KR101114188B1 - Polymeric waveguide biosensors and method of manufacturing the same - Google Patents

Abstract

The present invention relates to a label-free biosensor with improved sensitivity using a low refractive index polymer. In detail, the present invention relates to a low-cost portable biosensor for detecting a phenomenon in which the wavelength of the Bragg lattice structure to satisfy the Bragg reflection condition is changed when the biomaterial is bonded to the sensing portion of the sensor. The present invention is composed of a lower cladding and an optical waveguide core formed of a low refractive index polymer on a silicon substrate, and the optical waveguide core includes a Bragg grating structure.

Bio sensor, Label-free bio sensor, Low reflective index polymer, Polymeric waveguides, Bragg reflection optical waveguides

Description

Polymer waveguide biosensors and method for manufacturing the same

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a biosensor required for early diagnosis of a disease by a specific biomaterials, drug research, environment and food inspection, and a method of manufacturing the same. The present invention relates to a technique for providing a sensitivity, having a simple structure composed of a polymer optical waveguide device and a light source, and having a simple measurement method.

Label-free biosensors using optical waveguides known to date include Bragg grating engraved optical waveguides and Young interferometers. This is because, as the biomaterial is bonded to the sensing site on one of the two optical waveguides, the effective refractive index of the mode traveling along the optical waveguide is changed, so that the interference phenomenon of light traveling through the two optical waveguides at the output end is changed. It is a way to observe the change of optical power at the output stage by using the principle.

However, the label-free biosensors of the above technology have a disadvantage of requiring a high cost due to a complicated manufacturing process, a very complicated measurement method, and a variety of equipments.

On the other hand, there is an optical waveguide biosensor of a relatively simple measuring method compared to the above technology. Integrate the Mach-Zehnder interferometer into the optical waveguide and observe the interference phenomena caused by the combination of biomaterials in a similar manner as above. Measurement equipment and methods are simplified because light is input at one end of the optical waveguide and the optical power output at the other end is sensed.

However, the above technique is limited in the measurement range, such as the concentration of bio-materials, there is a disadvantage that the manufacturing process is somewhat complicated. In addition, since the silica material is used, it is not easy to improve and control the sensitivity by using a material having various refractive indices.

On the other hand, there is a Bragg grating optical waveguide biosensor, which is very simple in fabrication process compared to the above technology. It has a structure in which an optical waveguide is manufactured and a Bragg grating is formed on the optical waveguide using a nano imprinting process. When the biomaterial is bonded to the sensing area where the Bragg grating is located, the reflected wavelength or the reflected angle varies depending on the concentration of the biomaterial.

However, even in the above technique, the measurement method is complicated to measure the reflected wavelength or the angle of the reflected light, and thus, it is disadvantageous that expensive measurement equipment is required.

Meanwhile, development of a vertically incident Bragg grating biosensor having a relatively simple manufacturing process and measuring method is also in progress. By forming only the Bragg grating structure on the transparent substrate, the manufacturing process is very simple. In addition, since the light is incident in the vertical direction under the transparent substrate and the light having the wavelength satisfying the Bragg reflection condition is measured to be returned to the incidence direction, the measurement method is very simple.

In addition, in order to improve the sensitivity of the label-type biosensor based on the vertical incidence method described above, a number of studies have been conducted applying the surface plasmon phenomenon. This is formed by forming an optical waveguide and a Bragg grating structure on a transparent substrate, and then forming a metal thin film thereon. The effective refractive index of the surface plasmon mode changes sensitively to the refractive index change occurring at the sensing site, thereby improving the sensitivity of the biosensor to the binding of the biomaterial.

However, when a biolabeled biosensor is manufactured using a silica material as in the above techniques, it is not easy to improve and adjust the sensitivity of the sensor according to the biomaterial to be measured.

Meanwhile, in order to improve the sensitivity of the Bragg grating type optical waveguide biosensor, a number of techniques for improving the optical waveguide structure have been developed. In this method, the reverse symmetry optical waveguide structure was introduced to fabricate an optical waveguide having a relatively lower refractive index than the material to be measured. In addition, nanoporous silica materials are used to reduce the refractive index, which improves the sensitivity of the biosensor.

