KR101461235B1 - Probe for acquiring polarization-sensitive optical coherence image and a method for managing a system for acquiring polarization-sensitive optical coherence image - Google Patents

Abstract

The present invention relates to a probe for acquiring a polarization-sensitive optical interference image for measuring the distribution of cells in the body or in vitro for various biopsy diagnosis, and more particularly, to a probe for obtaining a polarization-sensitive optical interference image by measuring scattering of light in a cell distribution near the epidermis, Optical interference image acquisition probe capable of performing biopsy by measuring the birefringence of collagen fibers distributed in the vicinity of the dermis or cartilage of the skin.

A probe for polarized light sensitivity-optical interference image acquisition for biopsy of the present invention comprises a probe and a body, wherein the probe detects an optical signal including information of a living tissue, A probe for obtaining a polarization-sensitivity optical interference image, the probe for obtaining a polarization sensitivity-optical interference image comprising the probe in a polarization sensitivity-optical interference image system comprising a computer system for measuring and analyzing a living tissue from an optical signal, A polarizing sensitivity capable of detecting light scattering and birefringence by the biotissue using a Fourier domain technique capable of performing vertical axis scanning without movement of a mirror by performing horizontal axis scanning using the optical scanner of the living tissue - optical interference coherence And an optical system constituting a tomograph (PS-OCT).

Biological Tissue, Polarization Sensitivity - Optical Interference, Imaging System, Probes, Fourier Domain Technique

Description

[0001] The present invention relates to a polarization-sensitive optical coherence image and, more particularly, to a polarization-sensitive optical coherence image and a polarization-sensitive optical coherence image. a method for managing a system for acquiring polarization-sensitive optical coherence image}

The present invention relates to a probe for acquiring a polarization-sensitive optical interference image for measuring the distribution of cells in the body or in vitro for various biopsy diagnosis, and more particularly, to a probe for obtaining a polarization-sensitive optical interference image by measuring scattering of light in a cell distribution near the epidermis, Optical interference image acquisition probe capable of performing biopsy by measuring the birefringence of collagen fibers distributed in the vicinity of the dermis or cartilage of the skin.

Optical Coherence Tomography (OCT) is a device that captures living tissue or cells in high resolution in real time using light harmless to the human body. It is used for non-contact and non-invasive observation of the inside of a living body. And it is possible to distinguish the difference between the soft tissues, so that a more precise image can be obtained. The OCT system is based on a broad-band light source and a Michelson interferometer, and is currently being applied to a wide range of applications including ophthalmology, dermatology, digestive medicine, and dentistry due to the development of light sources and technologies.

However, the OCT technique is disadvantageous in that a precise image can not be obtained at a deep part (for example, a deep part of 500 mu m or more from the epidermis) in the precise measurement. This is because the scattered light amount can not be sufficiently collected at a deep portion.

Polarization Sensitive-Optical Coherence Tomography (hereinafter referred to as PS-OCT) is a technique that acquires depth-dependent changes in backscattered light caused by scattering of living tissue, The PS-OCT is based on the principle of Michelson interferometer. The PS-OCT is based on the Michelson interferometer principle. Compared with OCT, which implements only a structural image, PS-OCT is a functional information that is difficult to obtain when imaging by a general optical method, for example, .

The PS-OCT system consists of a light source, an optical system unit, a detection system unit, and a computer system as one large system. Since the optical fiber is sensitive to polarization, it can transmit light from the light source to the optical system part of the PS- However, it is difficult to use optical fiber in the optical system parts, that is, in the reference arm or the sample arm.

Therefore, it is difficult to apply the OCT system and the probe, which are used in the general OCT system, to the PS-OCT system independently of each other.

In particular, a biopsy diagnostic imaging system is required to use a computer system and a probe independently of each other.

The present invention provides a polarization sensitivity-optical interference image acquisition probe for biopsy diagnosis in which a light source and a computer system are integrated into one system, and an optical system constituting a PS-OCT and a detection system are integrated in a probe. In the present invention, since the light from the light source is transmitted only to the probe using the optical fiber, the PS-OCT system in the actual probe is not affected by the polarization.

The present invention also provides a polarization-sensitivity optical interference image acquisition probe for biopsy diagnosis in which a computer system is constituted of a single system and an optical system constituting a light source and a PS-OCT and a detection system are integrated in a probe. In the present invention, since the light source is generated in the probe, the influence of the noise is small, and more accurate results can be obtained.

The present invention also relates to a polarization-sensitive optical interference imaging system for biopsy diagnosis, which can detect a change in polarization component depending on a biological tissue characteristic using a PS-OCT using a Fourier domain technique for fast image acquisition of a living tissue to provide.

The present invention can acquire images non-invasively without cutting or dissecting cell tissue, and can acquire not only two-dimensional cross-sectional structural images of living tissue but also polarized component images with high resolution in real time, Since real-time acquisition is possible, 3D image can be implemented in a short time.

Since the present invention integrates a system into a hand-held probe, the system can be miniaturized and can be manufactured as a portable device, and a plug-and-play function against a main body including a computer system , And it is possible to objectively quantify the extent of the lesion stage and the therapeutic effect on the biotissue.

SUMMARY OF THE INVENTION The present invention provides a polarization-sensitivity optical interference image acquisition probe for measuring a living tissue in which a light source unit and a computer system are integrated into one system and an optical system constituting a PS-OCT and a detection system are integrated in a probe I have to.

According to another aspect of the present invention, there is provided a polarization-sensitive optical interferometer (PLL) capable of detecting a change in a polarization component depending on a biological tissue characteristic, using a PS-OCT using a Fourier domain technique, And to provide an acquisition probe.

Another object of the present invention is to acquire images in a non-invasive manner, acquire real-time two-dimensional high-resolution images of the longitudinal and transverse axes of living tissue through polarized component images, And to provide a polarization-sensitivity-optical interference image acquisition probe capable of realizing a three-dimensional image according to acquisition and quantitatively objectizing a lesion stage and a therapeutic effect on a living tissue.

In order to achieve the above object, the present invention provides a polarization-sensitivity optical interference image acquisition probe for biopsy of the present invention, which comprises a probe and a body, wherein the probe comprises a light source And a main body for detecting a signal, and the main body comprises a computer system for measuring and analyzing a living tissue from an optical signal received from the probe, the probe for obtaining a polarization sensitivity-optical interference image constituting the probe in a polarization sensitivity-optical interference , The probe for acquiring the polarization sensitivity-optical interference image is a probe for acquiring the polarization sensitivity-light interference image by using a light scanner of the living tissue and performing a horizontal axis scanning and a vertical axis scanning without a mirror movement, And an optical system that forms a polarization-sensitive coherent tomography (PS-OCT) capable of detecting birefringence. .

The probe includes a linear polarizer (LP) for producing one light from the light source unit only in the horizontally polarized light; An optical splitter (BS) for causing the purely horizontally polarized light output from the linear polarizer to travel through a sample arm as a reference arm in one path and as a sample arm in the other path; A first quarter wave plate (QWP) for transmitting light that has flowed from the linear polarizer to the reference arm to a reference mirror and returning light reflected from the reference mirror to the optical splitter; To allow the light from the linear polarizer to pass through the mirror installed in the galvanometer and to reflect or backscatter from the sample tissue to return the light reflected by the mirror mounted on the galvanometer back to the optical splitter A second quarter wave plate (QWP); A second quarter wave plate (QWP) to the galvanometer to focus the light reflected by the mirror mounted on the galvanometer and to enter the sample tissue; An objective lens for transmitting the emitted light to a mirror installed in the galvanometer; A PBS (PBS) in which the interference signal generated from the light collected by the optical splitter is divided into a horizontal component light and a vertical component light; And optical detectors for converting the horizontal component light emitted to the polarized light splitter and the vertical component light into respective electrical signals.

