JP2004248699A - X-ray imaging method and x-ray imaging system - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To obtain a phase contrast image using an X-ray tube widely used on a normal medical site and to adapt this imaging technique to enhance the image quality of an X-ray image. <P>SOLUTION: Magnifying imaging is performed using X rays having no coherence radially diffused from an orifice having a finite size and a phase contrast is formed on the half-shadow blur on the magnified imaged surface. <P>COPYRIGHT: (C)2004,JPO&NCIPI

Description

[0001]
TECHNICAL FIELD OF THE INVENTION
The present invention relates to an X-ray image capturing method and an X-ray image capturing system for capturing an X-ray image used for medical image diagnosis, and more particularly to an X-ray image capturing method and an X-ray image capturing system capable of obtaining high image quality.
[0002]
In X-ray imaging, when X-rays are transmitted through an object, an image can be obtained due to a change in X-ray intensity due to a decrease in the X-ray intensity due to absorption or scattering by constituent elements of the object. This is called an absorption contrast X-ray image. Further, since X-rays are electromagnetic waves, refraction and diffraction occur similarly to visible light. The image contrast obtained by this property is called phase contrast. An X-ray image obtained by adding this phase contrast to the above absorption contrast is referred to as a phase contrast X-ray image. The present invention relates to a phase contrast X-ray image capturing method and an X-ray image capturing system capable of realizing this high-quality X-ray phase contrast image widely in general medical facilities.
[0003]
[Prior art]
Improving the image contrast by adding a phase contrast to an X-ray image based on absorption contrast is disclosed in V.A. A. Somenkov, A .; K. Tjalic and S.M. Sh. According to Shil's shtein, the scientific journal Sov. Phys. Tech. Phys. 36, No. 11, pages 1309-1311 (1991) (Patent Document 1). However, the refractive index of X-rays is extremely small, on the order of one millionth of visible light. That is, as shown in FIG. 1, when the refractive index of a plastic resin for visible light in air is greater than 1 and about 1.2 to 1.7, and the refractive index for X-rays is represented by n = 1−δ, For glass in air for 20 keV X-rays, δ = 1.3 × 10^-6Therefore, n = 0.99999987.
[0004]
The wavelength of X-rays is about one thousandth of visible light. Therefore, it is not easy to obtain a phase contrast by X-ray refraction or diffraction.
[0005]
Therefore, the above paper and Victor N. US Patent 5,319,694 by Ingal (Patent Document 2) and PCT WO95 / 05725 by Stephan Wilkins (Patent Document 3) disclose a phase contrast X-ray image using two crystals for Bragg reflection and for analysis. An imaging method is disclosed. That is, as shown in FIG. 2, an X-ray beam narrowed down to a point or a line is made obliquely incident on a silicon single crystal plane, and a monochromatic X-ray beam is obtained by Bragg reflection. After the X-ray beam has passed through the subject, it is again subjected to Bragg reflection with a silicon crystal analyzer, and the refraction contrast is superimposed on the conventional absorption contrast by selectively extracting the X-ray refracted when passing through the subject. Improve image contrast. The X-ray source disclosed here uses a copper Kα ray or a Mo anode X-ray tube, and uses an X-ray beam whose X-ray beam diameter is narrowed down by a collimator. Therefore, with this method, the X-ray intensity is extremely low, and it is possible to capture only a small subject such as a fly, and it is difficult to capture an X-ray image of a human body.
[0006]
On the other hand, a phase contrast X-ray imaging technique capable of imaging a larger object by using a synchrotron radiation X-ray to obtain a strong point-like or linear X-ray beam has been developed by, for example, Fulvia Arfeli et al. This is reported in Magazine Radiology Vol. 215, No. 1, pp. 286-293 (2000) (Patent Document 4).
[0007]
However, synchrotrons that produce synchrotron radiation X-rays are extremely large facilities and require enormous construction costs. That is, with this method, it is extremely difficult to take a phase contrast X-ray image widely in the medical field.
[0008]
Also, as shown in FIG. 3, monochromatic X-rays with good parallelism from the radiation X-rays are used. However, by separating the X-ray detector from the subject without using an analyzer as described above, the X-rays can be reduced. A method of obtaining a phase contrast image by enlarging a shift caused by refraction is disclosed in Scientific Magazine Radiology Vol. 195, No. 1, pp. 239 to 244 (1995) by Emilo Burattini (Patent Document 5). However, this method also uses synchrotron radiation X-rays in the same manner as described above, so that it is difficult to put it to practical use in medical practice.
[0009]
On the other hand, as shown in FIG. 4, a phase contrast imaging technique using a microfocus X-ray tube and not using Bragg reflection is disclosed in US Pat. No. 5,802,137 (Patent Document 6) and 6,018 by Stephan Wilkins. 564 patent (patent document 7). This is to obtain a phase contrast by enlarging a shift that transmits and refracts a subject by applying enlarged X-ray imaging. In this method, a window from which X-rays are projected, that is, an X-ray tube having an X-ray focal size of 20 μm or smaller is used in order to avoid so-called geometric unsharpness in magnified imaging.
[0010]
For example, as described in the scientific journal Imaging & Therapeutic Technology, Vol. 18, No. 5, pp. 1257 to 1267 (1998) reported by Dachao Gao et al., A collaborator of Stephan Wilkins (Patent Document 8). As shown in FIG. 5, when the focus size is large, a phase contrast image cannot be obtained due to blurred penumbra caused by the focal size. However, an X-ray tube having a small X-ray focus size such that blurring of penumbra can be ignored is too weak to transmit through a human body, so that the X-ray beam cannot be widely used in medical sites and the like.
[0011]
[Patent Document 1]
Scientific magazine Sov. Phys. Tech. Phys. Vol. 36, No. 11, pp. 1309-1311 (1991)
[0012]
[Patent Document 2]
US Patent No. 5,319,694
[0013]
[Patent Document 3]
PCT WO95 / 05725
[0014]
[Patent Document 4]
Scientific Magazine Radiology Vol. 215, No. 1, pp. 286-293 (2000)
[0015]
[Patent Document 5]
Scientific Magazine Radiology Vol. 195, No. 1, pp. 239-244 (1995)
[0016]
[Patent Document 6]
US Patent No. 5,802,137
[0017]
[Patent Document 7]
6,018,564 Patent
[0018]
[Patent Document 8]
Scientific Magazine Imaging & Therapeutic Technology, Vol. 18, No. 5, pp. 1257 to 1267 (1998)
[0019]
[Problems to be solved by the invention]
As described above, in the related art, a phase contrast image cannot be obtained unless a method using X-rays having extremely weak X-ray intensity or a method using a huge synchrotron radiation X-ray source facility is used. This makes it impossible to obtain a phase contrast X-ray image in a general medical facility.
[0020]
Therefore, in the present invention, a phase contrast image can be obtained using an X-ray tube generally used widely in a medical field, and an X-ray image capturing method for improving the image quality of an X-ray image by applying the technique and a method thereof. Is provided.
[0021]
[Means for Solving the Problems]
The above problem can be solved by the following X-ray image capturing method and X-ray image capturing system.
"X-ray imaging in which magnified radiography is performed using non-coherent X-rays radiating radially from an aperture of finite size, and a phase contrast is generated on a semi-shadow blur on the image plane of the magnified radiography Method and its photographing system ”.
[0022]
More specifically, "When the distance (R1) from the X-ray focal point to the subject is 0.2 m or more and 2.5 m or less, and the wavelength (λ) of the X-ray used is 8 pm to 71 pm, 0.16 × 10⁶× λL <r <2.5 × 10^-4using a thermionic X-ray tube with a focal size of 2r, which is a^-5X-ray imaging method using an X-ray image detector of m or more and an imaging system thereof ”.
[0023]
“The X-ray image detector is an X-ray image capturing method in which the X-ray image detector is a screen film system or a digital X-ray image capturing system and its capturing system”.
[0024]
"When the distance (R1) from the X-ray focal point to the subject is 0.2 m or more and 2.5 m or less, and the wavelength (λ) of the X-ray used is 8 pm to 71 pm, 0.16 × 10⁷× λL <r <2.5 × 10^-4Using a thermionic X-ray tube with a focal size of 2r and a spatial resolution of 5 × 10 m (250 μm)^-5X-ray imaging method using an X-ray image detector of m or more and an imaging system thereof ”.
[0025]
"X-ray imaging method and X-ray imaging system in which the X-ray image detector is a digital X-ray imaging system".
[0026]
"If the width of the blur on the image plane is B and the width of the edge enhancement caused by the phase contrast is E, an X-ray image capturing method and a capturing system satisfying 9E ≧ B”.
