CN115931784A - Biological detector based on weak measurement method - Google Patents
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Abstract
The invention discloses a biological detector based on a weak measurement method, which comprises a light source, a front selection polarizing film, a prism, a detection chip, a quarter wave plate, an optical rotation sheet, a rear selection polarizing film and an imaging module, wherein light emitted by the light source enters the prism after passing through the front selection polarizing film, total internal reflection is realized at a weak measurement interface formed by the prism and the detection chip in an attaching manner, a phase difference exists between P light polarization and S light polarization, and reflected light enters the imaging module after passing through the quarter wave plate, the optical rotation sheet and the rear selection polarizing film to realize image imaging of the total internal reflection interface; the detection chip is provided with a detection channel and a reference channel at the position corresponding to the weak measurement interface, so that self-reference differential noise reduction is realized by using the difference of relative light intensity of the detection channel and the reference channel. The invention improves the anti-interference performance and stability of the high-sensitivity biosensor and improves the performance of the weak measurement biosensor.
Description
Technical Field
The invention relates to the technical field of biological detection instruments, in particular to a biological detector based on a weak measurement method.
Background
Life is a permanent subject of scientific research, and a series of great breakthroughs obtained by scientific research in the new century is actually improving human life. Simultaneously, the biosensor field formed by the fusion and intersection of life science and multidisciplinary also emits striking vitality. Biosensors generally combine a bio-sensitive material with a specific enzyme, antibody, antigen, protein, nucleic acid, small molecule, etc. bioactive analyte to be detected, and convert the combination process into distinguishable changes of characteristics such as electricity, optics, heat, etc. through a physicochemical transducer, and then convert biochemical signals into quantifiable electric signals through an amplification means. The optical signal has the advantages of high sensitivity, external interference resistance, good stability, low noise and the like in comparison with other physical signals. Therefore, the optical biosensor shows good performance in the aspect of biological detection systems, and promotes scientific research and industrial fields such as biomedicine, food health, substance characterization, medicine research and development, medical instruments, environmental protection and the like to obtain great breakthrough.
The conventional optical biosensor technologies mainly include a Bio-membrane interference technology (BLI), a surface plasmon resonance technology (SPR), and a Bio-fluorescence technology.
The biomembrane interference technology adopts a probe type biosensor to directly detect a sample, and utilizes the displacement of a reflection spectrum to represent the molecular dynamics process combined on the surface of the sensor in real time. However, the technology has the defects that only single-point detection of the spectrometer can be carried out, and high-flux parallel monitoring cannot be realized; the single-point detection mode makes environmental factors such as temperature, pressure and the like and errors brought by the system difficult to eliminate, and greatly limits the application range and the scene of the system.
The surface plasma resonance technology is to plate a metal film with the thickness of about dozens of nanometers on a carrying plane, usually takes a gold film and a silver film as main materials, and the technology has higher sensitivity and high-flux technology realization capability. However, the technical disadvantage is that the coating requirement of the metal film is high, precise nano-scale coating precision is required, reproducibility is restricted, and the difference of the thickness of the metal film between repeated measurements affects the detection result; the physical principle has higher requirements on the incident angle of light, so that the hardware precision requirement of the whole detection system is harsher, and the robustness of the environment is poorer; because the metal film is a core carrier of the technology, during detection, the pH value of a biological system to be detected can corrode the metal film, so that result errors and poor recycling capability can be caused; gold and silver are the most commonly used plating films, the silver film is easily oxidized, and the gold film easily catalyzes certain reactions, so the interpretation of the results is easily questioned.
The bio-fluorescence technology is a commonly used technology, but the influence of the labeling of fluorescent molecules on a system to be detected is always a controversial problem, so that the labeling problem is a problem which cannot be bypassed by biological detection. The advantages of label-free labeling are also apparent, and although fluorescence labeling techniques bring high sensitivity, they also bring a series of newly derived problems such as fluorescence quenching, fluorescence background noise, etc.
