JP2004002746A - Needle-shaped columnar storage phosphor crystal - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a phosphor crystal having a dimension suitable for getting the optimum relationship between the light transparency and the resolution or sharpness in the use of the phosphor crystal as a flat storage phosphor plate or panel and provide a computerized radiophotography system forming an image in high signal/noise ratio even in the case of using a low X-ray dose having low energy for mammography, etc. <P>SOLUTION: The needle-shaped columnar CsBr:Eu<SP>2+</SP>storage phosphor particle has an average cross-section diameter of 1-30μm and an average length of 100-1,000μm measured along the column of the crystal. The invention further provides a radiation storage phosphor screen or panel produced by using the storage phosphor particle. <P>COPYRIGHT: (C)2004,JPO

Description

[0001]
Field of the invention
The present invention relates to storage phosphors having dimensions that give excellent resolution after being stimulated by stimulating radiation.
[0002]
In particular, the present invention relates to the packing density of the acicular storage phosphor in the form of a cylinder after being coated, the role of the dye (in the stacked layer of said acicular storage phosphor or in the vicinity of said layer). If) and the need to avoid the presence of alkaline earth metals, rare earth metals and impurities such as oily compounds or debris.
[0003]
Background of the Invention
Radiation image storage, referred to as stimulable phosphor screens, sheets or panels, in contrast to conventional radiography in which a sensitized luminescent phosphor screen emits luminescent lines directly and said screen is not a storage medium Radiation image recording and reproduction techniques utilizing panels have been provided with stimulable phosphors. Using radiation image recording and reproduction techniques, the stimulable phosphor of the radiation image storage panel carries image information of the subject or absorbs radiation emitted from the sample. The stimulable phosphor is exposed to a stimulating ray, such as visible light or infrared light, which causes the stimulable phosphor to emit light in proportion to the amount of energy stored thereon during the irradiation exposure.
[0004]
The emitted fluorescence is then detected opto-electrically to obtain an electrical signal. The electrical signal is further processed and the processed electrical signal is used to reproduce a visible image on a recording material. This method of operation, utilizing a storage phosphor sheet or panel as an intermediate storage medium, is also referred to as "computer radiography". This working method is CsI, KI, RbI, CdS, Zn_xCd_1-xS, CdWO₃, Ga_ySe, Gd₂O₂S, La₂O₂Luminescent materials such as S or PbOz are distinctly different from radiation detection by scintillator materials which emit radiation directly upon irradiation with X-rays, for example as described in US-A-5171996.
[0005]
In radiography, it is important for a radiologist to have good image quality in order to make an accurate assessment of a patient's condition, so the important aspects of image quality are resolution and image signal to noise ratio.
[0006]
For computer radiography, the signal to noise ratio depends on a number of factors.
[0007]
First, the number of X-ray quanta absorbed by the storage phosphor screen is important. The signal-to-noise ratio will be proportional to the square root of the number of quantum absorbed.
[0008]
Second, so-called fluorescent noise is important. This noise is caused by the fact that the number of photostimulated light (PSL) quanta detected for the absorbed X-ray quanta is small. Fluorescence noise has a significant contribution to the signal-to-noise ratio because a lot of PSL light is lost in the detection process in computer radiography. It is important that on average at least one photon is detected for each absorbed X-ray quantum. If not, many absorbed X-ray quanta will not contribute to the image and the signal-to-noise ratio will be very poor.
[0009]
This situation is most critical in mammography where X-ray quanta are used at low energies. Soft X-rays will produce fewer PSL centers (PSL @ centers) and therefore will produce fewer PSL photons than hard X-rays.
[0010]
In computer radiography, the number of PSL centers is created by absorbed X-ray quanta. Not all PSL centers are stimulated in the readout process due to the limited time available for pixel stimulation and the available laser power. In practice, only about 30% of the PSL centers are stimulated to produce PSL photons. As these photons are emitted and scattered in all directions, only 50% of the PSL photons are emitted above the storage phosphor screen, where they can be detected by the detection system. The emitted PSL photons are guided by a light guide towards a detector. This light guide may consist of an array of optical fibers forming a rectangular detection area on the storage phosphor screen and having a circular cross section on the detection side. Light guides of this type have only a 30% numerical aperture, which means that only one out of three emitted PSL photons is guided to the detector. A filter is placed between the light guide and the detector, which blocks the stimulating light reflected by the storage phosphor screen and transmits the PSL light emitted by the screen. This filter also has small absorption and reflection of PSL light and transmits only about 75% of the PSL photons. In computer radiography, photomultipliers are commonly used to transform PSL signals into electrical signals. At 440 nm, the photomultiplier has a quantum efficiency of about 20%. This means that only one of the 5 PSL quanta reaching the photomultiplier is detected.
[0011]
In short, for a 1000 PSL center made with a storage phosphor screen only:
1000 × 0.3 × 0.5 × 0.3 × 0.75 × 0.2 = 6.75
Of PSL photons are detected.
