JP3353499B2 - X-ray imaging device - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to an X-ray imaging apparatus for obtaining an X-ray image of a subject, and more particularly to an X-ray diagnostic apparatus for correcting an X-ray intensity distribution and providing an X-ray image having excellent density uniformity. .

[0002]

2. Description of the Related Art An X-ray imaging apparatus that irradiates a subject with X-rays and directly irradiates transmitted X-rays to an X-ray film to obtain an X-ray image has been used for many years. Digital radiography system called “radiology” has been developed.

FIG. 3 is a schematic diagram showing an example of such a digital X-ray image system. X-rays emitted from an X-ray tube apparatus 11 and transmitted through a subject 12 are detected by a detection unit 13 in which detection elements are arranged in a matrix. , And the detection data for each element is stored in the storage element matrix 15 as digital data. The detection data for each element once stored is read out to the display circuit 17 and displayed on the display device 18 as an X-ray image.

Further, an X-ray dose is once stored in a film using a fluorescent photostimulable film, and then read out as a two-dimensional matrix data by a laser beam with a digital signal, or an image of an image intensifier is read by a CCD element. Using a readout or a semiconductor sensor, read out an ionization signal due to X-rays in the semiconductor,
There are also digital X-ray imaging systems configured to individually count ray photons.

FIG. 4A is an overall schematic view of an X-ray tube apparatus 11 used in such an X-ray imaging apparatus. In the figure,
The anode 11d and the cathode 11e are covered with a vacuum valve 11b made of glass or metal and maintained in a vacuum state. Since the vacuum valve 11b is immersed in the insulating oil 11c in order to absorb the heat generated in the anode 11d, the X-rays generated from the anode 11d are attenuated by glass or oil interposed therebetween, and the difference in the intensity is reduced. Occurs.

[0006] Also, as shown in FIG.
Thermal electrons e emitted from the cathode 11e ^- but is being generated by striking the anode 11d, the intensity distribution of X-rays, as shown in FIG. 4c, the anode of the electron flow, commonly referred to as a real focal point It depends on the above distribution and the apparent focal point for extracting it, and is determined by the shape of the anode 11e and the direction in which X-rays are extracted. Therefore, in FIG. 4B, the X-ray intensity shows a distribution that decreases in the θ direction, and no X-ray is generated at θ (θ> φ) larger than the angle of the target shown in φ in FIG.

In such a case, the difference in the intensity of the X-rays appears as shading of the picked-up image. In particular, the image at the end of the image is very low in density compared to the center and has a poor contrast ratio, so that an accurate diagnosis cannot be made. For this reason, in the conventional X-ray imaging apparatus,
By setting a sufficient distance between the X-ray tube device 11 and the detection unit 13, the influence on the captured image due to the difference in the intensity distribution of the X-rays has been reduced.

[0008]

However, the anode 1
The ratio of the energy of the electrons colliding with 1d that is converted to X-rays is very small, about 1%, and the generated X-rays are diffused in all directions, thus contributing to actually obtaining an X-ray image. The ratio of X-rays to be emitted becomes even smaller. For this reason, when the distance between the X-ray tube apparatus 11 and the detection unit 13 is sufficiently large, it is necessary to generate a considerably strong X-ray. In such a case, however, most of the input power is converted to heat. In order to perform imaging, a function capable of sufficiently cooling the X-ray tube device is required, and there has been a problem that the entire device becomes large.

[0009] When an X-ray tube device having a small heat capacity is used,
Since the cooling time becomes extremely long and a great deal of time is required for imaging, an X-ray tube device having a relatively large heat capacity has to be used for performing rapid imaging.

The present invention has been made in order to solve the above-mentioned problem, and is intended to provide a low-density X-ray tube even when a small and small-heat-capacity X-ray tube apparatus is used. It is an object of the present invention to provide an X-ray imaging apparatus that can obtain an image and can quickly perform X-ray imaging.

[0011]

An X-ray imaging apparatus according to the present invention comprises an X-ray tube device for generating X-rays and a plurality of detecting elements, and irradiates the irradiated X-rays as data for each element. X-ray detection means for outputting, a storage element matrix for storing the data of each of the X-rays irradiated in a state where no subject exists, and a storage element for storing the data of each of the X-rays transmitted through the subject. A correction means for normalizing with the corresponding data stored in the matrix, and a display means for displaying the normalized data as an X-ray image are provided.

