1. Introduction

Real time label-free optical transduction-based bio and chemical sensors have become increasingly popular in the last years [1]. Efficient methods for controlling the course of biological (chemical) binding/recognition events on the sensor surface normally employ evanescent waves, as in the cases of integrated planar [2] and fibre waveguides [3], or Surface Plasmon Resonance (SPR) biosensors [4]. The detection limit of commercial SPR devices with angular or spectral interrogation methods are about 1 × 10⁻⁶ refractive index units (RIU) and can be further improved down to 10⁻⁸ RIU using phase-sensitive SPR schemes [5]. However, the coincidence of the extremely sharp phase jump at SPR that provides very high sensor sensitivity with the minimum of reflected intensity poses certain limitation on the SPR method implementation. For this reason, TIR geometry, even with much smoother phase dependences, could be rather effective for the surface thin film investigation due to the fact that all the light intensity is totally reflected and available for the phase measurement. In addition, conventional SPR method usually cannot resolve the similar influence on the sensor response of the surface thin film formation or bulk RI changes that often occur during biological or chemical reactions. In order to overcome this drawback, additional surface characterization or more complicated SPR methods are required [6,7].

The main difference between SPR and TIR geometries consists in the presence of a thin gold film on the transparent substrate. For the bulk RI testing, SPR angular shift is slightly higher than the critical angle change, but TIR shows higher sensitivity in terms of change in reflected amplitude [8]. Absence of the surface metal film is also beneficial when the TIR method for surface testing is combined with transmission imaging for DNA chip, cell culture and microscopy investigation. Furthermore, as surface chemistry of glass and plastic is well known, TIR method could provide simple, efficient and low cost solution to portable biosensing.

In this paper, we propose polarimetric TIR biosensing set-up, where angular dependences of both phase and intensity of the reflected light allow to measure simultaneously the medium bulk refractive index and dynamic of deposition and thickness of biological thin films. The proposed methodology opens opportunities for the implementation of low-cost phase-sensitive TIR sensor designs for field and multi-sensing applications.

2. Instrumental methodology

In the proposed polarimetric set-up, light from a 5 mW stabilized He-Ne laser with 632.8 nm wavelength is passed through a Glan-Taylor polarizer to provide a 45° linearly polarized beam, as shown in Fig. 1 . The light is directed to a Photoelastic Modulator (PEM, HINDS I/FS50) after passing through a Soleil-Babinet compensator (SBC), which serves to optimize the initial phase retardation. The PEM is used to sinusoidally modulate the phase of the p-component at a frequency of 50 kHz. Due to the normal incidence of the light, no spatial displacement between s- and p- polarized component is introduced by PEM. However, static phase retardation in PEM should be considered in theoretical modeling [9]. The beam is then directed to a TIR sensing block that consists of a SF11 or BK7 coupling prism placed on the high precision θ-2θ goniometer (Thorlabs NR360S). In the experiments in air, a 45° prism was used whereas all tests in liquids were performed with a 60° prism connected to the flow injection cell with a 20 μl working volume. Glass slides with functionalized surface, in immersion contact (Cargile Lab) with the prism, were used for biosensing whereas in the simplified set-up the base of the prism was used for tests.

 

Fig. 1 Optical set-up for polarimetric measurement.

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The final periodic signal obtained by the detector with 45° oriented analyzer is decomposed into harmonics using a lock-in amplifier. Since the time domain signal is periodic and continuous, the harmonics of a frequency spectrum is modeled using a Fourier transform method. Thus, the first two harmonics are given by [10]:

Here, α is the initial phase retardance introduced by the SBC. R_p, ϕ_p and R_s, ϕ_s are reflection and phase under p- and s- polarization, respectively. J_n indicates a Bessel function of order n and M, the modulation amplitude. Phase retardation ϕ = ϕ_p-ϕ_s and intensity, I, can be derived from Eq. (1) and (2) and are given by [11]:

Using the Fresnel Formalism [12], we have performed theoretical modelling for the angular dependence of the first harmonic with different initial phase retardance α. The most characteristic curves are presented on Fig. 2(a) . We obtain for α = 0 a behaviour similar to the angular dependence of phase retardation between s- and p- component for TIR [12]. For the surface thin film testing in TIR configuration, application of phase difference as sensing parameter looks very promising [13]. At α = 90°, the curve is similar to the TIR angular intensity dependence where critical angle position follows the medium bulk RI. But, in this paper, we propose a method to test bulk RI as well as properties of thin films by measuring angular polarimetric curve (APC) obtained at α = -Δϕ _max, where Δϕ _max is the maximum possible phase difference at total reflection for given prism material. Indeed, this curve is a combination of the two parameters: intensity and phase, where angular position of the peak is sensitive only to the bulk RI and absolute value at the peak follows the dynamic of the surface film deposition (Fig. 2(b)). It is impossible to obtain similar results for SPR or TIR responses.

