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In this paper, unique surface sensing property and enhanced sensitivity in microring resonator biosensors based on subwavelength grating (SWG) waveguides are studied and demonstrated. The SWG structure consists of periodic silicon pillars in the propagation direction with a subwavelength period. Effective sensing region in the SWG microring resonator includes not only the top and side of the waveguide, but also the space between the silicon pillars on the light propagation path. It leads to greatly increased sensitivity and a unique surface sensing property in contrast to common evanescent wave sensors: the surface sensitivity remains constantly high as the surface layer thickness grows. Microring resonator biosensors based on both SWG waveguides and conventional strip waveguides were compared side by side in surface sensing experiment and the enhanced surface sensing capability in SWG based microring resonator biosensors was demonstrated.
© 2016 Optical Society of America
1. Introduction
Micro- and nano-scale photonic biosensors have become a fast growing research topic driven by the need of portable bio-detection systems with high sensitivity, high throughput, real-time and label-free detection [1–3]. Various devices, including surface plasmon devices [4–6], microring resonators [7–9], silicon nanowires [10], nanoporous silicon waveguides [11], one-dimensioanl (1D) and two-dimensional (2D) photonic crystal (PC) microcavities [12–15], have been proposed and demonstrated. Most of the proposed structures are based on the interaction between the evanescent wave and the biomolecules that are absorbed or immobilized on the sensor surface. For example, microring resonators built on silicon-on-insulator (SOI) substrate can detect biomolecule layers immobilized on the surface of the microring through the induced resonance shift in the transmission spectrum [7]. Extensive efforts have been made on this type of sensors focusing on increasing the sensitivity and lowering the detection limit [16–20]. However, this type of evanescent wave sensing mechanism faces limitation in surface sensing: the sensitivity drops inevitably with increasing thickness of the surface layer accumulated on the sensor surface. In real applications, this layer includes necessary oxide and chemical layers generated by surface treatment, probe proteins, target proteins and any other reagents that are used to enhance the signal. These can amount to a total layer thickness ranging from several nanometers to a few tens of nanometers, within which the sensitivity of the evanescent wave could drop considerably before it reaches the final target to be detected [21–25].
Recently, novel subwavelength grating (SWG) based waveguides and photonic devices were proposed and demonstrated [26–29]. The SWG waveguide consists of periodic silicon pillars in the propagation direction with a period much smaller than the operating wavelength. Within such a structure, the wave propagates in a similar way to conventional strip waveguides, but the interaction region between light and the cladding materials is greatly extended compared to the aforementioned evanescent wave based biosensors. Therefore, SWG structure shows great promise in integrated optical biosensors. In [24,30], microring resonators based on SWG waveguides were first demonstrated with bulk sensitivity (the ratio of resonance shift to the change of surrounding refractive index) greater than 400 nm/RIU, which is several times higher than conventional microring resonators based on strip waveguides [7]. However, the enhanced surface sensing capability in SWG waveguide, which is a unique advantage in this structure, has never been revealed and carefully studied.
In this paper, we analyzed and demonstrated enhanced surface sensing capability in microring resonator biosensor based on SWG waveguide structures. In the SWG waveguides, effective sensing region includes not only the top and side of the waveguide, where evanescent wave exists, but also the space between silicon pillars on the light propagation path. This leads to greatly increased sensitivity as well as a unique property of thickness-independent surface sensitivity in comparison to conventional microring resonator biosensors. The surface sensitivity (the ratio of resonance shift to the change of surface layer thickness) remains constantly high in SWG microring resonator even when surface layer thickness grows. Simulation shows that the surface sensitivity remains around 1.0 nm/nm in the first 25-nm thick layer upon the surface in the studied case. In the experimental demonstration, microring resonator biosensors based on both SWG waveguides and conventional strip waveguide were fabricated and compared side by side in a biosensing experiment. Special tuning of the pillar shape in the SWG was utilized to minimize the bending loss in the SWG microring and a high quality factor of 9100 in water environment was achieved. A comparison between the two types of sensors in the biosensing test verified the superior surface sensing capability in the sensors constructed by SWG waveguides. The SWG microring sensor was also used to detect microRNA with concentration as low as 1 nM.
