JPH0232893B2 - - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION The present invention relates to a radiological diagnostic apparatus capable of obtaining diagnostic images of a subject with excellent contrast resolution.

X-ray diagnostic equipment is often used for transmission diagnosis of a subject. The diagnostic image of the subject projected onto the X-ray film by this X-ray diagnostic device generally has a very high spatial resolution of 6 to 10 lines/mm, that is, excellent resolution, but its contrast resolution is relatively poor. It has the property of That is, the contrast resolution of X-ray film is as low as about 6 bits, and for this reason, it could not be expected to be sufficiently effective in diagnosing internal tissues such as the lungs, liver, bile, pancreas, and spleen.

For example, if we assume a specimen model of a uniform absorber with thickness x and X-ray absorption coefficient μ as shown in Figure 1, the number of photons incident on the specimen Io and the number of transmitted photons I
There is the following relationship between:

I=Io・exp(−μx) ………(1) Also, the absorption coefficient (μ+Δμ), the number of photons transmitted through the object (I+ΔI), is I+ΔI=Io・exp {−(μ+Δμ)x}
......(2) is shown as. Therefore, from these relationships, the degree of change in the number of transmitted photons can be expressed as (I+ΔI)/I=exp(-Δμx), and approximately as ΔI/I≒−Δμ・x (3) Can be done. Therefore, with the aforementioned X-ray film having a contrast resolution of about 6 bits, changes up to about .DELTA.I/I=0.05 can be identified. This also applies to item (3) above.
From the relationship shown in the equation, it can be seen that in order to ensure contrast resolution, the following condition should be satisfied: Δμ・≧0.05 (4). However, X
The linear absorption coefficient of a subject subjected to radiological diagnosis is generally
It is about 0.2, and for this reason, the thickness of the specimen x is 1 cm.
In this case, the ratio of Δμ to μ must be 25% or more different. Further, when the above-mentioned thickness x is 10 cm, a ratio of 2.5% or more must be ensured.

However, the linear absorption coefficients of the liver, pancreas, and spleen are approximately μ = 0.214, 0.211, and 0.204, respectively, and Δμ has only a 1% difference in linear absorption coefficient, so the X-ray film described above cannot absorb such minute rays. There was a problem that the difference in absorption coefficient could not be identified.

On the other hand, in recent years, attempts have been made to image-process X-ray detection signals to obtain a desired diagnostic image of a subject without using such an X-ray film. That is, while scanning the radiation beam in a fan shape, the subject is moved in a direction perpendicular to the fan-shaped scanning plane, the radiation beam passing through the subject in a certain region is detected by a radiation detector, and the detection output is image-processed to detect the subject. This is to obtain a specimen diagnostic image. However, in order to ensure the spatial resolution of the diagnostic image of the subject obtained in this way, it is sufficient to provide the detection elements constituting the radiation detector at a high density, but there is a limit to the density increase;
The influence of interference between adjacent detection elements and scattering of irradiated radiation in the object is large, and as a result, there is a problem in that it is difficult to obtain a detected image signal with high resolution.

The present invention was made in consideration of these circumstances, and its purpose is to provide a highly practical radiological diagnostic apparatus that can obtain diagnostic images of a subject using radiation with excellent spatial resolution and contrast resolution. It is about providing.

The outline of the present invention is to irradiate a subject with a radiation beam in a fan-shaped scan manner, and detect the transmitted radiation beams with a radiation detector array formed by arranging a plurality of radiation detection elements. In a radiological diagnostic apparatus that moves a scanning plane of a radiation beam in a direction perpendicular to the scanning plane and obtains a subject diagnostic image signal from the detection output of each of the radiation detection elements, the detection output of each of the radiation detection elements is weighted. By then combining the images to remove interference between detection elements and the effects of radiation scattering from the object, a diagnostic image of the object with excellent spatial resolution and contrast resolution can be obtained, effectively achieving the above objectives. This is what I did.

Hereinafter, one embodiment of the present invention will be described with reference to the drawings.

