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Abstract
A high-sensitivity surface plasmon resonance (SPR) sensor based on the coupling of Au grating and Au film is investigated through simulations and experiments. The SPR sensor is designed by using a hybrid method composed of genetic algorithm (GA) and rigorous coupled wave analysis (RCWA). The numerical results indicate the sensor has an angular sensitivity of 397.3°/RIU (refractive index unit), which is approximately 2.81 times higher than the conventional Au-based sensor and it is verified by experiments. Theoretical analysis, by finite-difference time-domain (FDTD) method, demonstrates the co-coupling between surface plasmon polaritons (SPPs) propagating on the surface of Au film and localized surface plasmons (LSPs) in the Au grating nanostructure, improving the sensitivity of the SPR sensor. According to the optimized structural parameters, the proposed sensor is fabricated using e-beam lithography and magnetron sputtering. In addition, the proposed sensor is very sensitive to the detection of small molecules. The limit of detection (LOD) for okadaic acid (OA) is 0.72 ng/mL based on an indirect competitive inhibition method, which is approximately 38 times lower than the conventional Au sensor. Such a high-sensitivity SPR biosensor has potential in the applications of immunoassays and clinical diagnosis.
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1. Introduction
Surface plasmon resonance (SPR) technique has been widely used due to the particularly advantages of label-free, real-time, high sensitivity, and quantitative detection [1–7]. SPR biosensors exploit surface plasma waves generated at the interface between a thin metal (Ag or Au) surface and the sensing medium to detect the biomolecules. The surface plasma waves are very sensitive to the changes in the refractive index (RI) of the external media [8–10]. Various biomolecular interactions, such as nucleic acid-protein, protein-protein and DNA hybridization occurring on the sensor surface, will cause a change in the surrounding RI, ultimately resulting in a significant shift of SPR signal. In general, the angular sensitivity of conventional SPR sensors consisting of a single metal film via Kretschmann coupling is only 50-150°/RIU [11]. SPR sensors with low sensitivity will hinder the detection of small molecular masses or low molecular concentrations due to the small shift of SPR signal to the RI changes. To improve the detection sensitivity, various sensitivity-enhancement approaches have been employed in SPR biosensors. Many researchers use surface chemistry methods to efficiently immobilize bioreceptors on the sensing layer or utilize nanoparticles to amplify SPR signal [12,13]. In addition, enhancing the electromagnetic field in the sensing area is also an effective approach to improve the sensors’ sensitivity [14,15].
In general, bioreceptors, such as antibodies, nucleic acids, or other recognition molecules, have poor immobilization on the sensor surface, limiting the sensitivity of the sensors [16]. Many chemical methods have been developed to immobilize bioreceptors on the surface. Among them, the self-assembly monolayer (SAM) has been widely proven to be particularly effective because of its good stability under extreme temperature and pH conditions, as well as reuse and applicability in flow systems [17–20]. For example, Kawaguchi et al. [19] employs the PEG based self-assembled layer for detecting TNT with a detection limit of 0.008 ng/mL. Wang et al. [20] reports a detection limit of 14 nM for small molecule drugs by using a nitrilotriacetic acid (NTA) SAM covered on the SPR chip as a high density of oriented protein bioreceptors.
To further improve the sensing performance of the SPR sensors, nanoparticles, such as Au nanoparticles (Au NPs) and magnetic nanoparticles (MNPs), are used as SPR signal amplification tags. Nanoparticles increase the RI of the analyte in the analyte-NP conjugates, resulting in an enhanced shift in the SPR signal [21].
In addition to as SPR signal amplification tags, Au NPs may improve the sensing performance by the electric field coupling effect between the surface plasmon wave propagating on the metal-dielectric interface and localized surface plasmons (LSPS) oscillating around the Au NPs [ 22]. At the same time, by adjusting the length of the chemical bond between the metal film and the nanoparticles, the electromagnetic field can be further enhanced [23]. In recent years, various AuNPs with different shapes and sizes have been proposed to improve the detection ability of SPR sensors. However, these approaches have several disadvantages including the lack of precise control over shapes and sizes. Therefore, periodic structures with uniform and controllable shapes, have a distinct optical property, and are expected to provide stronger electromagnetic field enhancement in the sensing area.
