JP4594208B2 - Functional confocal image display device - Google Patents

Description

  The present invention relates to a functional confocal image display device capable of measuring and displaying the shape of a subject inside a living body and measuring and displaying biological information of a part corresponding to the measured shape of the subject. .

  2. Description of the Related Art In recent years, attention has been focused on the use of optical coherence tomography (Optical Coherence Tomography) technology as a device and method that can easily and non-invasively measure the inside of a living body. According to this optical coherence tomography technique, it is possible to form an image in the micron order in the close region by using near infrared light, more specifically, near infrared low interference light. This optical coherence tomography technology is being put into practical use particularly in the field of intravascular catheters and endoscopes. For example, Patent Document 1 below combines optical coherence tomography technology and confocal microscopy technology. A tomographic imaging apparatus is disclosed.

By using this tomographic imaging apparatus, the doctor can perform an examination based on the tomographic image in the depth direction of the affected part using optical coherence tomography with excellent depth of penetration by switching operation, and has excellent resolution. The affected area can be examined at the cellular level using a confocal microscope. For this reason, for example, cancer, tumor, etc. can be discovered at an early stage, an accurate and quick diagnosis is attained, and the burden on the patient can be reduced. On the other hand, since these techniques can be accurately diagnosed in a non-invasive manner, the use of these techniques for eye disease diagnosis has been actively studied.
JP-A-2005-230202

  By the way, in the tomographic imaging apparatus described in Patent Document 1 described above, although a tomographic image of an affected area can be obtained satisfactorily, a doctor can obtain only information on a cross-sectional shape obtained from the tomographic image. Therefore, in diagnosing the medical condition and its progress, it depends on the experience and knowledge of the doctor, and the burden on the doctor increases. Further, in the diagnosis of eye diseases, particularly in the diagnosis of eye diseases in the vicinity of the retina of the eyeball, it is necessary to observe a very fine region, which increases the burden on the doctor. Further, for example, in an eye disease in which a photoreceptor cell is necrotized such as glaucoma, there is a possibility that an accurate diagnosis cannot be made only by information on a cross-sectional shape obtained from a tomographic image. Therefore, in particular, in the diagnosis of eye diseases, the practical application of a measuring device that can provide doctors with more accurate information, specifically, biological information such as blood volume and oxygen saturation, is desired. ing.

  The present invention has been made in order to solve the above-described problems, and its purpose is to provide a functional confocal image capable of observing in detail the internal state of a living body in a non-invasive manner using biological information associated with the metabolism of the living body. It is to provide a display device.

  A feature of the present invention is that a functional confocal image display device is operated by a user and outputs various signals based on instructions from the user, and at least based on a predetermined drive signal supplied from the controller. A non-modulated light generating means for generating near-infrared light or visible light having one specific wavelength, and a secondary drive signal obtained by modulating a predetermined drive signal supplied from the controller, and based on the secondary drive signal A light emitting portion having modulated light generating means for generating near infrared light having at least two specific wavelengths, and a spot shape emitted from the light emitting portion via a first pinhole plate having a predetermined hole diameter Near-infrared light or visible light transmitted through the subject and the near-infrared light reflected by the subject or visible light transmitted through the subject Alternatively, an optical path separating unit that separates from the optical path of visible light, and an optical scanning unit that scans near-infrared light or visible light transmitted from the optical path separating unit toward the subject in an in-plane direction with respect to the subject. A light receiving means for receiving near infrared light or visible light separated by the optical path separating section through a second pinhole plate having a predetermined hole diameter, and a near infrared light or visible light received by the light receiving means. A surface shape information calculating means for calculating surface shape information representing a shape in the vicinity of the surface of the subject based on the amount of near infrared light or visible light generated by the non-modulated light generating means of the light emitting portion, Demodulating means for demodulating the secondary drive signal of the near-infrared light generated by the modulated light generating means of the light emitting unit out of the near-infrared light or visible light received by the light-receiving means into the predetermined drive signal; Biological information calculation for calculating biological information of the subject accompanying metabolism of the living body based on the amount of near infrared light generated by the modulated light generating means of the light emitting unit and the amount of near infrared light demodulated by the demodulating means A light detection unit comprising: means; and image data generation means for generating visible image data based on the surface shape information calculated by the surface shape information calculation means and the biological information calculated by the biological information calculation means; Based on the image data generated by the light detection unit, a surface shape image of the subject, a biological information image of the subject, or a composite image obtained by combining the surface shape image of the subject and the biological information image is displayed. And a display unit. Here, the near-infrared light emitted from the light emitting unit may have a wavelength of 600 nm to 1000 nm, for example.

  In this case, the display unit matches the position specified by the surface shape image of the subject with the position specified by the biological information image of the subject, and displays the surface shape image and the biological information image. It is good to display the synthesized composite image. In this case, for example, the biological information calculated by the biological information calculation unit of the light detection unit is one of blood volume, blood flow volume, blood flow change, and oxygen saturation in the blood vessel of the subject. It is good to be. Further, in this case, the subject may be, for example, the fundus of an eyeball.

  Further, the modulated light generating means of the light emitting unit includes spread spectrum modulation means for generating a secondary drive signal by performing spread spectrum modulation on a predetermined drive signal supplied from the controller, and each light source includes the secondary drive signal. And a light synthesizing means for optically synthesizing near-infrared light having different specific wavelengths generated simultaneously, and the demodulating means of the light detection unit is included in the near-infrared light received by the light-receiving means The secondary drive signal may be demodulated by despreading the predetermined drive signal.

  Further, the modulated light generating means of the light emitting unit includes frequency division multiplex modulation means for generating a secondary drive signal by frequency division multiplex modulation of a predetermined drive signal supplied from the controller, and each light source includes the secondary light Light synthesizing means for optically synthesizing near-infrared light having different specific wavelengths generated all at once based on the drive signal, and the demodulating means of the photodetecting unit is adapted to the near-infrared light received by the light-receiving means. The secondary drive signal included may be demodulated into the predetermined drive signal.

  According to these, the functional confocal image display device according to the present invention operates as follows. That is, when the controller is operated by the user, the non-modulated light generating means of the light emitting unit emits near-infrared light or visible light, and the modulated light generating means of the light emitting unit emits near wavelengths of different specific wavelengths from each light source. Infrared light is emitted. At this time, the emitted near-infrared light or visible light passes through the first pinpoint plate and is emitted as a spot-like point light source. The optical path separation unit can employ, for example, a beam splitter, a dichroic mirror, etc., and transmits near-infrared light or visible light emitted from the light emitting unit toward the subject, that is, the fundus of the eyeball, The near-infrared light or visible light reflected from the fundus is optically separated from the transmitted near-infrared light or visible light path and emitted toward the light detection unit. At this time, the optical scanning unit can appropriately change the optical axis of the near-infrared light or visible light that passes through the optical path separation unit and travels toward the subject. The arrival point in the specimen can be scanned in the in-plane direction. Thereby, near infrared light or visible light can be reflected within the measurement region of the subject.