The apparatus further includes a uniaxial galvanometer for longitudinal axis scanning and transverse axis scanning of the biotissue sample along the longitudinal direction of the biopsy sample.

The light source of the light source unit is composed of a high-speed wavelength-converted low coherence light source.

The first quarter wave plate is a 1/4 wave plate of 22.5 DEG and the second quarter wave plate is a quarter wave plate of 45 DEG.

The photodetectors comprise photodiodes.

The computer system measures the light scattering and birefringence of the biotissue to obtain a quantitative value of the tumor tissue distribution and the lesion stage.

The computer system measures quantitative values of tumor tissue distribution and lesion stages by measuring light scattering of the biomedical surface layer region.

The computer system measures quantitatively the tumor tissue distribution and the lesion stage by measuring the birefringence at the dermal layer of the living body.

The computer system converts a signal from a detector into a digital signal and receives the digital signal. The digital signal is Fourier-transformed with respect to the vertical and horizontal components to obtain a magnitude and a phase value. From the values obtained by the Fourier transform, Obtain the variable value.

The signals from the photodetectors are transmitted to the computer system of the main body via the wires of the transmission path.

According to the present invention, there is provided a computer system for measuring light scattering and birefringence by a living body tissue by converting an output signal transmitted from the probe into a digital signal, Wherein when the output signal of the probe is converted into a digital signal and a signal of a vertical component and a signal of a horizontal component are input to the computer system, And a Fourier transform step of obtaining a magnitude and a phase value by applying Fourier transform to each signal of the horizontal component; A Stokes parameter calculating step of obtaining a Stokes parameter value using the values obtained by the Fourier transforming step; A step of constructing an image of a living body tissue structure and a 3-D image with Stokes parameters obtained in the Stokes parameter operation step; A skin part scattering measurement step of measuring a cell distribution at a skin tissue part of the biotissue from the image implemented at the image implementation step; A dermal portion birefringence measuring step of measuring a birefringence at a dermal biopsy site from the image implemented at the image forming step; From the results of the scattering measurement step of the epidermal region, the change of the scattering coefficient according to the cell distribution or the cell size of the biotissue epidermis is numerically quantified, and the quantitative value of the cell distribution of the cell distribution of the epidermis of the biotissue is quantitatively A value extraction step; The quantitative value of the tumor cell lesion stage which quantitatively determines the change of the tumor cell according to the lesion stage by detecting the birefringence possessed by the tumor cells distributed in the dermal region of the biotissue from the result of the dermal site birefringence measurement step Extraction step; And a control unit.

The epidermal scattering measurement step may include: an S ₃ image obtaining step of obtaining an S ₃ image among Stokes parameters; Plane fitting is performed to extract a size of a pixel corresponding to a cell distribution of a biotissue epidermis by performing a plane correction operation process to remove a shape error occurring in an inclination of a sample caused by a surface curvature phenomenon caused by an objective lens, Computing step; A size extracting step of extracting a slope with a size value of a pixel corresponding to a cell distribution of a biom tissue epidermis extracted in the planar correction calculating step; .

The dendritic site birefringence measurement step may include calculating S ₃ images among Stokes parameters and determining _three factors of wavelength (k ₀ ), birefringence (Δn), and sample depth (z) as S ₃ = cos (2k ₀ Δnz) An S ₃ image implementation step represented by a function having a function S ₃ ; An arccosine operation step of taking an output signal of the S ₃ image implementation step as an arc cosine; A slope value extracting step of extracting a slope from an output of the arcsine calculation step to obtain a value of birefringence for one line (x); Applying the S ₃ image implementation step or the slope value extraction step to all lines (x) of the image to obtain a birefringence and averaging the averaging key.

The present invention relates to a polarization-sensitivity-optical interference image acquisition method for biopsy diagnosis that can detect changes in polarization components depending on biological tissue characteristics using a PS-OCT using a Fourier domain technique for fast image acquisition of living tissue For example.

The probe for polarizing sensitivity-optical interference image acquisition for biopsy diagnosis of the present invention comprises a light source and a computer system as one system, and an optical system and a detection system constituting a PS-OCT can be integrated into a probe. In the present invention, since the light from the light source is transmitted only to the probe using the optical fiber, the PS-OCT system in the actual probe is not affected by the polarization.

In addition, the probe for polarizing sensitivity-optical interference image acquisition for biopsy diagnosis of the present invention comprises a computer system as one system, and the optical system and the detection system constituting the light source unit and the PS-OCT can be integrated into the probe. In this case, since the physical noise on the path through which the light source is transmitted is not intervened, there is an advantage that a more accurate diagnosis can be performed.

The probe for polarizing sensitivity-optical interference image acquisition for biopsy diagnosis of the present invention acquires images non-invasively without incision or removal of tissue, and has high resolution as well as two-dimensional cross- Dimensional images can be realized in a short period of time, and the cell distribution, tumor distribution and lesion stage information of the biotissue can be quantitatively objectified.

Further, since the system is integrated into a hand-held probe, the present invention can miniaturize the system, can be manufactured as a portable device, quantitatively objectify information on cell distribution, tumor distribution, and lesion stage of a living tissue have.

Hereinafter, the configuration and operation of the polarization-sensitivity-optical interference image acquisition probe for biopsy diagnosis according to an embodiment of the present invention will be described in detail with reference to the accompanying drawings. BRIEF DESCRIPTION OF THE DRAWINGS The above and other objects, features and advantages of the present invention will become more apparent from the following detailed description of the present invention when taken in conjunction with the accompanying drawings. It should be noted that the same or similar parts in the drawings are denoted by the same or similar reference numerals for convenience of explanation and understanding. The terms "epidermis of living tissue" and "epidermis of living tissue" referred to in the present invention are only a division according to the relative depth difference from the skin surface and do not necessarily mean "medical epidermis" and "medical dermis" Should be.

1 is a schematic view for explaining a PS-OCT of a polarization sensitivity-optical interference image acquisition probe for biopsy according to an embodiment of the present invention.

High-Speed Wavelength Conversion One light from a low coherence source 110 passes through a linear polarizer (LP) 120 (or a polarized light splitter (PBS)) to produce only horizontally polarized light .

The purely horizontally polarized light travels through a beam splitter (BS) 130 and proceeds as a reference arm in one path and a sample arm in the other path.

(Explanation of vertical axis scanning using reference mirror)

The light that has flowed into the reference arm passes through a 1/4 wave plate (QWP) 140 of 22.5 占 with respect to the horizontal. The light having passed through the 1/4 wavelength plate (QWP) 140 of 22.5 ° is made to pass through the reference mirror 150 and is reflected and passes through the 1/4 wave plate (QWP) 140 of 22.5 ° again And returns to the optical splitter (BS) 130. At this time, the light returned to the optical splitter (BS) 130 becomes linearly polarized at 45 degrees with reference to the horizontal component.

The light that is polarized in the light source through the polarized light splitter (PBS) 250 and polarized in the horizontal component is described by Jones vector as follows. Jones vector is a concise technique that can apply polarized light to an incoherent light source as another expression using an electrical vector, as shown in equation (1).

The polarized light of Equation (1) is divided into a reference arm and a sample arm by an optical splitter (BS), and the light transmitted to each of the reference arm and the sample arm has the same size (Equation 2).

