JP2007218769A - Nuclear medicine imaging equipment - Google Patents

Abstract

The present invention provides a nuclear medicine imaging apparatus capable of correcting scattered radiation with high accuracy in accordance with an actual imaging environment.
Emission data detected by a coincidence circuit is discriminated by an energy discriminating means into SEW data within a standard energy window (SEW) and UEW data within a high energy window (UEW). . Scattering component removing means 13 removes scattered radiation from the UEW data, performs predetermined amplification, obtains estimated SEW data, subtracts the estimated SEW data from the SEW data, and performs smoothing to obtain accurate SEW data. Acquire high scattered radiation data. Then, by subtracting the scattered r-ray data from the SEW data, SEW data from which the scattered component has been removed is obtained and reconstructed by the amplification factor storage means 14 to display a high-quality RI distribution image with few scattered rays. It is displayed on the means 16.
[Selection] Figure 1

Description

  The present invention relates to a nuclear medicine imaging apparatus, and more particularly to a nuclear medicine imaging apparatus such as a scintillation camera (gamma camera) or a SPECT (Single Photon Emission Computed Tomograph) apparatus configured using this camera.

Nuclear medicine imaging devices such as scintillation cameras and SPECTs, when radioactive isotopes (eg 15O, 18F, 11C, etc.) (RI) (RI) are administered to a subject and accumulated in a specific organ, etc., from inside the body to outside the body The emitted radiation is detected and emission data is collected to obtain an image.

Among such nuclear medicine imaging devices, PET (Positron Emission
Tomography) detects radiation using a detector array consisting of a number of detectors arranged in a ring around the area to be examined, and uses a computer to calculate the radiation source in the same manner as a normal CT apparatus. It is specified in the plane and an image of the test part is created. In addition, in order to capture a three-dimensional image of a test part within a short time, a PET apparatus that forms a cylindrical detector stack in which a large number of ring-shaped detector arrays are stacked is also widely used.

In such a PET apparatus, two gamma rays radiated in opposite directions by 180 ° due to pair annihilation with positron electrons generated by decay of radioisotopes in the body are sent to a large number of detectors arranged around the test portion. Emission data is collected by counting simultaneously.

However, the detected emission data includes not only those based on true events, which are two gamma rays radiating in the opposite directions of 180 °, but also those based on scattering events based on gamma rays due to Compton scattering, and the same pair annihilation. There are those based on contingent events that are gamma rays that do not occur and are coincidentally counted. These scattering and accidental events cause noise in nuclear medicine imaging.

For this reason, corrections are made for scattering events that cause noise. For example, a method is used in which the scattered radiation characteristics of a point source are obtained in advance and the scattered components are subtracted by deconvolution integration with measured projection data. . As shown in FIG. 6, such scattered radiation characteristics are obtained by disposing a point source R in a known position in advance and detecting gamma rays generated from the point source R around the detector 20 for a predetermined period. Can be obtained.

That is, FIG. 7 illustrates a state in which gamma rays are incident on a detector 20 including a scintillator block 20a and a photomultiplier (FP-PMT) 20b, and Compton scattering occurs. Gamma rays are detected at two locations in the block 20a. In such a case, the gamma rays detected after Compton scattering become a scattering component, but as shown in FIG. 6, since the position of the point source R is known, whether or not the detected data is a true event and a scattering event. Can be determined.

In addition to the methods described above, there are methods for measuring the wave height distribution of each detector and subtracting the scattering window count value from the photoelectric peak, and correction methods such as obtaining and subtracting the scattering distribution from the simulation based on the model. Exists (Patent Document 1, Patent Document 2).

Scattercorrection for three-dimensional PET based on an analytic model dependent onsource and attenuating objectL G Hiltz et al 1994 Phys. Med. Biol. 392059-2071 Scattermodelling and correction strategies in fully 3-D PETH.ZAIDI * Nuclear Medicine Communications, 2001,22, 1181 ± 1184

All of the above-described methods for correcting scattered radiation used in conventional nuclear medicine imaging apparatuses are approximate methods, and are based on data acquired in an environment different from the actual imaging environment. It ’s impossible. In particular, in recent years, a three-dimensional positron emission CT apparatus (3D-PET) in which an inter-ring septum has been removed has been proposed in order to improve sensitivity. In such a 3D-PET, a large amount of scattering components from outside the field of view are mixed. In addition, the conventional method described above has a problem in that it cannot correct the scattered component from outside the field of view.