  The light detection unit receives near-infrared light or visible light generated by the non-modulated light generating means and reflected by the subject, and represents the surface shape of the subject based on the received near-infrared light or visible light amount Surface shape information is calculated. The light detection unit demodulates the near-infrared light generated by the modulated light generation unit and reflected by the subject, and converts the near-infrared light amount generated by the modulated light generation unit and the demodulated near-infrared light amount. Based on the biological information of the subject, for example, blood volume, blood flow volume, blood flow change, oxygen saturation, and the like are calculated.

  At this time, since the light detection unit receives near-infrared light or visible light through the second pinhole plate, it is in a specific position of the subject (fundus), more specifically, inside the subject (fundus). Only near infrared light or visible light appropriately reflected from the arrival point of the incident light can be detected. In other words, it is possible to detect only near-infrared light or visible light necessary for calculating surface shape information and biological information by eliminating near-infrared light or visible light diffusely reflected inside the subject (fundus).

  The light detection unit generates visible image data based on the calculated surface shape information and biological information. On the other hand, the display unit displays a surface shape image based on the calculated surface shape information, a biological information image based on the calculated biological information, or a composite image obtained by combining these surface shape images and the biological information image. At this time, the display unit can display a composite image obtained by synthesizing the surface shape image and the biological information image by matching the position specified by the surface shape image with the position specified by the biological information image. .

  For this reason, the functional confocal image display device according to the present invention can calculate the surface shape and biological information of the subject (fundus), and displays the calculated surface shape and biological information on the display unit. be able to. Therefore, more accurate information can be provided to the doctor. In particular, for example, an image of biometric information of a part matching the same part is synthesized (overlaid) and displayed on a part observed by a doctor using a surface shape image displayed. it can. Thereby, the doctor can diagnose correctly based on the displayed surface shape and biological information. In addition, as biological information necessary for diagnosis of a medical condition, for example, blood volume, blood flow volume, blood flow change, and oxygen saturation can be easily calculated and displayed. Diagnosis can be made very easily and accurately.

  Further, since the modulated light generating means of the light emitting section has at least two light sources and can emit near infrared light having different specific wavelengths, a suitable specific wavelength is selected when calculating biological information. Can also be emitted. Thereby, biometric information can be calculated more accurately and a doctor's diagnosis can be more suitably assisted. In this case, near infrared light can be emitted simultaneously (simultaneously) based on the modulated secondary drive signal. Then, the light combining means (for example, an optical fiber or the like) can optically combine near-infrared lights having different specific wavelengths emitted simultaneously (simultaneously) and emit the light to the optical path separation unit. The near-infrared light separated by the optical path separation unit is demodulated by the light detection unit, and biological information is calculated.

  In this way, near-infrared light having a plurality of specific wavelengths is emitted at the same time, and each near-infrared light reflected by the subject (fundus) is detected, so that the state change with time can be made extremely small and the living body can be reduced. Information can be calculated. That is, for example, when calculating oxygen saturation in an artery or arteriole as biological information, it is necessary to calculate it based on the amount of near-infrared light that varies with the pulse wave of blood flow. At this time, since the pulse wave changes very quickly, for example, when near-infrared light is emitted sequentially, the amount of near-infrared light detected by the light detection unit corresponding to the order of emission of each near-infrared light, respectively, This is due to different pulse wave conditions. For this reason, the accuracy of the calculated biological information may be inferior. On the other hand, when near-infrared light is emitted at the same time, the amount of near-infrared light detected by the light detection unit depends on substantially the same pulse wave state, and the calculated biological information is accurately calculated. Can do. Therefore, the biological information can be calculated more accurately, and the doctor's diagnosis can be more suitably assisted.

  Furthermore, another feature of the present invention is that a confocal aperture movable in the optical axis direction of near-infrared light or visible light separated by the optical path separation unit is adopted as the second pinhole plate. is there. According to this, by moving the confocal aperture in the optical axis direction, it is possible to easily select near-infrared light or visible light that reaches the light detection unit. Therefore, more preferably, near-infrared light or visible light irregularly reflected inside the subject can be eliminated, and as a result, the accuracy of the calculated surface shape and biological information can be further increased. Thereby, a more accurate surface shape and biological information can be provided, and a doctor's diagnosis can be assisted more suitably.

a. First Embodiment Hereinafter, a first embodiment of the present invention will be described in detail with reference to the drawings. FIG. 1 schematically illustrates the configuration of a functional confocal image display apparatus M that measures the inside of a living body, for example, the surface shape of the fundus and biological information, according to the first embodiment of the present invention. The functional confocal image display apparatus M includes a light emitting unit 1, an optical path separating unit 2, an optical scanning unit 3, a light detecting unit 4, and a display unit 5. Furthermore, the functional confocal image display apparatus M includes a controller 6 whose main component is a microcomputer including a CPU, a ROM, a RAM, and the like.

  As shown in FIG. 2, the light emitting unit 1 includes a light generator 11 and two light generators 12 for generating near-infrared light or visible light having different specific wavelengths. The light generator 11 as non-modulated light generating means emits near infrared light or visible light having a specific wavelength without being modulated. In the following description, near infrared light or visible light generated by the light generator 11 is also referred to as non-modulated light. In order to emit this non-modulated light, the light generator 11 includes a light source driver 111 that acquires a drive signal supplied from the controller 6. The light source driver 111 drives (emits light) the light source 112 based on the drive signal acquired from the controller 6. The light source 112 includes, for example, a near infrared light emitting element such as a laser diode (LD) or a super luminescence diode (SLD). Accordingly, the light source 112 emits near infrared light or visible light having a specific wavelength.

  Here, the specific wavelength of near-infrared light emitted from the light source 112 can be appropriately selected from 600 nm to 1000 nm, and for example, 785 nm and 817 nm may be adopted. Further, the specific wavelength of visible light emitted from the light source 112 can be appropriately selected from 400 nm to 700 nm, and for example, 488 nm may be adopted. For this reason, in this embodiment, it demonstrates as what emits visible light which has a specific wavelength of 488 nm.

  The light generator 12 serving as the modulated light generating means emits light by performing spread spectrum modulation on near-infrared light having a specific wavelength. In the following description, near infrared light generated by the light generator 12 is also referred to as modulated light. In order to emit this modulated light, the light generator 12 generates a spread code sequence for generating, for example, a PN (Pseudorandom Noise) sequence consisting of “+1” and “−1” having a length of 128 bits as a spread code sequence. A container 121 is provided. The spreading code sequence generator 121 generates, for example, a Hadamard sequence, an M sequence, or a Gold code sequence as a PN sequence.

  The Hadamard sequence, the M sequence, or the Gold code sequence is the same as that generally used for spread spectrum modulation, and thus a detailed description of the generation method is omitted, but a brief description is given below. Keep it. The Hadamard sequence is a sequence obtained by extracting each row or each column of the Hadamard matrix composed of “+1” and “−1”. The M series uses a shift register in which n stages of 1-bit registers for storing a state of “0” or “+1” are arranged, and an exclusive OR of a value fed back from the middle of the shift register and a value in the last stage Is a binary sequence obtained by connecting to the first stage. However, in order to make this binary sequence a PN sequence, level conversion is performed to convert the value “0” into “−1”. The Gold code sequence is basically a code sequence obtained by preparing two types of M sequences and adding them. For this reason, the Gold code sequence is a sequence that can significantly increase the number of sequences as compared to the M sequence. As a feature of these sequences, different sequences have the property of being orthogonal to each other, and by performing a product-sum operation, “0”, that is, the correlation becomes “0” except for self. .

