JP2002267757A - Nuclear medical imaging device - Google Patents

Abstract

PROBLEM TO BE SOLVED: To prevent the occurrence of a lack of distribution or slippage of distribution in an RI distribution image. SOLUTION: This nuclear medical imaging device is provided with a non- integrating structure capable of determining the optical energy incident on a photomultiplier 2 from the signal intensity of a photoelectric conversion signal detected at a timing after the lapse of a predetermined prescribed time from the start of the input of the photoelectric conversion signal outputted from a γ-ray detector 3 earlier than the integration by an integrator. Since the processing of the next photoelectric conversion signal can be immediately started after the detection of the signal intensity of the necessary photoelectric conversion signal, an account omission of γ-ray can be prevented to suppress the lack of distribution, and the error by the offset or thermal drift of the integrator can be avoided to precisely determine the optical energy quantity because no integrator is used. Consequently, the precision of γ-ray incident position data is improved, and the slippage of distribution can be suppressed.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to the detection of γ-rays emitted from a subject to which a radioactive isotope (RI = radioisotope) has been administered, and the detection of R at a site of interest in the subject.
The present invention relates to a nuclear medicine imaging apparatus for creating an I distribution image, and more particularly to a technique for suppressing the occurrence of distribution loss or distribution deviation in an RI distribution image.

[0002]

2. Description of the Related Art Conventionally, as a nuclear medicine imaging apparatus,
Radioisotopes to be administered to a subject (hereinafter referred to as “RI”
PET (Pos) using positron emission type RI
ItronEmission Tomography) device. As shown in FIG. 10, the PET apparatus has a scintillator block 51 that generates light when a γ-ray emitted from a subject (not shown) to which a positron emission type RI is administered enters.
And a photomultiplier 52 that receives light emitted from the scintillator block 51 and outputs a photoelectric conversion signal.

[0003] In addition, this conventional device calculates the amount of light energy incident on each photomultiplier 52 from the photoelectric conversion signal output from each photomultiplier 52, as shown in FIG. An incident position data obtaining unit 54 for obtaining γ-ray incident position data for creating an RI distribution image in each scintillator block based on the energy amount is provided. Then, an RI distribution image of a site of interest of the subject is created according to the γ-ray incident position data obtained by the incident position data obtaining unit 54.

In the case of a PET device, a γ-ray detector 53
In the figure, a large number of scintillator blocks 51 are arranged in a ring shape so as to surround the subject along the periphery of the subject, and a large number of γ-ray incident positions obtained and collected by the incident position data obtaining unit 54. Some apparatuses are configured so that an image reconstruction process is performed according to data and an RI distribution CT image (RI distribution computed tomographic image) is created.

[0005]

However, in the case of the conventional PET apparatus, there is a problem in that the RI distribution image is likely to have a missing distribution or a deviation in the distribution. That is, in the RI distribution image, the distribution is likely to be lost due to omission of counting of γ rays (counting omission), and the displacement of the distribution is likely to occur due to insufficient accuracy of the γ-ray incident position data.

In the case of a conventional PET apparatus, as shown in FIG. 12, an incident position data obtaining section 54 amplifies a pulse-like photoelectric conversion signal output from a photomultiplier 52 by a preamplifier section 55 and then amplifies the signal. The amount of light energy incident on each photomultiplier is obtained by integrating the photoelectric conversion signal obtained by the integrator 56. further,
The γ-ray incident position data for creating the RI distribution image in each scintillator block 51 is calculated and calculated based on the obtained light energy amount.

Accordingly, during the time when one photoelectric conversion signal is being integrated by the integrator 56 (integration time = dead time), the next photoelectric conversion signal cannot be received, so that γ-rays that cause count omission are generated. Come out. Further, the integration time in the integrator 56 is substantially equal to the time length of the photoelectric conversion signal, and the length of the photoelectric conversion signal is about 2 to 3 times the light (fluorescence) half time τ of the crystal of the scintillator block 51. In particular, B often used for the scintillator block 51
Since the GO crystal has a long light half-life τ, the integrator 5
6, the integration time, that is, the dead time becomes longer, and the count omission of γ-rays increases.

