JP2014137279A - Nuclear medicine diagnosis device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To remove scattered radiation noise intruding due to insufficient energy resolution in a nuclear medicine diagnosis device, and to enable a radiation detector that has not been used in a conventional nuclear medicine diagnosis device because of insufficient energy resolution to be used in the nuclear medicine diagnosis device.SOLUTION: The nuclear medicine diagnosis device includes both a radiation detector used mainly for measurement and a radiation detector whose energy measurement accuracy is higher than that of the former radiation detector. Thus, the information amount used by a scattered radiation correction operation of removing the scattered radiation noise is increased, thereby improving the correction accuracy.

Description

  The present invention relates to a method for removing scattered radiation noise from an image measured by a nuclear medicine diagnostic apparatus.

  In Patent Document 1, a high S / N ratio is realized by synchronously measuring γ-ray signals measured at the same time by an organic scintillator and an inorganic scintillator and bundling them as one gamma-ray signal.

WO2009 / 157526

  In Patent Literature 1, two types of scintillator signals are combined and noise is removed from the simultaneous measurement signals of the two types of scintillators using a detection window by synchronous measurement. However, other than the method using the simultaneous measurement is considered. It has not been.

  In the present invention, synchronous measurement is not performed between the signals of the first detector and the second detector, and they are handled independently, and the scattered radiation noise that is mixed because the energy resolution is not sufficient is removed. In addition, since the energy resolution is not sufficient, a radiation detector that has not been used in a conventional nuclear medicine diagnostic apparatus can be used in the nuclear medicine diagnostic apparatus.

  In the nuclear medicine diagnostic apparatus, the first type radiation detector mainly used for measurement and the second type radiation detector having higher energy measurement accuracy than the first type radiation detector are used in combination. Synchronous measurement is not performed between the signals of the seed radiation detector and the second type radiation detector, and they are handled independently, and the measurement data of the second type radiation detector is obtained by the scattered radiation correction calculation for removing scattered radiation noise. It is characterized by using.

  In the nuclear medicine diagnostic apparatus according to the present invention, the scattered radiation noise that is mixed because the energy resolution is not sufficient, among the multiple types of detectors used, between the first type radiation detector and the second type radiation detector. Synchronous measurement is not performed between signals, but they are handled independently, and correction can be made based on information measured by a detector having a relatively high energy resolution. In addition, since the energy resolution is not sufficient, a radiation detector that has not been conventionally used in a nuclear medicine diagnostic apparatus can be used in the nuclear medicine diagnostic apparatus.

Outline diagram of scattered radiation Nuclear medicine imaging system Overview of the first and second embodiments Detector unit internal configuration example 1 Detector unit internal configuration example 2 Detector unit internal configuration example 3 Outline diagram of the third embodiment Example of detector arrangement Energy spectrum example Fourth embodiment outline diagram Overview diagram of the fifth embodiment and overview diagram of the sixth embodiment Detector head internal configuration example 1 Detector head internal configuration example 2 Detector head internal configuration example 3 Overview of the seventh embodiment Overview of the eighth embodiment

  Embodiments will be described below with reference to the drawings.

  Nuclear medicine diagnostic devices such as positron emission tomography (PET) and single photon emission tomography (SPECT) are derived from radioactive drugs injected into a subject. A gamma ray emitted from the subject's body is detected by a radiation detector, and the distribution of the radiopharmaceutical in the subject's body is tomographically imaged using measurement information obtained based on the detected gamma ray detection signal. . The obtained tomographic image information is used for diagnosis of metabolic function and physiological function.

Examination using a PET device (PET examination) is a radiopharmaceutical (hereinafter referred to as PET) that is labeled with a positron emitting nuclide ( ¹⁵ O, ¹³ N, ¹¹ C, ¹⁸ F, etc.) and specifically accumulates at a specific site in the body. (Referred to as drug) is administered to the subject, and the distribution of the PET drug in the body of the subject is examined. The positrons emitted from the PET drug administered to the subject combine with electrons in the vicinity of the body to annihilate positrons and emit a pair of gamma rays having energy of 511 KeV (hereinafter referred to as anti-gamma rays). Since each gamma ray is emitted in almost opposite directions, both gamma rays are measured by the respective radiation detectors, and the information obtained based on the respective gamma ray detection signals output from these detectors is obtained. By using and simultaneously measuring, it is possible to specify on which straight line the positron annihilation event occurred. After detecting a statistically sufficient number of anti-gamma rays in this way, using an image reconstruction algorithm such as OS-EM method, the frequency distribution of anti-gamma rays, that is, the PET drug in the body of the subject. The distribution can be tomographically imaged.

  The above-described measurement of gamma rays caused by PET in the body is called emission measurement, and image information reconstructed based on the gamma ray detection signal obtained by the emission measurement is called an emission image. In the emission measurement, ideally, a direct gamma ray pair incident on the detector without interference with other substances is detected simultaneously in the time window, and is simultaneously counted. However, there is a case where the gamma ray causes an interaction such as Compton scattering with a substance such as an object, and the traveling direction is refracted and the scattered ray measured is erroneously detected (FIG. 1). Since the scattered rays drop energy when they are refracted, the incident energy to the detector is reduced, but what enters the same time window and the same energy window becomes an erroneous LOR as shown by the broken line in FIG. Since such scattered radiation noise has a finite energy resolution, it is inevitable that it is mixed in data obtained during emission measurement in principle. For this reason, a scattered radiation correction is generally used in which the scattered radiation is appropriately estimated and subtracted from the measurement data.

As described above, the scattered rays interact with substances inside and outside the subject, so that the energy is lower than the original 511 keV. Therefore, it is possible to suppress mixing of scattered radiation by narrowing the setting of the energy window of the detector for a portion lower than 511 keV. For this reason, a radiation detector suitable for a PET apparatus is required to have high levels of sensitivity, time resolution, and energy resolution. Radiation detectors that have become mainstream in recent years include Lu ₂ SiO ₅ : Ce (LSO), LuYSiO ₅ : Ce (LYSO), etc., which satisfy the above-mentioned performance at a high level but are expensive. There's a problem. On the other hand, many of the above three main performances satisfy one or two. For example, Bi ₄ Ge ₃ O ₁₂ (BGO) has high sensitivity but low time resolution and low energy resolution, and sodium iodide ( NaI) has high energy resolution but low sensitivity. Semiconductor detectors such as cadmium telluride (CdTe) also have high energy resolution but low sensitivity and time resolution. As described above, if it becomes possible to combine long and short radiation detectors, not only will it be possible to realize a PET apparatus that also uses an inexpensive detector, but the use of detectors that have been excluded from application candidates in the past. Is possible, but the method is not known.

