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This paper proposes a label-free and spectrometer-free method for biological detection with high detecting resolution. Taking advantage of the optical properties of porous silicon microcavity, the refractive index changes caused by biological reaction can be detected by measuring the incident angle of the minimum reflected light intensity. Based on the above method, label-free eight-base pair DNA detection can be realized with a corresponding detection limit is as low as 87 nM. This method provides high detecting resolution at a low equipment cost, and can be further used to develop an advanced instrument for biological detection.
© 2015 Optical Society of America
1. Introduction
Porous Silicon (PS) is a promising substrate material for biosensors with many distinct advantages, including large specific surface area, good biological activity, compatibility and adsorption [1,2 ], selectivity based upon the size of entry biological molecules, good chemical stability [3] and the capability to form a variety of photonic structures. Therefore, PS is often used to build high detecting resolution biosensors. PS biosensors reported to date primarily utilize a multilayer grating structure or microcavity structure for biological detection [4–9 ], fluorescence detection [10,11 ], surface enhanced Raman Scattering [12], and other functions.
Among existing PS biosensors, the sensor based on the Porous Silicon Microcavity (PSM) structure is a porous silicon photonic crystal device with a one-dimensional defect mode. The reflection spectrum displays high defect peak transmittance and narrow Full Width at Half Maximum (FWHM) [13]. Some current research seeks to improve microcavity sensor detection by implementing a porous silicon microcavity multilayer ternary structure for insect antifreeze protein gene hybridization experiment detecting [14].
This paper proposes a novel biological detection method utilizing microcavity structure. A visible laser (e.g. a semiconductor laser) serves as the light source, with a wavelength equal to the transmission wavelength corresponding to the one-dimensional defect mode of porous silicon microcavity structure. Biological detection is accomplished by measuring the incident angle between the laser and the surface of the PSM corresponding to the minimum reflected light intensity, in order to obtain the refractive index change caused by biological reaction. This method is label-free and spectrometer-free. It possesses a high detecting resolution, and can detect changes in the refractive index as small as 10−4 orders of magnitude with ordinary angle-measuring devices.
2. Theory
In order to synthesize a one-dimensional photonic crystal porous silicon device with a band gap structure, we can obtain high (H) and low (L) refractive index layers by changing the current density in the process of anodic oxidation. Because the device is used for biological detection, the microcavity (MC) layer must be introduced into the periodic structure, thus forming the PSM structure [15–17 ] as shown later.
As the light incident angle on the surface of the PSM changes, the reflection spectrum position of the PSM will be changed. For a beam of Transverse Electric (TE) mode incident light, the refractive indices n H, n L and n C are 1.52, 1.21 and 1.21, respectively, where n H, n L and n C represent the high refractive index layer, low refractive index layer, and the defect layer, respectively. The optical thickness of two dielectric layers is equal and represented by n H d H = n L d L = λ C/4. The optical thickness of the MC layer n C d C = λ C/2, where d H, d L and d C are the thickness of high, low refractive index layer, and defect layer, respectively, and λ C is the reference wavelength, 633 nm. There are 27 total middle layers of material in the PSM. The reflection spectrum of the PSM device designed by the above parameters is calculated by the transfer matrix method with incident angle θ = 0°, 10°, 20° and 30°, as shown in Fig. 1(a) . From theoretical calculation, the resonance peaks corresponding to the incident angle θ of 0°, 10°, 20° and 30° occur at 633 nm, 628 nm, 611 nm and 584 nm, respectively. The calculation indicates that there is a blue shift in the reflection spectra of PSM as the incident angle increases.
Fig. 1 (a) PSM reflection spectra of incident angles θ = 0°, 10°, 20° and 30°, respectively; (b) Curve (a) represents the PSM reflection spectrum of vertical incidence (θ = 0°). Curve (b) represents the occurrence of a biological reaction in the PSM device, causing the refractive index to increase Δn = 0.01, and represents an identical PSM reflection spectrum to vertical incidence (θ = 0°). Curve (c) represents the PSM reflection spectrum after a certain rotary angle of incident light.