FIG. 2 is a schematic configuration diagram of a radiation (X-ray) diagnostic apparatus, and 1 is a bed (table) on which a subject 2 lies. The table 1 is moved in the direction of arrow A in the figure under the control of the table controller 3, and the subject 2 is subjected to radiological diagnosis.

Therefore, the X provided above the table 1
The ray generator 4 is controlled by an X-ray controller 5, and outputs a radiation (X-ray) beam in a fan-shaped manner across the subject 2. The fan-shaped scanning plane of this X-ray beam is orthogonal to the moving direction of the table 1, and therefore, the subject 2 is scanned in a fan-shaped manner by the X-ray beam and sequentially moved in the orthogonal direction through the scanning area. . It goes without saying that the X-ray beam is collimated and then scanned in a fan shape.

On the other hand, at a position below the table 1 and facing the X-ray generator 4, an X-ray detector 6 is provided to detect the fan-shaped X-ray beam. This X-ray detector 6 has an array structure of detection elements in which a plurality of X-ray detection elements made of silicon single crystal or Zenon gas cells are arranged to detect each fan-shaped scanning position (scanning angle) of the X-ray beam. It is what you do. The X-ray detection signals detected by each of these X-ray detection elements are guided to the data acquisition circuit 7, sampled and inputted, converted into a desired signal form, and output.

It is composed mainly of a signal processing section 10 and a CPU 11, and the table controller 3 and the X-ray controller 5 are operated under timing control by this CPU 11. Further, the X-ray detection signal extracted via the data acquisition circuit 7 is guided to the signal correction processing circuit 12 of the signal processing section 10 and subjected to correction processing, and then stored in the image memory 13. The above-mentioned signal correction processing circuit 12 is a feature of this apparatus, and as described later, it corrects detection output fluctuation components due to sensitivity deviation, interference amount, scattering, etc. due to the conditions of the detection element group and the subject 2, etc. It is. After undergoing this correction processing, the detection data stored in the image memory 13 is read out to the CPU 11, and is also supplied to the image processing device 14, where it is subjected to predetermined image processing, and then stored in the image memory 13 again. Stored. Then, the image is read out from the image memory 13 and displayed on the display 15. Incidentally, the disk memory 16 functions as an auxiliary storage device that semi-permanently stores and stores the image-processed X-ray diagnostic images as described above as an image file. Further, the control/display console 17 is used to input a series of these operation control command information.

According to the apparatus configured in this way, the X-ray generator 4 is driven under the control of the CPU 11, and the X-ray beam is scanned and irradiated onto the subject 2 on the table 1 in a fan shape. Ru. This fan-shaped scanning of the X-ray beam is performed only while the fan-shaped scanning plane intersects the region of interest of the subject 2, and after this fan-shaped scanning is performed, the table 1 is moved and the X-ray beam is applied to the next region of interest. A fan-shaped scan will be performed. Note that the X-ray beam irradiated at this time may be pulsed or continuous.

By the way, the spatial resolution of the detection signal in the region of interest of the subject 2 obtained by this fan-shaped scanning of the X-ray beam is determined by the unit cell length of the plurality of arranged X-ray detection elements constituting the X-ray detector 6. formed by. Further, the spatial resolution in the moving direction of the table 1 perpendicular to the arrangement direction of the X-ray detection elements is determined by the moving speed of the table 1 and the fan-shaped scanning period of the X-ray beam. Therefore, if these conditions such as the unit cell length are appropriately determined according to the specifications and the table 1 is moved at a predetermined speed, a diagnostic image of the subject can be obtained that has the same spatial resolution as when using conventional X-ray film. (X-ray transmission image) can be obtained.

The data acquisition circuit 7 samples the X-ray detection output at such a spatial resolution, converts it into digital data, and outputs it in a quantified manner.