On the other hand, a test using a SPECT apparatus (SPECT test) is a radiopharmaceutical (hereinafter referred to as “specific radiopharmaceutical”) that is labeled with a gamma-ray emitting nuclide ( ^99m Tc, ¹³³ I, ¹¹¹ In, ²⁰¹ Tl, etc.), , SPECT drug) is administered to the subject and the distribution of the SPECT drug in the body of the subject is examined. The SPECT drug administered to the subject emits gamma rays with energy determined for each nuclide. The gamma rays are detected by a radiation detector after restricting the direction of flight through a collimator formed of a radiation shielding material such as lead or tungsten. The position and orientation of the collimator and radiation detector are changed, for example, rotated, and the subject is imaged from a plurality of projection directions. Using an image reconstruction algorithm such as the OS-EM method, the frequency distribution of gamma rays, that is, the SPECT drug A tomographic image of the distribution in the body of the subject can be obtained.

  In SPECT imaging as well as PET imaging, there are cases where gamma rays interact with substances such as the subject such as Compton scattering, and the scattered rays measured by refraction of the traveling direction are erroneously detected. For this reason, in the case of SPECT imaging, similarly to PET imaging, scattered radiation correction in which scattered radiation is appropriately estimated and subtracted from measurement data is generally used.

  A radiation detector suitable for the SPECT apparatus is required to have high sensitivity and high energy resolution. In recent years, radiation detectors that have become mainstream in SPECT apparatuses include sodium iodide (NaI), cadmium telluride (CdTe), cadmium zinc telluride (CZT), and the like. In particular, a semiconductor detector with high energy resolution has a problem that the price is high. Similar to the PET apparatus, if it becomes possible to combine long and short radiation detectors, not only can a SPECT apparatus using an inexpensive detector be realized, but also a detector that has been excluded from the application candidates in the past. Can be used, but the method is not known.

  The present invention uses two types of detectors, a first type radiation detector and a second type radiation detector having higher energy measurement accuracy than the first type radiation detector, as a gamma ray detector. The second type radiation detector is arranged so as to measure substantially the same range, and the first type measured by the first type radiation detector based on the second type measurement data measured by the second type radiation detector. Scattered radiation that mixes in because the energy resolution is not sufficient by using a nuclear medicine diagnostic device with a scattered radiation correction arithmetic unit that removes scattered gamma rays from subjects, beds, detectors, etc. contained in species measurement data Among the plural types of detectors used, noise is treated independently without performing synchronous measurement between the signals of the first type radiation detector and the second type radiation detector, and the detector has a relatively high energy resolution. Correction can be made based on the information measured in step (b). In addition, since the energy resolution is not sufficient, a radiation detector that has not been conventionally used in a nuclear medicine diagnostic apparatus can be used in the nuclear medicine diagnostic apparatus.

<< First Embodiment >>
Hereinafter, a Bi ₄ Ge ₃ O ₁₂ (BGO) scintillator according to an embodiment of the present invention is used as a first type radiation detector 11 and a cadmium tellurium (CdTe) semiconductor detection having higher energy resolution than the first type radiation detector 11. A positron emission tomography apparatus 8 (PET apparatus) using the detector for the second type radiation detector 12 will be described with reference to the drawings as appropriate. In this embodiment, BGO is used for the first type radiation detector 11, but Lu ₂ SiO ₅ : Ce (LSO), LuYSiO ₅ : Ce (LYSO), Gd ₂ SiO ₅ : Ce (GSO) 2, PbWO ₄ ( PWO), LaBr ₃ : Ce (LaBr), or the like can also be used. The second type radiation detector 12 may be made of gallium arsenide (GaAs), lead iodide (PbI ₂ ), thallium bromide (TlBr), cadmium tellurium zinc (CZT), or the like. FIG. 2 shows a schematic diagram of the PET system, and FIG. 3 shows the relationship among the components of the PET apparatus 8. The PET system 1 includes a PET camera 8, a data processing device 3, a display device 22, and a bed 20. The PET camera 8 includes a plurality of detector units 10 arranged on the circumference of the gantry 6, and includes a signal discriminating device 30, a first coincidence device 31, a second coincidence device 32, a scattered ray correction arithmetic device 40, and the like. The imaging object 7 arranged at an appropriate position in the visual field R of the PET apparatus 8 is measured by the bed 20 or the like. Further, as shown in FIG. 4, the detector unit 10 is provided with a first type radiation detector 11 and a second type radiation detector 12 in proximity to each other, and in addition, a photodetector 13, a signal amplifying device 14, and a packet information generating device. Consist of 15. The total volume of the first type radiation detector 11 and the second type radiation detector 12 gives priority to the first radiation detector which is the main measurement, so the first radiation detector is larger, and in the example of FIG. 1. However, since the ratio of the total volume is determined in consideration of the sensitivity ratio of the material used for the first type radiation detector and the second type radiation detector to the measurement target gamma ray, an arbitrary ratio is used regardless of the ratio of the present embodiment. It is possible to use. Since the subject of the present invention is the detector configuration such as the first type radiation detector 11 and the second type radiation detector 12 and the function of the scattered radiation correction arithmetic unit 40 based on the measurement data, the detector and packet information generation Details of electronic circuits such as the connection circuit between the devices 15 and the power supply, as well as explanations and illustrations of light guides generally used when scintillators are used as detectors, electrodes for taking out electrical signals from semiconductor detectors, etc. Omitted.

  The gamma rays emitted from the imaging object 7 are light as an optical signal in the scintillator type detector based on the photoelectric effect or the Compton effect in the first type radiation detector 11 or the second type radiation detector 12 in the detector unit 10. In the detector 13, it is measured as an electric signal in the semiconductor type detector. The measured signal is amplified by the signal amplifying device 14, and the packet information generating device 15 becomes packet information (List-Data) to which the detector channel number, measurement energy, measurement time, detector type, and the like are added. It is sent to the processing device 30. In the data processing device 30, first, the signal discriminating device 30 determines whether the signal is detected by the first type radiation detector 11 or the second type radiation detector 12, and the signal is sent to the corresponding coincidence device. That is, the detection information of the first type radiation detector 11 is sent to the first coincidence counting device 31, and the detection information of the second type radiation detector 12 is sent to the second coincidence counting device 32, and is designated in advance for each detector type. The coincidence counting information is generated by combining the packet information within the time window. By counting this coincidence information for each Line of Response (LOR) determined by the combination of the detector number, that is, the detector position, the measurement data is recorded in the recording device 35 in a general measurement data recording format in a PET device called a sinogram. To be recorded. By the way, since the energy resolution is different between the first type radiation detector 11 and the second type radiation detector 12, the optimum energy window in each detector is different, and in the case of the present embodiment, the first radiation detector 11 is different. Is 350-650 keV, and the second radiation detector 12 is 490-540 keV. Therefore, the amount of scattered gamma rays that have been reduced in energy due to scattering in the subject 7 is smaller in the second type measurement data measured by the second radiation detector 12.