,

,

The Jones matrix expresses the transformation matrix of the optical component when the incident light polarized with the Jones vector is transmitted through the optical component and converted into a new vector. The Jones matrix of the QWP is expressed by Equation 3 along the horizontal and vertical axes.

The QWP rotated by? With respect to the horizontal axis is defined by Equation (4).

Here, R (Φ) is as follows.

The polarization change by QWP of 22.5 DEG according to Equation (4)

And a 45-degree linear polarization. The horizontal component and the vertical component have the same amplitude and phase.

 In this embodiment, light finally returned through the reference arm was confirmed to be 45 ° linearly polarized by using the QWP 140 and the polarized light splitter 250. Assuming that the final returning light is 45 ° linearly polarized, passing 45 ° QWP (140) once again produces fully polarized light on the horizontal axis or on the vertical axis. The polarized light can be confirmed by passing through the polarized light splitter 250. In this embodiment, the light passing through the polarization beam splitter 250 is detected only in the vertical component, and the 45-degree linear polarization of the reference arm is verified.

On the other hand, depending on the case, a light ground twisted pair (RSOD) may be used. In the case of using only the reference mirror (Ref. Mirror), it is the PS-OCT in the frequency domain in which depth information can be obtained by performing Fourier transform on the spectrum signal received from the detector without scanning in the depth direction.

(Description of sample path configuration - transverse scanning)

A transverse scanning technique is required to obtain a two-dimensional cross-sectional image of the sample. The horizontal axis scanning method is largely divided into a method of moving the sample and a method of moving the horizontal axis scanner.

 In this embodiment, the galvanometer mirror 220 is moved to scan in the transverse direction and the light is focused using the objective lens 230 between the sample and the scanning mirror 220.

The light that flows to the sample arm, which is one path other than the reference arm, passes through a 1/4 wave plate (QWP) 210 of 45 degrees with respect to the horizontal and passes through a sample 240 ), And the light at this time is circularly polarized based on the horizontal component. The light of the right circular polarization component is scattered or reflected back in the sample 240 and changes to elliptical polarization according to the optical characteristic of the sample 240. The reflected light then passes through a 1/4 wavelength plate (QWP) 210 of 45 ° and returns to the optical splitter (BS) 130.

That is, light passing through the 1/4 wavelength plate (QWP) 210 at 45 degrees in the sample arm is reflected by the mirror installed in the galvanometer 220 and focused on the objective lens 230, . Light reflected or backscattered from the sample (240) tissue is returned to the same path again. At this time, the horizontal axis scanning is performed by each rotation of the galvanometer 220.

In this embodiment, the light incident on the sample through the small movement of the galvanometer mirror 220 has a length of about 2.6 mm, and the focal length of the objective lens 230 used is set to 11 mm. The lateral resolution (x) is determined by the central wavelength of the light source and the numerical aperture of the objective lens as shown in Equation (6).

The transverse resolution calculated by Equation (6) is 20 占 퐉.

 In the PS-OCT, the light of the sample arm passes through the QWP 210 of 45 °, and circularly polarized light is incident on the sample. The light incident on the sample can be expressed as a Jones vector and a Jones matrix as described in the reference arm. The Jones matrix equation of the sample arm is the average phase delay caused by the characteristics of the sample B (z, DELTA n, alpha) is added and represented. Therefore, the QWP rotated by? On the horizontal axis in the sample arm is defined as shown in Equation (7).

Where kzn expresses light with a refractive index n and transmitted in the depth (z) direction.

Therefore, in the sample arm, the change in polarization due to 45 ° QWP 210 and B (z, Δn, α) is given by Equation 8, which indicates a polarization state that changes when the circularly polarized light incident on the sample returns backscattered .

As described above, in this embodiment, in order to acquire a large number of two-dimensional images in a short time, a spectral signal is acquired using a detector to obtain a Fourier domain, so that a vertical axis signal is acquired and a galvanometer mirror 220, The horizontal axis scanning was performed. In order to acquire an accurate range of images, it is obvious that the driving signals of the spectrum signal acquisition and the horizontal axis scanning must be synchronized and the 3D image can be realized through the synchronization. However, since the synchronization means is not a core matter of the present invention, a detailed description thereof will be omitted.

On the other hand, another galvanometer may be included so that the longitudinal axis scanning and the transverse axis scanning of the biotissue sample can be performed along the longitudinal direction of the biological tissue sample. The use of a galvanometer eliminates the need for the patient and the medical staff to move the body and allows precise longitudinal movement in synchronization with each scanning cycle, thus enabling more sophisticated image acquisition.

The light beams returned from the reference arm and the sample arm are collected by the optical splitter (BS) 130, and if the paths of the two lights are equal to each other by the reference arm, an interference phenomenon occurs. The generated interference signal passes through the polarized light splitter (PBS) 250 and is divided into a horizontal component and a vertical component, and the respective components are received by the optical detectors 260 and 270 to obtain desired data. The photodetectors 260 and 270 are spectroscopes constituting the detector 50, and may be photodiodes.

The signals detected at the photodetectors 260 and 270 are transmitted to the computer system 70. [ A collimator is disposed at the front end of the photodetectors 260 and 270. That is, the light of the horizontal component and the vertical component, which are generated by passing through the polarized light splitter (PBS) 250, And may be transmitted through a collimator to optical detectors 260 and 270 through an optical fiber.

FIGS. 2A and 2B are explanatory diagrams schematically illustrating a configuration of a polarization-sensitivity-light interference image acquisition probe for biopsy according to an embodiment of the present invention. FIG. 3B is an explanatory diagram for schematically explaining the configuration of the probe of FIG. 2B. FIG.

As shown in FIG. 2A, the main body 10 includes a power supply unit 20, a light source unit 30, and a computer system 70.

The power supply unit 20 supplies power to the light source unit 30, the detection unit 50, the computer system 70, the probe 90, and the like.

The light source unit 30 includes a high-speed wavelength-converted low coherence light source 110 and transmits the light to the optical system of the probe 90. The average time interval in which a light wave arriving at one point in a certain space oscillates predictably and maintains a sinusoidal form without a phase change is called a coherence time which determines temporal coherence . When observing at a fixed point in space, the proceeding light wave oscillates in sinusoidal form only for a time interval in which the phase remains constant. The spatial length of a light wave that vibrates regularly until the phase changes arbitrarily is called a coherence length and is a measure of how pure a spectrogram is spectrally pure. The low coherence light source is widely used in OCT recently because it has a wide spectrum band as a light source with short interference distance of light wave.

The transmission path 80 is composed of an optical fiber and wires, and serves as a channel for transmitting light between the main body 10 and the probe 90, and at the same time, transmits electric power.

3A, the probe 90 is an optical system that forms a PS-OCT and includes a linear polarizer (LP) 120, an optical splitter (BS) 130, a 1/4 wavelength plate (QWP A quartz wave plate 210, a galvanometer 220, an objective lens 230, a polarizing beam splitter (PBS) 250, And photodetectors 253 and 257.

One light from the low coherent light source 110 of the light source 30 passes through the linear polarizer (LP) 120 to produce only the light polarized in the horizontal direction.

The purely horizontally polarized light emitted from the linear polarizer (LP) 120 passes through the optical splitter (BS) 130 and travels as a reference arm in one path and to a sample arm in the other path.

The light that has flowed into the reference arm passes through a 1/4 wave plate (QWP) 140, i.e., the first quarter wave plate, of 22.5 占 with respect to the horizontal. The light having passed through the 1/4 wavelength plate (QWP) 140 of 22.5 占 is reflected by the reference mirror 150 and then passes through the 1/4 wavelength plate (QWP) 140 of 22.5 占 and passes through the optical splitter (130).

