JPH08136655A - Data collecting circuit for emission ct device - Google Patents

Abstract

PURPOSE: To eliminate the need of generating a picture due to an event by identifying the energy level of a α-ray event and, when the level is higher than a fixed level, adding '1' to the data at a designated address in a memory and, when the level is equal to or lower than the fixed level, subtracting '1' in accordance with the output of a frequency dividing circuit which makes weighting in accordance with the occurring frequency of events. CONSTITUTION: Two facing detectors input single event signals 1, 2, and 3 to an identifying circuit 4, simultaneous counting circuit 5, and address circuit 6, respectively. When the circuit 4 identifies that an event occurs in an energy window W2 a multiplexer 9 outputs the signal 10 of the circuit 5 as a signal 11 and adds '1' to the data at the address designated from the circuit 6 in a memory 20. When the circuit 4 identifies that an event occurs in another energy window W1 , the multiplexer 9 outputs the signal 10 as an event signal 12 and a frequency dividing counter 13 divides the frequency of the signal 12 at the value corresponding to a correction factor. When the counter 13 has an output, '1' is subtracted from the data in the memory 20. Therefore, the need of data transfer for correcting scattered light rays, reconstruction of a picture drawn by scattered light rays, etc., can be eliminated.

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a data collecting circuit for an emission CT apparatus, and more particularly to a data collecting circuit which is convenient for performing scattered ray correction.

[0002]

2. Description of the Related Art Single photon emission CT (S
PECT) equipment and positron emission CT (PE
It is extremely difficult to rigorously correct scattered radiation in an emission CT apparatus such as T) equipment because the distribution of scattered radiation differs depending on the shape of the subject and the radioisotope (RI) distribution in the subject. is there. SPECT
As one of the scattered radiation correction methods in the apparatus, Jazczak et al. Have a photoelectric peak (127 to 153 keV:
Second from the reconstructed image (P) to the projection data within ⁹⁹ mTc)
The reconstructed image (S) based on the projection data of the scattered radiation component within the energy window (92 to 125 keV) is subtracted to obtain the scattered radiation corrected image. Jezak et al. Multiply S by a correction factor of 0.5 obtained from an experiment in a phantom, and subtract this value from P. That is, the image (SPECT) after the scattered radiation correction is SPECT = P-0.5S
Is obtained. The literature is as follows. Jaszczak R.
J. et al ； Improved SPECT quantitation using compens
ation for scattered photons.J.Nucl.Med., 25,89
3, 1984.

[0003]

The above-mentioned prior art is PE
When applied to the T-apparatus, it is necessary to create a scattered ray image (S) by projection data composed of scattered ray components. When creating the scattered ray image (S), these projection data must be stored in a memory. Furthermore, reconstructing the scattered radiation image (S) takes about twice as long as the conventional image processing time. Therefore, there has been a demand for a PET device capable of performing these processes in real time (this was also true of the SPECT device).

[0004]

In order to achieve the above object, in the present invention, in a data collecting circuit for an emission CT apparatus which collects data in a histogram mode, a predetermined value is determined based on energy information of a measured gamma ray event. An energy discriminating circuit for discriminating whether or not the energy level is exceeded, and only when the energy discriminating circuit discriminates a gamma ray event exceeding a predetermined energy level, 1 is added to data of a predetermined address in the memory and recorded, and A memory circuit that subtracts 1 from the data of a predetermined address in the memory and records it only when the energy identification circuit identifies a gamma ray event that does not exceed a predetermined energy level, and switching between the 1 addition process and 1 subtraction process. And the frequency of occurrence of gamma ray events during the 1-subtraction process. And a frequency dividing circuit for giving an arbitrary weight in accordance with the frequency division, and the frequency dividing circuit arbitrarily divides a gamma ray event that does not exceed a predetermined energy level to generate an output signal of the frequency dividing circuit. The 1-subtraction process is executed only occasionally.

[0005]

In the present invention, when a scattered gamma ray event is identified, 1 is subtracted from the corresponding address data in the memory, a frequency divider circuit is inserted in the middle of the address data, and the scattered gamma ray is divided by the frequency division ratio corresponding to the scattered ray correction coefficient f. By dividing the event data, processing is achieved in real time.
Since the scattered radiation correction coefficient f is the same as that described in the section of the prior art, the image after the scattered radiation correction is performed in real time.
(For example, SPECT) also SPECT = P-
It is obtained by f × S. As a result, since it is not necessary to create an image by the scattered gamma ray event, it is possible to omit the memory for storing the projection data of the scattered ray component, and further it is not necessary to directly handle the projection data of the scattered ray component.
Data transfer time and scattered ray image reconstruction time are not required.