With the development of nanofabrication technology, periodic nanostructures are widely used in the design of SPR sensor. For example, Cai et al. [24] numerically investigates a highly sensitive SPR sensor based on trapezia metallic gratings of sensitivity 237°/RIU. Bijalwan et al. [25] proposes a sensor composed of Au-Al2O3 grating and aluminum layer. It has a refractive index sensitivity of more than 270°/RIU. In addition, an Au-Ag bimetallic grating configuration has also been reported by Bijalwan et al. [26] to improve the sensitivity of conventional grating-assisted SPR sensors, and its maximum sensitivity can reach 346°/RIU. However, these methods have not been implemented on the prism-based SPR sensor. Furthermore, the prism-based systems are widely used in commercial SPR devices due to its ease of use and high sensitivity [27].
To design high-performance prism-based SPR sensors with periodic structures, several parameters, such as the grating period, its width, and the thickness of metallic layer, should be adjusted simultaneously. Small changes in these parameters may change the characteristics of the sensors, resulting in lower efficiency of conventional manual optimization methods. Genetic algorithm (GA) is a simple, powerful and efficient global optimization algorithm that can achieve multi-parameter and multi-objective optimization simultaneously [28,29]. At present, several researchers used it for determining the parameters of SPR sensors.
Here, a high-sensitivity SPR biosensor based on a periodic Au grating adhered to Au film structure is theoretically and experimentally investigated. The physical mechanisms responsible for improving sensitivity of the devices, including the plasmon coupling of SPR and local surface plasmon resonances (LSPR) modes, have been analyzed clearly by the finite difference time domain (FDTD) method. According to the optimized structural parameters, the designed sensor is fabricated by using e-beam lithography and magnetron sputtering. Based on the experimental results, the RI sensitivity of this SPR sensor is 2.95 times that of the conventional Au SPR sensor, which is basically consistent with theoretical calculation results. In addition, the proposed sensor is used to detect the OA concentration by indirect competitive inhibition method. The results indicate that immobilizing Au grating on the Au layer can greatly reduce the limit of detection (LOD) of the SPR biosensor.
2. Theoretical simulation
2.1 Structure of the proposed SPR sensor
The designed SPR sensor is based on a prism, adhesion layer, Au layer and Au grating, whose configuration is shown in Fig. 1(a). In the sensor, we use BK7 glass as coupling prism and select Au as metal material to excite the SPR or LSPR. When the resonance frequencies of surface plasmon polaritons (SPPS) and LSPs are equal to each other, the coupling effect between the two plasmas will produce stronger local electric field. In theory, SPR-LSPR coupling effect can improve the sensitivity of the sensor. Figure 1(b) is corresponding cross-section view of the proposed sensor. The ${h_1}$ and ${h_2}$ are the thicknesses of Au gratings and Au layers, respectively. The $\Lambda $ and w denote the period and width of Au gratings, respectively. In the simulation, the Cr layer is as an adhesion layer to match the real experimental scenario with a thickness of 5 nm. The wavelength of the incident light source is 632.8 nm. At this wavelength, the RI of BK7 glass is 1.516 [30]. The RI of Cr and Au are obtained from the Drude-Lorentz model [31,32].
2.2 Optimization of the proposed SPR sensor
For a SPR biosensor, in order to achieve better sensing performance, a single parameter, combined sensitivity factor (CSF) [33,34], is used in GA.
Table 1. Ranges of the design parameters
Table 2. Parameters of GA
Table 3. Designed parameters of proposed sensor optimized through GA.
2.3 Simulation results of the proposed SPR sensor
To evaluate the sensitivity performance, we compare the conventional Au-based sensor and the proposed sensor. The resonance curve of the two structures as the function of incident angles is shown in Fig. 3(a) and Fig. 4(a), respectively. Each curve has a sharp dip at a specific incident angle, known as SPR resonance angle. Figure 3(a) is the reflectivity curves for the conventional Au-based sensor, and the thickness of Au film is 50 nm. From the Fig. 3(a), the resonance angle moves towards to the higher angle side (red shift), when ${n_s}$ varies from 1.330 to 1.345 with a step of 0.0025. The reason for such red shift is the propagating wave vector of the surface plasmon wave increases due to increasing the RI of sensing medium [39]. Figure 3(b) shows that the resonance angle for the Au-based sensor has a linear relationship with RI, and the linearity is 0.9997. Therefore, the sensitivity of this sensors is 141.1°/RIU. Similarly, the sensing performance of the proposed sensor is investigated in Fig. 4(a-b). The result indicate that the RI sensitivity is greatly improved, up to 397.3°/RIU, which is 2.81 times that of the conventional Au SPR sensor. Furthermore, the RI sensitivity obtained here is significantly improved compared to other reported work shown in Table 4. Although the proposed SPR sensor has a high sensitivity, we observe that the SPR curves in Fig. 4(a) exhibit a relatively shallower and broader resonance band, compared to those in Fig. 3(a). This due to the fact that localized plasmon modes generated by Au grating increase the absorptive damping in plasmonic wave. In practical application, this wide and shallow resonance band may affect the detection accuracy. To improve detection accuracy, some researchers use the prism of higher RI material in conventional Au-based sensor and obtain very deep and sharp resonance peaks [40]. Therefore, in future research, the use of higher RI substrate material in the proposed sensor is expected to be a possible direction for future improvement.