[0027]
"Using the thermionic X-ray tube having a focal size of 50 μm or more and 500 μm or less, the distance (R1) from the X-ray focus to the subject is 0.2 m or more and 2.5 m or less, and the X-ray image detector X-ray image capturing method and its capturing system having a distance of 0.15 m or more and 2.5 m or less.
[0028]
"When performing enlarged imaging using non-coherent X-rays radiating from an aperture of finite size, an X-ray image in which the edge enhancement width caused by phase contrast is enlarged by penumbra blur is converted to digital X-rays. An X-ray image capturing method and an image capturing system for capturing an image with an image detector and displaying the image on the digital X-ray image detector surface in a size smaller than the image size of the subject or in an actual size or a size close thereto.
[0029]
BEST MODE FOR CARRYING OUT THE INVENTION
Hereinafter, embodiments of an X-ray image capturing method and an X-ray image capturing system of the present invention will be described in detail with reference to the drawings, but the present invention is not limited to these embodiments.
[0030]
First, an X-ray and an X-ray source used for X-ray imaging according to the present invention will be described.
[0031]
X-rays were discovered by Dr. Roentgen in 1895 as "mysterious light that does not refract and has chemical and fluorescent effects." In 1912, Laue's crystal lattice diffraction experiment revealed that X-rays were short-wavelength electromagnetic waves. Electromagnetic waves are classified into γ-rays, X-rays, ultraviolet rays, visible light rays, infrared rays, and radio waves in the order of their wavelengths, and the wavelength of X-rays is as short as about 1 °.
[0032]
Electromagnetic waves cause refraction and diffraction when passing through an object. These phenomena are treated in electromagnetic optics, refraction is treated in geometric optics, and diffraction is treated in wave optics. In geometrical optics, a straight trajectory of light is represented by a line, and, for example, refraction of light by a lens or geometrical instability in X-ray magnified imaging is described. Regarding diffraction, for example, a phenomenon in which sunlight casts a shadow on the end of a high column blurred is described as a diffraction phenomenon caused by the wraparound of light, which is a property of the wave of an electromagnetic wave, and the wave optics must be repeated.
[0033]
However, optics teaches that wave interference that causes a diffraction phenomenon discussed in wave optics does not occur in all electromagnetic waves, but is a phenomenon that occurs in highly coherent electromagnetic waves. The coherence includes vertical coherence (temporal coherence) and horizontal coherence (spatial coherence). The vertical coherence refers to coherence due to a shift between two points along the traveling direction of a light ray, and is remarkable, for example, in a monochromatic electromagnetic wave. The lateral coherence is a coherence that focuses on a cross section perpendicular to the light traveling direction, and is a case of an electromagnetic wave whose wavefronts are aligned. This can be obtained with a light source that is an ideal point light source (that is, it has a "position" but no "size"), or a light source of a very small size that can be considered as a point light source in practice. Further, in the case of a light source having a size equal to or greater than a certain value, lateral coherence can be imparted by moving the light source far from the subject. For example, when the distance from the shadow object to the projection surface of the shadow is more than a certain value, such as the sun, when the distance from the shadow object to the projection surface of the shadow is more than a certain value, the shadow is blurred due to light sneaking in due to lateral coherence Is observed.
[0034]
Here, when attention is paid to the coherence in the lateral direction, the diffraction phenomenon is generally experienced as a natural phenomenon in the case of sunlight. However, with regard to artificial light sources, conversely, there are few light sources having this coherence. For example, in the case of a visible light source in the surroundings, in the case of a light bulb, light emitted from a heated filament is randomly emitted, so that the lateral coherence is low. The same applies to fluorescent lamps. In these cases, most optical phenomena can usually be explained by geometrical optics, and there is no need to wait for wave optics. A visible light source having both the horizontal coherence and the vertical coherence, which is widely known at present, is a laser light source, and optical elements utilizing the coherence have been widely put to practical use. For example, a hologram display is a typical example.
[0035]
X-ray sources generally used are all artificial sources, and the principle of X-ray generation is to convert kinetic energy of flying electrons into X-rays, which are electromagnetic waves. For example, in the case of a commonly used thermoelectron X-ray tube, thermoelectrons are generated from a filament, accelerated by applying a high voltage, and collide with a metal surface (target) to generate X-rays. Alternatively, in the case of a synchrotron, X-rays are obtained by accelerating electrons to near the speed of light and rapidly changing the flight direction of the electrons with a magnetic field or the like. A mechanism generated by forcibly changing the traveling direction of X-rays is called bremsstrahlung, and X-rays in a wide wavelength range are obtained, which is applicable to a thermionic X-ray tube and a synchrotron radiation X-ray source. When electrons collide with a metal surface, the X-rays that strike out the inner electrons of the atoms that make up the metal and fall into the vacant electron orbits to generate X-rays are called characteristic X-rays. Thus, strong X-rays having a narrow wavelength width can be obtained. It is obtained by thermionic X-rays, and the characteristic X-rays have different wavelengths depending on the type of target metal.
[0036]
Here, it is generally said that X-rays generated by causing thermal electrons to collide with a metal surface (target) have no or low lateral coherence. The generation of thermoelectrons from the filament occurs at random, and the thermoelectrons collide with the target again at random, making it difficult to obtain lateral coherence. This is easily understood from the fact that it is not possible to obtain lateral coherence with a light bulb that emits light from a filament. Since X-rays have a wavelength shorter than visible light by about one thousandth, it is particularly difficult to obtain lateral coherence. As described above, in the thermionic X-ray tube, lateral coherence cannot be obtained unless the size of the X-ray source is considerably reduced.
[0037]
Regarding the coherence of lateral X-rays, Sadao Aoki reported in Science Optics Vol. 27, No. 5, pp. 273-278 (1998), "In the case of a planar incoherent light source such as an electron beam excited X-ray source, The coherent region can be defined in a classical way. The distance D at which two points on the plane of interest have sufficient coherence is
D = 0.16λL / r
Is represented by Here, L is the distance from the light source to the plane of interest, and r is the radius when the light source is considered to be circular. " Here, λ is the X-ray wavelength.
[0038]
Here, “two points on a plane to be noted” is the size of the object to be resolved, and this is considered as the resolution of the detector. For example, when the resolution of the image detector is 100 cycles / mm, the distance between the above “two points on the plane of interest” can be calculated as 5 μm. The characteristic X-ray wavelength of the molybdenum tube for breasts is 71 pm (0.71 °), and when the distance from the X-ray tube to the “plane of interest” (the surface crossing the center of the subject) is L = 1 m, r = 2.7 μm. When L = 0.65 m, r = 1.8 μm. That is, in order to obtain a lateral coherence on a subject using a 17 keV characteristic X-ray obtained from a molybdenum anode X-ray with a detector having a resolution of 100 cycles / mm, the X-ray and the subject need to be combined. When the distance is 1 m, the radius of the X-ray source must be 2.7 μm or less, and when the normal mammography distance is 0.65 m, the radius must be 1.8 μm or less.
[0039]
Even though the resolution of an industrial X-ray film can realize 5 μm / mm, it is generally considered to be at most 10 μm / mm in a screen film system used for medical image diagnosis, and the pixel size of a digital X-ray detector for medical image diagnosis Since it is at most 50 μm, its spatial resolution is 10 cycles / mm. In the case of a digital system, since the distance D between “two points on the plane of interest” is 50 μm, if the distance from the X-ray tube to the “plane of interest” is L = 1 m, r = 0.27 μm. When L = 0.65 m, r = 0.18 μm.
[0040]
In the case of a molybdenum anode X-ray tube for mammography, the size of the X-ray source is generally designed to be a minimum and a square of 100 μm on a side, and the X-ray source is much larger than the above calculation result. Therefore, from the above calculation, it is theoretically possible that when the X-ray image detector is a screen film system or a digital system, the coherence in the horizontal direction is hardly obtained when the distance to the subject is about 0.65 m or 1 m. Is guided by
[0041]
In addition, the coherence in the horizontal direction was determined by Max Born and Emil Wolf using the degree of coherence (degree of coherence | μ).₁₂|: 0 ≦ | μ₁₂| ≦ 1). That is, 1 is an ideal case, the upper limit of the coherence, and 0 has no coherence. In ordinary natural phenomena, 1 is impossible, and is at most about 0.88.
