A biological detector based on weak measurement method

Abstract

This invention discloses a biosensor based on a weak measurement method, comprising a light source, a pre-selective polarizer, a prism, a detection chip, a quarter-wave plate, a rotator, a post-selective polarizer, and an imaging module. Light emitted from the light source passes through the pre-selective polarizer and enters the prism. Total internal reflection occurs at the weak measurement interface formed by the prism and the detection chip, causing a phase difference between the P-wave and S-wave polarizations. The reflected light passes through the quarter-wave plate, the rotator, and the post-selective polarizer before entering the imaging module to achieve image formation at the total internal reflection interface. The detection chip has a detection channel and a reference channel at the corresponding weak measurement interface to achieve self-reference differential noise reduction using the difference in relative light intensity between the detection channel and the reference channel. This invention improves the anti-interference and stability of high-sensitivity biosensors and enhances the performance of weak measurement biosensors.

Description

Technical Field

[0001] This invention relates to the technical field of biological detection instruments, and in particular to a biological detection instrument based on a weak measurement method. Background Technology

[0002] Life is an eternal theme in scientific research. Since the beginning of the new century, life science research has flourished, and a series of significant breakthroughs are tangibly improving human lives. Simultaneously, the field of biosensors, formed by the integration and intersection of life sciences and multiple disciplines, has also shown remarkable vitality. Biosensors generally combine biologically sensitive materials with specific bioactive analytes such as enzymes, antibodies, antigens, proteins, nucleic acids, and small molecules. The binding process is then converted into distinguishable changes in electrical, optical, and thermal properties by a physicochemical transducer, and further amplified to convert the biochemical signals into quantifiable electrical signals. Among these, optical signals offer advantages over other physical signals, including high sensitivity, resistance to external interference, good stability, and low noise. Therefore, optical biosensors exhibit excellent performance in biological detection systems and have promoted significant breakthroughs in scientific research and industry fields such as biomedicine, food health, material characterization, drug development, medical devices, and environmental protection.

[0003] Existing optical biosensing technologies mainly include bio-layer interferometry (BLI), surface plasmon resonance technology (SPR), and biofluorescence technology.

[0004] Biomembrane interferometry employs probe-type biosensors to directly detect samples, using the shift in reflectance spectra to characterize the molecular dynamics processes bound to the sensor surface in real time. However, this technique has drawbacks: it can only perform single-point detection by the spectrometer and cannot achieve high-throughput parallel monitoring; the single-point detection method makes it difficult to eliminate errors introduced by environmental factors such as temperature and pressure, as well as system-related errors, greatly limiting its application scope and scenarios.

[0005] Surface plasmon resonance (SPR) technology involves depositing a metal film, typically tens of nanometers thick, onto a substrate plane, often using gold or silver films. This technique boasts high sensitivity and high throughput. However, its drawbacks include stringent requirements for the metal film deposition, demanding precise nanometer-level accuracy. Reproducibility is also limited, as variations in film thickness between repeated measurements can affect results. Furthermore, its physics demands a high angle of light incidence, resulting in stringent hardware precision requirements for the overall detection system and poor environmental robustness. Since the metal film is the core carrier, the pH of the biological assay system can corrode it, leading to errors and reduced reusability. While gold and silver are the most commonly used coatings, silver films are easily oxidized, and gold films readily catalyze certain reactions, raising questions about the interpretability of the results.

[0006] Biofluorescence is a commonly used technique, but the impact of fluorescent molecule labeling on the target system remains a highly controversial issue, making labeling an unavoidable problem in biological detection. The advantages of label-free methods over labeling are obvious. While fluorescent labeling technology brings high sensitivity, it also introduces a series of new and derivative problems such as fluorescence quenching and fluorescence background noise.