 

Fig. 2 (a) Theoretical angular polarimetric curve (APC) with different initial retardance α. (b) Comparison of APC (α=-Δϕ _max solid lines) with SPR (dashed lines). Black curves are for tests in water and red curves after 10⁻² RIU bulk changes; blue curves show influence of surface thin film deposition (10 nm, n=1.45).

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Figure 2(b) presents theoretical results for the biosensing application of the proposed polarimetric TIR geometry (solid curves) in comparison with SPR method (dashed curves). As one can conclude, angular shift of the APC peak is comparable with SPR response for a bulk refractive index change equal to 10⁻² RIU. Nevertheless, almost identical response could be obtained for SPR sensor after depositing a 10 nm thin film with 1.45 RI (dashed red and blue line in Fig. 2(b)). This makes very difficult to test independently bulk refractive index changes and formation of thin biofilms on the surface. In contrast, APC preserves angular peak position (after bulk change) and shows only intensity changes after thin film deposition (red and blue solid curves in Fig. 2(b)). Here we propose to use these properties for biosensing, where angular peak position corresponds to the bulk refractive index changes (X displacement) and integral amplitude changes of APC or absolute value at the peak follows the dynamic of surface thin films formation (Y displacement). Note that APC is sharper than SPR resonance curve and, as sensing parameter, we measure the angular position and absolute value at maximum amplitude, as opposed to the SPR method where position of the minimum value of the curve is detected. Finally, TIR method sensitivity could be optimized with different prism materials, where higher prism refractive index provides higher phase changes and correspondingly, higher thin film sensitivity but lower angular shift to the bulk RI changes [12].

3. Experimental set-up optimization and phase sensitivity

In order to estimate the efficiency of the polarimetric experimental set-up for the proposed TIR biosensing methodology, several calibration tests were performed. In the absence of the sensing block, we measured responses of the first and second harmonic to the small phase retardation introduced with a SBC. Optimal signal-to-noise ratio for the phase measurements (Eq. (3)) was obtained at 154° phase modulation depth (M), which provides equal Bessel function J ₁ = J ₂ that after tuning α results in equal harmonics intensities and therefore optimal conditions for the noise elimination (Fig. 3(a) black line). Phase response provides information about surface films and obtained high phase sensitivity in 5.2×10⁻³ deg. confirms efficiency of our polarimetric set-up that also could be used for direct phase measurements at fixed incident angle.

 

Fig. 3 (a) Phase sensitivity experiment (phase: black line, polarimetric signal: red line) (b) Experimental results for first harmonic with the same retardance as in Fig. 2(a).

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As follows from Eq. (1) and (2) for independent harmonic measurements, best results were obtained for phase modulation depth providing a maximum detected signal at M = 105° for the first harmonic and at M = 175° for the second. Due to the overall better performance of the electronic equipment at 50 kHz and better phase sensitivity equal to 0.02° for the first harmonic, all further tests were performed at M = 105° (Fig. 3(a) red line). Should be noted that while net phase measurements provides better sensitivity for the thin film testing, more complete biosensing information could be obtained using independent harmonic testing, where combined intensity-phase signal allows simultaneous investigation of bulk RI and surface thin films.

Experimental measurement has been performed to confirm our theoretical consideration shown in Fig. 2(a). The initial retardance α = Δϕ _max is a key component to obtain an optimal APC. Experimentally, the angular phase dependence was measured (Eq. (3)) and the system was fixed at an incident angle corresponding to the maximum phase difference between s- and p-polarization (for SF11 prisms: 43° and 68.59° in air and in water, respectively). By adjusting manually the SBC, the amplitude of the first harmonic of the polarimetric signal was set equal to zero and resulting APC was measured (Fig. 3(b), curve 3). Position of the SBC was then kept for others experiences. Polarimetric curve obtained with phase retardance equal to 0° and 90° are also shown in Fig. 3(b) (curve 1 and 2 respectively).

The sensitivity of the proposed TIR methodology depends on the precision of the APC peak amplitude and angular position determination. Angular resolution is provided by the high precision rotating stages while peak sharpness and half-width depends on angular phase retardance and requires application of a well collimated incident beam. In this work, collimated laser light (collimation test was performed on a distance 20 m with a precision 0.01 deg.) with 2 mm spot was used and corresponding APC is presented on Fig. 4(a) . Then, 0.1 deg. convergent or divergent beam was employed and equal decreasing of the APC peak amplitude accompanied with a peak broadening was obtained (Fig. 4(a)). These results show the importance of the proper set-up calibration and could be used as initial beam calibration method: maximum reflected amplitude at fixed angle near the peak corresponds to the condition of well collimated beam on the sensor surface.