[0012]
If it is required to produce at least one detected PSL photon per X-ray quantum, the number of PSL centers created by the X-ray quantum should be large enough. Alternatively, conversely, the X-ray energy required to create a PSL center should be small enough.
[0013]
In mammography, the usual setting of the X-ray source is 28 kVp. This leads to an X-ray spectrum in which the average energy of the X-ray quanta is of the order of 15 keV. For an X-ray quantum having this energy, to generate at least one detected PSL photon, the energy required to create the PSL center is:
15000 × 6.75 / 1000 = 100 eV
Should be less than.
[0014]
It is well known that many X-ray medical imaging systems, especially mammography, require fine detail visualization and high resolution and high contrast images. The resolution of X-ray film / screen and digital mammography systems is currently limited to 20 line pairs / mm and 10 line pairs / mm, respectively. One of the main reasons for this limitation relates to the phosphor particle size in currently used X-ray screens. In particular, light scattering by the phosphor particles and their particle boundaries results in loss of spatial resolution and contrast in the image. In order to increase resolution and contrast, scattering of visible light must be reduced. Scattering can be reduced by reducing the phosphor particle size while maintaining the phosphor's emission efficiency. In addition, X-ray to light conversion efficiency, quantum detection efficiency (e.g., the portion of absorbed X-rays that can be converted to light emitted after stimulation) and screen efficiency (e.g., emission that escapes from the screen after irradiation with stimulation light) (The fraction of light emitted) should not be negatively affected by the reduction in phosphor particle size. Computer radiographic recording and reproduction techniques, given below as special benefits, show radiological images containing a large amount of information obtained at significantly lower radiation doses than in conventional radiography. Radiation image recording and reproduction techniques are particularly efficient for direct radiography, such as X-ray image recording for medical diagnostics.
[0015]
For clinical diagnosis and routine screening of an asymptomatic female population, screen-film mammography still represents the state of the art for early detection of breast cancer. However, screen-film mammography has limitations that reduce its effectiveness. Due to the very low difference in radiation absorption density in breast tissue, the image contrast is inherently low. Film noise and scattered radiation further reduce contrast, making detection of microcalcified images difficult in displayed images. Thus, for example, film gradations must be balanced against the need for a wide latitude.
[0016]
Computer radiography (CR) systems can be broadly classified as first digital systems and second digital systems. The first digital system refers to the direct conversion of incident radiation on a sensor into usable electrical signals to form a digital image. On the other hand, the second digital system includes an intermediate step in the conversion of radiation into a digital image. For example, in digital fluoroscopy, an image intensifier is used to intermediately convert X-rays to a visible image, which is then converted to a digital image using a video camera. For example, the scintillator material described in US-A-5171996 does not require a scanning procedure to read out. Because there is no storage of energy as in stimulable or storage phosphor materials. The X-ray quanta are converted directly into the emitted light, which further creates a charge in the photoconductive layer, thereby providing a visible image. The scatter of the directly emitted light is sharper than the sharpness of the storage phosphor used in digital X-ray systems, where the scatter of the stimulating light is very important as a source of photostimulation of the stored energy. Is decisive. Digital X-ray images generated by systems utilizing photostimulated luminescence (PSL) plates first store the visible image as energy. In a second step, the stored energy is converted to electrical signals and a laser is used to scan the PSL plate to form a digital image.
[0017]
In addition, various schemes using silicon photodiode arrays in scan mode for CR systems are used. However, these photodiode arrays require an intermediate phosphor screen to convert X-rays into visible light due to the sharp drop in quantum efficiency (sensitivity) of the array at energies of 10 keV and higher.
[0018]
The second digital system has several disadvantages, including loss of resolution. However, recent technological advances can overcome these difficulties by enabling semiconductor X-ray detectors to generate usable X-ray images. Although high quality solid state x-ray detectors have been known for many years, these detectors require a very sensitive preamplifier to produce a signal suitable for use. Due to recent advances in high density analog complementary metal oxide semiconductor (CMOS) integrated circuit technology and high density interconnects between semiconductor chips, those having a preamplifier on a single hybrid integrated circuit called a sensor chip. Thousands of sensing element integrations are currently possible. A semiconductor detector having an X-ray semiconductor sensor and an absorption layer located between X-rays from a subject is disclosed, for example, in US Pat. No. 4,905,265.