The light passing through the sample arm, which is the other path, passes through a 1/4 wavelength plate (QWP) 210 of 45 degrees to the horizontal, that is, a mirror installed on the galvanometer 220 through the second quarter- And is incident on the sample 240 after being focused on the objective lens 230. The galvanometer 220 is used as a single-axis or dual-axis optical scanner. At this time, the light is circularly polarized based on the horizontal component. The light of the right circular polarization component is scattered or reflected back in the sample 240 and changes to elliptical polarization according to the optical characteristic of the sample 240. Light reflected or backscattered from the sample (240) tissue is returned to the same path again. The reflected light then passes through a 1/4 wavelength plate (QWP) 210 of 45 ° and returns to the optical splitter (BS) 130.

The light reflected from the reference arm and the sample arm are collected by the optical splitter (BS) 130. When the path difference of the two lights is the same by the reference arm, an interference phenomenon occurs. The interference signal thus generated is transmitted through a polarizing beam splitter ) 250 to be divided into a horizontal component and a vertical component, and the horizontal component light and the vertical component light are incident on the detectors 50, 260, and 270, respectively.

The detection unit 50 is composed of photodetectors 260 and 270. The horizontal component light and the vertical component light from the polarized light splitter (PBS) 250 pass through the collimators 253 and 257, respectively, (50) to acquire data, and the obtained data is transferred to the computer system (70).

The computer system 70 converts a signal from the detector 50 into a digital signal and receives the digital signal. The received digital signal is Fourier-transformed with respect to the vertical and horizontal components to obtain a magnitude and a phase value. Obtain the Stokes parameter value from the values. From this, it is possible to quantify the size, depth, and spacing of normal cells and tumor cells in the vicinity of the biotissue epidermis, or to quantify the number of scattered areas changed in the vicinity of the epidermis of the biotissue, thereby quantitatively evaluating the tumor- And the change of the tumor cells according to the lesion stage can be quantitatively measured by measuring the birefringence depending on the stage of the lesion of the normal cells and the tumor cells distributed in the vicinity of the dermal tissue of the living tissue.

2B is different from FIG. 2A in that the light source unit 30 is embedded in the probe 90 without the light source unit 30 being present in the main body 10. FIG. FIG. 3B corresponds to the system of FIG. 2B, which shows that the light source incident on the linear polarizer (LP) 120 is not transmitted from the body but is generated in the probe 90 itself. In the case where the light source unit 30 is built in the probe 90, the optical noise is reduced as compared with the case where the light source unit 30 is received from the main body 10, thereby enabling a more accurate diagnosis. 2B and Fig. 3B, the other matters are exactly the same as those of the previous embodiment, and a duplicate description thereof will be omitted below.

The following is a brief description of the Stokes parameters.

The computer system 70 obtains the Stokes parameter value, which is a representation of the polarized light as a function of the observable quantities of the electromagnetic wave, and can be expressed as I, Q, U, V (or S0, S1, S2, and S3, respectively. These four variables can be obtained by using photodiodes and linear and circular polarizers. More information on Stokes parameters can be found in Eugene Hecht et al., "Optics, Fourth edition," published by Addison Wesley (2002).

I _{0 °} , I _{90 °} , I _{45 °} , and I _{45 °} are the linear angles of 0 °, 90 °, + 45 ° and -45 ° with respect to the horizontal axis, respectively. Is a light transmitted by a linear polarizer and is defined as linear polarized light, and I _rc and I _lc are light transmitted by a circular polarizer and are defined as right circular polarization and left circular polarization. Eox is the magnitude of the horizontal component, Eoy is the magnitude of the vertical component, εox is the phase of the horizontal component, and εoy is the phase of the vertical component. The four Stokes parameters representing the polarization state are defined as Equation (9).

I = It, Q = (I 0 ° - I 90 °), U = (I + 45 ° - I -45 °), V = (I rc - I _lc )

When we normalize the Stokes parameters of equation (9) to I, Q describes the amount of polarization along the horizontal (Q = + 1) or vertical (Q = -1) +1) or -45 ° (U = -1) direction. And V represents the polarized amount of the right circular (V = + 1) or left circular (V = + 1).

Expressing the light proceeding with the average angular frequency? As an x component (horizontal component) and a y component (vertical component) in an electromagnetic field,

The Stokes parameter expressed in the electromagnetic field is expressed by Equation (11).

This is summarized in Equation (12).

Figure 4 is a schematic flow diagram of the computer system of Figure 2a.

In the Fourier transform step, when the output signals of the photodetectors 260 and 270 are converted into digital signals and the signals of the vertical component and the horizontal component are input to the computer system, the signals of the vertical component and the horizontal component are respectively The magnitude and phase values are obtained by Fourier transform (S100).

In the Stokes variable operation step, the Stokes variable value is obtained by using the values obtained by the Fourier transform step (S150).

In the image implementation step, a tissue section structure image and a 3-D image are implemented with Stokes parameters (S200)

In the skin epidermal scattering measurement step, the distribution of cells in the epidermis of the biotissue is measured from the image implemented in the image implementation step (S250).

At the dermal site birefringence measurement step, the birefringence at the dermal tissue site of the living tissue is measured from the image implemented at the image implementation step (S300). That is, the birefringence according to the lesion stage of the tumor cells distributed in the dermal region of the biotissue is measured.

The quantitative value extraction step for the cell distribution of the epidermis of the biotissue is based on the results of the step of measuring the scattering of the epidermis of the biotissue (S250), and the change of the scattering coefficient according to the cell distribution or the cell size of the biotissue epidermis is quantified , And a quantitative value of the cell distribution of the epidermis of the biotissue is obtained (S350).

The quantitative value extraction step of the tumor cell lesion stage of the dermal region of the biotissue can be performed by detecting the birefringence possessed by the tumor cells distributed in the dermal region of the biotissue from the result of the measurement of the dermal site birefringence stage, Quantitative values for changes in tumor cells are obtained (S400).

FIG. 5 is a schematic flow chart of a step of measuring scattering of a biomicroscopic region of the tissue of FIG.

S _{3 In the} image implementation step, the S ₃ image is obtained from the Stokes parameters (S 260).

In the plane fitting operation step, a plane correction calculation process is performed to remove a shape error occurring in the inclination of the sample and the curvature of the surface caused by the objective lens (S280). Here, by performing the plane correction, the size of the pixel corresponding to the cell distribution of the epidermis of the biotissue can be extracted.

In step S290, the size of the pixel corresponding to the cell distribution of the biomedical tissue epidermis is extracted (S290), and the gradient is extracted (linear fitting) (S295) using the size value.

FIG. 6 is a schematic flow chart of the birefringence measurement step of the living tissue dermis region in FIG.

S _{3 In the} image implementation step, we obtain the S ₃ image among the Stokes parameters, where S ₃ The value can be represented by a function having three factors of a first wavelength k ₀ , a second birefringence Δn, and a third sample depth z, as shown in Equation 13 (S310).

S ₃ = cos (2k _0? Nz)

In the arc cosine operation step, the output signal of the S ₃ image implementation step is taken as arc cosine (S320). Here, S ₃ Taking the value as an arc cosine results in a linear function with a slope.

In the slope value extraction step, the slope is extracted from the output of the arccosine operation step (linear fitting), and the birefringence value for one line (x) is obtained (S330).

In the averaging step, it is applied to all lines (x) of the image to obtain a birefringence and averages the obtained birefringence (S340).