Under the condition of vertical incidence for the incoming light beam (θ = 0°), the wavelength of the resonance peak (λ C) is 633 nm. The biological reaction occurring in the PSM device leads to an increase in the refractive index (Δn = 0.01) for each layer of porous silicon, which causes a red shift in the reflection spectrum. The wavelength of transmission light corresponding to the defect mode has an increase (Δλ C) of 5 nm, labeled (a) and (b) as shown in Fig. 1(b). When the incident angle of the light source increases to 2.9°, the resonance peak (λ C) of the PSM remains unchanged from the condition of vertical incidence, indicating that these two resonance peaks overlap as shown by curve (c) and curve (a) in Fig. 1(b).
For the present PSM device, it can be concluded from Fig. 1 that Δn = 0.01 results in Δλ C of 5 nm. This is the common method to obtain the refractive index change from reflection spectrum. Here we propose a new angle measurement method to determine the refractive index variation of a PSM device. In this method, a single-wavelength visible laser (e.g. an He - Ne laser, a semiconductor laser), with wavelength (λ i) equal to the designed λ C, is used to irradiate the PSM device without reflection at normal incidence. Changes in the refractive index values caused by the biological reaction in vertical incidence to the PSM again will reflect light. If the laser incident beam is adjusted to a certain angle of θ, the reflected light will disappear again. Therefore, we can determine the refractive index variation Δn caused by the change of the incident angle of the laser Δθ.
Our calculations indicate that Δθ = 0.046° and Δλ C = 0.048 nm when Δn = 0.0001. For angle spectrum measurements, although the resolution of the general angle measuring instrument is 1ˊ, i.e. 0.0167°, it is difficult to measure the angle shift Δθ due to FWHM of angle spectrum in experiment. For spectroscopic measurements, the resolution of the optical experiment equipment for laboratory spectrum detection (UV-VIS spectrophotometer, Hitachi U - 4100) is generally 0.1 nm, and detection of a refractive index value change of Δn = 0.0001 can only be completed by an instrument with a resolution higher than 0.1 nm. In the case of Δn = 0.0005, Δθ = 0.22° and Δλ C = 0.21 nm. The angle shift Δθ is measurable, considering the uncertainty of our optical detector.
The plane wave expansion method is utilized to calculate the change of the forbidden band of PSM with the TE wave incident angle as shown in Fig. 2 . The change of the band gap width of the TE wave is negligible compared to the change in position. The red line in Fig. 2 represents the position of 633 nm wavelength at the forbidden band corresponding to the different incident angles. The black line corresponds to the position of incident angle θ of approximately 67°, indicating that the 633 nm wavelength will be beyond the forbidden band when the incident angle increases above 67°.
It can be determined from the above analysis that the experiment equipment is inexpensive in order to achieve the angle spectrum measurement method. The only required elements are a Semiconductor Laser Diode (LD), light detector, angle measuring instrument and optical elements, thus it is possible to make a portable device. Such a device based on porous silicon microcavity structure consisting of a biological sensor and array detection will be efficient in terms of inspection time.
3. Experiments
3.1 The preparation of PSM
PSMs are prepared by electrochemical anodization of highly doped p-type silicon (boron doped, 0.03–0.04 Ω·cm resistivity) in a mixed solution composed of 40% aqueous hydrofluoric acid and 99% alcohol (1:1 by volume). The electrolytic corrosion groove for Poly Tetra Fluoro Ethylene (PTFE) possesses a single groove structure, with a silicon anode and a copper cathode.
The silicon wafers should be cleaned before the experiment. In preparation, the wafers are soaked in cleaning fluid (volume ratio H2SO4: H2O2 = 3:1) at room temperature until the reaction completes, meaning that organic pollutants on the surface of the wafers have been completely removed; the wafers are then rinsed with deionized water. Next, to remove residual surface impurities, the silicon wafers are cleaned with acetone, alcohol and deionized water, respectively, in the ultrasonic instrument for 10 minutes each.