Accordingly, an object of the present invention is to provide a nuclear medicine imaging apparatus capable of correcting scattered radiation with high accuracy according to an actual imaging environment.

The nuclear medicine imaging apparatus according to claim 1, wherein a detection means for detecting radiation generated from a subject to which a medicine composed of a radioisotope (RI) is administered, and radiation detected simultaneously by the detection means as emission data Detecting simultaneous counting means, SEW data within a standard energy window (SEW) for obtaining an RI distribution image with respect to the emission data detected by the simultaneous counting means, and a reference value set within the SEW range Energy discriminating means for discriminating into UEW data within the above high energy window (UEW), storage means for storing an amplification factor which is an event ratio between the SEW data and the UEW data, and a scattered component from the UEW data The scattered component removing means 1 for removing, and the UEW data from which the scattered component has been removed After amplifying with the amplification factor stored in the storage means, subtracting from the SEW data and performing a smoothing process, a scattering component calculation means for obtaining a scattering component included in the SEW data, and the scattering from the SEW data A scattering component removing unit 2 that removes the scattering component using the scattering component calculated by the component calculating unit, and obtaining a radioisotope distribution image based on the SEW data from which the scattering component has been removed. Features.

The nuclear medicine imaging apparatus according to claim 2, wherein a detection means for detecting radiation generated from a subject to which a medicine composed of a radioisotope (RI) is administered, and radiation detected simultaneously by the detection means as emission data A coincidence unit for detection, and a high energy window (UEW) equal to or higher than a reference value set within a standard energy window (SEW) for obtaining a normal RI distribution image for the emission data detected by the coincidence unit ) Energy discriminating means for discriminating the UEW data within, a storage means for storing an amplification factor which is an event ratio between the SEW data and the UEW data within a standard energy window (SEW) range, and the UEW data The scattering component removing means 1 for removing the scattering component and the scattering component are removed. The UEW data, calculating means for amplifying an amplification factor stored in the storage means, and obtaining a distribution image of a radioisotope on the basis of the UEW data processed by said arithmetic means.

The nuclear medicine imaging apparatus according to claim 3 is characterized in that the calculation means according to claim 2 performs smoothing processing after amplifying the UEW data.

The nuclear medicine imaging apparatus according to claim 4, wherein the reference value set within the SEW range is an energy value at which a scattered radiation component of the radiation decreases. To do.

According to a fifth aspect of the present invention, in the nuclear medicine imaging apparatus according to the first to third aspects, the reference value set within the SEW range is a photopeak value of the radiation.

A nuclear medicine imaging apparatus according to a sixth aspect of the present invention is the nuclear medicine imaging apparatus according to the first to third aspects, wherein the radiation is a gamma ray, and the SEW is 300 to 400 kev or more and 600 to 800 kev or less, and is set within the range of the SEW. The reference value is from 500 kev to 600 kev.

A nuclear medicine imaging apparatus according to claim 7, wherein the radiation is gamma rays, the SEW is 300 to 400 kev or more, 600 kev to 800 kev or less, and the reference value of the UEW is 511 kev. 600 keV.

The nuclear medicine imaging apparatus according to claim 8 is the nuclear medicine imaging apparatus according to any one of claims 1 to 7, wherein the scattering component removing unit 1 performs a Fourier transform on the UEW data, performs a low frequency filter process, It is characterized in that scattered radiation is removed by inverse Fourier transform.

The invention of claim 1 removes scattered radiation from the UEW data within the high energy window (UEW) with less scattered radiation from the emission data detected by coincidence counting, and therefore removes scattered radiation with higher accuracy than usual. Is possible. Therefore, it is possible to obtain highly accurate scattered radiation data for SEW data by amplifying UEW data from which scattered radiation has been removed, estimating SEW data, subtracting this from SEW data, and performing smoothing. it can. By subtracting the scattered radiation data from the SEW data, correction using the scattered radiation information included in the actual imaging data can be performed, so that SEW data from which scattered radiation is accurately removed according to the actual imaging environment is obtained. Therefore, a high-quality RI distribution image with few scattered rays can be obtained in the nuclear medicine imaging apparatus.

The invention of claim 2 can amplify UEW data from which scattered radiation has been removed with high accuracy and estimate SEW data, so that an RI distribution image can be easily obtained from relatively highly accurate SEW data from which scattered radiation has been removed. It is done.

