KR20090093368A - Probe for acquiring polarization-sensitive optical coherence image - Google Patents

Abstract

A probe for acquiring polarization-sensitive optical coherence image is provided to acquiring the image on real time basis to the longitudinal axis and the horizontal axis through the polarizing component image. A probe for acquiring polarization-sensitive optical coherence image comprises a probe(90) and a main body(10). The probe detects an optical signal including the information of a biological tissue. The main body includes a computer system(70) measuring and analyzing the biological tissue from the light source. The probe scans horizontally using an optical scanner of the biological tissue. The probe uses a Fourier domain method which can scan vertically without movement.

Description

Probe for acquiring polarization-sensitive optical coherence image}

The present invention relates to a polarization-sensitive optical interference image acquisition probe for measuring the distribution of cells in or outside the body for the diagnosis of various biological tissues, and more particularly, the scattering of light in the cell distribution near the epidermis, such as living tissue or lung The present invention relates to a probe for obtaining polarization sensitivity-optical interference imaging, which enables measurement of a birefringence of collagen fibers distributed near the dermis, cartilage, and the like of the skin, thereby enabling a biopsy.

Optical Coherence Tomography (hereinafter referred to as OCT) is a device that captures living tissues or cells in high resolution in real time using light that is harmless to the human body. Not only can you distinguish between soft tissues, you can get more precise images. OCT system is composed of wide band light source and Michelson interferometer, and the application range of Ophthalmology, Dermatology, Gastroenterology, Dentistry is expanding due to the development of light source and technology.

However, the OCT technology has a disadvantage in that a precise image cannot be obtained at a deep portion (for example, 500 μm or more deep from the epidermis) in the precise measurement. This is because the scattered light amount is not sufficiently recollected at the deep portion.

Polarization Sensitive-Optical Coherence Tomography (hereinafter referred to as PS-OCT) acquires the variation of backscattered light caused by scattering of biological tissues by depth and implements the tissue in basic OCT PS-OCT is based on the Michelson interferometer principle. Compared to OCT, which only implements a structured image, PS-OCT provides more accurate functional information that can be difficult to obtain when imaging with a general optical method, for example, an image of a human body at least deep in amount of light. There is this.

PS-OCT system combines light source, optical system part, detection system part, and computer system into one big system. Because optical fiber is sensitive to polarization, it can transmit light from light source to optical system part of PS-OCT. However, it is difficult to use optical fibers in optical system parts, that is, reference arm or sample arm.

Therefore, it was difficult to apply the method of using the OCT system and the probe independent of each other, which is the method used in the general OCT system, to the PS-OCT system.

In particular, in a biotissue imaging system, a computer system and a probe need to be used independently of each other.

The present invention comprises a light source and a computer system as one system, and provides a polarization-sensitized optical-interfering probe for diagnosing a biological tissue in which an optical system and a detection system forming a PS-OCT are integrated in a probe. In the present invention, since the light emitted from the light source is transmitted only to the probe using the optical fiber, polarization is not affected in the PS-OCT system in the actual probe.

In another aspect, the present invention provides a probe for a polarization sensitivity-optical interference image acquisition for the diagnosis of a biological tissue that consists of a computer system, a light source and an optical system and PS-OCT integrated in the probe. In the present invention, since the light source is generated in the probe, the noise is less affected, and thus more accurate results can be obtained.

In addition, the present invention provides a polarization-sensitive optical interference imaging system for diagnosing a biological tissue using PS-OCT applying the Fourier domain technique for fast image acquisition of biological tissue, which can detect the change in polarization components that vary depending on the characteristics of the biological tissue to provide.

The present invention can acquire non-invasive image without cutting or detaching the cell tissue, and can obtain not only the two-dimensional cross-sectional structure image of the biological tissue but also a polarization component image in real time with high resolution, and also the two-dimensional image Since real-time acquisition is possible, it is possible to realize a 3D image in a short time.

The present invention integrates the system into a small hand-held probe, so the system can be miniaturized and can be made portable, and can have a plug & play function against a main body including a computer system. In addition, the degree of lesion stage and treatment effect on biological tissues can be quantitatively objective.

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 formed as one system, and an optical system and a detection system that form a PS-OCT are integrated into the probe. There is.

Another technical problem to be achieved by the present invention is to use a PS-OCT applied Fourier domain technique for fast image acquisition of biological tissue, polarization sensitivity-optical interference image that can detect the change in polarization components that vary depending on the characteristics of the biological tissue It is to provide an acquisition probe.

Another technical problem to be achieved by the present invention is to acquire an image in a non-invasive manner, to obtain a two-dimensional high-resolution image of the longitudinal and transverse axis of the biological tissue in real time through a polarization component image, in real time at a fast time along the longitudinal direction According to the present invention, a three-dimensional image can be implemented, and a polarization sensitivity-optical interference image acquisition probe capable of quantitatively objectively identifying a lesion stage and a treatment effect on a living tissue.

The present invention has been made to achieve the above object, the polarization sensitivity-optical interference image acquisition probe for a biotissue examination of the present invention comprises a probe and the body, the probe is a light containing information of the biological tissue In the polarization sensitivity-optical interference image acquisition probe forming the probe in the polarization sensitivity-optical interference imaging system consisting of a computer system for detecting a signal, the computer system for measuring and analyzing the biological tissue from the optical signal received from the probe , The polarization sensitivity-optical interference image acquisition probe, the horizontal scanning using the optical scanner of the biological tissue, the scattering of light by the biological tissue using a Fourier domain technique capable of vertical scanning without the movement of the mirror And a system of polarization sensitivity-optical coherence tomography (PS-OCT) for detecting birefringence. It characterized by comprising.

The probe may include a linear polarizer (LP) for generating one light emitted from the light source unit using only light polarized in a horizontal direction; An optical splitter (BS) for transmitting the purely horizontally polarized light output from 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 the light flowing from the linear polarizer to the reference arm and returning the light reflected from the reference mirror to the optical splitter; The light flowing from the linear polarizer to the sample arm is transmitted to the mirror mounted on the galvanometer, and the light reflected by the mirror mounted on the galvanometer is reflected or reverse scattered from the sample tissue to return to the optical splitter. A second quarter wave plate QWP; The light transmitted from the second quarter wave plate (QWP) to the galvanometer focuses the light reflected by the mirror installed in the galvanometer to be incident on the sample tissue, and is reflected or reverse scattered from the sample tissue. An objective lens for transmitting the reflected light back to a mirror installed in a galvanometer; A polarization optical splitter (PBS) in which the interference signal generated from the light returned from the reference arm and the sample arm and collected in the optical splitter is divided into a horizontal component light and a vertical component light; And light detectors for converting the horizontal component light emitted by the polarized light splitter and the vertical component light into electrical signals.

