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Abstract
Plasmonic devices can modulate light beyond the diffraction limit and thus have unique advantages in realizing an ultracompact feature size. However, in most cases, external light coupling systems are needed, resulting in a prohibitively bulky footprint. In this paper, we propose an integrated plasmonic biosensor on a vertical cavity surface emitting laser (VCSEL) platform. The plasmonic resonant wavelength of the nanohole array was designed to match (detune) with the emission peak wavelength of the VCSEL before (after) binding the molecules, thus the refractive index that represents the concentration of the molecule could be measured by monitoring the light output intensity. It shows that high contrast with relative intensity difference of 98.8% can be achieved for molecular detection at conventional concentrations. The size of the device chip could be the same as a VCSEL chip with regular specification of hundreds of micrometers in length and width. These results suggest that the proposed integrated sensor device offers great potential in realistic applications.
© 2021 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
1. Introduction
Label-free and real-time biosensing is regarded as a promising diagnostic tool for diseases. Critical aspects for practical application of such biosensors are portable, cost-effective, easy-to-use, while also enables rapid and accurate diagnosis [1–8]. Recently, plasmonic devices have attracted significant attention in realizing ultracompact components, because surface plasmons can modulate light in nanoscale spatial region beyond the diffraction limit. Plasmonic sensing components with micro-nanoscale sizes were reported [9–15]. Most recently, bio-sensors that based on LSPR using optical fiber for efficient detection were reported [16–18]. In most cases, however, external light coupling systems are needed, resulting in a prohibitively bulky footprint. As expected, efforts have been made to satisfy the growing demand for miniaturize the dimension of the systems. Im et al. [19] described a label-free plasmonic sensor system with portable size of 10 cm×5 cm×5 cm. The system containing an external laser diode with collimating lens and square pattern diffuser for illumination. Cetin et al. [20] reported a lightweight biosensor adopting an external LED light source, and the system was 60 g in weight and 7.5 cm in height. To further reduce the size of the system, the functional components could be compact integrated on a chip [21–26].
On the other hand, research works have shown that optical imaging biosensing technology is regarded as a promising tool to increase the accuracy in medical diagnostics [27–32]. Recently, a number of sensing approaches based on optical imaging technology have been proposed and achieved excellent results [33–39]. For example, Yesilkoy et al. designed dielectric metasurfaces and used hyperspectral imaging to develop a label-free analytical biosensing platform [33]. Jahani et al. introduced an imaging platform with microarrays and microfluidics for real time breast cancer detection [37].
In this paper, we propose an integrated plasmonic biosensor with small feature size as well as high contrast on a vertical cavity surface emitting laser (VCSEL) platform. VCSEL has superior characteristics such as low emission divergence angle, cost-effective, power-efficient, and reliable, thus hold considerable promise in compact integration. Plasmonic components can be established on top of VCSEL to realize an on-chip integrated sensor that can directly monitor the optical imaging. Here, the proposed integrated sensor is based on 662 nm wavelength VCSEL directly coupled to a plasmonic Au nanohole array. The optical properties and sensing performances were systematically investigated based on Lumerical FDTD solutions. The plasmonic resonant wavelength of the nanohole array was designed match with the emission peak wavelength of the VCSEL. After binding molecules, the plasmonic resonant wavelength detuning with the VCSEL emission wavelength, thus the light excitation efficiency was changed. Therefore, the concentration of the molecule could be measured by monitoring the light output intensity. It revealed that high contrast with relative intensity difference of 98.8% can be achieved for molecular detection at conventional concentrations. The volume of the device chip could be the same as a VCSEL chip with regular specification of hundreds of micrometers in length and width. These results suggest that the proposed integrated sensor device offers great potential in realistic applications.