(Performance Evaluation of PS-OCT for Biopsy Diagnosis through Image Implementation)

The polarization state of light passing through a living tissue can be specified by a mechanism of scattering and birefringence. Scattering changes the polarization state in an arbitrary direction. As the scattering particle size increases, scattering becomes more anisotropic and incident polarization is maintained well. In general, the polarized light is preserved by the isotropic scattering rather than the anisotropic scattering, and the linearly polarized light is preserved as the particle size increases. When the light incident on the cell tissue is scattered many times, a change effect of the single scattering is accumulated, and the polarization state is completely changed arbitrarily, so that the incident polarization information is lost. If the biotissue is of arbitrary structure and shape, the polarization state of the light will change arbitrarily.

 Birefringence is an optical property that occurs due to a linear structure that is surrounded by materials with different refractive indices. Birefringence changes the polarization state of light due to the difference in refractive index between the optical axis of the material or the polarized light of the vertical axis, and the difference in refractive index is defined as the phase delay between the horizontal component and the vertical component of light. Equation (14) is an expression of the phase delay and varies proportionally to the refractive index difference and the moving distance.

In general, birefringence occurs in many parts of living tissue, and typically includes hard tissues such as bones, cartilage, and teeth, and soft tissues such as tendons, nerves, and muscles.

(Example for confirming birefringence of hard tissue)

In order to confirm the polarization sensitivity of the designed PS-OCT, we measured the birefringence phenomenon in the hard tissue such as bone in living tissue. Using ribs of animals that can be easily obtained from everyday life, we have confirmed the ability to distinguish images of specific layers with different refractive indices. The image of the PS-OCT is obtained by the Stokes parameters (S ₀ , S ₁ , S ₂ , S ₃ ) and the Stokes variable S ₀ image is a structural cross-sectional image expressed as the size of backscattered light as a normal OCT image , And S ₃ shows the degree of birefringence of the sample. The images of the acquired Stokes variable S ₃ are normalized by dividing by the S ₀ value. FIG. 7 shows an image of S ₀ obtained with PS-OCT, and it can be structurally confirmed that there is no special layer at a depth of about 500 μm. FIGS. 8A and 8B show images of acquired S ₃ , and the result of performing the horizontal axis scanning in the other direction at the same position. FIG. 8A is an image of S ₃ extracted from the direction of the S ₀ image, and FIG. 8B is an image obtained when the horizontal axis is scanned in the orthogonal direction at the same position. Comparing two images shows different phenomena even though they are measured at the same position. 8A shows that the phase retardation caused by the birefringence changes rapidly. In contrast, FIG. 8B shows that the phase retardation caused by the birefringence changes slowly. Therefore, it is possible to discriminate the structural layer which can not be confirmed in the structural image of S ₀ functionally in S ₃ . These results indicate that there are layers with different refractive indices, which are differently represented by the arrangement of the tissues inside the bone and the direction of the incident light. In each figure, the arrows in black and white indicate the reference point, which means that the scanning positions are the same.

In addition, the depth of information acquisition was also different. In the S ₀ image, a clear structure can be identified at a depth of up to about 500 μm. However, in S ₃ , polarization information is provided even at a position deeper than S ₀ . This is because it is difficult to obtain structural information because the light intensity is remarkably reduced at a depth of 500 μm or more. However, the S ₃ information extracted through the change of polarization means that it can be acquired sensitively even at a small light intensity.

(Example of birefringence confirmation of soft tissue through image implementation)

Next, we measured the bottom of the finger (fingerprint) and the back of the finger to verify its applicability in soft tissue such as human skin. First, the image of S ₀ and S ₃ was analyzed by measuring the light above the back of the finger (FIGS. 9A and 9B). In the S ₀ image, the basement membrane was clearly distinguished and its position was measured at a depth of about 250 μm from the epidermis.

The next S ₃ image shows birefringence occurring in the basal layer and shows a layer curvature that can not be seen in the S ₀ image. Its shape is similar to that of the wrinkles on the back of the finger, and it is periodically repeated. This means that the pleomorphic tissues in the basal layer are arranged in the form of wrinkles. Figs. 9C and 9D show the result of measuring fingerprint portions. The image shows a distinctive basement membrane formed along the shape of the fingerprint. The bold line in black color on the epidermis is clearly expressed compared to other structures because the intensity of backscattering light is relatively large in the epidermis and the length of one period of periodically repeated bending is about 400 μm. FIG. 9D is a birefringence image of the fingerprint, and it is possible to confirm a change in the polarization component periodically appearing according to the bending of the fingerprint. However, in the S ₃ image, the basement membrane could not be clearly distinguished from the S ₀ , suggesting that the tissue existing in the basal layer of the fingerprint region is not formed like the fingerprint pattern unlike the back region.

(3D imaging embodiment of uterine tissue sample)

A synchronized signal of the system was verified by moving the 2D image of 2 mm (vertical axis scanning range) × 2.3 mm (horizontal axis scanning range) by 20 μm by moving the motion controller 310 100 times and implementing 3D image. Fig. 10 shows an image of 3D composed of a size of 2 mm x 2.3 mm x 2 mm of a uterine tissue sample.

(High-speed scanning PS-OCT for early diagnosis of cervix)

The cervical intraepithelial neoplasm is a term used to describe the continuous process of cervical cancer precursor lesions in the cervical epithelium. CIN I, CIN II, and CIN III are classified into three grades according to the extent of lesion progression. They begin with CIN I, a low-grade epithelium tumor, and progress to CIN III, a high grade epithelial tumor. There is a feature that can be.

Figure 11 shows the anatomical division of the cervix. Squamo-Columnar Junction is a site where the epithelium of the inner circumference and the circumference of the outer diameter meets, and it is a clinically important part where the epithelial tumor and cancer occur.

Cervical intraepithelial neoplasia (HCC) has the following histologic features as a stage progression, and it is a basis for diagnosis of lesion.

First, the epithelium is divided into multiple layers as a result of maturation and differentiation. The grade of epithelial lesion is determined by the thickness ratio of epithelial layer. The higher undifferentiated lesion, the thicker the undifferentiated cell layer, the more mature and differentiated cell layer becomes thinner or disappear.

Second, the rate of hyperchromasia, nuclear pleomorphism, nuclear-cytoplasm increases and the number of mitosis increases.

Third, keratinization develops and keratinized epithelium appears as leukoplakia. Generally, it is most severe in high-grade epithelial tumors but also in low-grade epithelial tumors.

When the uniformly polarized light passes through a medium having a high scattering coefficient such as a living tissue, the polarization component changes due to factors of particle size, dispersion degree and density of the scattering substance. The change in the polarization component is regarded as the only control mechanism in the optical positively imaging method of living tissue, and it uses light scattered repeatedly many times. Generally, the change of polarization is correlated with the wavelength of the incident light and the size of the medium. If the size of the scatterer is larger than the wavelength range, the property of the linear polarization is maintained longer than that of the circular polarization, and if the wavelength range is larger than the scatterer, the circularly polarized light will retain the polarization component longer than the linearly polarized light.

In epithelial tumors, the proportion of nuclei to the cells increases with the number of cells at which the mitosis occurs. As a result, the number of scattered light for the incident light will increase, which in turn affects the degree of polarization (DOP) of light and the depolarization of polarization.

The polarization change of the light incident on the sample by circularly polarized light can be described using a Jones matrix or a Mueller matrix. In general, the Jones matrix is used to describe phenomena that contain absolute values of the phase and amplitude of light and are either perfectly polarized or have no depolarization. The Mueller matrix contains the Jones matrix and is an expression for describing the interaction of polarization states and is used to explain the extinction of polarization.

