JP2008267913A - Nuclear medicine diagnostic apparatus and diagnostic system used for same - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a nuclear medicine diagnostic apparatus capable of improving scatter correction accuracy, and a diagnostic system used for the nuclear medicine diagnostic apparatus. <P>SOLUTION: A background region is extracted using transmission data to obtain a physical quantity which minimizes the difference quantity between emission data (an original image) related to the background region and a scatter component. The physical quantity derived via each means of a superposition integrating-subtracting section 14 is a value obtained in every imaging environment of a subject and is optimized for every imaging environment of the subject. Based on the thus obtained physical quantity, the difference quantity between the scatter component and the emission data (the original image) can thereby be obtained as final emission data according to each imaging environment of the subject, and scatter correction accuracy based on scatter correction by a low-pass filter (LPF) 14d and a subtractor 14g can be improved. <P>COPYRIGHT: (C)2009,JPO&INPIT

Description

  The present invention relates to a nuclear medicine diagnostic apparatus for obtaining nuclear medicine data of a subject based on radiation generated from the subject to which a radiopharmaceutical is administered, and a diagnostic system used therefor.

  As the above-described nuclear medicine diagnosis apparatus, that is, an ECT (Emission Computed Tomography) apparatus, a PET (Positron Emission Tomography) apparatus will be described as an example. The PET apparatus detects a plurality of gamma rays generated by annihilation of positrons, that is, positrons, and reconstructs a tomographic image of a subject only when the gamma rays are simultaneously detected by a plurality of detectors. It is configured

  In this PET apparatus, after a radiopharmaceutical is administered to a subject, the process of drug accumulation in the target tissue is measured over time, whereby quantitative measurement of various biological functions is possible. Therefore, the tomographic image obtained by the PET apparatus has functional information.

  Specifically, taking a human body as an example, a positron (positron) radioactive isotope (for example, 15O, 18F, 11C, etc.) is injected into the body of the subject, and the positron emitted from these is injected. Γ-rays generated when they are combined with electrons are detected. This detection of γ-rays is performed by a detector array comprising a number of γ-ray detectors arranged in a ring shape so as to surround the body axis that is the longitudinal axis of the subject. Then, calculation is performed by a computer in the same manner as in ordinary X-ray CT (Computed Tomography), and the image is specified in the plane, and an image of the subject is created.

  Usually, in order to obtain a three-dimensional image of a subject within a short time, a cylindrical detector stack is provided in which ring-shaped detector arrays are stacked in multiple stages in the body axis direction. The detector stack γ-ray detector configured as described above uses a scintillator, and a detection signal is amplified by a photomultiplier (photomultiplier tube) or the like to obtain a detection signal. In the collection, data (also referred to as “emission data”) are collected as a simultaneous count of two γ-rays radiating in the opposite directions of 180 ° by a number of γ-ray detectors arranged around the subject.

In order to simultaneously count γ-rays, each γ-ray is input to the coincidence counting circuit, and it is determined whether or not the time difference between the input γ-rays is within a predetermined time window. In an actual coincidence circuit, γ rays detected within a very short time window of about 4 ns to 20 ns (ns = 10 ⁻⁹ s) are generally regarded as “simultaneous”. Therefore, there is a possibility that one of γ rays generated at two different points is counted simultaneously. This is called “random coincidence”. On the other hand, when one or both of the pair of γ-rays are simultaneously counted after causing Compton scattering in the subject, this is called “scatter coincidence”. In addition, when both of a pair of γ rays are counted simultaneously, this is called “true coincidence”.

  Counts other than the true coincidence count (T) appear as noise. Therefore, in order to remove the component (scattering component) of the scattering coincidence count (S), scattered ray correction is performed (see, for example, Patent Document 1). For scattered radiation correction, the scattered radiation characteristics of the point source are obtained from a separate experiment, and the scattered component is subtracted by deconvolution and integration with the measured projection data. There is a method of subtracting the window count. There is also a correction method such as obtaining and subtracting the scattering distribution from a simulation based on the model.

Note that such scattered ray correction is usually performed by a technique called a convolution-subtraction method. A conventional superposition integral-subtraction method will be described with reference to FIG. FIG. 10 is a flowchart schematically showing a conventional superposition integral-subtraction method. As shown in FIG. 10, performs convolution using Fourier transform (Fast Fourier Transform) (denoted in FIG. 10, "FFT") on the collected emission data (original image) X _0. In the image X ₁ after the Fourier transform, the scattering component appears as a low frequency component. Therefore, the low-pass filter for the image X ₁ after the Fourier transform: extracting a (LPF Low Pass Filter) scattered component by performing (in FIG. 10 labeled "LPF"). The image of the scattered component and X _2. Perform deconvolution using (denoted by in FIG. 10, "inverse FFT") inverse Fourier transform for the image X ₂ of the scattering component, obtains the image X ₃ of the scattered component after inverse Fourier transform. By deriving the difference amount between the image X ₃ of the scattered component after inverse Fourier transform of the original image X _0, you are possible to obtain an image X ₄ removal of the component of the scattered coincidence count (S).
JP-A-10-206545

  However, in the case of the conventional example having such a configuration, all are approximate methods, and are different from the actual imaging environment of the subject, so that there is a problem that accurate correction is difficult.

  This invention is made | formed in view of such a situation, Comprising: It aims at providing the nuclear medicine diagnostic apparatus which can improve a scattering correction precision, and a diagnostic system used for it.

In order to achieve such an object, the present invention has the following configuration.
That is, the invention described in claim 1 is a nuclear medicine diagnostic apparatus for obtaining nuclear medicine data of a subject based on radiation generated from the subject to which a radiopharmaceutical is administered, and (A) an external of the subject. Morphological information acquisition means for acquiring, for each subject, morphological information based on radiation irradiated from the external radiation source and transmitted through the subject, and (B) acquired by the morphological information acquisition means. A first region extracting means for distinguishing and extracting a background region and other regions from the morphological information; (C) a scattering component estimating means for estimating a scattering component from the nuclear medicine data; and (D) the first Based on the background region extracted by the region extracting means, the background region relating to nuclear medicine data and the background relating to the scattering component estimated by the scattering component estimating means. A second region extracting means for extracting each of the undulated regions; and (E) a first difference for deriving a difference amount between the nuclear medicine data and the scattered component in each of the background regions extracted by the second region extracting means. A quantity derivation means; and (F) a physical quantity derivation means for deriving a physical quantity that minimizes the difference quantity derived by the first difference quantity derivation means, and (a) a physical quantity derived by the physical quantity derivation means. A scatter component extracting means for extracting a scatter component from nuclear medicine data on the basis of the above, and (b) a second difference amount derivation means for deriving a difference amount between the scattered component extracted by the scatter component extracting means and the nuclear medicine data And obtaining the difference amount derived by the second difference amount deriving means as final nuclear medicine data.