Fig. 3. (a) SPR curve for conventional SPR sensor with the variation of RI of sensing medium; Insert: Schematic illustrations of the conventional SPR sensor (b) Linear regression analysis between the resonance angle and RI of sensing medium.
Fig. 4. (a) SPR curve for proposed SPR sensor with the variation of RI of sensing medium; Insert: Schematic illustrations of the proposed SPR sensor (b) Linear regression analysis between the resonance angle and RI of sensing medium.
Table 4. Comparison of the performances of the proposed sensor and other reported sensors. (BP: black phosphorus; WSe2: tungsten diselenide)
To study the electric field enhancement for both the conventional Au film structure and the Au grating coupled structure, the electric field profiles of the structures are compared at resonance condition. The field distribution is investigated based on the FDTD method. In the FDTD algorithm, a plane wave is incident from the BK7 glass substrate side at a resonance angle with the z-axis. The electric field of the incident wave is set to be linearly polarized in the x direction (TM polarization). Bloch boundary condition is used in the x direction due to the light source oblique incident. Symmetric boundary condition is used in y direction to reduce the simulation time. A perfect matched layer (PML) boundary condition of 16 PML is used in the z direction, to eliminate the boundary scattering. The mesh size of 0.5 nm is set in all directions to improve the accuracy of the calculated results. Figure 5 (a-b) are the electric field distribution of the two sensors, with the field amplitude normalized to 10. Red regions represent the electric field-enhanced regions. Compared with the Au film structure, the electric field of the proposed structure in Fig. 5 (b) is significantly enhanced by coating Au grating on Au film. The LSPs present in Au grating are coupled to SPPs that propagate on the Au surface, when the resonance frequencies of SPPS and LSPs are equal to each other. The plasmon coupling effect results in an enhanced electric field distribution between the Au grating and the Au film, especially the Au grating surface [41]. In addition, the LSPs generated by grating can interacts with adjacent metal grating to produce the LSPs-LSPs coupling, creating an enhanced local electric field at the gap gratings [42]. Those electric field coupling regions can not only strengthen the interaction with external sensing media, but also have a more sensitive response to the change of external RI. This means that the proposed sensor is sensitive to the small change in sensing medium.
Fig. 5. Electric field distributions for (a) conventional sensor; (b) the proposed sensor. The field amplitude is normalized to 10.
3. Experiment
3.1 Sample fabrication
The proposed structure is fabricated by using e-beam lithography and magnetron sputtering. First, a 5 nm Cr layer and 41 nm thick Au layer are sequentially deposited on the polished BK7 glass substrate via magnetron sputtering, in which the Cr layer served as an adhesion layer. Then, the Au gratings are patterned on top of the Au layer by e-beam lithography. Figure 6 is the SEM image of fabricated grating structure. It is evident from the images that the Au gratings are evenly distributed over the sample. It is worth noting that the total area of the gratings successfully patterned on the Au layer is around 1 mm * 1 mm. The average period and width of the fabricated grating are 416.8 nm and 208.4 nm, respectively.
Fig. 6. Top view SEM images of the proposed sensor in different magnification. (a) 5000x magnification; (b) 30000x magnification.
