An X-ray generator projects X-rays into an object under examination. An X-ray detector detects the X-ray image data transmitted through the object. The transmitted X-ray image data is digitized by an A/D converter and visually displayed by a monitor. For effecting such display, an X-ray shield member having an X-ray shield section configured in a predetermined pattern is set in an X-ray projection area. The transmitted X-ray image data containing the X-ray shield section data obtained under this condition is supplied to a scattered X-ray intensity computing circuit through first and second switching circuits and a memory. The computing circuit computes the scattered X-ray component. Then, the X-ray shield member is retracted from the X-ray projection area. Under this condition, the transmitted X-ray image data is supplied through the first and second switching circuits to a subtracting circuit. The substracting circuit subtracts the scattered X-ray component previously obtained from the transmitted X-ray image data, thereby to send only the primary X-ray component to the monitor.

BACKGROUND OF THE INVENTION 
This invention relates to an X-ray diagnostic apparatus in which a 
transmitted X-ray image of an object to be examined, e.g., a patient, is 
available for diagnostic purposes, and more particularly to an X-ray 
diagnostic apparatus by which a transmitted X-ray image of the object is 
obtained, based only upon primary X-rays, without any adverse influences 
caused by the scattered X-ray as well as the systems structural factors. 
Generally, in the X-ray diagnostic apparatus set forth in the preamble, 
X-rays incident on an X-ray detector contain not only primary X-rays but 
also X-rays which are scattered by the object under examination. The 
scattered X-rays constitute one of the major causes of deteriorated 
contrast and resolution in the transmitted X-ray image. This makes it 
necessary to eliminate an image component on the scattered X-rays from the 
transmitted X-ray image data as sensed and provided by the detector. 
One of the approaches to eliminate the scattered X-ray component is to use 
a so-called "Buckey Blend" or an elimination grid for the scattered X-rays 
(referred to as a "grid"). This approach also involves a problem in that 
there is a limit in the scattered X-ray elimination because the grid per 
se scatters the X-rays incident thereupon. 
The elimination of the scattered X-rays is very significant in the field of 
X-ray diagnosis for the reasons that it improves an image quality, such as 
contrast and resolution, and thus allows a logarithm conversion of primary 
X-rays image data, thereby obtaining an accurate attenuation quantity of 
X-rays caused when the X-rays pass through the object. Many studies have 
been made on the scattered X-rays, aiming at their effective elimination. 
The complicated phenomena of the scattered X-rays impede or almost reject 
a theoretical approach to this problem. This is the present stage of 
technology in this field. 
SUMMARY OF THE INVENTION 
For the above background reasons, an object of the present invention is to 
provide, by introducing a novel technical idea, an X-ray diagnostic 
apparatus which can effectively eliminate the scattered X-ray image 
component from the transmitted X-ray image components as obtained by the 
X-ray detector. 
In an X-ray diagnostic apparatus according to the present invention, an 
X-ray shield means having an X-ray shield section configured in a 
predetermined pattern may be placed in and retracted from an X-ray 
projection area in visually displaying the digital image data of X-rays 
output by an X-ray detector, the data being transmitted through an object 
as the result of X-ray irradiation of the object under diagnosis. A means 
computes the scattered X-ray component using the X-ray transmission data 
containing the image data collected from the X-ray shield section when the 
X-ray shield means is placed in the X-ray projection area. A compensating 
means subtracts the scattered X-ray component, computed and produced by 
the scattered X-ray computing means, from the X-ray transmission data, 
produced on the condition that the X-ray shield means is not placed in the 
X-ray projection area, thereby to produce only the primary X-ray 
component. 
With such an arrangement, the scattered X-ray component may effectively be 
removed from the transmitted X-ray data. Therefore, the displayed image is 
based only on the primary X-ray component. A high quality image is 
produced which is free from blur and improved in contrast and sharpness. 
Such a high quality image enables a doctor, for example, to perform 
effective diagnosis.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
Before proceeding with the description of the various preferred embodiments 
of the present invention, the principle of the present invention will be 
described in detail. 