The weak measurement technology has shown great advantages in the field of high-precision measurement since the proposal in 1988. Particularly, the frequency domain weak measurement system and the realization of the weak measurement sensing system with general value enable the weak measurement sensor to show excellent wind collection in the field of biological detection. The application of weak measurement technology greatly improves the sensitivity of the sensing system. The biosensor based on the weak measurement technology is easily influenced by external factors such as temperature, vibration and the like, the anti-interference performance of the sensor is improved, and the improvement of the performance of the weak measurement biosensor is particularly important.
Disclosure of Invention
The invention aims to provide a biological detector based on a weak measurement method, which improves the anti-interference performance and stability of the biological detector.
The invention provides a biological detector based on a weak measurement method, which comprises a light source (1), a front selective polarizer (3), a prism (4), a detection chip (5), a quarter-wave plate (6), an optical rotation plate (7), a rear selective polarizer (8) and an imaging module (9), wherein light emitted by the light source (1) enters the prism (4) after passing through the front selective polarizer (2), total internal reflection is realized at a weak measurement interface formed by the prism (4) and the detection chip (5) in a fit manner, so that a phase difference exists between P light polarization and S light polarization, and reflected light enters the imaging module (9) to realize image imaging of the total internal reflection interface after passing through the quarter-wave plate (6), the optical rotation plate (7) and the rear selective polarizer (8); the detection chip (5) is provided with a detection channel and a reference channel at the position corresponding to the weak measurement interface, so that self-reference differential noise reduction is realized by using the difference of relative light intensity of the detection channel and the reference channel.
In some embodiments of the invention, the difference in relative light intensities is a difference between the mean of the light intensities detected by the detection channel and the reference channel.
In some embodiments of the invention, the direction of the fast axis of the quarter-wave plate (6) is set at an angle to the vertical
The achromatic quarter wave plate of (1).
In some embodiments of the present invention, the optical rotation plate (7) is a quartz optical rotation plate by which a weak coupling effect is achieved.
In some embodiments of the invention, the light source (1) is a superluminescent light emitting diode.
In some embodiments of the invention, further comprising a collimating lens (2) disposed between the light source (1) and the front selective polarizer (3).
In some embodiments of the invention, a motion control system is further included in communication with the rear selective polarizer (3).
In some embodiments of the invention, the detection chip (5) is a 3D printed chip or a glass based chip, the chip being arranged to match the refractive index of the prism.
In some embodiments of the present invention, data post-processing is performed by selecting a plurality of pixel points for averaging within respective inspection regions of the detection channel and the reference channel.
In some embodiments of the invention, in
Weak measurement ofDetecting the region, wherein tau is the coupling strength, and omega 0 Is the central frequency of the light source, ε is a small-value parameter ∈ is based on>
Is the phase difference between the P-light polarization and the S-light,
the invention has the following beneficial effects:
the invention provides a self-reference biological detector based on a weak measurement technology, wherein a detection channel and a reference channel are arranged on a detection chip at a position corresponding to a weak measurement interface, and self-reference differential noise reduction can be realized by utilizing the relative light intensity difference of the detection channel and the reference channel, so that the biological detector has good robustness to temperature change. In a preferred embodiment, the present invention further improves the stability of the system by averaging the pixel points. The biological detector has the advantages of simple structure, high robustness, wide application range, low detection cost, simple and convenient system and experiment operation and the like.