According to the third aspect of the present invention, since the smoothing process is performed when the SEW data is estimated after the UEW data is amplified, it is possible to obtain SEW data with less noise.

In the invention of claim 4, since the reference value set within the SEW range is an energy value that reduces the scattered radiation component of radiation, UEW data in which only minute scattered radiation is mixed can be obtained. By performing the scattered radiation removal process, UEW data from which scattered radiation has been removed with higher accuracy can be obtained.

In the invention of claim 5, since the reference value set within the range of SEW is the photo peak value of radiation, UEW data in which only minute scattered radiation is mixed can be obtained, and scattered radiation removal processing is applied to this. As a result, UEW data from which scattered radiation has been removed with higher accuracy can be obtained.

The invention of claim 6 is the invention according to claims 1 to 3, wherein the radiation is a gamma ray, the SEW is 300 kev to 400 kev or more, 600 kev to 800 kev or less, and a reference value set within the SEW range is a gamma ray Since the photo peak value is in the vicinity of 512 kev from 500 kev to 600 kev, UEW data in which only minute scattered rays are mixed can be obtained. And UEW data from which scattered radiation was removed with higher accuracy can be obtained by subjecting this to scattered radiation removal processing.

A seventh aspect of the present invention is the first to third aspects, wherein the radiation is a gamma ray, the SEW is from 300 kev to 400 kev and from 600 kev to 800 kev, and a reference value set within the SEW range is a gamma ray. Since the photo peak value of 511 kev is 600 kev, which is closer to the vicinity of 512 kev, which is the photo peak value, UEW data in which only smaller scattered rays are mixed can be obtained.

The invention according to claim 8 is the invention according to claims 1 to 7, wherein the scattering component removal means 1 performs a Fourier transform on the UEW data, performs a low frequency filter process, and further performs an inverse Fourier transform on these. Therefore, it becomes possible to remove scattered radiation more efficiently.

A nuclear medicine imaging apparatus according to an embodiment of the present invention will be described using a PET apparatus as an example. FIG. 1 is a schematic diagram showing a detection portion and a data processing portion of the PET apparatus of the present embodiment. In FIG. 1, a large number of detection units 20 are arranged around a subject 30. As shown in FIG. 2, a scintillator block 20a composed of 64 crystals and four photomultipliers when connected to the scintillator block 20a. It consists of pliers 20b. The gamma rays emitted from the subject 30 into which radioactive isotopes (for example, 15O, 18F, 11C, etc.) are introduced are incident on any crystal of the scintillator block 20a, and the position of the light emitted by the incidence of the gamma rays. Is detected by the photomultiplier 20b, the incident position of the gamma rays is detected as an event.

In FIG. 1, the gamma rays emitted from the subject 30 are detected by detectors 20 arranged around them, but the coincidence counting circuit 11 is used in the detectors facing each other at the same time or within a predetermined time window. Only the gamma rays detected in the above are regarded as true events, acquired as appropriate data, and stored in a storage means (not shown).

The energy discriminating means 12 uses the obtained emission data as SEW data within a standard energy window (SEW) for obtaining a normal radioisotope distribution image (hereinafter referred to as “RI distribution image”), and the SEW. Is discriminated from UEW data within a high energy window (UEW) equal to or higher than a reference value set within a range of.

  The energy discriminating means 12 itself has a well-known configuration. For example, by storing the relationship between the detector output and energy in advance, the detected emission data is discriminated in comparison with this. Yes.

FIG. 3 shows an example of a distribution diagram with the horizontal axis representing the energy of gamma rays and the vertical axis representing the number of events for the emission data acquired by coincidence. The standard energy window (SEW) is the lower limit. The value Sfrom is set from 300 kev to 400 kev, and the upper limit value Sto is set from 600 kev to 800 kev. In the high energy window (UEW), the lower limit value Ufrom is set in the vicinity of 512 Kev, and the upper limit value is set to the upper limit value Sto of the standard energy window (SEW). Note that the energy of the emitted gamma rays approximately shows the distribution shown in FIG.

