JPS6147535B2 - - Google Patents

Description

[Detailed description of the invention]

The present invention relates to an apparatus for reconstructing an X-ray tomographic image, and more particularly to an apparatus capable of reconstructing an X-ray tomographic image by performing desired data processing. In general, in the industrial and medical world, X-ray inspections are widely used to determine the internal structure of products and the human body.
Linear computerized tomography equipment (abbreviated as "CT") is being actively developed. By the way, in this type of conventional X-ray computerized tomography apparatus, first, as shown by the solid line in FIG. The X-ray projection distribution is measured by the X-ray detector c, and then, as shown by the chain line in FIG. 1, the X-ray generator a and the X-ray detector c are rotated by a predetermined angle α (for example, 1°). By irradiating the X-ray beam toward the X-ray inspection object b again from this position, the angle α is
After measuring the X-ray projection distribution when the distribution is moved by Computation processing is performed using the method, and then a tomographic image of the X-ray object b is reconstructed based on the processing results. However, such a conventional X-ray computer tomography apparatus requires a large number of X-ray projection distributions in order to reconstruct a tomographic image of the subject b, which causes the following problems. (1) Since it takes a long time to collect data (several seconds to several minutes), for moving object b,
The tomographic image cannot be reconstructed. (2) Since the amount of X-ray exposure increases, if the subject b is a living body such as a human body, there is a risk of adverse effects on the living body. (3) Since a large amount of data is required for processing, data processing becomes extremely complicated, and the equipment for processing this data is also large and expensive. The present invention aims to solve these problems, and allows highly accurate reconstructed images by simply using a single X-ray projection distribution as data for reconstructing a tomographic image of a subject. It is an object of the present invention to provide a single X-ray projection type tomographic image reconstruction device that can obtain a single X-ray projection type tomographic image reconstruction device. For this reason, the single X-ray projection type tomographic image reconstruction apparatus of the present invention uses a virtual tomographic plane S that includes the tomographic plane of the X-ray subject B and has m pixels vertically and n pixels horizontally. and , from one end to the other end of a single X-ray projection distribution D obtained by irradiating the plurality of unit X-ray beams from the X-ray generator A toward the X-ray inspection object B. N (=m×n) X-ray densities
X-ray concentration measuring device E that measures d ₁ to d _no .
The above N X-ray densities obtained by the linear densitometer E
Equipped with a data selection device DS that divides d ₁ to _{d no} into a plurality of measurement value groups consisting of a plurality of X-ray densities,
The virtual tomographic plane S is divided into a plurality of pixel groups consisting of a plurality of pixels based on a plurality of X-ray densities in each of the measurement value groups, and the
In order to calculate the linear absorption coefficient μ, the X-ray density d ₁ obtained by the unit X-ray beam passing only through basic pixel 1
Assuming that the X-ray absorption coefficient μ of the same basic pixel 1 is ₁ , for the pixels up to mnth, subtract the X-ray density d^ _k-1 (k^ = 2, 3, ... mn) from the X-ray density d^ _k . Accordingly, a projective transformation device F is provided, which includes an arithmetic device that sequentially calculates the X-ray absorption coefficient μ^ _k of each pixel, and a storage device G that can store the calculated X-ray absorption coefficient, and also performs the same projective transformation. In response to a signal from the device F, the X-ray absorption coefficient μ ₁ is applied to each of the N pixels 1 to mn constituting the virtual tomographic plane S.