However, since the thin film deposition process in a vacuum state such as dip-floating deposition is indispensable to the method for improving the sensitivity as described above, there is a disadvantage that expensive equipment and complicated process processes are required in the manufacturing process.

As described above, the unlabeled biosensor, which is widely used or studied in the past, has been studied in the form of using an optical interferometer or Bragg grating. In addition, various studies have been made to improve the sensitivity of the biosensor by forming a metal thin film and applying a surface plasmon phenomenon or applying a reverse symmetry optical waveguide structure. However, there is a problem that the manufacturing process is complicated, the measurement procedure is complicated, and expensive measurement equipment is required, which is not suitable for application to low-cost portable biosensors.

It is an object of the present invention to fabricate an unlabeled biosensor using a polymeric material and to introduce a low refractive index polymeric material to improve the sensitivity of the sensor.

Another object of the present invention is to implement a low-cost, label-free biosensor with low manufacturing cost and simple process.

Still another object of the present invention is to implement a portable biosensor capable of real-time measurement due to the simple measurement equipment and method.

The present invention relates to a substrate; A polymer cladding formed on the substrate; And a polymer core formed on the cladding, and a Bragg grating is formed on the core.

In the present invention, it is possible to detect the change in the wavelength of light that satisfies the Bragg reflection conditions due to the bonding of the biomaterial to the Bragg grating.

In the present invention, the cladding may be formed to have a refractive index close to the refractive index of the solution in which the biomaterial to be detected is dissolved.

In the present invention, when light is input into the polymer optical waveguide biosensor, light of a specific wavelength that satisfies the Bragg reflection condition according to the effective refractive index of the mode traveling through the polymer optical waveguide biosensor is reflected, and Light may be output by traveling along the polymer optical waveguide biosensor.

In the present invention, the refractive index of the cladding may be formed to be smaller than the refractive index of the core.

In the present invention, the cladding may include a thermosetting polymer, and the core may include a UV curing polymer.

In accordance with another aspect of the present invention, there is provided a method for forming a cladding on a substrate; Patterning an optical waveguide pattern on the cladding; Forming a core on the cladding in which the optical waveguide pattern is formed; And patterning a Bragg grating pattern on the core.

According to the present invention, the sensor sensitivity can be improved by using various refractive indices of the polymer material, and the sensitivity can be easily adjusted according to the biomaterial to be measured.

In addition, by fabricating a biosensor device using a polymer material, the manufacturing process is simplified, and the biosensor can be easily miniaturized and integrated, thereby lowering the unit cost and obtaining an effect that can be applied in various fields.

Furthermore, in the polymer optical waveguide fabrication process, by applying integrated optical technology, an optical device such as a light source, a tunable laser, an optical coupler, etc., in addition to a Bragg grating optical waveguide, can be easily integrated onto a substrate and applied to a portable biosensor. Can be obtained.

Hereinafter, with reference to the accompanying drawings will be described in detail the sensitivity of the biosensor using the low refractive index polymer of the present invention. The drawings introduced below are provided by way of example so that the spirit of the invention to those skilled in the art can fully convey. Therefore, the present invention is not limited to the drawings presented below and may be embodied in other forms. Also, throughout the specification, like reference numerals designate like elements.

Hereinafter, the technical and scientific terms used herein will be understood by those skilled in the art without departing from the scope of the present invention. Descriptions of known functions and configurations that may be unnecessarily blurred are omitted.

FIG. 1A is a plan view schematically illustrating a polymer optical waveguide biosensor 1 according to an exemplary embodiment of the present invention, and FIG. 1B is an enlarged side view of portion A of FIG. 1A.

1A and 1B, the polymer optical waveguide biosensor 1 of the present invention includes a core 13 on which a substrate 11, a cladding 12, and a Bragg grating 14 are formed. Here, in FIG. 1B, the receptor module 15 and the analyte module 16 are coupled to the Bragg grating 14 of the core 13.