In the scattered radiation correction arithmetic unit 40, the scattered radiation correction is performed using the fact that the second type measurement data has a lower amount of scattered radiation mixed than the first type measurement data. Since the second type radiation detector 12 has higher energy resolution than the first type radiation detector 11 and has a narrower energy window, the second type measurement data is mixed with scattered radiation than the first type measurement data. The amount is low. Therefore, it is possible to correct the scattered radiation of the second type measurement data with high accuracy by a generally known scattered radiation correction method such as Single Scatter Simulation (SSS). On the other hand, the scattered radiation corrected second type measurement data obtained in this way has a smaller volume of the second type radiation detector 12 than the first type radiation detector 11 and a narrow energy window. Scattered rays are removed due to low sensitivity, but statistical errors are large. By subtracting the scattering-corrected second type measurement data from the first type measurement data, the scattered radiation distribution included in the first type measurement data can be estimated (Formula 1). Α in Equation 1 is a coefficient for adjusting the sensitivity ratio between the first type measurement data and the second type measurement data. Incidentally, as shown in FIG. 4, since the first type radiation detector 11 and the second type radiation detector 12 are different in position, there are differences in the LOR positional deviation and the number of LORs indicated by the respective measurement data. However, since the scattered radiation distribution is generally spatially smooth, the position can be adjusted by linear approximation or the like. Next, statistical noise is removed from the scattered radiation distribution obtained as in Equation 1 using a low-pass filter or the like. Finally, the estimated scattered radiation distribution after removal of statistical noise is subtracted from the first type measurement data, thereby correcting the scattered radiation contained in the first type measurement data (Formula 2). In Equation 2, f is a kernel of the low-pass filter. The symbol * in Equation 2 means convolution. Although the second type measurement data has few scattered rays and can be used as it is, the spatial resolution of the image is lowered when a low pass filter is applied thereto. In the method of this embodiment, the scattered radiation distribution that is considered to be smooth is estimated first, and the estimated scattered radiation distribution is subjected to a low-pass filter to prevent a reduction in image resolution while suppressing statistical errors. It is out. This method can be performed in the same manner as the method of setting two energy windows with a single detector and correcting scattered radiation from two types of measurement data. However, in the present invention, the detector corresponding to each energy window is used. Because the materials are different, a detector with high energy resolution can be used for measurement data that requires less scattered radiation, and a detector with high sensitivity can be used for measurement data that requires at most high sensitivity for scattered radiation. It is expected to improve the accuracy of measurement images.
The image corrected as described above is recorded in the recording device 35 and displayed on the display device 22 via the information output device 21. In addition, arrangement | positioning with the 1st radiation detector 11 and the 2nd radiation detector 12 in the detector unit 10 is not restricted to FIG. 4, It is also possible to arrange | position like FIG.5 and FIG.6. In particular, when arranged as shown in FIG. 5, the first radiation detector 11 serves as a collimator for the second radiation detector 12, so that the amount of scattered radiation mixed in the second type measurement data is further reduced. In this embodiment, a scintillator detector is used for the first radiation detector 11 and a semiconductor detector is used for the second radiation detector 12. However, the second radiation detector 12 has an energy resolution higher than that of the first radiation detector 11. If the combination is high, a semiconductor or a scintillator may be used for both detectors.
(Equation 1)
sScatter = sSEW-α × sPEW
(Equation 2)
sTrue = sSEW-sScatter ＊ f

  In Patent Document 1 (WO2009 / 157526), the high time resolution of the organic scintillator is used (Time of Flight imaging). However, since the time resolution of the semiconductor is low in this embodiment, the TOF is basically changed. Not performed. Therefore, there is no common point in patent document 1 and a present Example other than using two types of detectors. Organic scintillators are generally low in sensitivity to high-energy gamma rays such as PET gamma rays (511 keV), and the signals of organic and inorganic scintillators are added together to form a single piece of detection data. In addition, due to this low sensitivity, it is difficult to apply the processing used in the present invention to the configuration of the known example with a single organic scintillator.

  As in the embodiment described above, two types of detectors, the first type radiation detector and the second type radiation detector having higher energy measurement accuracy than the first type radiation detector, are used as the gamma ray detector. The radiation detector and the second type radiation detector are arranged so as to measure substantially the same range, and independent measurement is not performed between the signals of the first type radiation detector and the second type radiation detector. Scattered ray correction arithmetic device that removes scattered gamma rays contained in the first type measurement data measured by the first type radiation detector based on the second type measurement data measured by the second type radiation detector The first-type radiation detector and the second-type radiation detector among the plurality of types of detectors used to detect scattered radiation noise that is mixed due to insufficient energy resolution by the nuclear medicine diagnostic apparatus comprising: Synchronizer between signals Independently handled without becomes possible correction on the basis of the information measured by relatively high energy resolution detectors. In the present embodiment, a radiation detector that has not been used in a nuclear medicine diagnostic apparatus because of insufficient energy resolution can be used in the nuclear medicine diagnostic apparatus.

  Note that the substantially same range when the first type radiation detector and the second type radiation detector are arranged so as to measure substantially the same range is caused by scattered radiation noise that is mixed because the energy resolution is not sufficient. Of the multiple types of detectors used, the signals of the first type radiation detector and the second type radiation detector are not subjected to synchronous measurement, but are handled independently and measured with a detector having a relatively high energy resolution. It suffices if it is within a range necessary for correction based on the information that has been corrected, and it is only necessary that the measurement target matches the correction target portion even if the measurement target does not match the non-correction target portion.

<< Second Embodiment >>
A second embodiment of the present invention will be described. The system components of the second embodiment are the same as those of the first embodiment (FIG. 3), but only the scattered radiation correction calculation method in the scattered radiation correction calculation unit 40 is different as follows.

  In the scattered radiation correction arithmetic unit 40, the scattered radiation correction is performed using the fact that the second type measurement data has a lower amount of scattered radiation mixed than the first type measurement data. Since the second type radiation detector 12 has higher energy resolution than the first type radiation detector 11 and has a narrower energy window, the second type measurement data is mixed with scattered radiation than the first type measurement data. The amount is low. By the way, in a scattered radiation correction method based on a generally known simulation such as Single Scatter Simulation (SSS), an amount proportional to the intensity from each point of the source intensity distribution in the initial image using the measurement image as an initial image. The gamma ray is calculated and the scattered radiation distribution is estimated by simulating the scattering probability and the scattering angle in the subject. However, since the initial image contains scattered radiation, it differs from the actual radiation intensity distribution. Therefore, conventionally, the simulation is repeatedly performed a plurality of times to estimate the actual radiation intensity distribution. In the present plan, second type measurement data with less scattered radiation is used for this initial image. By doing so, it is possible to perform highly accurate scattered ray correction even if the number of simulation iterations is reduced as compared with the prior art. In general, the accuracy and calculation time of a simulation calculation are determined by the range of physical phenomena considered in the calculation process, the resolution of an image used for the calculation, and the like. In order to improve the accuracy, a simulation incorporating a finer physical phenomenon with a high resolution is required, and the calculation time becomes longer. In this embodiment, as described above, since the number of simulation iterations can be reduced, even if the calculation time is the same, it is possible to improve the accuracy of a single simulation.