[0007] Since its introduction in 1988, weak measurement technology has demonstrated significant advantages in the field of high-precision measurement. In particular, the development of frequency-domain weak measurement systems and the realization of universally applicable weak measurement sensing systems have allowed weak measurement sensors to showcase their excellence in the field of biological detection. The application of weak measurement technology has greatly improved the sensitivity of sensing systems. However, biosensors based on weak measurement technology are easily affected by external factors such as temperature and vibration; therefore, improving the sensor's anti-interference capability is crucial for enhancing the performance of weak measurement biosensors. Summary of the Invention

[0008] The purpose of this invention is to provide a biological detector based on a weak measurement method, thereby improving its anti-interference and stability.

[0009] Therefore, this invention proposes a biological detector based on a weak measurement method, comprising a light source (1), a pre-selective polarizer (3), a prism (4), a detection chip (5), a quarter-wave plate (6), a rotator (7), a post-selective polarizer (8), and an imaging module (9). The light emitted by the light source (1) passes through the pre-selective polarizer (2) and enters the prism (4). Total internal reflection is achieved at the weak measurement interface formed by the prism (4) and the detection chip (5), causing a phase difference between the P-light polarization and the S-light polarization. The reflected light passes through the quarter-wave plate (6), the rotator (7), and the post-selective polarizer (8) and enters the imaging module (9) to achieve image imaging of the total internal reflection interface. The detection chip (5) is provided with a detection channel and a reference channel at the corresponding weak measurement interface so as to achieve self-reference differential noise reduction by utilizing the difference in relative light intensity between the detection channel and the reference channel.

[0010] In some embodiments of the present invention, the difference in relative light intensity is the difference between the mean values ​​of the light intensity detected by the detection channel and the reference channel.

[0011] In some embodiments of the present invention, the quarter-wave plate (6) is set at an angle of 0.5° to the vertical direction along the fast axis. Achromatic quarter-wave plate.

[0012] In some embodiments of the present invention, the optical rotator (7) is a quartz optical rotator, and a weak coupling effect is achieved through the quartz optical rotator.

[0013] In some embodiments of the present invention, the light source (1) is a superluminescent diode.

[0014] In some embodiments of the present invention, a collimating lens (2) is also included, disposed between the light source (1) and the pre-selective polarizer (3).

[0015] In some embodiments of the present invention, a motion control system connected to the post-selected polarizer (3) is also included.

[0016] In some embodiments of the present invention, the detection chip (5) is a 3D printed chip or a glass-based chip, and the chip is configured to match the refractive index of the prism.

[0017] In some embodiments of the present invention, data post-processing is performed by averaging multiple pixels within the respective inspection areas of the detection channel and the reference channel.

[0018] In some embodiments of the present invention, corresponding to The detection is performed in the weak measurement region, where τ is the coupling strength, ω0 is the center frequency of the light source, and ε is the minimum parameter. The phase difference between the P-polarized light and the S-polarized light.

[0019] The present invention has the following beneficial effects:

[0020] This invention proposes a self-referencing biosensor based on weak measurement technology. The detection chip has a detection channel and a reference channel at the corresponding weak measurement interface. Self-referencing differential noise reduction is achieved by utilizing the difference in relative light intensity between the detection channel and the reference channel, exhibiting good robustness to temperature changes. Thus, this invention ensures the extremely high sensitivity of the weak measurement-based sensor while maintaining system stability through dual-channel differential noise reduction, improving the performance of the weak measurement biosensor and enhancing its resistance to interference from external temperature changes. This results in a high-precision, high-throughput biosensor based on weak measurement. In a preferred embodiment, this invention further improves system stability by averaging the pixels. The biosensor of this invention has advantages such as simple structure, high robustness, wide application range, low detection cost, and convenient system and experimental operation. Attached Figure Description

[0021] Figure 1(a) is a schematic diagram of a two-channel chip in an embodiment of the present invention;

[0022] Figure 1(b) is a cross-sectional view of the two-channel chip in an embodiment of the present invention;

[0023] Figure 2 This is a schematic diagram of the optical path in an embodiment of the present invention;

[0024] Figure 3(a) is a schematic diagram of the prism combined with a detection chip in an embodiment of the present invention;

[0025] Figure 3(b) is a schematic diagram of the prism combined with another detection chip in an embodiment of the present invention;