In detail, a polymer optical waveguide including a cladding 12 and a core 13 is formed on the silicon substrate 11, and a Bragg grating 14 is formed on the core 13 of the optical waveguide. When light having a broad wavelength is input to an input of an optical waveguide, light having a specific wavelength that satisfies Bragg reflection conditions according to the effective refractive index of the mode in which the optical waveguide proceeds is reflected, and light of the remaining wavelengths is along the optical waveguide. Proceeds and is output from the output. When the biomolecule to be measured is bound to the surface of the biosensor in FIG. 1, the effective refractive index of the optical waveguide mode is changed, and the specific wavelength satisfying the Bragg reflection condition is changed.

Here, the sensitivity of the polymer optical waveguide biosensor 1 is defined as the ratio of the amount of change in the refractive index of the material in contact with the sensor surface and the amount of change in the effective refractive index accordingly. In the case of a sensor measuring the change in the Bragg reflection wavelength, the sensitivity may be defined as the ratio of the change in the refractive index of the material in contact with the sensor surface and thus the change in the Bragg reflection wavelength. The sensitivity of the polymer optical waveguide biosensor 1 increases as the difference in refractive index between the two polymer materials constituting the core 13 and the cladding 12 increases.

Therefore, the cladding 12 is composed of a low refractive index polymer material having a refractive index similar to that of a solution in which biomolecules are dissolved, and the core 13 is composed of a polymer material having a high refractive index, thereby maximizing the difference in refractive index between the two polymer materials. In this case, a polymer optical waveguide biosensor 1 having a very high sensitivity can be realized.

2 is a graph illustrating a change in sensitivity of a polymer optical waveguide biosensor for a change in thickness of a core and a change in refractive index of a cladding using an effective refractive index calculation method.

In order to calculate the sensitivity of the polymer optical waveguide biosensor, the refractive index of the material in contact with the surface of the sensor was applied to change from water (n = 1.333) to glycerin (n = 1.472), and the sensitivity depends on the difference in refractive index between the core and the cladding. In order to compare the changes, the refractive index of the core was 1.562, and the refractive index of the cladding material was 1.334 and 1.440, respectively.

According to FIG. 2, the sensitivity of the biosensor will be maximized when the refractive index of the cladding material is 1.440 for a core thickness of 1 μm or less. However, a polymer optical waveguide with a too thin core will have a coupling loss of light input from the optical fiber and a scattering loss occurring in the optical waveguide, which is approximately 1.8 times. The 0.9 μm of the expected improvement in sensitivity was limited to the minimum thickness of the core.

3A-3F illustrate a process for fabricating the polymer optical waveguide biosensor of the present invention using a photolithography process. 3A to 3D are front views of the polymer optical waveguide biosensor of FIG. 1A, and FIGS. 3E and 3F are side views of the polymer optical waveguide biosensor of FIG. 1A.

The polymer optical waveguide biosensor according to an embodiment of the present invention uses a UV curable polymer and a thermosetting polymer to form an optical waveguide. UV-curable polymers are simple to form multi-layered structures using coatings, but thermosetting polymers are hydrophobic, so they are surface treated with oxygen plasma and coated with an adhesion promoter when coating other polymers on top of them. Processing is required.

In order to fabricate a polymer optical waveguide biosensor according to an embodiment of the present invention, first, as shown in FIG. 3A, a thermosetting polymer is coated on a silicon substrate 11 by about 5 to 10 μm and thermally cured. Make the lower cladding 12. Next, in order to form an optical waveguide pattern on the lower cladding 12, as shown in FIG. 3B, a photoresist PR1 is coated and UV is irradiated on the photoresist M1 thereon. Then, by performing an etching process, the optical waveguide pattern 12a as shown in FIG. 3C is formed.

Next, after surface treatment of the surface of the thermosetting polymer, the UV curable polymer is coated and cured with the optical waveguide core 13 as shown in FIG. 3D.

Next, as shown in FIG. 3E, to form a Bragg grating on the optical waveguide core 13, a laser interferometer coated with a photoresist PR2 and equipped with a 488 nm argon laser (Ar laser). The photolithography process is performed to form a Bragg grating pattern having a 510 nm period. Similarly, the pattern portion is removed by a dry etching process by 100 to 300 nm to form a Bragg grating 14 as shown in FIG. 3F.

4 is a graph showing the spectrum of light output from the two polymer optical waveguide biosensors 1 having different heights of Bragg gratings.