In addition, the lower limit value Ufrom of UEW may be appropriately selected as an energy value at which the gamma ray scattering component is reduced to some extent, such as a value larger than around 512 kev which is a photo peak value of gamma rays, for example, 500 kev, 511 kev, or 600 kev. This is because when the gamma rays cause Compton scattering, the energy of the gamma rays after the scattering decreases, so that if the energy is larger than around 512 kev which is the photo peak value of the gamma rays, the scattered rays are further reduced. The photo peak value is an energy value at which the number of emitted radiation events reaches a peak.

The SEW data and UEW data discriminated by the energy discriminating means 12 are stored in the SEW memory 12a and the UEW memory 12b of the energy discriminating means 12. FIG. 4 shows the SEW of one of the simultaneously counted detector pairs when the horizontal axis is the energy of gamma rays detected by one detector and the vertical axis is the energy of gamma rays detected by the other detector. The relationship between the gamma ray discriminated as data and the gamma ray discriminated as UEW data is shown. As shown in FIG. 4, the UEW data is discriminated when both the simultaneously counted gamma rays exceed the lower limit value Ufrom of the high energy window (UEW).

  The scattering component removal means 13 removes the scattering component based on a scattered ray from SEW data using UEW data. Specific processing contents will be described with reference to FIG.

FIG. 5 shows a profile of the emission data of the phantom, which was obtained by photographing a test phantom in which a radioisotope was previously injected with the PET apparatus shown in FIG. Yes. In the left column of FIG. 5, UEW data and SEW data of the one profile obtained by simultaneous counting by the coincidence counting circuit 11 and energy discrimination by the energy discriminating means 12 are shown as original data.

First, the scattering component removal means 13 performs the removal of the scattering component contained in UEW data by the Deconvolution method which is a kind of a simple correction method. Specifically, the UEW data is subjected to Fourier transform, a high frequency component is removed by a low frequency filter, and this is then subjected to inverse Fourier transform to obtain a scattered component (P1). And the true UEW data from which the scattering component was removed is obtained by subtracting the calculated | required scattering component from original UEW data (P2).

Next, the scattering component removing unit 13 reads the amplification factor f, which is the ratio of the number of events of the SEW data and the UEW data, from the amplification factor storage unit 14, and amplifies the true UEW data with the amplification factor f (P3). Here, the amplification factor f is obtained in advance as a ratio of the number of events of SEW data and UEW data, and is stored in the amplification factor storage means 14. For example, instead of the subject 30 in FIG. 1, the amplification factor f is set as shown in FIG. 6 by arranging a radioisotope point source R and detecting gamma rays emitted by the detector 20 for a predetermined time, The SEW data and the UEW data are obtained from the emission data simultaneously counted by the coincidence circuit 11 by the energy discriminating means 12, and can be easily obtained by measuring the event number ratio.

As described above, the amplification factor f is acquired based on the ratio of the number of events of the SEW data and UEW data acquired in advance, but the ratio of the obtained SEW data to UEW data is also the actual shooting. Therefore, the calculated amplified data f × UEW data can be considered as estimated data of SEW data from which scattered radiation has been removed. However, the estimated SEW data here amplifies the UEW data with a small number of events, and thus may include noise.

In order to solve the noise problem, a true scattered ray component is further obtained from the SEW data, and this is subtracted from the SEW data. That is, first, the scattering component included in the SEW data is obtained by subtracting the amplified data f × UEW data obtained in P3 from the SEW data (P4). At this stage, since the scattered component of the SEW data includes noise due to amplification processing, smoothing processing is performed to remove high-frequency components (P5). As a result, an accurate scattering component due to the scattered radiation contained in the SEW data is obtained. By subtracting this from the SEW data, highly accurate SEW data from which the scattering component has been removed is obtained (P6).

In FIG. 1, an image reconstruction unit 15 performs a known reconstruction process to obtain an RI distribution image. More specifically, the SEW data from which the scattering component has been removed by the scattering component removing unit 13 is subjected to convolution and back projection processing to obtain an RI distribution image. The obtained RI distribution image is displayed on the display means 16 such as a CRT.

If the estimated SEW data obtained at the stage P3 in FIG. 5 is subjected to image reconstruction by the image reconstruction means 15 and an RI distribution image is obtained, a simple process is performed although there is some noise component. Thus, an RI distribution image can be obtained based on SEW data from which scattered radiation has been removed. Further, if smoothing processing is performed on the estimated SEW data obtained in the P3 stage and then image reconstruction is performed, an RI distribution image with less noise can be obtained.