[0026] Figure 4 This is a diagram of the motion control system device in an embodiment of the present invention;

[0027] Figure 5 This is a diagram of a display system device in an embodiment of the present invention;

[0028] Figure 6 This is a screenshot of the user interface of the software "QWMImage" in this embodiment of the invention;

[0029] Figure 7(a) is a schematic diagram of the heat conduction process of the simulated "chip" in an embodiment of the present invention;

[0030] Figure 7(b) is a schematic diagram of the heat conduction process of the simulated prism in an embodiment of the present invention;

[0031] Figure 7(c) is a line graph showing the refractive index change of a point-like region during the high-temperature liquid cooling process in an embodiment of the present invention;

[0032] Figure 7(d) is a line graph showing the refractive index change of the light spot region during the heating process of the low-temperature liquid in an embodiment of the present invention;

[0033] Figure 7(e) is a comparison of the relative light intensity changes of the two channels from 39.6°C to 29.7°C in an embodiment of the present invention.

[0034] Figure 7(f) is a comparison of the relative light intensity changes of the two channels from 10.1℃ to 22.3℃ in an embodiment of the present invention;

[0035] Figure 8(a) is a schematic diagram of eight concentric rectangular regions with different areas in an embodiment of the present invention;

[0036] Figure 8(b) is a histogram of the variance of pixel values ​​in eight different regions in an embodiment of the present invention;

[0037] Figure 9(a) is a curve showing the result of differentiating the relative light intensities of the two channels in an embodiment of the present invention.

[0038] Figure 9(b) is a graph showing the change in relative light intensity of IgG at different concentrations in the embodiments of the present invention;

[0039] Figure 9(c) is a line graph of the light intensity of the NaCl solution in an embodiment of the present invention;

[0040] Figure 9(d) shows the fitting curve of the Langmuir adsorption model to the experimental results in the embodiments of the present invention;

[0041] Figure 9(e) is a bar chart comparing the relative light intensity caused by different target samples in the embodiments of the present invention.

[0042] The attached figures are labeled as follows:

[0043] 1 is the power supply, 2 is the collimating lens, 3 is the front polarizer, 4 is the prism, 5 is the detection chip, 6 is the quarter-wave plate, 7 is the optical rotation plate, 8 is the rear polarizer, 9 is the imaging module, 10 is the imaging lens, and 11 is the charge-coupled device. Detailed Implementation

[0044] The present invention will be further described below with reference to the accompanying drawings and preferred embodiments. It should be noted that, unless otherwise specified, the embodiments and features described in this application can be combined with each other.

[0045] It should be noted that the directional terms such as left, right, up, down, top, and bottom used in this embodiment are only relative concepts or are based on the normal use of the product, and should not be considered as restrictive.

[0046] Leveraging the advantages of weak measurement technology and the characteristics of evanescent field sensing, weak measurement-based interface biosensors can achieve high-precision acquisition of biological information on ordinary glass interfaces. This invention proposes a self-referencing biosensor based on a weak measurement method, establishing a reference channel within the same interface to effectively solve external interference problems without requiring expensive additional detection components. This invention further improves system stability by averaging multiple pixels within the inspection area of ​​each of the experimental and reference channels while performing differential noise reduction on both channels. The differential noise reduction method here involves subtracting the mean values ​​of the light intensity calculated from the two channels to eliminate additive noise common to the instrument. The experimental and reference channels are the flow channels designed in the detection chip. Because there are various types of chips, this invention uses a simple two-channel chip for illustration, where one channel serves as the experimental channel and the other as the reference channel, as shown below. Figure 1a and Figure 1b As shown.