2. Structural design
Figure 1 he designed sensor structure based on a VCSEL chip. The VCSEL include 50 pairs of Al0.95Ga0.05As/AlAs as N type DBR1 (thickness dAl0.95Ga0.05As=50.7 nm, dAlAs=51.1 nm, refractive index nAl0.95Ga0.05As=3.254,nAlAs=3.23 [40]), 36 pairs of Al0.95Ga0.05As/Al0.5Ga0.5As as P type DBR2 (dAl0.5Ga0.5As=47.8nm, nAl0.5Ga0.5As=3.45), 45 nm-thick GaAs medium layer (nGaAs = 3.68).The sensor structure contains active region, current injection aperture, and two electrodes. The periodic Au nanohole array was arranged on top of the DBR1 just above the current injection aperture. The Au layer containing Au nanohole array, also serves as the top electrode. Here, material Au is chosen over the commonly used material Ag is due to its better chemical stability. The DBR is with $\lambda /4$ optical thickness for each layer at the center wavelength of 662 nm, which can be described by the equation ${\mathrm{\lambda }_{_{{Bragg}}}}{ = }4{nl}$ where ${\mathrm{\lambda }_{{Bragg}}}$ is the center wavelength of the stopband, $n$ is the refractive index, and $l$ is the physical thickness of each layer.
To match the emission wavelength of the VCSEL, we simulated the transmission spectra of VCSEL as the function of GaAs thickness from 20 nm to 200 nm. The results show that the resonant wavelength is periodically modulated as shown in the Fig. S2 in Supplement 1. In this paper, the thickness of the medium GaAs layer is set as 45 nm to match the emission wavelength of 662 nm. Here, the optical cavity length of the VCSEL is calculated to be 8λ. As shown in Fig. 2(a), the blue dotted line shows the reflectance spectrum of the cavity and the red solid line describes the VCSEL emission spectrum with a peak wavelength of 662 nm and linewidth of 0.18 nm.
Fig. 2. Reflectance spectrum of VCSEL (blue dotted line) and normalized emission spectra of the VCSEL (red solid line)
The working principle of the designed sensor is as follows. As known from the SPR characteristics, the optical property of the plasmonic Au nanohole array varies with the small changes in the refractive index of the surrounding environment. For example, the transmission spectrum of the Au nanohole array shifts when adsorbing the target molecules. Thus, when the VCSEL is equipped under the Au nanohole array, the laser emission intensity could be modulated by the SPR. As the plasmonic resonant wavelength is coupling well with the VCSEL emission, the light excitation efficiency is high, and after absorbing the target molecules on the Au, the wavelength detunes which results in a decreased light excitation efficiency. In other words, the emission intensity from the VCSEL is sensitive to the target molecules on the Au.
To increase the coupling efficiency as well as the sensitivity of the integrated sensor, the plasmonic resonant wavelength that supported by the Au nanohole array should be optimized to match with the emission wavelength of the VCSEL. In the next part, we present the detailed design process of the integrated sensor.
To clarify the physical origin of the plasmonic resonant modes, we simplify the VCSEL structure by replacing the optical cavity with a uniform effective medium, marked as E-VCSEL. The refractive index of the effective substrate neff is 3.63, as calculated from the volume weighted average of the VCSEL materials. A 3D FDTD simulation with periodic boundary condition for x, y direction and perfectly-matched layer (PML) boundary for z direction were employed. The mesh size of simulation for the metal region was set as 2 nm along the x, y and z direction. The permittivity data of Au was taken from Johnson and Christy [41]. The nanohole array was arranged in square in Au film as shown in Fig. 1. In this case, as the reflective index of the E-VCSEL structure was fixed, the plasmon resonant wavelength of the nanohole array can be tuned by the Au thickness h, hole diameter d and lattice constant p.
Figure 3(a) shows the normalized transmission spectra of the E-VCSEL with Au nanohole array at periodicity of 600 nm with different diameter of 200, 230, 260, 290, 320 nm and the thickness is 120 nm. The peak wavelength exhibits a clear red-shift with increasing the diameter of the nanohole, indicating that LSPR plays an important role. When the diameter of the Au nanohole array further decreased to less than 100 nm, the transmission intensity of the device is very low. Thus, balancing the relative intensity difference and the transmission intensity, we give the optimal diameter of 200 nm. Figure 3(b) shows the transmission spectra as a function of the array period, in which the hole diameter is 200 nm and the thickness is 120 nm. It is noted that as the period increased, the resonance peak wavelength shows red shift with increasing the array period. Therefore, it is believed that the resonance peak is also affected by the period of the array.