  [Operation / Effect] According to the invention described in claim 1, the scattering component extracted by the scattering component extraction unit includes (a) the scattering component extraction unit and (b) the second difference amount derivation unit. The amount of difference from the nuclear medicine data is derived, and the amount of difference is obtained as final nuclear medicine data. For example, the above-described conventional superposition integral-subtraction method can be realized. When the scattering component is extracted from the nuclear medicine data by the scattering component extraction means, extraction is performed using various physical quantities, but the physical quantities differ depending on the imaging environment of the subject (individual subject, imaging location of the subject). . The background region (air region) and other regions (substantially subject region) also differ depending on the individual subject and the imaging region of the subject.

  Therefore, since the data obtained by irradiating the subject from an external radiation source outside the subject has morphological information, first, (A) the morphological information acquisition means is an external radiation source outside the subject. For each subject, morphological information based on radiation irradiated from the external radiation source and transmitted through the subject is acquired for each subject. Since the morphological information acquired by the morphological information acquisition means has information such as an address, the (B) first area extraction means extracts the background area and other areas from the morphological information by distinguishing them. It is possible. On the other hand, (C) scattering component estimation means estimates a scattering component from nuclear medicine data. Based on the background region extracted by the first region extraction means described above, a background region relating to nuclear medicine data and a background region relating to the scattering component estimated by the scattering component estimation means described above are: (D) second region Each extraction means extracts. (E) The first difference amount derivation means derives the difference amount between the nuclear medicine data and the scattered component in the background areas respectively extracted by the second area extraction means.

  If the physical quantity used to extract the scattering component by the above-described (a) scattering component extraction unit is an optimum value, the nuclear medicine data and the scattering component are respectively detected in the background region extracted by the second region extraction unit. Are considered equal. That is, in the background region, it is considered that the difference amount between the nuclear medicine data and the scattered component is almost “0”. Therefore, (F) the physical quantity deriving means derives the physical quantity that minimizes the difference quantity (the nuclear medicine data and the scattering component) derived by the first difference amount deriving means. The physical quantity obtained in this way is a value obtained for each imaging environment of the subject, and is optimized for each imaging environment of the subject. Therefore, based on the physical quantity thus obtained, (a) the scattering component extraction means extracts the scattering component from the nuclear medicine data, and the difference amount between the scattering component extracted by the scattering component extraction means and the nuclear medicine data (B) The second difference amount deriving unit derives the difference amounts as final nuclear medicine data according to the imaging environment of the subject. As a result, it is possible to improve the scattering correction accuracy by the scattering correction by the (a) scattering component extraction means and (b) the second difference amount derivation means.

  In one example of the present invention described above, the nuclear medicine diagnostic apparatus includes the external radiation source described above, and the external radiation source emits the same type of radiation as the above-described radiopharmaceutical (the invention according to claim 2). In the case of this example, the nuclear medicine diagnosis apparatus includes an external radiation source, and the external radiation source irradiates the same kind of radiation as the radiopharmaceutical and passes through the subject, thereby obtaining morphological information based on the radiation. Then, using this form information, the first area extracting means distinguishes and extracts the background area and the other areas, and performs calculations by the means (B) to (F), (a), and (b). I do.

  Moreover, you may apply to the diagnostic system used for the nuclear medicine diagnostic apparatus based on the invention mentioned above. That is, this diagnostic system is a diagnostic system used for a nuclear medicine diagnostic apparatus, and is configured to include a nuclear medicine diagnostic apparatus and an X-ray CT apparatus. The nuclear medicine diagnostic apparatus is administered with a radiopharmaceutical. The X-ray CT apparatus obtains X-ray CT data based on X-rays irradiated from the outside of the subject and transmitted through the subject based on radiation generated from the subject. The nuclear medicine diagnostic apparatus includes (A ′) morphological information acquisition means for acquiring X-ray CT data as morphological information for each subject, and (B) a background region from the morphological information acquired by the morphological information acquisition means. And a first region extracting means for distinguishing and extracting other regions, (C) a scattering component estimating means for estimating a scattering component from nuclear medicine data, and (D) a background extracted by the first region extracting means. A second area extracting means for extracting a background area relating to nuclear medicine data and a background area relating to the scattering component estimated by the scattering component estimating means based on the area; and (E) extracting each by the second area extracting means. A first difference amount deriving unit for deriving a difference amount between the nuclear medicine data and the scattered component in the background region, and (F) a physical quantity that minimizes the difference amount derived by the first difference amount deriving unit. A physical quantity deriving means for deriving, (a) a scattering component extracting means for extracting a scattered component from nuclear medicine data based on the physical quantity derived by the physical quantity deriving means, and (b) extracting by the scattering component extracting means. A second difference amount deriving means for deriving a difference amount between the scattered component and the nuclear medicine data, and the difference amount derived by the second difference amount deriving means, Obtained as final nuclear medicine data (the invention according to claim 3).

  In the case of this system, in the X-ray CT apparatus, X-ray CT data is obtained based on X-rays irradiated from the outside of the subject and transmitted through the subject, and the X-ray CT data is used as morphological information (A ′) The form information acquisition means acquires for each subject. Then, using this form information, the first area extracting means distinguishes and extracts the background area and the other areas, and performs calculations by the means (B) to (F), (a), and (b). I do.

  According to the nuclear medicine diagnostic apparatus and the diagnostic system used therefor according to the present invention, each means of (A) to (F) (in the case of a diagnostic system, (A ′) and (B) to (F)) The physical quantity derived through the process is a value obtained for each imaging environment of the subject and is optimized for each imaging environment of the subject. Therefore, the difference amount between the scattering component and the nuclear medicine data can be obtained as final nuclear medicine data according to the imaging environment of the subject based on the physical quantity thus obtained, respectively (a) It is possible to improve the scattering correction accuracy by the scattering correction by the scattering component extraction means and (b) the second difference amount derivation means.

