JP2000028731A - Nuclear medicine diagnostic apparatus - Google Patents

Abstract

PROBLEM TO BE SOLVED: To provide a nuclear medicine diagnostic apparatus, by which the density of an image can be made uniform even in the case of an RI diagnosis at a high counting rate. SOLUTION: In a nuclear medicine diagnostic apparatus, a detector 12 which detects a radiation radiated from a radioactive isotope (RI) dosed to a subject and which is composed of a plurality of channels is provided, a first position calculator 18 and a second position calculator 19 by which a radiation incident position and an energy value in the detector 12 are calculated on the basis of the detection output of the detector 12 are provided, and an acquisition memory 20 by which RI distribution data inside the subject is created on the basis of outputs of the first and second position calculators 18, 19 is provided. In this case, when a plurality of radiations are simultaneously or nearly simultaneously incident on the detector 12, incident positions of the respective radiations are calculated by the first and second position calculators 18, 19.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a nuclear medicine diagnostic apparatus for producing RI distribution data in a subject to which a radioisotope (RI) has been administered, and more particularly to a high counting rate RI.
The present invention relates to a nuclear medicine diagnostic apparatus suitable for diagnosis.

[0002]

2. Description of the Related Art A general structure of a scintillation camera is shown in FIG. 1 is a scintillator that generates scintillation light when radiation (gamma ray) emitted from a subject to which a radioisotope is administered enters, and 2 is a plurality of scintillators that are optically coupled to the scintillator 1 and convert the scintillation light into electric signals. Photomultiplier tube. The above-mentioned scintillator 1 and photomultiplier tube 2 constitute a detector 9 together with a collimator (not shown), a light guide and the like. The output signal from the detector 9 is supplied to the X, Y position calculation circuit 4 after passing through each preamplifier 3, where the position information of the scintillation light is calculated. That is, the position calculation circuit 4 is configured by, for example, a weighting circuit using a resistance matrix, and obtains information such as X ⁺ , X ⁻ , Y ⁺ , and Y ⁻ in X and Y orthogonal coordinates with the center of the scintillator 1 as the origin. . The output signal of the detector 9 is supplied to the peak analysis circuit 7 via the variable resistor 5 and the addition amplifier 6, where the peak value information is obtained. On the basis of the position information and the peak value information, the R
An I distribution image is displayed.

[0003]

With respect to the incident time interval of gamma rays emitted from the subject and incident on the detector 9, assuming that the average incident rate is r (count / sec), the incident time interval t (second) The probability density function f (t) of f (t) = r · e ^{−rt In} other words, when t = 0, f (t) takes the maximum value, so that the incidence time interval t is 0 and the probability of gamma ray incidence is the highest. Further, the probability density function f (t) becomes a sharp function as the average incident rate r increases.

As described above, particularly when the average incidence rate is high, the probability that a plurality of gamma rays enter the scintillator 1 at the same time or approach in time is increased. In this case, the incident positions of the incident gamma rays on the scintillator 1 are calculated by the position calculation circuit 4 constituted by a weighting circuit using a resistance matrix, and thus are determined as the middle of the plurality of incident positions, that is, the center position. Was. That is, the miscalculation was performed. Such erroneous calculations occur more frequently in the central portion of the detector 9, so that a high counting rate R
In the case of the I diagnosis, the central portion of the detector has a raised profile, and there is a disadvantage that the image density becomes non-uniform.

SUMMARY OF THE INVENTION The present invention has been made in view of the above-mentioned problems, and is free from erroneous calculations even when a plurality of radiations are incident on a detector at the same time or approaching in time. An object of the present invention is to provide a nuclear medicine diagnostic apparatus capable of equalizing the image density even in the case of (1).

[0006]

In order to achieve the above object, the present invention provides a radioisotope (R) administered to a subject.
I) based on a detection means comprising a plurality of channels for detecting radiation emitted from
A nuclear medicine diagnostic apparatus comprising: a calculating unit that calculates a radiation incident position and an energy value in the detecting unit; and a unit that creates RI distribution data in a subject based on an output of the calculating unit. Is characterized in that when a plurality of radiations are incident on the detecting means simultaneously or almost simultaneously, an incident position of each radiation is calculated.

According to the present invention, when a plurality of radiations are simultaneously or almost simultaneously incident on the detection means by the calculation means, the respective incident positions of the plurality of radiations are calculated, thereby enabling a high counting rate RI diagnosis. Even in such a case, the image density can be made uniform.