Drawings
FIG. 1 (a) is a schematic diagram of a two-channel chip in an embodiment of the invention;
FIG. 1 (b) is a cross-sectional view of a two-channel chip in an embodiment of the present invention;
FIG. 2 is a schematic diagram of the optical path in an embodiment of the present invention;
FIG. 3 (a) is a schematic diagram of a prism in combination with a detection chip according to an embodiment of the present invention;
FIG. 3 (b) is a schematic diagram of a prism combined with another detection chip in an embodiment of the present invention;
FIG. 4 is a diagram of a motion control system apparatus in an embodiment of the present invention;
FIG. 5 is a diagram of a display system in an embodiment of the present invention;
FIG. 6 is a screenshot of the operation interface of the software "QWMarge" in the embodiment of the present invention;
FIG. 7 (a) is a schematic diagram of a heat conduction process simulating a "chip" in an embodiment of the present invention;
FIG. 7 (b) is a schematic diagram of the heat conduction process of the pseudo prism in the embodiment of the present invention;
FIG. 7 (c) is a line graph showing the refractive index change in the dotted region during cooling of the high-temperature liquid in the embodiment of the present invention;
FIG. 7 (d) is a graph showing the refractive index change of a light spot region during the temperature rise of the cryogenic liquid in the example of the present invention;
FIG. 7 (e) is a graph comparing the relative intensity change of two channels from 39.6 deg.C to 29.7 deg.C in the example of the present invention;
FIG. 7 (f) is a graph comparing the relative intensity change of two channels from 10.1 deg.C to 22.3 deg.C in the example of the present invention;
FIG. 8 (a) is a schematic illustration of 8 concentric rectangular regions of different areas in an embodiment of the present invention;
FIG. 8 (b) is a histogram of variance of pixel values for 8 different regions in an embodiment of the present invention;
FIG. 9 (a) is a graph showing the difference between the relative light intensities of two channels according to the embodiment of the present invention;
FIG. 9 (b) is a graph showing the variation of light intensity versus IgG concentration in examples of the present invention;
FIG. 9 (c) is a line graph showing the light intensity of a NaCl solution in the example of the present invention;
FIG. 9 (d) is a plot of the Langmuir adsorption model fitted to the experimental results in an example of the invention;
FIG. 9 (e) is a bar graph of relative light intensity contrast caused by different target samples in an example of the invention.
The reference numbers are as follows:
the device comprises a power supply 1, a collimating lens 2, a front selective polarizer 3, a prism 4, a detection chip 5, a quarter-wave plate 6, a rotary plate 7, a rear selective polarizer 8, an imaging module 9, an imaging lens 10 and a charge coupled device 11.
Detailed Description
The invention will be further described with reference to the accompanying drawings and preferred embodiments. It should be noted that the embodiments and features of the embodiments in the present application may be combined with each other without conflict.
It should be noted that the terms of orientation such as left, right, up, down, top and bottom in the present embodiment are only relative concepts to each other or are referred to the normal use state of the product, and should not be considered as limiting.
Based on the advantages of the weak measurement technology and the sensing characteristics of the evanescent field, the interface type biosensor based on the weak measurement can realize high-precision biological information acquisition on a common glass interface. The embodiment of the invention provides a self-reference biological detector based on a weak measurement method, a reference channel is established in the same interface, the problem of external interference is effectively solved, and an expensive additional detection element is not required to be added. According to the embodiment of the invention, when the experimental channel and the reference channel are subjected to differential noise reduction, a plurality of pixel points are selected in respective inspection areas in the two channels to be averaged, so that the stability of the system is further improved. The differential noise reduction method is used for eliminating additive noise in the instrument by calculating the mean value of the light intensity of two channels and then subtracting the mean value. The experimental channel and the reference channel are the channels designed in the detection chip, and because the types of the chip are various, the embodiment of the invention designs the simplest two-channel chip to describe, i.e. one of the two channels is used as the experimental channel, and the other channel is used as the reference channel, as shown in fig. 1a and fig. 1 b.
The embodiment of the invention provides a biological detector based on a weak measurement method, which comprises a light source 1, a front selection polaroid 3, a prism 4, a detection chip 5, a quarter-wave plate 6, an optical rotation plate 7, a rear selection polaroid 8 and an imaging module 9, wherein light emitted by the light source 1 enters the prism 4 after passing through the front selection polaroid 3, total internal reflection is realized at a weak measurement interface formed by the prism 4 and the detection chip 5 in an attaching manner, so that a phase difference exists between P optical polarization and S optical polarization, and reflected light enters the imaging module 9 after passing through the quarter-wave plate 6, the optical rotation plate 7 and the rear selection polaroid 8 to realize image imaging of the total internal reflection interface; the detection chip 5 is provided with a detection channel and a reference channel at the position corresponding to the weak measurement interface, so that self-reference differential noise reduction is realized by using the difference of relative light intensity of the detection channel and the reference channel.