In the present invention, X-rays are projected through the object to be 
examined under the condition that an X-ray shield member with a 
predetermined pattern is placed in the X-ray projection area. In such a 
condition, the transmitted X-ray image data detected by the detector in 
the area of the X-ray shield member contains only the scattered X-rays 
because the primary X-rays are shielded by the X-ray shield member. The 
present invention is based on the above assumption. 
To embody the above idea, the scattered X-ray component is calculated using 
the X-ray transmission image data obtained in the above way. 
Then, an object is irradiated with X-rays based on the condition that the 
X-ray shield member is not placed in the X-ray projection area. Under this 
condition, the X-ray transmission image data containing both the primary 
X-rays and the scattered X-rays is detected by the X-ray detector. The 
scattered X-rays as already calculated are substracted from the 
transmitted X-ray image data. As a result, detection data based on only 
the primary X-ray component are derived. 
The above principle leads to the first and second application principles as 
given hereinafter. 
It is assumed that X-rays incident on an object under examination are 
generally classified into "primary X-rays" which directly transmit through 
the object and enter into the X-ray detector, and "scattered X-ray" which 
are absorbed or scattered by the object through interactions of the X-rays 
with atoms constituting the object. In the energy range of medical X-rays 
(radiated under 50 KVp to 120 KVp of the X-ray tube voltage), some causes 
of X-ray scattering are known, for example, photoelectric effects, Compton 
effects, Thomson effects, and the like. These phenomena cooperate to cause 
the scattered X-rays to have adverse effects on the transmitted X-ray 
image to be described later. In general, because the scattered X-rays 
incident on the X-ray detector experience multi-scattering within the 
object, it is very difficult to exactly grasp an intensity and a spatial 
spread of an incident X-ray beam. This phenomenon is explained as follows. 
FIG. 1 schematically illustrates how an X-ray radiated from an X-ray source 
11, such as an X-ray tube, is scattered within an object 12 under 
examination and reaches an X-ray detector 13, while depicting a spatial 
spread with respect to detecting positions of the X-ray detector. FIG. 2 
illustrates the X-ray intensity distribution over the detecting positions 
of the X-ray detector 13. As seen from FIG. 2, a narrow spread, or spatial 
distribution of a sharp peak (as indicated by character K), located 
substantially at the center of the distribution curve, is caused by an 
inherent matter of the diagnosis system, for example, an X-ray focal spot 
and a wide spread (as indicated by character L) is caused by the scattered 
X-rays. 
In accordance with the study on the scattered X-rays by the inventors of 
the present patent application, the following recognition is made: in the 
medical X-ray energy range, the intensity distribution of scattered X-rays 
emanated from an object with a thickness substantially equal to that of a 
human body is generally expressed by the following equation; 
##EQU1## 
where Isc(x, y) indicates the intensity distribution of the scattered 
X-rays over detecting positions of the detector. The character A 
designates a constant in the above equation (1). The integration intervals 
-a to a and -b to b in the above equation define an area projected by the 
X-rays (referred to as an "X-ray projection area" hereafter) on the 
detecting positions of the detector. More exactly, -a.ltoreq.x.ltoreq.a 
and -b.ltoreq.y.ltoreq.b. In the above equation, f(Ip(x,y)) is a function 
of the primary X-ray intensity distribution Ip(x, y), and g(x,y) is a 
function defining the spatial spread of the scattered X-rays with respect 
to the incident X-rays as a pencil beam, and is a so-called "impulse 
response function". 
It is readily understood from the above description that this "impulse 
response function" means a function for defining the spatial spread of the 
scattered X-rays with respect to the incident X-rays as a fan-shaped beam 
or a parallel beam. The function g(x, y) satisfies the following equation 
(2) 
##EQU2## 
Generally, A, f(Ip(x, y)), and g(x, y) are determined by the tube voltage 
and the tube current of the X-ray tube, thickness of the object, distance 
between the object and the detector, and grid conditions respectively. 