According to the longitudinal direction of the biological tissue sample, it further comprises a galvanometer of one axis to enable longitudinal and horizontal scanning of the biological tissue sample.

The light source of the light source unit is a high speed wavelength conversion low coherence light source.

The first quarter wave plate is a quarter wave plate of 22.5 degrees, and the second quarter wave plate is a quarter wave plate of 45 degrees.

The photo detectors are made of photodiodes.

The computer system obtains quantitative values for tumor tissue distribution and lesion stage by measuring light scattering and birefringence of the biological tissue.

The computer system obtains quantitative values for tumor tissue distribution and lesion stage by measuring light scattering at the epidermal layer of the biological tissue.

The computer system obtains quantitative values for tumor tissue distribution and lesion stage by measuring the birefringence of the biotissue dermal layer region.

The computer system converts and receives a signal from detectors into a digital signal, obtains magnitude and phase values by performing Fourier transform on the received digital signal for vertical and horizontal components, respectively, and uses Stokes from values obtained by the Fourier transform. Get the value of a variable.

The signals from the photo detectors are transferred to the computer system of the main body via the wire of the transmission path.

According to the present invention, a light source unit for generating and transmitting a light source to the probe, and a computer system for measuring the scattering and birefringence of light by the biological tissue by converting the output signal received from the probe into a digital signal, receiving and processing A method of driving a polarization sensitivity-optical interference imaging system having a main body, wherein the output signal of the probe is converted into a digital signal and a vertical component signal and a horizontal component signal are input to the computer system. And a Fourier transform step of performing Fourier transform on the signals of the horizontal component to obtain magnitude and phase values. A Stokes variable calculation step of obtaining a Stokes variable value using the values obtained by the Fourier transform step; An image realization step of implementing a biopsy cross-sectional structure image and a 3-D image using the Stokes parameters obtained in the Stokes variable calculation step; Epidermal region scattering measurement step of measuring the cell distribution in the epidermal region of the biological tissue from the image implemented in the image implementation step; A dermal region birefringence measuring step of measuring the birefringence in the dermal region of the biological tissue from the image implemented in the image forming step; Quantification of the cell distribution to quantitatively change the scattering coefficient according to the cell distribution or cell size of the biological tissue epidermis from the results of the scattering measurement step of the epidermal region, and to obtain a quantitative value for the cell distribution of the epidermal region of the tissue. Phosphorus value extraction step; Quantitative value of the tumor cell lesion step to obtain a quantitative value for the change of tumor cells according to the lesion stage by detecting the birefringence of the tumor cells distributed in the dermal region of the biological tissue from the result of the dermal site birefringence measurement step Extraction step; Characterized in that it comprises a.

The skin site scatterometry step, S ₃ S ₃ image implement obtaining images from the Stokes parameters; Plane fitting to extract the size of the pixel corresponding to the cell distribution in the epidermal area of the tissue tissue by performing a planar correction process to remove the shape curvature caused by the objective lens and the slope of the sample Operation step; Extracting a size of a pixel having a size value of a pixel corresponding to a cell distribution of the epidermal region extracted from the planar correction operation; It includes.

The dermal region birefringence measurement step is to obtain the S ₃ image from the Stokes parameters, such as S ₃ = cos (2k ₀ Δnz) three factors of the wavelength (k ₀ ), birefringence (Δn), the depth of the sample (z) S ₃ image implementation step represented by a function with; An arc cosine operation step of taking an arc cosine of the output signal of the S ₃ image implementation step; A slope value extraction step of obtaining a birefringence value for one line x by linearly fitting a slope from an output of the arc cosine operation step; And applying the S ₃ image implementing step to the gradient value extracting step to all the lines x of the image to obtain birefringence and averaging them.

The present invention, using the PS-OCT applying the Fourier domain technique for fast image acquisition of biological tissue, polarization sensitivity-optical interference image acquisition for the diagnosis of biological tissue that can detect the change in polarization components that vary depending on the characteristics of the biological tissue To provide a probe.

The polarization sensitivity-optical interference image acquisition probe for diagnosing a biological tissue of the present invention comprises a light source unit and a computer system as one system, and integrates an optical system and a detection system constituting the PS-OCT into the probe. In the present invention, since the light emitted from the light source is transmitted only to the probe using the optical fiber, polarization is not affected in the PS-OCT system in the actual probe.

In addition, the polarization sensitivity-optical interference image acquisition probe for diagnosing a biological tissue of the present invention may be a computer system, and an optical system and a detection system constituting the light source unit and the PS-OCT may be integrated into the probe. In this case, since physical noise on the path through which the light source is transmitted is not involved, there is an advantage that more accurate diagnosis is possible.

Polarization sensitivity-optical interference image acquisition probe for diagnosing a biological tissue of the present invention acquires non-invasive image without making an incision or detaching the tissue, and a high-resolution polarization component image as well as a two-dimensional cross-sectional image of the biological tissue With real-time acquisition of 2D images and real-time 2D images, 3D images can be realized in a short time, and information about cell distribution, tumor distribution, and lesion stage of biological tissues can be quantitatively objective.

In addition, since the present invention integrates the system into a probe (hand-held probe), the system can be miniaturized and manufactured in a portable manner, and quantitatively objective information on cell distribution and tumor distribution and lesion stage of biological tissues can be obtained. have.

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

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

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

3A is an explanatory diagram for schematically illustrating a configuration of the probe of FIG. 2A.

3B is an explanatory diagram for schematically illustrating a configuration of the probe of FIG. 2B.

4 is a schematic flowchart of the computer system of FIG. 2A.

FIG. 5 is a schematic flowchart of the scattering measurement step of the biological tissue epidermal region of FIG. 4.

6 is a schematic flowchart of the birefringence measurement step of the biological tissue dermal region of FIG.

7 is a S0 cross-sectional image of the animal bone.

8A is a birefringence image of the same scanning direction as in FIG. 7.

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

9A is a S ₀ image of the dorsal portion of the finger.

9B is an S ₃ image of the dorsal portion of the finger.

9C is a S ₀ image of a fingerprint portion of a finger.

9D is an S ₃ image of the dorsal portion of the finger.

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

11 is an anatomical structure diagram of the cervix.

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

12B is a table showing the distribution of patients and samples used in this example.

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

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

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 histological image of high grade epithelial tumor tissue (H-SIL).

17A is a light intensity image of normal tissue.

17B is a graph analyzing light intensity signals of normal tissues.

18A is a graph re-expressed in DOCP to analyze Stokes variable S3.

18B is a liner fitting of the DOCP graph.

19A is a graph comparing average values of DOCP changes in normal tissues and epidermis from 200 μm to H-SIL tissues.

19b is a graph comparing average values of DOCP changes in normal tissues and epidermis from 300 μm to H-SIL tissues.

