1. Introduction

Accurate and efficient detection of disease-related biomarkers can save more effective time for disease treatment [1–4]. The development of new biosensors has brought new hope for the rapid and sensitive detection of biomarkers. In particular, optical sensors have outstanding applications in disease detection because of their advantages of high sensitivity, label-free, and fast [5–8]. They are expected to replace or augment the electrochemical luminescence detection used in traditional laboratories [9]. Among these biosensors, localized surface plasmon resonance (LSPR)-based biosensing shows an immense potential to satisfy the needs of liquid biopsy detection for low-cost, fast, and portable automated systems, highly sensitive and real-time detection, multiplexing and miniaturization [10–12]. However, despite the high performance of LSPR-based biosensors, conventional components of the LSPR system often rely on a benchtop or laboratory scale instruments, thereby limiting their applications in remote detection and long-time monitoring of diseases. At present, the miniaturization of devices is one of the hot topics in the research of nanoplasma sensing. Belushkin et al. built a nanoplasmonic imager through optical components to realize a portable sensing system. The realization of on-chip integration will further miniaturize the device to make it more convenient [13].

Vertical-cavity surface-emitting laser (VCSEL) is a miniaturized light source that has developed rapidly in recent years [14,15]. The VCSEL-based sensors have been applied to detect gas [16,17], volatile particles [18], pressure [19], proteins [20,21], etc. VCSELs can integrate with silicon photodiodes to achieve fluorescence sensing [22]. LSPR biosensors that integrate with VCSELs represent an alternative sensing platform to achieve miniaturization and portability [23]. However, this method is only functional integration. VCSELs replace traditional light sources, and LSPR chips are independent of VCSELs. Liu et al. proposed an integrated biosensor based on a nanohole array VCSEL platform that monitors light output [24].

The recent worldwide outbreak of severe acute respiratory syndrome related coronavirus 2 (SARS-CoV-2) has [25] led to unprecedented pressure on national healthcare systems [26,27]. The World Health Organization (WHO) advised the international community to perform extensive diagnostic tests to detect infection cases early and reduce the spreading of the virus [28]. The test biomarkers include nucleic acid and antibodies. Nucleic acid is suitable for early screening, and the antibody is more stable than nucleic acid. The serological analysis is essential for identifying asymptomatic infected persons to further control the spread of the virus and evaluate the effectiveness of the patient's immune system against infection. Therefore, new technologies are needed which could be quickly implemented and integrated into sensors to detect a range of viral antigens [29]. For traditional methods, such as enzyme-linked immunosorbent assay, operators need to be trained, and it takes hours or even days to analyze. On the other hand, rapid diagnosis methods are urgently needed. New diagnostic sensors have been developed on demand, including LSPR [30,31], field effect transistor (FET) [32], and electrochemical methods [33]. The purpose of these efforts is to find a reliable and rapid detection method. Although these methods are promising, each has its own limitations. For example, LSPR-based techniques usually require signal amplification to enhance sensitivity and improve the detection limit. The FET-based techniques require specialized analytical equipment. The electrochemical-based techniques require a complex electrode modification process. Thus, the development of label-free, reliable detecting methods is necessary.

In order to address the above challenges, a miniaturized VCSELs sensor was developed to detect the RBD protein of SARS-CoV-2, which was integrated with gold nanograting arrays based on the LSPR technique. The excitation light source (VCSELs), sensing sensitive layer (gold nanograting), and microfluidic channel are integrated to develop a small-size, high-sensitivity biosensor. VCSEL is used to replace the traditional light source of the LSPR system, which realizes the on-chip integration of the light source, and excites the local plasma resonance effect of the gold nanograting array. Gold nanograting arrays were prepared on VCSELs. By correlating the wavelength shift of the interaction between the LSPR peak of gold nanograting and the RBD protein of SARS-CoV-2, the power change of VCSELs exciter was caused. Then, the concentration of RBD protein was quantified. The RBD aptamer was modified on the gold nanograting array to realize the sensing detection of the RBD protein. It provides another new platform for the highly sensitive and rapid detection of biomarkers.

2. Methods

2.1 VCSEL fabrication

The preparation process is as follows: firstly, a VCSEL chip with a wavelength of 850 nm is prepared. A layer of 500 nm SiO₂ is passivated on the prepared VCSEL surface by plasma-enhanced chemical vapor deposition (PECVD) to insulate VCSEL from a gold nanograting array. The gold nanograting array as a sensing layer was prepared on the surface of VCSEL by electron beam evaporation for the gold layer and focused ion beam (FIB) etching for the nanograting array.