2. Simulation and analysis
The structure of the SWG microring resonator is shown in Fig. 1(a). It is constructed by replacing the strip waveguides in a regular microring resonator with SWG waveguides. The silicon SWG microring resonator sits on top of the buried oxide layer and is covered by sensing medium (water or other biological buffers, assuming refractive index n = 1.32). Figure 1(b) shows top view of the ring-waveguide coupling region. In this paper, the SWG in the microring resonator uses trapezoidal pillars to minimize bending loss, thus to achieve better quality factors [31]. The SWG in the bus waveguide still uses regular rectangular pillars [24].
Fig. 1 (a) Schematic of the studied SWG microring resonator biosensor; (b) Top view of the coupling region (yellow rectangular in (a)); (c) Cross section view of the SWG waveguides (purple cut line in (a)).
The structural parameters of the SWG microring are also marked in Fig. 1. The radius of the microring R is set to 10 μm to achieve high intrinsic quality factor and compact size at the same time. The grating period of the SWG is Λ = 200 nm, which satisfies the condition (λ = 1550 nm), so the waveguide operates in the subwavelength regime and behaves like a conventional waveguide [27]. Duty cycle of the grating (ratio of silicon pillar length l to grating period Λ) η and waveguide width w are determined through simulation to achieve large optical field overlap with the sensing medium. The gap between microring and the bus waveguide is d = 50 nm, which is determined by a parameter scanning in fabrication to achieve high quality factor and extinction ratio. For the trapezoidal pillars in the SWG microring, the grating period are the same as the straight waveguide. Widths of the top base (a) and bottom base (b) are determined by minimizing the bending loss in the microring using 3D finite difference time domain (FDTD) simulation [32]. The height of the silicon layer (see Fig. 1(c)) is h = 220 nm.
To find a proper duty cycle η and waveguide width w, the optical mode profile is simulated using the 3D plane wave expansion (PWE) method and the overlapping factor σ (defined as the ratio of the electric field inside the low refractive index medium region) is calculated. The results are plotted in Fig. 2, in which the waveguide width and duty cycle are scanned in x and y axis, respectively. The region inside the dashed blue curve indicates modes that are above the silicon dioxide light line and are not well-confined. The absorption loss of light in water also contributes to the total loss in the SWG microring and deteriorates its quality factor. To achieve a high overlapping factor while taking absorption loss into consideration, η = 0.65 and w = 450 nm are chosen, resulting in the calculated overlapping factor σ = 0.4. So the length of the rectangular pillar is l = ηΛ = 130 nm. The top and bottom bases (a and b) are then determined through the bending loss simulation as described above. The optimized a and b are 100 nm and 150 nm, respectively.
The mode profile simulated with the above parameters is shown in Fig. 3. Figure 3(a) illustrates the schematic of the SWG structure and position of cut planes 1-3 where the mode profiles in Fig. 3(b)-(d) are obtained. Figure 3(b) is the electric field distribution at the middle height of the pillar (xy plane at y = h/2); Fig. 3(c) shows the electric field between pillars (xy plane at z = constant); Fig. 3(d) shows the field between pillars cutting close to the edge of the pillar (yz plane at x close to edge of the pillar). From the mode profile, it can be seen that in contrast to the evanescent field on the top surface and sidewalls of the waveguide, there is significantly stronger mode field existing on the light propagation path between silicon pillars. In Fig. 3(d), it is especially clear that the field is strongly confined between silicon pillars. This gives SWG based microring biosensors extended surface sensing region on the propagation path and thus unique advantage in surface sensing over conventional microrings.
Fig. 3 Electric field intensity distribution at different cross sections. (b), (c) and (d) correspond to the cut positions marked with 1, 2 and 3 in the schematic (a), respectively.