A reconstruction device H that reconstructs the tomographic image of the X-ray subject B by applying ~μ _no , and displays the tomographic image of the X-ray subject B based on the output from the reconstruction device H. It is characterized by being provided with a display device J. Embodiments of the present invention will be described below with reference to the drawings. FIGS. 2 to 4 show a single X-ray projection type tomographic image reconstruction apparatus as the first embodiment, and FIG. FIG. 3 is a schematic diagram showing the distribution measuring means, FIG. 3 is a system configuration diagram thereof, and FIG. 4 is a schematic diagram for explaining its operation. As shown in FIG. 3 (a diagram corresponding to the claims), an X-ray inspection object B is positioned between an X-ray generator A and an X-ray detector C. This X-ray inspection object B can be irradiated with X-rays from a predetermined direction, and as a result, the X-rays that have passed through the X-ray inspection object B form a single X-ray projection distribution D (Fig. ). Note that, as shown in FIG. ), a plurality of unit X-ray beams are irradiated toward the X-ray inspection object B from a predetermined direction θ so that the unit X-ray beams pass through at least one lower left corner of each pixel 1 to mn. It is designed to generate X-rays with radiation quality (penetrating power) and dose that are compatible with the X-ray inspection object B, which is the object of inspection, and the wavelength of the generated X-rays depends on the applied voltage. In addition, the quality of X-rays is determined by their wavelength. The applied voltage varies depending on the application, ranging from 50,000 volts to 120,000 volts when used for medical diagnosis, and from 100,000 volts to 300,000 volts when used for non-destructive testing. Ru. X-ray inspection object B refers to an object that is irradiated with X-rays and whose transmitted dose distribution (X-ray projection distribution) is measured to reconstruct an image in a desired tomographic plane. For example, in the case of medical diagnosis, it is a human body (generally a living body), and in the case of non-destructive testing, it is a so-called industrial product. Further, as the X-ray detector C, an X-ray film, a scintillation detector, a semiconductor detector, a xenon gas detector, or the like is used. By the way, the single X-ray projection distribution data D obtained by the X-ray detector C is measured by an X-ray projection distribution measuring means (X
The above X-ray projection distribution D is measured by the line density measuring device) E.
Equal intervals w from one end to the other
Each value dk at mn positions k separated by (these measured values d ₁ to d _no can be divided into n sets of measured value groups, if m measured values are one measured value group) ) is now being measured (see Figure 4). When the X-ray detector C is an X-ray film, the X-ray projection distribution measuring means E may be used to measure a plurality of X-ray density distributions (so-called X-ray photographs) obtained as shading of the degree of blackening on the film. A microdensitometer is used that can measure the value of . In this way, from one end of the X-ray projection distribution D to the other end, there are multiple positions k spaced at equal intervals w.
In addition to the above-mentioned example, as a means for measuring each value dk in In the case of a scintillation detector, one scintillation detector may be combined with a mechanism that can move it from one end of the X-ray projection distribution D to the other, or multiple scintillation detectors may be combined to The arrangement may be performed over the entire width of the projection distribution D. In addition, if the X-ray detector C is a semiconductor detector that also receives X-rays transmitted through the X-ray inspection object B and outputs a signal corresponding to the X-ray concentration, the above-mentioned scintillation detection can be used as the above-mentioned means. As in the case of detectors, one semiconductor detector and a mechanism for moving it may be combined, or a large number of semiconductor detectors may be arranged. Furthermore, when the X-ray detector C is a xenon gas detector, the above means includes one xenon gas detector and a mechanism for moving it, as in the case of the scintillation detector or semiconductor detector described above. Combinations and the arrangement of multiple xenon gas detectors may be used. The signals obtained in this way by the microdensitometer, scintillation detector, and semiconductor detector are analog signals, so in order to input them to a digital computer, this signal is converted to an analog-to-digital converter E' ( Below, “A/D
called a converter. ) is converted into a digital signal and then stored on a disk or the like. Next, the data output (digital signal) from this A/D converter is sequentially input to the projective conversion device F via the data selection device DS. The data selection device DS selects n measurement value groups consisting of m X-ray densities for N (=mn) X-ray densities.
