JPH04159879A - Method and device for converting radiation picture - Google Patents

Abstract

PURPOSE:To prevent the deterioration of the picture quality of radiation pictures by scattered radiations by taking the radiation picture of an object by placing a structure constituted of alternately arranged radiopaque and radiotransparent members between the object and a stimulable phosphor and correcting the reading value of the shadow part of the radiotransparent members near the shadow part of the radiopaque members on the basis of the reading value of the shadow part of the radiopaque members. CONSTITUTION:A structure 4 constituted of alternately arranged radiopaque members 2 and radiotransparent members 3 is placed between an object 12 and stimulable phosphor 1. At the time of reading the latent image formed in the phosphor 1, the amount of scattered radiations reaching the part of the phosphor 1 in the shade of the radiopaque members 2 is found directly or through prescribed arithmetic operation from the reading value of the part. Then the found amount of scattered radiations is subtracted from the reading value of the shadow part of the radiotransparent members 3 near the shadow part of the members 2. Therefore, clear radiation pictures can be obtained by removing the influence of the scattered radiations from the reading value of the stimulable phosphor 1.

Description

[Detailed Description of the Invention] [Summary] The present invention relates to a radiation image conversion method and device using a stimulable phosphor, and aims to improve image quality deterioration due to scattered radiation without increasing the amount of radiation exposure. A radiation image that obtains image information by converting the stimulated luminescent light emitted when excitation light is irradiated onto a photostimulated phosphor that absorbs a portion of the radiation energy that has passed through the subject and accumulates it as a latent image into an electrical signal. In the conversion method, a structure in which radiopaque members and radiotransparent members are arranged alternately is placed between the subject and the photostimulable phosphor, and the subject is photographed, and the structure formed on the photostimulable phosphor is When reading the latent image, the reading value of the shadow part of the radiopaque member is corrected based on the read value of the shadow part of the radiopaque member.

[Industrial application field]

The present invention relates to a radiation image conversion method and apparatus using a photostimulated phosphor.

[Conventional technology]

Radiographic images such as X-ray images are used for diagnosing diseases and the like. A known method for obtaining X-ray images is to irradiate X-rays that have passed through the subject onto a fluorescent layer (fluorescent screen) to generate visible light, and then irradiate this visible light onto a silver halide film to obtain an X-ray image. ing.

In addition to obtaining two-dimensional images of indirect or direct radiation on silver halide films, when excitation light is irradiated onto a latent image formed on a stimulable (photostimulable) phosphor by X-rays that have passed through the subject. 2. Description of the Related Art X-ray digital diagnostic devices such as computed radiography that detect emitted stimulated luminescence light using a photoelectric converter or the like to obtain a digital image are known.

Here, the stimulable phosphor is one that accumulates a part of the energy when it receives the energy of radiation such as X-rays, and the accumulated energy is retained for a certain period of time. When the stimulable phosphor in this state is irradiated with the first light that acts as excitation light, the stored energy is released as second light (stimulated luminescent light). At this time, the first light is not limited to visible light, but has a wide wavelength range from infrared to ultraviolet light, and the selection thereof is determined by the phosphor material used. Further, the second light emitted as stimulated luminescence light includes various types of light ranging from infrared rays to ultraviolet rays, and depends on the phosphor material used.

Next, the configuration of a digital X-ray imaging device will be described.

Figure 7 shows the system configuration when the imaging section and reading section are configured as separate devices, and Figure 8 shows the system configuration when the imaging section and reading section are integrated. ing.

In FIGS. 7 and 8, X-rays emitted from the X-ray generating unit 11 and transmitted through the subject 12 cause a stimulable phosphor plate (or sheet) provided in the photographing table 13 or the imaging device [14]
A latent image is formed on 15.

A system in which the photographing table 13 and the reading section are separated (7th
In FIG. 1, after the photostimulated phosphor plate 15 on which the latent image has been formed is inserted into the insertion section 17 of the read-only device 16, the reading section 1
8, and the latent image is read. Further, in a system in which a photographing section and a reading section are integrated (FIG. 8), the stimulable phosphor plate 15 is directly sent to the reading section 18 and the latent image is read.

The read image information is subjected to image processing in an image processing section 19, and then displayed as a CRT image on an image display section 20 or hard-coded onto an X-ray film.

The stimulable phosphor plate 15 that has been read has its afterimage erased in the erasing section 21, and is sent to the photographing section (in the case of the system shown in FIG. 8) or the take-out section 22 (in the case of the system shown in FIG. 7).