2.2 PDMS channel fabrication

After preparing the gold nanograting array, a polydimethylsiloxane (PDMS) microfluidic channel with an injection chamber needs to be bonded to it. The PDMS prepolymer and curing agent were evenly mixed at a ratio of 10:1, and the mixture was left to be bubble-free. Then the PDMS prepolymer was poured into the master mold with a pattern and dried at 80°C for half an hour. To ensure the quality of light transmission, we should avoid residual bubbles. After bonding the microfluidic channel, the complete VCSEL biosensor is ready which defines the inlet and outlet of the injection. The circular cavity of the microfluidic channel should cover the laser mesa area on the chip.

2.3 Functionalization of gold nanograting

The effective functioning of the chip is very important for detection performance. Because the protein cannot directly bind to the gold nanograting, the experiment was carried out by modifying a layer of aptamer (SARS-CoV-2 RBD aptamer) on the surface of the gold nanograting that specifically binds to RBD protein. One end of the aptamer has a sulfhydryl group, which can be coupled with the gold nanograting through the Au-S bond. The functional modification process is as follows: 5 µM SARS-CoV-2 RBD aptamer is activated in tris(2-carboxyethyl) phosphine (TCEP) solution for 1 hour. The activated aptamer solution was passed into the microfluidic channel with a microinjection pump and was modified overnight at 4 °C for 10 hours to fully combine the thiol group on the aptamer. After rinsing with phosphate buffer solution (PBS) buffer, the biosensor was stored at 4°C for further use. The biosensor can be functionalized in one step.

2.4 Optical characterization of gold nanograting

Refractive index sensitivity is one of the important parameters for biosensors. Sucrose solution with mass percentages of 0, 10, 20, 30, 40, and 50% was used as a refractive index solution to test the refractive index sensitivity of gold nanograting. The refractive index sensitivity was calculated based on the spectral responses of nanograting chips to different concentrations of refractive index solutions.

2.5 Measurement of SARS-CoV-2 RBD protein

The SARS-CoV-2 RBD protein is used to detect the potential for biological applications of the sensor chips. Firstly, SARS-CoV-2 RBD protein was diluted into 100 µg/mL and divided into 10 µL tubes, that is, 500 µL of glycerol and 500 µL of sterile deionized water were added to 100 µg of protein. Dilute to the desired concentration according to the concentration during use. The P-I-V test system detects the change of the output light power after the VCSEL chip captures the antigen. The read power at saturation current was recorded as the sensed signal.

3. Results and discussion

3.1 Schematic design of the integrated biosensor

The VSCEL sensor integrated nanograting array is shown in Fig. 1(a)-(c). This sensor uses VCSEL with a laser band of 850 nm as the sensing light source. The design of the whole sensor chip and detection system is shown in Fig. 1(a)-(g). The sensor’s main structures include bottom contact, bottom distributed Bragg reflector (DBRs), oxide aperture, top DBRs, top contact, and nanograting array, taking a single laser as an example. VCSELs are used as the light source. A gold nanograting array is fabricated on the surface of VCSELs directly which is fabricated by FIB technology, and the sensing area is encapsulated by a PDMS microfluidic channel.

 

Fig. 1. Schematic illustration of integrated microfluidic plasmonic biosensor from VCSEL for SARS-CoV-2 RBD protein detection. (a) schematic illustration of integrated microfluidic plasmonic biosensor from VCSEL. (b) the VCSEL sensor is integrated with a microfluidic chip. VCSEL includes top contact, grating, top DBRs, bottom contact, and bottom DBRs. (c) enlarged schematic diagram of VCSEL structure. The grating array was fabricated on the VCSEL surface. (d) photos of integrated microfluidic plasmonic biosensor from VCSEL. (e) SEM image of gold grating array integrated with VCSEL. The scale is 10 µm. (f) functionalization process of integrated microfluidic plasmonic biosensor from VCSEL. (g) output power variation (sensing signal) of the sensor.