«Third embodiment»
A third embodiment of the present invention will be described. The positron emission tomography camera 8 (PET camera) of the third embodiment is substantially the same as that of the first embodiment, but handles the second type measurement data measured by the second type radiation detector 12 and scatter correction. The scattering correction method in the arithmetic unit 40 is different.

In this embodiment, a Bi ₄ Ge ₃ O ₁₂ (BGO) scintillator is used for the first type radiation detector 11, and a cadmium tellurium (CdTe) semiconductor having a higher energy resolution than the first type radiation detector 11 for the second type radiation detector 12. A detector was used. The first type radiation detector 11 includes Lu ₂ SiO ₅ : Ce (LSO), LuYSiO ₅ : Ce (LYSO), Gd ₂ SiO ₅ : Ce (GSO), PbWO ₄ (PWO), LaBr ₃ : Ce (LaBr). It is also possible to use gallium arsenide (GaAs), lead iodide (PbI ₂ ), thallium bromide (TlBr), cadmium tellurium zinc (CZT), etc. Is also possible. FIG. 2 shows a schematic diagram of the PET system, and FIG. 7 shows the relationship among the components of the PET apparatus 8. The PET system 1 includes a PET camera 8, a data processing device 3, a display device 22, and a bed 20. In the PET camera 8, a plurality of detector units 10 are arranged on the circumference of the gantry 6, and a signal discriminating device 30, a first coincidence device 31, an energy spectrum generating device 33, a scattered fraction estimating device 34, and a scattered ray correction calculating device. 40, and the imaging target 7 arranged at an appropriate position in the visual field R of the PET apparatus 8 is measured by the bed 20 or the like. Further, as shown in FIG. 4, the detector unit 10 is provided with a first type radiation detector 11 and a second type radiation detector 12 in proximity to each other, and in addition, a photodetector 13, a signal amplifying device 14, and a packet information generating device. Consist of 15. The total volume of the first type radiation detector 11 and the second type radiation detector 12 gives priority to the first radiation detector which is the main measurement, so the first radiation detector is larger, and in the example of FIG. 1. Since the ratio of the total volume is determined in consideration of the sensitivity ratio of the material used for the first type radiation detector and the second type radiation detector to the measurement target gamma ray, an arbitrary ratio should be used regardless of the ratio of this embodiment. Is possible. Since the subject of the present invention is the detector configuration such as the first type radiation detector 11 and the second type radiation detector 12 and the function of the scattered radiation correction arithmetic unit 40 based on the measurement data, the detector and packet information generation Details of electronic circuits such as the connection circuit between the devices 15 and the power supply, as well as explanations and illustrations of light guides generally used when scintillators are used as detectors, electrodes for taking out electrical signals from semiconductor detectors, etc. Omitted.

  The gamma rays emitted from the imaging object 7 are light as an optical signal in the scintillator type detector based on the photoelectric effect or the Compton effect in the first type radiation detector 11 or the second type radiation detector 12 in the detector unit 10. In the detector 13, it is measured as an electric signal in the semiconductor type detector. The measured signal is amplified by the signal amplifying device 14, and the packet information generating device 15 becomes packet information (List-Data) to which the detector channel number, measurement energy, measurement time, detector type, and the like are added. It is sent to the processing device 30. In the data processing device 30, first, the signal discrimination device 30 determines whether the signal is detected by the first type radiation detector 11 or the second type radiation detector 12, and the data of the first type radiation detector 11 can be obtained. For example, the data of the second type radiation detector 12 is sent to the first coincidence device 31 to the energy spectrum generator 33. The first coincidence device 31 generates coincidence information by combining packet information within a time window designated in advance. By counting this coincidence information for each Line of Response (LOR) determined by the combination of the detector number, that is, the detector position, the measurement data is recorded in the recording device 35 in a general measurement data recording format in a PET device called a sinogram. To be recorded. In addition, the energy spectrum generation device 33 generates an energy spectrum of gamma rays detected by the respective second type radiation detectors 12 and records it in the recording device 35.

  By the way, the mixing ratio of scattered radiation in the measured signal, that is, the scattered radiation fraction, changes depending on the drug distribution in the subject and the substance density distribution (physique etc.) of the subject. It is necessary to know the scattered dose to be corrected for each imaging. The energy spectrum actually measured by the second radiation detector 12 with respect to a point source having no scatterer is used as a direct gamma-ray response characteristic, and data deviating from this response characteristic is considered to be scattered radiation. FIG. 9 shows an energy spectrum of a phantom having a dotted line source and a scatterer. When the two spectra in FIG. 9 are compared, they are almost the same in the vicinity of the total absorption peak, but there is a difference in the distribution in the other ranges. Therefore, this difference is used as the amount of mixed scattered radiation, and the scattering fraction (SF) is obtained by dividing the amount of mixed scattered radiation by the total count.

In the scattering fraction estimation apparatus 34, a scattering fraction is estimated from the energy spectrum produced | generated in the energy spectrum production | generation apparatus 33 with the above-mentioned method. In addition, estimation of this scattering fraction may be calculated | required about each 2nd radiation detector 12 about a single detector, or may be calculated | required after combining the several detector which adjoins. In order to obtain the scattering fraction for each LOR by the simultaneous counting in the first type radiation detector 11 based on the scattering fraction in the second type radiation detector 12 obtained as described above, the following processing is performed. First, one or a plurality of average scattering fractions of the second type radiation detector 12 in the vicinity of an arbitrary first type radiation detector 11 are calculated and set as the scattering fraction of the first type radiation detector 11. Thereafter, the respective scattering fractions of the first type radiation detector pair (detector A and detector B) forming the LOR, that is, the scattering fraction SF _A at the detector _A and the scattering fraction SF _B at the detector _B Based on the above, the SF _LOR of the LOR is obtained using Equation 3.
(Equation 3)
SF _LOR = 1- (1- SF _A ) (1- SF _B ) = SF _A + SF _B -SF _A x SF _B

In the scattered radiation correction calculation device 40, the SF _LOR is multiplied by the count number measured by each LOR of the first type measurement data, thereby _obtaining and subtracting the scattered radiation count number to be removed. Note that the processing of the scattered radiation correction calculation device 40 is not based on the above, and when subtracting the estimated scattered radiation distribution from the first type measurement data in simulation scatter correction such as Single Scatter Simulation (SSS), Instead of tail-fitting that adjusts the intensity of the scattered radiation distribution so that the profile of the periphery of the imaging target matches the profile of the estimated scattered radiation distribution, the scattered radiation count to be removed and the estimated scattering estimated from the SF _LOR It is also possible to adjust the line distribution count to match.