Further, the integrator 56, normally, there are offset and thermal drift, offset and thermal drift of the integrator 56 error of the integrated value, i.e. appears as a light energy amount of the error. That is, the error in the amount of light energy causes a decrease in the accuracy of the γ-ray incident position data. Error of the light energy by the integrator offset and the thermal drift, is proportional to the integration time. Therefore, when a BGO crystal having a long integration time is used, the error in the amount of light energy increases, and the accuracy of the γ-ray incident position data becomes insufficient.

The present invention has been made in view of such circumstances, and a nucleus capable of producing an accurate RI distribution image by suppressing a distribution loss or a deviation in distribution due to omission of counting of γ-rays. It is a primary object to provide a medical imaging device.

[0010]

In order to solve the above-mentioned problems, a nuclear medicine imaging apparatus according to the first aspect of the present invention comprises a gamma ray emitted from a subject to which a radioisotope (RI) is administered. Γ-ray detecting means in which a scintillator block which generates light by light, a photomultiplier which receives light emitted from the scintillator block and outputs a photoelectric conversion signal, and photoelectric conversion output from each photomultiplier Based on the signal, the amount of light energy incident on each photomultiplier is obtained, and based on the obtained amount of light energy, γ-ray incident position data for creating an RI distribution image in each scintillator block is obtained. (A) the incident position data obtaining means includes a photoelectric conversion signal. Is configured to calculate the amount of light energy incident on each photomultiplier based on the signal intensity of the photoelectric conversion signal detected at a timing after a predetermined time elapses from the start of input of the photomultiplier.

According to a second aspect of the present invention, there is provided the nuclear medicine imaging apparatus according to the first aspect, wherein (b) the incident position data obtaining means detects the photoelectric conversion signal before detecting the signal intensity of the photoelectric conversion signal. There is provided filtering means for performing filtering processing for noise removal on the converted signal.

According to a third aspect of the present invention, in the nuclear medicine imaging apparatus according to the first or second aspect,
(C) In the γ-ray detector, the scintillator blocks are arranged facing each other across the patient, the two traveling toward the opposite direction occur simultaneously with the positron annihilation of positron-emitting radioisotope (RI) While two annihilation γ-rays are configured to be simultaneously detected, the RI distribution image is formed only according to the γ-ray incident position data when the two annihilation γ-rays are simultaneously incident on and detected by the scintillator blocks disposed opposite to each other. It is configured to be created.

[Operation] Next, the operation of the nuclear medicine imaging apparatus of the present invention will be described. According to the nuclear medicine imaging apparatus of the first aspect of the present invention, in the case where an RI distribution image is created, the γ-ray detecting means causes γ-rays emitted from the subject to which RI has been administered to enter the scintillator block. The generated light is received by the photomultiplier and output as a photoelectric conversion signal. In addition, the incident position data obtaining means detects each photomultiplier based on the signal intensity of the photoelectric conversion signal detected at a timing after a predetermined time has elapsed from the start of input of each photoelectric conversion signal from the γ-ray detection means. The amount of light energy incident on the pliers is determined. Further, γ-ray incident position data for creating an RI distribution image in each scintillator block is obtained based on the obtained light energy amount. Then, an RI distribution image of a site of interest in the subject is created according to the γ-ray incident position data obtained by the incident position data obtaining means.

Therefore, in the nuclear medicine imaging apparatus according to the first aspect of the present invention, unlike the conventional integration method in which the photoelectric conversion signal is integrated by the integrator to determine the amount of light energy incident on the photomultiplier, the photoelectric conversion signal is Non-integration (without integrator) for obtaining the amount of light energy incident on the photomultiplier from the signal intensity of the photoelectric conversion signal detected at a timing after a predetermined time has elapsed from the start of input.
System. In other words, after detecting the signal strength of the necessary photoelectric conversion signal, before obtaining all the photoelectric conversion signals, the processing of the next photoelectric conversion signal is started, so that the count omission of γ-rays can be suppressed. . Further, since there is no integrator, errors due to offset and thermal drift of the integrator are avoided, and the amount of light energy is accurately obtained, so that the accuracy of the γ-ray incident position data can be improved.