[0019]
Significant improvements in clinical image quality have been realized to eliminate the need for repeated exposure due to poor film image quality caused by factors such as radiation scatter noise, fog, blur, speckles and artifacts, where Digital radiography techniques allow physicians to perform quantitative radiography by image digitization, which can be performed by useful augmentation techniques such as transmitting mammograms to remote sites on a computer network and edge enhancement of microcalcified images. A group of 35 years or older by advantageously reducing the absorbed radiation dose received by the patient at least 7 times compared to screen-film mammography and further significantly increasing the benefit to risk equalization Mammography for routine screening of asymptomatic populations in Japan To be significantly reduced an additional 10 times absorbed dose to the patient the needle local production in Kenho that may require exposure. Another advantage of the invention, as described in U.S. Pat. No. 5,596,200, is to provide improved storage and retrieval of image data through the use of standard magnetic or optical disk media instead of photographic film, and Competitive with current X-ray imaging systems in capital costs, reducing material costs and processing time by eliminating photographic film and related chemicals, darkrooms and other peripherals and reducing technician time for film processing It is to provide a device. In determining the desired semiconductor material, aspects such as manufacturability, X-ray absorption, and operating temperature must be considered. For mammography applications, two alternative detection materials, silicon and gallium arsenide, are preferred. Silicon detectors are much easier to manufacture than GaAs detectors, but the silicon X-ray photon quantum absorption coefficient is much lower than GaAs. For applications in primary X-ray digital imaging systems having X-ray energies greater than about 25 keV, sensor materials with much higher X-ray absorption properties are required. As a result, GaAs, cadmium telluride, CdZnTe, indium antimonide and germanium sensing materials should be used at energies greater than 25 keV. The number of rows and columns of detectors and their length and separation can be further modified by the specific design conditions of the x-ray imaging system. For example, the length of the rows and the number of rows can be any desired value below the limit of the mechanical scan. It is also conceivable to place the sensor chips in an array-like manner.
[0020]
Since the generated image is isomorphic to the matrix of Arabic numerals generated during scanning, it can be processed by the signal processing unit 60 with suitable software. For example, the signal-to-noise ratio of the signal can be improved through processing. The data from the signal processing is advantageously stored and stored on standard magnetic or optical disk media instead of photographic film. The data from the storage station is then sent to an image processing unit, where various processing operations can be performed on the image. For example, the image processing unit may include (1) enlargement; (2) contrast enhancement and window generation; (3) sharpness and edge gradation enhancement; (4) gray or color scale attachment for enhanced image quality; (5) Image operation of image subtraction can be performed. The image produced by the image processing can be displayed on a video display, printer, film, or sent via an image transmission network, which provides substantially real-time image interpretation and diagnostics. A satellite or computer network for transmitting image data from a remote radiation laboratory to a central radiologist can be included. Applications in other areas of clinical imaging include, for example, low cost, low cost applications in breast computed tomography (CT); use in intelligent software for computer-aided diagnostic systems (CAD); stereotactic computer settings of biopsy needles; It is possible as a radiation control, visual and non-invasive imaging system for applications in nuclear medicine. An apparatus for imaging a patient's breast by scanning the imaging signal and receiver across the patient's breast and then constructing a scan-based time-delay integrated compound image is described in US-A 5,526,394. . The receiver includes an array of radiation sensitive detection elements, where the readout of the array is synchronized with the scanning motion of the receiver based on the output from the position encoder, so that synchronization is maintained regardless of scan drive changes. An apparatus for allowing selection of an appropriate radiation filter based on specific imaging conditions is also disclosed therein.
[0021]
US Pat. No. 6,300,640 specifically uses a nanocrystal-sized phosphor (nanophosphor) arranged in very small grooves (microgrooves) etched in a support for detecting radiation, especially X-rays. And a composite phosphor screen.
[0022]
Further improvements utilizing the particular detector, as always desired in x-ray imaging, especially in mammography applications, are described below.
[0023]
Summary of the Invention
It is an object of the present invention to optimize optimized optical properties, more particularly between light transparency and resolution, i.e. sharpness, when using phosphor crystals in flat storage phosphor plates or panels. It is to provide a phosphor crystal having the desired dimensions to obtain the relationship.
[0024]
It is a further object of the present invention to provide a computer radiography system which allows imaging with a good signal-to-noise ratio even when using X-ray quanta having low energies as in mammographic imaging. That is.
[0025]
The above objectives are measured along an average cross-sectional diameter in the range 1 μm to 30 μm (more preferably 2 μm to 15 μm), and a cylindrical casing in the range 100 μm to 1000 μm (more preferably 100 μm to 500 μm). CsBr: Eu in the form of a cylinder having an average length²⁺This is achieved by providing storage phosphor particles.
[0026]
Particular features of preferred embodiments of the invention, in particular with regard to the size of the acicular phosphor crystals, are set out in the dependent claims.
[0027]
Further advantages and embodiments of the present invention will become apparent from the following description and drawings. The above-mentioned advantageous effects are realized by having the special features stated in claim 1. Particular features of preferred embodiments of the invention are set out in the dependent claims. Further advantages and embodiments of the present invention will become apparent from the following description (and drawings).
[0028]
Brief description of drawings
FIG. 1 shows a photograph of a typical acicular storage phosphor panel having phosphor particles in the form of cylinders, with dimensions indicated.