  The image corrected as described above is recorded in the recording device 35 and displayed on the display device 22 via the information output device 21. In addition, arrangement | positioning with the 1st radiation detector 11 and the 2nd radiation detector 12 in the detector unit 10 is not restricted to FIG. 4, It is also possible to arrange | position like FIG.6 and FIG.8. In this embodiment, a scintillator detector is used for the first radiation detector 11 and a semiconductor detector is used for the second radiation detector 12. However, the second radiation detector 12 has an energy resolution higher than that of the first radiation detector 11. As long as the combination is high and the energy spectrum stability of the second radiation detector 12 is high, a semiconductor or a scintillator may be used for both detectors.

<< Fourth Embodiment >>
A fourth embodiment of the present invention will be described. Although the system components of the fourth embodiment are substantially the same as those of the third embodiment (FIG. 10), a maximum likelihood estimated image reconstruction device 41 is used instead of the scattered radiation correction calculation unit 40.

Similar to the third embodiment, the scattering fraction estimation device 34 estimates the scattering fraction from the energy spectrum generated by the energy spectrum generation device 33. The process up to the estimation of the scattering fraction is the same as that of the third embodiment. In order to obtain the scattering fraction for each LOR by the simultaneous counting in the first type radiation detector 11 based on the obtained scattering fraction in the second type radiation detector 12, the following processing is performed. First, one or a plurality of average scattering fractions of the second type radiation detector 12 in the vicinity of an arbitrary first type radiation detector 11 are calculated and set as the scattering fraction of the first type radiation detector 11. Thereafter, the respective scattering fractions of the first type radiation detector pair (detector A and detector B) forming the LOR, that is, the scattering fraction SF _A at the detector _A and the scattering fraction SF _B at the detector _B Based on the above, the SF _LOR of the LOR is obtained using Equation 3. The maximum likelihood estimated image reconstruction apparatus 41 performs weighting by assuming that (1-SF _LOR ) is a true coincidence count among the LOR counts of the first type measurement data, and performs maximum likelihood estimated image reconstruction. In contrast to this method, in this embodiment, a plurality of types of radiation detectors are used, and the energy spectrum measured by the second type radiation detector 12 having excellent energy resolution is used. As described above, as a radiation detector suitable for the PET apparatus, it is assumed that a detector that requires high levels of sensitivity, time resolution, and energy resolution is used. A combination of three main performances satisfying one or two can be implemented, and not only a PET apparatus using an inexpensive detector can be realized, but also has been excluded from application candidates in the past. The detector can be used.

<< Fifth Embodiment >>
Hereinafter, a sodium iodide (NaI) scintillator according to an embodiment of the present invention is used as a first type radiation detector 11 and a cadmium tellurium (CdTe) semiconductor detector having a higher energy resolution than the first type radiation detector 11 is used. A single photon emission tomography apparatus 9 (SPECT apparatus) used for the two-type radiation detector 12 will be described with reference to the drawings as appropriate. In this embodiment, NaI is used for the first type radiation detector 11, but other scintillators as described in the first embodiment can also be used. The second type radiation detector 12 may be made of gallium arsenide (GaAs), lead iodide (PbI ₂ ), thallium bromide (TlBr), cadmium tellurium zinc (CZT), or the like. FIG. 2 shows a system outline diagram, and FIG. 11 shows the relationship among the components of the SPECT apparatus 9. The SPECT system 1 includes a nuclear medicine camera 2 (SPECT camera 9 in this example), a data processing device 3, a display device 22, and a bed 20. The SPECT camera 9 includes a single or a plurality of detector heads 16, and the detector head 16 rotates 180 degrees or 360 degrees around the imaging field of view R by a rotation mechanism (not shown). Note that this rotation angle is arbitrary because it depends on the resolution requirement in the angular direction. The data processing device 3 includes a signal discriminating device 30, a first counting device 38, a second counting device 39, a scattered radiation correction calculating device 40, and the like, and is arranged at an appropriate position in the visual field R of the SPECT device 9 by the bed 20 or the like. The imaged object 7 is measured. Further, as shown in FIG. 4, the detector head 16 is provided with the first type radiation detector 11 and the second type radiation detector 12 in proximity to each other, and in addition, the photodetector 13, the signal amplifying device 14, and the packet information generating device. 15 and a collimator 17. The total volume of the first type radiation detector 11 and the second type radiation detector 12 gives priority to the first radiation detector which is the main measurement, so the first radiation detector is larger, and in the example of FIG. 1. However, since the ratio of the total volume is determined in consideration of the sensitivity ratio of the material used for the first type radiation detector and the second type radiation detector to the measurement target gamma ray, an arbitrary ratio is used regardless of the ratio of the present embodiment. The total volume ratio of the second radiation detector 12 may be larger than that of the first radiation detector 11. Since the subject of the present invention is the detector configuration such as the first type radiation detector 11 and the second type radiation detector 12 and the function of the scattered radiation correction arithmetic unit 40 based on the measurement data, the detector and packet information generation Details of electronic circuits such as the connection circuit between the devices 15 and the power supply, as well as explanations and illustrations of light guides generally used when scintillators are used as detectors, electrodes for taking out electrical signals from semiconductor detectors, etc. Omitted.

The gamma rays emitted from the imaging object 7 are light as an optical signal in the scintillator type detector based on the photoelectric effect or the Compton effect in the first type radiation detector 11 or the second type radiation detector 12 in the detector head 16. In the detector 13, it is measured as an electric signal in the semiconductor type detector. The measured signal is amplified by the signal amplifying device 14, and the packet information generating device 15 adds the detector channel number, measurement energy, measurement time, detector type, detector head rotation angle, rotation radius, and the like. Packet information (List-Data) is sent to the data processing device 30. In the data processing device 30, first, the signal discriminating device 30 determines whether the signal is detected by the first type radiation detector 11 or the second type radiation detector 12, and the signal is sent to the corresponding counting device. That is, the detection information of the first type radiation detector 11 is sent to the first counting device 38, and the detection information of the second type radiation detector 12 is sent to the second counting device 39. In the first counting device 38 and the second counting device 39, the detector number, that is, the detector position and the rotation angle and rotation radius of the detector head are recorded as projection data in the recording device 35. This projection data is edited for each line of response (LOR) determined by the combination of the rotation angle and rotation radius of the detector head, and converted into a general measurement data recording format by a PET or SPECT apparatus called a sinogram and recorded. It doesn't matter. By the way, since the energy resolution is different between the first-type radiation detector 11 and the second-type radiation detector 12, the optimum energy window in each detector is different. For example, when ^99m Tc is imaged in the present embodiment, the first radiation detector 11 has 120-170 keV, and the second radiation detector 12 has 130-160 keV. Therefore, the amount of scattered gamma rays that have been reduced in energy due to scattering in the subject 7 is smaller in the second type measurement data measured by the second radiation detector 12.

