JP2005098708A - Positron emission computed tomography equipment - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a positron emission computed tomography equipment which meets various test conditions for objects to be tested and carries out tests in a short period of time. <P>SOLUTION: Four insertion holes are formed in septa 13 apart from each other in the circumferential direction, and four guide bars 15 which are disposed between outside septa 14, 14, and set apart from each other in the circumferential direction of an annular support section 12, and both ends of which are fixed and attached thereto, and the guide bars 15 are inserted into the insertion holes of the septa 13, respectively. Respective septa 13 which are positioned near the outside septa 14, are moved toward a body axis via the guide bars 15, thereby preventing radiation which is emitted from the outside of an interested region of the object to be measured 5, from entering a radiation detector 11 by using the septa 13. <P>COPYRIGHT: (C)2005,JPO&NCIPI

Description

  The present invention relates to a positron emission tomography apparatus (Positron Emission Computed Tomography, hereinafter referred to as PET).

  The inspection technique using radiation can inspect the inside of a subject nondestructively. In particular, PET is an example of a radiation inspection technique for the human body. PET is a technique for measuring a physical quantity of an integral value (flight direction) of radiation and calculating the physical quantity of each voxel in the subject by projecting the integral value and imaging it. According to this technology, although it is necessary to process a huge amount of data, it has become possible to provide high-speed and high-definition images with the rapid development of computer technology in recent years.

PET is positron emission ^{(15 O, 13 N, 11} C, 18 F , etc.), and materials having a property to collect in specific cells in the body (hereinafter, referred to as a labeling substance) radiopharmaceutical comprising (hereinafter, pharmaceutical PET Is administered to a subject, and a site where a large amount of PET drug is consumed is examined. An example of a PET drug is fluorodeoxyglucose (2- [F-18] fluoro-2-deoxy-D-glucose, 18FDG). Since 18FDG is highly accumulated in tumor tissue by sugar metabolism, it is used to identify the tumor site. A positron emitted from a positron emitting nuclide contained in a PET drug accumulated at a specific location is combined with an electron in a nearby cell and disappears, and a pair of γ-rays having energy of 511 keV (hereinafter referred to as a γ-ray pair). ). Since γ-ray pairs are emitted in opposite directions (180 ° ± 0.6 °), if these γ-ray pairs are detected by a radiation detector, a positron is emitted between any two radiation detectors. You can see what was done. By detecting these many gamma ray pairs, a place where a lot of PET drug is consumed can be found. For example, 18FDG collects in cancer cells with intense glucose metabolism as described above, and thus it is possible to detect cancer foci by PET. The obtained data is converted into the radiation generation density of each voxel by the filtered back projection method described in Non-Patent Document 1, and the generation position of γ-rays (the position where the radionuclide accumulates, that is, the position of the cancer cell) is determined. Contributes to imaging. ¹⁵ O, ¹³ N, ¹¹ C, and ¹⁸ F used in PET are radioisotopes with a short half-life of 2 to 110 minutes.

Current PET requires tens of minutes for whole body examination. In PET, inspection time is one of the very important performances. By shortening the examination time, not only the burden on the subject is reduced, but also the movement of the subject under examination is reduced and the image quality is improved. In order to shorten the inspection time, it is only necessary to improve the sensitivity of the radiation detector. For this reason, a new type of scintillator has been studied. Moreover, substantial sensitivity can be improved by increasing the size of the gantry (detector ring) (increasing the body axis length). This increases the solid angle of the radiation detector with respect to the subject to improve the effective detection sensitivity.
IE Transaction on Nuclear Science NS-21 Volume pp. 228-229

  Incidentally, it is effective to increase the length of the gantry in the body axis direction (hereinafter referred to as gantry length) in order to shorten the PET examination time. However, if the gantry length is simply increased, the detected γ Lines increase, accidental coincidence or circuit data loss increases, and the effective sensitivity of the device saturates.

  The saturation point varies depending on the inspection conditions. The test conditions include the level of radioactivity administered to the subject, the time from drug administration to the test, the physique (height, weight) of the subject, and the like. For example, when the administered radioactivity level is high, increasing the gantry length increases data loss and does not increase substantial sensitivity. Conversely, when the administered radioactivity level is low, increasing the gantry length increases the substantial sensitivity without increasing data loss. That is, there is a problem that the optimum gantry length must be changed in accordance with various inspection conditions of the PET inspection, and the inspection time cannot be uniformly reduced with respect to various inspection conditions for the subject.

  In view of the above problems, an object of the present invention is to provide a positron emission tomography apparatus having a short examination time corresponding to various examination conditions for a subject.