A group is selected. Further, the projective conversion device F converts the virtual tomographic plane S into n pixels consisting of m pixels based on the digital data output (one-dimensional data) inputted from the A/D converter via the data selection device DS. group
_{Q 1 , Q 2 ,} . Unit X that passes only through basic pixel 1 to calculate and output
Assuming that the X-ray density d ₁ obtained by the ray beam is the X-ray absorption coefficient μ ₁ of the same basic pixel 1, for the pixels up to mnth, the X-ray density d^ _k to the X-ray density d^ _k-1 (k^
=
2, 3, ..., mn) for sequentially calculating the X-ray absorption coefficient μ^ _k of each pixel, and a storage device G for storing the calculated X-ray absorption coefficient (as will be described later). A dimensional internal structure storage device serves as this storage device.) A specific example thereof is a digital computer containing a desired program. Here, m measured values d _k (k=1, 2,
...mn) as a component of a virtual tomographic plane S (this plane S consists of n groups of pixels) including the tomographic plane of the X-ray subject B. A method for determining each X-ray absorption coefficient μ _k (two-dimensional data) of each pixel will be explained. As shown in FIG. 4, assuming that there are mn unit X-ray beams, the first unit X-ray beam reaches the X-ray detector C through the lower left corner of the first pixel, Similarly, by determining the position of each pixel so that the beam reaches the X-ray detector C through the lower left corner of the m-th pixel, until the m-th unit X-ray beam is reached. 1st
Determine the pixel group Q ₁ of . Next, the position of each pixel is determined so that the m+1 to 2m-th unit X-ray beams reach the X-ray detector C through the lower left corners of the m+1 to 2m-th pixels, respectively. 2 pixel group
Determine Q ₂ . Similarly, by determining the position of each pixel so that m unit X-ray beams pass through the lower left corner of each m pixel and reach the X-ray detector C, Determine the n-th pixel group Q ₃ to Q _o . By determining each pixel group Q ₁ to Q _o in this manner, the virtual tomographic plane S is also determined. in this case,
The unit X-ray beams are assumed to have the same mutual spacing w. In this way, the first unit X-ray beam passes only through the first pixel 1 (this pixel 1 is called the basic pixel), so the X-ray density obtained by this first unit X-ray beam is d ₁ and 1st pixel 1
If the relationship with the X-ray absorption coefficient μ ₁ is known,
This X-ray absorption coefficient _μ1 can be uniquely determined by a calculation device. Furthermore, the second unit X-ray beam is
Since the second unit X-ray beam passes through only the second pixel, the X-ray density d ₂ obtained by this second unit X-ray beam is based on both the X-ray absorption coefficients μ ₁ and μ ₂ of the first and second pixels. However, as mentioned above, since the X-ray absorption coefficient μ ₁ in the first pixel is already known, the X-ray absorption coefficient μ ₂ in the second pixel is also found by the calculation device. It is. By repeating almost the same operations one after another, it is possible to obtain up to the X-ray absorption coefficient of the m-th pixel, and thereby the first pixel group Q ₁ consisting of m pixels is obtained. Each pixel 1 in
The X-ray absorption coefficients μ ₁ to μ _n of ~m can be determined. The X-ray absorption coefficients μ ₁ to _{μ n} obtained by calculation in this manner are transferred to a memory G as a storage device, which will be described later. At this time, the m X-ray densities necessary for the calculation are selected by the data selection device DS. Furthermore, the m+1th unit X-ray beam is
Since only the m+1th pixel is passed through, this
The X-ray density d _n + ₁ obtained by the +1st unit X-ray beam is
It has each information of linear absorption coefficient μ ₂ to _{μ n+1} , but as mentioned above, X of the 2nd+m to 1st pixel
Since the linear absorption coefficients μ ₂ to _{μ n} are already known by calculating the X-ray absorption coefficients μ ₁ to μ n of each pixel 1 to m in the first pixel group Q ₁ , the X-ray absorption coefficients μ 2 to μ _n stored in the memory G By appropriately calling the information on the linear absorption coefficients μ ₁ to _{μ n} , the X-ray absorption coefficient μ _n+ 1 of the pixel m+1 constituting the second pixel group Q ₂ can be determined. Thereafter, by calling up information from the memory G as appropriate in substantially the same manner, the X-ray absorption coefficient μ _n of pixels m+2 to 2m constituting the second pixel group Q _2