[0006]

DESCRIPTION OF THE PREFERRED EMBODIMENTS First, the principle of the present invention will be described.
FIG. 3 is a conceptual diagram showing the energy spectrum of incident gamma rays obtained by the detector of the PET device. In FIG. 3, the horizontal axis represents the energy E and the vertical axis represents the count number n. E _p is the energy of the photopeak of annihilation radiation and corresponds to 511 keV. Usually, the energy window W ₂ is in a range including the photopeak (here, E ₂ ~
E ₃ ) is set. On the other hand, the scattered ray window W ₁ is in the range below the Compton Valley (here, E
_{1 to} E ₂ ). At present, the following two types of coincidence counting events measured between the opposite detectors of the PET device are considered to be effective. A. Coincidence events between single events that occurred within the energy window W ₂ , ie all coincidence events (region S _p +
S ₂ ) B. Coincidence events between single events occurring within the energy window W ₁ , i.e. scatter coincidence events (region S ₁ ). As shown by Jezak et al., Between S ₁ and S ₂ S ₂ =
Assuming that there is a relationship of fS ₁ (f is a correction coefficient, f <1), since the true coincidence event t is equal to the region S _p , t = (S _p + S ₂ ) −fS ₁ = S _p + S _{2 -S 2 = S p ... (} 1) in obtained it can be seen.

The scattered radiation coincidence counting events occurring within the energy window W ₁ are divided by an appropriate value corresponding to the correction coefficient f to produce the same effect as the multiplication by the correction coefficient f (f <1). . For example, if f = 0.5, by dividing the frequency by 1/2, the total number of scattered ray coincidence counting events accumulated in the set data acquisition time becomes 1/2. Such processing is performed for all different addresses, but when the count value is large, it becomes equivalent to the one in which the weight of the correction coefficient f = 0.5 is uniformly added.

Normal measurement is performed in a histogram mode (a method of sequentially adding measured projection data to corresponding addresses in a memory). In this mode, when storing all coincidence events in memory, each time this event is measured, the value stored in the pixel of the memory corresponding to the address associated with that event is read out. 1 is added and the original pixel is written again.
On the other hand, in the subtraction of the scattered ray component, the numerical value in the pixel of the memory corresponding to the address of the divided scattered ray coincidence counting event is read out, and -1 is added (that is, 1 is subtracted), The original pixel is written again. Here, the value of the correction coefficient f is 0.5 for convenience, but
The accuracy of the correction can be further improved by applying the threshold value obtained by applying the method of Jesac et al. To the PET apparatus.

FIG. 1 shows a first embodiment of the present invention. FIG.
Is an example of a PET apparatus. In FIG. 1, reference numeral 1
Is the energy information signal of a single event of opposite detectors,
2 is the timing signal of the single event, 3 is the address signal of the single event, 4 is the energy information signal of the event is a pair of energy windows W ₁ , W ₁ , or W ₂ ,
It is an energy discriminating circuit for discriminating whether or not it is a pair of W ₂ . This energy discriminating circuit 4 can be usually composed of a ROM. Further, 5 is a coincidence counting circuit, 6 is an address circuit, and 7 is one of the outputs of the energy discriminating circuit 4, and it is determined whether the determination result corresponds to the total coincidence counting event or the scattered radiation coincidence counting event. A distinguishing signal, 8 is an enable signal, and when the coincidence counting event that has occurred is the energy level E ₁ or E ₂
Indicates that the A multiplexer (MPX) 9 directs whether to add 1 or subtract 1 when recording the generated coincidence counting event in the memory 20. 10 is a signal indicating that the coincidence counting event is confirmed, 1
1 is a signal when it is determined that the output signal 10 is based on a total coincidence counting event, 12 is a signal when it is determined that the output signal 10 is based on a scattered ray coincidence counting event, 13
Is a frequency dividing counter, 14 is a frequency-divided output signal, and 15 is 1
Adder circuit, 16 subtraction circuit by 1, 17 data bus, 18
Is a circuit for designating an address of the memory 20, and 19 is an address bus.

Next, the operation of the circuit shown in FIG. 1 will be described. Now, it is assumed that gamma rays are incident on the opposite detector.
The two opposing detectors input the single event signals 1, 2 and 3 to the energy discriminating circuit 4, the coincidence counting circuit 5 and the address circuit 6, respectively. Now, the two opposing detectors output a single event that has occurred within the energy window W ₂ , and as a result, their coincidence counts
Assuming that the coincidence count is detected in step S1, the energy discriminating circuit 4 instructs the multiplexer 9 to output the coincidence count confirmation signal 10 as the all coincidence count event signal 11. After that, the numerical value of the pixel in the memory 20 corresponding to the address created by the address circuit 6 is read out by the 1 addition circuit 15, is incremented by 1, and is written in the same pixel again.

Next, assuming that the two detectors facing each other output a single event occurring in the energy window W ₁ and the coincidence counts thereof are detected by the coincidence counter circuit 5 in the same manner, the energy discrimination circuit. 4 instructs the multiplexer 9 to output the coincidence counting confirmation signal 10 as the scattered radiation coincidence counting event signal 12. The output signal 12 is then divided into 1 / N by the frequency division counter 13. The frequency dividing value is appropriately determined according to the correction coefficient f described above. Frequency division counter 13
When there is an output signal 14 from the memory (note that the output signal 14 is not generated for each input because it is divided), the memory 2 corresponding to the address created by the address circuit 6 is generated.
The numerical value of the pixel within 0 is read out by the 1 subtraction circuit 16,
After being added by -1, the value is written again in the same pixel. In this way, the processing shown in the sequential equation (1) is executed in real time in response to the arrival of the incident gamma ray.