Further, the signal intensity of the photoelectric conversion signal detected at a timing after a predetermined time elapses from the start of input, and the signal profile (signal waveform) of the photoelectric conversion signal
Since there is a correlation between and, based on this correlation,
It is not necessary to completely integrate the photoelectric conversion signal from the beginning to the end. In other words, if the correlation between the signal strength of the photoelectric conversion signal and the signal profile of the photoelectric conversion signal is used, the signal strength of the photoelectric conversion signal detected at a timing after a predetermined time has elapsed from the start of input. , The integrated value of the photoelectric conversion signal, that is, the amount of light energy incident on the photomultiplier is obtained.

According to the nuclear medicine imaging apparatus of the present invention, the incident position data obtaining means includes:
By the filtering process of the filtering means, noise is removed from the photoelectric conversion signal before the signal strength of the photoelectric conversion signal is detected. Then, the signal strength of the photoelectric conversion signal having no noise is detected at a timing after a predetermined time elapses from the start of input, so that the signal strength of the photoelectric conversion signal is accurately detected.

Furthermore, according to the nuclear medicine imaging apparatus of the invention described in claim 3, two annihilation γ-rays traveling toward the opposite direction occur simultaneously with the disappearance of positron RI is released, the subject An RI distribution image is created only according to the γ-ray incident position data when the light is simultaneously incident on and detected by the scintillator blocks arranged opposite to each other.

[0018]

An embodiment of the present invention will be described below with reference to the drawings. FIG. 1 is a block diagram showing a configuration of a PET apparatus according to an embodiment of the nuclear medicine imaging apparatus of the present invention, FIG. 2 is a schematic diagram showing a main configuration of a γ-ray detector of the embodiment, and FIG. FIG. 4 is a schematic view showing the arrangement of a scintillator block and a photomultiplier in FIG. 4, FIG. 4 is a plan view showing a scintillator block of a γ-ray detector of the embodiment, and FIG. FIG.

As shown in FIG. 1, the PET apparatus according to this embodiment includes a scintillator block 1 in which light is generated when gamma rays emitted from a subject M to which a radioactive isotope (RI) is administered enter. A γ-ray detector 3 combined with a photomultiplier 2 that receives light emitted from the block 1 and outputs a photoelectric conversion signal, and each photomultiplier based on the photoelectric conversion signal output from each photomultiplier 2 together determine the amount of light energy incident on the 2, the incident position data Motomede unit 4 for determining the γ-ray incident position data for creation of RI distribution image at each scintillator blocks based on the amount of light energy is the calculated
And an image reconstruction unit for executing an image reconstruction process in accordance with the γ-ray incidence position data from the incidence position data obtaining unit 4 to create an RI distribution CT image (RI distribution computed tomographic image) of a site of interest in the subject M. 5 and an image display monitor 6 for displaying the created RI distribution CT image and the like. Hereinafter, the configuration of each unit of the embodiment apparatus will be specifically described.

In the γ-ray detector 3, as shown in FIGS. 2 and 3, four photomultipliers 2 are arranged on the side 1 B of the scintillator block 1 opposite to the γ-ray incidence surface 1 A. The block 1 and the photomultiplier 2 are arranged in a ring shape so as to surround the subject M around the subject M. Further, the scintillator blocks 1 are arranged opposite to each other with the subject M interposed therebetween, and two (extinction) γ-rays generated by the annihilation of the positron of the positron emission type RI and traveling in the opposite direction are in the opposite arrangement. The light enters the scintillator block 1 and is detected simultaneously.

That is, the incident position data obtaining section 4 checks the position and the detection timing of the scintillator block 1 (as will be described in detail later), and determines the relationship between the γ-ray detector 3 and the opposing arrangement across the subject M. image reconstruction unit 5 two simultaneously γ-rays in the scintillator block 1 only γ-ray incident position data when it is detected as the proper data in
Send to At this time, when γ-rays are detected by only one of the two scintillator blocks 1, the γ-rays are treated as noise instead of γ-rays generated by the disappearance of the positron, and are discarded without being sent to the image reconstruction unit 5. It has a configuration.