  As described above, the PN sequence generated by the spread code sequence generator 121 is output to the controller 6 and also to the multiplier 122. The multiplier 122 calculates the product of the drive signal supplied from the controller 6 (hereinafter referred to as the primary drive signal in particular) and the PN sequence supplied from the spread code sequence generator 121. Thereby, the primary drive signal can be subjected to spread spectrum modulation. Then, the multiplier 122 supplies the light source driver 123 with the secondary drive signal subjected to the spread spectrum modulation.

  The light source driver 123 drives (emits light) the light source 124 based on the secondary drive signal supplied from the multiplier 122. Similarly to the light source 112 described above, the light source 124 is also composed of, for example, a near infrared light emitting element such as an LD or SLD, and emits near infrared light having a specific wavelength. Here, the specific wavelength of the near-infrared light emitted from the light source 124 can be appropriately selected from, for example, a wavelength range of about 600 nm to 1000 nm. For this reason, in this embodiment, one light source 124 emits near infrared light having a specific wavelength of, for example, 780 nm, and the other light source 124 emits near infrared light having a specific wavelength of, for example, 830 nm. It will be explained as a thing.

  Here, in the present embodiment, the light emitting unit 1 is composed of a light generator 11 that generates unmodulated light and two light generators 12 that generate modulated light. In other words, the light emitting unit 1 Is configured to generate near infrared light or visible light having three specific wavelengths. However, the number of light generators constituting the light emitting section 1, that is, the number of specific wavelengths of emitted near-infrared light or visible light is not limited to these. Accordingly, for example, the light emitting unit 1 is configured to generate two or more non-modulated lights by two or more light generators 11, or three or more modulated lights by three or more light generators 12. Needless to say, the present invention can be configured and implemented. Thus, by providing a large number of the light generators 11 and 12, the surface shape of the fundus oculi to be described later can be calculated in more detail, and sufficient quantitativeness can be ensured in the calculation of oxygen saturation as biological information to be described later.

  Returning again to FIG. 1, near-infrared light or visible light emitted from the light emitting unit 1, that is, unmodulated light or modulated light, enters the optical fiber H <b> 1 as a light combining unit. In this case, when predetermined modulated light is emitted from each light generator 12 at the same time, the respective modulated lights are combined by being conducted through the optical fiber H1. Then, the modulated light that has been conducted through the optical fiber H1 is collected by the condenser lens S1, passes through the pinhole plate A1, and then reaches the optical path separation unit 2.

  The optical path separation unit 2 transmits unmodulated light or modulated light that has passed through a pinhole having a predetermined hole diameter provided in the pinhole plate A1 toward the fundus and transmits unmodulated light or modulated light reflected by the fundus. It separates from the optical path of the unmodulated light or modulated light. For this reason, the optical path separation unit 2 includes a beam splitter 21 as shown in FIG. The beam splitter 21 is disposed so as to be inclined by, for example, 45 ° with respect to the optical axis of the unmodulated light or the modulated light that has passed through the pinhole plate A1. In the first embodiment, the beam splitter 21 is used as the optical path separating unit 2. However, near-infrared light or visible light emitted from the light emitting unit 1 is transmitted and reflected from the fundus. Needless to say, the present invention can be implemented by employing, for example, a dichroic mirror that can optically separate near-infrared light or visible light from the optical path of transmitted near-infrared light or visible light.

  The optical path separation unit 2 configured by the beam splitter 21 operates as follows. That is, the non-modulated light or the modulated light that has passed through the pinhole of the pinhole plate A1 passes through the beam splitter 21 and reaches the fundus through the condenser lens S2 and the optical scanning unit 3 described later. Then, the non-modulated light or the modulated light reflected from the fundus is emitted (reflected) by the beam splitter 21 toward the pinhole plate A2 provided on the light detection unit 4 side described later.

  The optical scanning unit 3 optically scans the non-modulated light or the modulated light transmitted through the beam splitter 21 of the optical path separating unit 2 and condensed by the condenser lens S2 in the vertical and horizontal directions of the fundus. For this reason, the optical scanning unit 3 is composed of a pair of galvanometer mirrors 31 and 32 as shown in FIG. The galvanometer mirror 31 is disposed on the optical axis of the non-modulated light or the modulated light collected by the condenser lens S2, and is provided to be rotatable around a rotation axis extending in a predetermined direction. The galvanometer mirror 31 changes the optical axis, that is, the reflection direction, of the unmodulated light or the modulated light transmitted through the condenser lens S2 by rotating around the rotation axis. The galvanometer mirror 32 is disposed on the optical axis of the unmodulated light or the modulated light reflected by the galvanometer mirror 31 and is provided to be rotatable about a rotation axis orthogonal to the rotation axis of the galvanometer mirror 31. . The galvanometer mirror 32 changes the optical axis, that is, the reflection direction of the unmodulated light or the modulated light reflected by the galvanometer mirror 31 by rotating about the rotation axis.

  With this configuration, the non-modulated light or the modulated light collected by the condenser lens S2 is reflected by the galvanometer mirrors 31 and 32 that rotate about the respective rotation axes, so that the optical axis, that is, the traveling direction of the reflected light is changed. Be changed. In this way, by appropriately changing the traveling direction of the reflected light, the arrival point of the non-modulated light or the modulated light that reaches the fundus via the objective lens T is moved in the in-plane direction within a predetermined measurement region on the fundus. Can do. Further, the non-modulated light or the modulated light reflected from the fundus is transmitted through the objective lens T, then reflected by the galvanometer mirrors 31 and 32, and transmitted through the condenser lens S2 to reach the optical path separating unit 2, that is, the beam splitter 21. To do.

  The light detection unit 4 detects unmodulated light or modulated light that is reflected by the optical path separation unit 2 and arrives through the pinhole plate A2, the condensing lens S3, and the optical fiber H2. The light detection unit 4 outputs an image signal representing the state of the fundus based on the detected signal corresponding to the detected non-modulated light or modulated light. Therefore, the light detection unit 4 includes a light receiver 41 for receiving non-modulated light or modulated light, as shown in FIG. The light receiver 41 has sensitivity to the entire wavelength region of near-infrared light or visible light that can be emitted from the light emitting unit 1, and includes, for example, a PIN diode (P-Intrinsic-N Diode), a photomultiplier, and the like. (Photo multiplier), APD (Avalanche Photo Diode), etc. are the main components. When unmodulated light or modulated light is received from the optical path separation unit 2, an electrical detection signal is output to the AD converter 42 in time series. The AD converter 42 converts the electrical detection signal (analog signal) output from the light receiver 41 into a digital signal and outputs it.