[0042]
"Principles of Optics III" (published by Tokai University, 1975), a scientific book written by them, p. From equation (28) of 766, the complex coherence degree μ by the quasi-single light color₁₂Is
μ₁₂  = (2J₁(V) / v) exp iψ
Coherence of the absolute value | μ₁₂|
｜ μ₁₂| = | 2J₁(V) / v | (28 ')
Also, when transforming equation (29),
v = 2πρΔX / λR1 (29 ′), (28 ′), (29 ′)
From the equation, the degree of coherence (degree of coherence | μ)₁₂|) Can be obtained. In addition, J₁(V) is approximately “Principles of Optics II (1975 Tokai University Press)” p. 608 can be read from FIG.
[0043]
For example, as shown in FIG. 6, if the focal size of a molybdenum anode thermoelectron tube is 100 μm and the region of interest is 10 μm, ρ = 100 μm = 1 × 10^-4m, R1 = 1m, λ = 0.7Å = 0.7 × 10^-10m, ΔX = 10 μm = 1 × 10^-5m, and v = 90 is obtained. As a result, the coherence degree | μ₁₂| ≒ 0. That is, the coherence is theoretically equal to zero. That is, when R1 is about 1 m, the above-mentioned calculation result that the coherence in the lateral direction cannot be obtained with the thermionic X-ray tube having the focal size of 100 μm is consistent with the above calculation result. The same applies when R1 is 0.65 m.
[0044]
The present invention uses an X-ray source capable of obtaining X-rays that can pass through a human body, using a thermionic X-ray tube widely used in medical practice. In this case, the size of the X-ray source is a square having a side length of 100 μm (corresponding to 2r or ρ in the above theory) at a minimum of 300 μm or 600 μm. In general, the distance from the X-ray tube to the subject is 0 μm. .2 m to 2.5 m. Therefore, since it is impossible to obtain X-rays having high lateral coherence, the present invention utilizes refraction instead of utilizing diffraction of electromagnetic waves, and is positioned as a technology that can be handled in the category of geometrical optics. Can be
[0045]
Next, the X-ray tube will be described.
[0046]
Recently, new X-ray source technologies such as a synchrotron radiation synchrotron X-ray source and an X-ray tube using carbon nanotubes have been developed. For example, a new X-ray source technology is introduced in “Special feature: New trends in X-ray sources and X-ray imaging technology” published in the Photographic Society of Japan, Vol. 65, No. 7 (2002), which is a scientific journal. Although the present invention uses a thermionic X-ray tube, the principle of the present invention can be applied to X-ray imaging using a synchrotron radiation synchrotron X-ray source.
[0047]
X-rays discovered X-rays in a Crookes tube for cathode ray experiments. In this case, the electrons jumped out of the cathode accelerated and went straight, and collided with the glass wall, so that X-rays were generated from the glass surface. At this time, if the amount of electrons to be collided was increased, the glass could be melted by the heat due to the collision with the glass, so that a large tube current could not flow. The amount of X-rays generated can be greatly improved by colliding electrons with a heat-resistant metal surface such as platinum. After that, the cathode that generates electrons is changed from a metal plate to a filament, and the amount of X-rays generated is further increased by heating the filament in a vacuum and accelerating the hot electrons to collide with the metal surface. Was completed. This X-ray tube is called a thermionic X-ray tube or a Coolidge tube. In order to further increase the amount of X-ray to be generated, a rotating anode X-ray tube that rotates a metal surface against which thermoelectrons collide has been devised, and is now widely used in medical practice. Typical metals that generate X-rays by generating thermoelectrons include gold, platinum, silver, copper, tungsten, molybdenum, and rhodium.
[0048]
The anode applied voltage of the thermionic X-ray tube used in the present invention, that is, the tube voltage is in the range of 1 kVp to 500 kVp. This tube voltage is selected depending on the subject, and when the human body is the subject, a voltage of 20 kVp to 150 kVp is appropriate. This set tube voltage further differs depending on the part of the human body to be imaged. 28 kVp or 34 kVp for mammography, 50 kVp to 80 kVp for limb bone imaging, 60 kVp to 110 kVp for built-in gastrointestinal, etc., and lung imaging for 90 kVp to 150 kVp. These tube voltage settings are set by a radiologist or doctor in order to obtain an appropriate image based on the characteristics of the X-ray detector in consideration of the physique and the like of the patient who is the subject.
[0049]
Tungsten, molybdenum, rhodium and the like are used as the metal constituting the anode of the thermal X-ray tube used in the present invention. Molybdenum and rhodium are generally used for mammography, and tungsten is generally used for other parts. The anode of the X-ray tube used in the present invention may be a fixed anode or a rotating anode. However, it is preferable to use a rotating anode through which a large amount of X-ray tube current can flow. In the present invention, for example, as described in Japanese Patent Application Laid-Open Publication No. H11-135044, it is possible to change the X-ray region to be emitted by giving a plurality of inclinations to the target surface of the rotating anode. Can be used. In addition, when taking a breast X-ray image, the technique disclosed in Japanese Patent Application Laid-Open No. 2001-91479, which is a published patent publication, can be used.
[0050]
The size of the opening of the thermal X-ray tube used in the present invention, that is, the focal point size is preferably 50 μm or more and 500 μm. The focal point of a thermionic X-ray tube is generally a square, and the length of one side is called a focal point size. If the shape of the focal point is a circle, it indicates its diameter, and if it is rectangular, it indicates its short side. Methods for measuring the focal size include a method using a pinhole camera, a method using a micro test chart, and the like are described in JIS Z 4704. Normally, the focal size is indicated by a product specification based on the measurement of the X-ray manufacturer, and the focal size is displayed by the X-ray manufacturer. This focal size is a “nominal focal size”, which is different from the focal size referred to in the present invention. The size of the nominal focal size is an allowable range of about ± 50%. Therefore, the effective focal size measured by the method described in JIS Z4704 is defined.
[0051]
In order to operate the thermionic X-ray tube, a high voltage device for applying a high voltage to the anode is required. In order to supply a high voltage and a high current in a short time, a capacitor type high voltage device is used. In addition, a high voltage generator that generates a single-phase current, a three-phase current, and an inverter-type high-frequency current through a rectifier is used. In the present invention, an inverter type high voltage generator is preferable.
[0052]
The thermal X-ray tube includes a tube attached to an X-ray generator fixed on the floor and a tube attached to a movable X-ray generator. There is also a C-arm or U-arm type X-ray generator that can simultaneously fix an X-ray tube and an X-ray detector. In the present invention, any X-ray generator can be used.
[0053]
When a microtron that is a small synchrotron is used, the technique disclosed in Japanese Patent Application Laid-Open No. 2002-172109 can be used. In addition, when using a synchrotron that is a large-sized device, JP-A-2002-336229, JP-A-2002-336230, JP-A-2002-336231, JP-A-2002-336232, which are patent publications. The technique can be applied to the technology disclosed in JP-A-2002-336233.
[0054]
Next, the X-ray imaging conditions will be described.
[0055]
A wave is represented by an amplitude representing the height of the wave and a phase representing the position of the wave. When the X-rays pass through the object, the height of the waves, that is, the X-ray intensity is reduced by the absorption of the X-rays by the object. This is called absorption contrast, and a conventional X-ray image has been drawn only with this X-ray contrast. On the other hand, when X-rays pass through the interface between objects having different densities, a phase change occurs, and X-ray refraction occurs. The contrast of the X-ray image caused by this refraction is called phase contrast (FIG. 3). The present invention uses this phase contrast to improve an X-ray image.
[0056]
Reported a theoretical formula for calculating the strength of phase contrast using geometrical optics in Scientific Magazine, Optical Review, Vol. 7, No. 6, pp. 566-572 (2000). As shown in FIG. 7, when the intensity of the phase contrast is defined by the half width E,
E = 2.3 (1 + R2 / R1)^1/3｛R2δ (2r)^1/2｝^2/3
Can be represented by At this time, R1 is the distance from the X-ray focal point to the object, R2 is the distance from the object to the image plane, δ is the refractive index of the object n = 1−δ, and r is the radius of the object when the object is a cylinder. Represent. This is an approximate expression when assuming that the X-ray source is a point light source.
[0057]
Here, in order to enlarge the shift of refracted X-rays, it is necessary to increase the distance R2 between the subject and the image plane. However, if a thermionic X-ray tube having a finite focal size is used, apart from an ideal point light source, as shown in FIG. If the width of this geometrical unsharpness is B, it can be expressed as B = D × (R2 / R1). D is the focal size of the thermionic X-ray tube. At this time, the half width of the phase contrast can be represented by E + B.