As seen from equation (1), the intensity distribution of the scattered 
X-rays is given by a convolution integration of the function f(Ip(x, y)) 
of the primary X-ray intensity distribution and the function g(x, y) of 
the impulse response. The experiment conducted by the inventors showed 
that a specific form of the equation (1), as given by the following 
equation (3), well described the intensity distribution of the scattered 
X-rays. 
##EQU3## 
Our study further showed that in equation (3), A, n and g(x, y) depend on 
the tube voltage, the tube current, grid conditions, and the distance 
between the object and the detector, but scarcely depend on the thickness 
of the object. Of the above factors, n is selected between 0.5 and 1.5. 
The present invention is based largely on equation (3) defining the 
scattered X-ray intensity distribution. 
Equation (3) will further be discussed referring to FIG. 3. By convolution 
integrating the input signal Ip.sup.n (x') of the primary X-rays as shown 
in FIG. 3(A) and the function g(x, y) slightly varying with respect to the 
positions as shown in FIG. 3(B), we have an intensity distribution (Isc(x) 
of the scattered X-rays as shown in FIG. 3(C). For purposes of simplicity 
of illustration, the functions in FIGS. 3(A) to 3(C) are depicted in a 
two-dimensional manner. 
As seen from FIGS. 3(A) to 3(C), variation of the impulse response function 
g(x, y) is more slight than the variation of the primary X-ray 
distribution Ip(x, y) with respect to depositions. Based on this fact, 
equation (3) can approximately be rewritten into 
##EQU4## 
The total X-ray intensity distribution Im(x, y) incident on the detector is 
the sum of the primary X-ray intensity distribution Ip(x, y) and the 
scattered X-ray intensity distribution Isc(x, y), and is given by 
EQU Im(x, y)=Ip(x, y)+Isc(x, y) (5) 
and 
EQU Ip(x, y)=Im(x, y)-Isc(x, y) (5A) 
The equation (5A) implies that if the intensity distribution of the 
scattered X-rays is known, the primary X-ray distribution can be obtained 
because the total X-ray distribution can easily be obtained by measuring 
the X-rays directly incident on the X-ray detector. Fortunately, the 
intensity distribution of the scattered X-rays component Isc(x, y) is 
substantially equal over the entire projection area except at its 
peripheral edge. Therefore, an X-ray image formed by the primary X-rays 
can be obtained by subtracting the intensity of the scattered X-rays in a 
portion of substantially the center of the X-ray projection area, which is 
previously measured, from the transmitted X-ray image data derived from an 
object under examination. This is a first application principle of the 
invention. 
If the impulse response function g(x, y) can be expressed by a Gaussian 
function, it can be expressed as the following equation (2A) 
##EQU5## 
In the above equation, .sigma..sub.x and .sigma..sub.y, like the factors A 
and n, depend on the grid condition and the distance between the object 
under examination and the X-ray detector, but little depends on the tube 
voltage of the X-ray tube and the thickness of the object under 
examination. 
Therefore, equation (3) can be approximated to 
##EQU6## 
Also in this case, the above equation (5) holds, as a matter of course. 
Therefore, if the factors C, .sigma..sub.x and .sigma..sub.y are known, 
Isc(x, y) can be calculated using equation (4A) describing the intensity 
of the scattered X-rays. If Isc(x, y) is obtained, Ip(x, y) can be 
obtained by the equation (5A). Therefore, the scattered X-ray component is 
removed from an image composed of superposed images by primary X-rays and 
the scattered X-rays, using equation (4) describing the scattered X-ray 
model. This is a second application principle of the invention. 
The following describes a method for obtaining the factors C, .sigma..sub.x 
and .sigma..sub.y. 