 The degree of circular polarization (DOCP) represents the change in polarization that occurs when circularly polarized light passes through a living tissue. Factors of DOCP change in biological tissues include polarization extinction due to scattering, phase delay due to birefringence, and noise.

 Changes in the DOCP basically indicate a change in the polarization component in the tissue. However, the influence of the DOP on the variation of the polarization component occurring when passing through an optical device and the area size to be detected and the persistence should be considered. In a system using a heterodyne detection system such as a PS-OCT, DOP is maintained irrespective of the size of the photodetector despite the large scattering, and the polarization component changed by the optical device is less than 10% of the total polarization change It is known that there is little influence on acquiring the change of the polarization component occurring in the biotissue. This is an important basis for distinguishing the characteristics of living tissue by measuring DOCP changes in PS-OCT.

High-speed scanning To obtain clinical data on PS-OCT, cervical tissue sections were sampled. As shown in FIG. 12A, the cervical tissues removed from the body for pathological diagnosis were measured in a drug pretreatment process, and the age of the patients varied from 20 to 30 years old (FIG. 12B). We obtained 50 ~ 200 cross - sectional images at the time of acquiring one data from the samples obtained for each age group and measured at 16 sites expected to be normal tissues and 9 sites expected to be lesion tissues. Among the measured sample sites, the diseased tissue was part of the tissue removed from the 20 patient's body and was diagnosed as a high-grade epithelial tumor by an obstetrician. The measurement area of the sample shown in the photograph of FIG. 12A was measured at 3:00, 6:00, 9:00, and 12:00, where the frequent occurrence of intraepithelial neoplasms was taken, and the direction of the entrance of the uterus was 12:00.

First, the tissues extracted from the cervix of a patient without a lesion were scanned and imaged with a size of 2 mm × 2.3 mm (longitudinal axis × transverse axis) and then compared with a histology image (FIG. 13 c). Since the sample used was a normal tissue and no mitosis was observed, there is a boundary line between the epithelial layer (e) and the basal layer (s), which can be confirmed by a histological image.

In the light intensity image obtained using the PS-OCT (FIG. 13A), the boundary line between the skin layer and the basal layer was clearly distinguished, and the depth of the skin layer was measured to be about 200 μm. 13B shows an image showing the disappearance of polarization. As the light reaches from the epidermis to the basal layer of the uterine tissue, the change of the polarization is not seen, and the change of the polarization occurs at the portion below the base layer.

Figures 14a, b and 15a, b compare the light intensity image and the polarized image of different normal tissues. Light intensity images show different phenomena despite being normal tissues. In the light intensity image of FIG. 14A, the discrimination of the basement membrane layer could be confirmed clearly, but in the image of FIG. 15A, it could not be distinguished from the epidermal layer and the basal layer. In general, it is known that the thickness of the epithelium becomes thinner and the distinction between the epithelial layer and the basal layer disappears as the lesion of the cervical intraepithelial tumor progresses. However, in practice, there is no basement membrane in normal tissues without lesions, and since they are formed differently according to the characteristics of individuals, this difference in the same normal tissues is a clear basis for the presence of basement membrane in the determination of cervical lesions It means that you can not.

Next, images of normal tissue and high-grade epithelial tumor (H-SIL) were compared (FIG. 16). High-grade epithelial tumors have an increased scattering ratio of cells and nuclei, It is expected that the depth of transmitted light is shorter than that of normal tissue. 16A to 16D, it can be seen that a signal can be obtained to a depth of 500 mu m or more in a normal tissue, but a signal in the base layer region is weakly detected in an image of H-SIL . These results suggest that the light penetrates through the normal tissue in H-SIL, resulting in a change in the depth of penetration due to the effect of larger scattering. However, in the light intensity image of H-SIL, the layer is divided at the lower part of the epidermis and the depth of the transmitted light can not be distinguished from the normal tissue. Therefore, It is not enough to distinguish between tissue and normal tissue.

Cervical intraepithelial neoplasia develops in the depth of 400 μm from the epidermis and proceeds. Measurement of cervical tissues by the designed PS-OCT system showed that the depth of the longitudinal axis of the obtained images could be used to identify the basement membrane and to include information on the range of possible lesions and to observe the polarization change information (FIG. 16B and FIG. 16e). However, since lesions of the cervical intraepithelial neoplasia can be observed in the histological images of FIGS. 16C and 16F, only the light intensity image which can observe only the structural features can be observed, -It is difficult to distinguish SIL organization. Therefore, in order to more accurately distinguish the characteristics between normal tissue and lesion tissue, it is necessary to analyze not only the structural image but also the change of polarization component.

Thus, as another example, we have analyzed the polarization component signals of acquired images to determine the exact state of the cervix. The signal intensity of the light intensity image was analyzed to measure the thickness of the basal membrane and the DOCP change was confirmed using the Stokes parameter information. FIG. 17A shows a 2D image represented by the light intensity, and the presence of the basement membrane was confirmed by contrast intensity intensity difference. The total depth that can be confirmed by the intensity of the signal on the image was about 600 μm, and signals below that are expressed at the same level as the air area, so that it can be determined that the amount of detected light is almost zero. In order to measure the exact position of the basement membrane identified in the image, vertical axis signals were expressed according to the intensity of light. The highest signal on the surface of the sample tissue was measured and the second largest signal in the basement membrane was detected. The exact location of the basement membrane measured using this was about 250 μm from the epidermis, which means the thickness of the epithelial layer of the sample tissue.

FIG. 18 is a graph for analyzing the Stokes parameter signal. FIG. 18 shows Stokes parameters obtained through the demodulation process and the signal process of the interference signal detected by the detector, and a graph re-expressed by DOCP is shown. FIG. 18A shows the DOCP of the epidermis expressed as 1 through a generalization process, and is a relative representation of a change with depth. It can be seen that the DOCP decreases with a depth of about 330 μm, which is caused by a decrease in the amount of scattered light. In general, lesions occurring in the uterine cervix are located near the epidermis and basement membrane, so the signals from the epidermis to the depth of 300 μm are selected to calculate the DOCP reduction ratio. The reduction rate obtained through the linear fitting (first linear function) can determine how fast the polarization component disappears as the incident circularly polarized light proceeds in the depth direction (FIG. 18B). In this figure, the decrease rate of polarization shows the degree of depolarization at 1 mm movement and 0.76204 depolarization / mm result. This value is a measure of normal tissue as a sample, and if the lesion is a tissue with altered cells, the larger the linear fitting result can be predicted because the polarization will disappear faster with depth.

Next, we compared the mean values of DOCP changes in normal tissues and H-SIL tissues (FIG. 19). The average value was extracted after performing the linear fitting from the epidermis to the depths of 200 μm and 300 μm. First, the slope of the DOCP according to the depth shows that the DOCP slope value increases as the depth range analyzed in both normal and H-SIL tissues increases. This means that as the circularly polarized light passes through the cervical tissue and travels in the depth direction, the polarization becomes more lossy. Considering that the light intensity is maintained at a depth of up to about 350 占 퐉 in FIG. 18, signal analysis in the range of 200 占 퐉 and 300 占 퐉 from the skin shown in this embodiment can be independently judged It can be grounds for doing. Comparison of DOCP slope for normal tissue and H-SIL tissue shows that DOCP slope in H-SIL is larger than that in normal tissue, and slope difference at 300 μm depth is smaller than at 200 μm depth ≪ / RTI > These results are consistent with the expectation that changes in the DOCP will increase as the number of scattering units increases as the cervix progresses to high-grade epithelial tumors such as H-SIL. Thus, changes in the DOCP extracted from PS-OCT can be a parameter to determine the presence of cervical intraepithelial neoplasia.