Now, what is found here to quantify the transmission image of the object 2 corresponds to the term (μx) in the above equation (1), and therefore, μx=ln(Io/I ) ......(1)' The logarithm value of the ratio between the number of input photons Io and the number of output photons I can be found. Specifically, the ratio between the intensity of the X-ray beam irradiated to the subject 2 and the intensity of the transmitted X-ray beam may be determined, and the logarithm value thereof may be determined. Then, by A/D converting this logarithm value, data quantifying the transmission image of the subject 2 is obtained.

By the way, in order to obtain a clear X-ray diagnostic image with improved spatial resolution and concentration resolution, it is necessary to shorten the unit cell length of the detection element and implement high-density packaging as described above. However, when such high-density packaging is used, mutual interference between adjacent detection elements becomes a problem.

For example, as shown in the model of the X-ray detector 6 in FIG. 3, a plurality of detection elements 6 _o-2 arranged in a straight line,
Assume that X-rays having a value of "100" are incident on each of 6 _o-1 , 6 _o , 6 _o+1 , and 6 _o+2 . At this time, if the detection element _6o receives interference of 5% and 10% from its adjacent detection elements 6o _-1 and 6o ₊₁ , respectively,
The value of the detection output is "115". In addition, the detection elements 6 _o-1 and 6 _o+1 will obtain detection outputs of values "95" and "90", and the interference will act as the number of photons I and have a linear adverse effect. Become.

The signal correction processing unit 12 uses multipliers 21a, 21b, 21c and an adder 22 as shown in FIG. After weighting an appropriate amount of coefficients (weighting), find the total sum,
The device is configured to correct for said interference. In this example, the output of element 6 _o is multiplied by a coefficient "1" and
_Coefficients “−0.05” and “ ₋
0.1" and summation of these output values yields a detection output corresponding to the element 6 _o with a value of "101.3".

Further, for the adjacent element 6 _o-1 , weighting is performed using coefficients ``0'', ``1'', and ``0.05'', and then the sum is calculated to obtain a corrected output with a value of ``100.8''. Therefore, according to this example, the output of the detection element _6o , which includes an error of 15%, is corrected to reduce the error to 1.3.
%, and the output of the detection element 6o _-1 containing an error of 5% can be corrected to a signal with an error of 0.8%.

In addition to such interference, there are also the effects of scattered radiation in the object 2, sensitivity deviations specific to the detection element, etc.
For variation factors that act linearly, the error can be suppressed by similar correction processing. However, in this case, it goes without saying that the handled signal has linearity, and this linearity has an important meaning in the correction process.

FIG. 5 is a configuration diagram showing an example of the signal correction processing circuit 12 that performs such correction processing. That is,
Each detection element of the X-ray detector 6 ~ _{6 o-2} , 6 _o-1 , 6 _o ,
Multipliers (coefficient units ₎ ~21 _{o-2 ,}
21 _o-1 , 21 _o , 21 _o+1 , 21 _o+2 ~ are respectively provided, and these multipliers ~ 21 _o-2 , 2
Each output of 1 _o-1 , 21 _o , 21 _o+1 , 21 _{o+2 ~} is configured to be input to an adder 22 . The output of the adder 22 is taken out as a corrected output signal, and is led to a comparator 23, where it is compared with a reference value and its error is determined. A coefficient setter 24 to which the error signal obtained by the comparator 23 is inputted is connected to the multipliers ˜21 _o-2 , 21 _o- according to a predetermined algorithm.
₁ , 21 _o , 21 _o+1 , 21 _o+2 - coefficient values ~ _{a o-2} , a _o-1 , a _o , a _o+1 , a _o+2 ~ are determined. Furthermore, these coefficient values ~ _{a o-2} , a _o-1 ,
a _o , a _o+1 , a _o+2 ~ are determined for linear errors such as the amount of interference to be corrected. For example, when X-rays at a reference level are emitted to the X-ray detector 6 It is determined through a calibration process.

Therefore, according to such a signal correction processing unit 12, when a detection element to be corrected is specified, the correction coefficients ~ _{a o-2} , a _o-1 , a _o ,
a _o+1 , a _o+2 ~ are determined, and the corrected output Σa _i I _i is determined.