[0029]
Detailed description of the invention
Preferred CsBr: Eu in cylindrical form²⁺As a starting point for the resolution obtained from needle-shaped phosphors such as storage phosphors or stimulable phosphors, needle-shaped crystals with a very small cross section cannot pass the stimulated radiation obtained during stimulation, while It has been experimentally demonstrated that needle-shaped crystals having a large diameter allow radiation to pass easily, but lead to sharp images. Therefore, needle-shaped CsBr: Eu in the form of a cylinder²⁺Storage phosphor particles are selected, said phosphor particles having an average cross-sectional diameter in the range of 1 μm to 30 μm to provide sufficient sharpness, more preferably 2 μm to 15 μm, 100 μm to 1000 μm to provide sufficient speed, More preferably, it has an average length measured along the cylindrical portion of the cylinder in the range of 100 μm to 500 μm.
[0030]
One of the major challenges in reducing the particle size below 1 μm is in the sharp decrease in phosphor luminous efficiency. This is due to a surface-related non-radiative process that is predominant in the region from 1 μm to 0.01 μm: as a theoretical boundary the diameter should never be smaller than 440 nm, representing the actual wavelength, and therefore well deviate from that wavelength The choice of 1 μm is justified as the boundary value to be used.
[0031]
There should also be a sufficient number of needle-like crystals so that the needle-like structures are drawn. This is because interference at the pixel size is not excluded. In mammography, for example, a pixel size of 60 μm is usually used so that a diameter of 30 μm gives one half of a pixel for one needle crystal. Four needles are provided per pixel, which is consistent with the minimum at which "aliasing" occurs. In a further preferred example according to the invention, the needle-shaped phosphor should have an average diameter in the range from 2 μm to 15 μm.
[0032]
During the manufacture of needle crystals by the preferred chemical vapor deposition (CVD) technique, it is observed that a slightly amorphous layer (amorph) deviating from the needle structure is formed in contact with the support: This layer referred to has an average thickness of about 30 μm, depending on the deposition parameters in the CDV process. In order to obtain a working system that exhibits the most favorable compromise between sensitivity and signal-to-noise ratio, it is recommended to provide needles having a needle length greater than at least twice the thickness of the amorphous nucleation layer. You. As a result, 100 μm is almost the boundary. Because smaller needle crystal lengths lack sensitivity and sufficiently high signal-to-noise values.
[0033]
Further, a needle crystal length greater than 1000 μm is not desirable because the intensity of the laser beam is weakened and cannot reach the bottom of the needle crystal. Needle layers exceeding 1000 μm cannot be read out and do not contribute to the desired signal. Many of the light emitted by the stimulated radiation will not be able to escape from the acicular layer having a thickness greater than 1000 μm. In a further preferred example according to the invention, the needle-shaped phosphor should have an average length in the range from 100 μm to 500 μm.
[0034]
Further, CsBr: Eu²⁺A further preferred (dimensionless) parameter for the stored needle phosphor population is the average ratio of length to cylinder diameter in the needle population, which ranges from 5 to 200, more preferably from 20 to 100. is there. To calculate this range, the average length of the acicular crystals is calculated similarly to the average length of a number of acicular phosphor cylinders, and a corresponding figure is obtained from both ratios.
[0035]
The walls of the cylinder acting as microgrooves are arranged to reflect the light emitted by the nanophosphor below the microgroove cylinder to a suitable light collection device, such as an electronic device. It is recommended to provide a wall with a smooth surface within the cylindrical coating, since the light is well guided without extreme aberrations or deflections.
[0036]
CsBr: Eu²⁺Another important parameter that characterizes the stored needle phosphor population is the coefficient of variation for the average cross-sectional diameter and the average barrel length, which indicates the efficiency and sharpness of light output. In a preferred example, the coefficient of variation is less than 0.30, in a more preferred example less than 0.20, most preferably a maximum of 0.10.
[0037]
Acicular CsBr: Eu according to the present invention²⁺17 mg / cm storage phosphor particles²~ 400mg / cm²Preferably, the storage phosphor plate or panel according to the present invention is coated in an amount for said storage phosphor particles in the range of.
[0038]
Acicular CsBr: Eu having dimensions as stated in the description of the invention²⁺Storage phosphors have high conversion efficiencies, so that low amounts of X-ray energy are required to create PSL centers. Phosphors have been invented in which the amount of X-ray energy required to create a PSL center is less than 100 eV and further has the desired dimensions as set forth in the claims. Furthermore, a manufacturing method is provided which leads to this phosphor with high quality.
[0039]
According to the invention, a storage phosphor screen or panel is provided, said screen or panel emitting radiation after stimulation of phosphor particles having stored energy from X-rays impinging on said screen or panel, said panel or panel comprising: Is:
A support having a flat surface;
-A number of cylindrical needle-shaped phosphors extending on the surface of the support, said phosphors being needle-shaped needle-shaped CsBr: Eu²⁺Storage phosphor particles).
[0040]
In a preferred embodiment, the multiple needle phosphors form a binderless phosphor layer.