[0058]
According to the present invention, a magnified image is taken using a thermionic X-ray tube having a finite focal size, and a phase contrast image can be taken when geometric instability occurs. When the half width E and the blur width B are used, 9E ≧ B. More specifically, the distance (R1) from the X-ray focal point to the subject is 0.2 m or more and 2.5 m or less, and the X-ray detector This is an X-ray imaging condition in which the distance to the camera is 0.15 m or more and 2.5 m or less, and an X-ray imaging system that realizes the condition.
[0059]
This invention is disclosed in Japanese Patent Application Laid-Open Nos. 2001-91479 and 2001-311701, and is a scientific journal by Honda et al., Medical Physics Magazine, Vol. 22, No. 1, pp. 21-29 (2002); Further, it is reported by Ohara et al. In Konica Technical Report Vol. 15, pp. 41-44 (2002).
[0060]
From these reports, conditions for photographing a phase contrast image according to the present invention or photographing conditions under which edge contrast is easily obtained are as follows.
[0061]
The focal size of the thermionic X-ray tube is 50 μm or more and 500 μm or less, and the smaller the focal size, the easier it is to obtain a phase contrast. An X-ray tube voltage of 20 kVp to 150 kVp is appropriate for imaging of a human body, and the lower the tube voltage, the easier it is to obtain a phase contrast. The imaging tube voltage is set in consideration of the appearance of the phase contrast in accordance with the thickness of the subject and the imaging region. As in an X-ray tube having molybdenum as an anode, the phase contrast is more easily obtained as the characteristic X-rays are stronger and the bremsstrahlung component is smaller. The focus size and the distance to the subject (R1) are 0.2 m or more and 2.5 m or less, and the distance from the subject to the X-ray detector is 0.15 m or more and 2.5 m or less. In both cases, the greater the distance, the easier it is to obtain phase contrast. That is, since the geometric unsharpness is B = D × (R2 / R1), the blur becomes smaller as R1 becomes larger. In addition, an increase in R2 increases a blur while increasing a refraction shift. Therefore, as R2 increases, phase contrast is more easily obtained. The enlargement ratio is expressed by (R1 + R2) / R1 by R1 and R2. The photographing conditions are set based on the magnification and the thickness or region of the subject. In order to magnify the image, an X-ray image detector having a size of six slices, a half slice, and a large side of 1 m or more is used. The size of the subject is large and spherical, and the greater the difference between the refractive index of the subject and the refractive index of the material surrounding the subject, the easier it is to obtain a phase contrast. All of the conditions disclosed herein in which the phase contrast is easily obtained can be explained by the theory of refraction of X-rays by geometrical optics.
[0062]
Next, the X-ray image detector will be described.
[0063]
Conventionally, an X-ray film or a screen film system (SF system) configured by a line intensifying screen (screen) and an X-ray film (film) has been used for detecting an X-ray image. This is an analog image capturing system, which is described in detail in the technical book "Revised Basics of Photographic Engineering", edited by The Photographic Society of Japan, page 718 (1998), published by Corona Publishing Co., Ltd. Recently, digital X-ray imaging apparatuses have been widely used. In particular, computed radiography (CR) has been widely used since the 1980's, and its usefulness has been recognized. For example, Honda et al. Describe in detail in the scientific journal of the Photographic Society of Japan, Vol. 64, No. 2, pp. 105-118 (2001). In the 1990's, flat panel X-ray detectors (FPD) have begun to be used. This technique is described in John A. The details are described in detail in a technical book "Handbook of Medical Imaging", Vol. 1, Chapter 4, pp. 223-328 (2000), SPIE Press (USA) by Rowlands et al. In the present phase contrast X-ray imaging technique, phase contrast X-ray imaging is performed using an analog SF system and a digital X-ray imaging system such as CR and FPD.
[0064]
Next, an analog X-ray imaging system will be described.
[0065]
A silver halide photographic film is used for analog X-ray imaging. This is an industrial X-ray film that directly irradiates the film with X-rays to expose the silver halide particles, and a visible light is generated by X-ray irradiation to increase the sensitivity, and the silver halide particles are converted by the visible light. There is an SF system that exposes light. Although both are characterized by high spatial resolution, industrial X-ray films that are directly exposed to X-rays generally have higher resolution because they do not involve the process of generating visible light. The analog X-ray imaging system has a disadvantage that a wet process called a development process must be performed after exposure.
[0066]
In an industrial X-ray film, silver halide grains constituting the silver halide film absorb X-rays to form a latent image, and an X-ray image is obtained by a developing process. However, since the amount of X-ray absorption of a silver halide film is generally small, the amount of silver on both surfaces of the resin film is converted to 10 g / m²And many silver halide grains must be used, which is expensive. The silver halide grains used are mixed crystals of silver iodide and silver bromide, and have an average grain size of 0.2 μm to 1 μm. The film support is made of polyethylene terephthalate (PET) resin having a thickness of 175 μm. Can be After taking the X-ray image, the film is developed with a developing solution designated by a film manufacturer or a similar developing solution by vat development or an automatic developing machine. The resolution of the image obtained with this film is very high, and it is widely used for nondestructive inspections and the like. However, when particularly high resolution is required for the images, industrial X-ray films can be used for phase contrast radiography techniques.
[0067]
To capture a general medical diagnostic X-ray image, a highly sensitive SF system is used to reduce the exposure dose of the patient. In this method, a screen and a silver halide film are superimposed, loaded into a cassette, and X-ray imaging is performed. That is, the screen emits light (fluorescence) having a wavelength in the near-ultraviolet to visible light range by the irradiated X-rays, and the light exposes the silver halide grains of the X-ray film. Then, an X-ray image can be obtained by vat development or automatic development using a developing solution designated by a film manufacturer or a similar developing solution.
[0068]
The screen is obtained by dispersing phosphor particles that emit visible light when irradiated with X-rays in a polymer binder and applying the phosphor particles on a support. Further, a protective film for preventing abrasion or the like is provided thereon, but in order to improve the sharpness of an image, the thickness thereof is preferably as thin as possible. That is, the screen has a rare earth phosphor such as calcium tungstate or gadolinium oxysulfide, and converts the energy of the irradiated X-rays into blue or green light. In particular, the technique disclosed in JP-A-6-67365 may be used for an intensifying screen using a rare earth phosphor. Further, as the silver halide photographic film, it is preferable to use one having a photosensitive emulsion coated on only one side of a support or one having a photosensitive emulsion coated on both sides of a support. The thickness of the photosensitive emulsion layer is preferably about 1 μm to 5 μm. The coated silver amount is 3 to 8 g / m²The silver halide grains are preferably silver iodobromide containing 0.5 to 3 mol% of iodine, and spherical, potato-like and tabular grains are used. In particular, in the case of a double-sided film, it is a preferred embodiment to use a photographic light-sensitive material having different photographic characteristics for each emulsion layer sandwiching a film support. It is also preferable to use a photographic film having a layer for absorbing cross-over light between the respective emulsion surfaces of the double-sided film. That is, it is most preferable that the crossover cut rate is at least 80% or more, preferably 90% or more and 100%. The size of the single-sided and double-sided films used in the present invention can be any size from six-cut size to half-cut size. These silver halide photographic light-sensitive materials are outlined in JP-A-6-67365 and, for example, "Revised Basics of Photographic Engineering-Silver halide photography-" (edited by The Photographic Society of Japan, Corona, 1998). As for the development processing, the average gradation can be increased by raising the development processing temperature or extending the processing time, but when performing the automatic development processing, in principle, processing is performed under the development processing conditions specified by the film manufacturer. Is preferred.
[0069]
In phase contrast X-ray imaging, so-called enlarged imaging is performed. In order to perform enlarged photographing, it is preferable to use a high-sensitivity SF system. That is, when a green light-emitting screen is used using a double-sided coated film, the standard sensitivity is the sensitivity of the SF system combining an SRO250 screen (manufactured by Konica Corporation) and an SRG film (manufactured by Konica Corporation). And It is preferable to use an SF system that combines a screen and a film having this sensitivity and higher than this sensitivity. The image contrast of the SF system is 2 or more and 5 or less. In order to obtain high sharpness, it is preferable to use an SF single-sided system in which a single-sided coated film and one screen are combined. In particular, it is generally used for mammography. In this case, the sensitivity is preferably 200 or more and 750 or less.
[0070]
In the present invention, the sensitivity of the SF system for mammography is measured as follows. That is, a screen / film system is exposed to a molybdenum target tube operated by a three-phase power supply of 28 kVp using X-rays transmitted through 1 mm of beryllium, 0.03 mm of molybdenum and 2 cm of an acrylic filter, and a silver halide photographic material is exposed. Is performed. The development process is performed using Konica SRX-502 automatic developing machine, the developer is Konica XD-SR, and the fixing solution is Konica XF-SR at 34 ° C. for 90 seconds (note that the film is specified by the manufacturer of the film used). , Or other developing processes). Here, the sensitivity refers to the reciprocal of the X-ray irradiation dose required to give an optical density of fog + 1.0, and is expressed as a relative value. That is, the sensitivity of the SF system of MD100 (manufactured by Konica) and CMH film (manufactured by Konica) is set to 100 as a reference.