A digital X-ray imaging device using this stimulable phosphor uses a silver halide film sandwiched between conventional intensifying screens (S/F method).
Compared to imaging devices that use It was no better than an X-ray image.

One of the factors that affects image quality is the X that occurs within the subject.
Due to the scattering of vA, an image that should originally be formed due to the difference in the transmittance of X-rays of the subject may become blurred due to the influence of the scattered X-rays.

A grid is generally used to remove the scattered X-rays. The grid is already used in both the S/F method and an imaging system using a stimulable phosphor, and is placed between the subject 12 and the stimulable phosphor plate I5 as shown in FIG.

As shown in the plan view of FIG. 10(a) and the cross-sectional view of FIG.
The line-transparent materials 25 are arranged alternately at equal intervals,
These materials are arranged on the entire surface of the grid 23. Specifically, the radiopaque material 24 is often lead, and the radiolucent material 25 is often aluminum. In addition, the thickness of the grid 23 is about 1 to 21111, the width of the lead foil (thickness when viewed as a foil) is about 50 μm, the ratio of the width of lead to aluminum is 172 to 1/4, and the pitch is 0.17.
~o, about 34 pengs.

Next, the effect of preventing scattered X-rays by the grid 23 will be explained using FIG. 11, taking the case of a silver halide film as an example.

When the emitted X-rays 26 are irradiated almost parallel to the object (X-ray scatterer) 12, scattered X-rays 27 are generated from the object 12 as shown by dotted lines in FIG. Scattered X! S! 27 arrives. However, when the grid 23 is provided, the scattered X-rays 27 having a large angle are blocked by the X-ray opaque material 24, and the scattered X-rays can be prevented from spreading over a wide range. Incidentally, a sensitizing layer 29 is provided on both sides of the silver salt film 28 to increase the sensitivity of the film to X-rays.

However, in the grid 23 described above, the subject 12
It was not possible to completely block the scattered X-rays 27 reaching the silver salt film 28 near the .

In order to solve this problem, it is conceivable to narrow the pitch of the X-ray opaque material 24 and to increase the aspect ratio of the X-ray opaque material 24. In fact, improvements to the grid 23 have been made in that direction.

Furthermore, the grid 23 shown in FIG. 10 is one in which an X-ray opaque material 24, for example, lead is arranged in one direction (the vertical direction in FIG. 10), but lead is arranged in both the vertical and horizontal directions. This is also being attempted.

[Problem to be solved by the invention]

However, all of the above-mentioned measures are
In order to reduce the amount of X-rays reaching 8, a highly sensitive intensifying screen (sensitizing layer) 29 and a silver salt film 28 are required.

However, in reality, it is difficult to realize a highly sensitive intensifying screen 29 and silver halide film 28, so measures are taken to increase the amount of X-ray irradiation, and as a result, the amount of radiation exposure to the human body increases. There was a problem with letting it work.

In addition, instead of narrowing the interval between the X-ray opaque materials 240, a method has been implemented in which the grid 23 is moved left and right in synchronization with the timing of X-ray irradiation, but this method also increases the amount of X-ray irradiation. This did not completely solve the above problems.

The above-mentioned problems are not limited to X-ray imaging devices using the silver halide film 28, but are the same in digital X-ray imaging devices using stimulable phosphors, and it has been desired to solve these problems.

An object of the present invention is to improve image quality deterioration caused by scattered radiation without increasing the amount of radiation exposure.

[Means to solve the problem]

FIG. 1(a) is a diagram illustrating the principle of the invention according to claim 1.

Image information is obtained by converting the stimulated luminescent light emitted when excitation light is irradiated onto the stimulated phosphor 1, which absorbs a portion of the radiation energy that has passed through the subject 12 and accumulates it as a latent image, into an electrical signal. In the radiation image conversion method, in the invention according to claim 1, a structure 4 in which radiopaque members 2 and radiotransparent members 3 are arranged alternately is installed between the subject 12 and the photostimulable phosphor 1. When reading the latent image formed on the Rinjin phosphor 1, based on the read value of the shadowed portion of the radiopaque member 2, the shadowed portion of the radiotransparent member 3 in the vicinity thereof is determined. Correct the reading.

As the above correction method, for example, if the radiopaque member 2 is thick and the radiopaque member 2 is
If the radiation that flies straight into the shadowed area of 2 is almost negligible, use the read value of the shadowed area of the radiopaque member 2 to calculate the shadowed area of the nearby radiolucent member 3. Correct the reading.