In a microring resonator biosensor, the resonance wavelength shifts as the biomolecules interact with the optical field on the surface of the microring. Specific biomolecules can be detected by immobilizing them on the surface through specific biochemical interactions. Figure 4(a) shows a schematic of this scenario for a sensor based on the SWG structure. The surface layer can add up to a thickness from a few nanometers to several tens of nanometers [21–25], including silicon dioxide layer (~5nm), chemical layer (several nanometers, generated after surface treatment with (3-Aminopropyl) triethoxysilane (APTES), glutaraldehyde, etc.) and protein layers (antibodies, antigens, etc., several nanometers per layer). Therefore, surface sensitivity is an important figure of merit. In resonance based sensing method, surface sensitivity Ss can be defined as the resonance wavelength shift in according to the change of surface layer thickness [30]:
where ng is group index and t is the thickness of surface layer. As shown in Fig. 4(a), we assume the sensing medium is water (n = 1.32) and the surface layer has uniform thickness across the surface with uniform refractive index of n = 1.48 [24,27,33]. Then the susceptibility in the periodic SWG structure is calculated from effective index () simulation using the 3D PWE solver. The SWG structure has the same parameters as presented above. in conventional strip waveguide (w = 500 nm, h = 220 nm, used to form regular microring resonator) is calculated in an eigenmode solver using finite element method (FEM). The simulation results are shown in Fig. 4(b). The in SWG waveguide is 4-6 times larger than that in a regular strip waveguide due to large mode overlapping factor (σ ~0.4). Furthermore, the value remains constantly high in SWG structure for the first 25nm of the surface layer, while in regular strip waveguide, drop monotonically with the accumulation of surface layer. This simulation result coincides with the above mode profile analysis that the field confined between silicon pillars extends the surface sensing region and the surface sensitivity becomes insensitive to surface layer thickness. It shows that SWG structure has superior surface sensing capability over evanescent wave based sensors like conventional microring resonator, in terms of both absolute surface sensitivity Ss and the ability to maintain high surface sensitivity when the thickness of surface layer grows.Fig. 4 (a) Schematic of the SWG structure covered by thin layers of silicon dioxide, chemicals and immobilized protein in water environment; (b) Comparing dneff/dt as the thickness of surface layer grows in SWG microring and conventional microring.
3. Experimental methods
3.1 Device fabrication and measurement
Devices were fabricated on a silicon-on-insulator (SOI) wafer with 220 nm thick top silicon layer. The wafer was first patterned by e-beam lithography with ZEP 520A resist (Zeon Chemicals). Then a single reactive ion etching (RIE) process using HBr and Cl2 is used to transfer the pattern to silicon. Resist was then removed. Before biosensing test, the devices went through thermal oxidation at 950 °C for 5 min to grow about 5 nm silicon dioxide on the surface. The silicon dioxide layer serves as the base for subsequent chemical treatment steps used to immobilize probe proteins covalently.
Devices were characterized on an optical test platform integrated with microfluidic system [24]. Transmission spectra of the fabricated devices were obtained using broadband LED light source and optical spectrum analyzer. Light was coupled into and out of the device with fiber arrays. The test stage is integrated with microfluidic channels in which sensing medium is pumped and flowed onto the sensor at a controlled flow rate. The stage is also thermally controlled with Newport 3040 Temperature Controller to avoid temperature-induced resonance shift in the biosensors. The whole system is connected to a computer and the transmission spectra of the device is monitored automatically.
3.2 Sensing experiment methods
To characterize the bulk refractive index sensitivity of the fabricated SWG microring resonator, different concentrations of glycerol in water solution (0%, 5%, 10%, 20%, v/v) were prepared and flowed onto the chip through microfluidic channels. The resonance wavelengths were recorded and the sensitivity was calculated by , the resonance shift versus refractive index change. Refractive index data was obtained based on [34].