  In the above embodiment, the case where the present invention is applied to a PET apparatus has been described. However, the present invention is generally applied to nuclear medicine imaging apparatuses such as a gamma camera apparatus, which detect gamma rays emitted from a subject and acquire an RI distribution image. It goes without saying that the invention is applied. Moreover, although gamma rays are shown as radiation emitted from the subject, it goes without saying that the present invention can be applied to other radiation.

SPECT (Single) composed of a scintillation camera (gamma camera) and this camera
The present invention can be used in nuclear medicine imaging apparatuses such as Photon Emission Computed Tomograph apparatus.

It is the schematic which shows the PET apparatus which is one form of the nuclear medicine imaging apparatus concerning this invention. It is a figure which shows the detector of the gamma ray used for a nuclear medicine imaging apparatus. It is a figure which shows the energy distribution figure of emission data. It is a figure which shows the distribution map of the event with respect to the energy of emission data. It is a flowchart for demonstrating the scattered-radiation removal process of this invention. It is a figure which shows a mode that the gamma ray emitted by a point source is detected with a detector. It is a figure for demonstrating the Compton scattering of a gamma ray.

Explanation of symbols

11
Simultaneous counting circuit 12
Energy discrimination means 13
Scattered ray removing means 14
Amplification rate storage means 15
Image reconstruction means 16
Display means 20
Gamma ray detector

Claims (8)

A detection means for detecting radiation generated from a subject to which a drug composed of a radioisotope (RI) is administered;
Simultaneous counting means for detecting radiation simultaneously detected by the detection means as emission data;
Regarding the emission data detected by the coincidence means, SEW data within a standard energy window (SEW) for obtaining an RI distribution image, and a high energy window (UEW above a reference value set within the range of the SEW) Energy discriminating means for discriminating into UEW data within;
Storage means for storing an amplification factor which is an event ratio between the SEW data and the UEW data;
Scattering component removal means 1 for removing scattering components from the UEW data;
After the UEW data from which the scattered component has been removed is amplified with the amplification factor stored in the storage means, the scattered component is obtained by subtracting from the SEW data and performing a smoothing process to obtain the scattered component included in the SEW data A calculation means;
Scattering component removal means 2 for removing the scattering component from the SEW data using the scattering component calculated by the scattering component calculation means,
A nuclear medicine imaging apparatus characterized in that a radioisotope distribution image is obtained based on SEW data from which the scattering component has been removed. A detection means for detecting radiation generated from a subject to which a drug composed of a radioisotope (RI) is administered;
Simultaneous counting means for detecting radiation simultaneously detected by the detection means as emission data;
Regarding the emission data detected by the coincidence means, UEW data within a high energy window (UEW) that is equal to or higher than a reference value set within a standard energy window (SEW) range for obtaining a normal RI distribution image. Energy discrimination means for discrimination;
Storage means for storing an amplification factor which is an event ratio between the SEW data and the UEW data within a standard energy window (SEW);
Scattering component removal means 1 for removing scattering components from the UEW data;
Arithmetic means for amplifying the UEW data from which the scattering component has been removed at an amplification factor stored in the storage means;
A nuclear medicine imaging apparatus, wherein a radioisotope distribution image is obtained based on UEW data processed by the computing means. The nuclear medicine imaging apparatus according to claim 2, wherein the arithmetic means performs smoothing processing after amplifying the UEW data. 4. The nuclear medicine imaging apparatus according to claim 1, wherein the reference value set within the SEW range is an energy value at which a scattered radiation component of the radiation decreases. 4. The nuclear medicine imaging apparatus according to claim 1, wherein the reference value set within the SEW range is a photo peak value of the radiation. 2. The radiation according to claim 1, wherein the radiation is a gamma ray, the SEW is 300 kev to 400 kev or more and 600 kev to 800 kev or less, and a reference value set within the SEW range is 500 kev to 600 kev. 3. The nuclear medicine imaging apparatus according to 3. 2. The radiation according to claim 1, wherein the radiation is gamma rays, the SEW is 300 kev to 400 kev or more, 600 kev to 800 kev or less, and a reference value set within the SEW range is 511 kev to 600 kev. 3. The nuclear medicine imaging apparatus according to 3. 2. The scattered component removing unit 1 performs a Fourier transform on the UEW data, performs a low frequency filter process, and further performs an inverse Fourier transform on the UEW data to remove scattered rays. The nuclear medicine imaging apparatus according to claim 7.