Here, as described above, the light emitting unit 1 can emit the non-modulated light by the light generator 11 and the modulated light by the light generator 12. Therefore, the light detection unit 4 generates non-modulated light generated by the light generator 11 and reflected from the fundus (hereinafter, the reflected non-modulated light is referred to as non-modulated reflected light) and each light generator 12 generates. Then, modulated light reflected by the fundus (hereinafter, the reflected modulated light is referred to as modulated reflected light) is received. In this case, in order to accurately calculate the oxygen saturation SO ₂ described later, each light generator 12 simultaneously generates modulated light. Therefore, the light detection unit 4 needs to receive and separate each modulated reflected light. . Therefore, the light detection unit 4 includes a spread code sequence acquisition unit 43 for selectively separating the modulated reflected light based on the modulated light emitted from the specific light generator 12. The spread code sequence acquisition unit 43 is connected to the controller 6 and acquires from the controller 6 a spread code sequence, that is, a PN sequence included in the modulated light from the specific light generator 12 to be received. Then, the spread code sequence acquisition unit 43 supplies the acquired PN sequence to each multiplier 44.

  The multiplier 44 calculates the product of the detection signal acquired from the AD converter 42 and the PN sequence supplied from the spreading code sequence acquisition unit 43. Then, the multiplier 44 outputs the calculated product value of the detection signal and the PN sequence to the accumulator 45. The accumulator 45 adds the supplied product value over one period or more of the supplied PN sequence. Then, the accumulator 45 outputs a detection signal corresponding to the modulated reflected light based on the modulated light generated by the specific light generator 12 to the calculator 46.

The computing unit 46 uses the detection signal output from the AD converter 42, more specifically, the detection signal of the non-modulated reflected light, based on the light amount distribution of the non-modulated reflected light reflected in the vicinity of the fundus. A surface shape signal representing the shape of the measurement area (ie, the surface shape of the fundus) is calculated. In addition, about the calculation of a surface shape signal, since a well-known calculation method can be employ | adopted, the detailed description is abbreviate | omitted. Further, the computing unit 46 uses the detection signal output from the accumulator 45, that is, the detection signal of the modulated reflected light reflected at the fundus vicinity (for example, a position advanced about 300 μm from the fundus surface), Based on the light amount of the modulated light emitted from each light generator 12 and the light amount of the received modulated reflected light, the oxygen saturation SO ₂ of the blood flowing in the capillary in the predetermined measurement region on the fundus is calculated.

Here, the calculation of the blood oxygen saturation SO _{2 by} the calculator 46 will be described. The absorption characteristics of near-infrared light in hemoglobin in blood, more specifically, oxygenated hemoglobin combined with oxygen and reduced hemoglobin not combined with oxygen, are described in literature (for example, Hitachi Medical Corporation, MEDIX, vol. 29, etc.). As is generally known, it can be expressed by the following equation 1 according to Lambert-Beer's law.
−ln (R (λ) / Ro (λ)) = εoxy (λ) × Coxy × d + εdeoxy (λ) × Cdeoxy × d + α (λ) + S (λ) Equation 1

  However, R (λ), Ro (λ), and d in the equation 1 respectively represent the detected light amount of the wavelength λ, the emitted light amount of the wavelength λ, and the optical path length of the detection region, as schematically shown in FIG. It represents. Further, εoxy (λ) in the above formula 1 represents the molecular extinction coefficient of oxygenated hemoglobin with respect to the wavelength λ, and εdeoxy (λ) represents the molecular extinction coefficient of reduced hemoglobin with respect to the wavelength λ. In the formula 1, Coxy represents the concentration of oxygenated hemoglobin, and Cdeoxy represents the concentration of reduced hemoglobin. Further, α (λ) in the above formula 1 represents the attenuation due to light absorption of a pigment other than hemoglobin in blood (for example, cytochrome aa33 reflecting the supply and demand of oxygen in mitochondria in cells). (λ) represents the attenuation due to light scattering of the living tissue.

In this way, based on the light absorption characteristics of hemoglobin in blood expressed according to the above equation 1, for example, by taking into account the blood flow change in the blood vessel and taking into account the difference before and after the blood flow change, the oxygen saturation level of the blood SO ₂ can be calculated. More specifically, if the light absorption characteristic before blood flow change is expressed according to the above-described equation 1 for capillaries existing near the fundus surface, the light absorption characteristic after blood flow change can be expressed as shown in the following equation 2. .
−ln (growthR (λ) / Ro (λ)) = εoxy (λ) × growthCoxy × d + εdeoxy (λ) × growthCdeoxy × d + growthα (λ) + S (λ) Equation 2
However, growthR (λ), growthCoxy, growthCdeoxy and growthα (λ) in Equation 2 above represent values that have increased or decreased due to changes in blood flow, respectively, It represents the oxygenated hemoglobin concentration after the blood flow change, the reduced hemoglobin concentration after the blood flow change, and the attenuation due to light absorption of a pigment other than hemoglobin after the blood flow change.

Here, since the light absorption amount of hemoglobin in the blood is extremely larger than the light absorption amount of a dye other than hemoglobin, α (λ) in the above equation 1 is expressed as α (λ) = growthα (λ). be able to. Thus, subtracting the formula 1 from the formula 2 yields the following formula 3.
−ln (growthR (λ) / R (λ)) = εoxy (λ) × ΔCoxy + εdeoxy (λ) × ΔCdeoxy Equation 3
Here, ΔCoxy and ΔCdeoxy in Formula 3 are represented by Formula 4 and Formula 5 below, respectively.
ΔCoxy = (growthCoxy−Coxy) × d Equation 4
ΔCdeoxy = (growthCdeoxy−Cdeoxy) × d (Formula 5)

Then, as schematically shown in FIG. 7 as a light absorption spectrum of hemoglobin, as a specific wavelength at which the contrast ratio of light absorption characteristics becomes clear, for example, the result of measurement using near infrared light of λ = 780 nm or 830 nm Based on this, by solving Equation 3, it is possible to relatively calculate oxygenated hemoglobin concentration change ΔCoxy, reduced hemoglobin concentration change ΔCdeoxy, and total hemoglobin concentration change (ΔCoxy + ΔCdeoxy). Then, by calculating these values, the relative oxygen saturation SO ₂ represented by the following formula 6 can be calculated.
SO ₂ = ΔCoxy / (ΔCoxy + ΔCdeoxy) (Formula 6)

Here, the oxygenated hemoglobin concentration change ΔCoxy, the reduced hemoglobin concentration change ΔCdeoxy total hemoglobin concentration change (ΔCoxy + ΔCdeoxy) and the oxygen saturation level SO ₂ in the above formula 6 are calculated as follows: It is calculated using the detected light quantity of near-infrared light reflected by hemoglobin, that is, modulated reflected light. By the way, the detected light amount of the modulated reflected light indicates the reflection intensity (refractive index change, etc.) at a predetermined measurement depth (for example, corresponding to the thickness of the capillary), but the influence of the absorption of the modulated reflected light is the same. It is affected by the hemoglobin concentration in the entire optical path through which light passes. That is, the amount of modulated reflected light is absorbed twice in the vicinity of the fundus surface.