In a preferred embodiment, the difference in relative light intensities is a difference between the mean values of the light intensities detected by the detection channel and the reference channel.
In the preferred embodiment, the quarter-wave plate 6 is set to have a fast axis direction at an angle with the vertical direction
The achromatic quarter waveplate of (1).
In a preferred embodiment, the optical rotation plate 7 is a quartz optical rotation plate, and a weak coupling effect is realized through the quartz optical rotation plate.
In a preferred embodiment, the light source 1 is a superluminescent light emitting diode.
In a preferred embodiment, a collimating lens 2 is further included, arranged between the light source 1 and the front selective polarizer 3.
In a preferred embodiment, a motion control system is also included in communication with the rear selective polarizer 8.
In a preferred embodiment, the detection chip 5 is a 3D printed chip or a glass based chip, the chip 5 being arranged to match the refractive index of the prism 4.
In a preferred embodiment, the data post-processing is performed by selecting a plurality of pixel points for averaging within respective inspection regions of the detection channel and the reference channel.
In the preferred embodiment, in the corresponding to
Is detected, where τ is the coupling strength, ω is 0 Is the center of the light sourceFrequency ε a small value parameter->
Is the phase difference between the P light polarization and the S light>
The self-reference biological detector based on the weak measurement technology provided by the following embodiments of the invention realizes the research of the interaction of biological molecules. According to the embodiment of the invention, on the premise that the sensor has extremely high sensitivity through a weak measurement technology, the stability of the system is ensured through a difference and pixel averaging method, so that a high-precision high-throughput biological detector based on weak measurement is developed, and the detection precision of the system on the mouse antibody reaches 2.709ng/mL. In addition, the biological detector in the embodiment of the invention has the characteristics of simple structure, high robustness, wide application range, low detection cost, simple and convenient system and experiment operation and the like of a common optical sensor.
The following are specific examples:
the light source 1 used in this embodiment is a superluminescent diode, and the spectral wave function approximately satisfies the central frequency of ω 0 Normal distribution with standard deviation of delta
Where ω is the light frequency and ξ represents the light source emission state. Setting the polarization direction of the pre-selected polarizer to be at an angle ^ with the vertical direction>
The polarization state can be written as->
|H>And | V>Respectively representing a horizontal and a vertical linear polarization state. The light beam is then reflected at an angle greater than the total reflection angle theta 0 Is incident on the inner surface of the prism, a certain phase difference is added between p light and s light in the reflected light>
From the fresnel theorem, it can be known that:
wherein
n 1 =1.73 refractive index of ZF6 prism, n 2 Is the equivalent refractive index on the other side of the reflective interface, so that the polarization state of the reflected light can be written as ^ er>
Where i denotes the imaginary symbol in the complex domain. Then, the direction of the light beam passing through the fast axis is set to be at an angle with the vertical direction>
The achromatic quarter wave plate 6 with which the interaction of the light beam can be written as->
Wherein<V|、<H | represents a vertical component and a horizontal component of light, respectively. The coupling state of the front-selected rear beam can thus be written as:
here, the optical rotation sheet 7 in the present embodiment is a quartz optical rotation sheet, and the weak coupling effect is realized by the quartz optical rotation sheet i Represented by (a) is a pre-selection state of the system. The coupling effect between the polarization state and the frequency domain can be represented by a unitary operator, wherein U = e -iτAω Where τ is the coupling strength, which is related to the thickness d of the optically active plate, where the thickness of the optically active plate suspended in this embodiment is d =1mm. Omega is the optical frequency, A is the polarization operator A = -i | H > V | and i | V > H |.