  The positron emission CT apparatus of the present invention uses a septa to prevent radiation from entering the radiation detector from the region of interest of the subject. Therefore, by moving or exchanging the septa according to the test conditions of the subject (height, weight, dose of drug, time since the drug was administered, site where the drug is accumulated, etc.), the region of interest The size of can be changed.

  According to the present invention, it is possible to provide PET with a short examination time corresponding to various examination conditions by moving or exchanging the septa according to various examination conditions for the subject.

  Hereinafter, embodiments of the present invention will be described with reference to the accompanying drawings of FIGS.

  FIG. 1 is an overall configuration diagram showing a PET according to the present embodiment, and FIG. 2 is a plan view showing a septum in FIG. 1 alone. FIG. 3 is an enlarged view showing the imaging apparatus in FIG. 1, and FIG. 4 is an enlarged view similar to FIG. 3 showing a state in which one septa is provided.

  As shown in FIG. 1, the PET according to the present embodiment includes an imaging device 1, a subject holding device 2, a signal processing device 3, and a scepter moving device 4. The subject holding device 2 includes a support member 21 and a bed 22 provided on the upper end portion of the support member 21. The bed 22 holds the subject 5 and is movable in the longitudinal direction. The signal processing device 3 includes a signal amplification circuit 31, a signal discrimination circuit 32, a coincidence counting circuit 33, a computer (for example, a workstation) 34, a storage device 35, and a display device 36.

  The imaging device 1 is installed in a direction perpendicular to the longitudinal direction of the bed 22 and includes a radiation detector 11, an annular holder 12, septa 13, 13,. A large number of radiation detectors 11 are installed on the inner peripheral side of the annular holder 12 in the circumferential direction and the axial direction, and tens of thousands are arranged in an annular shape so as to surround the subject 5. The radiation detector 11 is a semiconductor detector and is made of 5 mm cubic cadmium tellurium (CdTe). The radiation detector 11 may be made of gallium arsenide (GaAs), cadmium tellurium zinc (CZT), or the like. The annular holding part 12 is installed on the support member 21 and is installed on the floor of the examination room.

  As shown in FIG. 2, the scepter 13 has four insertion holes 13A, 13A,. Further, between the outer septa 14, 14, both ends of four guide rods 15, 15,... Are fixedly attached at an interval in the circumferential direction of the annular holding portion 12, and each guide rod 15 is attached to the septa 13. Is inserted into the insertion hole 13A. For this reason, as shown in FIG. 3, each scepter 13 shifted to the outer scepter 14 side is configured to move in the body axis direction via the guide bar 15 as shown in FIG. The scepter 13 and the outer scepter 14 are formed using ring-shaped lead and shield γ rays.

  FIG. 3 shows an example in which one scepter 13 is used. When the number of gantry divisions is two, the single scepter 13 shown in FIG. 3 may be moved to the intermediate portion in the longitudinal direction of the annular holding portion 12. In the present embodiment, the maximum number of gantry divisions is four. In this case, as shown in FIG. 4, all three scepters 13 may be moved from the outer scepter 14 side in the body axis direction. Of course, when it is not necessary to divide the gantry, the scepters 3 need not be moved in the body axis direction.

  The septa moving device 4 includes a scepter holding unit driving device 41, a scepter holding unit 43, a guide rail 44, and a scepter moving device holding unit 42. The septa holder 43 is held by the scepter holder drive device 41 via the guide rail 44. The scepter holding unit driving device 41 is held on a scepter moving device holding unit 42, and the scepter moving device holding unit 42 is connected to the support member 21. The septa holder 43 moves up and down and expands and contracts in the body axis direction, and changes the number of gantry divisions by moving and replacing the septa 13. Inside the septa holding unit drive device 41, there are first and second drive mechanisms including a motor and a speed reduction mechanism (both not shown). The septa holder 43 moves up and down by the first drive mechanism and expands and contracts in the body axis direction by the second drive mechanism.

  Each radiation detector 11 is connected to a corresponding signal amplifying circuit 31. The signal amplifying circuit 31 amplifies a minute signal from each radiation detector 11 and generates a peak value corresponding to the detected energy of γ-rays. The timing signal which detected the gamma ray with the analog signal which has is each output.

  In FIG. 1, the signal amplifying circuit 31 is illustrated as being installed away from the imaging device 1, but in practice it is preferably installed near the radiation detector 11. The signal when the radiation detector 11 detects γ-rays is a very small signal, and the signal amplification circuit 31 should be installed near the radiation detector 11 from the viewpoint of noise.