<Description of the symbols for the main parts of the drawings>

10: main body 20: power supply

30: light source 70: computer system

80: transmission path 90: probe

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

130: optical splitter (BS) 140: quarter wave plate (QWP) at 22.5 °

150: reference mirror

160: diffraction grating 170: optical delay lens

180: galvanometer 185: double pass mirror

210: 1/4 wave plate at 45 ° (QWP) 220: galvanometer

230: objective lens 240: sample

245: cervical contact portion 250: polarized light splitter

260, 270: light detector 310: motion controller

Hereinafter, the configuration and operation of a polarization sensitivity-optical interference image acquisition probe for diagnosing a biological tissue according to an embodiment of the present invention will be described in detail with reference to the accompanying drawings. The objects and constructional features of the present invention will become more apparent from the accompanying drawings and the description of the detailed and preferred embodiments of the present invention described below. It should be noted that the same to similar parts in the drawings are described with the same to similar reference numerals for convenience of explanation and understanding. Note that the term 'epidermis of biological tissue' and 'dermis of biological tissue' referred to in the present invention is merely a classification based on a relative difference in depth from the surface of the skin, and does not necessarily mean 'medical epidermis' and 'medical dermis'. Should be.

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

One light from the fast, wavelength-converted low coherence source 110 passes through a linear polarizer (LP) 120 (or polarized light splitter (PBS)) to produce only light that is horizontally polarized. .

Purely horizontally polarized light passes through a beam splitter (BS) 130 and travels to a reference arm in one path and to a sample arm in the other path.

(Explanation of longitudinal scanning using reference mirror)

Light flowing to the reference arm passes through a quarter wave plate (QWP) 140 at 22.5 ° with respect to the horizontal. The light passing through the quarter wave plate (QWP) 140 at 22.5 ° creates a light path by the reference mirror 150 and is reflected and passes again through the quarter wave plate (QWP) 140 at 22.5 °. To be returned to the optical splitter (BS) 130. At this time, the light returned to the optical splitter (BS) 130 is 45 ° linearly polarized light on the basis of the horizontal component.

Light polarized horizontally through the polarizing light splitter (PBS) 250 at the light source is depicted by the Jones vector as follows. Jones vector is a concise technique that can be applied to the coherent light source as another representation of the polarized light using the electric vector, which is represented by the equation (1).

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

,

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

QWP rotated by φ about the horizontal axis is defined as in Equation 4.

Where R (Φ) is

According to Equation 4, the change in polarization due to the QWP of 22.5 °

Expressed as the Jones vector of, the horizontal and vertical components have the same amplitude and phase with 45 ° linear polarization.

 In the present embodiment, 45 ° linearly polarized light was confirmed using the QWP 140 and the polarized light splitter 250 for the light finally returned through the reference arm. Assuming that the returned light is 45 ° linearly polarized light, passing through the 45 ° QWP 140 once more results in fully polarized light on the horizontal or vertical axis. The polarized light may be confirmed by passing through the polarized light splitter 250. In this embodiment, the light passing through the polarization light splitter 250 is detected only in the vertical component to verify the 45 ° linear polarization of the reference arm.

In some cases, an optical delay line RSOD may be used. In case of using only the reference mirror, it is PS-OCT in the frequency domain that can obtain the depth direction information by Fourier transforming the spectral signal received from the detector without scanning in the depth direction.

(Description of Sample Path Configuration-Horizontal Scan)

In order to obtain a two-dimensional cross-sectional image of a sample, a transverse scanning technique is required. The horizontal scanning method is divided into a method of moving a sample and a method of moving a horizontal scanner.

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

Light that flows to the sample arm, which is one path away from the reference arm, passes through a quarter wave plate (QWP) 210 at 45 ° to the horizontal, using a light concentrator such as an objective lens 230 ) And the light is circularly polarized with respect to the horizontal component. The light of the right polarization component is scattered or retroreflected in the sample 240 and causes a change in elliptical polarization according to the optical characteristics of the sample 240. The reflected light passes through the quarter wave plate (QWP) 210 at 45 ° and returns to the optical splitter (BS) 130.

That is, light passing through the quarter wave plate (QWP) 210 at 45 ° from the sample arm is reflected by a mirror installed on the galvanometer 220 and focused on the objective lens 230, and then the sample 240 tissue is formed. Will be joined. Light reflected or backscattered from the tissue of the sample 240 returns back to the same path. 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 made a range of about 2.6mm length and the focal length of the objective lens 230 used was set to 11mm. The lateral resolution Δx is determined by the central wavelength of the light source and the numerical aperture of the objective lens as shown in Equation (6).

The horizontal axis resolution calculated by Equation 6 is 20 μm.

 In the PS-OCT, the light of the sample arm passes through the QWP 210 at 45 ° and light of circular polarization is incident on the sample. The light incident on the sample can be represented by a Jones vector and a Jones matrix as described in the reference arm. The Jones matrix equation of the sample arm is B, the average phase delay caused by the characteristics of the sample. (z, Δn, α) are added and represented. Therefore, the QWP rotated by α about the horizontal axis in the sample arm is defined as in Equation 7.

Where kzn represents the light having a refractive index n and transmitted in the depth z direction.

Therefore, the polarization change by 45 ° QWP 210 and B (z, Δn, α) in the sample arm is expressed by Equation 8, which represents the polarization state that changes when the circularly polarized light incident on the sample is backscattered back. .

As described above, in the present embodiment, in order to acquire a large number of two-dimensional images in a short time, by obtaining a Fourier domain by acquiring a spectral signal using a detector, a longitudinal axis signal is obtained and at the same time the galvanometer mirror 220 The transverse scanning was performed using. 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, thereby realizing a 3D image. However, since the synchronization means is not the core of the present invention, a detailed description thereof will be omitted.

On the other hand, according to the longitudinal direction of the biological tissue sample, another galvanometer may be included to enable longitudinal and horizontal scanning of the biological tissue sample. The galvanometer eliminates the need for patients and medical personnel to move the body, and enables accurate longitudinal movement in synchronization with each scanning cycle, allowing for a cleaner image acquisition.

The light returned from the reference arm and the sample arm, respectively, is collected in the optical splitter (BS) 130, and if the two paths are the same by the reference arm, interference occurs. The generated interference signal is again passed through a polarized light splitter (PBS) 250 to be divided into a horizontal component and a vertical component, and each component is received by the photo detectors 260 and 270 to obtain desired data. The photo detectors 260 and 270 are spectrometers constituting the detector 50 and may be formed of photodiodes.

The signal detected at the photo detectors 260, 270 is passed to the computer system 70. For more accurate light transmission, a collimator is placed in front of the photo detectors 260 and 270, i.e., the light of the horizontal and vertical components generated through the polarized light splitter (PBS) 250, respectively. Via a collimator, it may also be delivered to optical detectors 260 and 270 via an optical fiber.