The reflectance by Bragg grating structure is strongly dependent on the thickness of the core and the height of the grating structure. In particular, in the case of the optical waveguide structure having a large refractive index difference, it has a very large reflectance even at a low grating height. As shown in FIG. 4, when the reflectance increases, the end of the Bragg reflection wavelength is lowered below the noise level in the spectrum output from the sensor, and the bandwidth of the Bragg reflection wavelength becomes thick, thereby making accurate measurement. It becomes difficult to do Therefore, it is desirable to limit the Bragg grating of the polymer optical waveguide biosensor 1 to 50 nm in height and 1 mm in length. At this time, the light spectrum output from the sensor is as shown in Figure 4 and the 3-dB bandwidth (bandwidth) is 0.9 nm.

FIG. 5 is a graph illustrating a mode profile of light traveling through the polymer optical waveguide biosensor 1 using a charge coupled device (CCD).

According to FIG. 5, the core thickness of the polymer optical waveguide biosensor 1 is 0.9 μm, but the size of the mode profile at the output end is 6.5 μm due to the deep penetration of the evanescent field into the cladding. do. The total optical loss measured by connecting single-mode fiber to the input and output of the biosensor is 20 dB, which is caused by the waveguide scattering loss and the mode mismatch loss. Therefore, the use of lenticular fiber instead of single-mode fiber at the input of the biosensor can reduce the total optical loss by 15 dB.

FIG. 6 is a graph illustrating measurement and change of Bragg reflection wavelength by dropping water and glycerin on the surface of a biosensor with two polymer optical waveguides having claddings having different refractive indices of 1.440 and 1.334.

As the material on the surface of the polymer optical waveguide biosensor is changed from water to glycerin, the effective refractive index of the optical waveguide mode is changed to change the wavelength of light that satisfies the Bragg reflection condition. In the case of the biosensor manufactured by the cladding having a refractive index of 1.334 in FIG. 6 (b), the Bragg reflection wavelength is shifted by 11.4 nm with respect to the biosensor manufactured by the cladding having a refractive index of 1.440 in FIG. It was confirmed that the wavelength shifted by 21.9 nm, and the sensitivity was improved by 1.9 times in the biosensor using a low refractive index polymer of 1.334.

In order to measure the binding of biomolecules, a fluidic channel structure was introduced to control very small amounts of biomaterials. First, a UV curable polymer was formed on a silicon substrate in a lithography process to form a pattern of fluid channels of 500 μm in width and 50 to 100 μm in height. Then, PDMS (Polydimethylsiloxane) was poured on the pattern, cured, and then peeled off, and then attached to the sensor surface using a UV curable epoxy, and a hose was connected.

FIG. 7 is a graph illustrating the measurement of anti-biotin at 33.2 and 531.2 nM concentration using the polymer optical waveguide biosensor according to the present invention.

According to FIG. 7, it is evident that the Bragg reflection wavelength changes as the effective refractive index of the optical waveguide is changed by biomolecule bonding to the surface of the biosensor. When observing the shift of the Bragg reflection wavelength, it can be seen that the binding of the biomolecule proceeds rapidly during the first 10 minutes and then slowly slows down, and after about 20 minutes, the binding of the biomolecules is saturated.

8 is a graph illustrating the measurement of anti-biotin at the same concentration according to the surface state of the polymer optical waveguide biosensor according to the present invention.

Anti-biotin (anti-biotin) used in the bio-detection in the present invention is an antigen antibody reaction with biotin (biotin), the reaction between the two is made of a specific binding. In order to attach biotin to the polymer surface, an amine group is formed on the polymer surface. This results in non-specific binding of amine groups and anti-biotin on the polymer surface, which can lead to unwanted shifts in Bragg reflection wavelengths. In addition, anti-biotin may be physically adsorbed on the polymer surface to allow the shift of the Bragg reflection wavelength. It can be suppressed by introducing a blocking buffer process, which is a buffer solution containing bovine serum albumin (BSA) and Tween 20 (Polysorbate 20) to the polymer surface. The biotin is adsorbed to an amine group to which no biotin is attached, thereby removing the space where the anti-biotin is adsorbed.