[0047] This invention proposes a biological detector based on a weak measurement method, comprising a light source 1, a pre-selective polarizer 3, a prism 4, a detection chip 5, a quarter-wave plate 6, an optical rotator 7, a post-selective polarizer 8, and an imaging module 9. Light emitted from the light source 1 passes through the pre-selective polarizer 3 and enters the prism 4. Total internal reflection occurs at the weak measurement interface formed by the prism 4 and the detection chip 5, causing a phase difference between the P-wave polarization and the S-wave polarization. The reflected light passes through the quarter-wave plate 6, the optical rotator 7, and the post-selective polarizer 8 before entering the imaging module 9 to achieve image imaging of the total internal reflection interface. The detection chip 5 has a detection channel and a reference channel corresponding to the weak measurement interface, enabling self-reference differential noise reduction by utilizing the difference in relative light intensity between the detection channel and the reference channel.

[0048] In a preferred embodiment, the difference in relative light intensity is the difference between the mean values ​​of the light intensity detected by the detection channel and the reference channel.

[0049] In a preferred embodiment, the quarter-wave plate 6 is positioned with its fast axis at an angle of [value missing] to the vertical direction. Achromatic quarter-wave plate.

[0050] In a preferred embodiment, the optical rotator 7 is a quartz optical rotator, through which a weak coupling effect is achieved.

[0051] In a preferred embodiment, the light source 1 is a superluminescent diode.

[0052] In a preferred embodiment, a collimating lens 2 is further provided between the light source 1 and the pre-selective polarizer 3.

[0053] In a preferred embodiment, a motion control system connected to the post-selected polarizer 8 is also included.

[0054] In a preferred embodiment, the detection chip 5 is a 3D printed chip or a glass-based chip, and the chip 5 is configured to match the refractive index of the prism 4.

[0055] In a preferred embodiment, data post-processing is performed by averaging multiple pixels within the respective inspection areas of the detection channel and the reference channel.

[0056] In a preferred embodiment, in the corresponding The detection is performed in the weak measurement region, where τ is the coupling strength, ω0 is the center frequency of the light source, and ε is the minimum parameter. The phase difference between the P-polarized light and the S-polarized light.

[0057] The following embodiments of the present invention present a self-referenced biodetector based on weak measurement technology, enabling the study of biomolecular interactions. The embodiments of the present invention ensure extremely high sensor sensitivity through weak measurement technology, while simultaneously guaranteeing system stability through differential and pixel averaging methods, thereby developing a high-precision, high-throughput biodetector based on weak measurement. The system achieves a detection accuracy of 2.709 ng / mL for mouse antibodies. Furthermore, the biodetector in the embodiments of the present invention also possesses the advantages of simple structure, high robustness, wide application range, low detection cost, and ease of system and experimental operation, similar to ordinary optical sensors.

[0058] The following are specific examples:

[0059] In this embodiment, the light source 1 is a superradiative diode, whose spectral wavefunction approximately satisfies a normal distribution with a center frequency of ω0 and a standard deviation of Δ. Here, ω represents the light frequency, and ξ represents the emission state of the light source. The polarization direction of the pre-selected polarizer is set to make an angle of [angle missing] with the vertical direction. The polarization state can be written as |H> and |V> represent the horizontal and vertical linear polarization states, respectively. The beam then strikes the inner surface of the prism at an incident angle θ greater than the total internal reflection angle θ0, resulting in a phase difference between the p-ray and s-ray in the reflected light. According to Fresnel's theorem:

[0060]

[0061] where n1 = 1.73 is the refractive index of the ZF6 prism, n2 is the equivalent refractive index on the other side of the reflection interface. Therefore, the polarization state of the reflected light can be written as where i represents the imaginary unit in the complex number field. After that, the light beam passes through the achromatic quarter-wave plate 6 with the fast axis direction set at an angle of with the vertical direction. The interaction between the light beam and it can be written as where <V| and <H| represent the vertical and horizontal components of the light respectively. Therefore, the coupled state of the light beam after pre-selection can be written as:

[0062]

[0063] Here, the optical rotation sheet 7 in this embodiment is a quartz optical rotation sheet, and the weak coupling effect is realized through the quartz optical rotation sheet. | i 〉 represents the pre-selection state of the system. The coupling effect between the polarization state and the frequency domain can be represented by a unitary operator U = e -iτAω , where τ is the coupling strength, which is related to the thickness d of the optical rotation sheet. Here, the thickness of the optical rotation sheet selected in this embodiment is d = 1 mm. ω is the light frequency, and A is the polarization operator A = -i|H〉V| + i|V〉H|.