As a result, according to the PET apparatus of this embodiment, a RI distribution CT image can be created by administering a positron emission type RI as the RI to the subject M.

In the case of a PET device, usually, during photographing,
Moves in a direction parallel to the body axis Z of the subject M with the subject M mounted thereon according to the movement of the top driving unit 8, so that the imaging section (slice plane) of the subject M changes (scans).
Configuration.

The γ-ray detector 3 may be of a rotating type in which the scintillator block 1 and the photomultiplier 2 detect γ-rays while rotating around the subject M during imaging, or may be a scintillator block 1 and a photomultiplier. It may be a stationary type that detects γ-rays while the pliers 2 are stationary.

The control of the cooperative operation of the γ-ray detector 3, the incident position data obtaining section 4, the image reconstructing section 5, the image display monitor 6, the top board driving section 8 and the like in the apparatus of this embodiment is performed by a console. 9 with the input operation of 9 and the progress of the shooting
It has a configuration which is performed in accordance with a command signal outputted in time from 0.

Further, in the γ-ray detector 3, as shown in FIGS. 2 and 3, the rectangular scintillator block 1
A light guide 11 for light diffusion is interposed between the scintillator block 1 and the photomultiplier 2 while the line incident surfaces 1A are arranged adjacent to each other in the plane direction.

As a specific arrangement of the scintillator blocks 1 in the γ-ray detector 3, for example, two scintillator blocks 1 are arranged in a direction parallel to the body axis Z of the subject M, and the body axis Z of the subject M In the direction around, an arrangement form in which many scintillator blocks 1 are arranged is exemplified. In addition, the arrangement of the scintillator is not limited to the above arrangement.

The individual scintillator blocks 1 are
As shown in FIG. 4, and the scintillator crystal piece 1P of a large number of squares is constituted by together closer to the aspect, the number of crystal pieces 1P per scintillator block, for example, the case of FIG. 4, the vertical 8 X = 8 in total and 64 in total. The shape and the number of crystal pieces 1P are not limited to those illustrated here.

As a specific material of the crystal piece 1P, BGO (Bi ₄ Ge ₃ O ₁₂ ) crystal or the like is used, but LSO (Lu ₂ SiO ₅ : Ce) crystal can also be used. Are not limited to those exemplified here.

In the case of the γ-ray detector 3 of the embodiment,
The line of the crystal piece 1P that continues in a ring shape in the direction around the body axis Z of the subject M corresponds to one slice section, and in the case of the embodiment apparatus, γ of eight slice sections is used.
The structure is such that lines can be detected simultaneously.

Next, a description will be given of the incident position data obtaining unit 4 which is a characteristic component of this embodiment. In the case of the incident position data obtaining unit 4, as shown in FIG. 5, a preamplifier unit 4A for amplifying the photoelectric conversion signal, and a low-pass (low frequency) for removing the high-frequency noise with respect to the photoelectric conversion signal amplified by the preamplifier unit 4A. A filtering unit 4B for performing a filtering process, and converting the analog photoelectric conversion signal from which noise has been removed by the filtering unit 4B into a digital signal at a timing after a predetermined time has elapsed from the start of input of the photoelectric conversion signal. An AD converter 4C for detecting the signal intensity of the photoelectric conversion signal, and the amount of light energy incident on each photomultiplier 2 is obtained based on the digital signal output from the AD converter 4C (that is, the signal intensity of the photoelectric conversion signal). The required number of signal processing systems including the energy amount calculation unit 4D are provided. Further, an incident position calculating unit 4E for obtaining γ-ray incident position data for creating an RI distribution image in the scintillator block 1 based on the amount of light energy obtained by the energy amount calculating unit 4D is provided downstream of the energy amount calculating unit 4D. ing.

Further, in the incident position data calculating section 4, the γ-ray incident position data output from the incident position calculating section 4E is generated by the disappearance of the positron and the two γ rays traveling in opposite directions are arranged in a facing relationship. Are provided to the two scintillator blocks 1 respectively, and a coincidence unit 4F for checking whether or not they are simultaneously detected is provided. In addition, the photoelectric conversion signal is directly input from the photomultiplier 2, the AD conversion timing of the photoelectric conversion signal by the AD conversion unit 4 </ b> C is controlled based on the input time point of the photoelectric conversion signal, and the energy amount calculation unit 4 </ b> D A timing control unit 4G is provided to control the calculation timing of the position calculation unit 4E and further control the check execution timing of the coincidence unit 4F.