+2 to μ _2n can be obtained. Furthermore, the X-ray absorption coefficients μ _2n+1 to μ _no of each pixel in the subsequent third to nth pixel groups Q ₃ to Q _o consisting of m pixels can be determined in substantially the same manner. . At this time, this information is transferred to the memory G each time m X-ray absorption coefficients are determined. In this way, the X-ray absorption coefficient can be found sequentially for each pixel group, but in addition to the qualitative explanation of the projective transformation described above, we will explain this projective transformation method in detail quantitatively using mathematical formulas, etc. Then, it becomes as follows. First, a virtual tomographic plane S as a tomographic image reconstruction plane of the X-ray subject B is defined as a column partial plane of the tomographic image reconstruction plane consisting of m pixels divided into small sections, as shown in FIG. Consider n groups of pixels, and let the center of the plane S be x-. Take it as the origin of the y coordinate. Note that it is assumed that each pixel has the same X-ray absorption coefficient, and for convenience of explanation, m and n are assumed to be even numbers, and the size of one pixel is assumed to be a square of △×△. In addition, the X-ray beam passing through this virtual plane S shall be irradiated in parallel from the direction θ that satisfies θ = tan ^-1 m, and the beam diameter of each unit X-ray beam shall be sufficiently small compared to each pixel. do. Now, each pixel 1 to m in the first pixel group _Q1
If m projected densities d ₁ to d _n are used to determine the X-ray absorption coefficients μ ₁ to μ _n of , the following equation is obtained. 〓〓＝〓 _Q１ ……(1) However 〓=(μ ₁ , μ ₂ , μ ₃ ..., μ _n ) ^T 〓 _Q1 = (d ₁ , d ₂ , d ₃ ..., d _n ) ^T where L is an m×m square matrix, a is the length of the X-ray beam passing through the pixel, a=△√1+
cot ² θ, where the symbol _T represents transposition. Solving this equation (1) yields the following equation. μ ₁ = p ₁ }…(3) μ _k = p _k −p _k-1 (k=2, 3,…
..., m) However, p = (p ₁ , p ₂ , p ₃ , ... p _n ) ^T = 〓 _Q1 /a. Here, the reason why the length a of the X-ray beam that passes through one pixel is used to determine μ ₁ and μ _k is to determine the accurate X-ray absorption coefficient of each pixel, and therefore the coefficient matrix Even if a is not included in L, 1
Two-dimensional data μ _k can be obtained from dimensional data d _k . In this case, the solution is as follows. μ ₁ = d ₁ } …(3)′ μ _k = d _k −d _k-1 (k=2, 3,…, m) By the way, as mentioned above, the absorption coefficient within one pixel is constant. If this assumption is established and there is no measurement error, the X-ray absorption coefficients μ ₁ to μ _n of each pixel 1 to m in the first pixel group Q ₁ are calculated from equation (3), but in reality It is difficult for this assumption to hold true for the X-ray inspection object B, and measurement errors are generally included, so even when the above assumption slightly collapses or a small amount of measurement error is introduced, Note that it is difficult to expect that the two-dimensional data μ ₁ to _{μ n} calculated using equation (3) will give good results. Therefore, using mathematical programming, the non-negative correction value r ₁ ,
Introduce r ₂ , r ₃ ..., r _n . Furthermore, since the absorption coefficient of an X-ray beam transmitted through an object is generally non-negative and does not exceed a certain positive upper limit U, equation (1) can be expressed as follows by adding these constraints.

[Table] Furthermore, the non-negative slack variables s ₁ , s ₂ , s ₃ , ...,
When s _3n is introduced, the simultaneous linear inequalities in equation (4) become the simultaneous linear equations in the following equation.

[Table] Objective function based on the constraint condition equation (5) above. Using mathematical programming, we find a solution that minimizes
After a finite number of calculations, the optimal 2
Dimensional data μ ₁ to _{μ n} are calculated. In the above example, the two-dimensional data μ ₁ to μ _n were calculated based on the objective function that minimizes the sum of the absolute values of the correction values in the constraint condition expression. There is also a method of minimizing the objective function in equation (8).