3.2 RI sensing experiments
Here, the RI sensing experiments are carried out to compare and analyze the sensitivity of the two sensors. We use an Abbe refractive index meter to calibrate the RI of the NaCl solutions at room temperature (26°C). The relationship between the RI and concentration of NaCl can be represented by a linear regression line: n = 0.000158648*c + 1.3315, where c represents the concentration in g/L. Therefore, the refractive indices of concentrations of NaCl solutions (0.1 g/L, 0.25 g/L, 0.5 g/L, 1 g/L, 2.5 g/L, 5 g/L, 10 g/L) are 1.33152, 1.33154, 1.33158, 1.33165, 1.33189, 1.33229, 1.33308 respectively. In the experiment, deionized water is first injected into the sensor and maintains for sufficient time to ensure the stability of the baseline. Then, NaCl solutions, and deionized water are repeatedly injected in a single cycle to test different concentrations of NaCl solutions. Every concentration is measured three times, and the experimental data is recorded in real time with software. Figure 7(a) show the response signal for the conventional Au-based biosensor in different RI solution. Figure 7(b) shows the corresponding linear fit curve, and the RI sensitivity is about 13951.2 pixel/RIU.
Fig. 7. (a) The response spectrum of NaCl solution with different concentrations for the conventional Au-based biosensor. (b) The corresponding fitted line of the RI to the response signal shift.
Similarly, in order to obtain the RI sensitivity of the proposed sensor, NaCl solutions with different concentrations are injected into the sensing region. The response spectrum is shown in Fig. 8(a). From the linear fitting curve of Fig. 8(b), the RI sensitivity is 41202.4 pixel/RIU with linearity of 0.97832. The experimental results show that the enhancement of the SPR response is due to the introduction of Au grating. In addition, the RI sensitivity of the proposed sensor is 2.95 times that of conventional sensor. A good agreement is observed between the experimental results and the theoretical calculation results.
Fig. 8. (a) The response spectrum of NaCl solution with different concentrations for the proposed biosensor. (b) The corresponding fitted line of the RI to the response signal shift.
3.3 Okadaic acid detection experiment
3.3.1 Reagent and materials
Okadaic acid (OA) with >95% purity (HPLC-grade), the okadaic acid-bovine serum albumin bioconjugate (OA-BSA), and anti-OA monoclonal antibody (OA-mAb) are purchased from Anti Biological Technology Co., Ltd. (Shenzhen, China). 11-Mercaptoundecanoic acid (11-MUA), N-hydroxysuccinimide (NHS), 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride (EDC), ethanolamine and phosphate-buffered saline (PBS) are purchased from Shanghai Aladdin Biochemical Technology Co., Ltd. Ethyl alcohol, NaCl, H2SO4, HCl, NaOH and H2O2 are obtained from Sinopharm Chemical Reagent Co., Ltd. (Shanghai, China). A 0.2 M NaOH solution is used as the regenerant. All reagents and solvents are analytical grade and are used without further purification. Deionized water (18.2 MΩ·cm) is used throughout the work.
3.3.2 Fabrication of SPR immunosensor
Initially, the chip is cleaned using a piranha solution (H2SO4/H2O2 = 7:3, v/v) to remove all organic substances. Then, the chip is rinsed thoroughly with plenty of ethanol and water, and finally dried with nitrogen. The clean chip is left immersed overnight in 1 mM 11-MUA/ethanol to form a stable SAM [43]. After cleaning with ethanol and drying with nitrogen, the chip is activated by NHS/EDC mixture solution (M:M = 4:1), incubated for 30 min, and followed by washing in plenty of water. After that, the OA-BSA conjugate is dropped onto the chip surface and incubated for an hour. Finally, 1 mol/L ethanolamine solution (pH 8.5) is used to block the non-specific interaction on the biosensor surface.
3.3.3 Determination of OA
The detection of OA is based on an indirect competition inhibition method, which is widely used for the detection of small molecule [44,45]. Figure 9(a) show the process of the OA-mAb and the OA-BSA specifically bind on the sensing surface. After reaching a stable baseline with running PBS buffer, the OA-mAb solution (10 µg/mL) is injected into the reactor. A steady increase in SPR response signal indicates specific binding of the OA-mAb and OA-BSA. At the end of the immunoreaction, the SPR response signal do not decrease during the flow of running PBS buffer, suggesting a stable binding of OA-mAb with the OA-BSA immobilized on sensor chip. After each detection cycle, the sensor surface is regenerated by injection NaOH (0.2 M) solution for cost-effective and multiple analyses.