In the scattered radiation correction arithmetic unit 40, the scattered radiation correction is performed using the fact that the second type measurement data has a lower amount of scattered radiation mixed than the first type measurement data. Since the second type radiation detector 12 has higher energy resolution than the first type radiation detector 11 and has a narrower energy window, the second type measurement data is mixed with scattered radiation than the first type measurement data. The amount is low. By subtracting the second type measurement data from the first type measurement data, the scattered radiation distribution included in the first type measurement data can be estimated (Formula 1). Α in Equation 1 is a coefficient for adjusting the sensitivity ratio between the first type measurement data and the second type measurement data. As shown in FIG. 12, the first type radiation detector 11 and the second type radiation detector 12 have different positions, and therefore there are differences in the LOR positional deviation and the number of LORs indicated by the respective measurement data. However, since the scattered radiation distribution is generally spatially smooth and the detector head rotates, it is possible to adjust the position by linear approximation or the like. Next, statistical noise is removed from the scattered radiation distribution obtained as in Equation 1 using a low-pass filter or the like. Finally, the estimated scattered radiation distribution after removal of statistical noise is subtracted from the first type measurement data, thereby correcting the scattered radiation contained in the first type measurement data (Formula 2). In Equation 2, f is a kernel of the low-pass filter. Although the second type measurement data has few scattered rays and can be used as it is, the spatial resolution of the image is lowered when a low pass filter is applied thereto. In the method of this embodiment, the scattered radiation distribution that is considered to be smooth is estimated first, and the estimated scattered radiation distribution is subjected to a low-pass filter to prevent a reduction in image resolution while suppressing statistical errors. It is out. For example, this method can be implemented in the same manner as the method in which two energy windows are set with a single detector in PET and the scattered radiation correction is performed from two types of measurement data. In the present invention, each energy window is supported. Because the materials of the detectors to be used are different, a detector with high energy resolution is required for measurement data that requires a small amount of scattered radiation, and a detector with high sensitivity is required for measurement data that requires at most high sensitivity for scattered radiation. It can be used, and the accuracy of measurement images is expected to improve.
The image corrected as described above is recorded in the recording device 35 and displayed on the display device 22 via the information output device 21. The arrangement of the first radiation detector 11 and the second radiation detector 12 in the detector head 16 is not limited to FIG. 12, and when there are a plurality of detector heads 16, the first radiation detector 11 as shown in FIGS. 13 and 14. Combining detector heads consisting of only the seed radiation detector and the second kind radiation detector in an arbitrary ratio, that is, for example, when the detector head is two SPECT cameras, one is as shown in FIG. 13 and the other is as shown in FIG. It is also possible to do. In this embodiment, a scintillator detector is used for the first radiation detector 11 and a semiconductor detector is used for the second radiation detector 12. However, the second radiation detector 12 has an energy resolution higher than that of the first radiation detector 11. If the combination is high, a semiconductor or a scintillator may be used for both detectors. The collimator is not limited to the parallel collimator as illustrated in the present embodiment, and may be a collimator shape such as a fan beam, a pinhole, or a multi-pinhole. Furthermore, although the rotation imaging is used in the present embodiment, an imaging method such as moving the detector so as to be shaken or moving in parallel may be used.

<< Sixth Embodiment >>
A sixth embodiment of the present invention will be described. The system components of the sixth embodiment are the same as those of the fifth embodiment (FIG. 11), but only the scattered radiation correction calculation method in the scattered radiation correction calculation unit 40 is different and is as follows.

  In the scattered radiation correction arithmetic unit 40, the scattered radiation correction is performed using the fact that the second type measurement data has a lower amount of scattered radiation mixed than the first type measurement data. Since the second type radiation detector 12 has higher energy resolution than the first type radiation detector 11 and has a narrower energy window, the second type measurement data is mixed with scattered radiation than the first type measurement data. The amount is low. By the way, in a scattered radiation correction method based on a generally known simulation such as Single Scatter Simulation (SSS), an amount proportional to the intensity from each point of the source intensity distribution in the initial image using the measurement image as an initial image. The gamma ray is calculated and the scattered radiation distribution is estimated by simulating the scattering probability and the scattering angle in the subject. However, since the initial image contains scattered radiation, it differs from the actual radiation intensity distribution. Therefore, conventionally, the simulation is repeatedly performed a plurality of times to estimate the actual radiation intensity distribution. In the present plan, second type measurement data with less scattered radiation is used for this initial image. By doing so, it is possible to perform highly accurate scattered ray correction even if the number of simulation iterations is reduced as compared with the prior art. In general, the accuracy and calculation time of a simulation calculation are determined by the range of physical phenomena considered in the calculation process, the resolution of an image used for the calculation, and the like. In order to improve the accuracy, a simulation incorporating a finer physical phenomenon with a high resolution is required, and the calculation time becomes longer. In this embodiment, as described above, since the number of simulation iterations can be reduced, even if the calculation time is the same, it is possible to improve the accuracy of a single simulation.

<< Seventh embodiment >>
A seventh embodiment of the present invention will be described. The single photon emission tomography camera 9 (SPECT camera) of the seventh embodiment is substantially the same as that of the fifth embodiment, but handles the second type measurement data measured by the second type radiation detector 12, and The scattering correction method in the scattering correction arithmetic unit 40 is different, and is as shown in FIG.