Download Full Size | PDF

PDMS is an ideal material for microchips with good light transmission and thermal stability [34]. The inlet and outlet were fabricated in the microfluidic channel (Fig. 1(b)). The inlet is used to connect the microinjection pump to introduce the regent into the sensor chip. The outlet is connected to a tube to collect waste liquid. The photograph of the sensor is shown in Fig. 1(d). The size of the sensor chip is only 1 cm × 1 cm. The scanning electron microscope (SEM) image of the gold nanograting array is shown in Fig. 1(e). The structure area of the whole gold nanograting array is 24 µm × 24 µm. The width of nano grating is 520 nm and the period is 570 nm. The surface plasmon resonance wavelength of the gold nanograting structure is the same as that of the VCSEL, i.e., 850 nm. Electrodes of the VCSELs are exposed outside so that the electrodes can be easily connected to add current to the sensor chip (Fig. S1, Supplement 1). The detector is used to collect optical power signals. The display is used to detect the change of power in real-time (Fig. S2, Supplement 1). After the preparation of the biosensor, a specific aptamer of RBD which coupled with thiol was introduced into the microchannel to functionalize the sensor. Then, different concentrations of RBD protein were successfully detected (Fig. 1(f)). The coupling of substances at each step will cause the change of sensor power which can establish the relationship between the concentration of the detected substance and the power (Fig. 1 (g)).

3.2 Characterization of the integrated biosensor

We use the finite difference time domain (FDTD) method to simulate the LSPR effect of metal nanostructures. First, the physical structure is modeled. Then, the simulation area and boundary conditions, the light source, and the monitor are set. After completion, the operation calculation can be carried out. The parameters of the gold grating are set as slit width (W) = 50 nm, and thickness of gold (H) = 50 nm. The periods (P) of the gold gratings were set at 550, 570, and 590 nm, respectively. In the experiment, the simulation time is set to 1000 fs to ensure that the simulation time is long enough and there is enough attenuation time. The detection wavelength range is set to 700-1100 nm, and the frequency number is set to 800. The periods of the gold gratings were set at 550, 570, and 590 nm, respectively. Through the simulation spectrum (Fig. S3, Supplement 1), the resonance valley wavelength of the 570 nm nanograting is the closest to the wavelength of the VCSEL light source. The optimal period of gold nanograting is 570 nm. The structure diagram and simulation results of the gold grating are shown in Fig. 2(a). With the increase of the background refractive index of the solution, the transmission spectrum line is red shift. The valley with the wavelength closest to 850 nm of the VCSEL light source was selected for linear fitting. Through fitting calculation, the transmittance sensitivity is 597.91 nm/RIU, as shown in Fig. 2(b). At 850 nm, the surface electric field of the grating structure is highly localized, showing the surface plasma polariton characteristics (Fig. 2(c)). The LSPR response spectrum of the grating corresponds to the output spectrum of the VCSELs prepared, i.e., 850 nm (Fig. 2(d)). The spectral structure can be integrated on VCSELs to achieve LSPR sensing.

 

Fig. 2. Design of the integrated sensing chip. (a) simulation results of gold nanograting transmittance under different refractive index environments: 1.332, 1.347, 1.363, 1.380, 1.399. (b) simulated refractive index sensitivity. The dotted line is the result of the linear fitting of all points. (c) electric field distribution of cross-sections (XY plane) of the gold grating structure. (d) output spectra of VCSEL.

Download Full Size | PDF

In order to test the sensing performance of the nanograting obtained by simulation, the gold nanograting array structure was prepared on the quartz plate. Different refractive solutions were tested and introduced into the microfluidic channel in sequence at a rate of 1 mL/h. After each concentration was tested, the microfluidic channel was rinsed with deionized water to prevent the residual solution from affecting the test results. As shown in Fig. 3(a), the transmission spectrum of the gold nanograting array on the quartz shows that the resonance wavelength of the gold nanograting array on the quartz shifts red as the concentration of sucrose solution increases. That is, the refractive index of the solution increases. In order to verify that the spectral shift is caused by the gold nanograting array, the transmittance spectrum of the quartz without the gold nanograting array was tested. The results are shown in Fig. S4 (Supplement 1). As the refractive index of the solution increases, the transmittance of the quartz does not change. In agreement with the simulation results, the spectral valley is closest to the wavelength of the VCSEL source at 850 nm. The relationship between the refractive index of the solution and the valley position is plotted, as shown in Fig. 3(b). The refractive index sensitivity can be obtained by linear-fitting the peak wavelength of the transmission spectrum with a wavelength between 750-850 nm and the refractive index of the corresponding refractive index solution. The calculated refractive index sensitivity is 571.07 nm/RIU.

 

Fig. 3. Sensing performance of nanograting structure and the integrated VCSEL biosensor. (a) transmission test results of gold nano gratings with different refractive solutions. (b) sensitivity calculation of gold nano gratings. (c) P-I-V characteristics of the VCSEL biosensor with different refractive solutions. (d) sensitivity calculation of the VCSEL biosensor. The dotted line is the result of the linear fitting of all points.