The state of the system after weak coupling can be written as follows:
the present embodiment realizes high-precision measurement of other physical quantities by measuring the shift of the center wavelength. The post-selection process is effected by means of a further polarizer, assuming the post-selection state is
Where epsilon is a small value and the front and back polarization states are nearly orthogonal. From the theory of weak amplification, it is known that the amplification of the center wavelength shift is related to the weak value.
| f Represents the post-selection state of the system when
The latter selection can amplify the shift of the center wavelength. It is known that the highest magnification can be obtained in this embodiment in the region of the inverse linearity. However, in this embodiment, much of the work is shown at->
Is still in the weak measuring region, the relative light intensity and phase selected later at this time are analyzed>
The relationship between:
the light intensity and phase position selected after passing can be seen
There is a quadratic relationship between them, and an approximately linear relationship can be obtained when measurements are made over a relatively small range. The present embodiment experimentally regards the above range as the detection interval of the present embodiment.
Fig. 2 is a schematic diagram of the optical path of the present embodiment. The power supply 1 is a Super Luminescent Diodes (SLD), which emits light that is collimated by a collimating lens 2. The collimated light passes through the front selective polarizer 3 and achieves total internal reflection at the inner surface of the prism 4 of ZF6, causing a phase difference between the P-and S-light polarizations. The collimated light then passes through a superachromatic 1/4 wave plate 6, so that the phase difference is converted into optical rotation. Finally, the collimated light enters an imaging module 9 after passing through a quartz polariscope and a rear selective polarizer 8, the imaging module includes an imaging lens 11 and a Charge Coupled Device (CCD) 12, and an image of the total internal reflection interface is imaged onto the CCD12 by the imaging lens 11.
The biosensing part is mainly realized by a coupling part prism in an optical path. The existing abundant biochip technology is combined with the prism 4 to form a weak measurement interface sensor, in this embodiment, a 3D printed multi-channel chip may be used, or a microfluidic chip may be used, and schematic diagrams of the combination with the prism 4 are shown in fig. 3 (a) and 3 (b).
In order to make the optical path sufficiently coupled with the chip reaction site, the chip 5 and the prism 4 are attached in two ways. First, as shown in fig. 3 (a), a chip is obtained by a chip manufacturing technique such as a 3d printing technique, and is bonded by a glue such as epoxy resin, so that the chip and the prism can be bonded; secondly, as shown in fig. 3 (b), a glass-based porous plate, a multi-channel, a microfluidic chip, etc. are used, where there is a requirement that the refractive index coefficient of the glass at the bottom of the chip needs to be similar to that of the prism glass, a refractive index matching fluid is dripped on the surface of the prism, and then the chip is tightly pressed and fixed on the prism, so that the glass-based chip and the prism can be attached to each other, in this embodiment, ZF6 glass with a refractive index of 1.73 is used.
The control system and the display system of the embodiment are mainly composed of hardware and software.
In the hardware part, the embodiment uses a stepping motor, an encoder and a single chip microcomputer to form a set of motion control system for the rear selection polaroid 8. The system and the optical path system are arranged on the same fixed plate, and the nesting of the stepping motor and the polaroid is realized after passing through the light-avoiding cylinder. The display system is a telescopic display structure formed by nesting a main cylinder and a shading cylinder on a CCD11 camera. The device is shown in figures 4 and 5. The software part is developed based on C # language and named 'QWMimege', and the software interface has the functions of controlling the camera, controlling the stepping motor and storing data as shown in FIG. 6.
Performance testing of biological testing instruments
Since in biological assays, the biomolecular interaction assay is a minimum biological scale reaction assay, measuring the detection capability of an instrument can be characterized by the detection capability of the biomolecular interaction.
For the performance test of this example, the ability of detecting mouse IgG (Immunoglobulin G) in a target solution using recombinant protein a as a probe was used for characterization. The specific detection scheme and the related results are as follows:
1. high temperature change suppression capability of the embodiment
The prism-based element used in this example is used as a sensor "chip" unit, on one hand because of the low cost of glass, and on the other hand, the glass surface modification process has been widely used in the preparation of biomolecular level sensors. However, the refractive index of glass is sensitive to the outside temperature. Ambient temperature variations can introduce large errors into the system.
