Nuclear medicine diagnostic device

Abstract

To perform normalization while reducing worker exposure or to perform normalization with high accuracy.SOLUTION: A nuclear medicine diagnosis device includes a scintillator, a storage unit, a calculation unit, and a generation unit. The scintillator emits self-radiation. The storage unit stores first detection efficiency correction data generated on the basis of an external radiation source or a simulation and first detection efficiency data for each scintillator calculated on the basis of radiation emitted from the scintillator. The calculation unit calculates second detection efficiency data for each scintillator, which is calculated on the basis of radiation emitted from the scintillator. The generating unit generates second detection efficiency correction data on the basis of the first detection efficiency correction data, the first detection efficiency data, and the second detection efficiency data.SELECTED DRAWING: Figure 1

Description

Technical Field

[0001] The embodiments disclosed in this specification and the drawings relate to nuclear medicine diagnostic devices.

Background Art

[0002] The calibration of a PET device includes a process called normalization, which measures the detection efficiency for each scintillator (crystal) and creates correction data used in clinical image creation. Normalization is generally performed by collecting gamma rays from a calibration radiation source placed inside the gantry.

[0003] However, when performing normalization using a calibration radiation source, operator exposure may occur due to the handling of the calibration radiation source. In addition, accuracy variations may occur due to work variations in the installation position accuracy of the radiation source and non-uniformity of the isotope distribution inside the radiation source.

Prior Art Documents

Patent Documents

[0004]

Patent Document 1

Summary of the Invention

Problems to be Solved by the Invention

[0005] One of the problems to be solved by the embodiments disclosed in this specification and the drawings is to perform normalization while reducing operator exposure or to perform normalization with high accuracy. However, the problems to be solved by the embodiments disclosed in this specification and the drawings are not limited to the above problems. It is also possible to position the problems corresponding to the respective effects of each configuration shown in the embodiments described later as other problems.

Means for Solving the Problems

[0006] The nuclear medicine diagnostic apparatus according to the embodiment includes a scintillator, a storage unit, a calculation unit, and a generation unit. The scintillator emits self-radiation. The storage unit stores first detection efficiency correction data generated based on an external radiation source or a simulation, and first detection efficiency data for each scintillator calculated based on the radiation emitted from the scintillator. The calculation unit calculates second detection efficiency data for each scintillator calculated based on the radiation emitted from the scintillator. The generation unit generates second detection efficiency correction data based on the first detection efficiency correction data, the first detection efficiency data, and the second detection efficiency data.

Brief Description of the Drawings

[0007]

Figure 1

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Embodiments for Carrying Out the Invention

[0008] Hereinafter, embodiments of a nuclear medicine diagnostic apparatus will be described in detail with reference to the drawings.

[0009] FIG. 1 is a diagram showing the configuration of a PET device 100 according to an embodiment. As shown in FIG. 1, the PET device 100 according to the embodiment includes a gantry device 10 and a console device 20.

[0010] The gantry device 10 includes a detector 1, a timing information acquisition circuit 102, a top plate 103, a bed 104, and a bed drive unit 105.

[0011] The detector 1 is a detector that detects radiation by detecting scintillation light (fluorescence), which is light re-emitted when a substance in an excited state due to the interaction of annihilation gamma rays emitted from positrons in the subject P with a phosphor (scintillator) transitions back to the ground state. The detector 1 detects the energy information of the radiation of the annihilation gamma rays emitted from the positrons in the subject P. A plurality of detectors 1 are arranged in a ring shape surrounding the periphery of the subject P and are composed of, for example, a plurality of detector blocks.

[0012] As an example of a specific configuration of the detector 1, it is a photon counting type, Anger type detector and has, for example, a scintillator, a photodetection element, and a light guide. That is, each of the pixels included in the detector 1 has a scintillator and a photodetection element that detects the generated scintillation light.