Next, a process from when the photoelectric conversion signal is input to the incident position data calculating section 4 to when the γ-ray incident position data is output from the incident position data calculating section 4 will be described. Gamma rays are incident on the scintillator block 1, and a photoelectric conversion signal S as shown in FIG. 6 is output from the photomultiplier 2 to the preamplifier unit 4A and the timing control unit 4G. Photoelectric conversion signal S the preamplifier 4A is sent is amplified to the filtering unit 4B. On the other hand, in the timing control unit 4G, the timing control operation starts in response to the input of the photoelectric conversion signal S.

In the filtering section 4B, high-frequency noise is removed from the photoelectric conversion signal S by low-pass filtering, so that the photoelectric conversion signal S becomes a smooth signal without noise as shown in FIG. Converter 4
Sent to C.

In the AD converter 4C, the timing controller 4
According to the control by G, the photoelectric conversion signal S is converted into a digital signal at a timing after a predetermined time has elapsed from the start of the input of the photoelectric conversion signal S. This digital signal is
This is the signal strength Sd detected at a timing after a predetermined time has elapsed from the start of the input of the photoelectric conversion signal S. Usually, one photoelectric conversion signal S is converted into a digital signal several times at different timings. A plurality of signal intensities Sd
Is to be obtained.

When the signal passes through the filtering unit 4B, FIG.
As shown in the figure, since the rise of the photoelectric conversion signal S is slightly dull, the first A
Perform D conversion. For example, it is preferable that the timing of the first AD conversion comes after a time Ts which is twice the time Tp from the start of input of the photoelectric conversion signal S to the signal peak Sp.

Further, in order to prevent omission of the count of the next incident γ-ray, it is desirable to set the timing of the last AD conversion as early as possible. That is, the timing at which the AD conversion is performed several times while the remaining signal is not attenuated and a plurality of signal intensities Sd can be detected from the photoelectric conversion signal is appropriate. For example, the timing of the last AD conversion is preferably within twice the optical half-life τ of the scintillator block 1, and more preferably within 1.5 times the optical half-life τ.

The number of signal intensities Sd to be detected for one photoelectric conversion signal S differs depending on the length of the photoelectric conversion signal S, the state of noise, and the like. For example, the number is preferably 5 or less.

The signal intensity Sd detected by the AD converter 4C
Is sequentially sent to the energy amount calculating unit 4D, and the energy amount calculating unit 4D
There obtaining the amount of light energy Eg incident on the photomultiplier 2, based on the total signal strength Sd.

As an operation method in the energy amount operation unit 4D, for example, after performing a signal reproduction process of reproducing the signal profile of the photoelectric conversion signal from the obtained signal intensity Sd by a digital operation process, and further integrating the reproduced signal. A method of calculating the light energy amount Eg by calculating the value by digital operation processing is exemplified. At this time, if the number of signal strengths Sd is large, the reproduction accuracy of the signal is improved. In addition, the amount of light energy Eg is obtained earlier by using digital arithmetic processing than when integrating using an integrator. The energy amount calculation unit 4D that performs such calculation includes, for example, commercially available DS.
P (Digital Signal Processor) can be used.

The above-mentioned operation method is, for example, as follows. The photoelectric conversion signal S can be substantially expressed by the following equation (1). S = Aexp (−t / τ) (1) where A is a signal intensity constant, exp is an exponential function e, t is the elapsed time from the start of light emission, and τ is the half-life of light of the crystal.

On the other hand, the light energy amount Eg is represented by the following equation (2). Eg = ∫Aexp (−t / τ) dt (from t = 0 to t = ∞) (2) From equation (1), since A = S / [exp (−t / τ)], light emission By measuring the signal strength Sd at the timing t1 after the elapse of a predetermined time from the start, A = (S
d) / [exp (−t1 / τ)] and A is immediately obtained.
Also, the crystal half-life τ is crystal piece 1
It is determined from the material of P. Therefore, in equation (2),
Since A and τ are known, the light energy amount Eg can be calculated by arithmetic processing without integrating the actual photoelectric conversion signal.