Fig. 3. Transmission spectra and electric field pattern. (a) Transmission spectra as the hole diameter changing from 200 nm to 320 nm (p=600 nm, h=120 nm);(b) transmission spectra as the function of Au nanohole array period (h=120 nm, d=200 nm); (c) transmission spectra as the function of Au nanohole array thickness (p = 600 nm, h=200 nm);(d) Electric field intensity distribution at points I, II in (c).
In Fig. 3(b), the resonance wavelength of the periodic nanohole array, which varies as a function of the array period, can be described by the Braggs coupling equation when the momentum matches between this mode and the incident photons [42]:
where, ${k_0}$ and ${k_{\bmod e}}$ are the wavevector of the incident light and the wavevector of the specific mode (SPP or WA)of the structure, respectively. The parameter $\theta$ is the incident angle, $i$ and $j$ are the gating order for lattice vector ${G_x}$ and ${G_y}$, where $|{G_x}|= |{G_y}|= 2\pi /a$ is the lattice constant in x and y direction. The wavevector of the SPP${{k}_{{WA}}}$[42] mode can be described as follows [43]:We calculated the ${\lambda _{{spp}}}$ from Eq. (3) and plotted it in dash line as shown in Fig. 3(b). It can be seen that the ${\lambda _{{spp(1)}}}$ (Au,Air) shows the same trend with the transmission peak wavelength. The mismatch could be due to the coupling between SPP and the strong LSPR mode.
To further confirm the existence of LSPR mode, we performed FDTD simulation to observe the electric field intensity distributions of the structure with hole diameter of 200 nm and Au thickness of 100 nm, as shows in the left of Fig. 3(d). The electric field distribution at P1 (653 nm) is mainly localized around the upper edge of the hole, exhibiting a LSPR characteristic. According to the Mie theory, the change of the nanostructure size could lead to a wavelength movement of LSPR mode. In this case, when increasing the hole diameters, the red shift of the resonant wavelength is due to the reduced Coulomb force (oscillation restoring force) between the electron cloud and the nucleus, which is coincidence with Fig. 3(a). Thus, it was confirmed that the origin of strong resonant mode was the interaction between SPP and LSPR of the Au holes.
As shown in Fig. 3(c), while increasing the thickness of Au nanohole layer, the resonant wavelength remains almost the same, however the transmission intensity was suppressed. The field distribution of the nanohole with 120 nm-Au-thickness and 220 nm-Au-thickness at wavelength of 656 nm were showed in Fig. 3(d) (point I and point II), respectively. We can see that, the electric field was enhanced and coupled to the air through the hole for the thin Au layer, but suppressed for the thick Au layer, thus the transmission intensity changed accordingly. In this case, relatively thin Au thickness should be applied. Ultimately, the optimized structure parameters of the Au nanohole array was set as 620 nm in period, 120 nm in thickness and 200 nm in diameter to achieve strong electric field enhancement as well as high transmittance.
3. Results and discussion
To demonstrate the potential of the device for biosensing applications, the target molecular layer is approximate to a film with a constant thickness of 10 nm and refractive index of 1.45 [44], and covered on the around Au, as shown in the inset of Fig. 4(a). Figure 4(a) and (b) show the normalized emission spectra of E-VCSEL with the optimized Au nanohole array structure after and before covering the molecules, respectively. The parameters of the Au nanohole array are: p = 620 nm, h=120 nm d = 200 nm. The plasmonic resonant wavelength P1 that supported by the structure were optimized to match well with the VCSEL emission spectrum (Fig. 4(c)) at 662 nm. As we aforementioned that the change of refractive index surrounding the Au nanohole array will lead to a shift of the plasmonic resonance peak wavelength. In this case, for the E-VCSEL, the plasmonic resonance peak wavelength shifts from P1’ at 662 nm in Fig. 4(b) to P1 at 683 nm in Fig. 4(a).
Fig. 4. Normalized transmission spectrum of E-VCSEL with Au nanohole array (a) after and (b) before covering the molecules (10 nm-thick layer with refractive index n=1.45); (c) Normalized emission spectrum of VCSEL without Au nanohole array. (d) Emission spectra of the VCSEL with Au nanohole array before (the black dashed line) and after (the red solid line) covering molecules. Far filed intensity distribution of VCSEL with Au nanohole array (e) before and (f) after covering molecules.