[0008]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS The configuration of an embodiment of the present invention will be described with reference to FIGS. In FIG. 1 showing the overall configuration of this embodiment, a detector 12 for detecting radiation (gamma rays) emitted from a subject 11 to which RI 10 has been administered is provided. The detector 12 includes a collimator 13 that passes only gamma rays in a predetermined direction, a scintillator 14 that converts gamma rays that have passed through the collimator into scintillation light, a light guide 15 that guides scintillation light generated by the scintillator, and a light guide 15 And a plurality of photomultiplier tubes 16 for receiving scintillation light from the scintillator 14. The plurality of photomultiplier tubes 16 are, for example, as shown in FIG.
One photomultiplier tube is arranged so that a hexagonal light receiving plane is formed.

Each output terminal of each photomultiplier tube 16 is connected via a preamplifier 17 to a first position calculator 18 for calculating a rough incident position of incident radiation, which will be described in detail later. In addition, the output signal X ₀ of the position calculator 18,
Y ₀ is supplied to a second position calculator 19 for calculating a more accurate incident position of the incident radiation, which will be described later in detail.

[0010] An acquisition memory 20 is provided, to which more accurate incident positions X, Y for the incident radiation output from the second position calculator 19 are supplied. This acquisition memory has a two-dimensional memory area designated by the X and Y position signals. When the second position calculator 19 receives, for example, the X1 and Y1 position signals, they are stored in the memory units X1 and Y1. Is added to the existing data. Such an operation is performed each time an output is supplied from the first position calculator 19, and after sufficient data collection, the RI distribution image data in the subject is stored in the collection memory 20. In order to use the RI distribution image data as a display image, a display memory 21 is provided, data is temporarily transferred to the display memory 21, and the transferred data is supplied to a display 23 via a D / A converter 22.

As shown in FIG. 3, the first position calculator 18 includes a specific coordinate position detecting circuit 24 and a group selecting circuit 2
5. An absolute coordinate system position calculation circuit 26. As shown in FIG. 2, the specific coordinate position detection circuit 24
With respect to the photomultiplier tubes 16 grouped into the axis and the N-axis, the signals are output as a specific coordinate position signal calculation based on the output signal of the photomultiplier tube 16 for each group. The group selection circuit 25 is a circuit that controls the selection of each group in the grouping of the photomultiplier tubes 16. The absolute coordinate system position calculating circuit 26 calculates the absolute coordinate system position signal X based on the specific coordinate position signal for each group output from the specific coordinate position detecting circuit 24 and the group selection signal output from the group selecting circuit 25. ₀ and Y ₀ are calculated and output.

As shown in FIG. 4, the specific coordinate position detecting circuit 24 has an independent circuit configuration for each group of photomultiplier tubes. Here, for example, as shown in FIG. 2, the photomultiplier tube 16 has an L axis (L1, L2,..., Ln) and an M axis (M
1, M2..., Mn), 3 of the N axis (N1, N2.
Divided into two groups. Since each of the groups has the same circuit configuration, only the L-axis group will be described. Each output of the photomultiplier is connected to the OR circuit 27,
This output terminal is connected to the A / D converter 28. As a result, the A / D-converted output signal is output as a specific coordinate position signal by the comparator 29 and the encoder 30.

The second position calculator 19 is configured as shown in FIG. That is, a decoding circuit 31 is provided to which the approximate incident position signals X ₀ and Y _{0 of the} incident radiation input from the first position calculator 18 are supplied. The decoding circuit receives the signals X ₀ and Y _0, and sets a total of seven photomultiplier tubes 16 around the photomultiplier tube 16 located at the positions X ₀ and Y ₀ in a hexagonal shape. A signal for selecting a group is supplied to the seven analog switches 32.

FIG. 6 shows how such photomultiplier tubes 16 are grouped. For example, when the photomultiplier tube of the serial number 4 is selected by the first position calculator, the group G1 of the photomultiplier tubes of the serial numbers 1, 2, 6, 7, 8, 13, and 14 surrounding it is selected. Be elected. Next,
When the serial number 5 is selected, the group G2 of the serial numbers 2, 3, 7, 8, 9, 14, and 15 surrounding it is selected. Others are selected in the same manner as above.
An arithmetic unit 33 is provided for receiving each output of the seven analog switches 32, that is, each output of the photomultiplier tube 16 selected by the decode circuit 31. This arithmetic unit calculates more accurate radiation incident position signals X and Y based on all the received signals, and calculates the energy value of the incident radiation from the total value of all the received signals, and this value is within the set range. When
It is output as incident position signals X and Y of the incident radiation.

Next, the operation of the embodiment having the above configuration will be described. In FIG. 1, a subject 11 who has been administered RI10
Gamma rays emitted from the detector 12 are detected by the detector 12. At this time, an electric signal is output from the output terminal of the photomultiplier tube 16 of the detector 12 according to the incident position of the gamma ray. The peak value of these output electric signals becomes higher as the photomultiplier tube 16 is disposed closer to the gamma ray incident position. The output electric signal from each photomultiplier tube 16 is amplified by each preamplifier 17 and then supplied to a first position calculator 18.

