JP2626461B2 - Nuclear medicine imaging equipment - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a nuclear medicine imaging apparatus and, more particularly, to static imaging with a detector stationary and SPECT (single photon emission computer tomography) by rotating the detector.
The present invention relates to an improvement of a scintillation camera used for performing photographing or the like.

[0002]

2. Description of the Related Art In a nuclear medicine imaging apparatus, radiation emitted from a radioactive substance administered into a subject is detected outside the subject. Therefore, Compton scattering is caused by tissue in the subject before the radiation is emitted to the outside. And absorption, which need to be corrected. Various proposals have been made for this correction (for example, see Japanese Unexamined Patent Publication No.
218886, Japanese Patent Application No. 60-298164,
See Japanese Patent Application No. 4-280475).

[0003]

However, conventionally, there are problems that the quantitativeness is inferior and the amount of calculation processing is enormous. That is, in Japanese Patent Application Laid-Open No. 62-218886, since the energy weighting function is approximated as being constant regardless of the position in order to perform scattered radiation removal correction in real time, the quantitativeness is poor. In addition, Japanese Patent Application No. 60-298
No. 164 has excellent quantitative performance because the scattered radiation removal correction coefficient and the energy weighting function are changed using the information of the energy spectrum for each position (pixel).
Is limited to a matrix of about 64 × 64, and the matrix 512 × 512 has a large statistical error and a large amount of calculation processing as in static imaging, which is not practical.

The inventors of the present invention have made an invention to solve these problems, and the present invention is disclosed in Japanese Patent Application No. 4-28047.
We applied for No. 5, but there were still insufficient points. That is, there is a scattered radiation removal correction X, Y, Z
And a weighted scattered radiation removal correction process in which a weight function is applied only to the raw data of the three-dimensional image. However, in a normal clinic, only data with a small number of counts can be obtained, and in such correction processing of raw data itself, a calculation error is rather increased due to a statistical error, and as a result, a background area or a cold spot is increased. The disadvantage is that the image becomes noisy in some parts.

[0005] In view of the above, the present invention reduces the influence of scattered radiation and reduces the influence of scattered radiation even when the number of counts is small, such as clinical data, so as to reduce the influence of scattered radiation. It is an object of the present invention to provide a nuclear medicine imaging apparatus in which the contrast of an image is improved and the spatial resolution of an image is improved.

[0006]

In order to achieve the above object, in the nuclear medicine imaging apparatus according to the present invention, the address of the storage means is stored in the digital value of each of the two-dimensional position signal and the energy signal output from the radiation detection means. To collect three-dimensional image data,
A two-dimensional spatial filter process or a smoothing process is performed on the two-dimensional image data, and an energy weighting function stored in advance is used. For an energy range in which the value of the function is equal to or greater than a predetermined value, the energy process before the filter process or the smoothing process is performed. Applying the function to the data, and performing the energy-weighted scattered radiation removal correction for each pixel by applying the function to the data after the above filter processing or smoothing processing in the energy range where the value of the function is less than the predetermined value. Is the feature.

[0007]

The function of the energy weight function is compared with a predetermined value.
For a larger energy range, the weighting function is applied to the data before smoothing, and for an energy range smaller than the predetermined value, the weighting function is applied to the data after smoothing. Since the energy-weighted scattered radiation removal correction is performed on the data, the data after the smoothing is used for the data having a small number of counts, thereby preventing a calculation error due to a statistical error due to the small number of counts. As a result, errors in the scattered radiation part can be reduced. In addition, since the scattered ray portion is originally data with poor spatial resolution, there is no problem even if data after smoothing is used. Therefore, even when the number of counts is small as in clinical data, the noise in the background area and the cold spot is not increased, the influence of scattered radiation is reduced, the contrast of the cold spot is improved, and the spatial Resolution can be increased.