[0013] The scintillator converts the annihilation gamma rays emitted from the positrons in the subject P and incident thereon into scintillation photons (optical photons) and outputs them. The scintillator is formed of a scintillator crystal suitable for TOF measurement and energy measurement, such as LaBr3 (Lanthanum Bromide), LYSO (Lutetium Yttrium Oxyorthosilicate), LSO (Lutetium Oxyorthosilicate), LGSO (Lutetium Gadolinium Oxyorthosilicate), or BGO, and is arranged two-dimensionally, for example. Note that the scintillator constituting the detector 1 emits self-radiation, for example. As an example, the scintillator constituting the detector 1 contains 176 Lu, a radioactive nuclide, and emits the gamma rays emitted from the 176 Lu as self-radiation.

[0014] As the photodetector element, for example, a SiPM (Silicon photomultiplier) or a photomultiplier tube is used. The photomultiplier tube has a photocathode that receives scintillation light and generates photoelectrons, a multi-stage dynode that applies an electric field to accelerate the generated photoelectrons, and an anode that is the outlet of the electron flow, and multiplies the scintillation light output from the scintillator and converts it into an electrical signal.

[0015] In addition, the gantry device 10 generates count information from the output signal of the detector 1 by the timing information acquisition circuit 102 and stores the generated count information in the storage unit 130 of the console device 20. Note that the detector 1 is divided into a plurality of blocks and includes the timing information acquisition circuit 102.

[0016] The timing information acquisition circuit 102 converts the output signal of the detector 1 into digital data and generates count information. This count information includes the detection position, energy value, and detection time of the annihilation gamma rays. For example, the timing information acquisition circuit 102 identifies a plurality of photodetector elements that have converted scintillation light into electrical signals at the same timing. Then, the timing information acquisition circuit 102 identifies the scintillator number (P) indicating the position of the scintillator where the annihilation gamma rays have entered. The means for identifying the position of the scintillator where the annihilation gamma rays have entered may be identified by performing a centroid calculation based on the position of each photodetector element and the intensity of the electrical signal. Also, when the element sizes of each of the scintillator and the photodetector element correspond, the scintillator corresponding to the photodetector element that has obtained an output may be identified as the position of the scintillator where the annihilation gamma rays have entered.

[0017] Also, the timing information acquisition circuit 102 identifies the energy value (E) of the annihilation gamma rays incident on the detector 1 by performing an integral calculation on the intensity of the electrical signals output from each photodetector element. Further, the timing information acquisition circuit 102 identifies the detection time (T) at which scintillation light due to the annihilation gamma rays has been detected by the detector 1. Note that the detection time (T) may be an absolute time or an elapsed time from the start time of imaging. In this way, the timing information acquisition circuit 102 generates count information including the scintillator number (P), energy value (E), and detection time (T).

[0018] Note that the timing information acquisition circuit 102 is realized by a circuit such as a CPU (Central Processing Unit), a GPU (Graphical Processing Unit), an application specific integrated circuit (ASIC), or a programmable logic device (e.g., a simple programmable logic device (SPLD), a complex programmable logic device (CPLD), and a field programmable gate array (FPGA)). The timing information acquisition circuit 102 is an example of a timing information acquisition unit.

[0019] The top plate 103 is a bed on which the subject P is placed and is disposed above the bed 104. The bed drive unit 105 moves the top plate 103 under the control of the bed control function 105d of the processing circuit 105. For example, the bed drive unit 105 moves the subject P into the imaging port of the gantry device 10 by moving the top plate 103.

[0020] The console device 20 receives operations of the PET device 100 by an operator, controls the imaging of PET images, and reconstructs PET images using the count information collected by the gantry device 10. As shown in FIG. 1, the console device 20 includes a processing circuit 105, an input device 110, a display 120, and a storage unit 130. Note that each unit included in the console device 20 is connected via a bus.