[0064] The state of the system after weak coupling can be written in the following form:

[0065]

[0066] In this embodiment, high-precision measurement of other physical quantities is achieved by measuring the shift of the central wavelength. The post-selection process is realized through another polarizer. Assume the post-selection state is where ε is a small value, and the front and rear polarization states are nearly orthogonal. According to the weak value amplification theory, the amplification factor of the central wavelength shift is related to the weak value.

[0067]

[0068] | f 〉 represents the post-selection state of the system. When , the post-selection can amplify the shift of the central wavelength. It is known that the highest amplification factor can be obtained in the antilinear region in this embodiment. However, in this embodiment, since many works show that it is still in the weak measurement region at , analyze the relationship between the relative light intensity and the phase of the post-selection at this time:

[0069]

[0070]

[0071] The selected light intensity and phase can be seen. There exists a quadratic function relationship between them, and when measurements are performed within a relatively small range, an approximately linear relationship can be obtained. In this embodiment, the above range is used as the detection interval in the experiment.

[0072] like Figure 2 The diagram shows the optical path of this embodiment. Power supply 1 is a superluminescent diode (SLD), whose emitted light is collimated by collimating lens 2. The collimated light passes through pre-selective polarizer 3 and undergoes total internal reflection on the inner surface of prism 4 of ZF6, causing a phase difference between P-polarization and S-polarization. The collimated light then passes through super-achromatic quarter-wave plate 6, converting the phase difference into optical rotation. Finally, after passing through quartz optical rotator and post-selective polarizer 8, the collimated light enters imaging module 9. Imaging module includes imaging lens 11 and charge-coupled device (CCD), where the imaging lens 11 images the total internal reflection interface onto CCD 12.

[0073] The biosensing part mainly relies on the prism in the optical path for coupling. The existing abundant biochip technology is combined with prism 4 to form a weak measurement interface sensor. In this embodiment, either a 3D-printed multi-channel chip or a microfluidic chip can be used. The schematic diagrams of its combination with prism 4 are shown in Figure 3(a) and Figure 3(b).

[0074] In this embodiment, to ensure full coupling between the optical path and the chip reaction site, two methods are used to bond the chip 5 to the prism 4. The first method, as shown in Figure 3(a), utilizes chip manufacturing technologies such as 3D printing to obtain the chip, and then uses adhesives such as epoxy resin for bonding, thus achieving chip-prism bonding. The second method, as shown in Figure 3(b), utilizes glass-based porous plates, multi-channel systems, microfluidic chips, etc. Here, it is required that the glass at the bottom of the chip has a refractive index similar to that of the prism glass. A refractive index matching liquid is dropped onto the prism surface, and then the chip is pressed and fixed onto the prism to achieve bonding between the glass-based chip and the prism. In this embodiment, ZF6 glass with a refractive index of 1.73 is used.

[0075] The control and display systems in this embodiment mainly consist of two parts: hardware and software.

[0076] In the hardware section, this embodiment uses a stepper motor, encoder, and microcontroller to form a motion control system for the rear-selected polarizer 8. This system and the optical path system are mounted on the same fixed plate, with the stepper motor and polarizer nested after passing through a light-shielding tube. The display system consists of a main tube and a light-shielding tube on the CCD11 camera nested together to form a retractable and adjustable display structure. The device diagram is shown below. Figure 4 , 5 As shown. The software is developed using the C# language and named 'QWMImage'. The software interface is shown below. Figure 6 As shown, it has functions for controlling the camera, controlling the stepper motor, and storing data.

[0077] Biometric Detector Performance Testing

[0078] Because in biological detection, the detection of biomolecular interactions is the smallest biological-scale reaction detection, the detection capability of an instrument can be characterized by its ability to detect biomolecular interactions.