[0008]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS Hereinafter, a preferred embodiment of the present invention will be described in detail with reference to the drawings. As shown in FIG. 1, in the nuclear medicine imaging apparatus according to one embodiment of the present invention, two-dimensional position signals X and Y and an energy signal Z, which are output from the radiation detector 1 and indicate the incident position of radiation, are respectively provided. A / D converter and latch circuits 2, 3,
4 via the digital signal Dig. X, Dig. Y, Di
g. After being converted to Z, it is sent to the multiplexer 5.
The multiplexer 5 is provided in the Dig. X, Dig. Y, Di
g. Z is selected and sent to the image acquisition memory 6 as an address signal.

The detector 1 is a standby wave height analyzer (not shown).
Is sent to the timing circuit 9 to trigger the unblank signal obtained by the determination.
The timing circuit 9 generates appropriate timing signals a and b each time it is triggered, that is, for each event of radiation incidence, and outputs these signals to the A / D converter and latch circuit 2,
3, 4 and the image acquisition memory 6, respectively.

[0010] In the image acquisition memory 6, for every event that generates an unblank signal, the Dig. X, Dig. 2 specified by Y
Dimensional address or Dig. X, Dig. Y, Di
g. +1 is added at the three-dimensional address of Z. In this way, the image acquisition memory 6 acquires, for example, three-dimensional image data of a 64 × 64 × 64 matrix as shown in FIG. FIG. 2 shows a case where three-dimensional image data is collected. Here, Dig. Z '
Is given by Dig. This is the result of converting Z. Dig. Z ′ = m (Dig.Zn) Here, m is a constant representing a gain, and n is a constant representing an offset. The gain and offset are determined by the multiplexer 5 of the signal from the A / D converter and the latch circuit 4.
It can be changed at the input stage to. Also collect 3
The matrix size of the dimensional image data is set in advance by switching selection of the multiplexer 5.

As a result, the image acquisition memory 6 collects, for example, an energy spectrum as shown in FIG. 3A for the pixel (x1, y1) and an energy spectrum as shown in FIG. 3B for the pixel (x2, y2). A spectrum is collected. From this energy spectrum, the pixel (x2, y
2) has more scattered radiation than the pixel (x1, y1), that is, the scatterer is thicker.

The data collected by the memory 6 is transferred to an image processing memory 7, and the memory 7 and the image processing circuit 8 perform a two-dimensional spatial filtering process or a smoothing process and an energy-weighted scattered radiation removal correction process. Various image processing such as is performed. The energy weighting function used for the energy weighted scattered radiation removal correction processing is stored in the magnetic disk 11 in advance, and is loaded into the memory 7 and used for calculation. The multiplexer 5, the image collection memory 6, the image processing memory 7, the image processing circuit 8, and the magnetic disk 11 are controlled by the CPU 10.

The three-dimensional image data collected by the image collection memory 6 is transferred to an image processing memory 7, and the memory 7 and the image processing circuit 8 first use the following equation (1).
Undergoes two-dimensional spatial filtering or smoothing as represented by

(Equation 1) Here, C (x, y, z ′) is the energy Dig. For the pixel (x, y). Count number for each Z ', Csm
(X, y, z ′) represents the pixel (x,
y) energy Dig. The count number for each Z ′, S (x, y) is a two-dimensional smoothing function in the X and Y directions (for example, 3 × 3 9-point smoothing is performed).

Next, energy-weighted scattered radiation removal processing is performed in units of 64 × 64 pixel positions in the X and Y directions as shown by the following equation (2).

(Equation 2) Here, Cc (x, y) is the corrected count number for the pixel (x, y), and k (z ′, r) is the pixel (x, y).
R obtained from the energy spectrum information of
(X, y) (for example, the count ratio between the photoelectric peak region and the Compton scattering region), the energy Dig.
The weight function for Z ', p (k) is 1 (k (z', r)
≧ 0) or 0 (k (z ′, r) <0),
q (k) is 0 (when k (z ′, r) ≧ 0) or 1
(When k (z ′, r) <0).