2A and 2B are explanatory diagrams schematically illustrating a configuration of a polarization sensitivity-optical interference image acquisition probe for biopsy according to an embodiment of the present invention, and FIG. 3A illustrates the configuration of the probe of FIG. 2A. It is explanatory drawing for demonstrating schematically, FIG. 3B is explanatory drawing for demonstrating schematically the structure of the probe of FIG. 2B.

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

The power supply device 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 conversion low coherence light source 110 and transmits the light to the optical system of the probe 90. The average time interval at which an optical wave reaching a point in any space oscillates predictably and maintains a sinusoidal shape without phase change is called coherence time, which determines the temporal coherence of the light wave. To become a measure of Observing from a fixed point in space, the advancing light waves oscillate in the form of a sine function only during time intervals in which the phase remains constant. The spatial length of the regularly oscillating light waves until the phase changes arbitrarily is called the coherence length and is a measure of how purely the light waves are spectroscopically. The low coherence light source is a light source with a short interference distance of light waves and has been widely used in OCT recently because of its wide spectral band.

On the other hand, the transmission path 80 is composed of optical fibers and wires, and serves as a channel for transmitting light between the body 10 and the probe 90 and at the same time transmits power.

As shown in FIG. 3A, the probe 90 is an optical system forming a PS-OCT, and includes a linear polarizer (LP) 120, an optical splitter (BS) 130, and a quarter wave plate (QWP) of 22.5 °. 140, reference mirror 150, 1/4 wave plate (QWP) 210 of 45 °, galvanometer 220, objective lens 230, polarized light splitter (PBS) 250, light Detectors 253 and 257.

Low coherence of the light source unit 30 One light from the light source 110 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 proceeds to the reference arm in one path and to the sample arm in the other path.

Light flowing to the reference arm passes through a quarter wave plate (QWP) 140, ie a first quarter wave plate, at 22.5 ° with respect to the horizontal. The light passing through the quarter wave plate (QWP) 140 at 22.5 ° is reflected by the reference mirror 150 and passes through the quarter wave plate (QWP) 140 at 22.5 ° and is further divided by the optical splitter (BS). Return to 130.

The light that flows to the sample arm, the other path, passes through a quarter wave plate (QWP) 210, i.e., a second quarter wave plate at 45 ° to the horizontal, and is installed on the galvanometer 220. The light is reflected by the objective lens 230 and focused on the sample 240. The galvanometer 220 is used as an optical scanner of one axis or two axes. At this time, the light is circularly polarized based on the horizontal component. The light of the right polarization component is scattered or retroreflected in the sample 240 and causes a change in elliptical polarization according to the optical characteristics of the sample 240. Light reflected or backscattered from the tissue of the sample 240 returns back to the same path. The reflected light passes through the quarter wave plate (QWP) 210 at 45 ° and returns to the optical splitter (BS) 130.

The light returned from the reference arm and the sample arm, respectively, is collected in the optical splitter (BS) 130. If the path difference between the two lights is the same by the reference arm, an interference phenomenon occurs. The generated interference signal is again a polarized optical splitter (PBS). After passing through the (250) is divided into a horizontal component and a vertical component, the horizontal component light and the vertical component light is incident to the detection unit (50, 260, 270).

The detector 50 is composed of photo detectors 260 and 270, and the horizontal component light and the vertical component light from the polarized light splitter (PBS) 250 pass through the collimators 253 and 257 through the optical fiber. The data is entered into 50 to acquire data, and the obtained data is transferred to the computer system 70.

The computer system 70 converts and receives a signal from the detection unit 50 into a digital signal, obtains magnitude and phase values by performing Fourier transform on the vertical and horizontal components of the received digital signal, and obtains the magnitude and phase values. Get the Stokes variable value from the values. From this, the size, depth and spacing of normal cells and tumor cells in the vicinity of the epidermis of the tissue can be quantified, or the scattering coefficients changed in the vicinity of the epidermis of the tissue can be quantified. Phosphorus value can be presented and the change of tumor cells according to the lesion stage can be presented as a quantitative value by measuring the birefringence depending on the lesion stage of the normal cells and the tumor cells distributed near the biological tissue dermis.

2B is similar to FIG. 2A in that the light source unit 30 does not exist in the main body 10 and the light source unit 30 is embedded in the probe 90. FIG. 3B is a diagram corresponding to the system of FIG. 2B, showing that the light source incident on the linear polarizer (LP) 120 is generated by the probe 90 itself rather than being transmitted from the body. When the light source unit 30 is embedded in the probe 90 itself, optical noise is reduced as compared with the case where the light source unit 30 is received from the main body 10, and thus the diagnosis can be more accurately performed. With respect to the present embodiment shown in Figs. 2B and 3B, other matters are exactly the same as in the previous embodiment, and the overlapping description thereof will be omitted below.

The following briefly describes Stokes variables.

The computer system 70 obtains Stokes variable values, which represent the polarized light as a function of the observable quantities of electromagnetic waves and are a complete description of the state of polarization, as I, Q, U, V (or S0, S1, It consists of four variables of S2 and S3). These four parameters can be found using photodiodes and linear and circular polarizers. For further details on Stokes parameters, see Eugene Hecht, entitled "Optics, Fourth edition", published Addison Wesley (2002).

If the total dose of light incident on the photodetector is It, then I _{0 °} , I _{90 °} , I _{+ 45 °} and I _{-45 °} are linear at 0 °, 90 °, + 45 ° and -45 ° angles with respect to the horizontal axis, respectively. The light transmitted by the linear polarizer is defined as linear polarization, and I _rc and I _lc are the light transmitted by the 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. Four Stokes parameters representing the polarization state are defined as in Equation 9.

I = It, Q = (I _{0 °} -I _{90 °} ), U = (I _{+ 45 °} -I _{-45 °} ), V = (I _rc -I _lc )

Normalizing the Stokes variables in Equation 9 to I, Q describes the amount of polarization along the horizontal (Q = + 1) or vertical (Q = -1) axis, and U is + 45 ° (U = +1) or -45 ° (U = -1) indicates the amount of polarization. And V represents the polarized amount of right circle (V = + 1) or left circle (V = + 1).

When the light traveling with the average angular frequency ω is expressed as x component (horizontal component) and y component (vertical component) in the electromagnetic field, Equation 10 is expressed.

The Stokes variable expressed in the electromagnetic field using the equation is represented by Equation 11.

This is summarized as in Equation 12.

4 is a schematic flowchart of the computer system of FIG. 2A.

In the Fourier transform step, when the output signal of the photo detectors 260 and 270 is converted into a digital signal and the vertical component signal and the horizontal component signal are input to the computer system, the vertical component signal and the horizontal component signal are respectively The Fourier transform is performed to obtain magnitude and phase values (S100).

As a Stokes variable calculation step, a Stokes variable value is obtained by Equation 4 using the values obtained by the Fourier transform step (S150).

