CA1147111A - Distributed phosphor scintillator structures - Google Patents
Distributed phosphor scintillator structuresInfo
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- CA1147111A CA1147111A CA000394702A CA394702A CA1147111A CA 1147111 A CA1147111 A CA 1147111A CA 000394702 A CA000394702 A CA 000394702A CA 394702 A CA394702 A CA 394702A CA 1147111 A CA1147111 A CA 1147111A
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Abstract
ABSTRACT OF THE DISCLOSURE
Distributed phosphor scintillator structures providing superior optical coupling to photoelectrically responsive devices together with methods for fabricat-ing said scintillator structures are disclosed. In accordance with one embodiment of the invention relat-ing to scintillator structures, the phosphor is distrib-uted in a layered fashion with certain layers being optically transparent so that the visible wavelength output of the scintillator is better directed to detect-ing devices. In accordance with another embodiment of the invention relating to scintillator structures, the phosphor is distributed throughout a transparent matrix in a continuous fashion whereby emitted light is more readily transmitted to a photodetector. Methods for fabricating said distributed phosphor scintillator structures are also disclosed.
Distributed phosphor scintillator structures providing superior optical coupling to photoelectrically responsive devices together with methods for fabricat-ing said scintillator structures are disclosed. In accordance with one embodiment of the invention relat-ing to scintillator structures, the phosphor is distrib-uted in a layered fashion with certain layers being optically transparent so that the visible wavelength output of the scintillator is better directed to detect-ing devices. In accordance with another embodiment of the invention relating to scintillator structures, the phosphor is distributed throughout a transparent matrix in a continuous fashion whereby emitted light is more readily transmitted to a photodetector. Methods for fabricating said distributed phosphor scintillator structures are also disclosed.
Description
DISTRIBUTED PHOSPHOR SCINTILLATOR STRUCTURES
sackground of the Invention . .
This application is a division of Canadian SN 313,333 filed October 13, L978 o This in~ention relates to scintillator structures and methods for fabricating such structures.
More particularly, this invention relates to a ~ethod for distributing the scintillator phosphor in such a way as to enhance the escape of the visible wavelen~th radiation that would otherwise be dissipated within a scintillator body. Two embodiments of the present invention are disclosed; one in which the phosphor is distributed in a layered structure and another in which the phosphor is dispersed throughout a trans-paren~ matrix.
In general, a scintillator is a material which emits electromagnetic radiation in the visible or near-visible spectrum when excited by high energy electro-magnetic photons such as those in the x~ray or gamma-ray regions of the spectrum, hereinafter referred to as supra-optical frequencies. Thus, these materials are excellent choices for use as detectors in industrial or medical x-ray or gamma-ray equipment. In most typi-cal applications, the output from scintillator materials is made to impinge upon photoelectrically responsive materials in order to produce an electrical output signal which is directly related to the intensity of the initial x-ray or gamma-ray bombardment.
Scintillator materials comprise a major portion of those devices used to detect the presence and intensity of incident high energy photons. Another commonly used de~ector is ~he high pressure noble gas ~ .
~ 7~ RD~12803 ioniza~ion device. This other form of high energy photon detector typically contains a gas, such as xenon, at a high pressure (density), which ionizes to a certain extent when subjected to high energy x-ray or gamm-ray radiation. This ionization causes a certain amount of current flow between the cathode and the anode of these detectors which are kept at relatively high and opposite polarities from one another. The current that flows is sensed by a current sensing circuit whose output is reflective of the intensity of the high energy radiation. Since this other form of detector operates on an ionization principle, after the termination of the irxadiating energy ~here still persists the possibility that a given ioni~ation path remains open. Ilence, these detectors are peculiarly sensitive to their own form of "after-glow" which results in -the blurring, in the time dimension, of information contained in the irradiating signal as a result of its passing through a body to be examined, as in computerized tomography applications.
As used herein and in the appended claims, the term "light" means ~hose electromagnetic radiations in the visible region of the spectrum and also near-visible wavelengths given off by certain fluorescent materials. Also, as used herein, and in the appended claims, the term "optical" encom~asses the same spec-tral region as the term "light" does.
In general, it is desirable that the amount of light output from these scintillators be as large as possible for a given amount of x-ray or gamma-ray energy. This is partlcularly true in the medical ~ 7~ RD-12803 tomography area where it is desired that the energy intensity of the x-ray be as smaLl as possible to minimize any danger to the patient.
Another important property that scintillator materials should possess is that of short afterglow or persistence. This means that there should be a relatively short period of time between the termination of the high energy radiating excitation and the cessa-tion of light output from the scintillator. If this is not the case, there is resultant blurring, in time, of the information-bearing signal generated~ for example, when the scintillator is used to produce tomographic imaging data. Furthermore, if rapid tomographic scan-ning is desired, the presence of the afterglow tends to severely limit ~he scan rate, thereby rendering difficult the viewing of moving bodily organs, such as the heart or lungs.
A scintillator boay or substance, in order to be effective, must be a good converter of high energy radiation (that is, x-rays and gamma-rays). Typically, present scintillator bodies consist of a phosphor in a powder, polycrystalline, or crystalline form. In these forms, the useful light that is produced upon high energy excitation is limited to that which can escape the interior of the scintillator body and that genera-ted in the surface regions. The escape of light is difficult due to the optical absorption resulting from multiple internal reflections~ each such reflection further attenuating the amount of light available to exterrlal detectors. Thus, it is necessary that not only the phosphors themselves have a good luminescent efficiency but it is also n~cessary that the light output be available for detection.
In the medical tomography area, where the intensity of x-radiation is modulated by the body through which lt passes, and which modulated radiation is then converted into electrical signals, it is important to have x~ray detection devices which have a good overall energy conversion efficiency. For devices with low efficiency, a higher x-ray flux radia-tion must be applied to produce the same light and electrical output from the overall system. In a medical tomographic context, this means that such a system has a low signal-to~noise ratio.