Therefore, when calculating the oxygenated hemoglobin concentration change ΔCoxy, reduced hemoglobin concentration change ΔCdeoxy, total hemoglobin concentration change (ΔCoxy + ΔCdeoxy) and oxygen saturation SO ₂ taking into account the absorption of modulated reflected light inside the fundus A ratio between the light quantity of the modulated reflected light at the depth and the light quantity of the modulated reflected light at the change Δ from the predetermined measurement depth may be obtained. At this time, the light intensity of the modulated reflected light is a combination of wavelengths (for example, 780 nm and 830 nm) in which the reflection intensity at a predetermined measurement depth and the reflection intensity at the change amount Δ are substantially the same and the absorption attenuation amount by hemoglobin is different. The ratio of

  In addition, in the combination of these different wavelengths, the refractive index that determines the reflection intensity can be ignored because the difference between the two wavelengths is small in the living body constituent material. As a result, the absorption attenuation ratio at the two wavelengths of the modulated reflected light within the width Δ can be obtained, and each hemoglobin concentration can be calculated using this absorption attenuation ratio.

Therefore, it is possible to calculate the oxygenated hemoglobin concentration change ΔCoxy, the reduced hemoglobin concentration change ΔCdeoxy, the total hemoglobin concentration change (ΔCoxy + ΔCdeoxy), and the oxygen saturation SO ₂ only at a predetermined measurement depth. As described above, when the calculator 46 calculates the surface shape of the fundus and the oxygen saturation level SO ₂ , the calculator 46 performs image processing on the surface shape signal indicating the calculated shape and the oxygen saturation level signal indicating the oxygen saturation level SO _2. Output to the unit 47.

  As shown in FIG. 8, the image processor 47 is a circuit including a frame control circuit 47a, a frame memory 47b, a multiplexer 47c, and an image generation circuit 47d. The frame control circuit 47a is a circuit that controls the operation of each frame memory 47b and the multiplexer 47c. The frame memory 47b outputs the surface shape signal or the oxygen saturation signal output from the computing unit 46 to the image generation circuit 47d under the control of the frame control circuit 47a. The image generation circuit 47d, based on the output surface shape signal or oxygen saturation signal, displays image data to be displayed in a predetermined manner on the display unit 5, that is, surface shape image data and oxygen saturation for displaying the surface shape. This generates oxygen saturation image data for displaying. In this embodiment, the surface shape signal or the oxygen saturation signal output from the computing unit 46 is temporarily stored in the frame memory 47b. However, if necessary, the signals are multiplexed in the multiplexer 47c. You may implement so that it may output directly.

  As shown in FIG. 9, the display unit 5 includes a display image data storage circuit 51, a conversion circuit 52, and a monitor 53 such as a liquid crystal display. The display image data storage circuit 51 synthesizes the surface shape image data and the oxygen saturation image data as necessary, and includes additional information on the surface shape image data, the oxygen saturation image data, and the composite image data. Data such as certain numbers and various characters are superimposed and temporarily saved. The conversion circuit 52 performs, for example, D / A conversion and TV format conversion on the image data stored in the display image data storage circuit 51. The display unit 5 displays the surface shape of the fundus or the oxygen saturation as it is based on the image data output from the image processor 47 of the light detection unit 4, or synthesizes (superimposes) these image data. To display.

  Next, the operation of the functional confocal image display apparatus M according to the embodiment configured as described above will be described by exemplifying a case where the fundus of the patient is observed.

  First, a doctor or an operator arranges the functional confocal image display device M so that the patient's eyeball is positioned on the optical axis of near infrared light or visible light irradiated by the light emitting unit 1. The doctor or operator adjusts the focus with respect to the fundus by moving the objective lens T in the direction of the arrow in FIG. 1, and then operates an input device (not shown) of the controller 6 to start emission of near infrared light or visible light. Instruct. Thereby, the controller 6 supplies a drive signal or a primary drive signal to the light generator 11 and the light generator 12 constituting the light emitting unit 1 at predetermined short time intervals. Thereby, the light generator 11 and the light generator 12 start the operation alternately at a predetermined short time interval. It is also possible to operate the light generator 11 and the light generator 12 at the same time.

  That is, the light generator 11 acquires the drive signal supplied from the controller 6 at a predetermined short time interval by the light source driver 111. Thereby, the light source driver 111 causes the light source 112 to emit pulses based on the acquired drive signal, and the light source 112 emits visible light of 488 nm.

  The controller 6 supplies primary drive signals to the two light generators 12 at predetermined short time intervals. As a result, the two light generators 12 start operating simultaneously, and emit near-infrared light of 780 nm and near-infrared light of 830 nm, respectively.

  Specifically, in each light generator 12, the spread code sequence generator 121 generates a gold code sequence as a PN sequence, for example. The spreading code sequence generator 121 outputs the generated PN sequence to the controller 6 and also outputs it to the multiplier 122. The multiplier 122 calculates the product of the primary drive signal supplied from the controller 6 and the PN sequence, and performs spread spectrum modulation on the primary drive signal. Then, the light source driver 123 obtains the spread spectrum modulated drive signal, that is, the secondary drive signal. Thereby, the light source driver 123 causes the light source 124 to emit pulses based on the acquired secondary drive signal. Accordingly, one light source 124 emits near-infrared light of 780 nm, and the other light source 124 emits near-infrared light of 830 nm.

  As described above, the near-infrared light and the visible light emitted from the light emitting unit 1 in a pulse manner, that is, the non-modulated light and the modulated light pass through the pinhole plate A1 through the optical fiber H1 and the condenser lens S1. Then, it reaches the optical path separation unit 2 as a spot-like point light source. Then, the unmodulated light and the modulated light that have reached the optical path separating unit 2 are transmitted through the beam splitter 21, scanned in the in-plane direction by the optical scanning unit 3, and reach the patient's eyeball. Then, the unmodulated light and the modulated light incident on the eyeball are reflected near the fundus surface and reach the beam splitter 21. Thus, the non-modulated light and the modulated light that have reached the beam splitter 21 are reflected by the beam splitter 21 and propagate toward the light detection unit 4. Thereby, the non-modulated light and the modulated light reflected by the fundus, that is, the non-modulated reflected light and the modulated reflected light reach the light detection unit 4 through the pinhole plate A2, the condensing lens S3, and the optical fiber H2. At this time, only the non-modulated reflected light and the modulated reflected light reflected by the objective lens T near the surface of the fundus in focus are detected by passing the non-modulated reflected light and the modulated reflected light through the pinhole plate A2. Part 4 is reached.

  In the light detection unit 4, the light receiver 41 receives the non-modulated reflected light and the modulated reflected light, and responds to the non-modulated reflected light or the modulated reflected light including the surface shape information or biological information of a predetermined measurement region on the fundus. The detected electrical signals are output to the AD converter 42 in time series. The magnitude of the electrical detection signal is proportional to the reflection intensity (light quantity) on the fundus. The AD converter 42 converts the output electrical detection signal into a digital signal and outputs the digitally converted detection signal.

  In this case, when the light generator 11 of the light emitting unit 1 generates unmodulated light, the computing unit 46 acquires a digital detection signal from the AD converter 42. Accordingly, the computing unit 46 calculates the surface shape of the fundus in a predetermined measurement region based on the digital detection signal corresponding to the 488 nm non-modulated reflected light generated by the light generator 11 and represents the calculated shape. Output surface shape signal.