Cervical cancer is the most common cancer in women, accounting for 42.1% of female cancer patients, but due to the nature of cervical cancer, it is difficult to diagnose early. Therefore, in this example, the possibility of early diagnosis of cervical cancer using PS-OCT was investigated.

PS-OCT is a new optical imaging technique that acquires high-resolution images in a non-invasive manner using light harmless to the human body. It acquires depth-dependent changes in backscattering caused by the scattering action of living tissue, And additionally acquires a change in the polarization component. In this embodiment, a PS-OCT system based on a high-speed scanning optical artifact (RSOD) technique was designed and manufactured using a light source of a near-infrared region (1296 nm). The system has a vertical resolution of 18 ㎛ and it takes 1 second each to acquire and store data, and a 3D image with a relatively small error range can be realized by using the synchronized signal.

First, to confirm the performance of the system that is sensitive to polarized light, the change in DOCP was observed in a highly scattering medium (intra-lipid). A mixture of 20% intra-lipid solution and distilled water was used to prepare 2.5%, 5%, and 10% mixed solutions. The concentration of DOCP in the solution was analyzed and the results showed that the concentration of DOCP increased linearly .

Next, structural form and birefringence images were measured using animal bones. We obtained a 2D cross-sectional image using the animal bone that can be easily obtained from the surrounding area, and obtained a signal that can confirm the structure at a depth of about 500 μm. Then, a birefringence image was obtained to observe the change of the polarization component. It was confirmed that the PS-OCT can detect the birefringent structure existing inside the bone through the birefringence image which is different from the polarization direction by observing the change of polarization component by performing the scanning direction at the same position.

We measured structural information and polarization changes in the soft tissue by measuring fingerprint area and fingerprint area of human finger. In the light intensity image, the structural form in the basal layer was observed, and the shape of the basement membrane formed by the wrinkles of the epidermis was clearly observed. In addition, the arrangement of layers with different refractive indices in the epidermis was confirmed in the birefringence image.

Finally, the changes of lesion tissue and normal tissue DOCP of cervix were extracted. The histological diagnosis of the lesion was made by measuring the rate of change of the DOCP using the section of the tissue removed from the patient's body. The H-SIL was examined for H-SIL in the normal tissue and high-grade epithelial tumor (H-SIL) The rate of change was measured largely. As a result, we could distinguish the histologic features of normal and epithelial tumors by changing the polarization components obtained from the cervix.

In addition, based on a deeper understanding of the scattering effect of light incident on living tissue, various opinions of gynecologists are gathered, and when data of normal tissues and stage epithelial tumors are collected and quantified through a large number of cervical biopsies, PS The change in DOCP acquired using -OCT allows early diagnosis of cervical lesions.

While the present invention has been particularly shown and described with reference to exemplary embodiments thereof, it is to be understood that the invention is not limited to the disclosed exemplary embodiments. Therefore, the present invention is not limited to the above embodiments, and the true scope of the present invention should be determined by the technical idea of the appended claims. It will be apparent to those skilled in the art that various modifications and equivalent embodiments may be made without departing from the spirit and scope of the invention as defined by the appended claims.

For example, the terms 'epidermis of living tissue' and 'epidermis of living tissue' referred to in the present invention are only a division according to the relative depth difference from the surface of skin, and they necessarily mean 'medical epidermis' and 'medical epidermis' It should be noted that it is not.

1 is a schematic view for explaining a PS-OCT of a polarization sensitivity-optical interference image acquisition probe for biopsy according to an embodiment of the present invention.

FIG. 2A is an explanatory view schematically illustrating a configuration of a polarization-sensitivity-light interference image acquisition probe and a main body for biopsy according to an embodiment of the present invention.

FIG. 2B is an explanatory view schematically illustrating the configuration of a probe for polarized light sensitivity-optical interference image acquisition and a main body for biopsy according to an embodiment of the present invention.

FIG. 3A is an explanatory view schematically illustrating the configuration of the probe of FIG. 2A.

FIG. 3B is an explanatory view schematically illustrating the configuration of the probe of FIG. 2B.

Figure 4 is a schematic flow diagram of the computer system of Figure 2a.

FIG. 5 is a schematic flow chart of a step of measuring scattering of a biomicroscopic region of the tissue of FIG.

FIG. 6 is a schematic flow chart of the birefringence measuring step of the dermal biopsy region of FIG.

Figure 7 is an S0 cross-sectional image of an animal bone.

8A is a birefringence image for the scanning direction as shown in FIG.

FIG. 8B is a birefringence image for the scanning direction intersected with FIG. 7. FIG.

9A is an S ₀ image on the back part of the finger.

9B is an S ₃ image on the back of the finger.

9C is an S ₀ image of the fingerprint area of the finger.

9D is an S ₃ image on the back of the finger.

10 is a 3D image of a 2 mm x 2.3 mm x 2 mm size uterine tissue sample using a probe for obtaining a polarization sensitivity-optical interference image for diagnosis of the cervix according to the present invention.

11 is an anatomical structure diagram of the cervix.

12A is a photograph of a real tissue of the cervix.

12B is a table showing the distribution of the patient and the sample used in the present embodiment.

13A is a light intensity image for comparison between normal tissues.

13B is a polarized image for comparison between normal tissues.

Figure 13c is a histological image for comparison between normal tissues.

14A is a light intensity image for comparison between normal tissues.

14B is a polarized image for comparison between normal tissues.

15A is a light intensity image for comparison between normal tissues.

15B is a polarized image for comparison between normal tissues.

16A is a light intensity image of normal tissue.

16B is a polarized image of normal tissue.

16C is a histological image of normal tissue.

16D is a light intensity image of high grade epithelial tumor tissue (H-SIL).

16E is a polarized image of high grade epithelial tumor tissue (H-SIL).

16F is a histological image of high grade epithelial tumor tissue (H-SIL).

17A is a light intensity image of normal tissue.

17B is a graph showing the analysis of the light intensity signal of the normal tissue.

18A is a graph re-expressed as DOCP for analyzing the Stokes variable S3.

18B is a graph obtained by Liner fitting the DOCP graph.

FIG. 19A is a graph comparing mean values of DOCP changes of normal tissue and H-SIL tissue from epidermis to 200 μm. FIG.

FIG. 19B is a graph comparing mean values of DOCP changes of normal tissue and H-SIL tissue from epidermis to 300 μm.