Typical scintillator phosphors which are used include barium fluorochloride doped with a europium activator (BaFCl:Eu). Other phosphors, for example, include bismuth germanate (Bi4Ge3O12), lanthanum oxy-bromide doped with terbium (LaOBr:Tb), cesium iodide doped with thalium (CsI:Tl~, cesium iodide doped with sodium (CsI:Na), calcium tungstate (Ca~O4), cadmium tungstate (CdWO4), zinc cadmium sulfide doped with silver (ZnCdS:Ag,Ni), gadolinium oxysulfide doped with terbium (Gd2O2S:Tb), and lanthanum oxybromide doped with dysprosium (LaOBr:Dy). Other host-crystal poss-ibilities for phosphors include the selenides o~ ~inc and cadmium, the tellurides of ~inc and cadmium, sodium iodide (NaI)~ and the oxysulfide o lanthanum (La2O2S).
SU~:r ~L o tlLe In~ Al~ iO1~
In accordance with one embodiment of the present invention, a suitable ph~sphQr is attached to both sides of an underlying substrate. The phosphor is attached either in the form of a powder~ in a more ~ RD-12~03 continuous form, or in a dispersed form in a trans-paren~ matrlx. The resulting substrates with the phD~sphor material attached are further layered with one another and with a transparent material between the phosphor-coated substrates. This structure ex-poses a greater area of phosphor material to a region from which the light output from the materiaL can easily escape to be detected. The resulting scintil-lator bodies are useful in tomographic detector arrays which require a high overa~ll energy conversion effici-ency for high resolution imagery and for patient safety.
In another embodiment of the present inven-tion, the phosphor material is distributed in a con-tinuous, rather than layered, fashion throughout a transparent matrix. Thus, the light that is produced upon x-ray or gamma-ray absorption deep within the scintillator body escapes from the body with a minimum amount of internal reflection and consequent light energy loss.
The phosphor material to be used in the above structure is inexpensive and the methods and other materials involved in fabricating the distributed phosphor scintillator structures are likewise inex-pensive.
In particular, scintillator detector bodies are qui~e inexpensive when compared to ionization type x-ray detectors which typically consist of a highly pressurized (25 atmospheres) noble gas surrounding electrodes maintained at high opposite polarities from one another.
Accordingly, it is an object of this invention to provide inexpensive scintillator structures and RD~ 03 methods of fabricating sa~le in which there is a maximum detectable light output.
Description of the Drawings FIG. l is a side elevation section view of one embodiment of the present invention showing the phosphor distributed in a layered fashion.
YIG. 2 is a top sectional view of the scintil- -lator body of FIG. 1 showing the phosphor distributed in a layered fashion.
FIG~ 3 is a side elevation view of a scintil-lator body surrounded by a ~acket which serves to convert the wavelength of the scintillator output to a more convenient wavelength.
FIG. 4 is a perspective view showing scintil-lator bodies with phosphor distributed in a layered fashion arranged as a portion of a tomographic x-ray detector.
FIG~ 5 iS a side elevation sectional view of a layered scintillator structure tilted for greater absorption.
FIG. 6 is a side elevation sectional ~iew of a scintillator body illustrating the continuous dis-persion of phosphor throughout the body.
FIG~ 7 iS a top sectional view of the scintil-lator body of FIGo 6 also showing a continuous disper-sion of phosphor material throughout the body.
FIG. 8 shows a portion of a tomographic x-ray detector employing scin~illator bodies with phosphor continuously distributed throughout.
~ 7~ RD-12~03 Detailed Description of the Invention The invention herein relates to a distributed phosphor scintillator structure which produces a greater optical coupling between the scintillator body and the photoelectrically responsive device, such as photodiode, to which it is typically coupled. There are two primary embodiments of the present invention: ~
one in which the phosphor is` distributed in a layered ~`
fashion, which will be discussed first, and the other embodiment in which the phosphor is distributed in a continuous fashion throughout the scintillator body.
It is to be noted, however, that sections of scintillator bodies with the phosphor distributed continuously throughout are in fact used in one embodiment of a multi-layered structure.
FIG. 1 shows a side elevation sectional view of a scintillator body 1 with a multi-layered struc-ture. In this embodiment, phosphor material 3 is attached to a substrate 5. The phosphored substrates are further layered with light channeLing laminates 4 between layers of phosphored substrates. Fig. l Eurther shows an x-ray photon 2 being absorbed by a phosphor particle at an absorption site 6 in the fourth layer of the phosphor encountered. The absorption of the high energy, high frequency x-ray photon causes the production of a multitude of lower energy, lower frequ-ency optical wavelength photons. The path 7 of a typical optical wavelength photon is shown as it is refIected back and forth within the light channeling laminate 4 between the phosphor layers, ultlmately escaping the scintilLator body whereby it may be more readily detec- ;
ted than if the absorption event had occurred deep ~ ~7~ . RD-1 2803 withinside a dense and much less optically transparent body.
FIG 2 shows a plan view of the same scintil-lator structure as shown in FIG. 1. It is to be noted that both FIG. 1 and FIG. 2 suggest that the number of phosphored substrate layers need not be ~ixed at four as shown.
FIG. 1 and 2 both illustrate a phosphor layer 3. In particular, this layer 3 applied to the substrate 5 is used in a variety of phosphor forms. The phosphor layer 3 consists of a phosphor either as a powder, as a single crystal, as a phosphor dispersed in a trans-parent matrix, or as a phosphor applied in a continuous layer, as for example, a phosphor material formed by a quench from its liquid form.
Any of the phosphors mentioned above are applicable to the substrate material with an appropriate adhesive and a suitable thickness. In par-ticuLar, ZnCdS:Ag has a particularly large particle size which enhances the escape of light from the powder layer.
The emission coLor of t-his particular phosphor is orange-red, which ma~es it well suited to detection by silicon semiconductor devices. A layer of 0.5 milli-meters of this phosphor absorbs between 20 and 25 percent of the ~-ray photons in a computerized tomography system in which the average x-ray energy is approximately 65 kev. Five or six layers of this thickness, ~or this phosphor, suffices to absorb 90 percent or more of the x-ray photons. When other more x-ray absorptive rare earth phosphors such as LaOBr:Tm, LaOBr:Tb, Gd2O2S:Tb, or La2O2S:Tb are employed, the number of layers to achieve the same total absorption is lower and losses ~ 7~ RD-12803 in the light guiding laminate layers are also lower.