[0071]
For example, the SF system is placed at a distance of 60 cm from a 1 mm beryllium, 0.03 mm molybdenum and 2 cm acrylic filter in a molybdenum target X-ray tube, and X-rays are irradiated from the film side. The irradiation X-ray dose adjustment changes the mAs value of the X-ray tube, calculates the reciprocal of the X-ray irradiation dose necessary to give an optical density of fog +1.0 after the development processing, and obtains an X-ray intensifying screen / silver halide photograph. Assume that the assembly sensitivity of the film is 100. Then, the X-ray intensifying screen instead of the MD100 to be measured next and the film CMH are combined, and the relative sensitivity of the intensifying screen is obtained in the same manner as described above. The developing process is performed using Konica SRX-502 automatic developing machine, the developing solution using Konica XD-SR, and the fixing solution using Konica XF-SR at 34 ° C. for 90 seconds (or another processing method specified by the manufacturer). But it doesn't matter).
[0072]
Here, in a screen film system, sharpness decreases as the sensitivity of the system increases. In the present invention, it is preferable to use an SF system for X-ray mammography which has high sensitivity and sharpness. That is, the sensitivity of the X-ray intensifying screen for mammography is 200 or more and 750 or less, and the contrast transfer function (CTF) of the X-ray intensifying screen is 0.3 or more and 1.00 or less at a spatial frequency of 5 lines / mm. The SF system is characterized in that: Here, the contrast transfer function (CTF) is one physical quantity representing the sharpness of an image obtained by the SF system used. The maximum value is 1.0, the minimum value is 0, and the higher the value, the better the sharpness. This measurement is performed roughly as follows. That is, a square wave chart made of lead is brought into close contact with a screen / film system, and X-ray irradiation is performed. The density of the rectangular wave image obtained after the development processing is measured with a microdensitometer, and a CTF for each spatial frequency is obtained.
[0073]
The sharpness of the SF system depends on the contrast of the silver halide photographic film used therein and the sharpness of the intensifying screen. An intensifying screen for X-rays having good sharpness is composed of a support, a phosphor layer containing a phosphor, and a protective layer, and the phosphor has an average particle diameter of 2 μm or more and 5 μm or less. It is preferable that the filling ratio of the phosphor layer is 60% or more and 80% or less, and the binder weight ratio of the phosphor layer is 0.1% or more and 5% or less. The binder preferably contains a resin having a hydrophilic polar group.
[0074]
The single-sided silver halide film for mammography used in the present invention has an emulsion surface on one side of a blue-colored support of a subbed polyethylene phthalate having a thickness of 100 to 200 μm, and the other side has a backing. A layer coated with a gelatin film is preferred. This backing layer is coated mainly for the purpose of preventing curling of the film, but a matting process for preventing sticking between the films is a preferred embodiment, and also contains an antistatic agent, an antihalation dye, and the like. Is a preferred embodiment. This film is preferably in the form of a sheet and has rounded corners in order to prevent injuries during handling, and preferably has a notch for discriminating the emulsion surface. In the present invention, it is preferable to use a conventionally used silver halide photographic light-sensitive material having a six-section size and a larger size in order to perform magnified photographing and to obtain a whole breast image.
[0075]
The single-sided silver halide film for taking a breast image used in the present invention may have one or more emulsion layers and layers each containing silver halide grains having a different average grain size. The sensitivity relationship between these photosensitive layers is not particularly defined. That is, when the photosensitive layer is composed of an upper layer and a lower layer, the upper layer sensitivity may be lower than the lower layer sensitivity or may be higher. This should be designed for diagnostic purposes.
[0076]
The silver halide grain composition contained in the emulsion layer of the silver halide film for mammography used in the present invention is preferably silver iodobromide, and the molar composition of iodine is preferably 2% or less. The shape of the silver halide grains may be cubic, octahedral, or plate-like, and may be a mixed state of them, or may be one of them. In the case of a plate, those having an average aspect ratio of 2 or more and less than 15 can be used. Further, it is preferable to use one in which silver halide emulsion grains are monodispersed, and it is also possible to use a mixture of several monodispersed emulsions. These silver halide emulsion grains are subjected to appropriate chemical sensitization such as gold-sulfur sensitization and selenium sensitization, and can be doped with metal ions such as iridium cations when forming silver halide grains. The sensitizing dye may be added when silver halide grains are formed. In particular, it is preferable to use the technology disclosed in Japanese Patent Application Laid-Open No. 2001-194738 for the breast imaging SF system used in the present invention.
[0077]
Next, a digital X-ray imaging system will be described.
[0078]
CR and FPD are used for digital X-ray imaging. This feature is that image processing can be easily performed in order to convert image information into digitized electric signals. In an analog X-ray imaging system, X-ray image detection and X-ray image display are performed on the same silver halide film. Therefore, when designing an image detection function, there is a restriction from the X-ray image display function. , Cannot fully demonstrate its ability. However, in the digital X-ray imaging system, since the X-ray image detector and the image display are functionally separated from each other, the capability can be sufficiently brought out as the X-ray image detector. For example, for both CR and FPD, the dynamic range of image detection is about twice that of SF systems. However, the digital X-ray imaging system has a drawback that the spatial resolution is smaller than that of the analog X-ray imaging system because the image is detected in pixels of a finite size. Generally, the pixels of a digital imaging system for medical image diagnosis have a size of 50 μm square to 300 μm square. That is, an image narrower than the pixel or an image smaller in size cannot be described.
[0079]
Here, the distance between the peak and the valley of the edge enhancement obtained by the phase contrast is about several μm when there is no blur on the image plane using an ideal point light source. Therefore, if the peaks and valleys of the edge enhancement are captured in one pixel of the detector, they cancel each other out and the edge effect cannot be detected. However, as shown in FIG. 9, the distance between the peak and the valley is widened by the phase contrast being on the blur due to the geometric instability, so that the peak and the valley can be captured by different pixels. , And the edge effect can be effectively detected.
[0080]
As shown in FIG. 10, blur (B value) on the image plane of a plastic fiber having a diameter of 8.5 mm photographed under the conditions of R1 = R2 = 1 m (double magnification photographing) and a tube voltage of an X-ray tube of 50 keV. 7 shows a simulation result of an X-ray intensity profile of edge enhancement when changed. Here, when the pixel size of the detector is 87.5 μm, the B value is 100 μm, the peaks and valleys of the edge enhancement are detected by different pixels, and the edge enhancement is efficiently obtained. As described above, when the digital X-ray imaging system is used, the phase contrast width is enlarged by the geometric instability, so that a phase contrast image can be captured.
[0081]
As described above, in the case of digital X-ray imaging, the output image size for performing image diagnosis can be easily changed. Depending on the purpose, enlarged display, actual size display, and reduced display can be performed. Since the phase contrast image photographing is an enlargement photographing, if it is displayed as it is, it is enlarged and displayed at the same magnification as that at the time of photographing.
[0082]
On the other hand, by displaying the same image size as conventional contact imaging, it is possible to make the image size familiar to the doctor, and viewing at the same image size as conventional contact imaging that has already been stored is important for image diagnosis. It is convenient. Therefore, the display of the actual size in this case means that the image is displayed at an enlargement ratio of the image obtained by the conventional close-up photographing. Since the subject is actually thick, it is in principle difficult to fix all the portions of the subject image to the same enlarged size. It is also possible to display a smaller size than the contact photographed image size, that is, the actual size. The small size allows the whole to be grasped at a glance, and the image quality can be improved by reducing the output pixels.
[0083]
For the CR system, see the scientific papers listed above and A. The technology is described in a recent scientific paper by Rowlands, Physics in Medicine and Biology, Vol. 47, pp. R123-R166 (2002).