That is, the light energy amount Eg is obtained earlier than the integration by the high-speed arithmetic processing, and the processing is started earlier for the next photoelectric conversion signal, so that the dead time is shortened.

The light energy Eg obtained by the energy calculator 4D is sent to the incident position calculator 4E. In the incident position calculating section 4E of this embodiment apparatus, the ratio of the amount of light energy Eg of the four photomultiplier 2 are disposed for each scintillator block 1,
The calculation of the incident position of the γ-ray in the scintillator block 1 is performed to obtain the γ-ray incident position data, which is then output to the coincidence unit 4F.

In the coincidence section 4F, the γ-ray detector 3
It is checked whether or not the γ-ray incident position data has been simultaneously obtained for the scintillator blocks 1 which are arranged so as to face each other across the subject M. When the γ-ray incident position data is obtained at the same time by the two scintillator blocks 1 that are in a facing relationship, the γ-ray incident position data is sent to the image reconstruction unit 5 as appropriate data. Otherwise, it is treated as noise, and the γ-ray incident position data is discarded without being sent to the image reconstruction unit 5.

Subsequently, the PET of the embodiment having the above configuration
The operation of the apparatus when creating an RI distribution CT image by the apparatus
This will be described with reference to the drawings. FIG. 8 is a flowchart showing a process of creating an RI distribution CT image by the apparatus of this embodiment.

[Step S1] Output of a photoelectric conversion signal [0047] As the γ-rays emitted from the subject M to which the positron emission type RI has been administered enter the scintillator block 1 of the γ-ray detector 3, a photomultiplier is generated. 2, the photoelectric conversion signal S is output to the incident position data obtaining unit 4. Since there are four photomultipliers 2 for one scintillator block 1, four photoelectric conversion signals S are normally output.

[0048] After the high-frequency noise is removed from the photoelectric conversion signal by [Step S2] High-frequency noise removal and the photoelectric conversion signal of the signal intensity of the detection incident position filtering section 4B of the data Motomede unit 4, the incident position data Motomede unit 4 by the A / D converter 4C, the A / D conversion of the photoelectric conversion signal S is performed at a timing after a predetermined time has elapsed from the start of the input of the photoelectric conversion signal S, and the photoelectric conversion signal S
Is detected.

[Step S3] Determining the amount of light energy of the photomultiplier Based on the detected signal intensity Sd of the photoelectric conversion signal S, the amount of light energy Eg of the photomultiplier 2 is obtained by the energy amount calculation unit 4D.

[Step S4] Determining γ-ray incident position data Based on the light energy Eg of the photomultiplier 2, the incident position calculator 4E calculates the γ-ray incident position in the scintillator block 1.

[Step S5] Confirmation of γ-ray incident position at opposing position Whether or not γ-ray incident position data has been determined for the scintillator blocks 1 that are arranged to oppose each other across the subject M by the coincidence unit 4F. A check is made. If the γ-ray incident position data is simultaneously obtained for the scintillator blocks 1 that are in a facing relationship, the process proceeds to the next step S6; otherwise, the process proceeds to the previous step S7.

[Step S 6] Outputting the γ-ray incident position data to the image reconstructing unit Since the γ-ray incident position data is proper data, it is sent to the image reconstructing unit 5.

[Step S7] Rejection of γ-ray incident position data Since the γ-ray incident position data is not proper data, it is rejected and eliminated (not collected) and is not sent to the image reconstruction unit 5.

[Step S8] Whether or not to collect γ-ray incident position data Whether or not to continue collecting γ-ray incident position data is checked. If the collection of the γ-ray incident position data is continued, the process returns to step S1 and the above-described operation is repeated. γ
When the collection of the line incident position data is completed, the process proceeds to the next step S9.

[Step S9] Creating an RI distribution CT image An RI distribution CT image is created according to the γ-ray incident position data collected by the image reconstruction unit 5.