In the image realization step, the biotissue cross-sectional structure image and 3-D image are implemented with Stokes parameters (S200).

In the epidermal region scattering measurement step, from the image implemented in the image implementation step, cell distribution in the epidermal region of the biological tissue is measured (S250).

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

A quantitative value extraction step for the cell distribution of the epidermal area of the tissue, quantifying the change of the scattering coefficient according to the cell distribution or the size of the cell epidermis from the result of the scattering measurement step (S250) of the biological tissue epidermal area From this, a quantitative value for cell distribution of the epidermal region of the biological tissue is obtained (S350).

In addition, the quantitative value extraction step of the tumor cell lesion stage of the dermal region of the biological tissue, the birefringence of the tumor cells distributed in the dermal region of the biological tissue from the result of the dermal region birefringence measurement step according to the lesion stage Obtain a quantitative value for the change of tumor cells (S400).

FIG. 5 is a schematic flowchart of the scattering measurement step of the biological tissue epidermal region of FIG. 4.

In the S ₃ image implementation step, an S ₃ image is obtained from the Stokes parameters (S260).

In the plane fitting operation, a plane correction operation is performed to remove a shape error caused by an image surface curvature caused by an objective lens and a slope of a sample (S280). If the plane correction is performed, the size of the pixel corresponding to the cell distribution of the epidermal region of the biological tissue can be extracted.

In the step of extracting the size of the pixel, the size of the pixel corresponding to the cell distribution of the epidermal region of the biological tissue is extracted (S290), and the slope is extracted using the size value (S295).

FIG. 6 is a schematic flowchart of the biological tissue dermal birefringence step of FIG. 4.

In the S ₃ image implementation step, we obtain the S ₃ image from the Stokes parameters, where S ₃ The value may be expressed as a function having three factors, firstly the wavelength (k ₀ ), secondly birefringence (Δn), and thirdly the depth (z) of the sample (S310).

S ₃ = cos (2k ₀ Δnz)

In the arc cosine calculation step, the arc cosine takes the output signal of the S ₃ image implementation step (S320). Where S ₃ as Equation 13 The arc cosine of the value gives a linear function with slope.

In the step of extracting the slope value, when the slope is extracted from the output of the arc cosine operation step (linear fitting), the value of the birefringence for one line x is obtained (S330).

In the averaging step, the birefringence is obtained by applying it to all lines x of the image and averaged (S340).

(Check the performance of PS-OCT for biotissue diagnosis through image implementation)

The polarization state of light passing through biological tissue can be characterized by a mechanism of scattering and birefringence. Scattering changes the polarization state in any direction. As the size of the scattering particles increases, the scattering becomes more anisotropic and the incident polarization is well maintained. In general, polarized light preserves incident polarization as particle size increases and linearly polarized light is preserved by isotropic scattering rather than anisotropic scattering. When the light incident on the cell tissue is scattered several times, the effect of the single scattering is accumulated, and the polarization state is completely changed at random, thus losing the incident polarization information. If the biological tissue is composed of arbitrary structures and shapes, the polarization state of light is randomly changed.

 Birefringence is an optical property caused by a linear structural form surrounded by materials with different refractive indices. Birefringence is the polarization state of the light is changed by the difference in refractive index between the polarized light of the optical axis or the vertical axis of the material, the difference in refractive index is defined as the phase delay between the horizontal and vertical components of the light. Equation 14 is a formula of the phase delay and changes in proportion to the difference in refractive index and the moving distance.

In general, birefringence occurs in many parts of biological tissues, and hard tissues such as bones, cartilage and teeth, and soft tissues such as tendons, nerves, and muscles are typical.

(Example for Checking Birefringence of Hard Tissue)

In order to confirm the polarization sensitivity of the designed PS-OCT, birefringence was measured in hard tissues such as bones. Animal ribs, which are readily available in everyday life, were used as samples to determine whether images of specific layers with different refractive indices could be identified. The image of PS-OCT is obtained by Stokes variable (S ₀ , S ₁ , S ₂ , S ₃ ), and Stokes variable S ₀ image is a structural cross-sectional image represented by the size of backscattered light as an image of a general OCT. , S ₃ will show the degree of birefringence of the sample. The images of the acquired Stokes variable S ₃ were expressed by normalizing them by dividing them by S ₀ values. 7 shows an image of S ₀ obtained by PS-OCT, and it can be structurally confirmed that no special layer exists at a depth of about 500 μm. 8A and 8B show the acquired S ₃ image, which is a result of performing horizontal scanning in different directions 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 when horizontal scanning is performed in the direction orthogonal to the same position. Comparing the two images shows different phenomena despite measuring at the same location. As shown in FIG. 8A, the phase delay caused by birefringence is rapidly changed. In contrast, the result shown in FIG. 8B shows that the phase delay caused by birefringence is slowly changing. As a result, the structural layers that could not be identified in the structural image of S ₀ were functionally distinguishable in S ₃ . This result is expressed differently according to the arrangement of tissues formed inside the bone and the direction of the incident light, indicating that layers having different refractive indices exist. The black and white arrows in each figure indicate the reference point and indicate that the scanning positions are the same.

There was also a difference in the depth at which information could be obtained. In the S ₀ image, a clear structure can be distinguished at a depth up to about 500 μm, but S ₃ provides polarization information even at a deeper position. This means that the light intensity is significantly reduced at a depth of 500 μm or more, so that it is difficult to obtain structural information. However, S ₃ information extracted through the change in polarization is sensitively obtained even at a small light intensity.

(Example of confirming birefringence of soft tissue through image realization)

Next, the fingertips and back of the fingers were measured to verify their applicability in soft tissues such as human skin. First, the back region of the finger was measured to analyze the images of S ₀ and S ₃ (FIGS. 9A and 9B). First, the S ₀ image clearly distinguished the basement membrane and the location was measured about 250 μm deep from the epidermis.

The next S ₃ image shows the birefringence in the base layer, and the curvature of the layer cannot be seen in the S ₀ image. Its shape is similar to the folds of wrinkles on the back of the finger and shows a repeating shape periodically. This means that the birefringent tissue in the base layer is arranged along the shape of the wrinkles. 9C and 9D show the results of measuring fingerprints. Looking at the image, the basement membrane formed along the shape of the fingerprint can be distinguished remarkably. The thick line in black on the epidermis is more clearly expressed than other structures because the intensity of backscattered light in the epidermis is relatively large, and the length of one cycle of periodically repeated bends is about 400 μm. 9D is a birefringence image of the fingerprint, and the change in the polarization component that appears periodically according to the curvature of the fingerprint was confirmed. However, S ₃ image, it can be deduced that no S ₀ than it did otherwise could not clearly differentiate between the base film is it unlike the through organizations such as site present in the basal layer of the fingerprint area array form is formed in the shape similar to the fingerprint.