There are nu~erous choices for the phosphor material. The three criteria to be used in selecting a particular phosphor for computerized tomographic applications are its high luminescent efficiency, short afterglow duration, and low absorbance for the emitted light. The decay rate ~afterglow) is partic-ularly important in those medical tomographic applica-tions where repeated scans are made, as for example, when dealing with images of moving bodily organs.
Choices for host-crystal lattices to be doped with an appropriate rare earth element ~that i9, those elements with an atomic number between 58 and 71 inclusive) or other activator include the sulfides, selenides, and tellurides of zinc and cadmium, the iodides of sodium and cesiumj the tungstates of calcium and cadmium, -~
lanthanum oxybromide, and the oxysulfides~of ;lanthanum and gàdolinium.
In addition to applying the phosphor material to the substrate in a powder form, the phosphor materiaL
is also applicable as a single crystal when such crystals exist. For example, a single crystal o~ cesium iodide doped with thallium is attachable to the substrate with a suitable adhesive, such as an epoxy.
In another embodiment of the present inven- -tion, the phosphor layer which is attached to the sub-strate consists of an appropriate phosphor continuously dispersed in a transparent matrix such as that which is more particularly described below as an embodiment of the entire scintillator body. This particular layer-ed embodiment combines features of both of the principal embodiments disclosed herein.
~ RD-12803 In another embodiment of the present invention, the phosphor layer 3 (FIG. 1) consists of a phosphor layer which is either evaporated onto the substrate, is melt solidified, hlgh pressure sintexed, or hot-forged. Such methods for producing these continuous phosphor layers are described in the Canadian applica-tion of Dominic A. Cusano, et al., Serial No. 313,418 dated October 13, 1978 which is assigned to the same assignee as the present invention.
The substrate material itself must be of a material which is not absorptive of x-rays whose frequencies are within the spectral region o~ interest.
Typically, the substrate is composed of a clear fused quartz material. It is desirable that this substrate also be optically transparent although this is not essential since the majority of the optical photons emitted find their way to the outside through the light channeling laminate layers. A typical thickness for this substrate in a tomographic application is 20 mils.
Between each phosphored substrate, there is a layer of light channeling laminate material. Typically, this laminate is an epoxy. The principal criteria for selection of this laminate materlal is its optical trans-parency. Other criteria for its selection include its chemical compatibility with the phosphor material, its siructural rigidity, its low x~ray absorptivity, and its ability to withstand prolonged x-ray bombardment.
The laminate layer also includes as a matter of choice, at least one wavelength conversion material such as a fluorescent dye which emits a visible wave-length photon at a wavelength susceptible to detection by the photoelectrically responsive sensing device ~ RD~12~03 when the wavelength conversion material itself is stlmulated by the opticaL wavelength photons produced by the phosphor material. The intimate contact which the laminate layers have with the phosphorea layers render them particularly useful and efficient for wave-length conversion.
If it is not desired that the wavelength conversion process take place within the scintillator - -body, then in another embodiment of the present inven-tion, the scintillator body is surrounded by a jacket containing a wavelength conversion substance. This embodiment is illustrated in FIG. 3 wherein multilayer-ed scintillator body lO, described below) is surrounded by a jacket 8 containing suitable wavelength conversion material such as certain fluorescent dyes. In FIG. 3, x-ray photon 2 is absorbed at absorption site 6 within the scintillator body l as a result of which multiple photons at a first lower wavelength are emitted, said photon paths being illustrated by a typical photon path 7. Said photon traverses a typical path 7 to a secondary absorptive site 6' within the jacket 8 sur-rounding the scintillator body l. At this site 6', the first wavelength photon is absorbed and another photon at a second wavelength is emitted traversing typical path 7l In this fashion, the scintillator wavelength can be matched to a more sensitive spectral region of the photoelectrically responsive detector.
If desired, several wavelength conversions may be made through the use of multiply matched fluorescent materials.
By way of example, the fabrica-tion of a typical multi-layered scintillator body is described. BaFCl:Eu phosphor is mixed with an equal weight of 1269A
STYCAS ~ epoxy (supplied by Emerson and Cuminy, Inc., of Canton, Mass.) containing rhodamine to the extent of 0.1 gm rhodamine per 30 ml of epoxy. The BaFCl:Eu phosphor is typically doped with Eu to the extent of approximately 1 mole percent but doping to as low as O.l mole percent and as high as 5 mole percent is also used. The phosphor is suspended in epoxy by tumbling for 16 hours in a glass container containing y1ass beads. The suspension is doctor bladea onto 20 mil CFQ (clear fused quartz) substrates to a thickness of 20 mils using a Gardner doctor blade. The film is cured for 18 hours at 88C and the substrate is then coated on the other side and cured. A block scintil-lator material is made by laminating the phosphored substrate with 40 mil cured epoxy spacers using the same epoxy as a cement. In another embodiment of this invention~ as described above, the organic dye rhodamine is incorporated in the epoxy spacers rather than in the phosphor itself.
There are a number of desirable features o~
this particular scintillator structure. For instance, the amount of x-ray radiation absorption is controlled by the number of layers of the phosphor. The number of layers is adjusted for the various absorptivity of the phosphors employed. If desired, certain substrates are coated with phosphor material on one side only. Another advantage of this structure is that it permits the use of phosphors other than the single crystal phosphors such as CsI.Tl. This structure permits much more flex-ible choices in the selection of an appropriate phosphor when such phosphor properties as light output wavelength, ` RD-12803 ~ ~7~
short afterglow duration, and luminescent efficiency must be balanced. In addition, there are certain phosphors such as CsI:Na, which while otherwise being good phosphor materials, suffer from the problem that they are hygroscopic and when exposed in an atmospheric environment absorb water and suffer a resultan~ deg-ràdation in performance. This problem is much less severe in a layered structure in which the phosphor material has a minimal exposure to the atmosphere than it is when the phosphor is applied to an exposed imaging screen. Another feature of the scintillator body stemming from the structure of the present inven-tion is its rigidity and ruggedness. A scintillator tomographic detector structure, such as that described in FIG. ~ does not suffer from the acoustic or micro-phonic noise pickup effects as do high pressure noble gas ionization type detectorsO Moreover, the scintil-lator bodies are manuEactured to a high degree of accuracy which enabLes them to be preciseLy aligned in a detector array such as that of FIG. 4. Moreover, the phosphor substrates may be angled as shown in FIG.