[Table] By solving the above equation, tomographic two-dimensional data μ ₁ to μ _n are calculated under the condition that the maximum correction value of the absolute value in the constraint condition equation is the minimum. Furthermore, the objective function is It is also possible to carry out the method by minimizing the objective function F=r ² . . . (10) under the constraint expression (7). Two-dimensional data μ ₁ ~μ obtained in this way
_n is transferred to a three-dimensional internal structure storage device (memory) G. Next, each pixel m+1~ in the second pixel group _Q2
In order to find the X-ray absorption coefficient μ _n+1 ~ μ _2n of 2m, m projected densities d _n+1 ~ d _2n and the X
If the linear absorption coefficients are employed, m X-ray absorption coefficients μ _n+1 to μ _2n can be determined as shown in equation (11) using an arithmetic device in substantially the same manner as the above-mentioned method. μ _k = p _k −p _k-1 + μ _kn …(11) (k=m+1, m+2, …2m) However, 〓=(p _n+1 , p _n+2 , p _n+3 , … p _2n ) ^T = 〓 _Q2 /a In this case as well, when calculating μ _k , it is not necessary to include the information about the length a of the X-ray beam that passes through one pixel, and the solution in this case is as follows. become. μ _k = d _k −d _k-1 + μ _kn ……(11)′ Furthermore, in the same manner as when calculating the X-ray absorption coefficient of each pixel in the first pixel group Q ₁ described above, mathematical programming is applied as appropriate. The optimal X-ray absorption coefficient (two-dimensional tomographic data) μ _{n+1 ~ μ 2n} is determined using
is transferred to memory G. After that, repeat almost the same operation to create the subsequent third
The X-ray absorption coefficient of each pixel in the ~n pixel groups _Q3 to _Qo is determined. Here, μ _k for 2m+1st and subsequent pixels
can also be expressed as follows. μ _k = p _k − p _k-1 + μ _kn ……(11)″ Or μ _k = d _k − d _k-1 + μ _kn ……(11)″ By the way, a three-dimensional internal structure storage device that also serves as a storage device As described above, G receives a set of m signals from the projective transformation device F and assembles each pixel group Q ₁ to Q _o in accordance with their mutual order, that is, the first to
This refers to a memory that can store signals from the projective transformation device F in order to assemble the n-th pixel group from the left in their ordinal order, and further refers to a memory that calculates three-dimensional internal structure data of the X-ray inspection object B. By the way, from the projective transformation device F, m pieces are sequentially n
Two-dimensional data μ _k (k=1, 2,
3, . 2 for other cross sections.
Since dimensional data μ _k ′ can be easily obtained almost in the same way as in the previous case, it is possible to obtain two-dimensional data μ _k , μ′ _k , μ″ _k , etc. for several different cross sections. This makes it possible to store the three-dimensional internal structure of the X-ray inspected object B, but in order to construct a complete three-dimensional internal structure, interpolation between each cross-sectional data is required. This memory G is also used as a memory device that has a calculation function for this purpose.An arbitrary tomographic image construction device (reconstruction device) H is connected to this memory G, and an arbitrary tomographic image construction device (reconstruction device) H is connected to this memory G. The apparatus H selects and extracts two-dimensional data regarding a specified arbitrary cross section of the X-ray subject B from the three-dimensional internal structure data of the X-ray subject B stored in the memory G, and generates a tomographic image. Note that the arbitrary cross section here refers to a horizontal, vertical, or diagonal cross section of the X-ray subject B. In this way, this arbitrary tomographic image constructing apparatus H The two-dimensional data of any tomographic image thus obtained has been consistently calculated mathematically and without contradiction based on the X-ray projection distribution obtained by the X-ray projection distribution measuring means E. is a suitable digital-to-analog converter F' (hereinafter referred to as "D/A converter").