Embodiment 1 of the present invention will be described below with reference to the drawings.
FIG. 1 is a side view and a block diagram of a PET (Positron Emission Tomography) apparatus according to the first embodiment, and FIG. 2 is a layout diagram of γ-ray detectors in the PET apparatus. In addition, this Example 1 including Example 2 mentioned later demonstrates taking a PET apparatus as an example as a nuclear medicine diagnostic apparatus.

  The PET apparatus according to the first embodiment includes a top plate 1 on which a subject M is placed as shown in FIG. The top plate 1 is configured to move up and down and translate along the body axis Z of the subject M. With this configuration, the subject M placed on the top 1 is scanned from the head to the abdomen and foot sequentially through the opening 2a of the gantry 2, which will be described later. Diagnostic data such as projection data and tomographic images are obtained. This diagnostic data corresponds to nuclear medicine data in the present invention.

  In addition to the top plate 1, the PET apparatus according to the first embodiment includes a gantry 2 having an opening 2a, a plurality of scintillator blocks (not shown) and a plurality of photomultipliers (not shown). The γ-ray detector 3 is provided. As shown in FIG. 2, the γ-ray detector 3 is arranged in a ring shape so as to surround the body axis Z of the subject M, and is embedded in the gantry 2. The photomultiplier is disposed outside the scintillator block. As a specific arrangement of the scintillator blocks, for example, two scintillator blocks are arranged in a direction parallel to the body axis Z of the subject M, and many scintillator blocks are arranged around the body axis Z of the subject M. Can be mentioned. The γ-ray detector 3 collects emission data to be described later.

  In addition, a point line source 4 and a γ-ray detector 5 for collecting absorption correction data (also referred to as “transmission data”) to be described later are provided. The γ-ray detector 5 for absorption correction data includes a scintillator block 5a (see FIG. 2) and a photomultiplier 5b (see FIG. 2) in the same manner as the γ-ray detector 3 for collecting emission data. Has been. The dotted line source 4 is a radiation source that irradiates a radioactive drug to be administered to the subject M, that is, a radiation of the same kind as the radioisotope (RI) (in this embodiment, γ rays), and is arranged outside the subject M. It is installed. The dotted line source 4 is embedded in the gantry 2. The dotted line source 4 rotates around the body axis Z of the subject M. The point source 4 corresponds to the external source in the present invention.

  In addition to performing absorption correction described later, the transmission data includes a background region (air region) BG (see FIG. 6) and other regions (substantially the region of the subject M) R (see FIG. 6). Used to distinguish and extract. The transmission data has form information such as an address. Transmission data corresponds to form information in the present invention.

  In addition, in Example 1, including Example 2 described later, the γ-ray detector 5 has a plurality of scintillator blocks on the body axis Z of the subject M as shown in FIG. 5a and a photomultiplier 5b are provided. That is, a detector array composed of a large number of γ-ray detectors 3 arranged in a ring shape is stacked in the body axis Z direction so as to be stacked in multiple stages. In the first embodiment, there are six rows.

  In addition, the PET apparatus according to the first embodiment includes a top plate driving unit 6, a controller 7, an input unit 8, an output unit 9, a coincidence counting circuit 10, a projection data deriving unit 11, an absorption correction data deriving unit 12, and an absorption. A correction unit 13, a superposition integration-subtraction unit 14, and a memory unit 15 are provided. The top plate driving unit 6 is a mechanism for driving the top plate 1 so as to perform the above-described movement, and is configured by a motor or the like not shown.

  The controller 7 comprehensively controls each part constituting the PET apparatus according to the first embodiment. The controller 7 includes a central processing unit (CPU).

  The input unit 8 sends data and commands input by the operator to the controller 7. The input unit 8 includes a pointing device represented by a mouse, a keyboard, a joystick, a trackball, a touch panel, and the like. The output unit 9 includes a display unit represented by a monitor, a printer, and the like.

  The memory unit 15 is composed of a storage medium represented by ROM (Read-only Memory), RAM (Random-Access Memory), and the like. In the first embodiment, the projection data obtained by the projection data deriving unit 11, the absorption correction data obtained by the absorption correction data deriving unit 12, the projection data subjected to absorption correction by the absorption correction unit 13, and the superposition integration − Various types of data described later processed by the subtracting unit 14 are written and stored in the RAM, and read from the RAM as necessary. The ROM stores in advance programs for performing various nuclear medicine diagnoses, and the controller 7 executes the programs to perform nuclear medicine diagnosis according to the programs.

  The projection data deriving unit 11, the absorption correction data deriving unit 12, the absorption correcting unit 13, and the superposition integral-subtracting unit 14 are, for example, a program or an input unit stored in a ROM of a storage medium represented by the memory unit 15 described above. This is realized by the controller 7 executing a command input by a pointing device represented by 8 or the like.

  The γ-rays generated from the subject M to which the radiopharmaceutical is administered are converted into light by the scintillator block of the γ-ray detector 3, and the converted light is photoelectrically converted by the photomultiplier of the γ-ray detector 3. Output to electrical signal. The electric signal is sent to the coincidence counting circuit 10 as image information (pixel).

  Specifically, when a radiopharmaceutical is administered to the subject M, two γ rays are generated due to the disappearance of the positron of the positron emission type RI. The coincidence counting circuit 10 checks the position of the scintillator block and the incident timing of the γ-ray, and only when the γ-ray is simultaneously incident on two scintillator blocks that are opposed to each other with the subject M interposed therebetween, The information is determined as appropriate data. When γ rays are incident only on one scintillator block, the coincidence circuit 10 treats the image information sent at that time as noise instead of γ rays generated by the disappearance of the positron, and rejects it. To do.

  Actually, even if such processing is performed in the coincidence circuit 10, noise cannot be completely removed, and noise components such as random coincidence or scattered coincidence count remain (especially scattering coincidence count). At the same time, it is sent to the projection data deriving unit 11. The projection data deriving unit 11 obtains the image information sent from the coincidence counting circuit 10 as projection data, and sends the projection data to the absorption correction unit 13. The projection data obtained by the projection data deriving unit 11 is also called “emission data”.