For obtaining these factors, an X-ray shield plate is used which can shield 
the X-rays along the X-axis and Y-axis in the X-ray projection area. A 
thin plate 32, made of acryl resin, for example, and having a thin lead 
cross 31 thereon, as shown in FIG. 4, may be used for the X-ray shield 
plate. As shown in FIG. 5 the, acryl plate 32 is placed on a phantom 12', 
with an X-ray tube 11 located above the acryl plate 32. An X-ray detector 
13 is further disposed just under the phantom 12'. In operation, X-rays 
are radiated by the X-ray tube, transmitted through the acryl plate 32 and 
the phantom 12', and detected by the detector 13. The intensity of the 
transmitted X-rays differs with locations on the acryl plate 32 bearing 
the lead cross 31, which is irradiated with X-rays. For example, along a 
line L1 on the acryl plate 32 in FIG. 5, the X-ray intensity is 
distributed as shown in FIG. 6(A). The X-ray intensity along lines L2 and 
L3 on the acryl plate 32 are as shown in FIGS. 6(B) and 6(C). The X-ray 
intensities shown in FIGS. 6(B) and 6(C) are those of only the scattered 
X-rays since the primary X-rays are shielded by the lead 31 on acryl plate 
32. Using the equation (4A), we can express the X-ray intensity along the 
line L2 as follows 
##EQU7## 
In the above equations, C and C1 are constants and F1(x) is a function 
describing the x-dependency of the scattered X-rays. In a similar way, the 
X-ray intensity distribution along the line L3 can be expressed 
EQU Isc(o, y)=C.multidot.C2.multidot.F2(y) (9) 
##EQU8## 
In these expressions, C and C2 are constants, and F2(y) is a function 
describing the y-dependency of the scattered X-rays distribution. 
By applying a fitting process to the X-ray intensity distribution curve of 
the X-rays detected on the line L2, values for CxC1 and F1(x) most closely 
approximating the intensity distribution curve, can be obtained. C1 and C2 
can respectively be obtained by using the equations (7) and (10) since 
.sigma..sub.x and .sigma..sub.y can be obtained by F1(x) and F2(y) already 
obtained. 
In this way, values for C, .sigma..sub.x and .sigma..sub.y are obtained. 
In the embodying the present invention based on the second application 
principle, to find the scattered X-ray intensity distribution, both the x- 
and y-dependencies are not necessarily obtained, and it is satisfactory to 
obtain either of them if necessary. 
A first embodiment of an X-ray diagnostic apparatus according to the 
present invention, which is based on the first principle, will be 
described referring to FIGS. 7 and 10. 
In FIG. 7, there is shown an illustrative X-ray diagnostic apparatus as a 
first embodiment of the present invention, and is designed on the first 
application principle. 
In the figure, reference numeral 45 designates an X-ray generator such as 
an X-ray tube. An X-ray shield plate 47 is located within an X-ray 
projection area and between object 46 under examination and the X-ray tube 
45. The X-ray shield plate 47 will be described in detail later. An X-ray 
detector 48 detects X-rays transmitted through the object 46. An A/D 
converter 49 converts to digital data analog signals output from the X-ray 
detector 48. First and second switching circuit means 52 and 53 are 
provided between the A/D converter 49 and a memory 50. Fixed contacts B 
and B of these switches have a substractor 51 connected therebetween which 
is additionally connected to a scattered X-ray intensity-computing circuit 
means 54 for receiving data from the memory 50. 
The switching circuit means 52 and 53 are connected with a controller 55. 
These switching means cooperate to form a path between the A/D converter 
and the memory 50 when the object 46 is irradiated under a condition in 
which the X-ray shield plate 47 is placed in the X-ray projection area. 
Those switches form another path containing the subtracting circuit means 
51 between the A/D converter 49 and the memory 50 when the object is 
irradiated under a condition in which the X-ray shield plate 47 is placed 
outside the X-ray projection area. The memory 50 is for storing the 
detected image data output from the A/D converter 49 through the 
respective paths. A circuit means 54 computes the intensity of the 
scattered X-rays while using the detected image data fetched from the 
memory 50. The subtracting circuit means 51 subtracts the computed 
scattered X-ray intensities from the data collected under the condition 
that the object 46 is irradiated after removal of the X-ray shield plate 
47 from the X-ray projection area. The controller 5 controls the timing of 
the operation of the first and second switching circuit means 52 and 53, 
and the X-ray shield plate 47. A monitor 56, coupled with the output of 
the memory 50, visualizes for monitoring the image data of only the 
primary X-rays stored in the memory 50. An X-ray aperture 57, defining the 
X-ray projection area, is disposed close to the X-ray tube 25. 