When VCSEL is adopted as the light emission source, the resonant cavity formed by two DBRs provide strong optical feedback and mode selection to ensure a strong and narrow linewidth emission. When the plasmonic resonant wavelength is coupling well with the VCSEL emission, the light excitation is high, as shown as the dash line in Fig. 4(d). At the same time, we can also see the high-intensity far-field distribution shown in Fig. 4(e). After covering the high index layer on the Au surroundings, the plasmonic resonant wavelength moves to longer wavelength, thus the light excitation efficiency of VCSEL was decreased due to the wavelength detuning, as plotted the solid line in Fig. 4(d). The far field of its weak intensity is distributed in Fig. 4(f). In other words, the emission intensity from the VCSEL is determined by the detuning between the peak wavelength of the VCSEL and the plasmonic resonance wavelength of the nanohole array. Thus, the VCSEL emission intensity is sensitive to the refractive index of the surroundings caused by the molecules. From Fig. 4(d), we could clearly see that the emission intensity of VCSEL is decreased dramatically after covering the molecular layer. To quantitively analysis the intensity change, a relative intensity difference $\eta $ is defined as [45]:
In order to explore the sensing characteristics of the device, the emission spectra of the integrated sensor were performed with changing the dielectric layer index, as shows in Fig. 5(a). The light emission intensity changes with the molecular concertation therefore different refractive index of the surrounding medium. When the intensity and corresponding refractive index are calibrated, the refractive index as well as the molecular concertation can be learned through the differentiation of intensity. There is a slight difference of the optimal Au nanohole parameters between the E-VCSEL and the VCSEL with DBRs. Here, the optimized period of Au nanohole is 611 nm. The normalized emission spectra of the integrated sensor with dielectric layer index n changing from 1 to 1.45 is shown in Fig. 5(b). The relative intensity difference is calculated of 98.8% in Fig. 5(b). The remarkable change in emission intensity is facilitate to the detection of protein molecular, which indicates that the proposed integrated sensor offers great potential in realistic applications. Besides, we could also design the resonant wavelength of the plasmonic structure detune with the VCSEL emission before covering the molecules, and then match after covering. The detailed analyses are presented in Supplement 1. Finally, the performance comparison of the reported plasmonic biosensors in recent years is listed in Supplement 1, Table S1. It is seen that some of sensors demonstrate high relative intensity difference of 70∼80%, however, the external light coupling systems are needed, resulting in a prohibitively bulky footprint. In our work, the proposed sensor is an integrated plasmonic bio-sensor with small feature size as well as high contrast in relative intensity difference of 98.8% on a VCSEL platform.
Fig. 5. (a) Emission spectrum at different refractive index (protein concentration) for the integrated sensor; (b) The relative intensity difference as a function of refractive index.
4. Conclusion
In conclusion, we proposed an integrated plasmonic sensor with high relative intensity difference on a VCSEL platform. The working principle and structural parameters were systematically investigated. It was found that the plasmonic resonance peak was generated by the interaction between the SPP on the metal surface and the LSPR on nanoholes. By means of rational design, the plasmonic resonant wavelength could be coupled well (or detuned) with the VCSEL emission before (or after) covering the molecular layer, thus the emission intensity varies. The highest relative intensity difference of the integrated sensor reached 98.8% for molecular detection at conventional concentrations, shows impressive sensing performance. Moreover, the size of the device chip could be the same as a VCSEL chip with regular specification of about several hundred microns in length and width, which exhibit great potential in portability and power saving. Our studies could offer guidance for miniaturization and integration of highly sensitive biomedical sensors.
Funding
National Key Research and Development Program of China (2018YFB1801001); National Natural Science Foundation of China (61975037, 62175039, U2001601, 12004445); Natural Science Foundation of Guangdong Province (2019A1515010905, 2019A1515011471); Program for Guangdong Introducing Innovative and Entrepreneurial Teams.
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.
Supplemental document
See Supplement 1 for supporting content.
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