[0084]
In the CR, the X-ray energy is temporarily stored in a stimulable phosphor, and fluorescent light having an intensity proportional to the irradiated X-ray intensity is extracted by light irradiation. The pixel size corresponds to the interval (sampling pitch) for reading this signal. The stimulable phosphor used here is not particularly limited, but europium-activated fluorofluorobarium halide is used. Halides are chlorine, bromine, iodine and their mixtures. Also, thallium-activated cesium iodide, europium-activated cesium bromide, and thallium-activated rubidium bromide are used. It is preferable to use a laser light source for the light that causes stimulable light emission, and the wavelength is selected from the physical properties of each stimulable phosphor. The stimulated emission is read by a photomultiplier tube, CCD, or the like. At this time, the stimulable luminescence can be read from both sides by using the substrate holding the stimulable luminous body as a transparent substrate. In the case of a columnar crystal, laser light irradiation is performed from the transparent substrate side, and stimulated emission can be read from the columnar crystal side. Alternatively, laser light irradiation can be performed from the columnar crystal side, and stimulated emission can be read on the same surface. Further, it is also possible to irradiate the laser beam a plurality of times to extract the stimulated emission a plurality of times. Using a plurality of, for example, four stimulable phosphor plates, a larger area, that is, a half-cut size or larger area can be photographed.
[0085]
Regarding the FPD, a direct type FPD that directly converts the energy of the irradiated X-rays into electric charges, reads out the amount of electric charges generated in proportion to the X-ray irradiation amount, and uses the read out as an image signal, Two types of indirect FPDs, which once convert energy into visible light and irradiate the light to an optical semiconductor to obtain an electric signal proportional to the light intensity and obtain an image signal, are typical. The latter also includes a method of reading visible light generated by a CCD without using an optical semiconductor.
[0086]
A-Se is generally used as a material for converting the energy of irradiated X-rays into electric charges used in the direct type FPD. A voltage close to 1000 V is applied above and below the a-Se plate to separate electrons and holes generated by X-ray irradiation. This electric charge is temporarily stored in a capacitor, and the electric charge is read as needed by a thin film transistor (TFT). Instead of a-Se used here, lead iodide, mercury iodide, lead oxide, and thallium bromide can be used.
[0087]
A pixel in this method is a unit read by a TFT, and has a side length of 50 μm to 300 μm for the purpose of medical image diagnosis. In order to apply a high voltage, it is preferable to use a Zener diode in order to prevent the TFT from being broken. As the TFT, an FET (Field Emission Transistor) using a-Si is used, and a TFT using low-temperature polysilicon, an organic semiconductor, or the like is preferably used.
[0088]
Further, for the substrate of the TFT, glass, resin, or a composite material thereof can be used. Further, it is preferable to form an FPD having a larger area in units of two or four FPDs. The pixel size of the direct type FPD used for mammography is preferably 100 μm or less, and 50 μm, 25 μm, and even 10 μm pixels are preferably used. The thickness of a-Se used for the direct type FPD used for mammography is preferably 1000 μm or less and more preferably 100 μm or more.
[0089]
The indirect FPD includes a scintillator for converting irradiation X-ray energy into light, an optical semiconductor for converting fluorescent light emitted from the scintillator into electric energy, and a TFT for reading an image signal from the optical semiconductor in a two-dimensional plane. The scintillator is one in which phosphor particles that emit visible light by X-ray irradiation are dispersed in a polymer film. The phosphor used is a rare earth phosphor such as calcium tungstate or gadolinium oxysulfide, or CsI. The phosphor may be dispersed in the polymer medium as particles, and it is preferable to use spherical phosphor particles having an average particle diameter of 3 μm or less and 1 μm or more. CsI and the like are preferably used as columnar crystals by a vapor deposition technique or the like. In particular, the technique disclosed in JP-A-6-67365, which is a phosphor used for an X-ray screen, may be used for a rare earth phosphor. As the optical semiconductor, those having a-Si as a main composition, those having crystalline Si as a main composition, and the like can be used. Further, it is more preferable to use a mixture of fullerene, a carbon nanotube, and a conductive polymer such as polythifen.
[0090]
As the optical semiconductor used here, an nip type optical semiconductor, a Schottky type optical semiconductor or a MIS type semiconductor is used. The pixel size of the indirect FPD is determined by the length of one side of the optical semiconductor. The pixel size of the indirect FPD used for medical image diagnosis is about 50 μm to 300 μm. In particular, in the case of a breast image, the size is preferably 100 μm or less, and more preferably 50 μm, 25 μm, and even 10 μm. As the TFT, an FET (Field Emission Transistor) using a-Si is used, and a TFT using low-temperature polysilicon, an organic semiconductor, or the like is preferably used. Further, for the substrate of the TFT, glass, resin, or a composite material thereof can be used. Further, it is preferable to form an FPD having a wider area of a half-cut size or more in units of two or four FPDs.
[0091]
As another technique of the indirect FPD, a technique in which light emitted by a scintillator is guided to each pixel of a CCD by condensing an optical fiber or a lens, and the CCD converts light energy into electric energy can be used. The pixel size is determined by the diameter of the fiber for collecting the fluorescent light from the scintillator or the diameter of the lens. Also in this case, the preferable range of the pixel size is the same as the above. In order to reduce the noise of the CCD, it is preferable to perform air cooling or water cooling. It is preferable to form an FPD having a larger area by using a plurality of FPD panels.
[0092]
Next, image processing will be described.
[0093]
When a digital X-ray imaging system is used for phase contrast imaging, it is important in the present invention to process the obtained image signal to form an easily viewable image.
[0094]
Since the phase contrast image photographing is 1.2 to 3 times magnification photographing, by returning to the conventional contact image size, diagnosis can be performed without discomfort from the past image. At this time, it is preferable to perform the reduction process by the enlargement factor because the interpolation process is unnecessary. Also, when reducing the image, it is preferable to match the minimum pixel size or the standard pixel size of an output device such as a laser imager. Alternatively, it is preferable to reduce the pixel size to an integral multiple of the minimum pixel size. When reducing the pixel size to an integral multiple of the output pixel size, it is preferable to perform phase contrast imaging at a magnification that is equal to or close to the conventional contact image size.
[0095]
When forming an output image using the obtained image signal, it is fundamental to form the output image with a linear image density for the image signal. Following the conventional SF system, the image can be converted into an S-shaped curve and an image output image can be drawn. It is also possible to form an image output image with a moderate contrast by selecting a portion having a large signal value or a portion having a small signal value. Further, it is preferable to output the maximum image density of an image output device or an image output film for an image signal area equal to or less than a certain value. In a conventional SF system image, a maximum image density contrast is drawn near a density of 1. The digital image contrast is preferably from 1 to 3, but especially in the case of a breast image, it is preferable to describe in the range of 2.5 or more and 5 or less. Further, it is preferable to output the image signal in the image area range so as to fall within the minimum density and the maximum density area of the output image from 70% to 100%.
[0096]
In the phase contrast image, the edge of the subject is emphasized. Further, it is preferable to further enhance the periphery by so-called blur mask processing. When performing image enhancement of the peripheral portion, an increase in image noise can be suppressed by increasing the enhancement as the image signal intensity change increases, and reducing the enhancement as the image signal intensity changes decrease. Further, in order to decompose the obtained image signal into spatial frequencies and obtain an image suitable for each diagnostic site, it is preferable to perform an emphasis process based on the spatial frequency.
[0097]
In phase contrast image capturing, the edges of the subject are emphasized. That is, the image signal change is enlarged. It is preferable to extract only a portion where the change in the image signal is large, and to draw an image whose main edge is main. It is also preferable to add the large portion of the extracted field image signal to the original image again to enhance the phase contrast.
[0098]
The phase contrast differs depending on the X-ray tube voltage at the time of X-ray imaging. That is, when the voltage is high, the phase contrast is small, and when the voltage is low, the phase contrast is large. Therefore, it is preferable to control the phase contrast component by adding or subtracting two captured image signals having different X-ray tube voltages.
[0099]
The obtained phase contrast image is superimposed on the past captured image to emphasize the change in the medical condition, so-called time-difference processing. The lesion information is input to a computer in advance, and the possibility of the lesion is compared with the obtained image information. It is preferable to use a so-called CAD (Computed Assisted Diagnostics) that displays on an image.
[0100]
Techniques such as JP-A-2001-238871, JP-A-2001-299733, JP-A-2002-85389, and JP-A-2002-159482 can be used for a digital imaging system for phase contrast X-ray images.
[0101]
Next, output of a phase contrast X-ray image will be described.
[0102]
In the digital X-ray imaging system, unlike the SF system, X-ray imaging and X-ray imaging are separated, and images can be easily displayed by various methods. That is, a hard copy of a transparent image displayed as an image similar to a conventional SF system, a reflective hard copy mainly used for reference, or a so-called soft copy such as a cathode ray tube (CRT) or a liquid crystal. Is output to The display depth of these images is preferably 8 bits or more, more preferably 10 bits or more. The maximum density of the image is preferably 2.5 or more and 5 or less. Particularly when displaying a breast image, the image density is preferably 3 or more, more preferably 3.4 or more.