In order to confirm the disturbance of the temperature change to the "chip", the heat transfer process is simulated by using comsol software in the present embodiment, fig. 7 (a) shows the heat transfer process simulating the "chip", and fig. 7 (b) shows the heat transfer process simulating the prism. The effect of temperature changes on the "chip" refractive index was simulated by varying the initial temperature of the liquid in the flow channel (below room temperature and above room temperature). As shown in fig. 7 (c), the refractive index change of the point region in the cooling process of the high-temperature liquid is shown, and as shown in fig. 7 (d), the refractive index change of the spot region in the heating process of the low-temperature liquid is shown, wherein the abscissa is time(s), and the ordinate is the refractive index, when the liquid different from room temperature enters the system through the flow channel, the refractive index near the inner surface of the prism can be rapidly converged to the refractive index at the ambient temperature, and a refractive index error is introduced; although the temperature of the "chip" at room temperature can converge to the room temperature, the error introduced in the sensor in the process cannot be eliminated.
To verify the robustness of this embodiment to temperature, an evaluation experiment was performed by passing deionized water at a temperature different from room temperature through the flow channel under room temperature conditions. The temperature variation was amplified in the experiment to show more clearly the immunity of the sensor to external temperature variations. Fig. 7 (e), 7 (f) show the signal variation curve after each channel of the sensor and the two-channel difference, wherein the abscissa is time (min) and the ordinate is relative light intensity (a.u.). Each channel can be considered herein as an independent conventional interfacial weak measurement sensor. As shown in fig. 7 (e) showing the relative light intensity of the two channels and the different results from 39.6 ℃ to 29.7 ℃, and fig. 7 (f) showing the relative light intensity of the two channels and the different results from 10.1 ℃ to 22.3 ℃, in the process that deionized water different from room temperature is introduced into the flow channel and becomes room temperature in the flow channel, the signals collected by the receiving end have obvious changes, namely the sensor chip is very sensitive to temperature disturbance; however, as the self-reference scheme proposed in this embodiment, after the difference between the two channels, the signal acquired by the receiving end does not change significantly, which also proves that the sensor system in this embodiment has good robustness to temperature change.
2. Instrument data post-processing method
The present embodiment further improves the stability of the system by averaging the pixel points. Fig. 8 (a) is a schematic diagram showing 7 concentric rectangular regions of different areas centered on point 1, where the area of the regions from 1 to 8 are 1, 16, 64, 256, 1024, 16384, 65536, 102400 respectively. The fluctuation variance of the sensor within 20 minutes after the pixel points of 8 regions (such as region 1) are averaged is compared in sequence. As shown in fig. 8 (b), the region 1 in fig. 8 (b) represents the point 1 in fig. 8 (a), the variance decreases with the increase of the number of pixel points, and when the average pixel point is more than 1024, the variance fluctuation of the system is stabilized at 11.7 within 20 minutes. This also provides a reference to the minimum range of each detection point for this embodiment when the system is later combined with microfluidic technology.