  More specifically, as described above, the non-modulated reflected light that is reflected by the fundus and reaches the light detection unit 4 is scanned by the optical scanning unit 3 in a predetermined measurement area on the fundus in the in-plane direction. ing. For this reason, the digital detection signal output from the light receiver 41 to the computing unit 46 via the AD converter 42 becomes a two-dimensional light amount distribution of non-modulated reflected light near the fundus surface. Here, the two-dimensional light amount distribution of the non-modulated reflected light changes according to the shape near the fundus surface. For this reason, the surface shape of the fundus can be calculated by using this light quantity distribution. Then, the calculator 46 outputs a surface shape signal representing the calculated surface shape of the fundus to the image processor 47.

  On the other hand, when the two light generators 12 of the light emitting unit 1 are simultaneously generating modulated light of 780 nm and 830 nm, respectively, a digital detection signal obtained by demodulating the modulated reflected light based on these modulated lights is calculated. Is output to the device 46. Specifically, as described above, the controller 6 supplies the primary drive signal to the light emitting unit 1 and then obtains the PN sequence (for example, from the spread code sequence generator 121 of each light generator 12, for example). (Gold code sequence) is supplied to the light detection unit 4. Thereby, the light detection unit 4 acquires the supplied Gold code sequence by the spread code sequence acquisition unit 43. Then, the spread code sequence acquisition unit 43 supplies the acquired Gold code sequence to the multiplier 44.

  The multiplier 44 calculates the product of the digital detection signal acquired from the AD converter 42 and the Gold code sequence supplied from the spreading code sequence acquisition unit 43. Then, the multiplier 44 outputs the calculated product value to the accumulator 45, and the accumulator 45 adds the output product value over one period of the PN sequence (ie, 128 bits long) or more. . In this way, the product-sum processing by the multiplier 44 and the accumulator 45 can correlate the digital detection signal with the supplied Gold code sequence, and the near-infrared light from the specific light generator 12, specifically, Specifically, only the digital detection signal corresponding to the modulated reflected light having a wavelength of 780 nm or 830 nm can be selected and output to the calculator 46.

  That is, as described above, the Gold code sequence (PN sequence) has a property that different sequences are orthogonal to each other, in other words, a property that the sum of products of different sequences is “0”. For this reason, when the spread code sequence acquisition unit 43 supplies the multiplier 44 with the gold code sequence (PN sequence) of the specific light generator 12, the digital detection signal output from the AD converter 42 The product value of the detection signal other than the digital detection signal corresponding to the near-infrared light generated by the specific light generator 12 and the supplied gold code sequence (PN sequence) is “0”. As a result, the value added by the accumulator 45 over one period of the Gold code sequence (PN sequence) is also “0”, and the correlation is “0”. Therefore, a digital detection signal that does not have (or does not match) the Gold code sequence (PN sequence) supplied from the spread code sequence acquisition unit 43, in other words, modulated light generated by a light generator other than the specific light generator 12. The modulated reflected light based on is selectively excluded, and only the detection signal corresponding to the modulated reflected light based on the modulated light emitted from the specific light generator 12 is output to the calculator 46.

Then, the arithmetic unit 46 uses the digital detection signal corresponding to the 780 nm modulated reflected light and the digital detection signal corresponding to the 830 nm modulated reflected light supplied from the accumulator 45 to calculate as described above. An oxygen saturation signal representing the oxygen saturation SO ₂ at the site matching the surface shape of the fundus is output. That is, the computing unit 46 uses the obtained digital detection signal corresponding to the modulated reflected light of 780 nm and 830 nm, in other words, using the light amount distribution of the portion matching the surface shape of the fundus oculi described above, and according to the above formulas 1-6. Calculate the oxygen saturation SO ₂ . Then, the calculator 46 outputs an oxygen saturation signal representing the calculated oxygen saturation SO ₂ to the image processor 47.

In the image processor 47, the frame control circuit 47a temporarily stores the surface shape signal and the oxygen saturation signal output from the calculator 46 in the frame memory 47b. Then, the frame control circuit 47a causes the multiplexer 47c to output the surface shape signal and the oxygen saturation signal temporarily stored in the predetermined storage position of the frame memory 47b to the image generation circuit 47d. The image generation circuit 47d generates surface shape image data representing the surface shape of the fundus oculi based on the output surface shape signal, and oxygen that matches the position of the surface shape of the fundus oculi based on the output oxygen saturation signal. Oxygen saturation image data representing saturation SO ₂ is generated. Then, the image generation circuit 47d outputs the generated surface shape image data and oxygen saturation image data to the display unit 5.

  In the display unit 5, the display image data storage circuit 51 temporarily stores the surface shape image data and the oxygen saturation image data supplied from the image generation circuit 47d. Then, the image data once stored in the display image data storage circuit 51 is converted by the conversion circuit 52, so that the monitor 53 displays or synthesizes the surface shape of the fundus and the oxygen saturation of the fundus. Display surface shape and oxygen saturation.

As can be understood from the above description, according to the functional confocal image display device M according to the first embodiment, the surface shape of the fundus in the predetermined measurement region is displayed, and the portion that matches the surface shape The oxygen saturation of SO ₂ can be calculated and displayed. Further, these surface shapes and oxygen saturation SO ₂ can be synthesized and displayed. As a result, for example, when a doctor examines an eye disease in which photoreceptor cells are necrotized such as glaucoma, it is possible to provide information on both the surface shape of the fundus and the oxygen saturation SO ₂ , so that the early stage of the disease state Can contribute to discovery. In other words, with optical coherence tomography and confocal microscopes that have been used for this type of diagnosis, the surface shape and cross-sectional shape of the fundus can be observed in detail. It was necessary to depend on knowledge and experience. However, according to the functional confocal image display apparatus M, since the oxygen saturation SO ₂ can be observed simultaneously with the surface shape of the fundus, for example, it is very easy to reduce the oxygen saturation SO ₂ due to necrosis of photoreceptor cells. Can be confirmed. Thereby, a doctor's diagnosis can be supported accurately and an appropriate treatment can be performed early on a patient.

In addition, the two light generators 12 of the light emitting unit 1 can perform near-infrared light having different wavelengths by performing spread spectrum modulation, so that the change in the oxygen saturation SO ₂ can be calculated in more detail. That is, although the time change of the oxygen saturation SO ₂ is relatively slow, strictly speaking, it changes with time. On the other hand, by simultaneously emitting near-infrared light having different wavelengths, modulated reflected light that reflects the state of oxygen saturation SO ₂ at substantially the same time point reaches the light detection unit 4. For this reason, the instantaneous oxygen saturation level SO ₂ can be calculated satisfactorily, and the change in the oxygen saturation level SO ₂ over time can also be calculated extremely accurately.

In the first embodiment, the equation 1) to Equation (6 (more specifically, the formula 6) in accordance with, was performed to calculate the oxygen saturation SO _2. Incidentally, the oxygenated hemoglobin concentration change ΔCoxy and the reduced hemoglobin concentration change ΔCdeoxy calculated in the first embodiment are calculated including the optical path length d, as is apparent from the equations 4 and 5. In general, it is extremely difficult to precisely measure or calculate the optical path length d of light incident on the inside of a living body. Therefore, the optical path length d in the equations 4 and 5 is used as a relative amount, and the oxygen saturation SO ₂ calculated according to the equation 6 using the oxygenated hemoglobin concentration change ΔCoxy and the reduced hemoglobin concentration change ΔCdeoxy is also relative. It becomes quantity.