  The point source 4 emits γ rays toward the subject M while rotating around the body axis Z of the subject M, and the irradiated γ rays are used as a scintillator block of the γ ray detector 5 for absorption correction data. The light 5a is converted into light, and the converted light is photoelectrically converted by the photomultiplier 5b of the γ-ray detector 5 and output to an electrical signal. The electrical signal is sent as image information (pixels) to the absorption correction data deriving unit 12 and the superposition integration-subtraction unit 14.

  Absorption correction data is obtained based on the image information sent to the absorption correction data deriving unit 12. The absorption correction data deriving unit 12 uses the calculation representing the relationship between the absorption coefficient of γ-rays or X-rays and the energy to convert the projection data for CT, that is, the distribution data of the X-ray absorption coefficient, to the γ-ray absorption coefficient. By converting to distribution data, the distribution data of the γ-ray absorption coefficient is obtained as absorption correction data. The derived absorption correction data is sent to the absorption correction unit 14. The absorption correction data obtained by the absorption correction data deriving unit 12 is applied to the projection data obtained by the projection data deriving unit 11 to correct the projection data in consideration of the absorption of γ rays in the body of the subject M. . The absorption-corrected projection data is sent to the superposition integration-subtraction unit 14.

  Based on the corrected projection data and the image information detected by the γ-ray detector 5, the superimposition integration / subtraction unit 14 performs various processes described later. A tomographic image obtained through various processes in the superposition integration-subtraction unit 14 is sent to the output unit 9 via the controller 7. The absorption correction data obtained by the absorption correction data deriving unit 12 and the image information sent to the superposition integration-subtraction unit 14 are also referred to as “transmission data”.

  Next, a specific function of the superposition integration-subtraction unit 14 will be described with reference to FIG. FIG. 3 is a block diagram of the superposition integration-subtraction unit 14 and its peripheral part. The superposition integral-subtraction unit 14 includes a transmission acquisition unit 14a, a first addressing unit 14b, a Fourier transform unit 14c, a low-pass filter 14d (indicated as “LPF” in FIG. 3), an inverse Fourier transform unit 14e, and a second addressing. 14f, a subtractor 14g, and a parameter deriving unit 14h.

  The transmission acquisition unit 14 a acquires transmission data (image information) corresponding to the morphological information sent from the γ-ray detector 5 for each subject M. The first addressing unit 14b distinguishes and extracts the background region (air region) and the other region (substantially the region of the subject M) from the transmission data acquired by the transmission acquisition unit 14a.

  On the other hand, the Fourier transform unit 14 c performs superposition integration using Fourier transform on the emission data (projection data) after absorption correction sent from the absorption correction unit 13. The low-pass filter 14d estimates or extracts the scattered data that appears as a low-frequency component with respect to the emission data after the Fourier transform by the Fourier transform unit 14c. The inverse Fourier transform unit 14e performs deconvolution integration using inverse Fourier transform on the scattered component estimated or extracted by the low-pass filter 14d. The second addressing unit 14f is based on the background region extracted by the first addressing unit 14b, and the background region and the low-pass type relating to the original image that is emission data after absorption correction sent from the absorption correction unit 13 Background regions relating to the scattered components after the inverse Fourier transform estimated by the filter 14d are extracted.

  The subtractor 14g derives the difference amount between the emission data (original image) after the absorption correction and the scattering component in the background area extracted by the second addressing unit 14f. The subtractor 14g derives a difference amount between the scattering component extracted by the low-pass filter 14d and the emission data (original image) after absorption correction. The parameter deriving unit 14h derives a parameter that minimizes the difference amount (related to the background area) derived by the subtractor 14g. The parameter corresponds to a physical quantity in the present invention. Further, the difference amount derived by the subtracter 14g is sent to the output unit 9 via the controller 7 as final emission data.

  The transmission acquisition unit 14a corresponds to the form information acquisition unit in the present invention, the first addressing unit 14b corresponds to the first region extraction unit in the present invention, and the low-pass filter 14d includes the scattered component estimation in the present invention. The second addressing unit 14f corresponds to the second region extracting means in the present invention, and the subtractor 14g is the first difference amount deriving means and the second difference amount deriving means in the present invention. The parameter deriving unit 14f corresponds to a physical quantity deriving unit in the present invention. In the first embodiment, the low-pass filter 14d also serves as the scattering component estimation means and the scattering component extraction means in the present invention, and the subtractor 14g has the first difference amount derivation means and the second difference in the present invention. Also serves as a quantity derivation means.

  Next, specific processing of a series of superposition integration-subtraction methods by the superposition integration-subtraction unit 14 will be described with reference to FIGS. 4 is a flowchart of a series of superposition integral-subtraction methods according to the first embodiment, FIG. 5 is a flowchart schematically showing the superposition integral-subtraction method according to the first embodiment, and FIG. FIG. 7 is an explanatory diagram schematically showing extraction of each region from data, FIG. 7 is a schematic diagram of a low-pass filter used for explanation of parameters, and FIG. 8 is a schematic diagram of parameter range setting. It is explanatory drawing represented.

(Step S1) Sinogram creation The emission γ-ray detector 3 collects emission data for each subject M, and the transmission γ-ray detector 5 collects transmission data.

  Specifically, the γ-rays generated from the subject M to which the radiopharmaceutical is administered are converted into light by the scintillator block, and the converted light is photoelectrically converted by the photomultiplier and output to an electrical signal. The γ-ray detector 3 detects γ-rays. The electrical signal is sent to the coincidence counting circuit 10 as image information (pixels), and γ rays are incident simultaneously on two scintillator blocks that are opposed to each other across the subject M, including the coincidence coincidence and the scattering coincidence. The coincidence circuit 10 determines that the sent image information is appropriate data only when the image information is sent. These pieces of image information are sent to the projection data deriving unit 11.

  Further, the γ-ray irradiated from the point source 4 and transmitted through the subject M is converted into light by the scintillator block 5a, and the converted light is photoelectrically converted by the photomultiplier 5b and output to an electric signal. Thus, the γ-ray detector 5 detects γ-rays. The electric signal is sent as image information (pixels) to the absorption correction data deriving unit 12 and is also sent to the transmission acquisition unit 14 a of the superposition integration-subtraction unit 14.