[Step S10] Display of the Created RI Distribution CT Image When the creation of the created RI distribution CT image is displayed on the screen of the image display monitor 6, the creation process of the RI distribution CT image ends.

As described above, the PET of this embodiment
In the case of the device, a configuration is provided in which the light energy amount Eg incident on the photomultiplier 2 in a non-integral (integrator-less) method can be obtained earlier than the integration by the integrator. In other words, after detecting the required signal strength of the photoelectric conversion signal,
Because lean on taken without waiting until the photoelectric conversion signal ends the processing of the next photoelectric conversion signals, can suppress the counting leakage γ rays, lack of distribution it can be suppressed. In addition, since there is no integrator, errors in the amount of light energy due to offsets and thermal drift of the integrator can be avoided, and the amount of light energy can be determined accurately, so that the accuracy of the γ-ray incident position data is improved and the deviation of the distribution is eliminated. Is done. As a result, the RI distribution image becomes accurate.

In particular, BGO (Bi ₄ Ge ₃ O) crystal crystal 1P has advantages such as low cost, but has a long light half-life τ.
₁₂ ) In the case of a crystal, the dead time is greatly reduced because the processing time for obtaining the amount of light energy is greatly reduced. That is, the count leakage of the γ-ray is sufficiently suppressed, and a large error due to the offset and the thermal drift of the integrator is avoided, so that the light energy amount can be obtained very accurately.
As a result, the accuracy of the γ-ray incident position data is significantly improved, and the deviation of the distribution is eliminated.

In the case of the PET device of this embodiment, the signal strength of the photoelectric conversion signal is detected after noise is removed from the photoelectric conversion signal by the filtering unit 4B, so that the signal strength of the photoelectric conversion signal is accurately determined. Is detected. That is, the amount of light energy incident on the photomultiplier is more accurately obtained, and the accuracy of the γ-ray incident position data is further improved.

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

(1) In the above embodiment, as shown in FIG. 5, the filling section 4B is provided at the preceding stage of the AD conversion section 4C to remove the high frequency noise from the photoelectric conversion signal S. For example, as shown in FIG. 8, the photoelectric conversion signal S amplified by the preamplifier unit 4A without providing the filtering unit 4B may be immediately sent to the AD conversion unit 4C in the case of the photoelectric conversion signal S having a small number. In addition,
Other configurations are the same as those of the above embodiment.

In the case of the device of this modification, since the rising of the photoelectric conversion signal S due to the removal of the high-frequency noise by the filtering unit 4B does not become dull, the detection timing of the signal intensity Sd of the photoelectric conversion signal S is advanced to reduce the dead time. It can be shorter.

(2) In the above embodiment, the CT type P
Although the nuclear medicine imaging apparatus of the present invention is not limited to a CT type PET apparatus, for example, a non-C
A T-type PET device may be used, or a SPECT device using a normal single-photon type RI whose RI is not a positron emission type may be used.

[0064]

As is apparent from the above description, according to the nuclear medicine imaging apparatus of the first aspect, the γ in which the scintillator block is combined with the photomultiplier.
The amount of light energy incident on each photomultiplier is determined by an integrator based on the signal intensity of the photoelectric conversion signal detected at a timing after a predetermined time has elapsed from the start of input of the photoelectric conversion signal output from the line detection means. It has a non-integral (integrator-less) configuration that is found faster than the integration. That is, after detecting the necessary signal strength of the photoelectric conversion signal, the processing of the next photoelectric conversion signal is started without waiting for the end of the photoelectric conversion signal, so that the count omission of γ-rays can be suppressed.

Further, since there is no integrator, errors due to offset and thermal drift of the integrator are avoided, and the amount of light energy is accurately obtained, so that the accuracy of the γ-ray incident position data can be improved.

Therefore, in the nuclear medicine imaging apparatus according to the first aspect of the present invention, the omission of the distribution is suppressed by suppressing the count leakage of the γ-rays, and the deviation of the distribution is eliminated by improving the accuracy of the γ-ray incident position data. An RI distribution image is created.