To demonstrate the enhanced surface sensitivity, both SWG microring resonator and conventional microring resonator were fabricated on the same chip and characterized with biosensing test. Before the test, the chip was first silanized by 2% (v/v) APTES in toluene. Then the chip was further treated with 2.5% (v/v) glutaraldehyde in phosphate buffer saline (PBS) to provide aldehyde group linker that is able to immobilize protein covalently [21,25]. Next, anti-streptavidin antibody (50 μg/mL, from Abcam), bovine serum albumin (BSA, 1 mg/mL), streptavidin (100 μg/mL, from Sigma-Aldrich), and biotinylated BSA (1 mg/mL, from Thermo Fisher Scientific) were flowed in sequence into the microfluidic channel containing both microring sensors. Anti-streptavidin antibody was immobilized on the sensor surface as probe protein. BSA was used as blocking buffer to block any vacant sites. Streptavidin binds to probe protein and later capture biotinylated BSA through biochemical interactions. Before switching reagent at each of the above steps, PBS buffer was flowed to remove any unbound biomolecules. Resonance wavelengths of both the SWG and conventional microring resonator were recorded and resonance shifts were compared.
The SWG microring resonator was also used to detect low concentrations of microRNA (miRNA). The chip containing SWG microring biosensors was chemically modified with APTES and glutaraldehyde as described above. Then capture DNA (1 mM) was flowed to cover the sensor surface followed by applying blocking buffer. The conjugate miRNA (1 nM and 100 nM) was then flowed in the microfluidic channels to conjugate with the capture DNA. Anti-DNA:RNA antibody was flowed last to amplify the signal.
4. Results and discussion
Scanning electron microscope (SEM) images of the fabricated SWG microring resonator are shown in Fig. 5(a) with the coupling region enlarged to show the trapezoidal pillars in the microring and the rectangular pillars in the bus waveguide. A transmission spectrum of the SWG microring is shown in Fig. 5(b), from which the free spectral range is measured to be 12.5 nm, corresponding to group index . The estimated quality factor () is as high as 9100 due to the use of trapezoidal pillars in the SWG microring. The trapezoidal shape induces pre-distorted refractive index profile to significantly reduce mode mismatch and radiation loss in the SWG bend [31]. It is also worth noting that the absorption loss in water contributes to the total loss in the SWG microring. Moving to a shorter wavelength such as 1310 nm for operation and adjust the design slightly to compensate for the wavelength change would potentially further improve the quality factor. The device would suffer about an order of magnitude less absorption loss [30].
Fig. 5 (a) Scanning electron microscope (SEM) image of the SWG microring resonator with its coupling region enlarged. (b) transmission spectrum of the fabricated SWG microring resonator; (c) Resonance wavelength shift during the bulk refractive index sensing test; (d) Resonance shift with respect to refractive index change. Linear fit shows a bulk sensitivity of 440.5 nm/RIU.
Next, bulk refractive index sensitivity of the SWG microring biosensor was characterized. The resonance peak wavelength was monitored during the experiment and plotted in Fig. 5(c). Thus, the bulk sensitivity can be estimated by a linear fit on the resonance shift versus refractive index change plot, as shown in Fig. 5(d). The fit curve shows a bulk sensitivity of Sb = 440.5 ± 4.2 nm/RIU, which is a typical value for SWG microring resonators [24], and is about 4 times of that of a conventional microring resonator [7,30]. Considering the high quality factor, the detection limit of the sensor isRIU.