  Based on the image information sent to the projection data deriving unit 11, a sinogram is created with the vertical axis as the γ-ray irradiation direction θ and the horizontal axis as the surface direction orthogonal to the body axis Z of the subject M. Similarly, based on the image information sent to the transmission acquisition unit 14a, a sinogram is created with the vertical axis as the γ-ray irradiation direction θ and the horizontal axis as the surface direction orthogonal to the body axis Z of the subject M.

(Step S2) Conversion in Coronal Direction The sinogram created in step S1 is converted into a cross section in the coronal direction (hereinafter referred to as “coronal cross section”). Here, the coronal section is a horizontal plane parallel to the top plate 1 (see FIG. 1). Therefore, the coronal section is parallel to the body axis Z and is orthogonal to the tomographic plane (slice plane). Of the image that has been coronally transformed from the sinogram, the emission data (projection data) sent to the projection data deriving unit 11 is subjected to absorption correction by the absorption correction unit 13 using transmission data, and the emission data (projection data) after the absorption correction. Is the original image (indicated as “P ₀ ” in FIG. 5), the following will be described.

(Step S3) Extraction of Background Region On the other hand, transmission data sent to the transmission acquisition unit 14a is acquired for each subject M from the images that have been coronally transformed from the sinogram in step S2, and first addressing is performed from the transmission data. The unit 14b extracts the background region (air region) and the other region (substantially the region of the subject M) by distinguishing them.

  Specifically, the transmission data is data obtained by irradiating the subject M from the point source 4 outside the subject M, and has information such as an address. Therefore, as shown in FIG. 6, the background region BG and the other region R can be distinguished and extracted. Since the background region BG is also an air region, the γ-rays irradiated from the point source 4 are detected by the transmission γ-ray detector 5 with almost no absorption. Therefore, the pixel value at the address (pixel) corresponding to the air region is high. On the other hand, since the other region R is also a substantial region of the subject M, the γ-rays irradiated from the point source 4 are absorbed to some extent and detected by the transmission γ-ray detector 5. Therefore, the pixel value at the address (pixel) corresponding to the substantial region of the subject M is low. Therefore, a predetermined threshold value is provided, and a pixel having a pixel value higher than the threshold value is applied as the background region BG, and a pixel having a pixel value lower than the threshold value is applied to other pixels. By applying as the region R, each region can be extracted as shown in FIG.

(Step S4) Fourier transform (FIG. 5 labeled "FFT") Fourier transform section 14c of convolution with the original image P ₀ after coronal converted absorbed correction in Fourier transform step S2 is performed.

(Step S5) Range setting of α and β The low-pass filter 14d estimates or extracts by passing the scattered component to the image of the emission data after the Fourier transform (indicated as “P ₁ ” in FIG. 5). To do. This low-pass filter 14d uses a kernel z (x) represented by the following equation (1) in the frequency domain after Fourier transform.

z (x) = 1.0 / [1.0 + α × exp (β × f × f)] (1)
Where f is an offset frequency (in the frequency domain), α is an amplitude control parameter, β is a band control parameter, and exp is an exponential function. The control parameter α and the band control parameter β are set by a low-pass filter as shown in FIG. The characteristics of the low-pass filter 14d are changed by setting these parameters α and β, and the scattering component estimated or extracted by the low-pass filter 14d is adjusted.

  For this purpose, the ranges of the parameters α and β are set as shown in FIG. If the number of parameter α ranges is m and the number of parameter β ranges is n, steps S6 to S10 described later are repeated m × n times to loop.

(Step S6) LPF Processing Under the condition of certain parameters α and β within the range shown in FIG. 8, the low-pass filter 14d uses the kernel z (x) expressed by the above equation (1) to reduce the low-pass filter 14d. A band-pass process (LPF process) (indicated as “LPF” in FIG. 5) is performed on the image P ₁ of the emission data after Fourier transform. That is, a low-pass process is performed by applying a kernel z (x) in which certain parameters α and β are set to each pixel in the image P ₁ of emission data after Fourier transform.

  When the low-pass processing is performed by sequentially setting the parameters α and β and repeating Steps 6 to S10, the parameters α and β are not reliably determined. The processing performed by the low-pass filter 14d in step S6 at this time is defined as “estimation of scattering components”. On the other hand, when the parameters α and β that minimize the difference amount related to the background area are determined in step S11 (derivation of the parameter that minimizes the difference amount) described later, the parameter α that minimizes the difference amount is determined. The processing by the low-pass filter 14d in step S6 under the condition of β is defined as “extraction of scattered components”.

  That is, when the low-pass processing is performed by sequentially setting the parameters α and β and repeating steps 6 to S10, the low-pass filter 14d in step S6 is the scattering component estimation means in this invention. Steps under the conditions of parameters α and β when the parameters α and β that minimize the difference amount related to the background region are determined in step S11 (derivation of the parameter that minimizes the difference amount). The low-pass filter 14d in S6 fulfills the function of the scattered component extraction means in this invention.

(Step S7) Inverse Fourier Transform Inverse Fourier transform (indicated as “inverse FFT” in FIG. 5) with respect to the scattering component estimated by the low-pass filter 14d in step S6 (indicated as “P ₂ ” in FIG. 5) The inverse Fourier transform unit 14e performs the deconvolution integration using (). By this inverse Fourier transform, in the next steps S8 and S9, the original image P ₀ after the absorption correction and the scattering component after the inverse Fourier transform (FIG. 5) to be extracted by the second addressing unit 14f and subtracted by the subtractor 14g. Then, “P ₃ ” is the same dimension.

(Step S8) Extraction of Background Region Background region BG related to the original image P ₀ after the absorption correction subjected to coronal transformation in step S2 based on the background region BG extracted by the first addressing unit 14b in step S3. and background region BG relates scattered component P ₃ after the inverse Fourier transform by the inverse Fourier transform unit 14e at step S7 estimated in the low-pass filter 14d in step S6 is the second addressing unit 14f extracts respectively .