The silicon wafers are then anodized using a computer controlled labview program. As shown later, the two DBRs are fabricated with a current density alternating between 60 mA/cm2 for 1.2 s and 110 mA /cm2 for 1.0 s. The optical cavity defect layer that separates the two DBRs is formed by employing a current density of 110 mA /cm2 for 2.0 s. In order to ensure that the corrosion is relatively uniform, there are timely supplements of fluoride concentration and a 3 seconds pause after each layer formation.
Figure 3 illustrates the Scanning Electron Microscope (SEM) images of the surface and the cross section of the porous silicon microcavity structure. Figure 3(a) displays the porous property from the top surface, demonstrating that the diameter of the mesh pores varies from 10 to 20 nm. As shown in Fig. 3(b), the entire thickness of the PSM structure is approximately 3.1 μm. The pore size of the porous silicon ensures a higher biological infiltration function than mesoporous and macroporous materials [18], allowing DNA molecules to better infiltrate the PSM and react, thus causing the change in the PSM refractive index to allow biological detection.
3.2 Functionalization
After preparation of the PSMs, the functional processing is integral to the function of the biosensor device. The fresh porous silicon is easily oxidized in the air; therefore, in order to stabilize the surface structure of the porous silicon, the PSMs are put into a concentration of 40% hydrogen peroxide at 60° C and placed in a vacuum drying oven for 3 hours. Then the PSMs are removed and repeatedly rinsed with deionized water before drying in the air. Next, the PSMs are soaked for one hour in 5% of APTES solution, composed of APTES (99%), alcohol (99.5%) and deionized water with a volume ratio of 1:10:10. To remove excess solvent, the PSMs are then flushed repeatedly with deionized water and baked at 100° C for 10 minutes in a vacuum drying oven. Finally, the PSMs are put into 2.5% solution of glutaraldehyde for 1 hour at room temperature, to promote connections with biological molecule. Finally, the PSMs are washed three times with Phosphate Buffer Solution (PBS, pH = 7.4) to remove excess glutaraldehyde.
3.3 Immobilization
The eight-base pair DNA fragments if sequence 5′-TGCAACGT-3′ are selected for the DNA probe. The sequences of complementary and non-complementary DNA are 5′-ACGTTGCA-3′ and 5′-TAGCCGAT-3′, respectively, with lengths of approximately 1.76 nm [19]. 40 µL about 10 µM of amino-modified DNA probe are applied to PSMs and incubated for 2 hours, followed cleaning with PBS. Finally, the PSMs are dipped into a 3M ethanolamine buffer with PH 9.0 for 1 hour at 37° C to close any unreacted aldehyde groups and minimize non-specific binding [20], followed by post-process cleaning with PBS, and then dried in the air.
3.4 Detection
The PSMs are dropped separately into 40 µL of solution containing different concentrations of complimentary DNA in buffer, 40 µL of non-complimentary DNA in buffer, as well as buffer solution alone, and incubated in thermostat for 1 hour. Afterwards, all samples are rinsed with buffer repeatedly and dried in the air.
3.5 Detection apparatus
The experimental device used to implement our proposed biological detection method based on the nanoporous silicon microcavity is shown in Fig. 4 . The experimental device is composed of a laser, polarizer, lens, apertures, and turntable with scale and detector. The laser utilized in this experiment is a high stability He-Ne laser with a divergence angle of 0.79 mrad and a wavelength of 633 nm. The laser is Transverse Electromagnetic (TEM) wave, obtained by polarizing prism P. The polarizing laser Gaussian beam is expanded to become an approximate plane wave for the PSM with a collimating beam expander through two lenses L.
4. Results and Discussion
Oblique incidence angle measurements are taken after each functionalization step to confirm chemical attachment. Figure 5(a) demonstrates the changes in angle for each functionalization step, in which they are from left to right: after oxidation, after silanization and after addition of glutaraldehyde. The angle change is due primarily to the change of the effective refractive index. After the above steps, the angle changes demonstrate that the functionalization process is successfully in this experiment. Figure 5(b) demonstrates the detection of DNA in the PSM. The angle spectrum shifts approximately 12.5° after the PSM in which the immobilized 10 µM DNA probe is exposed to 5µM of complimentary DNA. The effective refractive index change is primarily due to complementary DNA and immobilized probe DNA effectively combining in the PSM device. To demonstrate specificity, Figs. 5(c) and (d) show the angle spectra after exposure to 10 µM of non-complimentary DNA and after exposure to buffer solution, respectively, with negligible shift. A possible explanation for the negligible shifts in the angle spectra may be because the probe DNA and non-complimentary DNA in buffer do not bind in the hole of the PSM, thus resulting in no change in the effective refractive indices.