[0021] In the embodiment, each processing function performed by the specific function (simultaneous counting information generation function) 105a, the generation function 105b, the system control function 105c, the bed control function 105d, and the calculation function 105e is stored in the storage unit 130 in the form of a program executable by a computer. The processing circuit 105 is a processor that reads out and executes the program from the storage unit 130 to realize the functions corresponding to the respective programs. In other words, the processing circuit 105 in the state of having read out each program has each function shown in the processing circuit 105 of FIG. 1. In FIG. 1, although the processing functions performed by the specific function (simultaneous counting information generation function) 105a, the generation function 105b, the system control function 105c, the bed control function 105d, and the calculation function 105e are described as being realized by a single processing circuit 105, it is also possible to configure the processing circuit 105 by combining a plurality of independent processors, and each processor realizes the function by executing the program. In other words, each of the above-described functions may be configured as a program, and even when one processing circuit 105 executes each program. As another example, it may be a case where a specific function is implemented in a dedicated independent program execution circuit.

[0022] In FIG. 1, the specific function 105a, the generation function 105b, the system control function 105c, the bed control function 105d, and the calculation function 105e are each an example of a specific unit, a generation unit, a system control unit, a bed control unit, and a calculation unit.

[0023] As used in the above description, the term "processor" means, for example, a CPU (Central Processing Unit), a GPU (Graphical Processing Unit), or a circuit such as an application specific integrated circuit (ASIC), a programmable logic device (e.g., a simple programmable logic device (SPLD), a complex programmable logic device (CPLD), and a field programmable gate array (FPGA)). The processor realizes its functions by reading and executing the programs stored in the storage unit 130.

[0024] The processing circuit 105 generates coincidence count information based on the count information regarding the detector 1 acquired by the timing information acquisition circuit 102 by means of a specific function (coincidence count information generation function) 105a, and stores the generated coincidence count information in the storage unit 130.

[0025] The processing circuit 105 reconstructs a PET image by means of a generation function 105b. Specifically, the processing circuit 105 reads out the time series list of the coincidence count information stored in the storage unit 130 by means of the generation function 105b, and reconstructs a PET image using the read time series list. Further, the processing circuit 105 stores the reconstructed PET image in the storage unit 130.

[0026] The processing circuit 105 performs overall control of the PET device 100 by controlling the gantry device 10 and the console device 20 by means of a system control function 105c. For example, the processing circuit 105 controls imaging in the PET device 100 by means of the system control unit 105c.

[0027] The processing circuit 105 controls the bed drive unit 105 by means of a bed control function 105d.

[0028] The input device 110 is a mouse, keyboard, etc. used by the operator of the PET device 100 to input various instructions and various settings, and transfers the input various instructions and various settings to the processing circuit 105. For example, the input device 110 is used to input a shooting start instruction.

[0029] The display 120 is a monitor, etc. referred to by the operator, and displays the respiratory waveform and PET image of the subject or displays a GUI (Graphical User Interface) for receiving various instructions and various settings from the operator under the control of the processing circuit 105.

[0030] The storage unit 130 stores various data used in the PET device 100. The storage unit 130 is composed of, for example, a memory, and as an example, is realized by a semiconductor memory element such as a RAM (Random Access Memory), a flash memory, or a hard disk, an optical disk, etc. The storage unit 130 stores count information which is information in which a scintillator number (P), an energy value (E), and a detection time (T) are associated, coincidence information in which a set of count information is associated with a coincidence No. which is a serial number of the simultaneous count information, a reconstructed PET image, etc.

[0031] As shown in FIG. 2, the storage unit 130 includes a first storage unit 131 for storing data related to normalization at the time of manufacturing and factory shipment, and a second storage unit 132 for storing data related to normalization performed when the device is used in a medical institution. The first storage unit 131 stores a first detection efficiency correction data 131a and a first reference crystal efficiency 131b described later. The second storage unit 132 stores a second reference crystal efficiency data 132a, a detection efficiency change amount data 132b, and a second detection efficiency correction data 132c described later.