Download Full Size | PDF

The VCSEL biosensor was prepared by integrating a gold nanograting array. Therefore, the wavelength shift caused by the refractive index change of different solutions is shown as the change of VCSEL power on this integrated biosensor. The wavelength of the light source provided by VCSEL is consistent with the designed resonance wavelength of nanograting. The refractive index solution will change the plasmon resonant wavelength of the grating. Therefore, the sensing sensitivity of the integrated sensor was measured with different refractive solutions. The output optical power-current-voltage (P-I-V) characteristics of different refractive indexes of solutions were tested. Figure 3(c) shows that the output power of the VCSEL biosensor increases with the increase of the different refractive indexes of the solution. The output light power of 50 mA current in different refractive index solutions changes obviously. The relationship between the refractive index of the solution and the output power is plotted which is shown in Fig. 3(d). The refractive index sensitivity can be obtained by linear fitting the output optical power of 50 mA current with the refractive index of the corresponding refractive index solution. The sensitivity of the integrated VCSEL biosensor is 2.99 × 10⁶ nW/RIU.

3.3 Performance of the integrated biosensor

The good detection performance of the biosensor is inseparable from the efficient functionalization method of the hot spot. The traditional gold surface functionalization method is to modify 11-mercaptoundecanoicacid (11-MUA), activate by 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide-HCI (EDC) and sulfo-nhydroxysuccinimide (S-NHS), and then connect antibodies through carboxyl and amino binding. Many modification steps introduce uncertainties factors. In this study, sulfhydryl-modified aptamers are used to identify RBD protein which can achieve one-step functionalization of the biosensor. The process of biosensor modification and detection has been shown in Fig. 1(f). First, we used a synthesized aptamer with both the thiol group and green fluorescent group to detect the coupling between the aptamer and gold nanograting. From the bright field and fluorescent field in Fig. 4(a) and Fig. 4(b), there is no fluorescence signal on the gold substrate, but a green fluorescence signal appears after adding the aptamer. Then, we used the atomic force microscope (AFM) test to characterize the successful capture of the RBD protein. As shown in Fig. 4(c) and Fig. 4(d), when the protein is captured in the gold nanograting, the surface roughness of the gold nanograting becomes larger, from 5.97 nm to 12.26 nm. The increase of roughness and the change of surface morphology prove that RBD protein has been successfully modified on the surface of the gold nanograting array. After modifying the aptamer and protein on gold nano grating, the wavelength has a red shift. For the VCSELs sensor, the output power increases after protein modification. The successful biological modification caused the corresponding changes in the spectrum of the grating structure and the output light power of the VCSEL sensor. Power change is more sensitive than the spectral change which indicates that this VCSEL sensor can be applied to protein detection.

 

Fig. 4. Biomodification and performance characterization of the sensor. (a) bright-field and fluorescence-field image before aptamer modification. (b) bright-field and fluorescence-field image after aptamer modification. (c) AFM characterization of gold nanograting array before detection process. (d) AFM characterization of gold nanograting array after detection process. (e) transmission wavelength of the detecting process. (f) the output power of the detecting process.

Download Full Size | PDF

The specificity of the sensor is one of the important indicators for the accurate detection of proteins. In order to verify the specificity of the VCSEL sensor, 1× PBS, 1 µg/mL immunoglobulin G (IgG), and 0.1 µg/mL RBD proteins were added after the RBD aptamer was modified to test the output power change of the sensor. P-I-V characteristics of the VCSEL sensor under different solutions are shown in Fig. 5(a) and Fig. 5(b). Compared with the BSA-blocked chip (green curve), the response of the chip incubated with a high concentration of 1 µg/mL IgG (purple curve) was basically unchanged. Even though the concentration of IgG incubated on the RBD aptamer functionalized VCSEL sensor was much higher than that of the RBD aptamer, the incubation of non-specific IgG produced only small signal changes. The result illustrates an output power increases only on the positive sample, providing a qualitative indication of the presence of RBD protein.

 

Fig. 5. Specific characterization of the VCSEL sensor. (a) P-I-V characteristics of the sensor immobilized RBD aptamer to different analytes (PBS, 1 µg/mL IgG, and 0.1 µg/mL RBD). (b) output power variation of the sensor immobilized RBD aptamer to different analytes.