Fig. 9. (a) Response curve showing specific binding between OA-mAb (10 µg/mL) and the OA-BSA followed by regeneration step. (b) The shift of response signal with increasing concentration of OA-mAb. (c) Experimental diagram of OA solution at different concentrations for the conventional Au-based biosensor. (d) Calibration curve for the detection of OA by competitive immunoassay.
To investigate the appropriate concentration of OA-mAb, the sensor chip is exposed to the flow of OA-mAb at various concentrations (1-60 µg/mL). The net rise of resonance pixel with increasing concentration of OA-mAb is shown in Fig. 9(b). The resonant pixel shift increases rapidly before the antibody concentration reaches 25 µg/mL, and then increases slightly above this concentration. Thus, the use of approximately 25 µg/mL OA-mAb solution is considered to be the best choice for OA sensing experiments. For the indirect competitive immunoassay, each OA solutions are mixed with OA-mAb to prepare mixtures containing the antibody at a final concentration of 25 µg/mL and OA with different concentrations (16-1024 ng/mL). The mixtures are incubated for 30 min, and then flow through the immobilized sensor chip. The obtained SPR response signal is shown in Fig. 9(c).
From the Fig. 9(c), it can be inferred that the resonance pixel shift for the immune reaction is reduced in the presence of OA. The extent of the decrease in the resonance pixel shift is directly related to the concentration of OA. In addition, SPR signal response from triplicate injection at the same concentration shows an almost similar curve, indicating good reproducibility of the immune reaction. The sigmoidal calibration curve clearly shows that the binding interaction between OA-BSA and OA-mAb is inhibited in the presence of OA, as shown in Fig. 9(d). As can be seen from the curve, the linear detection range of the biosensor is from 64 ng/mL to 512 ng/mL. According to the LOD formula [46], the LOD of the conventional Au-based biosensor can be calculated to be 26.7 ng/mL.
The low-molecular weight analyte OA is also detected by the proposed SPR sensor by an indirect competition inhibition method, as shown in Fig. 10(a-c). According to the curve shown in Fig. 10(a), the 8 µg/mL OA-mAb concentrations is chosen to mix with standard solutions of OA. It is worth noting that the mixtures contain the antibody at a final concentration of 8 µg/mL and different concentrations of OA (0.25-2048ng/mL). Those mixtures are injected into the sensor chip, and the response signal is recorded in Fig. 10(b). When the mixture is exposed to the OA-BSA immobilized on the sensor chip, OA in mixed solution will compete with OA-BSA for binding to OA-mAb, resulting in a decrease in the response signal. The extent of reduction is directly proportional to the OA concentration. Figure 10(c) shows the dependence of the response signal shift against the concentration of OA. From the sigmoidal calibration curve, it can be inferred that the linear detection range of the proposed biosensor for OA is from 1 ng/mL to 1024 ng/mL, and the LOD is 0.72 ng/mL (S/N = 3). The high sensitivity observed with the proposed biosensor is attributed to the electric field intensity enhancement.
Fig. 10. (a) The shift of response signal with increasing concentration of OA-mAb. (b) Experimental diagram of OA solution at different concentrations for the proposed biosensor. (c) Calibration curve for the detection of OA by competitive immunoassay.
4. Conclusion
In summary, a high-sensitivity SPR biosensor based on SPR-LSPR coupling is proposed, and its performance is simulated and optimized by the RCWA and GA methods. The structure of the sensor is made of a periodic Au grating and Au film, and is fabricated using e-beam lithography and magnetron sputtering. The device is used to measure the RI sensitivity and detect OA concentration. The result of RI experiments show that the RI sensitivity of the proposed sensor is about 2.95 times higher than the conventional Au sensor. In addition, the LOD of OA for the proposed sensor is 0.72 ng/mL, which is about 38 times lower than the conventional Au sensor. This device may have great application prospects in chemical detection and clinical diagnostics.
Funding
National Natural Science Foundation of China (61775191); Science Fund for Distinguished Young Scholars of Fujian Province (2020J06025); Youth Talent Support Program of Jimei University (ZR2019002); Innovation Fund for Young Scientists of Xiamen (3502Z20206021); Xiamen Marine and Fishery Development Special Fund (20CZB014HJ03); Natural Science Foundation of Fujian Province (2020J01712); Youth Talent Support Program of Fujian Province (Eyas Plan of Fujian Province 2021).
Disclosures
The authors declare that there are no conflicts of interest related to this article.
Data availability
Data underlying the results presented in this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.
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