In contrast, by calculating the oxygen saturation SO ₂ in accordance with the formulas shown below, the oxygen saturation SO ₂ in the pulsating _component, in other words, to calculate the oxygen saturation SO ₂ arterial or arterioles Can do. Note that the oxygen saturation calculation method according to this modification is, for example, a calculation method disclosed in Japanese Patent Laid-Open No. 63-111837 and widely known in the art, and detailed description thereof is omitted. .

The infrared attenuation level in the living body can be calculated according to the following formula 7.
−log (I1 / I0) = E × C × e + A Equation 7
However, I1 in the formula 7 represents the amount of transmitted light that has passed through the arteriole, and I0 represents the amount of incident light. E in Equation 7 represents the extinction coefficient of hemoglobin, C represents blood hemoglobin blood concentration, e represents blood layer thickness (corresponding to the optical path length d in Equations 4 and 5), A Represents the degree of attenuation of the tissue layer. Here, Equation 7 is used to calculate the attenuation of infrared light transmitted through the living body, but it is known that even reflected infrared light exhibits similar characteristics.

If the blood layer thickness e is changed by Δe due to pulsation, the change in the infrared attenuation can be calculated according to the following equation (8).
− (Log (I1 / I0) −log (I2 / I0)) = E × C × e−E × C × (e−Δe) Equation 8
If the above equation 8 is arranged, the following equation 9 is obtained.
−log (I2 / I1) = E × C × Δe Equation 9
However, I2 in the above formulas 8 and 9 represents the amount of transmitted light after the change in the thickness of the blood layer.

Next, assuming that the wavelength of the infrared light having the transmitted light amount I1 is λ1, the wavelength of the infrared light having the transmitted light amount I2 is λ2, the light amounts of the transmitted light at λ1 at times t1 and t2 are I11, I21, λ2. Assuming that the amount of each transmitted light is I12 and I22, the change in the infrared attenuation at each time can be expressed by the following equations 10 and 11 according to the equation 9.
−log (I21 / I11) = E1 × C × Δe Equation 10
−log (I22 / I12) = E2 × C × Δe Equation 11
However, E1 in the formula 10 represents the extinction coefficient of hemoglobin with respect to the infrared light having the wavelength λ1, and E2 in the formula 11 represents the extinction coefficient of hemoglobin with respect to the infrared light having the wavelength λ2. Then, when the formula 11 is divided by the formula 10, the following formula 12 is established in which the blood layer thickness change Δe is eliminated.
log (I12 / I22) / log (I11 / I21) = E2 / E1 Equation 12
Therefore, if Equation 12 is modified, the following Equation 13 is established.
E2 = E1 × log (I12 / I22) / log (I11 / I21) ... Equation 13

Here, referring to the light absorption spectrum of hemoglobin corresponding to the oxygen saturation shown in FIG. 10, when 805 nm is selected as the absorption wavelength corresponding to the absorption coefficient E1 of hemoglobin, the oxygen saturation SO ₂ = 0% and the oxygen saturation Obtain the intersection of the curves of degree SO ₂ = 100%. As a result, the extinction coefficient E1 becomes a value that is not affected by the oxygen saturation. Then, for example, 750 nm is selected as the absorption wavelength corresponding to the absorption coefficient E2 of hemoglobin, and the absorption coefficient of hemoglobin when the oxygen saturation level SO ₂ = 0% is Ep and the oxygen saturation level SO ₂ = 100%. Assuming that the extinction coefficient of hemoglobin is E0, the current oxygen saturation SO ₂ can be calculated according to the following equation (14).
SO ₂ = (E2−Ep) / (E0−Ep) Equation 14
As a result, the oxygen saturation SO ₂ calculated according to the equation 14 is calculated without including a relative amount, so that the actual oxygen saturation can be obtained. Therefore, more accurate oxygen saturation SO ₂ can be provided in diagnosis by a doctor.

b. Second Embodiment In the first embodiment, the objective lens T is moved to adjust the focus with respect to the fundus, and the non-modulated reflected light and the modulated reflected light reflected by a predetermined measurement region are reflected by the beam splitter 21 to generate light. By passing the pinhole plate A2 fixedly disposed on the road, information in a predetermined measurement region, in other words, the focal plane of the objective lens T was selectively extracted. As a result, the non-modulated reflected light and the modulated reflected light from the portion not condensed on the focal plane are eliminated by the pinhole plate A2, and the surface shape and the oxygen saturation SO ₂ can be obtained with high accuracy. .

However, in this case, focus adjustment by the objective lens T is necessary, and it may be difficult to strictly cope with the complex surface shape of the fundus. To solve this problem, the pinhole plate A2 of the first embodiment is configured as a confocal aperture CA that is movably disposed on the optical path, and the focus adjustment of the objective lens T with respect to the fundus is omitted. be able to. Thereby, the surface shape and oxygen saturation SO ₂ can be obtained easily and accurately. Hereinafter, although this 2nd Embodiment is described, the same code | symbol is attached | subjected to the same part as the said 1st Embodiment, and the detailed description is abbreviate | omitted.

  In the second embodiment, as shown in FIG. 11, the confocal aperture CA is disposed on the optical path where the non-modulated reflected light and the modulated reflected light reflected by the beam splitter 21 are collected by the condenser lens S3. ing. The confocal aperture CA is provided so as to move in the optical axis direction of the unmodulated reflected light and the modulated reflected light reflected by the beam splitter 21 by a driving device (not shown).

  Also in the functional confocal image display device M according to the second embodiment configured as described above, unmodulated light and modulated light are emitted from the light emitting unit 1 as in the first embodiment. The emitted non-modulated light and modulated light become a point light source by the pinhole plate A1 and reach the fundus via the optical path separating unit 2 and the optical scanning unit 3. In this case, the focal plane of the objective lens T is set to a predetermined plane in the measurement region. Then, the non-modulated reflected light and the modulated reflected light reflected by the predetermined measurement region on the fundus are reflected by the beam splitter 21 and condensed by the condenser lens S3, and reach the confocal aperture CA. In this state, by moving the confocal aperture CA, the non-modulated light or the modulated light scanned in the in-plane direction by the light scanning unit 3 forms the most image on the curved fundus surface, that is, the most focused light. Only matching unmodulated reflected light or modulated reflected light can be selectively passed.

Thereby, the unmodulated reflected light and the modulated reflected light that have passed through the confocal aperture CA reach the light detection unit 4 via the optical fiber H2. Therefore, the light detection unit 4 can accurately detect the surface shape of the fundus based on the non-modulated reflected light, and can accurately detect the oxygen saturation SO ₂ of the fundus based on the modulated reflected light. Thereby, the display unit 5 can accurately display the surface shape of the fundus and the oxygen saturation level SO ₂ . As for other effects, the same effects as in the first embodiment can be expected.

c. Other Modifications The implementation of the present invention is not limited to the above-described embodiments and modifications, and various modifications can be made without departing from the object of the present invention.