A drive means (not shown) under control by the controller 55 drives the 
X-ray shield plate 47 to place it in and retract it from the X-ray 
projection area. The X-ray shield plate 47 is formed of a thin plate 47B 
made of acryl resin, for example, which has an X-ray shield member, such 
as a lead piece 47A, located at the center of the X-ray projection area, 
as shown in FIG. 8. The object 46 is irradiated by the X-ray tube 45 while 
the X-ray shield plate 47 is placed in the X-ray projection area, as shown 
in FIG. 9. In this test, the intensity distribution of the X-rays detected 
by the X-ray detector 48 along a line A--A' in FIG. 9 is plotted as shown 
in FIG. 10. It is seen from FIG. 10 that the distribution curve is 
substantially flat over the X-ray projection area, while having a local 
minimum point at the center of the X-ray projection area. The intensity of 
the scattered X-rays Isc at the minimum point is due to the scattered 
X-rays Isc and has substantially equal values over the entire X-ray 
projection area. 
The scattered X-ray intensity computing circuit means 54 averages the 
detected image data of picture elements in an area of the X-ray shield 
plate 47 covered by the lead piece 47A, and provides the averaged value as 
the scattered X-rays component input. 
The operation of the X-ray diagnostic apparatus thus arranged on the basis 
of the first principle will be described below. 
For obtaining the intensity of the scatttered X-rays, the X-ray shield 
plate 47 is placed in the X-ray projection area. Then, the switching 
circuit means 52 and 53 are set to the fixed contacts A and A to form a 
path directly connecting the A/D converter 49 to the memory 50. Then, the 
X-ray tube 45, is driven to irradiate the object 46 while the X-ray shield 
plate 47 is placed in the X-ray projection area. The X-ray detector 48 
detects the X-rays transmitted through the object 46 to produce analog 
output-image data. The image data generated by the portion of detector 48 
in line with lead piece 47A of shield plate 47 contains only the scattered 
X-rays. The image data is A/D converted by the A/D converter connected at 
the succeeding stage. The converted image data is stored in a first 
section of the memory 50 through the path A--A formed by the switching 
circuit means 52 and 53. The memory applies the image data of the picture 
elements corresponding to lead-piece covering portion on the X-ray shield 
plate 47 to the scattered X-ray intensity-computing circuit means 54. This 
circuit means 54 averages the stored data over the number of picture 
elements corresponding to shield section 47A, stored in the first section 
of memory 50, and produces it as the scattered X-ray component data. The 
averaging of such data is performed for reducing contained noise. 
In the next step, the X-ray shield plate 47 is retracted from the X-ray 
projection area. Then, the switching circuit means 52 and 53 are operated 
such that the output of the A/D converter 49 is connected through the 
subtraction circuit 51 to the memory 50. Succeedingly, the X-ray tube 
projects X-rays into the object 46, while the X-ray shield plate 47 is put 
outside the X-ray projection area. The X-ray detector 48 detects the 
X-rays transmitted through the object 46. The image data collected here 
contains both the primary X-rays and the scattered X-rays. The A/D 
converter 49 converts to digital data the image (analog) data output from 
the X-ray detector 48 and applies the digital data to the subtracting 
circuit means 51 through the first switch B. The subtracting circuit means 
51 subtracts from the image (digital) data thus obtained the detected data 
as the scattered X-ray component, which comes from the first section of 
memory 50 via the averaging means 54. Thus obtained image data of only the 
primary X-rays is applied to a second section of the memory 50 and then to 
the monitor 56. In this way, the monitor 56 displays the X-ray image based 
on only the primary X-rays. 