Description of the Related Art

10: main body 20: power supply

30: light source 70: computer system

80: transmission path 90: probe

110: low coherent light source 120: linear polarizer (LP)

130: optical splitter (BS) 140: quarter wave plate (QWP)

150: Reference mirror

160: diffraction grating 170: optical retardation lens

180: Galvanometer 185: Double-pass mirror

210: quarter-wave plate (QWP) 220 at a 45 DEG angle: galvanometer

230: objective lens 240: sample

245: cervical contact 250: polarized light splitter

260, 270: photodetector 310: motion controller

Claims (17)

The probe comprises a probe and a body, wherein the probe detects an optical signal including information on a living tissue, and the body comprises a computer system for measuring and analyzing a living tissue from an optical signal received from the probe, 1. A polarization sensitivity-optical interference image acquiring probe for forming the probe in an imaging system, The probe for obtaining a polarization sensitivity-light interference image is a probe for obtaining a polarization sensitivity-light interference image by using a light scanner of the living tissue to perform horizontal axis scanning and using a Fourier domain technique capable of vertical axis scanning without movement of a mirror, (PS-OCT), which is capable of detecting birefringence, and a polarization-sensitive optical coherence tomography (PS-OCT) The probe comprises: a linear polarizer (LP) for producing only one light from the light source in the horizontal direction; An optical splitter (BS) for causing the purely horizontally polarized light output from the linear polarizer to travel through a sample arm as a reference arm in one path and as a sample arm in the other path; A first quarter wave plate (QWP) for transmitting light that has flowed from the linear polarizer to the reference arm to a reference mirror and returning light reflected from the reference mirror to the optical splitter; To allow the light from the linear polarizer to pass through the mirror installed in the galvanometer and to reflect or backscatter from the sample tissue to return the light reflected by the mirror mounted on the galvanometer back to the optical splitter A second quarter wave plate (QWP); A second quarter wave plate (QWP) to the galvanometer to focus the light reflected by the mirror mounted on the galvanometer and to enter the sample tissue; An objective lens for transmitting the emitted light to a mirror installed in the galvanometer; A polarized light splitter (PBS) that returns from the reference arm and the sample arm, and the interference signal generated from the light collected by the optical splitter is divided into a horizontal component light and a vertical component light; And a photodetector for converting the horizontal component light emitted to the polarized light splitter and the vertical component light into respective electrical signals. The polarized light sensitivity-optical interference image acquisition probe according to claim 1, delete An optical splitter (BS) for causing the light output from the light source unit to pass through the linear polarizer to a reference arm in one path and a sample arm in the other path; A first quarter wave plate (QWP) for transmitting light that has flowed from the linear polarizer to the reference arm to a reference mirror and returning light reflected from the reference mirror to the optical splitter; To allow the light from the linear polarizer to pass through the mirror installed in the galvanometer and to reflect or backscatter from the sample tissue to return the light reflected by the mirror mounted on the galvanometer back to the optical splitter A second quarter wave plate (QWP); A polarized light splitter (PBS) that returns from the reference arm and the sample arm, and the interference signal generated from the light collected by the optical splitter is divided into a horizontal component light and a vertical component light; Photodetectors for converting the horizontal component light emitted to the polarized light splitter and the vertical component light into respective electrical signals; And a probe for obtaining a polarization sensitivity-optical interference image for biopsy. The method of claim 3, Light from the second quarter wave plate (QWP) to the sample arm is transmitted to the galvanometer, and the light reflected by the mirror installed on the galvanometer is focused to enter the sample tissue, An objective lens for transmitting light reflected or scattered back from the sample tissue to a mirror installed in the galvanometer; Further comprising: a polarization-sensitivity optical interference image acquisition probe for biopsy. 4. The method according to any one of claims 1 to 3, The apparatus according to claim 1, further comprising a uniaxial galvanometer for longitudinal axis scanning and transverse axis scanning of the biological tissue sample along the longitudinal direction of the biological tissue sample. . 4. The method according to any one of claims 1 to 3, Wherein the light source of the light source unit is composed of a high-speed wavelength-converted low coherence light source, and a polarizing sensitivity-optical interference image acquisition probe for biopsy. The method according to claim 1 or 3, Wherein the first quarter-wave plate is a 1/4 wavelength plate of 22.5 占 The polarizing sensitivity-optical interference image acquiring probe for biopsy. The method according to claim 1 or 3, And the second 1/4 wavelength plate is a 1/4 wavelength plate of 45 °. The method according to claim 1 or 3, Wherein the photodetectors are made of a photodiode. The polarimetric sensitivity-optical interference image acquisition probe according to claim 1, The method according to claim 1, Wherein the computer system measures the light scattering and the birefringence of the biotissue to obtain a quantitative value of the tumor tissue distribution and the lesion level, thereby obtaining a polarization sensitivity-optical interference image acquisition probe for biopsy. 11. The method of claim 10, Wherein the computer system obtains quantitative values of tumor tissue distribution and lesion stages by measuring light scattering of the biomedical skin layer region. 11. The method of claim 10, Wherein the computer system measures the birefringence at the site of the dermal layer of the living body to determine a quantitative value of the tumor tissue distribution and the lesion stage. The method according to claim 1, The computer system converts a signal from a detector into a digital signal and receives the digital signal. The digital signal is Fourier-transformed with respect to the vertical and horizontal components to obtain a magnitude and a phase value. From the values obtained by the Fourier transform, And a value of a variable is obtained. A probe for obtaining polarization sensitivity-optical interference image for biopsy. The method according to claim 1, And a signal from the photodetectors is transmitted to a computer system of the main body through an electric wire of a transmission path. A polarization sensitivity-optical interference image acquisition probe for biopsy. And a computer system for measuring scattering and birefringence of light by a living body tissue by converting an output signal received from the probe into a digital signal, receiving and processing the digital signal, A method of driving a polarization-sensitivity-optical interference imaging system, When the output signal of the probe is converted into a digital signal and a signal of a vertical component and a signal of a horizontal component are input to the computer system, a magnitude and a phase value are obtained by subjecting the signal of the vertical component and the signal of the horizontal component to Fourier transform A Fourier transform step; A Stokes parameter calculating step of obtaining a Stokes parameter value using the values obtained by the Fourier transforming step; A step of constructing an image of a living body tissue structure and a 3-D image with Stokes parameters obtained in the Stokes parameter operation step; A skin part scattering measurement step of measuring a cell distribution at a skin tissue part of the biotissue from the image implemented at the image implementation step; A dermal portion birefringence measuring step of measuring a birefringence at a dermal biopsy site from the image implemented at the image forming step; From the results of the scattering measurement step of the epidermal region, the change of the scattering coefficient according to the cell distribution or the cell size of the biotissue epidermis is numerically quantified, and the quantitative value of the cell distribution of the cell distribution of the epidermis of the biotissue is quantitatively A value extraction step; The quantitative value of the tumor cell lesion stage which quantitatively determines the change of the tumor cell according to the lesion stage by detecting the birefringence possessed by the tumor cells distributed in the dermal region of the biotissue from the result of the dermal site birefringence measurement step Extraction step; Wherein the polarization-sensitivity-optical interference imaging system comprises: 16. The method of claim 15, The skin portion scattering measurement step S ₃ S ₃ image implement obtaining images from the Stokes parameters; Plane fitting is performed to extract a size of a pixel corresponding to a cell distribution of a biotissue epidermis by performing a plane correction operation process to remove a shape error occurring in an inclination of a sample caused by a surface curvature phenomenon caused by an objective lens, Computing step; A size extracting step of extracting a slope with a size value of a pixel corresponding to a cell distribution of a biom tissue epidermis extracted in the planar correction calculating step; Wherein the polarization-sensitivity-optical interference imaging system comprises: 16. The method of claim 15, The dermal site birefringence measurement step S ₃ images were obtained from Stokes parameters, S ₃ = cos (2k _0? Nz) S ₃ image realization step represented by a function having three factors of wavelength (k ₀ ), birefringence (Δn), and sample depth (z) as shown in FIG. An arccosine operation step of taking an output signal of the S ₃ image implementation step as an arc cosine; A slope value extracting step of extracting a slope from an output of the arcsine calculation step to obtain a value of birefringence for one line (x); Obtaining the birefringence by applying the S ₃ image implementation step or the slope value extraction step to all the lines (x) of the image and averaging the averaging step; Wherein the polarization-sensitivity-optical interference imaging system comprises:

Images