FIG. 2 is a schematic diagram for explaining the mechanism of the 1 addition circuit and 1 subtraction circuit. In the figure, 26 is a data read circuit, 27 is a data write circuit, and 24 is an up-down counter with preset consisting of n bits. In the case of the 1 addition process, the numerical value of the pixel in the memory 20 corresponding to the address designated by the address designating circuit 18 is output to the data reading circuit 26. This numerical value is preset in the preset counter 24.
The output signal 11 from the multiplexer 9 is connected to the up-count clock input U of the counter 24 with presetts, and +1 is added to the preset value according to the input of the output signal 11. Then, the output signal 25 of the preset counter 24 is written in the same pixel in the memory 20 via the data writing circuit 27. on the other hand,
In the case of 1 subtraction processing, the output signal 14 of the multiplexer 9
Is connected to the down counter clock input D of the preset counter 24, and -1 is added to the preset value according to the input of the output signal 14. The other procedures are the same as in the case of the 1 addition process.

FIG. 4 shows a second embodiment of the present invention. FIG.
Is an example of a SPECT apparatus. In FIG. 4, those having the same mechanism as those in FIG. 1 are designated by the same reference numerals. From the position calculation circuit (not shown) of the SPECT device, a signal indicating the incident position coordinates (X, Y) of the gamma ray incident on the detector, and a timing signal indicating the incident timing,
A signal representing energy information is output. In FIG. 4, 51 is a signal indicating energy information of incident gamma rays,
Reference numeral 52 is an incident timing signal of the incident gamma ray, and 53 is a signal representing the incident position coordinate (X, Y) of the incident gamma ray.
In the energy discriminating circuit 4, the energy information signal 51 is separated by the energy windows W ₁ and W _{2 in the same} manner as in the first embodiment, and is divided into an event having a normal gamma ray energy and an event caused by scattered rays. To be
Reference numeral 54 is a signal for controlling the multiplexer 9 based on the identification result, which corresponds to 7 in FIG. When the incident timing signal 52 is input to the multiplexer 9, it is controlled by the output signal 54 of the energy discriminating circuit 4 to be divided into a signal 11 corresponding to an event due to a normal gamma ray or a signal 12 targeted for an event due to a scattered ray. In the case of 1, it is input to the 1 addition circuit 15, and in the latter case, it is input to the frequency dividing counter 16. The incident timing signal 52 is further input to the address circuit 6 and synchronized with the incident position coordinate signal 53,
Designate the corresponding address of the memory 20. Other parts operate as described in the first embodiment.

[0014]

The following effects can be obtained by applying the present invention to an emission CT device. (1) The scattered radiation correction can be executed in real time. (2) The scattered ray data memory required for the scattered ray correction processing is unnecessary. (3) Since the memory shown in (2) is unnecessary, it is possible to save the time required for the data transfer related to the scattered ray correction processing and the scattered ray image reconstruction.

[Brief description of drawings]

FIG. 1 is a diagram showing a first embodiment of the present invention and shows a mechanism for executing a subtraction process of a scattered radiation component of a PET device in real time.

FIG. 2 is a conceptual diagram for explaining a mechanism of a 1 addition circuit and a 1 subtraction circuit in FIG.

FIG. 3 is a conceptual diagram showing an energy spectrum of incident gamma rays obtained by a detector of a PET device.

FIG. 4 is a diagram showing a second embodiment of the present invention, which is SPECT.
It shows a mechanism for executing the subtraction processing of the scattered radiation component of the apparatus in real time.

[Explanation of symbols]

 4 Judgment Circuit 9 Multiplexer 13 Frequency Division Counter 15 1 Addition Circuit 16 1 Subtraction Circuit 20 Memory 24 Preset Counter 26 Data Read Circuit 27 Data Write Circuit

Claims (1)

[Claims] 1. An emission identification device data collection circuit for collecting data in a histogram mode, an energy identification circuit for identifying whether or not a predetermined energy level or more is determined based on energy information of a measured gamma ray event, and the energy identification. Only when the circuit identifies a gamma ray event that exceeds a predetermined energy level, the data of a predetermined address in the memory is added and recorded, and the energy identification circuit identifies a gamma ray event that does not exceed the predetermined energy level. Memory circuit that subtracts 1 from the data at the predetermined address in the memory and records only when this is done, a multiplexer circuit that switches between the 1 addition process and the 1 subtraction process, and a gamma ray event during the 1 subtraction process And a frequency dividing circuit for giving an arbitrary weight according to the occurrence frequency of Data for an emission CT device characterized in that the frequency dividing circuit arbitrarily divides a gamma ray event that does not exceed a predetermined energy level, and 1 subtraction processing is executed only when an output signal of the frequency dividing circuit is generated. Collection circuit.