5 to produce a greater absorption without increasing the thickness of the phosphor on the substrate, thereby permitting, if desired, fewer layers to produce the same amount of absorption.
In accordance with the other principal preferred embodiment of the present invention, the phosphor material is dispersed in a continuous and uniform fashion throughout a transparent supportive matrix. FIG. 6 shows such a scintillator body 10 being stimulated by a high energy x-ray photon 2.
In this configuration, phosphor particules Ll are ~ .
~a~
suspended in a rigid transparent matrix 12. The high energy x-ra~ photon 2 is converted to light by absorp~
tion at absorption site 6 from which a multitude of lower energy optical waveLength photons are emitted.
These photons are readily transmitted through the trans-parent matrix 12 being periodically reflected and scat tered by the phosphor particles 11 which are also pre-sent. However, most of the resultant light energy produced eventually reaches the exterior of the scin-tillator body where it is detected. Such a typical light path 7 is shown in FIG. 6.
FIG. 7 is a sectlonal plan view of FIG 6 showing the absorption event produced by x-ray photon
sackground of the Invention . .
This application is a division of Canadian SN 313,333 filed October 13, L978 o This in~ention relates to scintillator structures and methods for fabricating such structures.
More particularly, this invention relates to a ~ethod for distributing the scintillator phosphor in such a way as to enhance the escape of the visible wavelen~th radiation that would otherwise be dissipated within a scintillator body. Two embodiments of the present invention are disclosed; one in which the phosphor is distributed in a layered structure and another in which the phosphor is dispersed throughout a trans-paren~ matrix.
In general, a scintillator is a material which emits electromagnetic radiation in the visible or near-visible spectrum when excited by high energy electro-magnetic photons such as those in the x~ray or gamma-ray regions of the spectrum, hereinafter referred to as supra-optical frequencies. Thus, these materials are excellent choices for use as detectors in industrial or medical x-ray or gamma-ray equipment. In most typi-cal applications, the output from scintillator materials is made to impinge upon photoelectrically responsive materials in order to produce an electrical output signal which is directly related to the intensity of the initial x-ray or gamma-ray bombardment.
Scintillator materials comprise a major portion of those devices used to detect the presence and intensity of incident high energy photons. Another commonly used de~ector is ~he high pressure noble gas ~ .
~ 7~ RD~12803 ioniza~ion device. This other form of high energy photon detector typically contains a gas, such as xenon, at a high pressure (density), which ionizes to a certain extent when subjected to high energy x-ray or gamm-ray radiation. This ionization causes a certain amount of current flow between the cathode and the anode of these detectors which are kept at relatively high and opposite polarities from one another. The current that flows is sensed by a current sensing circuit whose output is reflective of the intensity of the high energy radiation. Since this other form of detector operates on an ionization principle, after the termination of the irxadiating energy ~here still persists the possibility that a given ioni~ation path remains open. Ilence, these detectors are peculiarly sensitive to their own form of "after-glow" which results in -the blurring, in the time dimension, of information contained in the irradiating signal as a result of its passing through a body to be examined, as in computerized tomography applications.
As used herein and in the appended claims, the term "light" means ~hose electromagnetic radiations in the visible region of the spectrum and also near-visible wavelengths given off by certain fluorescent materials. Also, as used herein, and in the appended claims, the term "optical" encom~asses the same spec-tral region as the term "light" does.
In general, it is desirable that the amount of light output from these scintillators be as large as possible for a given amount of x-ray or gamma-ray energy. This is partlcularly true in the medical ~ 7~ RD-12803 tomography area where it is desired that the energy intensity of the x-ray be as smaLl as possible to minimize any danger to the patient.
Another important property that scintillator materials should possess is that of short afterglow or persistence. This means that there should be a relatively short period of time between the termination of the high energy radiating excitation and the cessa-tion of light output from the scintillator. If this is not the case, there is resultant blurring, in time, of the information-bearing signal generated~ for example, when the scintillator is used to produce tomographic imaging data. Furthermore, if rapid tomographic scan-ning is desired, the presence of the afterglow tends to severely limit ~he scan rate, thereby rendering difficult the viewing of moving bodily organs, such as the heart or lungs.
A scintillator boay or substance, in order to be effective, must be a good converter of high energy radiation (that is, x-rays and gamma-rays). Typically, present scintillator bodies consist of a phosphor in a powder, polycrystalline, or crystalline form. In these forms, the useful light that is produced upon high energy excitation is limited to that which can escape the interior of the scintillator body and that genera-ted in the surface regions. The escape of light is difficult due to the optical absorption resulting from multiple internal reflections~ each such reflection further attenuating the amount of light available to exterrlal detectors. Thus, it is necessary that not only the phosphors themselves have a good luminescent efficiency but it is also n~cessary that the light output be available for detection.
In the medical tomography area, where the intensity of x-radiation is modulated by the body through which lt passes, and which modulated radiation is then converted into electrical signals, it is important to have x~ray detection devices which have a good overall energy conversion efficiency. For devices with low efficiency, a higher x-ray flux radia-tion must be applied to produce the same light and electrical output from the overall system. In a medical tomographic context, this means that such a system has a low signal-to~noise ratio.