[0041]
As a basic structure, a radiation image storage panel comprises a support and a stimulable phosphor layer lying on one surface of the support, unless the stimulable phosphor layer is self-supporting. If the stimulable phosphor layer is self-supporting, the radiation image storage panel need not necessarily be provided with a support. In another example, the acicular phosphor is deposited on a (structured or unstructured) support layer, such as when the support is comprised of so-called "hollow fibers". Generally, the stimulable phosphor layer comprises a binder layer and a layer containing the stimulable phosphor dispersed in the binder. However, after being deposited by the preferred chemical vapor deposition technique, the needle-shaped columnar phosphors having dimensions as in the present invention are stacked vertically on a primed support in a regularly arranged manner as a deposited layer, wherein said ordered The ordered layers are not predetermined by a pre-formed pattern as in US-A-5171996.
[0042]
Radiation image storage panels provided with acicular phosphors having dimensions as in the present invention have, apart from the support adjacent to the support material, a reflective or antireflective layer and / or a dye-containing antihalation layer. Is also good. The layer having antireflection properties is, for example, MgF₂However, the present invention is not limited thereto.
[0043]
The radiation image storage panel according to the present invention is further described below by taking a radiation image storage panel having a typical configuration comprising a support, optionally a light reflecting layer and / or a dye-containing antihalation layer, and a stimulable phosphor layer. Will be described.
[0044]
The support may be a film of a plastic material such as cellulose acetate, polyester, polyethylene terephthalate, polyamide, polyimide, triacetate, polycarbonate, Kevlar®, a metal sheet such as aluminum foil and aluminum alloy foil, plain paper, barita paper, resin. Consisting of materials selected from various types of support materials commonly used in known radiation image storage panels, such as coated paper, pigmented paper containing pigments such as titanium dioxide, and paper sized with polyvinyl alcohol and the like. You may.
[0045]
Alternatively, the support may be a glass support (including colored or uncolored) or a heat-resistant synthetic resin material, and may be provided with a deposited metal layer. In this connection, the materials commonly used for electronic circuits (PCBs) are very suitable for use as supports. By way of example, a synthetic material called PERTINAX® (on which a layer of copper having a thickness in the range of 10 μm to 200 μm is deposited) is very suitable for use. This synthetic material resists destruction by heating up to temperatures above 250 ° C., thus providing a deposition without degradation.
[0046]
Since the X-ray absorption of such a synthetic resin material is extremely low, an X-ray detector (so-called AEC detector) may be mounted after the synthetic support to measure the X-ray dose. As a result, aluminum is only suitable for use as a metal layer when its thickness is low enough.
[0047]
Copper as a deposited material provides a reflective metal layer with very good adhesion to synthetic support materials that do not exhibit the disadvantageous properties of glass supports such as brittleness and flare.
[0048]
In order to improve the adhesion between the phosphor layer of the support of the radiation image storage panel of the present invention, the intermediate dye-containing antihalation layer, or the formed reflection or antireflection layer, an adhesion-imparting layer is recommended. The layer is composed of a high temperature resistant high molecular weight material used in the needle phosphor deposition process.
[0049]
The light reflection (prevention) layer forms a coating composition containing a light reflection (prevention) substance, a binder, and a solvent, and uniformly applies the coating composition on the surface of the support to form a coating film of the coating composition. May be formed. The binder and the solvent for forming the light reflecting (preventing) layer may be selected from the binders and solvents used for forming the stimulable phosphor layer. Generally, the mixing ratio of the binder to the white pigment, for example, in the coating composition for forming the light reflecting layer may be selected from the range of 1: 1 to 1:50, and the individual figures are expressed as weight ratios. You. From the viewpoint of the reflection (prevention) characteristics of the light reflection (prevention) layer, the proportion of the binder is preferably as low as possible. Binders for the exemplified pigments in the coating composition for the formation of the light-reflective (anti-) layer when considering the ease of forming the light-reflective (anti-) layer and the mechanical and physical strength of the radiation image storage panel Is preferably selected from the range of 1: 2 to 1:20 (expressed as a weight ratio). The thickness of the light reflecting layer is preferably in the range of 5 μm to 100 μm. The coating composition for forming the light reflecting layer may be applied by conventional coating means such as a doctor blade, a roll coater, or a knife coater. After the coating film of the coating composition is formed on the surface of the support, the coating film is heated and dried little by little.