Download Full Size | PDF

The integrated VCSEL sensor was applied to detect the RBD protein. We tested the P-I-V characteristics and spectral characteristics of different concentrations of recombinant RBD protein. First, RBD protein solutions with different concentrations of 0.50 ng/mL, 31.25 ng/mL, 500 ng/mL, 1 µg/mL, 10 µg/mL, and 50 µg/mL were prepared and introduced into the microfluidic channel from low to high to test the P-I-V characteristics. To compare with the gold grating on quartz, we also detect the wavelength shift of different concentrations of protein (5 ng/mL, 10 ng/mL, 100 ng/mL, 1000 ng/mL). Results are shown in Fig. 6. The output power change of the VCSEL sensor is more sensitive than the wavelength shifts of gold grating on quartz. The resolution of spectral shift is limited by the spectrometer, and the peak of the transmission spectrum is smooth. Thus, it is easy to introduce errors when looking for the peak wavelength (Fig. 6(a)). In a small fluctuation range of protein concentration, the spectral shift is not obvious such as 100 ng/mL and 1000 ng/mL (Fig. 6(b)). At the same time, the laser-intense single-wavelength laser and the nW level P-I-V test system extend the concentration range of the nanograting structure response to SARS-CoV-2 RBD protein and become more sensitive. The P-I-V characteristic of the VCSEL sensor under different concentrations of RBD protein is shown in Fig. 6(c). As the RBD protein concentration increased, the output power of the VCSEL biosensor gradually increased. It can be seen from the V-I characteristic curve that the integrated VCSEL biosensor has high working stability, which allows for achieving accurate detection. In addition, it can also be seen from Fig. 6(c) that both the threshold current and the on voltage are small. That is, a small current can make the biosensor work.

 

Fig. 6. Detection capability of SARS-CoV-2 RBD with grating structure and integrated VCSEL sensor. (a) transmission wavelength shifts under different RBD protein concentrations. (b) histogram of transmission wavelength under different RBD protein concentrations. (c) P-I-V characteristics under different RBD protein concentrations. (d) the logistic curve of the relationship between different concentrations of RBD protein and output power variation of VCSEL. The red dashed line is the standard concentration fitted by the Hill equation.

Download Full Size | PDF

In order to further verify the reliability of this integrated VCSEL biosensor, we analyzed the RBD protein concentration range (0.50 ng/mL, 31.25 ng/mL, 500 ng/mL, 1 µg/mL, 10 µg/mL and 50 µg/mL). The average output power variation of the three tests is shown in Fig. 6(d). Output power variation (ΔP) generated by SARS-CoV-2 RBD is related to the concentration [RBD] and can be calibrated with the Hill equation, which is the most commonly used to describe antigen-antibody interaction [35,36]:

Among them, $\mathrm{\Delta }{\textrm{P}_{\textrm{max}}}$ is the maximum output power variation of the biosensor. k is the antibody-antigen affinity constant. According to the best fitting results, the integrated VCSEL biosensor is $\mathrm{\Delta }{\textrm{P}_{\textrm{max}}}\textrm{ = 0}\mathrm{.21\ \pm 0}\textrm{.02 mW}$ and $\textrm{k = 160}\mathrm{.43\ \pm 92}\textrm{.29 ng/mL}$. When the protein concentration increased from 0.50 ng/mL to 50 µg/mL, the output power of VCSEL gradually increased. The trend was consistent with the previous simulation and test results. For the detection of the SARS-CoV-2 RBD protein, the minimum detection concentration of the VCSEL biosensor is similar to that of the traditional LSPR sensor in the literature [11], that is, less than 1 ng/mL. When the RBD protein concentration increased to 50 µg/mL, the output power tends to be saturated, indicating that the test protein concentration tends to be saturated. This result shows that the integrated VCSEL biosensor can detect protein sensitively and has a wide detection range.

4. Conclusion

Here, we reported an integrated VCSEL biosensor that is portable and label-free. VCSEL is located at the bottom of the chip. When the laser is emitted from the top, the microfluidic channel is used to introduce the RBD recombinant protein solution with different concentrations into the gold nanograting sensing layer that has been modified to capture RBD aptamer. The optical power sensitivity of the biosensor can reach 2.99 × 10⁶ nW/RIU. With the combination of protein and aptamer on the sensing layer, the refractive index of the surrounding environment is changed, which causes the change of laser power. This biosensor has a wide detection range compared with the single nanograting sensor and can detect RBD protein from 0.50 ng/mL to 50 µg/mL. The biosensor effectively combines semiconductor laser, nanostructure-based localized surface plasmon resonance technology, and microfluidic chip technology. It provides an important reference value for the point-of-care testing (POCT) of the optical biosensor. By using these advantages of high sensitivity, less sample consumption, and high integration, the integrated VCSEL biosensor may have potential application in biomarker detection for disease diagnosis. In the future, the biosensor can also be integrated with the detection component, to realize a highly integrated and portable biosensor system and ultra-resolution imaging system.