For example, if the radiopaque member 2 is relatively thin and the amount of radiation that passes through the radiopaque member 2 and reaches its shadow area cannot be ignored, the subject 12 may not be placed or the radiation When a uniform phantom with little scattering is placed and photographed, the reading value of the shadowed part of the radiopaque member 2 is c, the reading value of the shadowed part of the radiotransparent member 3 in the vicinity is S, If the reading value of the shadowed part of the radiopaque member 2 when the subject 12 is placed and photographed is c, and the reading value of the shadowed part of the radiolucent member 3 in the vicinity is S, then C -S
c/s is subtracted from the read value S of the shadow portion of the radiation transparent member 3 as a correction value for the scattered radiation dose.

Alternatively, the shadow portions 5a and 5b of two adjacent radiopaque members 2 when the subject 12 is placed and photographed
The reading value of Ca is C11+1, and the reading value of the shadow portion 6c of the radiolucent member 3 sandwiched between these two radiopaque members 2 is S. , the reading values of the shadow parts 6a and 6b of the two radiotransparent members 3 sandwiching the two radiopaque members 2 are S n-1, Sri+1
, then ccn-c s, XH+,) 10C,,l −(S,,l
/2 [(tanshi, S, = (sn-+ +Sn)/2
, Ss+1''(Sn+Sn+1)/2, H,=c/s
) is subtracted from the reading pattern S1 as a correction value for scattered radiation.

Further, in the radiation image conversion method according to the fifth aspect, after correcting the temporal tailing of stimulated luminescence, the scattered radiation is corrected. As a method for correcting temporal tailing of stimulated luminescence,
For example, by determining the amount of light emitted by a photostimulable phosphor at a scanning position before the current scanning position from multiple readings in a time series, and subtracting that amount from the reading at that time,
The effect of temporal tailing of stimulated luminescence is corrected.

Next, FIG. 1(C) is a principle explanatory diagram corresponding to the invention set forth in claim 7.

In the same figure, in the structure 4, radiopaque members 2 and radiolucent members 3 are arranged alternately, and the subject (
12) and is installed near the photostimulable phosphor 1.

The reading means 7 reads the latent image formed by the radiation that reaches the photostimulable phosphor 1 via the structure 4.

Based on the read value of the shadow portion of the radiopaque member 2 read by the reading means 7, the correction means 8
The read value of the shadowed portion of the radiation transparent member 3 in the vicinity is corrected.

[For production]

In the above radiation image conversion method, first, from the read value of the portion of the photostimulable phosphor 1 that is shadowed by the radiopaque member 2,
The amount of scattered radiation reaching the shadow area is determined directly or by a predetermined calculation.

Then, the amount of scattered radiation of the shadowed portion of the radiopaque member 2 is subtracted from the read value of the shadowed portion of the radiation transparent member 3 in the vicinity.

Thereby, the influence of scattered radiation can be removed from the read values of the photostimulable phosphor 1, and a clear radiation image can be obtained.

Further, in the radiation image conversion method according to claim 5, since the scattered radiation is corrected for the value corrected for the temporal tailing of stimulated luminescence, a clearer radiation image is obtained by removing both errors. be able to.

Furthermore, in the radiation image conversion apparatus according to the seventh aspect, since the amount of scattered radiation included in the read value of the photostimulable phosphor 1 can be removed, a clearer radiation image without image blurring can be obtained.

〔Example〕

Embodiments of the present invention will be described below with reference to the drawings.

FIG. 2 is a configuration diagram of a reading section of a digital X-ray imaging apparatus based on the radiation image conversion method of the present invention. The overall configuration of the digital X-ray imaging apparatus of this embodiment is similar to that of the digital X-ray imaging apparatus shown in FIG. 7 or 8.

In FIG. 2, a stimulable phosphor plate (or sheet) 15 on which a χ latent image of a subject 12 is formed is movable in the direction of the arrow in the figure by a moving mechanism (not shown). Excitation light emitted from an excitation light source 31 such as a laser is scanned by a scanner 32 made of a polygon mirror or the like in a direction perpendicular to the moving direction (arrow direction) of the stimulable phosphor 15.
After the beam shape is further corrected by a lens 33, the beam passes through a reflection mirror 34 and is irradiated onto the stimulable phosphor plate 15.

Stimulated luminescence light is emitted from the χ-ray latent image of the stimulable phosphor plate 15 by scanning the excitation light, but this stimulated luminescence light is collected by a condenser 35 such as a fiber array and is converted into a photoelectron. A converter 36 converts it into an electrical signal. This electrical signal is amplified by the amplifier 37, then converted to a digital signal by the A/D converter 3B, and stored in the frame memory 39.

