure. After reflection from the sensor surface, the reflection profile of the first beam (which is close to the TIR angle) is recorded for both *s* - and *p* -polarizations, while the second one is recorded for *s* -polarization only. The reason for recording the additional *p* -polarization of the first beam is its use for direct registration of the critical TIR angle of the liquid, which immediately provides us the liquid RI. Moreover, the sharpness of the reflection near the critical angle and the measurement precision of the liquid RI herein are much high than the ones in standard critical-angle refractometers on uncoated prisms [7].

## Photonic crystal structure

The following 1D PC structure is used in experiments: *substrate* /$(LH)^{3}L'$/ *water*, where $L$ is a $SiO_2$ layer with thickness $d_1=186.4$ nm, $H$ is a $Ta_2O_5$ layer with $d_2=115.2$ nm and $L'$ is a $SiO_2$ layer with $d_3=776.8$ 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 $\lambda=632.8$ nm, are $n_0=1.515$, $n_1=n_3=1.47$, $n_2=2.07$ and $n_e=1.332$, correspondingly.

## Reagents

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 [10]. Biotin-X-X-NHS [Biotinamidohexanoyl-6-aminohexanoic acid N-hydroxysuccinimide ester; molecular weight – 567.7] dissolved in DMF [N,N-dimethylformamide] was used to biotinylate the amino-terminated surface. The streptavidin from Amersham (UK) [molecular weight $M_\mathrm{str}\sim 60\,000$] was deposited on the biotinylated surface. The free biotin [vitamin H; molecular weight $M_\mathrm{b}=244.31$] was used as a test to detect small molecule binding with a streptavidin monolayer. All experiments were carried out in PBS [phosphate-buffered saline; pH=7.2].

## Sample preparation

Samples were prepared as follows: first, glass slides were processed by ionized water vapor in a plasma cleaner for 5 minutes. Next, these precleaned slides (with expected $OH$ bonds on the ultra-hydrophilic $SiO_2$ surface) were immersed in 1% aminosilane solution in 95% acetone/water for 5 min. The glass slides, now with expected $NH_2$ bonds on the ultra-hydrophobic $SiO_2$ surface, were then dried by argon and baked in vacuum for 20 min at $50^0$C. To biotinylate the $NH_2$-terminated surface the slides were left overnight in solution, where Biotin-X-X-NHS was dissolved by DMF to concentration 12 mg/mL and then the PBS was added (1:1). Afterwards, the slides with biotinylated surface were sequentially sonicated and thoroughly rinsed with DMF and PBS to remove any excess of Biotin-X-X-NHS.

## Flow cell

The flow cell was made from a glass slide with two holes through which two glass tubes were fitted to serve as inlet and outlet, respectively. An internal surface of this glass slide was frosted to avoid reflection causing an additional interference of the refracted beam (at $\theta<\theta_\mathrm{TIR}$). The inlet tube was connected to a small tank filled with the solution under investigation. Flow velocity was controlled by an elevation difference of the inlet tank and the outlet end level. The height of the cell was determined by the thickness of a Teflon film, which served as a sealing gasket and as a spacer between the sample and the glass slide. We used the Teflon film with 35 $\mu$m thickness and corresponding flow cell volume was 3.5 $\mu$L. 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 rates up to 1 mL/min.

## Data handling

Data from the photodiode array was acquired, processed and presented using homemade software. Changes of the critical angle position $P_0$ and the resonance peak positions $P_i$ $_{(i=1,\,2)}$ on the photodiode array were converted to changes of the angle parameters $\rho_0$ and $\rho_i=n_0\sin(\theta_i)$.

The RI of the liquid was derived as in classical critical-angle Abbe refractometers, through the angle of total internal reflection $\theta_0=\theta_\mathrm{TIR}$. The liquid RI is then given by
``` math
\begin{equation}
   n_e=\rho_0=n_0\sin(\theta_0)\, ,
   
\end{equation}
```
where $n_0$ is the RI of the prism in which the critical angle $\theta_0$ is measured. To derive the changes of the adlayer thickness from the changes of the resonance angle $i$ and $\Delta n_e$ (known from (1)), we used the next relation:
``` math
\begin{equation}
\Delta d_a  =
 \frac { \Delta\rho_i - ({\partial \rho_i}/ {\partial n_e})\Delta n_e}
 {({\partial \rho_i}/{\partial d_a})}
 \; .
\end{equation}
```
The coefficients $({\partial \rho_i}/ {\partial n_e})$ and $({\partial \rho_i}/{\partial d_a})$ are able to be obtained, for example, from a theoretical simulation of the real 1D PC structure. The $\Delta n_e$ and $\Delta d_a$ data presented below were derived using eqs. (1) and (2), with the coefficients, which mean values are approximately equal: $<\!({\partial \rho_2}/{\partial n_e})\!>\simeq0.053$, $<\!({\partial \rho_2}/{\partial d_a})\!>\simeq0.057$ [1/$\mu$m] (precise values slightly depend on $n_e$ and $d_a$ and were derived for each particular point).