[0079] For the performance testing of this embodiment, the ability to detect mouse IgG (immunoglobulin G) in the target solution using recombinant protein A as a probe was characterized. The specific detection scheme and related results are as follows:

[0080] I. High suppression capability of temperature changes in this embodiment

[0081] In this embodiment, prisms are used as the sensor "chip" unit. This is because glass is inexpensive, and glass surface modification technology is widely used in the fabrication of biomolecular level sensors. However, the refractive index of glass is very sensitive to external temperature. Changes in external temperature can introduce significant errors into the system.

[0082] To confirm the impact of temperature changes on the "chip," this embodiment uses COMSOL software to simulate the heat transfer process. Figure 7(a) shows the simulated heat conduction process of the "chip," and Figure 7(b) shows the simulated heat conduction process in the prism. The effect of temperature changes on the refractive index of the "chip" is simulated by changing the initial temperature of the liquid in the flow channel (below room temperature and above room temperature). Figure 7(c) shows the refractive index change of a point region during the cooling process of the high-temperature liquid, and Figure 7(d) shows the refractive index change of a spot region during the heating process of the low-temperature liquid. The horizontal axis represents time (s), and the vertical axis represents the refractive index. When a liquid different from room temperature enters the system through the flow channel, the refractive index near the inner surface of the prism will rapidly converge to the refractive index at ambient temperature, introducing a refractive index error. Although the temperature of the "chip" at room temperature can converge to room temperature, the error introduced into the sensor during this process cannot be eliminated.

[0083] To verify the temperature robustness of this embodiment, an evaluation experiment was conducted by introducing deionized water at a temperature different from room temperature into the flow channel at room temperature. The temperature changes were amplified in the experiment to more clearly demonstrate the sensor's resistance to external temperature variations. Figures 7(e) and 7(f) show the signal change curves for each channel of the sensor, as well as the signal after differential oscillation of the two channels. The horizontal axis represents time (min), and the vertical axis represents relative light intensity (au). Each channel can be considered as an independent traditional interface-type weak measurement sensor. Figure 7(e) shows the relative light intensity of the two channels and the different results from 39.6℃ to 29.7℃, and Figure 7(f) shows the relative light intensity of the two channels and the different results from 10.1℃ to 22.3℃. During the process of introducing deionized water at a temperature different from room temperature into the flow channel and then allowing it to cool to room temperature within the channel, the signals collected by the receiver showed significant changes, indicating that the sensor chip is highly sensitive to temperature disturbances. However, as the self-reference scheme proposed in this embodiment—differential oscillation of the two channels—the signals collected by the receiver did not show significant changes, which also proves that the sensor system in this embodiment has good robustness to temperature changes.

[0084] II. Instrument Data Post-processing Methods

[0085] This embodiment further improves system stability by averaging the pixels. Figure 8(a) shows a schematic diagram of seven concentric rectangular regions with different areas centered on point 1. The areas of regions 1 to 8 are 1, 16, 64, 256, 1024, 16384, 65536, and 102400, respectively. The variance fluctuation of the sensor was compared within 20 minutes after averaging the pixels in the eight regions (e.g., region 1). As shown in Figure 8(b), region 1 in Figure 8(b) represents point 1 in Figure 8(a). The variance decreases as the number of pixels increases. When the average number of pixels is greater than 1024, the variance fluctuation of the system stabilizes at 11.7 within 20 minutes. This also provides a reference for the minimum range of each detection point when this embodiment is combined with microfluidic technology in the future.