[0050]
Pigments or dyes may be deposited in the interstices between the stimulable phosphor needles along with the stimulable phosphor, provided that dye deposition is possible. Therefore, it is required to have a dye or pigment that does not decompose, and its optical properties, together with the storage phosphor, need not change under the influence of increased temperature during the deposition. In such a case, the ratio (weight ratio) of the stimulable phosphor to the dye or pigment is preferably in the range of 1,000,000: 1 to 1000: 1. If a pigment is introduced into the stimulable phosphor layer together with the storage phosphor, a light reflecting layer reflecting the stimulating rays may be formed on one surface of the stimulable phosphor layer. An example of a dye suitable for use within the scope of the present invention is methylene blue (C₁₆H₁₈ClN₃), Azure B (C_FifteenH₁₇ClN₃S), Toluidine Blue 0 (C_FifteenH₁₆N₃SCl), thionine (C₁₂H₁₀ClN₃S), indocyanine green (IGC) (C₄₃H₄₇N₂O₆S₂Na), magnesium phthalocyanine, oxatricarbocyanine, indotricarbocyanine, zinc phthalocyanine, oxazine, cryptocyanine and tetra-1-butyl-naphthalocyanine, but the most preferred pigment dyes are the most chemically Β-copper phthalocyanine (CuC)₃₂N₆H₁₆) Is selected. Coating of a very thin homogeneously colored layer is achieved without undesired gaps between the acicular phosphor crystals. Preferred dye pigments are blue-tinted, so that a very good absorption of the red laser radiation and a very good transmission of the blue light are possible. The β-copper phthalocyanine is available from Ciba-Geigy, Switzerland.
[0051]
According to the present invention there is provided a method for incorporating a dye or pigment into a needle image phosphor plate. After being immersed in an alcohol solution in which the pigment or dye is dissolved, the phosphor plate is preferably air-dried to evaporate the alcohol present and to immobilize the dye in the phosphor plate. In a preferred embodiment according to the invention, the binder is present in the alcohol dye solution used in the process in a low amount, preferably in an amount of 1% to 10% by weight. The presence of the dye avoids the transfer of light from one needle to another (adjacent) needle and enhances the beneficial effect of the needle structure: although a slight speed reduction is detected, Produces even improved sharpness or resolution. According to the invention, the storage phosphor plate or panel is 1 μg / m²~ 1000 μg / m²Preferably contains one or more dyes.
[0052]
As already mentioned above, it is recommended to provide a wall with a smooth surface within the cylindrical coating. In a preferred embodiment, the walls of the cylinder are colored with a dye or pigment that absorbs impinging stimulating radiation.
[0053]
According to the invention, a storage phosphor screen is provided on the wall of the cylinder, arranged to at least partially (most preferably totally) reflect the radiation emitted by the phosphor after performing the stimulation. A light-reflective coating disposed along it.
[0054]
In another example, the reflective coating is disposed along a wall of the cylinder, more preferably the reflective coating is a metal coating. In the case of a metal coating, a metal selected from the group consisting of gold, silver, platinum, palladium, nickel and aluminum is most preferred.
[0055]
A transparent protective film composed of a plastic material for physically and chemically protecting the stimulable phosphor layer is generally formed on the surface of the stimulable phosphor layer, the surface being opposite to the support side surface. The radiation image storage panel according to the invention is preferably provided with such a transparent protective film. The protective film may be formed on the stimulable phosphor layer using, for example, a technique in which a plastic film is preformed and then adhered to the surface of the stimulable phosphor layer with an adhesive. Alternatively, the protective film may be formed on the stimulable phosphor layer using a technique in which the protective film material containing the coating composition is applied onto the surface of the stimulable phosphor layer and then dried.
[0056]
Fine particle fillers may be included in the protective layer to reduce interference non-uniformity and to enhance the image quality of the radiation image. Examples of suitable resins for the production of light transmissive plastic films include polyester resins such as polyethylene terephthalate and polyethylene naphthalate; and cellulose ester derivatives such as cellulose triacetate. Various resin materials such as polyolefins and polyamides may be used for the production of the light-transmitting plastic film. Preferably, the thickness of the protective film is in the range of approximately a few μm (eg 3 μm to about 20 μm for sharpness). An anti-reflective layer, eg MgF, adjacent to the protective coating₂May be further added.
[0057]
Preferably, the storage phosphor screen comprises radiation attenuating means coated on at least a portion of the acicular cylindrical phosphor particles arranged to selectively attenuate radiation. In a preferred embodiment, the screen is provided with radiation attenuating means comprising a coating arranged in the microgrooves.
[0058]
The storage phosphor screen according to the invention further comprises means for collecting the radiation emitted by the storage phosphor upon stimulation.
[0059]
Many attempts have been made to manufacture digital medical X-ray imagers, as is well known. The energy distribution of the X-rays in the image plane is determined by the illuminated object, human tissue and bone, resulting in an almost continuous energy spectrum on a spatial and intensity scale. The image is generally a low contrast image where the gray tones need to be protected for an accurate diagnosis. Such X-ray images are generally recorded continuously on film in analog fashion. However, digital recording devices such as the present invention determine the averaged intensity on a photographic element or pixel.