If the tomographic image is sent to an arbitrary tomographic image display device J for display via the Therefore, there is no guarantee that a suitable image will be obtained. Therefore, in order to correct the data from this optional tomographic image construction device H, this data is sent to the optional tomographic image quality improvement device I. This arbitrary tomographic image quality improvement device I improves the image quality by removing noise, smoothing, and emphasizing the sharpness of arbitrary tomographic image data sent from the arbitrary tomographic image construction device H. , a digital filter is used to remove noise,
A smoothing circuit is used for smoothing, and a differentiation circuit is used for emphasizing sharpness. The signal whose image quality has been improved in this way is D/A
It is sent to the arbitrary tomographic image display device J via the converter F'. This optional tomographic image display device J receives the signal sent from the optional tomographic image quality improvement device I and displays X on a color or black and white cathode ray tube (braun tube) monitor.
A device that displays an arbitrary tomographic image of the object B to be inspected as a visible image, and generally a cathode ray tube is used as described above. With the above configuration, in order to reconstruct a tomographic image of the X-ray subject B, first the X-ray generator A irradiates the X-ray subject B with X-rays from a predetermined direction, and then the X-ray detector C Single X-ray projection distribution D obtained with
Using the X-ray projection distribution measuring means E, each value d _k at mn positions k spaced at equal mutual intervals w from one end of the X-ray projection distribution D is calculated from the other end of the X-ray projection distribution D. By measuring until 1
Dimensional data d _k can be seen. Next, these one-dimensional data _dk are appropriately analog-digital converted, and then the X-ray absorption coefficient of each pixel is determined for each m pixel group using the method described above in the projective conversion device F. Furthermore, by sequentially transferring the X-ray absorption coefficients for each pixel group to the memory G, each X-ray absorption coefficient μ _k (two-dimensional data) of mn pixels on the virtual tomographic plane S is obtained, and these X Linear absorption coefficients are stored in memory G according to a predetermined order. Thereafter, these two-dimensional data μ _k are transmitted from the memory G to an arbitrary tomographic image constructing device H, an arbitrary tomographic image quality improvement device I, and a D/A converter F′, and then to an arbitrary tomographic image display device J. The line is reconstructed and displayed as a tomographic image of the object B to be inspected. FIG. 5 is a schematic diagram for explaining the operation of a single X-ray projection type tomographic image reconstruction apparatus as a second embodiment of the present invention. In the case of this second embodiment, the X-ray generator A and
In the case where the object B to be inspected is close to the X-ray beam, that is, the X-ray beam is a Fan beam, in this case, if the tomographic two-dimensional data is composed of pixels with modified polar coordinates as shown in Fig. 5. , Equation (1) is obtained, and the above theory can be applied as is. In FIG. 5, the X-ray source is the origin 0, the inner radius of the polar coordinate pixel is R _I , the outer radius R _O , and the pixels are composed of m×n. The pixels are numbered as shown in FIG. 5, and a straight line passing through OO' is used as the standard for angle. The radii R ₁ , R ₂ and angles ₁ , ₂ , ₃ , and ₄ are expressed as follows.

[Table] Now, let us draw a straight line passing through points R ₁ , ₁ and R ₂ , ₃ .
If L ₂ is a straight line passing through L 1 , points R ₁ , ₂ and R ₂ , ₄ , pixel k has an area _surrounded by radii R ₁ , R ₂ and two straight lines L ₁ , L ₂ . However, θ' is the apex angle far inside (radius R ₁ ) of the polar coordinate pixel. In this way, even when the X-ray beam is a Fan beam, two-dimensional data is obtained for each column partial plane of the tomographic image from m one-dimensional data in the same manner as in the case of a parallel beam as in the first embodiment. As a result, a tomographic image of the X-ray object B can be reconstructed from a single X-ray projection distribution D. Note that, instead of each pixel group consisting of m pixels forming a column partial plane of the tomographic image of the X-ray subject B as in each of the above-mentioned embodiments,
1) A pixel group consisting of 1) pixels may be appropriately combined to constitute a column partial plane of the tomographic image. In addition, in order for each of the above pixel groups to constitute a row partial plane of the above tomographic image, this pixel group is divided into m−m(n−1).