A sodium iodide (NaI) scintillator was used as the first type radiation detector 11, and a cadmium tellurium (CdTe) semiconductor detector was used as the second type radiation detector 12 having higher energy resolution than the first type radiation detector 11. In this embodiment, NaI is used for the first type radiation detector 11, but other scintillators as described in the first embodiment can also be used. The second type radiation detector 12 may be made of gallium arsenide (GaAs), lead iodide (PbI ₂ ), thallium bromide (TlBr), cadmium tellurium zinc (CZT), or the like. FIG. 2 shows a system outline diagram, and FIG. 15 shows the relationship among the components of the SPECT apparatus 9. The SPECT system 1 includes a nuclear medicine camera 2 (SPECT camera 9 in this example), a data processing device 3, a display device 22, and a bed 20. The SPECT camera 9 includes a single or a plurality of detector heads 16, and the detector head 16 rotates 180 degrees or 360 degrees around the imaging field of view R by a rotation mechanism (not shown). Note that this rotation angle is arbitrary because it depends on the resolution requirement in the angular direction. The data processing device 3 includes a signal discriminating device 30, a first counting device 38, a second counting device 39, a scattered radiation correction calculating device 40, and the like, and is arranged at an appropriate position in the visual field R of the SPECT device 9 by the bed 20 or the like. The imaged object 7 is measured. Further, as shown in FIG. 4, the detector head 16 is provided with the first type radiation detector 11 and the second type radiation detector 12 in proximity to each other, and in addition, the photodetector 13, the signal amplifying device 14, and the packet information generating device. 15 and a collimator 17. The total volume of the first type radiation detector 11 and the second type radiation detector 12 gives priority to the first radiation detector which is the main measurement, so the first radiation detector is larger, and in the example of FIG. 1. However, since the ratio of the total volume is determined in consideration of the sensitivity ratio of the material used for the first type radiation detector and the second type radiation detector to the measurement target gamma ray, an arbitrary ratio is used regardless of the ratio of the present embodiment. The total volume ratio of the second radiation detector 12 may be larger than that of the first radiation detector 11. Since the subject of the present invention is the detector configuration such as the first type radiation detector 11 and the second type radiation detector 12 and the function of the scattered radiation correction arithmetic unit 40 based on the measurement data, the detector and packet information generation Details of electronic circuits such as the connection circuit between the devices 15 and the power supply, as well as explanations and illustrations of light guides generally used when scintillators are used as detectors, electrodes for taking out electrical signals from semiconductor detectors, etc. Omitted.

  The gamma rays emitted from the imaging object 7 are light as an optical signal in the scintillator type detector based on the photoelectric effect or the Compton effect in the first type radiation detector 11 or the second type radiation detector 12 in the detector head 16. In the detector 13, it is measured as an electric signal in the semiconductor type detector. The measured signal is amplified by the signal amplifying device 14, and the packet information generating device 15 adds the detector channel number, measurement energy, measurement time, detector type, detector head rotation angle, rotation radius, and the like. Packet information (List-Data) is sent to the data processing device 30. In the data processing device 30, first, the signal discrimination device 30 determines whether the signal is detected by the first type radiation detector 11 or the second type radiation detector 12, and the data of the first type radiation detector 11 can be obtained. For example, the data of the second type radiation detector 12 is sent to the first counting device 38 to the energy spectrum generating device 33. In the first counting device 38, the detector number, that is, the detector position and the rotation angle and rotation radius of the detector head are recorded in the recording device 35 as projection data. This projection data is edited for each line of response (LOR) determined by the combination of the rotation angle and rotation radius of the detector head, and converted into a general measurement data recording format by a PET or SPECT apparatus called a sinogram and recorded. It doesn't matter. In addition, the energy spectrum generation device 33 generates an energy spectrum of gamma rays detected by the respective second type radiation detectors 12 and records it in the recording device 35.

The scattered radiation correction is performed by the scattered radiation correction arithmetic unit 40 based on the above measurement data, and the method is the same as that of the third embodiment. That is, first, the energy spectrum previously measured by the second radiation detector 12 with respect to the point source without the scatterer by the scattering fraction estimation device 34 and the energy spectrum generation device 33 by the above-described method when the subject 7 is imaged. The scattering fraction is obtained by comparing with the generated energy spectrum. In addition, estimation of this scattering fraction may be calculated | required about each 2nd radiation detector 12 about a single detector, or may be calculated | required after combining the several detector which adjoins. Thereafter, one or a plurality of average scattering fractions of the second type radiation detector 12 proximate to the arbitrary first type radiation detector 11 are calculated, and are calculated as the estimated scattering fraction SF _{EST of} the first type radiation detector 11. And

In the scattered radiation correction calculation device 40, the SF _EST is multiplied by the count number measured in the recording unit (channel or LOR) of the first type measurement data, thereby obtaining and subtracting the scattered radiation count number to be removed. Make corrections. Note that the processing of the scattered radiation correction calculation device 40 is not based on the above, and when subtracting the estimated scattered radiation distribution from the first type measurement data in simulation scatter correction such as Single Scatter Simulation (SSS), Instead of tail-fitting that adjusts the intensity of the scattered radiation distribution so that the profile of the periphery of the imaging target matches the profile of the estimated scattered radiation distribution, the scattered radiation count to be removed and the estimated scattering estimated from the SF _{EST are used.} It is also possible to adjust the line distribution count to match.

  The image corrected as described above is recorded in the recording device 35 and displayed on the display device 22 via the information output device 21. The arrangement of the first radiation detector 11 and the second radiation detector 12 in the detector head 16 is not limited to FIG. 12, and when there are a plurality of detector heads 16, the first radiation detector 11 as shown in FIGS. 13 and 14. Combining detector heads consisting of only the seed radiation detector and the second kind radiation detector in an arbitrary ratio, that is, for example, when the detector head is two SPECT cameras, one is as shown in FIG. 13 and the other is as shown in FIG. It is also possible to do. In this embodiment, a scintillator detector is used for the first radiation detector 11 and a semiconductor detector is used for the second radiation detector 12. However, the second radiation detector 12 has an energy resolution higher than that of the first radiation detector 11. If the combination is high, a semiconductor or a scintillator may be used for both detectors. The collimator is not limited to the parallel collimator as illustrated in the present embodiment, and may be a collimator shape such as a fan beam, a pinhole, or a multi-pinhole. Furthermore, although the rotation imaging is used in the present embodiment, an imaging method such as moving the detector so as to be shaken or moving in parallel may be used.

<< Eighth Embodiment >>
An eighth embodiment of the present invention will be described. The system components of the eighth embodiment are substantially the same as those of the third embodiment (FIG. 16), but a maximum likelihood estimated image reconstruction device 41 is used instead of the scattered radiation correction calculation unit 40. In this example, as in the seventh embodiment, the scattering fraction estimation device 34 estimates the scattering fraction from the energy spectrum generated by the energy spectrum generation device 33. The process up to the estimation of the scattering fraction, that is, the estimated scattering fraction SF _EST of the recording unit (channel or LOR) of the first type measurement data is the same as in the seventh embodiment.

Based on the obtained estimated scattering fraction SF _EST , the maximum likelihood estimated image reconstruction device 41 has a true count (1-SF _EST ) among the counts of the recording units (channel or LOR) of the first type measurement data. Is weighted and maximum likelihood estimation image reconstruction is performed. This method can be realized by applying the PET method to SPECT image reconstruction. In contrast to the maximum likelihood estimation image reconstruction method, in this embodiment, a plurality of types of radiation detectors are used, and the energy spectrum measured by the second type radiation detector 12 having excellent energy resolution is used. As described above, a radiation detector suitable for a SPECT apparatus is required to have high sensitivity and energy resolution, but the present invention satisfies one or two of the two main performances described above. It is possible to implement a combination of components, and not only a SPECT apparatus using an inexpensive detector can be realized, but also a detector that has been conventionally excluded from application candidates can be used.