[0060]
The physical detection process and pixel spacing determine the resolution of the device. The intensity at each pixel is typically put into one of a finite number of levels. Thus, the grayscale resolution is limited to this number of levels, rather than being a continuous spectrum value. All known devices are intended to optimize resolution and gray scale performance, but are usually limited to any aspect. Direct conversion devices involve the use of a layer of selenium to attenuate x-rays and create free electron hole pairs for collection by suitable electrodes. They also include other means such as increasing the emitted electrons in a pixelated CCD camera, converting and collecting charge following conversion of X-rays in a solid or gaseous medium. Indirect converters utilize photostimulable phosphors as described above for the conversion of x-rays to visible light, where the photostimulable phosphors image red light for later activation. Store. The phosphor can be "read out" by a scanning laser and emits blue light consistent with the stored image. Further emitted light is collected and detected by the photomultiplier.
[0061]
Apart from coating the pigment dye with the storage phosphor particles as previously disclosed, the pigment dye is applied to a very thin layer on a fiber optic plate (FOP), which preferably comprises clear glass. The collection efficiency of such FOPs, which may be coated and is generally known, is much better than other optical components. As a result, an opportunity is provided to provide the light emitted by the acicular phosphor after stimulation so as to be guided in an ideal manner to the CCD device via such a FOP covered with the pigment dye. The dye pigment can be coated or deposited from an alcohol dye solution as already described above.
[0062]
It is established that the absence of impurities in the needle phosphor and in the corresponding needle phosphor storage panel is highly desired.
[0063]
First, it is desirable that no impurities from the dopant exist. CsBr to be selected: Eu²⁺The phosphor is preferably free of alkali metals (ions), especially those selected from lithium, sodium, potassium and rubidium; alkaline earth metals such as strontium; and thallium. More preferably, samarium is present at the same time.
[0064]
However, if present, depending on the method of manufacture, analysis of those elements should indicate the presence of said elements in an amount of less than 0.01 mol%, more preferably less than 0.001 mol%. p. p. m. In terms of quantity, these impurities are 100 p. p. m. Less than 10 p. p. m. Should be present in less than an amount.
[0065]
According to the present invention, acicular CsBr: Eu²⁺The storage phosphor plate or panel having the storage phosphor particles may include any one of the alkaline earth metals Li, Na, K, and Rb; any one of the alkaline earth metals Sr or the lanthanides Sm, Tm, and Pr. If included, its 100p. p. m. The following amount, more preferably 10 p. p. m. The following amounts are preferred.
[0066]
It is highly desirable that there be no oil and debris apart from the impurities mentioned above. The presence of oil may be due to loss of oil from the deposition unit, thereby causing oil penetration between the needle crystals of the needle phosphor. These phenomena clearly show a change in reflection and scattering properties when the material produces an undesirable glare on its surface and as it becomes more transparent. While reduced scattering results in increased sharpness, the presence of non-renewable amounts of oil cannot be recognized as part of an uncontrollable manufacturing process. Otherwise add a small amount of oil (eg spraying) in a controlled manufacturing process and clean the surface from excess oil (avoid the presence of oil on the surface, thus penetrating the oil between the acicular phosphor particles) Is given). However, if present on a storage plate or panel according to the present invention, the oil or debris is about 10 p. p. m. Is preferably present in an amount not exceeding.
[0067]
Since the temperature of the support has an effect on the diameter of the needles during formation, it is clear that it also has an effect on the packing density. The preferred temperature of the support during deposition of the acicular phosphor particles is in the range of about 250 ° C.
[0068]
Preferably the binderless phosphor screen is made by vapor deposition under an inert gas atmosphere such as argon. By adjusting the temperature of the support and the pressure of the inert gas during the deposition, the crystal size and packing density of the acicular crystals can be adjusted to desired levels.
[0069]
Ar is preferred as an inert gas suitable for use during deposition. Advantageously, the temperature of the air stream entering the vapor deposition device is maintained at a temperature between 0C and 100C. Preferably, the air stream is maintained at room temperature, i. The cooling airflow entering the vapor deposition device can cool both the vapor and the support prior to vapor deposition. The support is preferably maintained at a temperature T of 50 ° C ≦ T ≦ 300 ° C., more preferably 90 ° C. ≦ T ≦ 200 ° C. The Ar pressure is preferably at most 10 Pa, more preferably 1 Pa to 3 Pa.
[0070]
It has been found that when the phosphor screen is made under the conditions described above, the macroscopic dimensions of the needles can be affected: when the above method is used, thin needles having a diameter of 2 μm to 10 μm are obtained. Obtained. This is because the deposition rate of the phosphor or phosphor precursor is 1 mg / cm.²-Especially when it was above min. The packing density of the needle crystals is 70% to 90%.
[0071]
For display screen coating, the phosphor should have a high packing density so that optimal brightness of the display screen can be obtained. Furthermore, the phosphor in the display screen coating should be homogeneously distributed. That is, the photosensitive phosphor suspension should not form "bubbles" and "pinholes" in the drying process. Individual pixels in the display screen coating should exhibit high resolution to accurately reproduce the shadow mask structure. Phosphor colors should retain their color purity during the coating process. In addition, the display screen coating should adhere well and the unexposed areas of the display screen should be free of fogging.