According to the nuclear medicine imaging apparatus of the second aspect, the incident position data obtaining means removes noise from the photoelectric conversion signal before the signal intensity of the photoelectric conversion signal is detected by the filtering means. After that, a configuration is provided in which the signal intensity of the photoelectric conversion signal is detected at a timing after a predetermined time has elapsed from the start of the input of the photoelectric conversion signal. That is, since the signal intensity of the photoelectric conversion signal is accurately detected, the amount of light energy incident on the photomultiplier is more accurately obtained, and the accuracy of the γ-ray incident position data is further improved.

Further, according to the nuclear medicine imaging apparatus of the third aspect of the present invention, two annihilated γ-rays which are generated simultaneously with the annihilation of the positron emitted by the RI and travel in the opposite direction are detected by the γ-ray detecting means. Since the RI distribution image is created only according to the γ-ray incident position data at the time of detection,
As described above, a positron emitting element can be administered to a subject to create an RI distribution image.

[Brief description of the drawings]

FIG. 1 is a block diagram illustrating an overall configuration of a PET apparatus according to an embodiment.

FIG. 2 is a schematic diagram illustrating a configuration of a main part of a γ-ray detector of the embodiment device.

FIG. 3 is a schematic diagram showing an arrangement state of a scintillator block and a photomultiplier in a γ-ray detector of the example device.

FIG. 4 is a plan view showing a scintillator block of the γ-ray detector of the example device.

FIG. 5 is a block diagram illustrating a configuration of an incident position data obtaining unit of the apparatus according to the embodiment.

FIG. 6 is a signal waveform diagram showing a photoelectric conversion signal before noise removal.

FIG. 7 is a signal waveform diagram showing a photoelectric conversion signal after noise removal.

FIG. 8 is a flowchart showing a process when an RI distribution CT image is created by the apparatus of the embodiment.

FIG. 9 is a block diagram illustrating a configuration of an incident position data obtaining unit of the modification example device.

FIG. 10 is a schematic diagram showing a configuration of a main part of a γ-ray detector of a conventional device.

FIG. 11 is a block diagram showing a γ-ray detector and an incident position data obtaining unit of the conventional device.

FIG. 12 is a block diagram showing a configuration around an integrator for integrating a photoelectric conversion signal in an incident position data obtaining unit of the conventional device.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 ... Scintillator block 2 ... Photomultiplier 3 ... Gamma ray detector 4 ... Incident position data obtaining part 4A ... Preamplifier part 4B ... Filtering part 4C ... AD conversion part 4D ... Energy amount calculating part 4E ... Incident position calculating part 4F ... Coincidence part M: subject S: photoelectric conversion signal Sd: signal intensity of photoelectric conversion signal

Claims (3)

[Claims] 1. A scintillator block for generating light upon incidence of gamma rays emitted from a subject to which a radioactive isotope (RI) has been administered, and receiving light emitted from the scintillator block to generate a photoelectric conversion signal. Γ-ray detecting means combined with an output photomultiplier, and the amount of light energy incident on each photomultiplier is determined based on the photoelectric conversion signal output from each photomultiplier. A nuclear medicine imaging apparatus comprising: an incident position data obtaining means for obtaining γ-ray incident position data for creating an RI distribution image in each scintillator block based on the amount; A signal of a photoelectric conversion signal detected at a timing after a predetermined time elapses from the start of input of the conversion signal. A nuclear medicine imaging apparatus configured to obtain the amount of light energy incident on each photomultiplier based on the intensity. 2. The nuclear medicine imaging apparatus according to claim 1, wherein (b) the incident position data obtaining means removes noise from the photoelectric conversion signal before detecting the signal intensity of the photoelectric conversion signal. A nuclear medicine imaging apparatus, comprising: a filtering unit for performing a filtering process. 3. The nuclear medicine imaging apparatus according to claim 1, wherein (c) in the γ-ray detecting means, a scintillator block is disposed to face the subject, and a positron emission type radioisotope is provided. A scintillator block configured to simultaneously detect two annihilation gamma rays generated simultaneously with the annihilation of the (RI) positron and traveling in the opposite direction, and the two annihilation gamma rays are arranged to face each other. At the time of simultaneous incidence and detection
A nuclear medicine imaging apparatus characterized in that an RI distribution image is created according to only line incident position data.