Specifically, the addresses (pixels) of the background region BG (see FIG. 6) extracted from the transmission data and the other region R (see FIG. 6) are stored in the original image P ₀ and the scattered component P ₃ . Apply each addressing. Then, the region identified as the background region BG (see FIG. 6) in the transmission data is extracted as the background region BG (see FIG. 5) also in the original image P ₀ and the scattered component P ₃ , The region identified as the region R (see FIG. 6) is extracted as the other region (see FIG. 5) in the original image P ₀ and the scattered component P ₃ .

(Step S9) background area respectively extracted background region BG in deriving step S8 of the difference amount with a second addressing unit 14f relates, emission data (original image) after absorption correction P ₀ and the scattered component P ₃ The subtractor 14g derives the difference amount (that is, the difference amount related to the background region). When obtaining the difference amount related to the background area, the difference (subtraction) of pixel values is performed between the same addresses (pixels).

(Step S10) Number of loops?
The parameters α and β are sequentially changed, and steps S6 to S10 are repeatedly looped to determine whether or not the m × n number of loops has been reached. If the number of loops of m × n has not been reached, a combination of parameters α and β is newly set assuming that there is a combination of parameters α and β not set within the range shown in FIG. Return to step S6. Then, step S6 (LPF processing), step S7 (inverse Fourier transform), step S8 (extraction of the background region), step S9 (derivation of the difference amount regarding the background region), and step S10 (loop number?) Are repeated. . If the m × n number of loops has been reached, it is determined that all combinations of the parameters α and β within the range shown in FIG. 8 have been set, and the process proceeds to the next step S11.

(Step S11) Derivation of parameter that minimizes difference amount The parameter deriving unit 14h derives the parameters α and β (combination) that minimize the difference amount related to the background area derived by the subtractor 14g in step S9. . Specifically, each time Steps 6 to S10 are repeated, the difference amounts relating to the background area for each combination of the parameters α and β obtained in Step S9 are respectively compared, and the difference amount for m × n times of the existing combination The parameter derivation unit 14h determines and derives the parameters α and β (indicated as “optimal α, β” in FIG. 5) that are the smallest among the parameters.

Under the conditions in which the parameters α and β are determined in this way, the low-pass processing (LPF processing) by the low-pass filter 14d using the kernel z (x) expressed by the above equation (1) ( In FIG. 5, “LPF” is applied to the image P ₁ of the emission data after the Fourier transform. That is, by applying the kernel z (x) in which the parameters α and β that minimize the difference amount (related to the background region) are applied to each pixel in the image P ₁ of the emission data after Fourier transform. Perform low-pass processing.

(Step S12) Selection from S6 and S7 Actually, step S6 (LPF processing) and step S7 (inverse Fourier transform) have already been performed under the conditions of parameters α and β that minimize the difference amount. Under the conditions of the parameters α and β that minimize the difference amount, the scattering component extracted by the low-pass filter 14d in step S6 (indicated by “P ₂ ′” in FIG. 5) is obtained in step S6. in selecting from the scattered component P ₂ of m × n times estimated by the low-pass filter 14d. Then, under the conditions of the parameters α and β that minimize the difference amount, the scattering component (indicated as “P ₃ ′” in FIG. 5) after the inverse Fourier transform by the inverse Fourier transform unit 14e in step S7, in step S7 in the inverse Fourier transform unit 14e selects from the inverse Fourier transformed m × n times the scattered component P _3.

(Step S13) Derivation of difference amount The subtractor 14g derives the difference amount between the scattering component P ₃ ′ thus selected and the emission data (original image) P ₀ after absorption correction. At this time, the background area is not extracted as in step S8, and is not performed only in the background area as in step S9. The difference amount is included in the entire area including the background area and other areas. Is obtained by the subtractor 14g.

The difference amount derived by the subtractor 14g in step S13 is sent to the output unit 9 via the controller 7 as final emission data (indicated as “P ₄ ” in FIG. 5). Note that when a tomographic image is obtained by reconstruction from projection data related to this emission data, since it is performed on the coronal section after step S2 except for the image after Fourier transformation, it is restored after returning to the sinogram. Make the configuration. This tomographic image is also sent to the output unit 9 via the controller 7.

  According to the PET apparatus according to the first embodiment having the above-described configuration, by including the low-pass filter 14d and the subtractor 14g, the scattering component and the emission data extracted by the low-pass filter 14d ( A difference amount with respect to the original image) is derived, and the difference amount is obtained as final emission data. For example, the conventional superposition integral-subtraction method described above can be realized. When the scattering component is extracted from the emission data (original image) by the low-pass filter 14d, extraction is performed using various physical quantities (parameters α and β in the first embodiment). β) differs depending on the imaging environment of the subject M (individual subject M, imaging location of the subject M). In addition, the background region (air region) and other regions (substantially subject region) vary depending on the individual subject M and the imaging region of the subject M.

  Therefore, since the transmission data obtained by irradiating the subject M from the point source 4 outside the subject M has the form information, first, the transmission acquisition unit 14a first selects the point source outside the subject M. 4, transmission data (morphological information) based on γ rays irradiated from the point source 4 and transmitted through the subject M is acquired for each subject M. Since the transmission data acquired by the transmission acquisition unit 14a has information such as an address, the first addressing unit 14b determines the background region BG (see FIG. 6) and other regions R from the transmission data. (See FIG. 6) can be distinguished and extracted. On the other hand, the low-pass filter 14d estimates the scattering component from the emission data (original image). Based on the background region BG extracted by the first addressing unit 14b described above, the background region BG related to emission data (see FIG. 5) and the background related to the scattering component estimated by the low-pass filter 14d described above. The second addressing unit 14f extracts the region BG (see FIG. 5). In the background regions BG (see FIG. 5) extracted by the second addressing unit 14f, the subtractor 14g derives a difference amount between the emission data (original image) and the scattered component.