[0086] III. Verification of biomolecular interactions and related parameters

[0087] To further verify the sensing capability of the instrument in this embodiment in biomolecular interactions, the specific binding reaction of IgG and protein A was detected. Taking advantage of the easy surface functionalization of silica, dopamine was first used to modify the surface of the "chip" in the experimental channel (dopamine self-polymerizes to form a thin, surface-adhesive film, which can be used for surface modification with various organic and inorganic materials). Then, protein A and a protein-free blocking solution were used sequentially to coat and block the functionalized "chip" surface. Finally, different concentrations of IgG solutions were sequentially passed through. While the above operations were performed in the experimental channel, phosphate-buffered saline (PBS) solution was continuously passed through the reference channel for self-reference noise reduction. Figure 9(a) shows the results after differential analysis of the relative light intensities of the two channels: I represents dopamine-'chip' functionalization; II represents protein A-'chip' modification; III represents protein-free blocking solution-'chip' blocking; IV represents PBS-baseline; and V represents mouse IgG-molecular interaction. As shown in Figure 9(b), IgG solutions with concentrations of 20 ng, 200 ng, 2000 ng, 20000 ng, and 40000 ng caused changes in the relative light intensity of the acquired signal. Figure 9(c) shows the light intensity of a 20 g / L NaCl solution over 10 minutes, with a standard deviation of 7.72. For Figures 9(a), (b), and (c), the horizontal axis represents time (min), and the vertical axis represents relative light intensity (au). Simultaneously, according to formula c... L =3×σ s Based on (ΔI / Δc) and the system fluctuations mentioned earlier, the detection limit for IgG can be calculated to be 2.709 ng / mL. In the formula, c... L σ s ΔI, Δc, and Δc represent the detection limit, the standard deviation when the detection signal stabilizes, the change in light intensity, and the change in refractive index, respectively. In Figure 9(d), the Langmuir adsorption model is used to fit the experimental results, where the horizontal axis represents mouse IgG concentration (ng / mL), the vertical axis represents relative light intensity (au), and ΔI = ΔC / Δt. max K al ρ l (1+K a ρ), K a ρ is the absorption constant, ρ is the concentration of IgG, and ΔI is the change in relative light intensity caused by a specific binding reaction. max =5757.u. Figure 9(e) shows the relative light intensity caused by different target samples, and the absorption constant was calculated to be 9.652×10-5mL / (ng·au).

[0088] The above description, in conjunction with specific preferred embodiments, provides a further detailed explanation of the present invention. It should not be construed that the specific implementation of the present invention is limited to these descriptions. For those skilled in the art, several equivalent substitutions or obvious modifications can be made without departing from the concept of the present invention, and all such modifications, achieving the same performance or purpose, should be considered within the scope of protection of the present invention.

Claims

1. A biological detector based on a weak measurement method, characterized in that, The system includes a light source, a pre-selective polarizer, a prism, a detection chip, a quarter-wave plate, a rotator, a post-selective polarizer, and an imaging module. Light emitted from the light source passes through the pre-selective polarizer and enters the prism. Total internal reflection occurs at the weak measurement interface formed by the prism and the detection chip, causing a phase difference between the P-wave and S-wave polarizations. The reflected light passes through the quarter-wave plate, the rotator, and the post-selective polarizer before entering the imaging module to achieve image formation at the total internal reflection interface. The detection chip has a detection channel and a reference channel corresponding to the weak measurement interface. The detection channel is used to introduce the biological sample to be tested, so as to achieve self-reference differential noise reduction by subtracting the average light intensity detected by the detection channel and the reference channel; multiple pixels are selected in the detection areas of the detection channel and the reference channel for averaging to further reduce noise; the quarter-wave plate is an achromatic quarter-wave plate with the fast axis direction set at an angle of π / 4 with the vertical direction; the optical rotator uses a quartz optical rotator to achieve weak coupling; the light source is a superluminescent diode; the detection chip is a 3D printed chip or a glass-based chip and its refractive index is matched with that of the prism; in the corresponding The detection is performed in the weak measurement region, where τ is the coupling strength, ω0 is the center frequency of the light source, and ε is the minimum parameter. The phase difference between the P-polarized light and the S-polarized light. To achieve high-precision detection of biomolecular interactions.

2. The biological detector as described in claim 1, characterized in that, It also includes a collimating lens disposed between the light source and the pre-selective polarizer.

3. The biological detector as described in claim 1, characterized in that, It also includes a motion control system connected to the post-selected polarizer.

Images