[0072]
In the present application, which is directed to further improvements in the construction of flat storage phosphor screens or panels, the manufacture and operation of those phosphor screens depends on the application of the coating or layer arrangement used to increase the light output of the screen. In particular, it is optimized by improving the dimensions of the deposited storage phosphor crystals having a cylindrical shape. Selective filtering and / or attenuation of the X-ray energy applied to the screen, which allows the X-ray energy to be detected in a single exposure, reflects light when desired, and is also transparent to X-rays. It may be further provided by a placed coating. Improvements on storage phosphor screens within this range as described above are described in EP-A Nos. 10655525, 1103846 and 111458; WO 01/003156 and U.S. Pat. No. 1,159,004; . 01000694, 01000695, 01000696, 01000697 (filed on December 3, 2001) and EP Application No. No. 01000711 (filed on Dec. 5, 2001).
[0073]
According to the present invention, there is further used a radiation image storage phosphor panel as described above, wherein in a method of imaging for storing and reproducing a radiation image, the method has passed through or emitted by an object. Exposing said radiation image storage panel with radiation energy and storing said radiation energy in the form of a latent image on said image storage panel; energy stored in the form of light by irradiation with stimulating rays in the visible or infrared region And thereby emit light from the ultraviolet or visible wavelength range; collecting the light emitted from the storage panel by light collection means; converting the collected light into a series of electrical signals; Generating an image corresponding to the latent image from the image.
[0074]
While the invention will be described in the following in connection with preferred embodiments thereof, it will be understood that it is not intended to limit the invention to those embodiments.
[0075]
Example
To incorporate the dye or pigment into the needle image phosphor plate, the pigment or dye was dissolved in an alcohol solution. The phosphor plate was dipped in an alcohol solution and air-dried: as the alcohol evaporated, the dye was immobilized in the phosphor plate. In the alcohol dye solution MOWILITH @ CT5 (copolymer of vinyl acetate and crotonic acid, Hoechst @ AG, Frankfurt, trade name of Germany) was present as a binder in an amount of 5% by weight.
[0076]
In another experiment, CsBr: Eu²⁺The acicular phosphor layer of the storage phosphor was deposited on glass. Sensitivity (speed) and homogeneity were measured.
[0077]
A drop of oil was spread over the surface of the phosphor layer with a soft texture to provide oil penetration between the deposited needles. An increase in surface glare and transparency was visually observed.
[0078]
Sensitivity (speed) and homogeneity were measured again: speed was reduced by a factor of 2-3, but homogeneity was adversely affected, as in the resulting image, and strongly different "cloudy" sites were observed. This indicates a local difference in the tissue as the effect of oil contamination is different between sites (sites having a maximum diameter in the range of 2-5 mm). The speed cannot be restored later to reach its highest original level, even after performing a new annealing step.
[0079]
Sensitivity (speed) and homogeneity were measured again after 14 days: identical results were obtained according to those measured immediately after addition of the oil droplets.
[0080]
From these measurements, the effects of oilIs permanentThat can be established.
[0081]
Having described preferred embodiments of the invention in detail, it will be apparent to those skilled in the art that numerous changes may be made therein without departing from the scope of the invention as defined in the appended claims. .
[Brief description of the drawings]
FIG. 1 shows a photograph of a typical acicular storage phosphor panel having phosphor particles in the form of cylinders, with dimensions indicated.

Claims (6)

Acicular CsBr: Eu ²⁺ storage phosphor particles in the form of a cylinder, the average cross-sectional diameter ranging from 1 to 30 μm and the average length measured along the cylinder of the cylinder ranging from 100 μm to 1000 μm. Needle-like CsBr: Eu ²⁺ storage phosphor particles having thickness. 2. The acicular CsBr: Eu2 ⁺ storage phosphor particles of claim 1, wherein the phosphor particles have an average aspect ratio of length to cross-sectional diameter in the range of 5-200. The acicular CsBr: Eu2 ⁺ storage phosphor particles of claim 1 or 2, wherein the phosphor particles have an average aspect ratio of length to cross-sectional diameter in the range of 20-100. The acicular CsBr: Eu ²⁺ storage phosphor particles according to any one of claims 1 to 3, wherein a coefficient of variation of the average cross-sectional diameter and the average cylindrical portion length is in a range of 0.05 to 0.30. A radiation image storage phosphor screen or panel, which emits radiation after stimulation of phosphor particles storing energy from X-rays impinging on the screen or panel, said support having a flat surface, said support A radiation image comprising a plurality of cylindrical needle-like phosphors extending on the surface of a phosphor, wherein the phosphor is needle-like CsBr: Eu ²⁺ storage phosphor particles in the form of a cylinder according to any of claims 1 to 4. Storage phosphor screen or panel. 6. The radiation image storage of claim 5, wherein a light reflective coating is applied along the wall of the cylinder and is arranged to at least partially reflect radiation emitted by the phosphor after stimulation. Phosphor screen.

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