As described above, in the embodiment of FIG. 7, in the first step for 
obtaining the scattered X-rays component, X-rays are projected into the 
object 46 under examination under the condition that the X-ray shield 
plate 47 is placed in the X-ray projection area on the X-ray incident side 
of the object 46. In the second step for obtaining the image data 
containing both the primary X-rays component and the scattered X-rays 
component, the X-ray shield plate is retracted from the X-ray projection 
area, and the object 46 is irradiated with X-rays radiated from the X-ray 
tube 45. In the third step, the scattered X-rays component is subtracted 
from the image data in the second step, thereby to obtain the image data 
containing only the primary X-rays. Using such image data, an X-ray image 
of the object is displayed by the monitor. Therefore, the displayed X-ray 
image is free from blur and good in contrast and sharpness. 
The above-mentioned embodiment may be modified into an arrangement as shown 
in FIG. 11 in which a logarithm-converting circuit means 58 is 
additionally provided following the subtracting circuit means 51. The 
logarithm-converting circuit means 58 converts the output signal Ip(x, y) 
from the subtracting circuit means 51 into a natural logarithmic value. 
The provision of the logarithm-converting processing circuit means enables 
one to obtain an attenuation quantity of X-rays 
##EQU9## 
in the following way. The X-ray intensity Ip(x, y) of only the primary 
X-rays can be expressed by 
##EQU10## 
In the above equation, Io is the intensity of X-rays incident on the X-ray 
detector 48, which depends on the conditions which the X-ray tube has when 
there is no object. And .mu.i(x, y) is an X-ray absorbing coefficient of 
tissues (i) of the object 46. Further di(x, y) describes the thickness of 
the object 46 as a function of x and y, which define the position on the 
tissue (i) of the object 46 corresponding to a position on the X-ray 
detector 48 hit with an X-ray beam. If the exponential expression (12) is 
expressed in the form of the natural logarithm, we have the following 
logarithmic expression (13). 
##EQU11## 
In the above equation, ln Io is a known value. Therefore, the attenuation 
quantity of the X-rays 
##EQU12## 
can be obtained using the equation (13). 
As seen from the foregoing description, by subtracting the scattered X-ray 
component as previously defined from the X-rays transmitted through the 
object, a formed transmitted X-ray image depends solely on the primary 
X-rays. Therefore, the following useful effects can be attained: 
(1) to improve contrast and resolution of the image of the patient. 
(2) to exactly obtain an X-ray attenuation quantity by logarithmically 
converting the image data. 
The effect (2) above is more effective particularly for the X-ray diagnosis 
carried out using an X-ray constrast medium. Specifically, in handling a 
subtraction image between the images before and after the contrast medium 
is administered, if the subtraction is performed after both of these 
images are logarithmically converted, it is possible to exactly obtain the 
product .DELTA..mu..multidot.d of a changed amount of .DELTA..mu. of an 
X-ray coefficient, which is caused by the contrast medium and the 
thickness "d" of the tissue under X-ray radiation. 
A second embodiment of an X-ray diagnostic apparatus according to the 
present invention, which is based on the second application principle, 
will be described referring to FIG. 12. In the figure, X-rays radiated 
from an X-ray generator 11 transmit through the acryl plate 32 and the 
object 21 under examination, and enter into an X-ray detector 22. The 
X-ray detector 22 detects the intensity of the transmitted X-rays. An A/D 
converter is used for converting to digital signal and the analog image 
signals output by the X-ray detector 22. An image processor 24 includes a 
memory for storing image data, and a computing means for removing the 
scattered X-ray component from the image data, as will subsequently be 
described. Reference numeral 25 designates a monitor for displaying the 
image data processed by the image processor 24. 
The acryl plate 32 has a structure as shown in FIG. 4. For X-ray 
irradiation to determine the scattered X-rays intensity or to obtain the 
parameters C, .sigma..sub.x and .sigma..sub.y, the acryl plate 32 is 
placed in the X-ray projection area. After these parameters are 
determined, it is retracted from the X-ray projection area. The movement 
of the acryl plate 32 is made by a drive means (not shown). 