(Example of implementing 3D image of uterine tissue sample)

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

(Example of early diagnosis of cervix through high-speed scanning PS-OCT)

The cervical intraepithelial neoplasm is a term used to collectively describe the continuous process of cervical cancer precursor lesions occurring in the cervical epithelium. It is generally divided into three grades of CIN I, CIN II, and CIN III according to the progression of the lesion stage, starting from CIN I, a low-grade epithelial tumor, and progressing to CIN III, a high-grade epithelial tumor. There are features that can be.

11 shows the anatomical divisions of the cervix. In general, squamo-columnar junction (Squamo-Columnar Junction) is the place where the circumferential epithelium of the inner diameter canal and the squamous epithelium of the outer diameter of the inner tube can be a place of intra-epithelial tumors and cancer has a clinical significance.

Cervical epithelial tumors show the following histological characteristics as stages of progression and provide a basis for diagnosing lesions.

First, as a result of maturation and differentiation, the epithelium is divided into layers. The epithelial lesions are graded using the thickness ratio of the epithelial layer. The higher grade lesions have thicker undifferentiated cell layers and thin or disappeared mature and differentiated cell layers on the surface.

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

Third, keratinization occurs and keratinized epithelium appears as leukoplakia. It is usually most severe in high grade epithelial tumors but also in low grade epithelial tumors.

When uniformly polarized light passes through a medium having a high scattering coefficient such as a tissue, the polarization component is changed by factors of particle size, dispersion degree and density of the scattering material. The change in polarization component is regarded as the only control mechanism in the optical tablet imaging method of biological tissues and uses scattered light repeated several times. In general, the change in 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 linear polarization property is maintained longer than the circularly polarized light. If the wavelength range is larger than the scatterer, the circularly polarized light maintains the polarization component longer than the linearly polarized light.

In epithelial tumors, the ratio of cell nuclei to cells increases with the number of cells in which mitosis occurs. This will increase the scattering coefficient for the incident light and consequently affect the degree of polarization (DOP) and the depolarization of the polarization.

The polarization change of the light incident on the sample by circularly polarized light can be described using the Jones matrix or the Muller matrix. In general, the Jones determinant contains the absolute value of the phase and amplitude of light and is used to describe phenomena that are either fully polarized or without depolarization. The Mueller matrix contains the Jones determinant and is a representation to describe the interaction of polarization states and is used to explain the disappearance of polarization.

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

 Changes in DOCP basically indicate a change in polarization component that changes in tissue. However, the influence and persistence of the DOP on the change in polarization component and the area size to be detected when passing through the optics should be considered. In the system using the heterodyne detection method such as PS-OCT, DOP is maintained regardless 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 in obtaining the change of the polarization component occurring in the living tissue. This is an important basis for measuring DOCP changes in PS-OCT to distinguish the characteristics of biological tissues.

Cervical tissue sections were measured as samples to obtain clinical data of high-speed scanning PS-OCT. As shown in FIG. 12A, cervical tissues removed from the body for pathological diagnosis were measured during drug pretreatment, and patients' ages were diversified into 20s, 30s, and 40s (FIG. 12b). 50-200 cross-sectional images were obtained from one sample of each age group, and the measurements were performed at 16 sites expected to be normal tissues and 9 areas expected to be lesions. Of the sample sites measured, the lesion predicted tissue was part of the tissue removed from the patient's body in their 20s and was diagnosed by a gynecologist as a high grade epithelial tumor. The measurement region of the sample shown in the photograph of FIG. 12A was collected at 3, 6, 9, and 12 o'clock, which are frequent areas of epithelial tumor, and the 12 o'clock indicates the uterine entrance direction.

First, the tissue extracted from the cervix of the patient without a lesion was scanned to realize an image of 2 mm × 2.3 mm (vertical axis × horizontal axis), and then compared with a histology image (FIG. 13C). Since the sample used did not undergo mitosis into normal tissue, there is a boundary between the epithelial layer (e) and the basal layer (s), which can be confirmed by histological images.

In the light intensity image (FIG. 13a) obtained using PS-OCT, the boundary line between the epidermal layer and the base layer was clearly distinguishable, and the depth of the epidermal layer was measured to be about 200 μm. 13B shows an image showing the disappearance of polarization. As light reaches the epidermis of the uterine tissue from the basal layer, no large polarization change is seen, and it can be seen that the polarization change occurs in the portion below the basal layer.

Figures 14a, b and 15a, b compare the light intensity image and the polarization image of different normal tissues. The light intensity image shows different phenomena despite normal tissue. In the light intensity image of FIG. 14A, the base layer was clearly distinguishable, but the epidermal layer and the base layer could not be distinguished from the image of FIG. 15A. In general, as the lesion of the cervical epithelial tumor progresses, the thickness of the epithelial layer becomes thinner or the distinction between the epithelial layer and the basal layer disappears. However, since there are no basement membranes in normal tissues without lesions, and they are formed differently according to individual characteristics, these differences in the same normal tissues may be a clear basis for the presence of basement membranes in judging cervical lesions. It means you can't.

Next, images of normal tissues and high grade epithelial tumor tissue (H-SIL) were implemented and compared (FIG. 16). Since high grade epithelial tumors increase the ratio of cells and nuclei and increase mitosis, the scattering action is increased. The depth of transmitted light is expected to be shorter than that of normal tissue. By comparing the light intensity images of FIGS. 16A to 16D, the signal can be obtained up to a depth of 500 μm or more in the case of normal tissue, but the signal in the base layer region is weakly detected in the H-SIL image. . These results can be judged that the change in penetration depth is due to the effect of greater scattering action when the light in the H-SIL passes through the normal tissue. However, in the light intensity image of H-SIL, the distinction between layers is made in the lower part of the epidermis and the difference in the depth of transmitted light is not clearly distinguished from normal tissue. It is not enough to distinguish between tissue and normal tissue.

Cervical intraepithelial tumors develop and progress within a 400 μm depth from the epidermis. As a result of measuring the cervical tissue with the designed PS-OCT system, the longitudinal depth of the acquired image was able to identify the basement membrane, include the extent to which the lesion could occur, and observe polarization change information (FIGS. 16B and FIG. 16e). However, the lesion of cervical epithelial tumor is normal tissue and H by using only the light intensity image that can observe only the structural features because the tissue changes and progresses as a cell unit as shown in the histological images of FIGS. 16C and 16F. Difficulties in identifying SIL organizations. Therefore, in order to distinguish the characteristics between the normal tissue and the lesion tissue more accurately, not only the structural image comparison but also the additional analysis of the change of the polarization component is required.