[0032] Subsequently, the background according to the embodiment will be described.

[0033] The calibration of a PET device includes a process called normalization, which measures the detection efficiency for each scintillator (crystal) and creates correction data used in clinical image generation. Normalization is generally performed by collecting gamma rays from a calibration radiation source placed inside the gantry.

[0034] However, when performing normalization using a calibration radiation source, operator exposure may occur due to the handling of the calibration radiation source. For example, when adjusting the installation position of the radiation source, the distance between the operator and the radiation source is likely to be short for a relatively long time, so operator exposure may occur during the insertion / removal of the radiation source or during phantom creation.

[0035] In addition, variations in accuracy may occur due to variations in the work of the installation position accuracy of the radiation source and the non-uniformity of the isotope distribution inside the radiation source.

[0036] For example, when installing the radiation source inside the gantry, variations may occur depending on the alignment accuracy. Also, if normalization is performed with the wrong phantom installation position, there is a risk of creating incorrect correction data. In addition, when using unsealed radiation sources such as FDG, the uniformity of the radiation source may not be sufficient due to work errors such as insufficient agitation, and even in the case of sealed radiation sources, the uniformity of the radiation source shipped from the radiation source vendor may not be sufficient.

[0037] Therefore, in view of such a background, in the nuclear medicine diagnostic device according to the embodiment, based on the self-radiation emitted from the scintillator, for example, 176 radiation such as Lu, the detection efficiency for each scintillator is calculated. Thereby, operator exposure can be reduced, the uniformity of the radiation source can be maintained, and the accuracy of normalization can be improved.

[0038] Regarding such a configuration, it will be described with reference to FIGS. 3 to 7. The PET device according to the embodiment uses gamma rays from the autoradiological nuclides contained in the scintillator, measures the detection efficiency for each crystal, and compares the measured detection efficiency for each crystal with reference data or history, thereby detecting the change over time in the detection performance of the detector 1 of the PET device 100 and detecting a detection function abnormality.

[0039] First, the case of detecting the change over time in the detection performance of the detector 1 and the detection function abnormality using the data regarding the gamma-ray events of the single-mode data from the autoradiation in the scintillator will be described. Regarding such a configuration, it will be described with reference to FIGS. 3 and 4. FIG. 3 shows the overall flow of such processing, and FIG. 4 details the processing of step S10 in FIG. 3.

[0040] First, in FIG. 3, in step S10, the processing circuit 105 measures the detection efficiency for each crystal using the gamma rays from the autoradiological nuclides contained in the scintillator by the calculation function 105e. Regarding such processing, it will be described with reference to FIG. 4. FIG. 4 details step S10 in FIG. 3 in more detail. That is, steps S11 to 14 in FIG. 4 correspond to step S10 in FIG. 3.

[0041] First, in step S11, the timing information acquisition circuit 102 performs data collection in single-mode data with the radiation source not placed in the PET gantry, and collects, for example, data on gamma rays from the autoradiological nuclides contained in the scintillator. Here, the single-mode data is data representing a single photon detection event and has information on the position, energy, and detection time of the detection element. The processing circuit 105 acquires the collected single-mode data from the timing information acquisition circuit 102 by the calculation function 105e.

[0042] Also, examples of the autoradiation contained in the scintillator include, for example, lutetium nuclides. For example, the scintillator of the nuclear medicine diagnostic device according to the embodiment is, for example 176The gamma rays emitted from Lu are emitted as radiation due to the self-radioactivity contained in the scintillator.

[0043] Subsequently, in step S12, the processing circuit 105 applies an energy window to the single-mode data acquired in step S11 by the calculation function 105e to extract events in a specific energy range. Specifically, the processing circuit 105 uses, by the calculation function 105e, an energy window that extracts only events in a specific energy range, for example, to remove circuit noise, scattered photons, simultaneous detection of multiple photons (pile-up), etc., thereby extracting only events in the specific energy range. As an example, 176 When using Lu as the self-radioactivity, 176 Lu emits gamma rays of 88, 202, 307, and 401 keV, but since the dominant ones are 202 keV and 307 keV, the processing circuit 105 applies, by the calculation function 105e, an energy window of, for example, 180 to 330 keV to extract gamma rays with an energy of 180 to 330 keV and reject the rest.