  If the physical quantities (parameters α, β) used to extract the scattered component by the low-pass filter 14d described above are optimum values, the emission data is emitted from the background region BG extracted by the two addressing unit 14f. It is assumed that the (original image) and the scattered component are equal. That is, in the background region BG, it is considered that the difference amount between the emission data (original image) and the scattered component is almost “0”. Therefore, the parameter deriving unit 14h derives the physical quantities (parameters α, β) that minimize the difference amount (the original image and the scattered component) derived by the subtractor 14g. The physical quantities (parameters α and β) obtained in this way are values obtained for each imaging environment of the subject M, and are optimized for each imaging environment of the subject M. Therefore, the low-pass filter 14d extracts a scattering component from the emission data (original image) based on the physical quantities (parameters α, β) thus determined, and the scattering extracted by the low-pass filter 14d. By subtracting the difference amount between the component and the emission data (original image) by the subtractor 14g, the difference amount can be obtained as final emission data according to the imaging environment of the subject M, respectively. As a result, it is possible to improve the scattering correction accuracy by the scattering correction by the low-pass filter 14d and the subtractor 14g.

  In the first embodiment, the PET apparatus includes a point source 4, and the point source 4 irradiates γ rays of the same type as the radiopharmaceutical. In the case of the first embodiment, the PET apparatus includes the point source 4, and the point source transmits the same kind of γ-ray as the radiopharmaceutical and transmits through the subject M, so that transmission data (morphological information) is based on the γ-ray. ) Then, using this transmission data (morphological information), the first addressing unit 14a distinguishes and extracts the background region and the other regions, and performs calculations by each means of the superposition integration-subtraction unit 14.

Next, Embodiment 2 of the present invention will be described with reference to the drawings.
FIG. 9 is a side view and block diagram of a PET-CT apparatus including a PET apparatus and an X-ray CT apparatus according to the second embodiment. The PET-CT apparatus corresponds to the diagnostic system in the present invention.

  In the first embodiment described above, the PET apparatus includes the point source 4, the point source 4 emits the same γ-ray as the radiopharmaceutical and passes through the subject M, and the γ-ray detector 5 detects the γ-ray. Thus, transmission data was collected as morphological information based on the radiation, but in the second embodiment, CT projection data obtained from an X-ray CT apparatus is used as morphological information. The projection data for CT corresponds to the form information in the present invention, and also corresponds to the data for X-ray CT.

  The X-ray CT apparatus includes a gantry 31 having an opening 31a, an X-ray tube 32, and an X-ray detector 33. The X-ray tube 32 and the X-ray detector 33 are arranged to face each other with the subject M interposed therebetween, and are embedded in the gantry 31. A large number of detection elements constituting the X-ray detector 33 are arranged in a fan shape around the body axis Z of the subject M.

  In addition, the X-ray CT apparatus includes a gantry driving unit 34, a high voltage generation unit 35, a collimator driving unit 36, and a CT reconstruction unit 37. The CT reconstruction unit 37 is realized by the controller 7 executing, for example, a program stored in a ROM of a storage medium represented by the above-described memory unit 15 or the like or an instruction input from the input unit 8. Note that CT projection data, which will be described later, and CT tomographic images processed by the CT reconstruction unit 37 are also written and stored in the RAM of the memory unit 15 in the same manner as in the first embodiment described above. Read from RAM.

  The gantry drive unit 34 is a mechanism that drives the X-ray tube 32 and the X-ray tube detector 33 to rotate around the body axis Z of the subject M within the gantry 31 while maintaining the opposing relationship with each other. The motor is composed of a motor (not shown).

  The high voltage generator 35 generates a tube voltage or tube current of the X-ray tube 32. The collimator driving unit 36 is a mechanism that sets an X-ray irradiation field and drives the collimator (not shown) adjacent to the X-ray tube 32 to move in the horizontal direction. It consists of

  In the case of the indirect conversion type X-ray detector 33, the X-ray irradiated from the X-ray tube 32 and transmitted through the subject M is converted into light by a scintillator (not shown) in the X-ray detector 33. The light-sensitive film (not shown) photoelectrically converts the converted light and outputs it as an electrical signal. In the case of the direct conversion type X-ray detector 33, the X-ray is directly converted into an electric signal by a radiation sensitive film (not shown) and output. The electrical signal is sent to the CT reconstruction unit 37 as image information (pixels). The image information sent to the CT reconstruction unit 37 is transmitted as projection data for CT.

  The projection data for CT has the same form information as the transmission data described in the first embodiment. In the second embodiment, the absorption correction data deriving unit is used to use the projection data for CT as transmission data. 12 and the transmission acquisition unit 14a (not shown in FIG. 9, refer to FIG. 3) of the superposition integral-subtraction unit 14 and also the CT reconstruction unit 37.

  Image information (CT projection data) sent to the CT reconstruction unit 37 is reconstructed to obtain a CT tomographic image. This CT tomographic image is sent to the output unit 9 via the controller 7.

  Since the functions of the processing unit (absorption correction unit 13) at the subsequent stage of the PET apparatus including the absorption correction data deriving unit 12 and the superposition integral-subtraction unit 14 are the same as those in the first embodiment, description thereof is omitted. The tomographic image for PET reconstructed through the superposition integration-subtraction unit 14 and the CT tomographic image reconstructed by the CT reconstruction unit 37 are superimposed and output by the output unit 9. Also good.

  However, in the first embodiment, the transmission acquisition unit 14a of the superimposition integration / subtraction unit 14 acquires, for each subject M, the transmission based on the γ rays irradiated from the point source 4 and transmitted through the subject M as morphological information. On the other hand, in the second embodiment, X-ray CT data is acquired as morphological information for each subject M.

  As described above, in the second embodiment, the projection data for CT obtained by being detected by the X-ray detector 33 of the X-ray CT apparatus is sent to the CT reconstruction unit 37 and the absorption correction data deriving unit 12 or The data is sent to the transmission acquisition unit 14a of the superposition integration-subtraction unit 14, and the projection data for CT is used as the form information.

  According to the diagnosis system of the PET-CT apparatus according to the second embodiment having the above-described configuration, as in the first embodiment, the physical quantities (parameters α and β) derived through the respective means of the superposition integral-subtraction unit 14 are used. ) Is a value obtained for each imaging environment of the subject M, and is optimized for each imaging environment of the subject M. Therefore, based on the physical quantities (parameters α, β) obtained in this way, a difference amount between the scattering component and the emission data (original image) is obtained as final emission data according to the imaging environment of the subject M, respectively. The scattering correction accuracy by the scattering correction by the low-pass filter 14d and the subtractor 14g can be improved.