Fig. 5 (a) Angle spectra corresponding with changes in the PSM functional steps: after oxidation, silanization, and addition of glutaraldehyde. (b) The PSM angle spectra changes for complementary DNA. (c) Negligible PSM angle spectra changes for non-complementary DNA. (d) No change in the PSM angle spectra in reaction to buffer solution.
The experiment results indicate that the PSM device is able to distinguish DNA hybridization between complementary and non-complementary DNA according to oblique incidence angle measurements. The shift quantity of the angle spectrum corresponds to the different concentrations of DNA. We expect that the ultimate detection limit is much smaller than 10 µM. Figure 6 shows the shift of the angle spectrum for complementary DNA at different concentrations: 0.3125, 0.6250, 1.250, 2.500, 3.125, 5.000, 6.250 and 10.00 µM. This indicates that the different angle variations corresponding to these concentrations are 2.9°, 4.7°, 7.2°, 11.2°, 13.1°, 16.3°, 19.4°, 25.8°, respectively. Figure 6 also demonstrates a strong linearity (R = 0.970) between the shift of the angle spectrum and the concentration of complementary DNA in the range from 0.3125 to 10.00 µM. The equation of linear regression is determined to be Y = 2.30X + 4.15, where Y represents the shift of the angle spectrum and X represents the concentration of complementary DNA. Therefore, the sensitivity of the sensor is 2.30°/µM, calculated by the slope of the regression line. With the resolution of the device at approximate 0.2°, the detection limit of the biosensor is 0.2deg/(2.30deg/µM) = 87 nM. The DNA detection limit based on the PSM biosensor is in the nM order of magnitude, and is comparable to that reported for other favorable PS biosensor based on silicon by spectroscopic measurement, such as the PS polybasic symmetrical structure for detecting 23-base pair DNA with a limit of 21.3 nM [14], the PSM biosensor utilizing SOI wafers for 19-base pair DNA with a limit of 43.9 nM [20], the PS membrane waveguide for 24-base pair DNA with a limit of 42.0 nM [21], and the PSM biosensor utilizing SOI wafers for 23-base pair DNA with a limit of 5.7 nM [22].
Fig. 6 Linear relationship diagram for angle spectrum changes with the concentration of DNA from 0.3125 to 10 μM.
In this experiment, the wavelength λ C corresponding to the defect mode of PSM is designed to equal the wavelength of the laser. However, the refractive index of each layer increases in the functionalization and biological probe preparation process of the PSM device, resulting in an increase of the resonance peak wavelength. Therefore, before biological detection, the laser incident angle is adjusted from 0° to θ 1 to make the resonance peak wavelength under the oblique incidence equal to the laser wavelength. When the refractive index of the PSM device changes due to biological reaction, we continue to adjust the laser incident angle from θ 1 to θ 2 to make the resonance peak wavelength equal to the laser wavelength again. The change of refractive index Δn can be determined based on the change of the laser incident angle Δθ = θ 2-θ 1.
5. Conclusions
In this study, a novel detection method is proposed to produce a biological sensor with high detecting resolution based on porous silicon microcavity. Based on the angle spectrum detection method, the porous silicon microcavity sensor can detect the refractive index changes of 5☓10−4 with ordinary angle measuring instruments. This method is applied to the detection of eight-base pair DNA molecule hybridization, with a corresponding detection limit of 87 nM. The biological detection device based on angle spectrum measurement provides high detecting resolusion at low cost, and can therefore be widely used.
Acknowledgments
This work was supported by the National Science Foundation of China (no. 61265009 and 11264038).
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