[0044] Subsequently, in step S13, the processing circuit 105 holds, by the calculation function 105e, the number of detections of the events extracted in step S12 for each crystal constituting the detector 1. As an example, the processing circuit 105 holds, by the calculation function 105e, the number of detections of the events extracted in step S12 for each crystal constituting the detector 1 as a histogram. Note that the storage unit 130 may hold the number of detections of the events extracted for each crystal in step S13 in the first storage unit or the second storage unit 132.

[0045] Subsequently, in step S14, the processing circuit 105 divides, by the calculation function 105e, the number of detections [count] held for each crystal in step S13 by the acquisition time [seconds] to calculate data in the unit of counts / second [cps].

[0046] Thus, the processing circuit 105 calculates the detection efficiency for each scintillator based on the radiation emitted from the scintillator by the calculation function 105e.

[0047] Returning to FIG. 3, in step S20, the processing circuit 105 calculates the detection efficiency after attenuation correction based on the detection efficiency of the detector before attenuation correction, the collection time of the reference data, and the collection time of the data to be corrected by the calculation function 105e. That is, since the self-radiation ability of the detector decreases with the passage of time as the radioactive substance decays, the self-radiation ability weakens with the passage of time. The processing circuit 105 corrects such an effect by the calculation function 105e. For example, the processing circuit 105 calculates the detection efficiency after attenuation correction according to the following formula (1) by the calculation function 105e.

[0048]

Equation

[0049] Here, p is the detector number, R corr.p is the detection efficiency after attenuation correction of detector p, R measure,p is the detection efficiency before attenuation correction (crystal efficiency) of detector p, T0 is the collection time of the reference data, T1 is the collection time of the new data, that is, the collection time when single-mode data was collected in step S11, and H is the half-life of the self-radioactive isotope contained in the scintillator. Here, the collection time of the reference data is, for example, the time when data was collected at the time of factory shipment.

[0050] Note that when the half-life of the self-radioactive isotope is long, the processing of this step can be omitted. For example, 176 The half-life of 10 Lu is as long as 3.78 * 10 176 years, so when normalization is performed using the self-radiation ability of

[0051] Subsequently, in step S30, the processing circuit 105 uses the calculation function 105a to compare the detected efficiency after attenuation correction calculated in step S20 with the existing data which is the reference data, calculates the change amount of the crystal efficiency, and detects an abnormality based on the calculated change amount of the crystal efficiency. That is, the processing circuit 105 uses the calculation function 105a to calculate the change amount of the crystal efficiency based on the radiation from the scintillator, and performs an abnormality detection process based on the calculated change amount of the crystal efficiency.

[0052] As an example, the processing circuit 105 uses the calculation function 105a to compare with the existing data, and determines an element without signal output as an abnormal element, thereby detecting an abnormality.

[0053] Also, as another example, the processing circuit 105 uses the calculation function 105a to divide the detected efficiency calculated in step S20 by the existing data for each element, obtains the change rate of the detected efficiency, and detects an abnormality based on the obtained change rate.

[0054] For example, based on the change rate of the detected efficiency calculated by the processing circuit 105 and the calculation function 105a, an abnormality in the system of the entire device or an abnormality in specific hardware is detected. For example, when the change rate of the detected efficiency calculated by the processing circuit 105 exceeds a threshold value, for example, ±5% across the entire detector, the processing circuit 105 determines that a system abnormality has occurred in the entire device.