In order to observe the effect of blocking buffer treatment, blocking buffer treatment was performed for 30 minutes, and the anti-biotin concentration of 100 ug / ml diluted in deionized water was micro Injection into the biosensor surface via a fluidic channel. In case of injecting anti-biotin into APTES and injecting NHS-PEO biotin into APTES and injecting anti-biotin, the presence of blocking buffer The measurement was carried out, and the results are shown in FIG. 8. In (a) of FIG. 8, the peak wavelength is shifted by non-specific binding between anti-biotin and APTES, whereas in FIG. 8 (b), blocking buffer is blocked. It was clear that the peak wavelength hardly shifted due to the non-specific binding caused by the buffer treatment. In addition, similar results were observed for FIGS. 8C and 8D, and in FIG. 8D, the shift of the peak wavelength due to non-specific binding after a certain amount of time was observed. It was confirmed that it continued to occur.

FIG. 9 is a graph showing measurement results of biosensors for various concentrations of anti-biotin.

Before injecting the anti-biotin solution into the sensor, the anti-biotin solution was injected through the microfluidic channel after the sensor surface treatment with a blocking buffer to prevent nonspecific binding. Characterization as a biosensor was measured using an anti-biotin solution prepared at various concentrations from 8.3 nM to 664 nM. After the measurement experiment, cleaning the surface of the sensor did not change the Bragg reflection wavelength, and it was found that the anti-biotin bond was not broken. In addition, it was confirmed that the Bragg reflection wavelength was linearly changed for the high concentration of anti-biotin, and the measurement was performed up to a concentration of 8.3 nM.

As described above has been described by the limited embodiment and the drawings with respect to the sensitivity of the polymer optical waveguide biosensor, biosensor using a low refractive index polymer of the present invention, which is provided only to help a more general understanding of the present invention, the present invention The present invention is not limited to the above embodiments, and various modifications and variations are possible to those skilled in the art to which the present invention pertains.

Therefore, the spirit of the present invention should not be limited to the described embodiments, and all of the equivalents and equivalents of the claims as well as the claims to be described later belong to the scope of the present invention. .

FIG. 1A is a plan view schematically illustrating a polymer optical waveguide biosensor 1 according to an exemplary embodiment of the present invention, and FIG. 1B is an enlarged side view of portion A of FIG. 1A.

FIG. 2 is a graph illustrating a change in sensitivity of a polymer optical waveguide biosensor for a change in thickness of a core and a change in refractive index of a cladding using an effective refractive index calculation method.

3A-3F illustrate a process for fabricating the polymer optical waveguide biosensor of the present invention using a photolithography process.

4 is a graph comparing spectra of light output from two polymer optical waveguide biosensors having different Bragg grating heights.

FIG. 5 is a graph illustrating a mode profile of light traveling through the polymer optical waveguide biosensor 1 using a charge coupled device (CCD).

FIG. 6 is a graph illustrating measurement and change of Bragg reflection wavelength by dropping water and glycerin on the surface of a biosensor having two refractive waveguides having different refractive indices of 1.440 and 1.334, respectively.

FIG. 7 is a graph illustrating the measurement and measurement of anti-biotin at 33.2 and 531.2 nM concentration using the polymer optical waveguide biosensor according to the present invention.

8 is a graph observed by injecting anti-biotin of the same concentration according to the surface state of the polymer optical waveguide biosensor of the present invention.

FIG. 9 is a graph showing measurement results of biosensors for various concentrations of anti-biotin.

Claims (4)

Board; A polymer cladding formed on the substrate; And A polymer core formed on the cladding, Bragg grating is formed in the core, The polymer optical waveguide biosensor for detecting a wavelength change of light that satisfies the Bragg reflection condition of the bonding of the biomaterial to the Bragg grating. delete The method of claim 1, The cladding is a polymer optical waveguide biosensor, characterized in that formed to have a refractive index close to the refractive index of the solution in which the biomaterial to be detected is dissolved. The method of claim 1, When light is input into the polymer optical waveguide biosensor, light of a specific wavelength that satisfies Bragg reflection conditions according to an effective refractive index of a mode traveling through the polymer optical waveguide biosensor is reflected, and light of the remaining wavelengths is emitted from the polymer. Polymer optical waveguide biosensor, which is output along the optical waveguide biosensor.

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