[0103]
When making a hard copy, there are a method of drawing the obtained image on a CRT and printing it on a silver halide film through a lens, a method of printing it on a silver halide film with a laser beam having an intensity proportional to the image signal intensity, and the like. In addition, a so-called thermal head that generates heat in proportion to the image signal intensity, which draws an image corresponding to the change in the amount of heat, and inks with different ink ejection amounts and densities so that the density is proportional to the image signal intensity. There is an ink jet method in which an image is drawn by appropriately discharging an image. When a transparent medium is used, it is preferable to use a transparent resin base colored in blue or a colorless transparent resin base as in the conventional SF system. It is preferable that the thickness is designed around 175 μm. When forming a reflection image, it is preferable to use a white resin base or a base using paper. In the case of a silver salt image, a neutral black tone image is preferable, and a reddish image or an image with a yellow tint is not preferable. The silver salt image includes a film subjected to a conventional development process, that is, a wet process, a color image formed by overheating after laser exposure, and a method of forming an image only by heating. For image formation of a hard copy, dyes and pigments are preferable other than those mainly using silver. Also in this case, it is preferable to form an image with a black tone obtained by a silver halide image, and it is preferable that the arrow or the patient name displayed in the image is colored not red and yellow, but blue instead of black and white. . In addition, it is preferable to color the region of interest and the edge emphasized by the phase contrast. When an image is formed with a dye or pigment, a resin capsule containing a coloring component such as a dye or pigment is broken by overheating, and a compound that reacts with the coloring component is dispersed in a medium in which the capsule is dispersed. Image formation in which the destroyed portion develops color is preferred. In addition, there is a method of forming an image by sublimating the ink from a sheet coated with the ink by overheating and transferring the ink to a medium such as a resin, or forming an image by ejecting the ink from a nozzle. At the time of this image formation, it is possible to process the image signal input to the image output device, perform enlargement / reduction, dynamic range compression, image edge enhancement processing, etc., and then output the image. The pixel size of the output image is preferably equal to or an integer multiple of the image input pixel size. It is preferable that the size of the pixel is high definition from 300 μm to 40 μm and further to 10 μm.
[0104]
When displaying a phase-contrast X-ray image by soft copy, a self-luminous CRT or plasma display display, or a liquid crystal or e-ink display that displays an image by adjusting the transmitted light from a light emitting source can be used. . In the case of soft copy display, the screen brightness needs to be 400 nit or more, and preferably 600 nit or more. The contrast ratio is preferably 600 or more and 10,000 or less. The display pixel size is preferably from 300 μm to 40 μm. The image display may be for monochrome or color images. When displaying this image, it is possible to process the image signal input to the image output device, perform enlargement / reduction, dynamic range compression, image edge enhancement processing, and the like, and then output the image.
[0105]
Next, the X-ray imaging system will be described.
[0106]
The phase contrast X-ray imaging system is a medical X-ray imaging system that does not use conventional X-ray refraction. Therefore, the present system requires devices and equipment capable of capturing a phase contrast image.
[0107]
That is, in the case of general imaging, as shown in FIG. 11, a subject fixture 12 for fixing a patient is required between the X-ray tube 10 of the X-ray source and the X-ray image detector 11. The infrared light position detector 20 detects the distance R1 between the X-ray tube 10 and the X-ray image detector 11, and the operating device 22 detects the drive motor 21. To move the X-ray tube 10 on the distance marking rail 23. The operation device 22 includes an image display monitor 22a, a keyboard 22b, a control device 22c, and an X-ray source control device 22d, and outputs an X-ray image to the laser imager 24.
[0108]
In this X-ray imaging system, the size of the subject in the output image can be easily known by putting the scale of lead in the subject fixture 12. In addition, it is preferable that a material that hardly transmits X-rays is provided under the subject fixture 12 to prevent unnecessary exposure of the subject to the patient. This is a case in which the imaging is performed in the upright position. In the case of imaging in the lying state, that is, in the lying position, the X-ray source, the fixture, and the X-ray image detector 11 are installed in the vertical direction, and the X-ray source is Located above or below. At this time, oblique X-ray photographing is also performed for photographing the head image. These three elements, ie, the X-ray source, the fixture, and the X-ray image detector 11 can be incorporated and fixed in one device, or they can be separately installed and separately arranged as appropriate. When installed separately, the X-ray image detector 11 can have a side length of 50 cm or more. In such a large size, a column supporting the X-ray image detector 11 should be used to prevent vibration and the like. Is preferably plural. In phase contrast imaging, since the distance R2 between the subject fixed to the subject fixture 12 and the X-ray image detector 11 is set to 0.15 m or more, the X-ray grid conventionally used in close contact imaging is used. No need.
[0109]
In the case of mammography, a device in which an X-ray tube 10 of an X-ray source and an X-ray image detector 11 are fixed as a pair as shown in FIG. 12 is called a so-called C-arm or U-arm. General. This is because the distance between the X-ray source and the X-ray image detector 11 is fixed at 0.6 m to 0.65 m.
[0110]
The distance stamping column 31 rotates around the rotation column 30 as a fulcrum, and the distance stamping column 31 is provided with a grip bar 32 of the subject. An electric resistor 33 is provided on the distance marking support 31, and the position of the subject fixture 12 is detected by an electric resistance position detector 34.
[0111]
When taking a breast X-ray image having a phase contrast, it is necessary to separate the subject from the X-ray image detector 11, and it is desirable that the distance between the X-ray image detector 11 and the subject is large. Further, the imaging device needs to include an X-ray tube 10 as an X-ray source, a subject fixing device 12 as a breast fixing device, and an X-ray image detector 11, and is operated by an operation device 22. The operation device 22 includes a setting condition display panel 22e, a keyboard 22b, a control device 22c, and an X-ray source control device 22d.
[0112]
Therefore, unlike the conventional breast imaging apparatus, the imaging apparatus does not fix the X-ray image detector 11 at 0.6 m or 0.65 m. For example, even if the X-ray source and the subject fixture 12 are fixed at the conventional distance R1 of 0.6 m or 0.65 m, in phase contrast imaging, the X-ray image detector 11 sets the distance R2 from the subject to 0.15 m. It is necessary to separate more. When the X-ray image detector 11 is separated from the subject in this way, it is not necessary to use the X-ray grid conventionally used in close contact imaging. Further, in the conventional mammography apparatus, the patient is photographed in a standing position or in a sitting position such as in a wheelchair. However, the patient can be photographed in a lying position. In mammography, phase contrast X-ray imaging includes CC imaging in which X-rays penetrate vertically from the patient's head and MLO imaging in which X-rays are radiated obliquely.
[0113]
In the phase contrast imaging, since the distance R2 is set between the patient who is the subject and the X-ray image detector 11, it is preferable to provide a jig so that unnecessary things do not enter the path along which the X-ray flies. It is preferable that the distances between the X-ray tube 10, the subject fixture 12, and the X-ray image detector 11 are automatically measured. It is preferable to calculate an enlargement ratio based on the distance data and use the calculated magnification ratio for determining X-ray imaging conditions. A terminal for measuring the X-ray intensity is installed in the X-ray image detector 11 and the subject fixture 12, and the imaging conditions can be determined based on the values. Also, in order to distinguish between phase contrast imaging and conventional contact imaging, the X-ray generator is equipped with a device that does not emit X-rays if the imaging condition is set to contact imaging when the distance relationship is such that phase contrast imaging is performed. Is preferred.
[0114]
When an image is captured by the SF system, a cassette equipped with a silver halide film and an X-ray screen is used as an X-ray image detector. After taking the X-ray image, the silver halide film is taken out of the cassette and developed by an automatic developing machine or the like. An image is drawn on the obtained processed film, and the image is observed on a light box (Schaukasten) equipped with a fluorescent lamp or the like. After the image is observed, the image may be kept in a predetermined place for a certain period of time, and may be taken out and observed as needed. Further, the image information can be digitized by a film digitizer, and the image can be processed to display the image separately, and the image information is stored as digital information.
[0115]
When photographing with a digital system, in the case of CR, a cassette containing a stimulable phosphor plate is used as an X-ray image detector, and a CR device integrated with a stimulable plate reader is used as an X-ray image. May be used as a detector. When the cassette is used, the X-ray image is read by setting the cassette in the image reading device after photographing. It is used particularly when an integrated photographing device such as a C-arm or a U-arm is used. In the case of the photostimulable plate reader integrated type, the X-ray source and the subject fixture are separately installed. When the FPD is used, there are a case where the X-ray source-subject fixture-FPD is integrated with the three elements, and a case where the X-ray source-subject fixture-FPD is installed independently and photographed. At this time, the FPD can be used as an automatic X-ray exposure device.