In addition to the pixel group consisting of n to (m-1)n pixels, in order to constitute the row partial plane of the above tomographic image,
It may also be composed of a single pixel. Furthermore, it can also be configured with fewer pixels than each pixel group m, n. In other words, one pixel group can be made up of any number of pixels as long as it is less than mn, but the number of pixels that make up one pixel group depends on the processing for data processing. It is desirable to determine the optimal number based on the capacity of the device and the data processing time. As described in detail above, the single X-ray projection type tomographic image reconstruction apparatus of the present invention provides the following effects and advantages. (1) Since the tomographic image of the X-ray subject B can be reconstructed from a single X-ray projection distribution D, the data collection time is extremely short (30 milliseconds or less), and this makes it possible to The tomographic image of the X-ray object B (for example, the heart) can also be clearly reconstructed. (2) Also, since the amount of X-ray exposure is extremely small (several tenths to hundreds of times compared to conventional methods),
Even when the X-ray inspection subject B is a living body, there is no adverse effect. (3) For each pixel group, calculate the
Since the linear absorption coefficient can be determined, the amount of data required for one processing is greatly reduced, which greatly simplifies data processing, which in turn leads to the miniaturization and cost reduction of data processing equipment. Fully achievable.

[Brief explanation of the drawing]

FIG. 1 is a schematic diagram for explaining a conventional X-ray tomographic image reconstruction means, and FIGS. 2 to 4 are a single X-ray projection type tomographic image reconstruction apparatus as a first embodiment of the present invention. Fig. 2 is a schematic diagram showing the means for measuring the X-ray projection distribution, Fig. 3 is a system configuration diagram, Fig. 4 is a schematic diagram for explaining its operation, and Fig. 5 is a schematic diagram showing the means for measuring the X-ray projection distribution. FIG. 2 is a schematic diagram for explaining the operation of a single X-ray projection type tomographic image reconstruction apparatus as a second embodiment of the present invention. A...X-ray generator, B...X-ray inspected object, C
...X-ray detector, D...X-ray projection distribution, DS...
Data selection device, E...X-ray projection distribution measuring means (X-ray density measurement device), E'...A/D converter, F
...Projective conversion device, F'...D/A converter, G...
...Three-dimensional internal structure storage device (memory) that also serves as a storage device, H...Arbitrary tomographic image composition device (reconstruction device), I...Arbitrary tomographic image quality improvement device, J...Arbitrary tomographic image display device, Q ₁ , Q ₂ , Q ₃ , Q _o ... Pixel group, S ... Virtual tomographic plane.

Claims (1)

[Claims] 1. Regarding a virtual tomographic plane S that includes the tomographic plane of the X-ray inspected object B and has m pixels vertically and n pixels horizontally, the unit X-ray beam is located at the corner of each pixel 1 to mn. The above X should pass through at least one part of the
A plurality of units X from a predetermined direction toward the object B to be inspected
An X-ray generator A that irradiates a ray beam, and a single X-ray obtained by irradiating the plurality of unit X-ray beams from the X-ray generator A toward the X-ray inspection object B. N (=m×n) X-ray densities d ₁ to d from one end to the other end of the projection distribution D
X-ray concentration _measuring device E that measures _no .