The first position calculator 18 takes in the specific coordinate position detecting circuit 24 shown in FIG. These signals are collected by an OR circuit 27 and converted into digital signals by an A / D converter 28. These digital signals are sorted by the magnitude of the signal value by the comparator 29, and finally the L axis, the M axis, and the N
It is output as a specific coordinate position signal for each axis group. Which group the output signal belongs to is identified by a signal from the group selection circuit 25. The identification signal and the specific coordinate position signal are supplied to the absolute coordinate system position calculating circuit 26, which outputs the absolute coordinate position signals X ₀ and Y _{0 of the} incident gamma ray.

The signals X ₀ and Y ₀ output from the first position calculator 18 in this manner are supplied to a second position calculator 19. In the second position calculator, the decoder circuit 3
1 receives the signals X ₀ and Y _0, and receives the coordinate positions X ₀ and Y ₀ .
A group of a total of seven photomultiplier tubes surrounding the photomultiplier tube existing in ₀ and surrounding it in a hexagonal shape,
Each analog switch 32 is selected by operation control. The total output of the photomultiplier tube 16 selected in this way is supplied to the calculation unit 23, where the more accurate incident positions X and Y of the incident gamma ray are calculated, and the energy value of the incident gamma ray is calculated. When the calculated energy value falls within the set range, the more accurate incident position signals X and Y of the incident gamma ray are output.

The output signals X and Y are supplied to the acquisition memory 20, and one count data is added to the memory data at the memory position designated by X and Y. By repeating such an operation, the R in the subject is stored in the acquisition memory 20.
I distribution data is collected and stored. The acquired data is temporarily transferred to the display memory 21 and converted into an analog signal by the D / A converter 22, and then the RI distribution image in the subject is displayed on the display 23.

In addition, according to the apparatus having the above-described configuration, not only can the incident position of the radiation (gamma ray) incident one by one at a time interval be calculated, but also the photomultiplier tubes 16 are divided into a plurality of groups. Then, the output signals are processed, so that it is possible to calculate the respective incident positions of a plurality of gamma rays incident at the same time.

It should be noted that the present invention is not limited to the above-described embodiment, and that the design can be changed as appropriate without changing the gist of the present invention. For example, as shown in FIG. 7, the photomultiplier tubes may be grouped into nine groups surrounding the central photomultiplier tube in a diamond shape. That is, for example, the serial numbers 1, 2, 3, 6, 8, 12, 1 around the photomultiplier tube of serial number 7
A group g1 of groups 3 and 14, or a group g2 of serial numbers 2, 3, 4, 7, 9, 13, 14, and 15 centered on the serial number 8.

[0021]

As described above, according to the present invention, the position coordinates of each of a plurality of simultaneously incident radiations can be accurately calculated, so that even in the case of RI diagnosis with a high counting rate, the image density can be reduced. It can be made uniform.

[Brief description of the drawings]

FIG. 1 is a block diagram showing the configuration of an embodiment of the present invention.

FIG. 2 is an explanatory diagram showing a mode of grouping the photomultiplier tubes in the embodiment.

FIG. 3 is a block diagram showing a configuration of a first position calculator in the apparatus shown in FIG.

FIG. 4 is a block diagram showing a configuration of a specific coordinate position detection circuit in the first position calculator shown in FIG.

FIG. 5 is a block diagram showing a configuration of a second position calculator in the device shown in FIG.

FIG. 6 is an explanatory diagram showing grouping of photomultiplier tubes.

FIG. 7 is an explanatory diagram showing grouping of photomultiplier tubes.

FIG. 8 is a block diagram showing a configuration of a conventional device.

[Explanation of symbols]

 Reference Signs List 12 detector 16 photomultiplier tube 18 first position calculator 19 second position calculator 20 collection memory 21 display memory 22 D / A converter 23 display 24 specific coordinate position detection circuit 25 group selection circuit 26 absolute coordinates System position detection circuit 31 Decoding circuit 32 Analog switch 33 Operation unit

Claims (1)

[Claims] 1. A detecting means comprising a plurality of channels for detecting radiation emitted from a radioisotope (RI) administered into a subject, and a radiation incident position on the detecting means based on a detection output of the detecting means. And a calculating means for calculating the energy value, and a means for creating RI distribution data in the subject based on the output of the calculating means. A nuclear medicine diagnostic apparatus characterized in that when radiations of the same type enter simultaneously or almost simultaneously, the incident positions of the respective radiations are calculated.