Furthermore, the digital signal stored in the frame memory 39 is subjected to correction of scattered X-rays, etc., which will be described later, in the image processing section 41, and is subjected to gradation correction, frequency processing, etc. as necessary, and then processed to the magnetic disk 40, etc. is stored as image data.

Next, a method for correcting scattered X-rays based on the radiation image conversion method of the present invention will be described based on the digital X-ray imaging apparatus having the above configuration.

Now, a grid 23 (for example, as shown in FIG. (a)),
The X-ray 26 is placed near the stimulable phosphor 15 on the side where the radiation is generated.
irradiate.

At short distances, the X-rays 26 are parallel and straight to the grid 23.
Therefore, if the X-ray opaque material 24 is lead and its thickness is, for example, about 2 mm or more,
The X-rays 26 that pass through the subject 12 and travel linearly are blocked by the X-ray opaque material 24, and the X-rays 26 that travel linearly are blocked by the X-ray opaque material 24. I never reach it.

That is, in the shadow portion 42 of the radiopaque material 24,
It may be considered that only the scattered X-rays 27 (FIG. 4(a)) arrive. Therefore, the X-ray dose of the shadow part 42 of the X-ray opaque material 24 is calculated by
By subtracting it from the X-ray dose of 3, the influence of scattered X-rays can be removed from the read value of the stimulable phosphor plate 15.

Therefore, a clear digital X-ray image without image blur caused by scattered X-rays can be obtained.

However, in the case of lead, if the thickness is more than about 211ff1, the amount of X-rays reaching the stimulable phosphor plate 15 will decrease and the SZN ratio will decrease, so it is necessary to widen the pitch of the X-ray opaque material 24. be. In this case, the pitch does not necessarily have to be an integral multiple of one pixel, and the pitch may be changed depending on the location.

Next, the radiopaque material 24 is relatively thin (the subject 12
A method of correcting scattered X-rays when the radiation coming straight from the object cannot be ignored will be explained.

First, let's look at the case where the image is taken without a subject or with a uniform phantom with little scattering of radiation.
This will be explained with reference to Figures (a) to (C).

In this case, a slightly smaller amount of X-rays than the irradiated X-ray dose reaches the shadowed portion 43 of the X-ray transparent material 25 in FIG. , a much smaller amount of X-rays will be reached.

As a result, the distribution of the amount of X-rays absorbed by the photostimulable phosphor plate 15 becomes rectangular as shown in FIG.
A latent image is formed on 5 according to the distribution of the absorbed X-ray dose.

When excitation light is scanned over this X-ray latent image and the emitted stimulated luminescence light is read, a latent image with the absorption X-ray dose distribution shown in Figure 3 (C) is obtained as shown in Figure 3 (C). The value is read as shown.

Here, the reason why the read value in Fig. 3 (C) does not become a rectangle as shown in Fig. 3 (B) is because the beam diameter of the excitation light has a certain size, so for example, the scanning position of the excitation light from the shadow part 43 of the X-ray transparent material 25 to the X-ray opaque material 2
This is because even if the subject moves to the shaded area 42 of No. 4, the stimulated luminescence light in the two regions is detected, and the stimulated luminescence light does not immediately become zero.

As shown in FIG. 3(C), if the read value of the shaded part 42 of the X-ray opaque material 24 is c, and the read value of the shaded part 43 of the X-ray transparent material 25 is S, each reading position The read values C and S are approximately the same value. At least C/S has approximately the same value.

Next, the case where the object (X-ray scatterer) 12 is placed and photographed will be described with reference to FIGS. 4(a) to 4(C).

In this case, scattered X-rays 27 are generated inside the subject 12, and the scattered X-rays 27 and the X-rays that have linearly passed through the subject 12 reach the grid 23. Then, a part or most of those X-rays reach the stimulable phosphor plate 15. As a result,
The X-ray absorption amount distribution of the photostimulable phosphor plate 15 is shown, for example, in FIG.
As shown in b), latent images are formed according to the respective X-ray absorption amounts.

When excitation light is scanned over this X-ray latent image and the emitted stimulated luminescence light is read, the latent image has the X-ray absorption distribution shown in Figure 4 (B), as shown in Figure 4 (C). A reading is obtained.

In FIG. 4(C), the reading values of the shaded part 42 of the radiopaque material 24 are C□8, C1l from the left side of the figure.
, C1m+1, and the shadow part 4 of the X-ray transparent material 25
Similarly, from the left side, read 5n-1, Sn, S.
□1.