To demonstrate the enhanced surface sensitivity in SWG microring resonators, both SWG microring resonator and regular microring resonator were fabricated on the same chip and compared in a surface sensing test as described in the Methods section. The regular microring resonator has the same radius of 10 μm and the waveguide is 450 nm wide by 220 nm high (measured bulk sensitivity ~44.6 nm/RIU). The experimental results are shown in Fig. 6. Figure 6(a) presents the real time monitoring of the resonance shift in SWG microring biosensor during the sensing experiment. Regions with blue background indicate PBS buffer washing steps between the flow of different reagents. The reagent flow steps are marked in the figure with corresponding reagent names. The gradual red-shift of resonance during during steps reflects the continuous binding of biomolecules to the chemically treated surface or to its conjugated biomolecules. PBS buffer removes unbound biomolecules and create a background refractive index so that the resonances can be compared at each step. The resonance shifts of the conventional microring were also recorded at each buffer washing step. The resonance shifts for both microring biosensors are shown in Fig. 6(b). Resonance shift in SWG microring are several times larger than that in regular microring as expected, because of the much larger overlapping factors in SWG microring. It can also be seen in this figure that with more and more layers built on the surface, the resonance shift difference between the two microrings also becomes larger. To explicitly show this difference, surface sensitivity with respect to the thickness of surface layer is compared in both rings as shown in Fig. 6(c). The thickness of the surface layers is estimated by combining the simulated surface sensitivity and the experimental resonance shift in SWG microring. According to Eq. (1) and the simulation in Fig. 4(b), the surface sensitivity of the SWG ring is (λ = 1550 nm, ng = 3.0, assume n = 1.48 across all surface layer [24,27,33]) for the first 25 nm thick of surface layer. Therefore, the surface layer thickness can be estimated (). The surface sensitivity of the regular microring can then be calculated by the first part of Eq. (1) (). Figure 6(c) shows that the sensitivity of the microring resonator drops monotonically compared to that of the SWG ring as thickness of accumulated biomolecules grows continuously. It is worth noting that both devices were tested side by side in the same microfluidic channel, the surface layer thickness can be taken as the same, thus the resonance shift at each thickness can be compared. The estimated thickness in the x-axis of Fig. 6(c) also takes into account the initial thickness of silicon dioxide (~5 nm) and APTES (~5nm).
Fig. 6 (a) Real time monitoring of the resonance shift in SWG microring biosensor during the biosensing experiment; Blue region indicate buffer washing steps and other steps are marked with the corresponding reagents used. Anti-SA: anti-streptavidin antibody, SA: streptavidin, bio-BSA: biotinylated BSA. (b) Resonance shift in both SWG microring and regular microring; insets show SEM images of both microrings; GLU: glutaraldehyde (c) Surface sensitivity with respect to estimated thickness in both SWG microring and regular microring.
The fabricated SWG microring biosensor was also tested with low concentration biosamples to further demonstrate the enhanced sensitivity for biosensing applications. The experiment is performed as described in the Methods section. The test result is shown in Fig. 7. A net resonance wavelength shift of 0.11 nm was observed for 1nM miRNA with anti-DNA:RNA antibody amplification. A net resonance wavelength shift of 0.19 nm was observed for 100nM miRNA with antibody amplification. It shows that the SWG microring biosensor is promising in detecting low concentration of biomolecules in real applications.
5. Conclusion
In conclusion, we have shown that microring resonator biosensors based on SWG waveguides possess unique property of thickness-independent surface sensitivity and enhanced sensitivity compared to conventional microring resonators. Numerical simulation reveals that due to periodic pillar structure in the propagation direction, the effective sensing region includes not only top surface and sidewall of the waveguide, but also the space on the propagation path between the periodic pillars. It is the strong optical field between the periodic pillars that leads to significantly enhanced interaction with the sensing medium. Biosensing experiment on both SWG microring and conventional microring demonstated the superior surface sensing capability of the SWG waveguide. Along with the demonstration of miRNA detection at 1 nM concentration, SWG microring resonator is shown to be promising in real biosensing applications.
Funding
Department of Energy (DOE) (Contract #: DE SC-0013178); National Cancer Institute/ National Institutes of Health (NCI/NIH) (Contract #: HHSN261201500039C); National Natural Science Foundation of China (No. 61372038); Fund of State Key Laboratory of Information Photonics and Optical Communications (Beijing University of Posts and Telecommunications, IPOC2015ZC02); China and Postgraduate Innovation Fund of SICE, BUPT, 2015.
Acknowledgments
The idea is conceived by X. Xu and R. T. Chen. L. Huang acknowledges the China Scholarship Council (CSC) (NO. 201506470010) for scholarship support.
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