[0012]

The X-rays generated from the X-ray tube apparatus 1 and transmitted through the subject 2 are detected by the detecting section 3 composed of a large number of detecting elements, and are converted into digital data for each element by the second storage element matrix 6. Is stored. The correction means 9 irradiates X-rays in a state where the subject 2 is not present in advance, reads out corresponding data from the first storage element matrix storing the detection data for each element, and reads the second storage element matrix. 6 is normalized. Then, when the normalization of all the stored data is completed, the display circuit 7 reads out the normalized data of the second storage element matrix 6 based on the instruction of the correction means 9, and displays it on the display device 8 as an X-ray image. Display.

[0013]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS An embodiment of the present invention will be described with reference to FIGS. FIG. 1 is a schematic diagram of an X-ray imaging apparatus according to the present invention, and 1 is an X-ray tube apparatus having the configuration of FIG.
Generates X-rays. Reference numeral 2 denotes a subject to be imaged by X-rays.

Reference numeral 3 denotes a detection unit in which a large number of detection elements are arranged two-dimensionally, and is configured to output X-ray detection data as a digital signal 4 for each element.

Reference numeral 5 denotes a first storage element matrix which stores detection data obtained for each detection element of the detection section 3 when X-rays have been irradiated in a state where the subject 2 is not present. Reference numeral 6 denotes a second storage element matrix,
Is temporarily stored.

Reference numeral 7 denotes a display circuit which reads out the X-ray transmission data of the subject 2 stored in the second storage element matrix 6 and displays an X-ray image of the subject 2 on a display device 8 constituted by a CRT or the like.

Numeral 9 denotes a correction means for normalizing the X-ray transmission data of the subject 2 stored in the second storage element matrix 6 using the data stored in the first storage element matrix, and a display circuit 7. To read the normalized data.

Next, the operation of the present invention will be described with reference to the flowchart of FIG. First,
Before performing X-ray imaging of the subject 2, X-rays are irradiated in a state where the subject 2 does not exist, and the detection data A (i, j) obtained for each element of the detection unit 3 is stored in a first storage element. It is stored in the matrix 5. At this time, the correction means 9 specifies and stores the maximum value MAX {A} from the detection data A (i, j). Then, X-ray imaging of the subject 2 is performed, and the imaging data B (i, j) for each element detected by the detection unit 3 is stored in the second storage element matrix 5.

In this state, the correcting means 9 stores one image data B (i, j) for each element stored in the second storage element matrix 6 and the image data B (i, j) stored in the first storage element matrix 5. Detection data A (i,
j) is read (S1, S2).

Then, by performing the following arithmetic processing, the imaging data B (i, j) is converted to the detection data A (i, j).
(S3).

C (i, j) = B (i, j) × Max {A} / A (i, j) When the arithmetic processing is completed, the correcting means 9 converts the normalized data C (i, j) The data is stored in the corresponding storage element of the second storage element matrix 6 (S4).

Then, the above-mentioned operations S1 to S4 are repeated (S5), and when the processing is completed for all the matrices, the correcting means 9 gives a display instruction of the normalized data to the display circuit 7 and ends the processing. (S6). Display circuit 7
Receives the instruction, sequentially reads the normalized data C (i, j) from the second storage element matrix 6 and causes the display device 8 to display an X-ray image.

In the above-described embodiment, the X-ray image is displayed after all the storage data A (i, j) of the second storage element matrix 6 are normalized. (I, j) and its normalized data C (i, j)
Since j) is independent for each element, the normalized data C (i, j) may be directly output to the display circuit 7 without rewriting to the second storage element matrix 6. . Also, in advance, Max {A} / A (i, j)
Is stored in the first storage element matrix, the processing time for normalization is reduced.

Since the detection data A (i, j) stored in the first storage element matrix 5 is data captured in a state where the subject 2 does not exist, the adjacent data A (i, j) It is considered that no large difference occurs between the values of A (i + 1, j) and A (i + 1, j). Therefore, a plurality of elements, for example, B
(I, j), B (i + 1, j), B (i, j + 1), B
A (2 + 1) element of (i + 1, j + 1) is converted into one data,
For example, corresponding elements A (i, j), A (i + 1, j),
A (i, j + 1) or A (i + 1, j + 1), or the average value of them, allows the first storage element matrix 5 to have a memory size of 4
Can be reduced by a factor of one.

Furthermore, if the difference between the detection data A (i, j) is stored in the first storage element matrix 5, the number of bits per element can be reduced, so that the memory capacity is reduced. .