Typical scintillator phosphors which are used include barium fluorochloride doped with a europium activator (BaFCl:Eu). Other phosphors, for example, include bismuth germanate (Bi4Ge3O12), lanthanum oxy-bromide doped with terbium (LaOBr:Tb), cesium iodide doped with thalium (CsI:Tl~, cesium iodide doped with sodium (CsI:Na), calcium tungstate (Ca~O4), cadmium tungstate (CdWO4), zinc cadmium sulfide doped with silver (ZnCdS:Ag,Ni), gadolinium oxysulfide doped with terbium (Gd2O2S:Tb), and lanthanum oxybromide doped with dysprosium (LaOBr:Dy). Other host-crystal poss-ibilities for phosphors include the selenides o~ ~inc and cadmium, the tellurides of ~inc and cadmium, sodium iodide (NaI)~ and the oxysulfide o lanthanum (La2O2S).
SU~:r ~L o tlLe In~ Al~ iO1~
In accordance with one embodiment of the present invention, a suitable ph~sphQr is attached to both sides of an underlying substrate. The phosphor is attached either in the form of a powder~ in a more ~ RD-12~03 continuous form, or in a dispersed form in a trans-paren~ matrlx. The resulting substrates with the phD~sphor material attached are further layered with one another and with a transparent material between the phosphor-coated substrates. This structure ex-poses a greater area of phosphor material to a region from which the light output from the materiaL can easily escape to be detected. The resulting scintil-lator bodies are useful in tomographic detector arrays which require a high overa~ll energy conversion effici-ency for high resolution imagery and for patient safety.
In another embodiment of the present inven-tion, the phosphor material is distributed in a con-tinuous, rather than layered, fashion throughout a transparent matrix. Thus, the light that is produced upon x-ray or gamma-ray absorption deep within the scintillator body escapes from the body with a minimum amount of internal reflection and consequent light energy loss.
The phosphor material to be used in the above structure is inexpensive and the methods and other materials involved in fabricating the distributed phosphor scintillator structures are likewise inex-pensive.
In particular, scintillator detector bodies are qui~e inexpensive when compared to ionization type x-ray detectors which typically consist of a highly pressurized (25 atmospheres) noble gas surrounding electrodes maintained at high opposite polarities from one another.
Accordingly, it is an object of this invention to provide inexpensive scintillator structures and RD~ 03 methods of fabricating sa~le in which there is a maximum detectable light output.
Description of the Drawings FIG. l is a side elevation section view of one embodiment of the present invention showing the phosphor distributed in a layered fashion.
YIG. 2 is a top sectional view of the scintil- -lator body of FIG. 1 showing the phosphor distributed in a layered fashion.
FIG~ 3 is a side elevation view of a scintil-lator body surrounded by a ~acket which serves to convert the wavelength of the scintillator output to a more convenient wavelength.
FIG. 4 is a perspective view showing scintil-lator bodies with phosphor distributed in a layered fashion arranged as a portion of a tomographic x-ray detector.
FIG~ 5 iS a side elevation sectional view of a layered scintillator structure tilted for greater absorption.
FIG. 6 is a side elevation sectional ~iew of a scintillator body illustrating the continuous dis-persion of phosphor throughout the body.
FIG~ 7 iS a top sectional view of the scintil-lator body of FIGo 6 also showing a continuous disper-sion of phosphor material throughout the body.
FIG. 8 shows a portion of a tomographic x-ray detector employing scin~illator bodies with phosphor continuously distributed throughout.
~ 7~ RD-12~03 Detailed Description of the Invention The invention herein relates to a distributed phosphor scintillator structure which produces a greater optical coupling between the scintillator body and the photoelectrically responsive device, such as photodiode, to which it is typically coupled. There are two primary embodiments of the present invention: ~
one in which the phosphor is` distributed in a layered ~`
fashion, which will be discussed first, and the other embodiment in which the phosphor is distributed in a continuous fashion throughout the scintillator body.
It is to be noted, however, that sections of scintillator bodies with the phosphor distributed continuously throughout are in fact used in one embodiment of a multi-layered structure.
FIG. 1 shows a side elevation sectional view of a scintillator body 1 with a multi-layered struc-ture. In this embodiment, phosphor material 3 is attached to a substrate 5. The phosphored substrates are further layered with light channeLing laminates 4 between layers of phosphored substrates. Fig. l Eurther shows an x-ray photon 2 being absorbed by a phosphor particle at an absorption site 6 in the fourth layer of the phosphor encountered. The absorption of the high energy, high frequency x-ray photon causes the production of a multitude of lower energy, lower frequ-ency optical wavelength photons. The path 7 of a typical optical wavelength photon is shown as it is refIected back and forth within the light channeling laminate 4 between the phosphor layers, ultlmately escaping the scintilLator body whereby it may be more readily detec- ;
ted than if the absorption event had occurred deep ~ ~7~ . RD-1 2803 withinside a dense and much less optically transparent body.
FIG 2 shows a plan view of the same scintil-lator structure as shown in FIG. 1. It is to be noted that both FIG. 1 and FIG. 2 suggest that the number of phosphored substrate layers need not be ~ixed at four as shown.
FIG. 1 and 2 both illustrate a phosphor layer 3. In particular, this layer 3 applied to the substrate 5 is used in a variety of phosphor forms. The phosphor layer 3 consists of a phosphor either as a powder, as a single crystal, as a phosphor dispersed in a trans-parent matrix, or as a phosphor applied in a continuous layer, as for example, a phosphor material formed by a quench from its liquid form.
Any of the phosphors mentioned above are applicable to the substrate material with an appropriate adhesive and a suitable thickness. In par-ticuLar, ZnCdS:Ag has a particularly large particle size which enhances the escape of light from the powder layer.
The emission coLor of t-his particular phosphor is orange-red, which ma~es it well suited to detection by silicon semiconductor devices. A layer of 0.5 milli-meters of this phosphor absorbs between 20 and 25 percent of the ~-ray photons in a computerized tomography system in which the average x-ray energy is approximately 65 kev. Five or six layers of this thickness, ~or this phosphor, suffices to absorb 90 percent or more of the x-ray photons. When other more x-ray absorptive rare earth phosphors such as LaOBr:Tm, LaOBr:Tb, Gd2O2S:Tb, or La2O2S:Tb are employed, the number of layers to achieve the same total absorption is lower and losses ~ 7~ RD-12803 in the light guiding laminate layers are also lower.