Therefore, as another example, the polarization component signals of the acquired image were analyzed to determine the exact state of the cervix. The base layer thickness was measured by analyzing the signal of the light intensity image, and the DOCP change was confirmed using the information of Stokes parameters. FIG. 17A shows a 2D image expressed in light intensity, and the presence of the basement membrane was confirmed through the contrast of the light intensity intensities. The total depth that can be confirmed by the intensity of the signal on the image was about 600 μm, and since the signals below it are expressed at the same level as the air region, it can be determined that there is little amount of light detected. In order to measure the exact position of the basement membrane identified in the image, the longitudinal signal was expressed according to the light intensity. The highest signal at the surface of the sample tissue was measured and the second largest signal at the basement membrane was detected. The exact location of the basement membrane measured using this was about 250 μm deep from the epidermis, indicating the epidermal thickness of the sample tissue.

FIG. 18 is a graph for analyzing a Stokes variable signal, and illustrates a graph in which an interference signal detected by a detector is obtained through a demodulation process and a signal processing process, and is re-expressed in DOCP. FIG. 18A shows that DOCP in the epidermis is represented by 1, and the change in depth is relatively expressed through the generalization process. It can be seen that DOCP decreases based on a depth of about 330 μm, which is caused by a small amount of scattered light. Since lesions occurring in the cervix generally progress near the epidermal layer and the basement membrane, the rate of DOCP reduction was calculated by selecting signals from the epidermis indicated by the dotted line to the depth of 300 μm. The reduction rate obtained through the linear fitting (first linear function) can determine how quickly the polarization component disappears as the light of the incident circular polarization proceeds in the depth direction (FIG. 18B). In this figure, the reduction rate of polarization indicates the degree of disappearance of the polarization at 1 mm shift, showing a result of 0.76204 depolarization / mm. This value is a sample of normal tissue, and if lesions have changed cells, the polarization disappears more quickly depending on the depth, so large linear fitting results can be predicted.

Next, the average value of DOCP change between normal tissue and H-SIL tissue was compared (FIG. 19). The average value was extracted after linear fitting was performed in DOCP from the epidermis to the depth of 200 ㎛ and 300 ㎛. First of all, the change in DOCP slope according to depth shows that the DOCP slope value increases as the depth range analyzed in both normal tissue and H-SIL tissue increases. This means that the polarized light increases as the light incident by circularly polarized light passes through the cervical tissue and proceeds in the depth direction. Considering that the light intensity is maintained at a depth from FIG. 18 to about 350 μm, signal analysis in the 200 μm and 300 μm depth ranges from the epidermis presented in this embodiment is independently determined without affecting the magnitude of the light intensity. It can be a basis for doing so. Comparing the difference between DOCP slopes for normal tissues and H-SIL tissues, it can be seen that DOCP slopes in H-SIL are larger than those in normal tissues. Can be observed. This result is consistent with the expected result of DOCP change as the scattering coefficient increases as the cervix progresses to higher grade epithelial tumors such as H-SIL. Therefore, the change information of DOCP extracted from PS-OCT may be a parameter for determining the presence of tumors in the cervical epithelium.

Cervical cancer is a cancer that occurs in women so that it accounts for 42.1% of female cancer patients, but due to the nature of the cervical cancer, it is difficult to diagnose it early. Therefore, in this example, PS-OCT was used to investigate the possibility of early diagnosis of cervical cancer.

PS-OCT is a new optical imaging technique that uses non-invasive methods to obtain high-resolution images of living tissues using light that is harmless to the human body. It is a technique for acquiring and additionally obtaining a change in the polarization component. In this embodiment, a PS-OCT system based on a high speed scanning optical delay line (RSOD) technique was designed and fabricated using a light source in the near infrared region (1296 nm). The system takes 18 seconds to acquire and store data with a vertical resolution of 18 µm, and a 3D image with a relatively small error range can be realized by using a synchronized signal.

First, the change of DOCP in high scattering medium (intralipid) was confirmed to confirm the performance of the polarization sensitive system. 20% intra lipid solution and distilled water were mixed to make 2.5%, 5% and 10% mixed solution, and the DOCP values were extracted according to the concentration. It was confirmed that the increase.

Next, we measured the structural shape and birefringence images using the bones of animals. As a result of acquiring 2D cross-sectional images using the bones of animals that can be easily obtained from the surroundings, a signal capable of confirming the structure at a depth of about 500 μm was obtained. The birefringence image was then acquired to observe the change in polarization component. In the same position, the scanning direction was alternately observed to observe the change in polarization component, and the birefringence images appearing differently showed that the birefringent structure inside the bone could be detected using PS-OCT.

The back and fingerprint areas of the human finger were measured to obtain structural information and changes in polarization components in soft tissues. In the light intensity image, the structural shape in the basal layer region was observed and the shape of the basal membrane formed by the wrinkles of the epidermis was clearly identified. In addition, the birefringence images showed the arrangement of layers with different refractive indices in the epidermis.

Finally, DOCP changes of lesion and normal tissue of cervix were extracted. For the histological diagnosis of lesions, the rate of change of DOCP was measured using sections of tissues removed from the patient's body and divided into normal and high grade epithelial tumor (H-SIL) tissues. The rate of change was largely measured. As a result, the histological characteristics of the normal tissue and the epithelial tumor could be distinguished by the change of the polarization component obtained from the cervix.

In addition, based on a deeper understanding of the scattering of light incident on biological tissues, we gather various opinions from gynecologists and collect and quantify normal and staged epithelial tumor data from a large number of cervical biopsies. Changes in DOCP, obtained using OCT, allow early diagnosis of cervical lesions.

In the above, the best embodiments have been disclosed in the specification with reference to the drawings, but for the purpose of easily understanding the present invention. Therefore, the present invention is not limited to the above embodiments, and the true technical protection scope of the present invention should be defined by the technical spirit of the appended claims. In addition, one of ordinary skill in the art will appreciate that various modifications and equivalent other embodiments can be derived from the purpose and configuration of the present invention, but this is also within the scope of the present invention.

For example, the epidermis of biological tissue and the dermis of biological tissue, which are referred to in the present invention, are merely classified according to the relative difference in depth from the skin surface, and necessarily mean 'medical epidermis' and 'medical dermis'. Note that this is not the case.