  The signal discriminating circuit 32 receives the timing signal from the signal amplifying circuit 31 and measures the γ-ray detection time and the γ-ray energy via an analog-digital converter (not shown), and holds each as digital data. Then, the signal discriminating circuit 32 transfers the γ-ray detection time, the γ-ray energy, and the sensor number detecting the γ-ray to the coincidence counting circuit 33 as one packet data.

  The coincidence counting circuit 33 receives bucket data that is γ-ray detection data from the signal discriminating circuit 32, performs simultaneous measurement, generates a sensor number pair from the data pair determined to be simultaneous, and transfers it to the computer 34. . At the time of measurement by the coincidence circuit 33, the computer 34 stores data relating to the sensor number pair from the coincidence circuit 33 in the storage device 35. At the end of measurement, the computer 34 performs image reconstruction processing on the data stored in the storage device 35 and outputs a PET image to the display device 36.

  Before specifically describing the inspection in the present embodiment, a method of using the septa 13 will be described. This embodiment has been made based on the following examination by the inventors.

  FIG. 5 and FIG. 6 show the simulation results when the internally administered radioactivity is 15 mCi and 1 mCi. The subject 5 is simulated with a height of 170 cm and a weight of 60 kg. 5 and 6, the horizontal axis represents the gantry length, and the vertical axis represents NECR (Noise Equivalent Count Rate). NECR is an effective count rate in consideration of the noise component calculated from Equation 1 below. NECR and measurement time are inversely related.

NECR = C (t) / (1 + C (s) / C (t) + 2C (r) / C (t)) (Formula 1)

  C (t) in Equation 1 is the coincidence rate of true events (events that provide true information), and C (s) is the coincidence rate of scattered events. A scattering event is an event in the case of capturing gamma rays whose angle has changed due to internal scattering, and is a false event. Also, C (r) is the coincidence event coincidence rate. An incidental event is an event in which γ rays generated by another positron decay are captured accidentally, and is a false event.

  When the radioactivity administered into the body is large, specifically, as shown in FIG. 5, when the radioactivity administered into the body is, for example, 15 mCi, the NECR is saturated when the gantry length exceeds 200 to 250 mm. This is because the count rate of the incident event C (r) increases while the count rate of the detected γ rays increases as the gantry length increases.

  Therefore, in the case of FIG. 5, the septa 13 are arranged at three intervals of 150 mm, 300 mm, and 450 mm at equal intervals, and simultaneous measurement is performed only in an area partitioned by each septa 13. By arranging the septa 13 in this way, the region delimited by the septa 13, 13, 13 can block the incidence of γ rays from the region delimited by the other septa 13, 13, 13. Therefore, in each of the four areas partitioned by the septa 13, the NECR is 450 kcps, and the total is 1800 kcps (450 kcps × 4). When each septa 13 is excluded, it becomes 150 kcps as shown in FIG. 5, and the NECR can be increased 12 times by the septa 13. In other words, the measuring time can be reduced to 1/12 by the scepter 13.

  On the other hand, when the in-vivo administered radioactivity is small, specifically when the in-vivo administered radioactivity is, for example, 1 mCi, as shown in FIG. 6, it can be seen that NECR does not saturate even when the gantry length is 600 mm. This is because less γ rays are emitted from the body. Therefore, in the case of FIG. 6 (internally administered radioactivity 1 mCi), the septa 13 is not inserted. In this case, the NECR is 400 kcps. In the case of FIG. 6, if the septa 13 of FIG.

  In the present embodiment configured as described above, the attachment position of the septa 13 is variable in accordance with the magnitude of the in-vivo administered radioactivity, and the measurement time is obtained by attaching the ceptor 13 to the optimum position, removing the ceptor 13 or exchanging it. Can be shortened. The factor that determines the position of the septa 13 is the radioactivity level in the subject 5 in the gantry. The radioactivity level in the subject 5 in the gantry varies depending on the in-vivo radioactivity, rest time, height, weight, etc. of the subject 5. Rest time is the time from drug administration to measurement, and when a drug with a short half-life is used, the radioactivity level in the subject 5 at the time of measurement is lower than the dose. The height and weight of the subject 5 determine the degree of spatial drug condensation.

  A processing flow at the time of inspection by PET will be described. First, the height and weight of the subject 5 are measured, the in-vivo radioactivity and rest time are determined, the position of the septa 13 is determined from these conditions, and the septa 13 is inserted into the annular holder 12. deep. The method for determining the position of the scepter 13 from the various conditions may be based on a simulation result as in the present embodiment, or may be determined from a measured database by measuring the characteristic line in FIG. 5 in advance. Good.