In addition, this invention is not limited to an above-described Example, Various modifications are included. For example, the above-described embodiments have been described in detail for easy understanding of the present invention, and are not necessarily limited to those having all the configurations described. Further, a part of the configuration of one embodiment can be replaced with the configuration of another embodiment, and the configuration of another embodiment can be added to the configuration of one embodiment. Further, it is possible to add, delete, and replace other configurations for a part of the configuration of each embodiment.
Each of the above-described configurations, functions, processing units, processing means, and the like may be realized by hardware by designing a part or all of them with, for example, an integrated circuit. In addition, each of the above-described configurations, functions, and the like may be realized by software by interpreting and executing a program that realizes each function by a computer processor. Information such as programs, tables, files, measurement information, and calculation information for realizing each function is stored in a recording device such as a memory, a hard disk, or an SSD (Solid State Drive), or a recording medium such as an IC card, an SD card, or a DVD. Can be put in. Therefore, each process and each configuration can realize each function as a processing unit, a processing unit, a program module, and the like.
Further, the control lines and information lines indicate what is considered necessary for the explanation, and not all the control lines and information lines on the product are necessarily shown. Actually, it may be considered that almost all the components are connected to each other.

1 Nuclear medicine imaging system
2 Nuclear medicine camera
3 Data processing device
4 Scattered ray correction calculation device
5 Imaging targets
6 Gantry 7 Imaging object 8 Positron emission tomographic imaging camera (PET camera)
9 Single photon emission tomography camera (SPECT camera)
DESCRIPTION OF SYMBOLS 10 Detector unit 11 1st type radiation detector 12 2nd type radiation detector 13 Optical detector 14 Signal amplifier 14 Packet information production | generation apparatus 16 Detector head 17 Collimator 20 Bed 21 Information output device 22 Display apparatus 23 Gamma ray 24 Radiation Detector 30 Signal discriminating device 31 First coincidence device 32 Second coincidence device 33 Energy spectrum generation device 34 Scattered fraction estimation device 35 Storage device 36 First measurement image creation device 37 Second measurement image creation device 38 First count device 39 Second counting device 40 Scattered ray correction computing device 41 Maximum likelihood estimation image reconstruction device

Claims (12)

  The first type radiation detector and the second type radiation are used as the gamma ray detector, the first type radiation detector and the second type radiation detector having higher energy measurement accuracy than the first type radiation detector. The detectors are arranged to measure substantially the same range, and the signals of the first type radiation detector and the second type radiation detector are handled independently without performing synchronous measurement, and the second type radiation detection is performed. A nuclear medicine diagnostic apparatus comprising: a scattered radiation correction arithmetic device that removes scattered gamma rays included in the first type measurement data measured by the first type radiation detector based on the second type measurement data measured by the instrument .   The nuclear medicine diagnosis apparatus according to claim 1, wherein a scintillator is used for the first type radiation detector and a semiconductor detector is used for the second type radiation detector.   The nuclear medicine diagnosis apparatus according to claim 1, wherein a scintillator is used for the first type radiation detector and the second type radiation detector.   The nuclear medicine diagnostic apparatus according to claim 1, wherein a semiconductor detector is used for the first type radiation detector and the second type radiation detector.   5. The nuclear medicine diagnostic apparatus according to claim 1, or claim 2, or claim 3 or claim 4, wherein the first type radiation detector and the second type radiation detector perform simultaneous counting measurement. Positron emission tomography system.   In the nuclear medicine diagnostic apparatus according to claim 1, or claim 2, or claim 3 or claim 4, the first type radiation detector performs coincidence measurement, and the second type radiation detector A positron emission tomography system that records energy information of detected gamma rays and creates an energy spectrum.   5. The nuclear medicine diagnostic apparatus according to claim 1 or claim 2 or claim 3 or claim 4, wherein the position of the first type radiation detector and the second type radiation detector or the position of the collimator or Single-photon emission tomography device that captures images while changing posture.   5. The nuclear medicine diagnostic apparatus according to claim 1 or claim 2 or claim 3 or claim 4, wherein the position of the first type radiation detector and the second type radiation detector or the position of the collimator or A single photon emission tomography apparatus that performs imaging with the first type radiation detector while changing the posture, records energy information of detected gamma rays in the second type radiation detector, and creates an energy spectrum.   In the scattered radiation correction arithmetic unit of the nuclear medicine diagnostic apparatus according to any one of claims 5 and 7, a sinogram or an image obtained from a measurement result of the first type radiation detector as the first type measurement data. Using the same type as the first type measurement data among the sinogram or the image obtained from the measurement result of the second type radiation detector as the second type measurement data, and using the first type measurement data and the After adjusting the measurement sensitivity with the second type measurement data, the scattered radiation distribution is estimated by subtracting the second type measurement data from the first type measurement data, and the statistical noise of the estimated scattered radiation distribution is filtered. A nuclear medicine diagnostic apparatus that corrects scattered radiation by subtracting the estimated scattered radiation distribution after noise removal from the first-type measurement data after removing by the above.   In the scattered radiation correction arithmetic unit of the nuclear medicine diagnosis apparatus according to any one of claims 5 and 7, when the scattered radiation distribution included in the first type measurement data is estimated by simulation, the second type A nuclear medicine diagnostic apparatus that performs scattered ray correction by subtracting an estimated scattered ray distribution from the first type measurement data using a second type measurement image obtained from measurement data as an initial value.   The scattered radiation correction arithmetic unit of the nuclear medicine diagnostic apparatus according to claim 5 or claim 6 or claim 7 or claim 8, wherein the first-type radiation detector and the second-type radiation detector are The ratio of scattered rays to the gamma ray counts measured by the first type measurement detectors that are arranged close to each other and based on the second type measurement energy spectrum obtained from the second type measurement data. A nuclear medicine diagnostic apparatus characterized in that a certain scattered radiation ratio is estimated, and the scattered radiation ratio is used when estimating a scattered radiation distribution included in the first type measurement data by simulation.   The scattered radiation correction arithmetic unit of the nuclear medicine diagnostic apparatus according to claim 5 or claim 6 or claim 7 or claim 8, wherein the first-type radiation detector and the second-type radiation detector are The ratio of scattered rays to the gamma ray counts measured by the first type measurement detectors that are arranged close to each other and based on the second type measurement energy spectrum obtained from the second type measurement data. An apparatus for nuclear medicine diagnosis characterized by estimating a certain scattered ray ratio and using the scattered ray ratio when reconstructing a tomographic image from the first type measurement data by a maximum likelihood estimation image reconstruction method.