Claims (17)

A probe and a main body, the probe detects an optical signal including information of the biological tissue, the body is a polarization sensitivity-optical interference consisting of a computer system for measuring and analyzing the biological tissue from the optical signal received from the probe In the polarization sensitivity-optical interference image acquisition probe constituting the probe in an imaging system, The polarization sensitivity-optical interference image acquisition probe is a horizontal axis scanning using the optical scanner of the biological tissue, using a Fourier domain technique capable of vertical scanning without the movement of the mirror, scattering light by the biological tissue and A polarization sensitivity-optical interference imaging probe for biopsy, characterized in that the optical system consisting of a polarization sensitivity-optical coherence tomography (PS-OCT) capable of detecting birefringence. The method of claim 1, The probe, A linear polarizer (LP) for generating one light emitted from the light source unit using only light polarized in a horizontal direction; An optical splitter (BS) for transmitting the purely horizontally polarized light output from 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 the light flowing from the linear polarizer to the reference arm and returning the light reflected from the reference mirror to the optical splitter; The light flowing from the linear polarizer to the sample arm is transmitted to the mirror mounted on the galvanometer, and the light reflected by the mirror mounted on the galvanometer is reflected or reverse scattered from the sample tissue to return to the optical splitter. A second quarter wave plate QWP; The light transmitted from the second quarter wave plate (QWP) to the galvanometer focuses the light reflected by the mirror installed in the galvanometer to be incident on the sample tissue, and is reflected or reverse scattered from the sample tissue. An objective lens for transmitting the reflected light back to a mirror installed in a galvanometer; A polarization optical splitter (PBS) in which the interference signal generated from the light returned from the reference arm and the sample arm and collected in the optical splitter is divided into a horizontal component light and a vertical component light; Photodetectors for converting the horizontal component light and the vertical component light emitted by the polarization optical splitter into respective electrical signals; Probe for obtaining a polarization sensitivity-optical interference image for biopsy, characterized in that consisting of. 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 the light flowing from the linear polarizer to the reference arm and returning the light reflected from the reference mirror to the optical splitter; The light flowing from the linear polarizer to the sample arm is transmitted to the mirror mounted on the galvanometer, and the light reflected by the mirror mounted on the galvanometer is reflected or reverse scattered from the sample tissue to return to the optical splitter. A second quarter wave plate QWP; A polarization optical splitter (PBS) in which the interference signal generated from the light returned from the reference arm and the sample arm and collected in the optical splitter is divided into a horizontal component light and a vertical component light; Photodetectors for converting the horizontal component light and the vertical component light emitted by the polarization optical splitter into respective electrical signals; Probe for obtaining a polarization sensitivity-optical interference image for biological examination, characterized in that it comprises at least. The method of claim 3, Light flowing from the second quarter wave plate (QWP) to the sample arm is transmitted to the galvanometer to focus light reflected by a mirror installed in the galvanometer to enter the sample tissue. An objective lens for transmitting the reflected or backscattered light from the sample tissue back to a mirror mounted on a galvanometer; Probe for obtaining polarization sensitivity-optical interference image for biopsy, characterized in that it further comprises. The method according to any one of claims 1 to 3, According to the longitudinal direction of the biological tissue sample, the polarization sensitivity-optical interference image acquisition probe for biopsy, characterized in that it further comprises a galvanometer of one axis, to enable longitudinal and horizontal scanning of the biological tissue sample . The method according to any one of claims 1 to 3, And a light source of the light source unit is a high speed wavelength conversion low coherence light source. The method according to claim 2 or 3, The first quarter wave plate is a quarter wave plate of 22.5 ° polarization sensitivity-optical interference imaging probe for biopsy. The method according to claim 2 or 3, The second quarter wave plate is a quarter wave plate of 45 ° polarization sensitivity-optical interference image acquisition probe for biopsy. The method according to claim 2 or 3, The light detector is a polarization sensitivity-optical interference image acquisition probe for biopsy, characterized in that consisting of a photodiode. The method according to claim 1 or 2, And the computer system obtains quantitative values of tumor tissue distribution and lesion stage by measuring light scattering and birefringence of the biological tissue. The method of claim 10, The computer system is a polarization sensitivity-optical interference image acquisition probe for biopsy, characterized in that to obtain a quantitative value for tumor tissue distribution and lesion stage by measuring the light scattering of the epidermal layer of the biological tissue. The method of claim 10, And the computer system obtains quantitative values for tumor tissue distribution and lesion stage by measuring the birefringence of the dermal layer of the biological tissue. The method according to claim 1 or 2, The computer system converts and receives a signal from detectors into a digital signal, obtains magnitude and phase values by performing Fourier transform on the received digital signal for vertical and horizontal components, respectively, and uses Stokes from values obtained by the Fourier transform. A probe for obtaining polarization sensitivity-optical interference image for biopsy, characterized in that obtaining a parameter value. The method of claim 2, And a signal from the photo detectors is transmitted to a computer system of the main body via a wire of a transmission path. And a main body including a light source unit for generating and transmitting a light source to a probe, and a computer system for converting an output signal received from the probe into a digital signal, receiving and arithmetic processing to measure scattering and birefringence of light by biological tissue. 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 vertical component signal and a horizontal component signal are input to the computer system, Fourier transforms are performed on the vertical component signal and the horizontal component signal to obtain magnitude and phase values. Fourier transform step; A Stokes variable calculation step of obtaining a Stokes variable value using the values obtained by the Fourier transform step; An image realization step of implementing a biopsy cross-sectional structure image and a 3-D image using the Stokes parameters obtained in the Stokes variable calculation step; Epidermal region scattering measurement step of measuring the cell distribution in the epidermal region of the biological tissue from the image implemented in the image implementation step; A dermal region birefringence measuring step of measuring the birefringence in the dermal region of the biological tissue from the image implemented in the image forming step; Quantification of the cell distribution to quantitatively change the scattering coefficient according to the cell distribution or cell size of the biological tissue epidermis from the results of the scattering measurement step of the epidermal region, and to obtain a quantitative value for the cell distribution of the epidermal region of the tissue. Phosphorus value extraction step; Quantitative value of the tumor cell lesion step to obtain a quantitative value for the change of tumor cells according to the lesion stage by detecting the birefringence of the tumor cells distributed in the dermal region of the biological tissue from the result of the dermal site birefringence measurement step Extraction step; A method of driving a polarization-sensitized optical interference imaging system for a biopsy, characterized in that it comprises a. The method of claim 15, The epidermal area scattering measurement step S ₃ S ₃ image implement obtaining images from the Stokes parameters; Plane fitting to extract the size of the pixel corresponding to the cell distribution in the epidermal area of the tissue tissue by performing a planar correction process to remove the shape curvature caused by the objective lens and the slope of the sample Operation step; Extracting a size of a pixel having a size value of a pixel corresponding to a cell distribution of the epidermal region extracted from the planar correction operation; A method of driving a polarization-sensitized optical interference imaging system for a biopsy, characterized in that it comprises a. The method of claim 15, The dermal region birefringence measuring step Obtain the S ₃ image from the Stokes parameters, S ₃ = cos (2k ₀ Δnz) S ₃ image implementation step as a function having three factors, such as wavelength (k ₀ ), birefringence (Δn), the depth of the sample (z); An arc cosine operation step of taking an arc cosine of the output signal of the S ₃ image implementation step; A slope value extraction step of obtaining a birefringence value for one line x by linearly fitting a slope from an output of the arc cosine operation step; An averaged step of obtaining birefringence by applying the S ₃ image implementation step to the gradient value extraction step to all lines x of the image and averaging them; A method of driving a polarization-sensitized optical interference imaging system for a biopsy, characterized in that it comprises a.