[0055] Also, the processing circuit 105 uses the calculation function 105a to calculate the average value of the change rate of the detected efficiency in units of hardware such as the scintillator, detector unit, power supply unit, data transmission path, etc. When the average value of the change rate of the detected efficiency exceeds a predetermined threshold value in a specific hardware part, the processing circuit 105 determines that an abnormality has occurred in the hardware part.

[0056] Next, an embodiment will be described in which normalization is performed by combining normalization with a radiation source during manufacturing and single-mode data measurement using the above-described self-radiation to create data for crystal efficiency correction. First, the background for performing normalization using normalization with a radiation source and single-mode data measurement using self-radiation will be briefly described.

[0057] A histogram of gamma-ray events from radionuclides in the scintillator does not completely match a histogram based on gamma rays from a subject placed in the PET gantry. This is because, while the former can count scintillation light as detection events in the same way in all detection elements, in the latter, non-uniformity in the number of events occurs between detection elements due to scattering in the detector chassis and the like.

[0058] Therefore, in each medical device according to the embodiment, normalization with a radiation source and crystal efficiency calculation are performed during manufacturing, and at a medical institution, when the device is in operation, the change in pure detection efficiency obtained from collection without a radiation source is applied to the correction data.

[0059] That is, as shown in FIG. 5, first, in step S100, typically during the manufacture or factory shipment of the PET device 100, the PET device 100 performs a first normalization.

[0060] Thereby, more accurate normalization can be performed in consideration of the geometric arrangement of the detectors in the gantry and the like.

[0061] Note that the details of the first normalization process performed in step S100 are shown in FIG. 6, and step S100 in FIG. 5 corresponds to steps S101 to S103 in FIG. 6. Also, in step S200, typically at a medical institution or the like, for example, when the PET device 100 is in use, the PET device 100 performs a second normalization. The details of the second normalization process performed in step S200 are shown in FIG. 7, and step S200 in FIG. 5 corresponds to steps S210 to S240 in FIG. 7.

[0062] First, the first normalization process performed in step S100 will be described with reference to FIG. 6.

[0063] First, in step 101, the PET device 100 arranges a radiation source on the gantry device 10 and performs normal normalization with a radiation source by an external radiation source. The processing circuit 105 acquires data related to normal normalization with a radiation source, and the calculation function 105e calculates crystal efficiency correction data (A) (first detection efficiency correction data). The crystal efficiency correction data (A) is data including information on crystal efficiency measured in a state where a radiation source is arranged on the gantry device 10 during manufacturing. As another example, the crystal efficiency correction data (A) is data obtained by simulation. The crystal efficiency correction data (A) is correction data for absorbing and correcting differences in the geometric arrangement of the radiation source, such as the difference between Coincidence mode data and single mode data, and whether the radiation source is inside the gantry or inside the detector.

[0064] Note that in step 101, instead of performing the process of normal normalization by an external radiation source, the processing circuit 105 may calculate the crystal efficiency correction data (A) by creating a look-up table (LUT (Look-up Table)) by simulation such as Monte Carlo simulation using the calculation function 105e.

[0065] Subsequently, in step S102, with the radiation source not arranged in the gantry device 10, the PET device 100 performs the processes from step S11 to S14 in FIG. 4 using its own radioactivity, and the processing circuit 105 calculates the first reference crystal efficiency data (B) (first detection efficiency data) by the calculation function 105e. The first reference crystal efficiency data (B) is crystal efficiency data obtained using its own radioactivity during manufacturing, for example, crystal efficiency data collected in single mode.

[0066] Subsequently, in step S103, the storage unit 130 stores the detection efficiency correction data (A) calculated in step S101 and the first reference crystal efficiency data (B) calculated in step S102 in the first storage unit 131 together with the measurement time. That is, the storage unit 130 stores the crystal efficiency correction data (A) (first detection efficiency correction data) generated based on an external radiation source or simulation and the first reference crystal efficiency data (B) (first detection efficiency data) for each scintillator calculated based on the radiation emitted from the scintillator.