Here, for example, the shadow portion 43c of the X-ray transparent material 25c
The amount of scattered X-rays included in the read value Sn is the amount of scattered
It can be considered to be approximately equal to the radiation dose. On the other hand, the read value C1 of the shadow part 42a of the X-ray opaque material 24a shows that the X-ray opaque material 2
It also includes the amount of X-rays that pass through 4a, and the amount is
Reading value s when photographing without subject weight 2,
c, and the reading value S of the shadow part 43c of the X-ray transparent material 25a when the photograph is taken with the subject 12 placed therein,
It can be approximated by S, , ×C/s.

Therefore, the value obtained by subtracting 5n×c/s from the read value C6 of the shadow portion 42a of the X-ray opaque material 24a becomes the scattered X-ray dose in that portion. Then, from the reading value S7 of the shadow part 43c of the X-ray transparent material 25c in the vicinity, the value obtained by eliminating the amount of scattered X-rays, that is, 5n-(c, -s.X
c/s) is a value proportional to the amount of X-rays that pass through the subject 12 and come in a straight line.

As a result of this, the influence of scattered X-rays can be removed from the read values of the stimulable phosphor plate 15, and clear image information can be obtained.

Further, the above correction method has the advantage that it is easy to implement, since it is sufficient to sequentially process the time-series read signals obtained when the photostimulable phosphor plate 15 is scanned.

Note that in the above correction method, the amount of scattered X-rays included in the read value Sn is determined from the amount of scattered X-rays in the shadow portion 42a of the X-ray opaque material 24a in the vicinity thereof; however, in the case of more accurate correction is the scattering X contained in the reading value Sn
The dose is determined by the reading values C and C1°1 of the shadowed portions 42a and 42b of the radiopaque materials 24a and 24b on both sides thereof.
It should be sought from Additionally, readings C1 and C
0,.

The amount of scattered X-rays contained in the adjacent X-ray transparent material 25
Shadow portions 43a, 43b and 4 of a, 25b and 25c
3c reading S n-1% Sr++1, S, .

In this case, S, −(sn−+ +Sn)/2, S, ,
When g = (SI, + 5rl-1) / 2, the amount of scattered X-rays (corrected value) included in the reading value S7 is (Cst-(s, x c/s) + C,,, -( S,,, ×
It can be expressed as c/s))/2.

Therefore, the influence of scattered X-rays can be removed from the read value of the shadowed part of the X-ray transparent material 25 at any position by subtracting the amount of scattered X-rays obtained from the above formula. It is possible to obtain clear X-ray images.

Note that the reading value C of the shadow part of the X-ray opaque material 24 when photographing without placing the subject 12 is the same as that of the subject 12.
When the reading value is very small compared to the reading value 07 etc. when taking a picture with the camera placed, (C.

+C,. I)/2 may be used as the correction value.

The method of correcting the scattered X-rays described above is based on the X-ray opaque material 24.
This is a correction method when one pitch corresponds to one pixel.

When it is desired to further reduce the amount of X-ray irradiation or to change the pixel density, it is not always possible to make one pitch correspond to one pixel. In this case, since the amount of scattered X-rays differs from pixel to pixel, it is necessary to perform correction for each pixel.

FIG. 5 shows the readings of the stimulable phosphor plate 15 when there are n pixels between the xig-opaque materials 24.

In FIG. 5, the reading values of the shadowed portions of the two X-ray opaque materials 24 are CII, C@+1, and the X-ray transparent material between the two X-ray opaque materials 24 is The reading value of the shaded part of 25 is Sl, S2 ・
...S, ...Sn, s, ll= (S-1+sI
)/2, S, (in this case, SI, l indicates the average value of SI and St). + = (Sn+S,,-
+ )/2, then the correction value (scattered X-ray dose) of the j-th pixel is ((n-j)(C,-311×c/s) +
It can be expressed as j(C#+1-8□, Xe/s))/n, or ((n-j)C, jxC,,+)/n.

Therefore, by subtracting the correction values determined by the above formula from the read values 31 to Sn of each pixel, it is possible to obtain X-ray image information from which the influence of scattered X-rays has been removed.

Providing a large number of pixels between the X-ray opaque materials can increase the amount of X-rays reaching the stimulable phosphor plate 15 without increasing the amount of chi-ray radiation.

Therefore, it is possible to suppress the amount of X-ray exposure to the human body and to improve image blurring caused by scattered X-rays, which has a great effect on improving the performance of the X-ray imaging device.

Incidentally, the above-mentioned correction methods are all correction methods for reading stimulated luminescence light at a relatively low scanning speed. Here, low speed has the following meaning.