# Results and discussion

## Streptavidin monolayer deposition

![Immobilization of streptavidin on a biotinylated surface (top left) with subsequent binding of free biotin to this streptavidin monolayer (top right) and corresponding changes of RI of the buffer during these injections (bottom). The measurement time was 1 second per point (no posterior data averaging and smoothing). In color inserts corresponding processes are illustrated.](media/BB2010/Fig2.eps)

*Immobilization of streptavidin on a biotinylated surface (top left) with subsequent binding of free biotin to this streptavidin monolayer (top right) and corresponding changes of RI of the buffer during these injections (bottom). The measurement time was 1 second per point (no posterior data averaging and smoothing). In color inserts corresponding processes are illustrated.*

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 $v_\mathrm{str}=0.3$ mL/min. Figure 2 illustrates that the increase of the adlayer thickness due to immobilization of streptavidin on a biotinylated surface (top left) occurs with kinetics different from those of the RI change of buffer during injection (see bottom left). This fact indicates that the volume and surface contributions from an analyte are indeed separated into different registration channels.

![The injection of a high concentration of free biotin: changes in the thickness of the streptavidin–biotin complex (left) and the corresponding change of RI of the buffer (right). In color insert the possible corresponding process is illustrated.](media/BB2010/Fig3.eps)

*The injection of a high concentration of free biotin: changes in the thickness of the streptavidin–biotin complex (left) and the corresponding change of RI of the buffer (right). In color insert the possible corresponding process is illustrated.*

## Biotin binding to the streptavidin monolayer

The right side of Fig. 2 presents the adlayer thickness changes observed during free biotin binding to the streptavidin monolayer (top right) and the RI changes of the analyte (bottom right) during this biotin solution injection. Biotin (in a concentration of $c_\mathrm{b (low)}=0.9~\mu$g/mL) was injected into PBS running through the flow cell with volumetric flow rate $v_\mathrm{b}=0.4$ mL/min. One can see that the biosensor reliably detects the increase in streptavidin monolayer thickness at free biotin binding ($\Delta d_a\simeq 0.45\,\mathrm{\AA}$).

It should be noted here that the behaviour of the streptavidin-biotin complex appeared to depend on the concentration of the injected biotin solution. In several published experimental analyses, the thickness of the streptavidin-biotin complex was slowly and monotonically decreased after an initial sharp increase [14, 8]. We found that such behaviour may take place in the presence of an excess concentration of free biotin. To illustrate this point, we injected free biotin at high concentration $c_\mathrm{b (high)}=80~\mu$g/mL in the flow cell. Indeed, after a short initial increase, the thickness of the streptavidin monolayer was monotonically decreased as shown in Fig. 3. One possible explanation for this process may be the fact that free biotin (at high concentrations) acts as a competitor for streptavidin sites that are bound to the surface [3], i.e., elution of the streptavidin molecules from the surface may be taking place.

The detected increase in the physical thickness of the streptavidin monolayer $\Delta d_a\simeq 0.45\,\mathrm{\AA}$ (upon injection of low biotin concentrations) is in agreement with previously published data, where an increase in optical thickness (i.e., physical thickness of the adlayer multiplied by the adlayer RI) $\Delta (d_a n_a)\simeq 0.6\div0.7\,\mathrm{\AA}$ was measured by reflectometric interference spectroscopy technique [12]. Taking into account the refractive index of the adlayer $n_a=1.43$, which was used in our computation, we obtain a reasonable agreement in measured values. It may be noted that the system presented here, wherein three independent angle parameters are recorded, may also be used to experimentally determine the adlayer RI as the third calculated value (together with the adlayer thickness and the RI of the liquid). However, this type of adlayer RI determination yields precise ($\le 0.01$) and reliable values only when considering rather large adlayer thicknesses. There is a rather fundamental reason behind this limitation, which will be discussed elsewhere.

# Conclusions

In this paper we have proposed and experimentally tested a new biosensor technique capable of direct and model-independent determination of the RI of a liquid. The liquid RI is determined by measuring the critical angle for *p* -polarization of the laser beam, while the *s* -polarization is used for the excitation of the adlayer-thickness-sensitive PC SW. The attained noise floor of the RI baseline is $\delta n_e=9\!\times\!10^{-8}~\mathrm{RIU}/\mathrm{Hz}^{1/2}$, while the baseline noise of the adlayer thickness is $\delta d_a=5\!\times\!10^{-13}~\mathrm{m}/\mathrm{Hz}^{1/2}=0.5~\mathrm{pm}/\mathrm{Hz}^{1/2}$. The potential to record surface and volume events independently may be an important advantage in a number of applications where temperature and composition of the liquid under study may vary over a wide range.

# Acknowledgments

This work was financially supported by the Grant 09-02-00366-a from the Russian Foundation for Fundamental Researches. The authors would like to thank D. Klinov for his help in sample preparation.

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