a data selection device DS that divides _{the no} . In order to divide into multiple pixel groups and calculate the X-ray absorption coefficient μ of each pixel constituting each pixel group, the X-ray density d ₁ obtained by a unit X-ray beam passing only through basic pixel 1 is calculated from the same basic pixel. The X-ray absorption coefficient μ of 1 _{is 1} , and for the mnth pixels, the X-ray density d _k to the X-ray density d _k-1
(k = 2, 3, ... mn) from an arithmetic device that sequentially calculates the X-ray absorption coefficient μ _k of each pixel, and a storage device G that can store the calculated X-ray absorption coefficient. A projective transformation device F is provided, and upon receiving a signal from the projection transformation device F, the X-ray absorption coefficients μ ₁ to μ are placed at the positions of each of the N pixels 1 to mn constituting the virtual tomographic plane S. _no
a reconstruction device H that reconstructs a tomographic image of the X-ray subject B by applying the above, and a display device J that displays the tomographic image of the X-ray subject B based on the output from the reconstruction device H. A single X-ray projection type tomographic image reconstruction device, characterized in that: 2. The X-ray concentration measuring device E includes a microdensitometer capable of measuring multiple X-ray concentration values on the X-ray concentration distribution obtained on the X-ray film, and an analogue from this microdensitometer. A single X-ray projection type tomographic image reconstruction apparatus according to claim 1, comprising an analog-to-digital converter capable of converting a signal into a digital signal. 3. The X-ray concentration measuring device E includes a scintillation detector with a moving mechanism that receives X-rays passing through the X-ray inspection object B and outputs a signal corresponding to the X-ray concentration, and scintillation detection of the same. 2. The single X-ray projection type tomographic image reconstruction apparatus according to claim 1, comprising an analog-to-digital converter capable of converting an analog signal from a device into a digital signal. 4 The X-ray concentration measuring device E includes a number of scintillation detectors that receive X-rays transmitted through the X-ray inspection object B and output signals corresponding to the X-ray concentration, and these scintillation detectors. 2. The single X-ray projection type tomographic image reconstruction apparatus according to claim 1, comprising an analog-to-digital converter capable of converting an analog signal from a device into a digital signal. 5 The X-ray concentration measuring device E includes a semiconductor detector with a moving mechanism that receives X-rays passing through the X-ray inspection object B and outputs a signal corresponding to the X-ray concentration, and a semiconductor detector from the semiconductor detector. 2. A single X-ray projection type tomographic image reconstruction apparatus according to claim 1, comprising an analog-to-digital converter capable of converting an analog signal into a digital signal. 6 The X-ray concentration measuring device E includes a large number of semiconductor detectors that receive X-rays transmitted through the X-ray inspection object B and outputs signals corresponding to the X-ray concentration, and from these semiconductor detectors. 2. A single X-ray projection type tomographic image reconstruction apparatus according to claim 1, comprising an analog-to-digital converter capable of converting an analog signal into a digital signal. 7 The X-ray concentration measuring device E includes a xenon gas detector with a moving mechanism that receives X-rays passing through the X-ray inspection object B and outputs a signal corresponding to the X-ray concentration, and a xenon gas detector. Claim 1 comprising an analog-to-digital converter capable of converting an analog signal from a device into a digital signal.
The single X-ray projection type tomographic image reconstruction device described in 2. 8 The X-ray concentration measuring device E includes a large number of xenon gas detectors that receive X-rays passing through the X-ray inspection object B and output signals corresponding to the X-ray concentration, and these xenon gas detectors. 2. The single X-ray projection type tomographic image reconstruction apparatus according to claim 1, comprising an analog-to-digital converter capable of converting an analog signal from a device into a digital signal. 9 The projective transformation device F includes a digital computer that calculates and outputs a digital signal corresponding to the X-ray absorption coefficient of each pixel based on the data output for each group of measured values. A single X-ray projection type tomographic image reconstruction device according to any one of items 1 to 8. 10 The storage device G also serves as a memory capable of storing tomographic image information for each tomographic portion of the plurality of tomographic portions in the X-ray subject B and displaying any tomographic image on the display device J. A single X-ray projection type tomographic image reconstruction apparatus according to any one of claims 1 to 9. 11 Claims 1 to 10, wherein a digital filter is interposed between the reconstruction device H and the display device J in order to improve the quality of the image displayed on the display device J. The single X-ray projection type tomographic image reconstruction device according to any one of the above. 12 Claims 1 to 10, wherein a smoothing circuit is interposed between the reconstruction device H and the display device J in order to improve the quality of the image displayed on the display device J. The single X-ray projection type tomographic image reconstruction device according to any one of the items. 13 Claims 1 to 10, wherein a differentiation circuit is interposed between the reconstruction device H and the display device J in order to improve the quality of the image displayed on the display device J. The single X-ray projection type tomographic image reconstruction device according to any one of the above.