The stimulated luminescence light emitted when a photostimulated phosphor is irradiated with excitation light does not immediately go to zero even after the excitation light irradiation is stopped.
It has the property of continuing to emit light while decaying exponentially. In other words, if the time when excitation light irradiation is stopped is O, the instantaneous amount of light emission is P, and the time constant of decay is τ (the time when P becomes 1/e), the amount of light emitted at time t, p(t), is as follows: It can be expressed as a formula.

When the time P (f,) = P
, is an integer multiple of τ, P (t) becomes a considerably large value.

For example, if we look at the case where a shadow part of the X-ray transparent material 24 is scanned from a shadow part of the X-ray opaque material 25, a large amount of X-rays reaches the former part, so
The amount of stimulated luminescence in that part increases. At this time, if the scanning speed of the excitation light is fast, the next position (X-ray opaque material 25
As a result, the relatively large amount of light emitted from the former part was added to the small amount of light emitted from the latter part, causing an error in the detected value of the X-ray dose.

Therefore, it is necessary to correct the temporal deviation of this stimulated luminescence before correcting the scattered X-rays.

FIG. 6 is a diagram showing the influence of tailing of stimulated luminescence light when scanning at high speed.

In the figure, the solid line shows the read value of stimulated luminescence light when scanning at low speed, and the dotted line shows the read value when scanning at high speed.

Now, if the reading values of the shadowed part of the X-ray opaque material 24 when scanning at high speed are respectively 57-1 and Sn, and their sampling time interval is Δt, then the reading value Sfi contains 5n- The amount of time reduction from 1 is 5n-I Xexp(-Δt/τ). By subtracting this value from the read value Sfi, it is possible to correct the trailing amount due to 5n-1.

If this correction is performed for all read values, it is possible to correct errors due to temporal differences in stimulated luminescence.

In this way, an accurate amount of luminescence is determined by correcting the temporal effect of stimulated luminescence in the shadow part of the X-ray opaque material 24,
Further, by correcting the read value of the shadowed portion of the X-ray transparent material 25 as described above based on the amount of light emitted from the shadowed portion, more accurate X-ray image information can be obtained.

In addition, in the above correction method, although the case of correcting the influence of temporal bias due to stimulated luminescence at the scanning position one sampling time ago, the tail effect due to stimulated luminescence at the scanning position n sampling time before is corrected. Pull can also be corrected in the same way.

When performing the above-described correction, it is necessary to know where on the stimulable phosphor plate 15 the X-ray opaque material 24 is placed. The most advantageous method in terms of positional relationship is the eighth
This is a device in which the photographing section and the reading section shown in the figure are integrated, and unlike the figure, the stimulable phosphor plate 15 is fixed, and the reading section is placed near the stimulable phosphor plate 15. , which has a structure in which the reading section moves.

In this case, the stimulable phosphor plate 15 and the X-ray opaque material 24
That is, since the positional relationship with the grid 23 does not change, correction can be performed very accurately.

Further, a stimulable phosphor plate 15 as shown in FIGS. 7 and 8 may also be used.
In a device having a structure in which the
By adhering or closely fixing a structure (grid) in which radio-opaque materials 24 and By integrally forming the transparent layer and the X-ray transparent layer, the positional relationship between the two can be maintained.

Furthermore, the position of the X-ray opaque material 24 of the above-mentioned structure is measured over the entire surface, and the reading of the stimulable phosphor plate 15 is performed at a low speed in the initial stage, and from the reading result, the X-ray opaque material 24 is measured. By accurately detecting the position of the transparent material 24, methods for determining the position of other parts can also be used.

Furthermore, since it is roughly known where the shadowed part of the X-ray opaque material 24 is located on the stimulable phosphor plate 15,
It is also possible to use a method of correcting tailing over the entire surface and then detecting the valley at that position.

In any case, correction is possible if the position of the X-ray opaque material 24 can be detected with a certain level of accuracy or higher. For example, correction is possible if the positional relationship can be known with an accuracy of about 172 or less of the width of the X-ray opaque material 24 used.

It is relatively easy to achieve an accuracy of about 10 μm for the mechanical parts of a normal conveyance system, so for example, if a lead foil with a radius of 50 μm is
When used with radiopaque material 24, sufficient correction is possible.

Note that the present invention can be applied to radiation other than X-rays, and can be applied to diagnostic devices and medical equipment other than X-ray imaging devices.

〔Effect of the invention〕

According to the present invention, the influence of scattered radiation from the subject can be removed using the read value of the shadowed portion of the radiopaque member. Thereby, a clear radiation image can be obtained without increasing the radiation dose irradiated to the subject.