[0116]
As shown in FIG. 13, the X-ray image information obtained by the digital X-ray image detector of the phase contrast X-ray imaging apparatus 90 is subjected to image processing, CAD processing if necessary, The data is output to a soft copy or a hard copy for display, image signal storage, etc., and is provided for image diagnosis. The image information can be transmitted to a local network in a hospital facility, a network line with another medical facility or research facility, or an Internet line, and the image can be diagnosed in a remote place where the image information is transmitted. When transmitting such image information as digital information, it is preferable to conform to a protocol such as the DICOM standard.
[0117]
When digital image information is displayed in soft copy or hard copy, it is necessary for diagnosis such as patient data and medical history, medical information, distance R1 and distance R2, magnification ratio, type and name of X-ray image detector used, etc. It is preferable to output the information together with information necessary for medical affairs such as medication, address, and name. Further, it is preferable to output not only a phase contrast image but also any medical image such as a past X-ray image or a close contact imaging image, an X-ray CT image, an MRI image, an endoscope image, a fundus photograph image, or the like. In particular, when displayed as a soft copy, it is preferable to use a plurality of screens to display all image information and the like necessary for diagnosis. After the diagnosis is completed, it can be output as a hard copy and saved, or the hard copy can be provided to a necessary department. In addition, an image signal can be stored in a magneto-optical recording material or the like as an electric signal and read out when necessary to reproduce an image.
[0118]
The present invention can be applied to the techniques disclosed in JP-A-2001-238871, JP-A-2002-102215, and JP-A-2002-162705.
[0119]
【The invention's effect】
As described above, the inventions according to claims 1 to 16 are extremely useful in providing high-quality diagnostic images in a wide range of medical facilities.
[0120]
In particular, in imaging using a contrast agent, the amount of use of a dangerous contrast agent can be reduced, and this can be applied to panoramic imaging in dentistry. In addition to such conventional medical images for diagnosis, it can be used for sample tests and the like, and can be applied to nondestructive tests. As an application of a medical image, the present invention can be applied to not only a still image but also a medical moving image. This phase contrast image capturing technique can also be applied to X-ray CT capturing.
[Brief description of the drawings]
FIG. 1 is a diagram illustrating refraction of an electromagnetic wave.
FIG. 2 is a diagram for explaining that an X-ray beam narrowed down into a point or a line is obliquely incident on a silicon single crystal surface to obtain a monochromatic X-ray beam by Bragg reflection.
FIG. 3 is a diagram for explaining that a phase contrast image is obtained by separating a subject from an X-ray detector without using an analyzer to enlarge a shift caused by small refraction of X-rays.
FIG. 4 is a diagram illustrating phase contrast imaging using a microfocus X-ray tube and not using Bragg reflection.
FIG. 5 is a diagram illustrating that when a focal size is large, a phase contrast image cannot be obtained due to blurred penumbra caused by the focal size.
FIG. 6 is a view for explaining that a coherence in the horizontal direction cannot be obtained with a molybdenum anode thermoelectron tube.
FIG. 7 is a diagram illustrating the strength of phase contrast.
FIG. 8 is a diagram illustrating that when a thermionic X-ray tube having a finite focal size is used, geometric instability due to blurring of penumbra occurs.
FIG. 9 is a diagram for explaining that a distance between a peak and a valley is widened by a phase contrast being on a blur due to geometrical unsharpness, and that an edge effect can be detected.
FIG. 10 is a diagram showing a simulation result of an X-ray intensity profile of edge enhancement.
FIG. 11 is a diagram illustrating an X-ray imaging system.
FIG. 12 is a diagram showing another embodiment of the X-ray imaging system.
FIG. 13 is a diagram illustrating an output of the X-ray imaging system.
[Explanation of symbols]
10. X-ray tube of X-ray source
11 X-ray image detector
12 subject fixture
20 Infrared light position detector
21 Drive motor
22 Operating device

Claims (16)

Enlarged radiography is performed using non-coherent X-rays radiating radially from an aperture of finite size, and a phase contrast is generated on a blurred penumbra on the image plane of the magnified radiography. X-ray imaging method. When the distance (R1) from the X-ray focus to the subject is 0.2 m or more and 2.5 m or less, and the wavelength (λ) of the X-ray used is 8 pm to 71 pm, 0.16 × 10 ⁶ × λL <r < An X-ray image detector having a spatial resolution of 10 ⁻⁵ m or more using a thermionic X-ray tube having a focal size of 2r of 2.5 × 10 ⁻⁴ m (250 μm). Item 1. The X-ray image capturing method according to Item 1. 3. An X-ray imaging method according to claim 1, wherein the X-ray image detector is a screen film system or a digital X-ray imaging system. When the distance (R1) from the focus of the X-ray to the object is 0.2 m or more and 2.5 m or less, and the wavelength (λ) of the X-ray used is 8 pm to 71 pm, 0.16 × 10 ⁷ × λL <r < It is characterized by using an X-ray image detector having a spatial resolution of 5 × 10 ⁻⁵ m or more using a thermionic X-ray tube having a focal size of 2r of 2.5 × 10 ⁻⁴ m (250 μm). The X-ray imaging method according to claim 1. The X-ray imaging method according to claim 3, wherein the X-ray image detector is a digital X-ray imaging system. The width of the blur on the image plane is B, and the width of the edge enhancement caused by the phase contrast is E, 9E ≧ B, wherein 6E ≧ B. X-ray imaging method. Using the thermionic X-ray tube having a focal size of 50 μm or more and 500 μm or less, a distance (R1) from the X-ray focus to the subject is 0.2 m or more and 2.5 m or less, and the distance from the subject to the X-ray image detector is The X-ray imaging method according to any one of claims 1 to 6, wherein the distance is 0.15 m or more and 2.5 m or less. When performing magnification photography using non-coherent X-rays radiating from an aperture of a finite size, an X-ray image in which the edge enhancement width caused by the phase contrast is magnified by the blur of a semi-shadow is converted to a digital X-ray image. 8. An image taken by a detector and displayed in a size smaller than the image size of the subject on the surface of the digital X-ray image detector, or the actual size or a size close to the actual size. 3. The X-ray image capturing method according to 1. Enlarged radiography is performed using non-coherent X-rays radiating radially from an aperture of finite size, and a phase contrast is generated on a blurred penumbra on the image plane of the magnified radiography. X-ray imaging system. When the distance (R1) from the X-ray focus to the subject is 0.2 m or more and 2.5 m or less, and the wavelength (λ) of the X-ray used is 8 pm to 71 pm, 0.16 × 10 ⁶ × λL <r < An X-ray image detector having a spatial resolution of 10 ⁻⁵ m or more using a thermionic X-ray tube having a focal size of 2r of 2.5 × 10 ⁻⁴ m (250 μm). Item 10. An X-ray imaging system according to Item 9. The X-ray imaging system according to claim 9 or 10, wherein the X-ray image detector is a screen film system or a digital X-ray imaging system. When the distance (R1) from the focus of the X-ray to the object is 0.2 m or more and 2.5 m or less, and the wavelength (λ) of the X-ray used is 8 pm to 71 pm, 0.16 × 10 ⁷ × λL <r < It is characterized by using an X-ray image detector having a spatial resolution of 5 × 10 ⁻⁵ m or more using a thermionic X-ray tube having a focal size of 2r of 2.5 × 10 ⁻⁴ m (250 μm). The X-ray imaging system according to claim 9. 13. The X-ray imaging system according to claim 11, wherein the X-ray image detector is a digital X-ray imaging system. 14. The image processing apparatus according to claim 9, wherein 9E ≧ B, where B is a width of the blur on the image plane, and E is a width of edge enhancement generated by the phase contrast. 15. X-ray imaging system. Using the thermionic X-ray tube having a focal size of 50 μm or more and 500 μm or less, a distance (R1) from the X-ray focus to the subject is 0.2 m or more and 2.5 m or less, and the distance from the subject to the X-ray image detector is The X-ray imaging system according to any one of claims 9 to 14, wherein the distance is 0.15 m or more and 2.5 m or less. When performing magnification photography using non-coherent X-rays radiating from an aperture of a finite size, an X-ray image in which the edge enhancement width caused by the phase contrast is magnified by the blur of a semi-shadow is converted to a digital X-ray image. 16. An image taken by a detector and displayed in a size smaller than the image size of the subject on the surface of the digital X-ray image detector, or the actual size or a size close to the actual size. 2. The X-ray image capturing system according to 1.

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