[0067] Thus, while the crystal efficiency correction data (A) (first detection efficiency correction data) is data generated based on the radiation emitted from the radiation source arranged in the gantry device 10, the first reference crystal efficiency data (B) (first detection efficiency data) and the second reference crystal efficiency data (C) (second detection efficiency data) described later are data calculated based on the data representing single photon detection events of the radiation emitted from the scintillator. In this way, both the data generated based on the radiation emitted from the radiation source arranged in the gantry device 10 during manufacturing and the data representing single photon detection events of the radiation emitted from the scintillator are acquired, so that it becomes possible to correct the effects of the geometric arrangement etc. in the gantry device 10 of the detector, and it also becomes possible to accurately correct the crystal efficiency without using an external radiation source except during manufacturing.

[0068] Subsequently, with reference to FIG. 7, the process of the second normalization performed in step S200 will be described.

[0069] First, in step S210, for example, in a medical institution or the like, with the radiation source not arranged on the gantry device 10, the processes from step S11 to S14 in FIG. 4 are performed using self-radiation ability. The processing circuit 105 calculates, by means of the calculation function 105e, the detection efficiency data before attenuation correction, which is the crystal efficiency data obtained using self-radiation ability during the use of the PET device 100. Except for the difference that step S210 is performed, for example, in a medical institution or the like during the use of the PET device 100, while step S102 is performed during manufacturing, factory shipment, etc., the processing of step S102 and the processing of step S210 are the same processing.

[0070] Subsequently, in step S220, the processing circuit 105 performs, by means of the calculation function 105e, the same processing as in step S20, and calculates, based on the detection efficiency data before attenuation correction calculated in step S210, the second reference crystal efficiency data (C), which is the crystal efficiency data after attenuation correction. The second reference crystal efficiency data (C) is the crystal efficiency data obtained using self-radiation ability and collected, for example, in single mode, during the use of the device in a medical institution or the like. That is, the processing circuit 105 calculates, by means of the calculation function 105e, the second detection efficiency data (C) for each scintillator calculated based on the radiation emitted from the scintillator.

[0071] Subsequently, in step S230, the processing circuit 105 calculates, from the generation function 105b, the crystal efficiency change amount data (D) by dividing the second reference crystal efficiency data (C) by the first reference crystal efficiency data (B). The second reference crystal efficiency data (C) and the first reference crystal efficiency data (B) have the same data collection method and only differ in the data collection time, and the influence due to the self-decay of the radionuclide has been corrected. Therefore, the crystal efficiency change amount data (D) is an amount indicating how the crystal efficiency has changed between the time of manufacturing when the first normalization was performed and the second normalization.

[0072] Subsequently, in step S240, the processing circuit 105 calculates the second detection efficiency correction data (E) by dividing the first detection efficiency correction data (A) by the crystal efficiency change amount data (D) using the generation function 105b. The first detection efficiency correction data (A) is data on the crystal efficiency when using a radiation source, and the crystal efficiency change amount data (D) is an amount indicating how the crystal efficiency has changed between the time of manufacture and the time when the second normalization was performed. Therefore, the second detection efficiency correction data (E) obtained through such a procedure becomes data indicating the crystal efficiency when assuming that normalization is performed using an external radiation source when the second normalization is executed. Therefore, by performing the processing from step S210 to step S240, appropriate normalization can be performed without using an external radiation source except during manufacture. In this way, the processing circuit 105 generates the second detection efficiency correction data (E) based on the first detection efficiency correction data (A), the first detection efficiency data (the first reference crystal efficiency data (B)), and the second detection efficiency data (the second reference crystal efficiency data (C)) using the generation function 105b.

[0073] Note that the PET device 100 can automatically execute the second normalization process in step S200. For example, the processing circuit 105 can automatically execute the processing in step S200 (i.e., the processing in steps S210 to 240) during the idle time of the PET device 100 using the calculation function 105e to automatically correct the crystal efficiency, or can automatically monitor the state of the PET device 100.