In this way, a detection signal with suppressed errors is obtained through the correction process, and it is possible to ensure a sufficiently high contrast resolution.
Then, image processing is performed using the radiation detection data with this high contrast resolution and high spatial resolution, and the image is displayed on the display 15.

By the way, as mentioned above, the data processing circuit 7 logarithmically transforms the detection output and converts it into a digital signal by A/D converting it, so the image processing device 1
When performing image processing such as reduction/enlargement using 4 together,
Preferably, the digital signal is subjected to anti-logarithmic transformation. For example, if logarithmic transformation is performed according to the relationship shown in FIG. 6a, anti-logarithmic transformation may be performed according to the relationship shown in FIG. 6b. Specifically, logarithmic transformation and anti-logarithmic transformation can be easily performed digitally by realizing the approximate characteristic curve as a digital polygonal line and performing linear interpolation on the data between them. In order to output the radiological diagnostic image stored in the image memory 13 after undergoing correction processing and image processing, the signal dynamic range is reduced, and in order to further simplify the subsequent image processing operation, the above-mentioned signal It is advantageous if the output is further digitally logarithmically transformed.

In the above embodiment, the correction process was performed before the signal was input to the image memory 13, but after the detection output is stored in the image memory 13, the correction process is performed on the stored detection signal. You may also do so. In addition, this correction processing is performed by CPU1
1 or the image processing device 14, etc., by software processing. Further, in the embodiment, the detection of a diagnostic image using X-rays has been described, but it is also possible to use gamma rays, and furthermore, the concept can also be used for ultrasonic diagnosis.
In short, the present invention can be implemented with various modifications without departing from the gist thereof.

As described in detail above, according to the present invention, the linear error in the detection output detected by the radiation detector is corrected by weighting processing and signal summation processing to reduce the amount of error. The sensitivity characteristics of the detector, interference between detection elements, and adverse effects such as radiation scattering on the subject can be linearly and extremely effectively removed. Therefore, the contrast resolution of the detection signal can be made sufficiently high, and in combination with the spatial resolution, an extremely good radiological diagnostic image with a remarkable diagnostic effect can be obtained. On top of that,
This has tremendous effects such as enabling high-speed digital diagnostic image processing.

[Brief explanation of drawings]

Fig. 1 is a diagram showing the principle of radiation detection, Fig. 2 is a schematic configuration diagram showing an embodiment of the present invention, Fig. 3 is a diagram showing a model of detection signal output, and Fig. 4 is the concept of detection signal correction. FIG. 5 is a block diagram showing an example of a signal correction processing section, and FIGS. 6a and 6b are diagrams showing log/anti-log conversion characteristics. 1... Bed (table) 2... Subject, 3...
Table controller, 4...X-ray generator, 5...
...X-ray controller, 6... X-ray detector, 7...
Data acquisition circuit, 10...Signal processing section, 11...
CPU, 12... Signal correction processing unit, 13... Image memory, 14... Image processing device, 15... Display, 16... Console, 21a, 21
b, 21c, 21 _o-2 , 21 _o-1 , 21 _o , 21 _o+1 ,
21 _o+2 ... Multiplier, 22... Adder, 23...
Comparator, 24...coefficient setter.

Claims (1)

[Scope of Claims] 1. A radiation source that outputs a radiation beam in a fan-shaped scan, and a plurality of radiation detection elements for detecting the amount of radiation transmitted through a subject by the radiation beam in the direction of the fan-shaped scan of the radiation beam. a detector array arranged in the detector array; and means for respectively correcting and outputting the radiation detection outputs from the radiation detection elements of the detector array, and correcting the radiation detection outputs from the radiation detection elements respectively. The outputting means adds radiation detection outputs from other radiation detection elements obtained by multiplying the radiation detection outputs obtained from the radiation detection elements facing the direction of the fan-shaped radiation beam by respective predetermined weighting coefficients. A radiological diagnostic apparatus comprising means for determining the weighting coefficients according to the amount of radiation leakage between the radiation detecting element and the other radiation detecting elements facing each other in the direction of the radiation beam.