  For example, in each of the above embodiments, the light generator 12 of the light emitting unit 1 is implemented so as to generate modulated light by performing spread spectrum modulation. On the other hand, other modulation schemes, for example, ASK (Amplitude Shift Keying) modulation, FDMA (Frequency Division Multiple Access) modulation, and the like can be employed. Even when these modulation methods are employed, the two light generators 12 can emit the modulated light without interfering with each other. Then, the photodetection unit 4 demodulates the digital detection signal output from the AD converter 42 by a predetermined demodulation method, and based on the modulated light emitted from the specific light generator 12 as in the above embodiments. Only the digital detection signal of the modulated reflected light can be output to the calculator 46. Therefore, even in this case, the same effects as those of the above embodiments can be expected.

In each of the above-described embodiments and modifications thereof, oxygen saturation as biological information is performed using the amount of modulated light emitted from the light emitting unit 1 and the amount of modulated reflected light detected by the light detection unit 4. The degree of SO ₂ was calculated. On the other hand, according to the functional confocal image display apparatus M according to the present invention, the light amount of the modulated light emitted from the light emitting unit 1 and the light amount of the modulated reflected light detected by the light detecting unit 4 are used. If it can be calculated, other biological information, for example, blood flow in blood vessels or blood flow changes can be calculated and displayed on the display unit 5. Thus, in each of the above embodiments, the functional confocal image display device M is applied to the examination of the fundus, but the functional confocal image display device M is used for the examination of other parts of the living body. Needless to say, this is possible.

1 is a block diagram illustrating an outline of a functional confocal image display device according to a first embodiment of the present invention. It is a block diagram which shows schematically the structure of the light-projection part of FIG. It is the schematic for demonstrating the structure of the optical path separation part of FIG. It is the schematic for demonstrating the structure of the optical scanning part of FIG. FIG. 2 is a block diagram schematically showing a configuration of a light detection unit in FIG. 1. It is a figure for demonstrating derivation | leading-out of oxygen saturation. It is the graph which showed roughly the change of the molecular extinction coefficient with respect to the wavelength of oxygenated hemoglobin and reduced hemoglobin. FIG. 6 is a block diagram schematically showing the configuration of the image processor in FIG. 5. FIG. 2 is a block diagram schematically showing a configuration of a display unit in FIG. 1. It is the graph which showed the change of the molecular extinction coefficient with respect to the wavelength according to the modification of 1st Embodiment with respect to the difference of oxygen saturation. It is a block diagram which shows the outline of the functional confocal image display apparatus which concerns on 2nd Embodiment of this invention.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 ... Light emission part, 11, 12 ... Light generator, 111, 123 ... Light source driver, 112, 124 ... Light source, 121 ... Spreading code sequence generator, 122 ... Multiplier, 2 ... Optical path separation part, 21 ... Beam splitter DESCRIPTION OF SYMBOLS 3 ... Optical scanning part, 31 ... Galvano mirror, 32 ... Galvano mirror, 4 ... Light detection part, 41 ... Light receiver, 42 ... AD converter, 43 ... Spreading code sequence acquisition device, 44 ... Multiplier, 45 ... Accumulation , 46 ... arithmetic unit, 47 ... image processor, 5 ... display unit, 6 ... controller, S1, S2, S3 ... condensing lens, A1, A2 ... pinhole plate, CA ... confocal aperture, M ... functional type Confocal image display device

Claims (7)

A controller that is operated by a user and outputs various signals based on the user's instructions;
Non-modulated light generating means for generating near-infrared light or visible light having at least one specific wavelength based on a predetermined drive signal supplied from the controller, and two modulated drive signals supplied from the controller A light emitting unit having a modulated light generating means for generating a next drive signal and generating near infrared light having at least two specific wavelengths based on the secondary drive signal;
The near-infrared light reflected from the subject while transmitting the spot-like near-infrared light or visible light emitted from the light emitting part through the first pinhole plate having a predetermined hole diameter toward the subject. An optical path separating unit that separates light or visible light from the optical path of the transmitted near-infrared light or visible light;
An optical scanning unit that scans near-infrared light or visible light transmitted from the optical path separation unit toward the subject in an in-plane direction with respect to the subject; and
Of the near-infrared light or visible light received by the light-receiving means, and the light-receiving means that receives the near-infrared light or visible light separated by the optical path separation unit through the second pinhole plate having a predetermined hole diameter Surface shape information calculating means for calculating surface shape information representing a shape near the surface of the subject based on the amount of near infrared light or visible light generated by the non-modulated light generating means of the light emitting section; and the light receiving means Demodulating means for demodulating the secondary drive signal of the near-infrared light generated by the modulated-light generating means of the light emitting part out of the near-infrared light or visible light received by the light into the predetermined drive signal; Biological information for calculating biological information of the subject accompanying metabolism of the living body based on the amount of near infrared light generated by the modulated light generating means and the amount of near infrared light demodulated by the demodulating means An optical detection unit comprising: a calculation unit; and an image data generation unit that generates visible image data based on the surface shape information calculated by the surface shape information calculation unit and the biological information calculated by the biological information calculation unit. When,
Display that displays a surface shape image of the subject, a biological information image of the subject, or a composite image obtained by combining the surface shape image of the subject and the biological information image based on the image data generated by the light detection unit And a functional confocal image display device. In the functional confocal image display device according to claim 1,
The modulated light generating means of the light emitting part is
Spread spectrum modulation means for generating a secondary drive signal by performing spread spectrum modulation on a predetermined drive signal supplied from the controller;
Each light source is composed of light synthesizing means for optically synthesizing near-infrared light having different specific wavelengths generated simultaneously based on the secondary drive signal,
The demodulating means of the light detection unit is
A functional confocal image display device, wherein the secondary drive signal contained in the near-infrared light received by the light receiving means is despread and demodulated to the predetermined drive signal. In the functional confocal image display device according to claim 1,
The modulated light generating means of the light emitting part is
Frequency division multiplex modulation means for generating a secondary drive signal by frequency division multiplex modulation of a predetermined drive signal supplied from the controller;
Each light source is composed of light synthesizing means for optically synthesizing near-infrared light having different specific wavelengths generated simultaneously based on the secondary drive signal,
The demodulating means of the light detection unit is
A functional confocal image display device, wherein the secondary drive signal contained in near-infrared light received by the light receiving means is demodulated into the predetermined drive signal. In the functional confocal image display device according to claim 1,
The functional confocal image display device, wherein the second pinhole plate is a confocal aperture movable in the optical axis direction of near-infrared light or visible light separated by the optical path separation unit. In the functional confocal image display device according to claim 1,
The display unit
Displaying a composite image obtained by synthesizing the surface shape image and the biological information image by matching the position specified by the surface shape image of the subject with the position specified by the biological information image of the subject. A functional confocal image display device. In the functional confocal image display device according to claim 1,
The biological information calculated by the biological information calculation means of the light detection unit is
A functional confocal image display device which is one of blood volume, blood flow volume, blood flow change and oxygen saturation in the blood vessel of the subject. In the functional confocal image display device according to claim 1,
The functional confocal image display device, wherein the subject is a fundus of an eyeball.