 

Fig. 4 (a) Angular Polarimetric Curve with different beam propagation geometry. (b) Working distance minimizing the influence of the convergence/divergence near the focal point; Inset: Measured propagation Gaussian beam near the focal point.

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Also, we have found that using focused Gaussian beam gives APC close to the optimal (Fig. 4(a) red line). Explanation follows from the fact that Gaussian beam is quasi parallel at the focal point (inset in Fig. 4(b)) [14]. Indeed, in our case we employed a low NA optic (NA = 9×10⁻⁴) that results in a long Rayleigh range of 285 mm that provides nearly collimated beam with a 400 µm spot size and a flat phase at the point of incidence. The inset in Fig. 4(b) shows results of theoretical calculations and experimental data for a beam waist measured with a CCD beam shaper (Ophir-Spiricon). We have calculated the influence of the TIR coupling prism material on the working distance (Fig. 4(b)) where the beam divergence/convergence has an effect on the sensitivity lower or comparable to sensitivity value of 0.005 deg. (dashed line) for the phase and 0.02 deg. (continuous line) for the first harmonic). In this work, a SF11 prism was used that provides a rather long working distance (6.72 mm) whereas TIR system build on the Si will require positioning of the sensor surface in less than 1mm from the beam focal point. Thus, focused beam phase sensitive method is important in the design of portable micro fluidic multi-sensor or imaging systems where spot size provides resolution and sensitivity.

4. Results and discussion

To evaluate the sensing performance of the TIR polarimetric system and to estimate the sensitivity to the bulk refractive index changes, we have performed a series of experiments where aqueous solutions of NaCl with different RI [15] were introduced into flow–injection cell in contact with glass slides cleaned by “piranha” solution. Figure 5(a) presents sensor response as the angular position of the APC peak was measured in real time. For numerical analysis of the obtained APC, polynomial curve fitting method of forth order was used. Taking into account the level of noises (1.6×10⁻³deg.), the experimental detection limit for the bulk RI was estimated better than 10⁻⁵ RIU, which is comparable to SPR sensitivity (10⁻⁵~10⁻⁶ RIU).

 

Fig. 5 (a) Angular variation of the peak position for different refractive index solutions. (b) Responses of the system with different concentrations of Bovine Serum Albumin.

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To demonstrate the proposed biosensing methodology, we have performed experimental tests where adsorption of the thin biofilms on the surface is accompanied by the bulk RI changes. In order to form a thin biolayer on the sensing surface, different concentrations of Bovine Serum Albumin (BSA) in the phosphate buffered saline (PBS) were prepared and introduced into flow-injection cell. Glass slides with different surface functionalization in immersion contact with the SF11 prism were used as sensing surfaces. Clean glass was first tested and only bulk refractive index change was detected. Indeed, the existence of an electrostatic repulsion between the protein and the clean glass surface [16] prevents BSA layer deposition. Then, experimental results for 20 µg/ml and 50 µg/ml concentrated BSA were obtained on the glass functionalized with 3-(aminopropyl)-triethoxysilane (Sigma-Aldrich) [17], where surface amino-groups promote the non-specific binding of the BSA (Fig. 5(b)). Here the black line represents data for the APC peak maximum value corresponding to the thin biolayer formation and the red line shows the angular position of the peak maximum which is linked to the bulk refractive index of the solution. These information (bulk and thin layer) are completely different in their dynamic (change in the red line are quicker than black line which follows binding dynamic of BSA on amino-group) and in their level (continuous line does not return to the initial baseline due to the adsorption of BSA compared to the red one where solution returns to PBS). We can conclude from this figure that both bulk and thin layer information could be obtained independently.

The proposed methodology enables to control the thin layer formation and bulk RI changes independently in contrast to the SPR sensor where usually the combination of responses is monitored. Sensitivity of the system could be estimated by considering the value of 1.45 of the BSA monolayer refractive index [18]. Using out theoretical prediction for 0.02 deg. phase resolution, the experimental detection limit is estimated to an average thickness of 0.5 nm biolayer and could be improved for direct phase measurements.

5. Conclusion

In this paper, we demonstrated polarimetric biosensing TIR system based on the method of temporal phase modulation. Measurements of the phase difference between s- and p- polarized light combined with their amplitudes allow to obtain information about medium refractive index changes as well as dynamics and thickness of the deposited thin biofilms. Obtained phase resolution of 0.02 deg. corresponds theoretically to a sensitivity of 0.5 nm in biolayer thickness and tested sensitivity for the bulk refractive index is better than 10⁻⁵ RIU. The higher refractive index platforms such as high refractive index dielectrics and silicon offer even higher sensitivities for biofilms tests. Microfabrication technology on the glass, optical polymers or silicon enables construction of low-cost integrated biosensing systems.