There are nu~erous choices for the phosphor material. The three criteria to be used in selecting a particular phosphor for computerized tomographic applications are its high luminescent efficiency, short afterglow duration, and low absorbance for the emitted light. The decay rate ~afterglow) is partic-ularly important in those medical tomographic applica-tions where repeated scans are made, as for example, when dealing with images of moving bodily organs.
Choices for host-crystal lattices to be doped with an appropriate rare earth element ~that i9, those elements with an atomic number between 58 and 71 inclusive) or other activator include the sulfides, selenides, and tellurides of zinc and cadmium, the iodides of sodium and cesiumj the tungstates of calcium and cadmium, -~
lanthanum oxybromide, and the oxysulfides~of ;lanthanum and gàdolinium.
In addition to applying the phosphor material to the substrate in a powder form, the phosphor materiaL
is also applicable as a single crystal when such crystals exist. For example, a single crystal o~ cesium iodide doped with thallium is attachable to the substrate with a suitable adhesive, such as an epoxy.
In another embodiment of the present inven- -tion, the phosphor layer which is attached to the sub-strate consists of an appropriate phosphor continuously dispersed in a transparent matrix such as that which is more particularly described below as an embodiment of the entire scintillator body. This particular layer-ed embodiment combines features of both of the principal embodiments disclosed herein.
~ RD-12803 In another embodiment of the present invention, the phosphor layer 3 (FIG. 1) consists of a phosphor layer which is either evaporated onto the substrate, is melt solidified, hlgh pressure sintexed, or hot-forged. Such methods for producing these continuous phosphor layers are described in the Canadian applica-tion of Dominic A. Cusano, et al., Serial No. 313,418 dated October 13, 1978 which is assigned to the same assignee as the present invention.
The substrate material itself must be of a material which is not absorptive of x-rays whose frequencies are within the spectral region o~ interest.
Typically, the substrate is composed of a clear fused quartz material. It is desirable that this substrate also be optically transparent although this is not essential since the majority of the optical photons emitted find their way to the outside through the light channeling laminate layers. A typical thickness for this substrate in a tomographic application is 20 mils.
Between each phosphored substrate, there is a layer of light channeling laminate material. Typically, this laminate is an epoxy. The principal criteria for selection of this laminate materlal is its optical trans-parency. Other criteria for its selection include its chemical compatibility with the phosphor material, its siructural rigidity, its low x~ray absorptivity, and its ability to withstand prolonged x-ray bombardment.
The laminate layer also includes as a matter of choice, at least one wavelength conversion material such as a fluorescent dye which emits a visible wave-length photon at a wavelength susceptible to detection by the photoelectrically responsive sensing device ~ RD~12~03 when the wavelength conversion material itself is stlmulated by the opticaL wavelength photons produced by the phosphor material. The intimate contact which the laminate layers have with the phosphorea layers render them particularly useful and efficient for wave-length conversion.
If it is not desired that the wavelength conversion process take place within the scintillator - -body, then in another embodiment of the present inven-tion, the scintillator body is surrounded by a jacket containing a wavelength conversion substance. This embodiment is illustrated in FIG. 3 wherein multilayer-ed scintillator body lO, described below) is surrounded by a jacket 8 containing suitable wavelength conversion material such as certain fluorescent dyes. In FIG. 3, x-ray photon 2 is absorbed at absorption site 6 within the scintillator body l as a result of which multiple photons at a first lower wavelength are emitted, said photon paths being illustrated by a typical photon path 7. Said photon traverses a typical path 7 to a secondary absorptive site 6' within the jacket 8 sur-rounding the scintillator body l. At this site 6', the first wavelength photon is absorbed and another photon at a second wavelength is emitted traversing typical path 7l In this fashion, the scintillator wavelength can be matched to a more sensitive spectral region of the photoelectrically responsive detector.
If desired, several wavelength conversions may be made through the use of multiply matched fluorescent materials.
By way of example, the fabrica-tion of a typical multi-layered scintillator body is described. BaFCl:Eu phosphor is mixed with an equal weight of 1269A
STYCAS ~ epoxy (supplied by Emerson and Cuminy, Inc., of Canton, Mass.) containing rhodamine to the extent of 0.1 gm rhodamine per 30 ml of epoxy. The BaFCl:Eu phosphor is typically doped with Eu to the extent of approximately 1 mole percent but doping to as low as O.l mole percent and as high as 5 mole percent is also used. The phosphor is suspended in epoxy by tumbling for 16 hours in a glass container containing y1ass beads. The suspension is doctor bladea onto 20 mil CFQ (clear fused quartz) substrates to a thickness of 20 mils using a Gardner doctor blade. The film is cured for 18 hours at 88C and the substrate is then coated on the other side and cured. A block scintil-lator material is made by laminating the phosphored substrate with 40 mil cured epoxy spacers using the same epoxy as a cement. In another embodiment of this invention~ as described above, the organic dye rhodamine is incorporated in the epoxy spacers rather than in the phosphor itself.
There are a number of desirable features o~
this particular scintillator structure. For instance, the amount of x-ray radiation absorption is controlled by the number of layers of the phosphor. The number of layers is adjusted for the various absorptivity of the phosphors employed. If desired, certain substrates are coated with phosphor material on one side only. Another advantage of this structure is that it permits the use of phosphors other than the single crystal phosphors such as CsI.Tl. This structure permits much more flex-ible choices in the selection of an appropriate phosphor when such phosphor properties as light output wavelength, ` RD-12803 ~ ~7~
short afterglow duration, and luminescent efficiency must be balanced. In addition, there are certain phosphors such as CsI:Na, which while otherwise being good phosphor materials, suffer from the problem that they are hygroscopic and when exposed in an atmospheric environment absorb water and suffer a resultan~ deg-ràdation in performance. This problem is much less severe in a layered structure in which the phosphor material has a minimal exposure to the atmosphere than it is when the phosphor is applied to an exposed imaging screen. Another feature of the scintillator body stemming from the structure of the present inven-tion is its rigidity and ruggedness. A scintillator tomographic detector structure, such as that described in FIG. ~ does not suffer from the acoustic or micro-phonic noise pickup effects as do high pressure noble gas ionization type detectorsO Moreover, the scintil-lator bodies are manuEactured to a high degree of accuracy which enabLes them to be preciseLy aligned in a detector array such as that of FIG. 4. Moreover, the phosphor substrates may be angled as shown in FIG.