The image processor 24 contains a scattered X-ray intensity computing 
section 24A and a compensating section 24B. 
In FIG. 13 illustrating the image processor, a first memory 24A-1 
temporarily stores, of the detected signals output from the X-ray detector 
22, a detected signal Isc.sup.ob (0, 0) representing X-ray intensity at 
positions (x, 0) and (0, y) on the lines L2 and L3 on the acryl plate 32. 
A computing means, for example, a low-pass filtering means 24A-2, cuts off 
the high frequency component in the detected signal. A normalizing means 
24A-3 noramlizes the detected signals Isc(x, 0) and Isc(0, y) after the 
high frequency component is cut off by a detected signal Isc(0, 0) 
representing an X-ray intensity at a center position (a cross point of the 
lines L2 and L3). A second memory 24A-4 stores functions F1.sup.cal (x) 
and F2.sup.cal (y) previously calculated for various parameters 
.sigma..sub.x and .sigma..sub.y, and the normalized functions Fa.sup.cal 
(x)/F1.sup.cal (0) and F2.sup.cal (y)/F2.sup.cal (0). A .sigma. parameter 
determining means 24A-5 compares F1.sup.ob (x)/F1.sup.ob (0) and F2.sup.ob 
(x)/F2.sup.ob (0) output from the normalizing means 24A-3, F1.sup.cal 
(x)/F1.sup.cal (0) F2.sup.cal (y)/F2.sup.cal (0) output from the second 
memory 24A-4, and determines parameters .sigma..sub.x and .sigma..sub.y to 
provide the most approximate F1.sup.cal (x)/F1.sup.cal (0) and F2.sup.cal 
(y)/F2.sup.cal (0). A C1, C2 determining means 24A-6 determines C1 and C2 
by the equations (7) and (10), using the parameters .sigma..sub.x and 
.sigma..sub.y. A C determining means 24A-7 determines C by the equation 
(6), using the parameters C1 and C2 and the detected signal Isc.sup.ob (0, 
0) stored in the first memory 24A-1. A scattered X-ray intensity-computing 
means 24A-8 computes the scattered X-ray intensity Isc(x, y) at a proper 
position by the equation 4A, using the parameters .sigma..sub.x, 
.sigma..sub.y and C, and stores the computed one. The compensating means 
24B contains a subtractor for subtracting, by the equation (5A), a signal 
Isc(x, y) representing the scattered X-ray intensity, obtained by the 
scattered X-ray intensity-computing means 24A-8, from a detected signal 
Im(x, y), produced from the X-ray detector 22, when the X-rays transmitted 
through the object 21 are detected by the X-ray detector with retraction 
of the acryl plate 32 from the X-ray projection area. An X-ray aperture 27 
(FIG. 12), defining the X-ray projection area, is disposed close to the 
X-ray generator 11. 
The operation of the X-ray diagnostic apparatus thus arranged will be 
described. 
To obtain the scattered X-ray intensity, the acryl plate 32 is set in the 
X-ray projection area and the X-ray tube projects X-rays into the object 
21. The X-rays transmitted through the object 21 are detected by the X-ray 
detector 22. A detected signal from the X-ray detector 22 is converted to 
digital data by the A/D converter 23 and is supplied to the image 
processor 24. Of those detected signals supplied to the image processor 
24, the detected signals Isc(x, 0) and Isc(0, y) of the X-rays on the 
lines L2 and L3 on the acryl plate 32 are stored in the first memory 
24A-1. As shown in FIG. 14(A), the detected signal Isc(x, 0) contains high 
frequency noise. The detected signal is passed through the low-pass 
filtering means 24A-2 where its high frequency component is cut off. The 
detected signal Isc(x, o), after passing the low-pass filter, is as shown 
in FIG. 14(B). Then, the detected signal Isc(x, 0) and the detected signal 
Isc(0, 0) at the center position are applied to the normalizing means 
24A-3 where the detected signal Isc(x, 0) is normalized by the detected 
signal Isc(0, 0). Here, the normalizing means performs the operation of 
Isc(x, 0)/Isc(0, 0). The result of the normalizing represents F1.sup.ob 
(x)/F1.sup.ob (0) according to the equation (6). The second memory 24A-4 
stores F1.sup.cal (x)/F1.sup.cal (0) (see FIG. 14C) and F1(x) is 
normalized by F1.sup.cal (x) and F1.sup.cal (0) obtaining by using various 
values of .sigma..sub.x according to the equation (8). The .sigma. 