As described above, according to the present invention, the distance between the X-ray tube device and the detection unit is made sufficiently short so that even if the intensity of the irradiated X-rays varies, an accurate X-ray image without shading is obtained.
A line image is obtained. As a result, the X-ray utilization can be improved, and the amount of heat generated in the X-ray tube device can be reduced accordingly.
Even when an X-ray tube device having a small heat capacity is used, the cooling time is shortened, and rapid X-ray imaging is possible.

The detecting section 3 here uses, for example, a film with a fluorescent stimulable film, that is, the incident X-ray.
A detector which forms a film with a substance which is in a metastable state according to the dose and reads it out by two-dimensional scanning with laser light, and which may take out a signal with an optical fiber or the like provided separately. However, since the signal to be output is an analog signal, each pixel is digitized by synchronizing the scanning cycle of the laser beam with the A / D converter through an A / D converter.

An X-ray image intensifier can also be used as a detector. The X-ray image intensifier is a detector that uses a scintillator that converts X-rays into light, converts this light into electrons using a photoelectric film, and applies an accelerating voltage to increase the energy of the electrons and make the output phosphor screen emit light strongly. is there. A digital image can be obtained by capturing the output phosphor screen with an image pickup tube or a CCD element and subjecting the signal to A / D conversion.

Further, the electrons generated in the photoelectric film are transferred to a TFT.
A method is also conceivable in which active elements such as elements, MIM elements, and MIS elements are provided in a matrix and the charge amount is read.

Further, a detector which reads out an ionization signal ionized by X-rays in a depletion layer of a semiconductor in an analog manner using a semiconductor sensor matrix, and digitizes the ionized signal using an A / D converter; A detector that directly extracts digital signals by counting the ionization signals individually without using the / D converter can be considered. For the former, for example, general semiconductor materials such as Si and GaAs can be used.
The latter is preferably one using HgI ₂ or dTe.

The first and second storage element matrices 5,
6 is preferable because a general so-called IC memory is the most economical and compact. The number of pixels (corresponding to the number of detection elements) required in the medical image is at least 51.
In order to obtain a resolution equivalent to that of a general X-ray film with a 2 × 512 matrix, the number of pixels of a 2048 × 2048 matrix or more is required, and the number of bits of one pixel is 12 bits or more, preferably at least 16 bits. is necessary. This means that the dynamic range required for one X-ray image is three or more decimal digits, that the change in X-ray dose during imaging, that is, the dose change due to X-ray tube current and voltage is one digit, and that the thickness of the subject Since there is more than one order of magnitude change, a dynamic range of five or more digits is required to satisfy these. 16-bit resolution is 2
Since the resolution is the 16th power of 65536, a dynamic range exceeding 5 digits is satisfied.

[0032]

According to the X-ray imaging apparatus of the present invention,
Since the intensity distribution of the X-rays generated from the X-ray tube apparatus is normalized and imaged, an image with excellent density uniformity can be obtained. Accordingly, even if the distance between the X-ray tube device and the detection unit is shortened, a good image with excellent density uniformity can be obtained, and as a result, the utilization rate of X-rays can be improved. . Therefore, the load on the X-ray tube apparatus is reduced, and the amount of heat generated can be suppressed, and the time required for diagnosis can be reduced by downsizing the apparatus and shortening the cooling time.

[Brief description of the drawings]

FIG. 1 is a schematic diagram of an X-ray imaging apparatus according to the present invention.

FIG. 2 is a flowchart illustrating an operation of a correction unit according to the present invention.

FIG. 3 is a schematic diagram of a conventional X-ray imaging apparatus.

FIG. 4 is a schematic diagram of an X-ray tube device.

[Explanation of symbols]

 DESCRIPTION OF SYMBOLS 1 ... X-ray tube apparatus 2 ... Subject 3 ... Detector 5 ... First storage element matrix 6 ... Second storage element matrix 7 ... Display circuit 8 ... Display Apparatus 9 ... Correction means

Claims (1)

(57) [Claims] 1. An X-ray tube device for generating X-rays, an X-ray detecting means comprising a plurality of detecting elements, and outputting irradiated X-rays as data for each element; A storage element matrix for storing the data of each element of the irradiated X-rays, and a correction means for normalizing the data of each element of the X-rays transmitted through the subject with the corresponding data stored in the storage element matrix An X-ray imaging apparatus comprising: a display unit configured to display the normalized data as an X-ray image.