3. Biomolecular interaction validation and related parameters
To further verify the sensing ability of the instrument of this example in biomolecular interactions, specific binding reaction detection of IgG and protein a was performed. By utilizing the characteristic that silicon dioxide is easily functionalized on the surface, the surface of a chip is modified in an experimental channel by using dopamine (dopamine is self-polymerized to form a thin film with a surface adhered to the surface, and the thin film can be used for surface modification with various organic and inorganic materials), then the functionalized surface of the chip is coated and closed by sequentially using protein A and protein-free sealing liquid, and finally IgG solutions with different concentrations are sequentially introduced. While the experimental channel was operated as described above, the reference channel was continuously fed with Phosphate Buffered Saline (PBS) for self-reference noise reduction. FIG. 9 (a) shows the difference between the relative intensities of the two channels, I being dopamine-chip functionalization; II is protein A- 'chip' modification; III is a protein-free confining liquid- 'chip' confining; IV is PBS-baseline; v is the mouse IgG-molecule effect. As shown in FIG. 9 (b), igG solutions at concentrations of 20ng, 200ng, 2000ng, 20000ng, and 40000ng caused relative intensity changes in the collected signals, respectively. FIG. 9 (c) is the light intensity of 20g/L NaCl solution over 10 minutes with a standard deviation of 7.72. For fig. 9 (a) (b) (c), the abscissa is time (min) and the ordinate is relative light intensity (a.u.). Simultaneously according to formula c L ＝3×σ s /(. DELTA.I/. DELTA.c) and the foregoing system fluctuations allow the system to calculate IgG detection to a limit of 2.709ng/mL, where c is L 、σ s Δ I, Δ c are the detection limit, the standard deviation when the detection signal is stable, the amount of change in light intensity, and the amount of change in refractive index, respectively. In fig. 9 (d), langmuir adsorption model was applied to fit the experimental results with mouse IgG concentration (ng/mL) on the abscissa and relative light intensity (a.u.) on the ordinate, Δ I = Δ I max K al ρ l (1+K a ρ)，K a Is the absorption constant, ρ is the concentration of IgG, Δ I is the relative change in light intensity due to a particular binding reaction, Δ I max =5757.u. As shown in FIG. 9 (e), the absorption constant was calculated to be 9.652X 10-5 mL/(ng.a.u.) for the relative light intensities caused by the different target samples.
The foregoing is a more detailed description of the invention in connection with specific preferred embodiments and it is not intended that the invention be limited to these specific details. For those skilled in the art to which the invention pertains, several equivalent substitutions or obvious modifications can be made without departing from the spirit of the invention, and all the properties or uses are considered to be within the scope of the invention.
Claims (10)
1. A biological detector based on a weak measurement method is characterized by comprising a light source, a front selection polarizing film, a prism, a detection chip, a quarter-wave plate, an optical rotation plate, a rear selection polarizing film and an imaging module, wherein light emitted by the light source enters the prism after passing through the front selection polarizing film, total internal reflection is realized at a weak measurement interface formed by the prism and the detection chip in an attached mode, a phase difference exists between P light polarization and S light polarization, and reflected light enters the imaging module to realize image imaging of the total internal reflection interface after passing through the quarter-wave plate, the optical rotation plate and the rear selection polarizing film; the detection chip is provided with a detection channel and a reference channel at the position corresponding to the weak measurement interface, so that self-reference differential noise reduction is realized by using the difference of relative light intensity of the detection channel and the reference channel.
2. The biological detector of claim 1, wherein the difference in relative light intensities is a difference between the mean of the light intensities detected in the detection channel and the reference channel.
3. The biological detector according to claim 1, wherein the quarter wave plate is oriented with its fast axis set at an angle to the vertical
The achromatic quarter wave plate of (1).
4. The biodetector of claim 1, wherein the polarimeter comprises a quartz polarimeter, and the quartz polarimeter is used to achieve a weak coupling effect.
5. The biodetector of claim 1, wherein the light source is a superluminescent light emitting diode.
6. The biological detector of claim 1, further comprising a collimating lens disposed between the light source and the front selective polarizer.
7. The biological detector of claim 1, further comprising a motion control system coupled to the rear selective polarizer.
8. The biodetector of claim 1, wherein the detection chip is a 3D printed chip or a glass-based chip, the chip being configured to match the refractive index of the prism.
9. The biodetector of claim 1, wherein the data post-processing is performed by selecting a plurality of pixel points within the respective examination regions of the detection channel and the reference channel for averaging.
10. As claimed in claimThe biological detector according to any one of claims 1 to 9, wherein the biological detector is characterized by being used in a biological detection apparatus for detecting a biological sample
Is detected, where τ is the coupling strength, ω is 0 Is the central frequency of the light source, epsilon is a small value parameter>
Is the phase difference between the P light polarization and the S light, ∈, <' > is present>
/>
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