determining means 24A-5 compares F1.sup.ob (x)/F1.sup.ob (0) output from 
the normalizing means 24A-3 and the various normalized functions 
F1.sup.cal (x)/F1.sup.cal (0) output from the second memory 24A-4, and 
determines .sigma..sub.x to provide the most approximate F1.sup.cal 
(x)/F1.sup.cal (0). The .sigma..sub.x determined by the .sigma. 
determining means 24A-5 is output to the C1/C2 determining means 24A-6. 
This means 24A-6 performs the operation of the equation (10) to determine 
C2 using .sigma..sub.x, and sends the determined C2 to the C determining 
means 24A-7. The C determines means 24A-7 determines C by the equation (6) 
using the value of C2, the function F1.sup.cal (0) stored in the second 
memory 24A-4, and the actually measured data Isc.sup.ob (0, 0) stored in 
the first memory 24A-1, and sends the C to the scattered X-ray 
intensity-computing means 24A-8. Similarly, .sigma..sub.y is determined by 
the detected signal Isc(0, y). The scattered X-ray intensity-computing 
means 24A-8 performs the operation of the equation (4A), using thus 
determined .sigma..sub.x, .sigma..sub.y and C to obtain a signal Isc(x, y) 
representing a scattered X-ray intensity at a proper position and stores 
this signal. 
Following the computing of the scattered X-ray intensity, the acryl plate 
32 is retracted from the X-ray projection area, and then the object 21 is 
irradiated with X-rays. The X-rays transmitted through the object 21 are 
detected by the X-ray detector 22. A signal output from the X-ray detector 
22 is then applied to the A/D converter 23 where it is converted to 
digital data. The converted digital data is applied, as the output signal 
Im(x, y), to the compensating processing section 24B in the image 
processor 24. In the compensating section 24B, the equation (5A) is 
performed to have the image signal Ip(x, y) containing only the primary 
X-ray component. 
In this way, the scattered X-ray component is removed from the transmitted 
X-ray data containing both the primary X-ray component and the scattered 
X-ray component, and the X-ray image based only on the primary X-ray 
component is visually displayed by the monitor. 
When a contrast phantom 26 is radiographed by using the x-ray diagnostic 
apparatus thus arranged, as shown in FIG. 15, the X-ray detector 22 
produces a detected signal Im(x, y) containing the scattered X-ray 
component, as shown in FIG. 16(A). A signal Isc(x, y) representing a 
scattered X-ray intensity distributed as shown in FIG. 16(B), which is 
produced from the scattered X-ray intensity-computing section 24A, is 
subtracted from the detected signal Im(x, y). Through this subtraction, a 
signal Ip(x, y) containing only the primary X-rays can be obtained as 
shown in FIG. 16(C). Therefore, the X-ray image displayed by the monitor 
25 is improved in contrast and sharpness, and is free from blur. 
Although not shown, the image processor 24 contains a controller containing 
a CPU for providing control signals to extract the detected signals Isc(x, 
0) and Isc(0, y) of the X-rays along the lines L2 and L3 and stores them 
in the related memory, and various control signals necessary for operating 
the scattered X-ray intensity-computing section 24A and the compensating 
processing section 24B. 
Having described specific embodiments, it is believed obvious that 
modification and variation of our invention are possible in light of the 
above teachings. 
For example, the scattered X-ray intensity-computing section 24A in the 
image processing unit 24 in the second embodiment may be replaced by a 
means in which .sigma. is determined by the least square method. For the 
function g(x, y), an exponential function corresponding relatively well to 
the results of experiment may be used in place of the Gauss function.