  In the case of the second embodiment, in the X-ray CT apparatus, X-ray CT data is obtained based on the X-rays irradiated from the outside of the subject M and transmitted through the subject M, and the X-ray CT data is obtained as morphological information. As described above, the transmission acquisition unit 14a of the superposition integration-subtraction unit 14 acquires for each subject M. Then, using this form information, the first addressing unit 14b (not shown in FIG. 9, refer to FIG. 3) distinguishes and extracts the background region and other regions, Calculation by means is performed.

  In the second embodiment, the PET apparatus and the X-ray CT apparatus are integrated into one diagnostic system. However, the X-ray CT apparatus is configured as an external apparatus of the PET apparatus, and the PET apparatus can be obtained from the X-ray CT apparatus. The form information (CT projection data used as the form information in the second embodiment) may be transferred. In this case, the X-ray CT apparatus is an external apparatus when viewed from the PET apparatus.

  The present invention is not limited to the above-described embodiment, and can be modified as follows.

  (1) In each of the above-described embodiments, the PET apparatus has been described as an example. However, the present invention detects a single gamma ray and reconstructs a tomographic image of the subject. It can also be applied to devices.

  (2) In each of the above-described embodiments, the γ-ray detector 5 for transmission is a stationary type that detects γ-rays while still, but the γ-ray detector 5 rotates around the subject M. A rotary type that detects γ rays may be used.

  (3) In each of the above-described embodiments, the form information (transmission data in the first embodiment and X-ray CT data in the second embodiment) is also used for the absorption correction. However, it is not always necessary to perform the absorption correction. . Further, although the mission data after absorption correction is used as an object of the scattering component subtraction process as an original image, the emission data that is not subjected to the absorption correction may be used as an original image and an object of the scattering component subtraction process.

  (4) In each of the embodiments described above, the low-pass filter 14d also serves as the scattering component estimation means and the scattering component extraction means in the present invention, and the subtractor 14g has the first difference amount derivation means and the Although the two difference amount derivation means is also used, a low-pass filter corresponding to the scattering component estimation means and a low-pass filter corresponding to the scattering component extraction means may be separately provided, or the first difference amount derivation may be provided. A subtractor corresponding to the means and a subtractor corresponding to the second difference amount derivation means may be provided separately.

1 is a side view and block diagram of a PET (Positron Emission Tomography) apparatus according to Embodiment 1. FIG. FIG. 3 is a layout diagram of γ-ray detectors in a PET apparatus. It is a block diagram of a superposition integral-subtraction part and its peripheral part. 3 is a flowchart of a series of superposition integral-subtraction methods according to the first embodiment. 3 is a flowchart schematically showing a superposition integral-subtraction method according to the first embodiment. It is explanatory drawing which represented typically extraction of each area | region from transmission data. It is a schematic diagram of the low-pass filter for explanation of parameters. It is explanatory drawing which represented the range setting of the parameter typically. It is the side view and block diagram of PET-CT apparatus provided with the PET apparatus and X-ray CT apparatus which concern on Example 2. FIG. It is the flow which showed the conventional superposition integral-subtraction method roughly.

Explanation of symbols

4 ... Point source 14a ... Transmission acquisition unit 14b ... First addressing unit 14d ... Low-pass filter (LPF)
14f ... second addressing unit 14g ... subtractor 14h ... parameter deriving unit BG ... background region α, β ... parameter M ... subject

Claims (3)

  A nuclear medicine diagnostic apparatus that obtains nuclear medicine data of a subject based on radiation generated from a subject to which a radiopharmaceutical is administered, and (A) an external radiation source that is external to the subject. Morphological information acquisition means for acquiring, for each subject, morphological information based on the radiation irradiated from and transmitted through the subject, and (B) a background region and others from the morphological information acquired by the morphological information acquisition means First region extracting means for distinguishing and extracting the region, (C) scattering component estimating means for estimating a scattering component from the nuclear medicine data, and (D) the background region extracted by the first region extracting means To extract a background region relating to nuclear medicine data and a background region relating to the scattering component estimated by the scattering component estimation means, respectively. (E) first difference amount deriving means for deriving a difference amount between the nuclear medicine data and the scattered component in the background region extracted by the second region extracting means, and (F) A physical quantity deriving means for deriving a physical quantity that minimizes the difference quantity derived by the first difference quantity deriving means, and (a) scattering from nuclear medicine data based on the physical quantity derived by the physical quantity deriving means. And (b) a second difference amount deriving unit for deriving a difference amount between the scattered component extracted by the scattering component extracting unit and the nuclear medicine data, the second difference A nuclear medicine diagnostic apparatus characterized in that the difference amount derived by a quantity derivation means is obtained as final nuclear medicine data.   The nuclear medicine diagnosis apparatus according to claim 1, further comprising the external radiation source, wherein the external radiation source emits radiation of the same type as the radiopharmaceutical.   A diagnostic system for use in a nuclear medicine diagnostic apparatus, comprising a nuclear medicine diagnostic apparatus and an X-ray CT apparatus. The nuclear medicine diagnostic apparatus detects radiation generated from a subject to which a radiopharmaceutical is administered. The X-ray CT apparatus obtains X-ray CT data based on X-rays that are irradiated from the outside of the subject and transmitted through the subject. (A ′) morphological information acquisition means for acquiring the X-ray CT data as morphological information for each subject, and (B) a background region and the others from the morphological information acquired by the morphological information acquisition means A first region extracting means for distinguishing and extracting the region, (C) a scattering component estimating means for estimating a scattering component from the nuclear medicine data, and (D) the background extracted by the first region extracting means A second area extracting means for extracting a background area relating to nuclear medicine data and a background area relating to the scattering component estimated by the scattering component estimating means based on the area; and (E) the second area extracting means respectively. A first difference amount deriving unit for deriving a difference amount between the nuclear medicine data and the scattering component in the extracted background region; and (F) the difference amount derived by the first difference amount deriving unit is a minimum. A physical quantity deriving means for deriving a physical quantity to become, (a) a scattering component extracting means for extracting a scattering component from nuclear medicine data based on the physical quantity derived by the physical quantity deriving means, and (b) the scattering component Second difference amount deriving means for deriving a difference amount between the scattered component extracted by the extracting means and the nuclear medicine data, and derived by the second difference amount deriving means. The obtained difference amount is obtained as final nuclear medicine data.

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