5 to produce a greater absorption without increasing the thickness of the phosphor on the substrate, thereby permitting, if desired, fewer layers to produce the same amount of absorption.
In accordance with the other principal preferred embodiment of the present invention, the phosphor material is dispersed in a continuous and uniform fashion throughout a transparent supportive matrix. FIG. 6 shows such a scintillator body 10 being stimulated by a high energy x-ray photon 2.
In this configuration, phosphor particules Ll are ~ .
~a~
suspended in a rigid transparent matrix 12. The high energy x-ra~ photon 2 is converted to light by absorp~
tion at absorption site 6 from which a multitude of lower energy optical waveLength photons are emitted.
These photons are readily transmitted through the trans-parent matrix 12 being periodically reflected and scat tered by the phosphor particles 11 which are also pre-sent. However, most of the resultant light energy produced eventually reaches the exterior of the scin-tillator body where it is detected. Such a typical light path 7 is shown in FIG. 6.
FIG. 7 is a sectlonal plan view of FIG 6 showing the absorption event produced by x-ray photon
2 within scintillator body 10. The principa] require-ments for the transparent matrix 12 are: that it be a good tran$mitter of light at the wavelength produced by the phosphor; that it not react with the phosphor;
and that it maintain the phosphor in a fixed suspension after the phosphor is thoroughly dispersed throughout it. ~ number of plastics are appropriate for this purpose, such as the polyimide silicone copolymers.
In the embodiment described in FIG. 6, the particular choice of phosphor material is a matter of design consideration to be based on such factors as absorptivity, afterglow duration, lumlnescent e~ficiency, and output wavelength. However, the concentration of phosphor within the transparent (typically plastic) matrix is controlled to effect changes in the overall absorption. Typically, in this embodiment, the disper-sed phosphor scintillator body consists of the phosphor in a concentration of 10 to 20 percent by volume.
Wavelength conversion is accomplished in the ~D-12803 dispersed phosphor structure, if desired, in one of two ways. In accordance with one embodiment, the scintillator body is surrounded by a jacket containing a suitabLe wavelength converter such as the oryanic dye rhodamine. Such a structure is illustrated in FIG. 3 which is equally applicable to the dispersed phosphor structure as lt is for the multilayered structure discussed above. In accordance with another embodiment of the invention, the wavelength conversion substance is mixed with the transparent matrix material.
The fabrication of dispersed phosphor scin-tillator bodies is particularly easy. The phosphor chosen is thoroughly mixed with an epoxy, plastic, or other polymer whose optical and chemical properties are not seriously affected by x-radiation and the resul-tant mixture is then solidified or if not actually solidified, the phosphor is fixed in suspension. The solidification is accomplished by a variety oE processes including chemical activation, slight temperature eleva-tion (curing), or ultraviolet irradiation. The scintil-lator bodies are produced either individualLy or en masse in a prefabricated detector array structure.
FIG. ~ describes such a scintillator structure compris-ing a front wall member 21 composed of a low Z (atomic number) material, such as beryllium or aluminum, which is highly nonabsorptive for x-rays, collimator members 20 composed of a high Z material such as tungsten or tantalum which are relatively impervious to x-rays, a floor membex 23, and a rear wall member 22. Detector members 20, 21, 22 and 23 define a series of volumes into which the dispersed phosphor material is inserted.
To insure an adequate and thorough fill of this collimator structure, the entire detector assembly is agitated, typically at ultrasonic frequencies, during the inser-tion process. The material is then solidified either through chemical activation, during, or by ultraviolet irradiation. This process results in an extremely rugged detector array which is much less susceptible to acoustic noise vibration during operation than is a typical ionization detector array. Again, a wave-length conversion material is added, if desired.
As with the multilayered structure, the dis-persed phosphor structure is also rigid, rugged, and is easy to align accurately. The hygroscopic problem suffered by certain phosphor materials is also greatly reduced. In addition, the use of non-single crystal phosphor material is possible. Furthermore, the dis-persed phosphor structure is not very sensitive to changes in the x-ray spectrum occurring as a result of any filtering effects of the matrix material.
Thus, it can be appreciated from the above that the principal preferred embodiments of the present invention describe scintillator structures with decided advantages over other scintillator structures and which in particularly greatly enhance the escape of light from the scintillator bod~.
While this invention has been described with reference to particular embodiments and examples, other modifications and variations will occur to those skilled in the art in view of the above teachings. Accordingly, it should be understood that within the scope of the appended claims, the invention may be practices other-wise than is specifically described.
and that it maintain the phosphor in a fixed suspension after the phosphor is thoroughly dispersed throughout it. ~ number of plastics are appropriate for this purpose, such as the polyimide silicone copolymers.
In the embodiment described in FIG. 6, the particular choice of phosphor material is a matter of design consideration to be based on such factors as absorptivity, afterglow duration, lumlnescent e~ficiency, and output wavelength. However, the concentration of phosphor within the transparent (typically plastic) matrix is controlled to effect changes in the overall absorption. Typically, in this embodiment, the disper-sed phosphor scintillator body consists of the phosphor in a concentration of 10 to 20 percent by volume.
Wavelength conversion is accomplished in the ~D-12803 dispersed phosphor structure, if desired, in one of two ways. In accordance with one embodiment, the scintillator body is surrounded by a jacket containing a suitabLe wavelength converter such as the oryanic dye rhodamine. Such a structure is illustrated in FIG. 3 which is equally applicable to the dispersed phosphor structure as lt is for the multilayered structure discussed above. In accordance with another embodiment of the invention, the wavelength conversion substance is mixed with the transparent matrix material.