  Next, a PET drug determined in advance by injection or the like is administered to the subject 5. The PET drug is selected according to the affected area to be examined. Thereafter, the subject 5 waits for a rest period until the PET drug diffuses in the body of the subject 5 and collects in the affected area (for example, an affected area of cancer) and becomes ready for imaging. After the rest time has elapsed, the bed 22 on which the subject 5 lies is inserted into the annular holding unit 12 together with the subject 5, and a PET examination is performed using the imaging apparatus 1. After the subject 5 to which the PET drug has been administered is inserted into the hole and a voltage is applied to each radiation detector 11 by a power source (not shown), each radiation detector 11 is released from the subject 5. Γ-rays are detected and PET inspection is started.

  The coincidence counting circuit 33 needs to divide a region to be simultaneously determined according to the attachment position of the scepter 13 and perform simultaneous measurement. Therefore, the area divided by the preset attachment position of the septa 13 is identified by the position information of the radiation detector 11, and the coincidence counting circuit 33 performs the simultaneous measurement with the data for each area.

  You may change the attachment position of the septa 13 during a measurement. In particular, when a radioisotope with a short half-life is used, the radioactivity is attenuated during the measurement. Therefore, the septa 13 is moved during the measurement according to the radioactivity level, and the optimum mounting position of the septa 13 is realized. It is also possible to do.

According to the present embodiment, the following effects can be obtained.
(1) According to the present embodiment, the effective detection sensitivity can be improved by inserting the septa 13 into the PET gantry and dividing the area to be simultaneously measured, and the measurement time is, for example, about 1/10. Can be shortened.
(2) According to the present embodiment, it is possible to determine the optimum attachment position of the scepter 13 corresponding to various inspection conditions, and to shorten the measurement time.
(3) According to the present embodiment, by changing the attachment position of the septa 13 during measurement, the optimum attachment position of the septa corresponding to the radioactivity level due to the half-life can be determined, and the measurement time can be shortened. Can do.
(4) According to the present embodiment, an optimum mounting position of the scepter 13 can be realized by using a plurality of scepters 13 corresponding to the number of divided gantry and moving or exchanging the scepters 13. Since the plurality of septa 13 are configured to be installed, the positioning operation of the septa 13 can be simplified, errors can be reduced, and the structure of the septa holding portion 43 and its moving mechanism can be simplified.
(5) According to the present embodiment, the individual septa 13 may be arranged at an optimum position. Accordingly, since the weight of the moving scepter 13 is reduced, the scepter holding portion 43 and its moving mechanism can be simplified. Further, it is not necessary to take out the septa 13 out of the gantry, and a space for replacing the septa 13 and a storage space for the unused septa 13 are not required. Therefore, the entire PET can be reduced in size.
(6) According to the present embodiment, the coincidence counting circuit 33 needs to divide and simultaneously measure the region to be simultaneously determined according to the attachment position of the scepter 13, and detect the region divided by the preset scepter position. The coincidence counting circuit 33 performs simultaneous measurement using data for each area. That is, it corresponds to the change of the attachment position of the septa 13 by signal processing. Therefore, it is possible to reduce the size and cost of the entire apparatus without increasing the coincidence counting circuit 33 that is hardware.

It is a whole lineblock diagram showing PET concerning an embodiment of the invention. It is a top view which shows the septa in FIG. 1 alone. It is an enlarged view which shows the scepter in FIG. It is an enlarged view which shows the other scepter which concerns on embodiment of this invention. It is a characteristic diagram which shows the relationship between the gantry length which concerns on embodiment of this invention, and NECR. It is another characteristic diagram which shows the relationship between the gantry length which concerns on embodiment of this invention, and NECR.

Explanation of symbols

DESCRIPTION OF SYMBOLS 1 Imaging device 2 Subject holding device 3 Signal processing device 4 Septa moving device 5 Subject 11 Radiation detector (semiconductor detector)
13 Septa

Claims (5)

  In a positron emission tomography apparatus that performs examination of a subject using radiation from within the subject, a radiation detector that detects radiation from within the subject, and a region of interest from outside the subject A positron emission tomography apparatus, comprising: a radiation detector that prevents radiation from entering the radiation detector; and a septum that separates the radiation detector and the region of interest into a plurality of regions, and images each separated region of interest. .   The positron emission tomography apparatus according to claim 1, wherein the size of the region of interest is changed by moving the septa according to the examination condition of the subject.   2. The positron emission tomography apparatus according to claim 1, wherein the size of the region of interest is changed by removing or replacing the septa according to the examination condition of the subject.   The positron emission tomography apparatus according to claim 1, wherein the radiation detector is a semiconductor detector.   The positron emission tomography apparatus according to claim 1, wherein the septa is formed using ring-shaped lead or tungsten.

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