[Brief explanation of drawings]

FIGS. 1(a) to (C) are diagrams explaining the principle of the present invention, FIG. 2 is a configuration diagram of a reading section of a digital X-ray imaging device according to an embodiment of the present invention, and FIG. 3(a) is An explanatory diagram when photographing without a subject, Figure 3 (c) shows the X of the photostimulable phosphor photographed without a subject.
Figure 3 (C) is a diagram showing the readings from a photostimulable phosphor photographed without a subject, and Figure 4 (a) is a diagram showing the radiation absorption distribution when photographed with an X-ray scatterer. Explanatory diagram: Figure 4(b) is an X-ray absorption distribution diagram of a photostimulated phosphor photographed with an X-ray scatterer placed, and Figure 4(C) is a photostimulated phosphor photographed with an X-ray scatterer placed. Figure 5 shows the readings from the phosphor. Figure 5 shows the readings of the photostimulable phosphor when there are a large number of pixels between the radiopaque materials. Figure 6 shows the readings taken at high speed. Figures 7 and 8 are system configuration diagrams of a boat-mounted X-ray imaging device; Figure 9 is a side view of an imaging platform with a grid; Figure 10 (a
) is a plan view of the grid, FIG. 10(b) is a sectional view of the grid, and FIG. 11 is an explanatory diagram of the effect of the grid. DESCRIPTION OF SYMBOLS 1... Stimulated phosphor, 2... Radio-opaque member, 3... Radio-transparent member, 4... Structure, 7... Reading means, 8... Correction means.

Claims (1)