[0074] That is, the PET device 100 automatically measures the radiation emitted from the scintillator constituting the detector 1, for example, during the idle time of the PET device 100. The processing circuit 105 automatically calculates the second detection efficiency data based on the radiation emitted from the scintillator that has been automatically measured using the calculation function 105e. The processing circuit 105 automatically generates the second detection efficiency correction data based on the calculated second detection efficiency data using the generation function 105b.

[0075] In the case of such an embodiment, since the correction of the crystal efficiency is automatically executed, the PET apparatus 100 can be made maintenance-free.

[0076] As described above, according to the embodiment, since the radiation source is not handled, the exposure of the operator in the medical institution can be reduced. Also, in normalization (calibration of detection efficiency), the variation in work due to the installation position accuracy of the radiation source is reduced. Further, the variation in accuracy due to the non-uniformity of the isotope distribution inside the radiation source is reduced.

[0077] In addition, in some embodiments, normalization can be performed without the need for manual operations such as inserting and removing the radiation source. As a result, by automatically calculating the amount of change from the crystal efficiency standard during the apparatus idle time, it can be utilized as a monitoring system for the apparatus state. Also, by connecting the monitoring result to the correction data update, a maintenance-free configuration can be achieved.

[0078] That is, according to at least one of the embodiments described above, normalization can be performed while reducing the exposure of the operator, or normalization can be performed with high accuracy.

[0079] Although several embodiments have been described, these embodiments are presented as examples and are not intended to limit the scope of the invention. These embodiments can be implemented in various other forms, and various omissions, replacements, changes, and combinations of the embodiments can be made without departing from the gist of the invention. These embodiments and their modifications are included in the scope and gist of the invention, and are also included in the invention described in the claims and the equivalent scope thereof.

Explanation of Reference Numerals

[0080] 105 Processing circuit 105a Specific function 105b Generation function 105e Calculation function

Claims

**Claim 1**: A nuclear medicine diagnostic apparatus, comprising: a scintillator that emits self-radiation; a storage unit that stores first detection efficiency correction data generated based on an external radiation source or a simulation, and first detection efficiency data for each scintillator calculated based on radiation emitted from the scintillator; a calculation unit that calculates second detection efficiency data for each scintillator calculated based on radiation emitted from the scintillator; a generation unit that generates second detection efficiency correction data based on the first detection efficiency correction data, the first detection efficiency data, and the second detection efficiency data; wherein: the generation unit calculates crystal efficiency change amount data by dividing the second detection efficiency data by the first detection efficiency data, and generates the second detection efficiency correction data based on the crystal efficiency change amount data and the first detection efficiency correction data; the nuclear medicine diagnostic apparatus automatically measures radiation emitted from the scintillator during an idle time of the nuclear medicine diagnostic apparatus; the calculation unit automatically calculates the second detection efficiency data based on the automatically measured radiation emitted from the scintillator; and the generation unit automatically generates the second detection efficiency correction data based on the calculated second detection efficiency data. **Claim 2** The nuclear medicine diagnostic apparatus according to claim 1, wherein the scintillator emits gamma rays emitted from lutetium as the self-radiation. **Claim 3** The first detection efficiency correction data is data generated based on radiation emitted from a radiation source disposed on a gantry device, and the first detection efficiency data and the second detection efficiency data are data calculated based on data representing single photon detection events of the radiation emitted from the scintillator. **Claim 4** The nuclear medicine diagnostic apparatus according to claim 1, wherein the calculation unit calculates a change amount of crystal efficiency based on radiation emitted from the scintillator, and performs an abnormality detection process based on the calculated change amount of crystal efficiency. **Claim 5** The nuclear medicine diagnostic apparatus according to claim 1, wherein the calculation unit generates the first detection efficiency correction data by creating a look-up table through Monte Carlo simulation.

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