The fabrication of dispersed phosphor scin-tillator bodies is particularly easy. The phosphor chosen is thoroughly mixed with an epoxy, plastic, or other polymer whose optical and chemical properties are not seriously affected by x-radiation and the resul-tant mixture is then solidified or if not actually solidified, the phosphor is fixed in suspension. The solidification is accomplished by a variety oE processes including chemical activation, slight temperature eleva-tion (curing), or ultraviolet irradiation. The scintil-lator bodies are produced either individualLy or en masse in a prefabricated detector array structure.
FIG. ~ describes such a scintillator structure compris-ing a front wall member 21 composed of a low Z (atomic number) material, such as beryllium or aluminum, which is highly nonabsorptive for x-rays, collimator members 20 composed of a high Z material such as tungsten or tantalum which are relatively impervious to x-rays, a floor membex 23, and a rear wall member 22. Detector members 20, 21, 22 and 23 define a series of volumes into which the dispersed phosphor material is inserted.
To insure an adequate and thorough fill of this collimator structure, the entire detector assembly is agitated, typically at ultrasonic frequencies, during the inser-tion process. The material is then solidified either through chemical activation, during, or by ultraviolet irradiation. This process results in an extremely rugged detector array which is much less susceptible to acoustic noise vibration during operation than is a typical ionization detector array. Again, a wave-length conversion material is added, if desired.
As with the multilayered structure, the dis-persed phosphor structure is also rigid, rugged, and is easy to align accurately. The hygroscopic problem suffered by certain phosphor materials is also greatly reduced. In addition, the use of non-single crystal phosphor material is possible. Furthermore, the dis-persed phosphor structure is not very sensitive to changes in the x-ray spectrum occurring as a result of any filtering effects of the matrix material.
Thus, it can be appreciated from the above that the principal preferred embodiments of the present invention describe scintillator structures with decided advantages over other scintillator structures and which in particularly greatly enhance the escape of light from the scintillator bod~.
While this invention has been described with reference to particular embodiments and examples, other modifications and variations will occur to those skilled in the art in view of the above teachings. Accordingly, it should be understood that within the scope of the appended claims, the invention may be practices other-wise than is specifically described.
Claims (14)
- The embodiments of the invention in which an exclusive property or privilege is claimed are defined as follows:
l. A method of manufacturing dispersed phosphor scintillator bodies to increase detectable output at optical wavelengths when excited by high energy photons at supra-optical frequencies, said method comprising the steps of:
A) mixing a suitable phosphor with a non-reactive liquid matrix substance to form a suspension, said liquid matrix substance being hardenable, and transparent to optical and supra-optical radiation after hardening;
B) inserting the suspension from step A into a mold of desired shape; and then C) causing the liquid matrix substance to harden, whereby the phosphor particles are fixed in position and whereby optical wavelength output produced within the resulting scintillator body is permitted to escape to the exterior of the body for detection. - 2. The method of claim l in which the mold in step B is an array with a plurality of substantially identical separate compartments, said array comprising:
a front wall member, said front wall member being permeable to radiation at supra-optical frequencies;
side wall members, said side wall members being opaque to radiation at supra-optical frequencies;
interior wall members defining a plurality of said compartments, said interior wall members being orient-ed parallel to said side wall members and being opaque to radiation at supra-optical frequencies;
a floor member; and a rear wall member;
whereby the resultant structure upon hardening is usable as a portion of a detector system. - 3. The method of claim 2 in which during the insertion step B, the structure is agitated, whereby a uniform distribution of suspension is achieved through-out each compartment.
- 4. The method of claim 3 in which the agita-tion occurs at ultrasonic frequency.
- 5. The method of claim 2 in which the side wall members and the interior members are selected from the group consisting of tungsten and tantalum.
- 6. The method of claim 1 in which the phosphor is selected from the group consisting of BaFCl:Eu, ZnCDS:Ag, ZnCdS:Ag,Ni, CsI:Tl, CsI:Na, CaF2:Eu, Gd2O2S:Tb, LaOBr:Dy, LaOBr:Tm, LaOBr:Tb, La2O2S:Tb, Bi4Ge3O12, ZnS, ZnSe, ZnTe, CdS, CdSe, CdTe, and NaI.
- 7. The method of claim 1 in which in step A, at least one wavelength conversion substance is also mixed in the liquid matrix substance, whereby the optical output of the phosphor is matched to the spectral response of a photoelectrically responsive detector.
- 8. The method of claim 7 in which the wave-length conversion substance is rhodamine.
- 9. The method of claim 6 in which the phosphor is BaFCl:Eu.
- 10. The method of claim 1 in which the liquid matrix is an epoxy.
- 11. The method of claim l in which the liquid matrix substance is a silicone polyimide copolymer.
- 12. The method of claim 1 in which the harden-ing is caused by chemical activation.
- 13. The method of claim 1 in which the hardening is caused by ultraviolet irradiation.
- 14. The method of claim 1 in which the hardening is caused by curing at a temperature sufficient to harden the matrix material.
Priority Applications (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CA000394702A CA1147111A (en) | 1978-10-13 | 1982-01-21 | Distributed phosphor scintillator structures |
Applications Claiming Priority (2)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CA313,333A CA1122335A (en) | 1977-11-21 | 1978-10-13 | Distributed phosphor scintillator structures |
| CA000394702A CA1147111A (en) | 1978-10-13 | 1982-01-21 | Distributed phosphor scintillator structures |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CA1147111A true CA1147111A (en) | 1983-05-31 |
Family
ID=25668790
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CA000394702A Expired CA1147111A (en) | 1978-10-13 | 1982-01-21 | Distributed phosphor scintillator structures |
Country Status (1)
| Country | Link |
|---|---|
| CA (1) | CA1147111A (en) |
Cited By (1)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US11047747B2 (en) | 2017-03-27 | 2021-06-29 | Firouzeh Sabri | Light weight flexible temperature sensor kit |
-
1982
- 1982-01-21 CA CA000394702A patent/CA1147111A/en not_active Expired
Cited By (1)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| US11047747B2 (en) | 2017-03-27 | 2021-06-29 | Firouzeh Sabri | Light weight flexible temperature sensor kit |
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