[Scope of Claims] 1) Stimulated luminescence light emitted when excitation light is irradiated to the stimulable phosphor (1), which absorbs a part of the radiation energy transmitted through the subject (12) and accumulates as a latent image. In a radiation image conversion method for obtaining image information by converting into electrical signals, a radiopaque member (2) and a radiotransparent member (3) are disposed between a subject (12) and a photostimulable phosphor (1). ) are arranged alternately to photograph the subject (12), and when reading the latent image formed on the photostimulable phosphor (1), the radiopaque member (2) A radiographic image conversion method comprising: correcting a reading value of a shadowed portion of a radiation transparent member (3) in the vicinity of the shadowed portion based on a readout value of a shadowed portion of the radiation transparent member (3). 2) Converting the stimulated luminescent light emitted when excitation light is irradiated onto the photostimulated phosphor (1), which absorbs a portion of the radiation energy that has passed through the subject (12) and accumulates it as a latent image, into an electrical signal. In the radiation image conversion method for obtaining image information, a radiation-opaque member (2) and a radiation-transparent member (3) are alternately arranged in the vicinity of the stimulable phosphor (1) on the radiation generating part side. The radiopaque member (1) of the photostimulable phosphor (1) when photographing is carried out with the structure (4) set up and no object (12) placed or a uniform phantom with little radiation scattering placed. 2) The reading value of the shadowed part is c, and the radiolucent member in the vicinity (
Let s be the read value of the shadowed part of 3), and let the subject (1
2) when the radiopaque member (
Let C be the reading value of the shadowed part of 2), and S be the readout value of the shadowed part of the radiation transparent member (3) in the vicinity, then from the reading value S, C-S×c/ A radiation image conversion method characterized in that a value obtained by subtracting s is used as a read value of a shadowed portion of the radiation transparent member (3). 3) Converting the stimulated luminescent light emitted when excitation light is irradiated onto the stimulated phosphor (1), which absorbs a portion of the radiation energy that has passed through the subject (12) and accumulates it as a latent image, into an electrical signal. In the radiation image conversion method for obtaining image information, a radiation-opaque member (2) and a radiation-transparent member (3) are alternately arranged near the photostimulable phosphor (1) on the side of the radiation generating part. The radiopaque member (2) of the photostimulable phosphor (1) when photographing is performed with the structure (4) installed and no object (12) or a uniform phantom with little radiation scattering placed. ) is the reading value of the shaded part c
, the reading value of the shadowed part of the radiopaque member (3) in the vicinity is s, which is the shadow of the two adjacent radiopaque members (2) when the subject (12) is placed and photographed. The reading values of the portions are respectively C_m and C_m_+_1, and the reading values of the portions that are in the shadow of the radiolucent member (3) sandwiching the two adjacent radiopaque members (2) are S_m, respectively.
n_-_1, S_n_+_1, and when the read value of the shadowed part of the radiolucent member (3) sandwiched between the two adjacent radiopaque members (2) is S_n, [C_n-(S_m ×H_m)+C_m_+_1-(S
_m_+_1×H_m_+_1)]/2 [However, S_m
=(S_n_-_1+S_n)/2, S_m_+1=(
A radiation image conversion method characterized by subtracting S_n+S_n_+_1)/2, H_m=c/s] from a read value S_n as a correction value. 4) Converting the stimulated luminescence light emitted when excitation light is irradiated onto the stimulated phosphor (1), which absorbs a portion of the radiation energy that has passed through the subject (12) and accumulates it as a latent image, into an electrical signal. In the radiation image conversion method for obtaining image information, a radiation-opaque member (2) and a radiation-transparent member (3) are alternately arranged near the photostimulable phosphor (1) on the side of the radiation generating part. A structure (4) is installed, and the photostimulated phosphor (1)
A plurality of pixels are set in the shadow part of the radio-transparent member (2), and when correcting the scattered radiation from the subject (12), two radio-opaque pixels sandwiching the radio-transparent member (3) are set. A radiation image characterized in that a value obtained by weighting according to the position of each pixel using a read value of a shadowed portion of the sexual member (2) is used as a correction value for the scattered radiation dose of each pixel. Conversion method. 5) In the radiation image conversion method according to claim 1 or 4, the read value of the photostimulable phosphor (1) is corrected for temporal tailing of stimulated luminescence, and then the scattered radiation is corrected. A radiation image conversion method characterized by: 6) When the two time-series reading values of the photostimulable phosphor (1) are S_n_-_1 and S_m, the time constant of temporal tailing of stimulated luminescence is τ, and the sampling time interval of reading is Δt. , S_n_-_1×exp(-Δt/τ) as a correction value for temporal tailing, which is subtracted from the read value S_n of the photostimulable phosphor (1). . 7) Converting the stimulated luminescent light emitted when excitation light is irradiated onto the stimulated phosphor (1), which absorbs a portion of the radiation energy that has passed through the subject (12) and accumulates it as a latent image, into an electrical signal. In a radiation image conversion device that obtains image information, a radiation-opaque member (2) and a radiation-transparent member (3) are arranged alternately, and a stimulable phosphor (1) on the side of the radiation generating part is arranged.
a structure (4) installed near the stimulable phosphor (1); a reading means (7) for reading the latent image formed on the photostimulable phosphor (1); and a radiopaque member read by the reading means (7). A radiation radiation source characterized by comprising: a correction means (8) for correcting a read value of the shadow portion of the radiation transparent member (3) in the vicinity based on a read value of the shadow portion of (2); Image conversion device. 8) Converting the stimulated luminescent light emitted when excitation light is irradiated onto the stimulated phosphor (1), which absorbs a portion of the radiation energy that has passed through the subject (12) and accumulates it as a latent image, into an electrical signal. In a radiation image conversion device that obtains image information, a radiation-opaque member (2) and a radiation-transparent member (3) are arranged alternately, and a stimulable phosphor (1) on the side of the radiation generating part is arranged.
a structure (4) installed near the photostimulable phosphor (1), a reading means (7) for reading the latent image formed on the photostimulable phosphor (1), and a radiation-transparent member ( When a plurality of pixels exist in the shadowed portion of 3), using the reading values of the shadowed portions of the two radiopaque members (2) sandwiching the radiolucent member (3), A radiation image conversion device comprising: a correction unit that calculates the amount of scattered radiation of each pixel by performing weighting according to the position of each pixel, and corrects the read value of each pixel using the amount of scattered radiation. . 9) Converting the stimulated luminescent light emitted when excitation light is irradiated onto the stimulated phosphor (1), which absorbs a portion of the radiation energy that has passed through the subject (12) and accumulates it as a latent image, into an electrical signal. In a radiation image conversion device that obtains image information, a radiation-opaque member (2) and a radiation-transparent member (3) are arranged alternately, and a stimulable phosphor (1) on the side of the radiation generating part is arranged.
a structure (4) installed near the photostimulable phosphor (1); a reading means (7) for reading a latent image formed on the photostimulable phosphor (1); a tailing correction means for correcting temporal tailing; and based on the read value of the shadowed portion of the radiopaque member (2) corrected by the tailing correction means, the radiation transmission in the vicinity thereof is corrected. A radiation image conversion device comprising: a correction means for correcting a read value of a portion that is a shadow of the sexual member (3). 10) The radiation image conversion device according to claim 7, wherein the structure (4) comprises a radiation-opaque layer and a radiation-transparent layer formed on the radiation-irradiated surface of the stimulable phosphor (1).