The weighting function k (z ', r) is obtained in advance with a phantom or the like, and
1 is stored. Then, a parameter r is obtained from the actual energy spectrum information of each pixel, and a weight is obtained by substituting the parameter r. The calculation is performed by the image processing circuit 8 and the CPU 10. FIG. 4 shows an example of the weight function thus obtained. 4A is a weighting function k (z ′, r1) for a pixel (x1, y1) having an energy spectrum as shown in FIG.
Is a weight function k (z ′, r2) for a pixel (x2, y2) having an energy spectrum as shown in FIG. 3B. Parameters r1 and r2 of energy spectrum information
Are different, so they are different. If the parameter r is the same, the same weight function k (z ′, r) is obtained for different pixels.

As shown in the above equation 2, the weight function k
The value of (z ′, r) is compared with a predetermined value “0”, and the weight function is applied to data before smoothing for an energy range larger than 0, and By applying the weight function to the data after smoothing, the energy-weighted scattered radiation removal correction is performed for each pixel (x, y). That is, since the data in the energy range where the weighting function k (z ′, r) is smaller than 0, that is, the data of the scattered ray portion has a small number of counts and a large error, the data after smoothing is used. Can be. Also,
Since the scattered ray portion is originally data having poor spatial resolution, there is no problem even if data after smoothing is used. On the other hand, in the energy range where the weighting function k (z ′, r) is larger than 0, the count number is large, and it is considered that data independent of scattered radiation is obtained. Therefore, energy weighting is performed using data before smoothing. Perform scattered radiation removal correction.

As described above, it is possible to prevent a calculation error due to a statistical error from increasing in a cold spot portion or the like where the count is small, so that an image with much noise in the cold spot portion or the like can be prevented.

From the data in the image acquisition memory 6, the count number C20 (x, y) of the 20% window W without performing the scattered radiation removal correction can be obtained for each pixel by using the following equation (3). it can.

(Equation 3)

The present invention is not limited to the above embodiment, but can be variously modified without departing from the spirit of the present invention. For example, a matrix size other than the above can be used. Further, the above energy weight function k (z ′, r)
May be a function k (z ′) of only energy and may not be dependent on the spectrum for each pixel. Further, in the above description, “0” is used as the predetermined value for comparing the value of the energy weighting function k (z ′, r). However, a value other than “0” may be used, or the value may be divided into two or more stages. The smoothing function can be different for each. Further, a filtering process using a two-dimensional filter such as a Butterworth filter may be performed instead of the smoothing process.

[0020]

As described above, according to the nuclear medicine imaging apparatus of the present invention, noise in the background area and the cold spot portion is prevented from increasing even when the count number is small as in clinical data. The effect of scattered radiation can be reduced, the contrast at a cold spot or the like can be improved, and an image with high spatial resolution can be obtained. In addition, by changing the energy weighting function based on the energy spectrum information for each position, that is, for each pixel, and performing the war turbulence correction, the quantitativeness can be particularly improved. Further, the present invention is applicable not only to SPECT imaging but also to static imaging and whole body scanning imaging. It is also possible to obtain both an image after correction and an image before correction.

[Brief description of the drawings]

FIG. 1 is a block diagram of one embodiment of the present invention.

FIG. 2 is a conceptual diagram of an image acquisition memory.

FIG. 3 is a graph showing an example of an energy spectrum.

FIG. 4 is a graph showing an example of a weighting function.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 Detector 2, 3, 4 A / D converter and latch circuit 5, Multiplexer 6 Image acquisition memory 7 Image processing memory 8 Image processing circuit 9 Timing circuit 10 CPU 11 Magnetic disk

Claims (1)

(57) [Claims] 1. A radiation detecting means for generating a two-dimensional position signal representing the position of incident radiation and an energy signal representing energy, and three-dimensional by being addressed by respective digital values of the two-dimensional position signal and the energy signal. Storage means for collecting image data, means for performing two-dimensional spatial filtering or smoothing processing on the collected three-dimensional image data, and using a previously stored energy weighting function, the value of the function is used. The function is applied to the data before the filtering or smoothing processing for the energy range equal to or more than the predetermined value, and the function is applied to the data after the filtering or smoothing processing for the energy range where the value of the function is less than the predetermined value. Energy-weighted scattered radiation removal correction for each pixel. Nuclear medicine imaging apparatus comprising: