Ceramic scintillator, photon counting x-ray detector, and method for manufacturing ceramic scintillator

Abstract

A ceramic scintillator according to an embodiment is formed of a garnet compound containing Pr. The absorptivity of the ceramic scintillator at a wavelength of 320 [nm] is 0.15 or less relative to the absorptivity at a wavelength of 280 [nm].

Description

TECHNICAL FIELD

[0001] Embodiments of the present application relate to a ceramic scintillator, a photon counting type X-ray detector, and a manufacturing method of a ceramic scintillator. BACKGROUND

[0002] Imaging systems employing radiation such as X-rays are widely used for industrial uses such as baggage inspection and non-destructive inspection, medical uses such as X-ray diagnostic apparatuses and X-ray computed tomography (CT) apparatuses. The mainstream of the detection method of the imaging system employing X-rays is currently the energy integrating type, which is generally configured by a combination of a luminescent material and a light detector.

[0003] However, in the energy integrating type, it is not possible to obtain energy information of X-rays, and the exposure dose becomes a problem. In recent years, in order to solve these problems, development of an X-ray detector employing a photon counting method is being conducted. The photon counting type X-ray detector is also called a photon detector or a photon counting type detector.

[0004] The photon counting method has a feature of performing pulse signal processing on incident X-ray photons one by one. Furthermore, in the photon counting method, there are classified a direct type employing a semiconductor such as CdTe to directly convert X-rays into an electric signal, and an indirect type converting X-rays into light by a luminescent material and then converting the light into an electric signal by a light detector.

[0005] In the photon counting method, it is possible to measure the number of carriers in the direct type, and it is possible to measure the energy of X-ray photons from the number of luminescent photons in the indirect type, and noise components are easily removed, so it is possible to obtain data with a high S / N ratio. According to these advantages, in medical uses, it is expected that a reduction in exposure dose will be achieved by new functions of diagnosis or low radiation measurement. The mainstream of the research and development in the photon counting method is the direct type that can obtain a high energy resolution. However, semiconductor materials such as CdTe are expensive, it is difficult to obtain uniform characteristics, it is difficult to manufacture a large-area detector, and a signal pile-up phenomenon caused by a low response speed of the material becomes a problem of the direct type.

[0006] On the other hand, in the photon counting method, as a light detector in the indirect type, a photomultiplier tube having a relatively high multiplication rate is generally employed. However, the photomultiplier tube is large in shape, and it is difficult to constitute a pixel with a narrow gap, which becomes a problem of the indirect type. A Si-based light detector operating in the Geiger mode, that is, a silicon photomultiplier tube (Si-PM) developed in recent years solves the problems of the photomultiplier tube described above, has a multiplication rate close to that of the photomultiplier tube, and is low in cost. Therefore, in the indirect type, it is expected that the use of the silicon photomultiplier tube will gradually expand in the future.

[0007] PRIOR ART DOCUMENTS PATENT LITERATURE Patent Literature 1: Japanese Patent Application Laid-Open No. 2018-2974 Patent Literature 2: Japanese Patent Application Laid-Open No. 2010-235388 SUMMARY

[0008] PROBLEMS TO BE SOLVED BY THE INVENTION A technical problem common to the direct type and the indirect type is that high count rate cannot be dealt with. The count rate indicates the number of photons of X-rays incident per unit area per unit time, and corresponds to the intensity of X-rays.

[0009] In the case of imaging based on photon counting type X-ray measurement in a large number in a short time, a detector corresponding to high count rate is required. Such a case is required in X-ray CT, in which the count rate required is 10 8 [cps / mm 2 ] or more. This corresponds to X-ray photons being incident at an average interval of 10 [nsec] to an area of 1 [mm 2 ]2. In the case where the count rate exceeds the capability of the detector, overlap of pulse signals occurs, and it is difficult to measure the correct X-ray photon energy.

[0010] On the other hand, measurement for a long time with the count rate suppressed causes problems of reduction in flux and reduction in time resolution due to the increase in the imaging time. The main factor that determines the count rate that the detector can deal with is the response speed of the material, and improvement is expected.

[0011] The problem to be solved by the present application is to provide a ceramic scintillator that can deal with high count rate, a photon counting type X-ray detector, and a method for manufacturing a ceramic scintillator.

[0012] MEANS FOR SOLVING THE PROBLEM The ceramic scintillator according to the embodiment is formed of a garnet compound containing Pr. The absorbance at a wavelength of 320 [nm] of the ceramic scintillator is 0.15 or less relative to the absorbance at a wavelength of 280 [nm]. BRIEF DESCRIPTION OF DRAWINGS

[0013] Figure 1 is a schematic diagram showing the configuration of the photon counting type X-ray detector according to the embodiment.

[0014] Figure 2 is an enlarged sectional view of the partial region shown in (B) in the photon counting type X-ray detector according to the embodiment. Figure 1

[0015] Figure 3 ​is a graph in which the relationship between the absorbance at a wavelength of 320 [nm] relative to the absorbance at a wavelength of 280 [nm] (320 nm absorbance / 280 nm absorbance) and the relative light emission amount relative to the relative decay time constant (relative light emission amount / relative decay time constant) is shown in table form.

[0016] Figure 4 is a graph in which the change in the absorbance spectrum in the example is shown in a graph.

[0017] Figure 5 is a graph in which the relationship between the absorbance at a wavelength of 320 [nm] relative to the absorbance at a wavelength of 280 [nm] (320 nm absorbance / 280 nm absorbance) and the relative light emission amount relative to the relative decay time constant (relative light emission amount / relative decay time constant) is shown in table form. DETAILED DESCRIPTION

[0018] Hereinafter, embodiments of a ceramic scintillator, a photon counting type X-ray detector, and a manufacturing method of a ceramic scintillator will be described in detail with reference to the drawings.

[0019] (Photon counting type X-ray detector) Figure 1 (A) to (C) are schematic diagrams showing the configuration of the photon counting type X-ray detector to which the embodiments relate. Figure 1 (A) is a plan view of the photon counting type X-ray detector to which the embodiments relate. Figure 1 (B) is a side view showing the channel direction CH of the photon counting type X-ray detector to which the embodiments relate. Figure 1 (C) is a side view showing the slice direction SL of the photon counting type X-ray detector to which the embodiments relate.

[0020] Figure 1 (A) to (C) show a photon counting type X-ray detector (hereinafter referred to simply as "X-ray detector") 1 to which the embodiments relate. Furthermore, Figure 1 (B) and (C) show a collimator device 3 in addition to the X-ray detector 1. Figure 1 (C) shows an X-ray tube 2 in addition to the X-ray detector 1.

[0021] The X-ray detector 1 can be provided on a rotating frame of a stand device. The X-ray detector 1 is provided with n (n: plural) X-ray detection elements In. The X-ray detection elements In are arranged in a matrix shape two-dimensionally in the channel direction and the slice direction SL. Further, the channel direction means the direction of expansion of the fan beam X-rays irradiated from the X-ray tube 2, and the slice direction is the thickness direction of the fan beam X-rays.

[0022] The X-ray incident surface of the X-ray detector 1 is formed by the X-ray incident surfaces of the X-ray detecting elements In. The X-ray detecting elements In are arranged, for example, 1000 or so in the channel direction CH and 64 in the slice direction SL.

[0023] The X-ray tube 2 is disposed in the rotating frame of the stand device in opposition to the X-ray detector 1. The X-ray tube 2 is a vacuum tube that generates X-rays by irradiating hot electrons from a cathode (filament) toward an anode (target) by applying a high voltage. For example, the X-ray tube 2 is a rotating anode type X-ray tube that generates X-rays by irradiating hot electrons against a rotating anode.

[0024] The collimator device 3 has a plurality of collimator plates that have a function of absorbing scattered X-rays. The plurality of collimator plates are constituted by plates extending in the slice direction SL, that is, plates that are disposed upright in a manner of zoning the X-ray detecting elements In in the channel direction CH (one-dimensional collimation). Alternatively, the plurality of collimator plates are constituted by plates extending in the slice direction SL, that is, plates that are disposed upright in a manner of zoning the X-ray detecting elements In in the channel direction CH and plates extending in the channel direction CH, that is, plates that are disposed upright in a manner of zoning the X-ray detecting elements In in the slice direction SL (two-dimensional collimation). The plate surfaces of the collimator plates are adjusted in inclination in a manner of being parallel to the direction of X-ray irradiation, that is, the X-ray irradiation direction E from the X-ray focus F of the X-ray tube 2. Figure 1 (C) shows a case where the collimator device 3 is a one-dimensional collimator. Also, at times, a detector in which the collimator device 3 is combined in the X-ray detector 1 is referred to as a "photon counting type X-ray detector".

[0025] Also, the X-ray detector 1 can be constituted by arranging a plurality of detector modules by modularizing a prescribed number of X-ray detecting elements in the X-ray detecting elements In. Also, the collimator device 3 can be constituted by arranging a plurality of collimator modules by modularizing a prescribed number of collimator plates.

[0026] Figure 2 is an X-ray detecting element In in the X-ray detector 1 Figure 1 (B) is an enlarged sectional view of the partial region R.

[0027] The X-ray detecting elements In are provided on the ceramic substrate 4. The X-ray detecting elements In each have a ceramic scintillator 11 and a photoelectric conversion element 12.

[0028] The ceramic scintillator 11 is an element that converts an incident X-ray into a photon and emits it. Here, the X-ray generally has a prescribed X-ray energy distribution. It can be considered that an X-ray having a specific X-ray energy is a group of X-ray particles in a number corresponding to the magnitude of the X-ray energy. The ceramic scintillator 11 converts an X-ray particle into a photon with a prescribed probability in a state in which the group of X-ray particles is maintained. That is, if an X-ray is incident in the ceramic scintillator 11, a group of photons in a number corresponding to the X-ray energy is emitted at substantially the same time in accordance with the X-ray energy.

[0029] The ceramic scintillator 11 is substantially a light conversion element having a rectangular shape or a cubic shape. The ceramic scintillator 11 is disposed in a manner that its X-ray incident surface is substantially perpendicular to an X-ray irradiation direction, that is, an X-ray incident direction E, and a side surface parallel to the X-ray incident direction E is substantially parallel to the channel direction CH and the slice direction SL.

[0030] The photoelectric conversion element 12 has a substantially plate-like rectangular shape, converts an incident photon into an electric signal, and outputs the electric signal. The electric signal is an electric pulse signal corresponding to each of the incident photons. If a group of photons is simultaneously incident in the photoelectric conversion element 12, a pulse signal having a wave height corresponding to the number of photons constituting the group of photons is output. The photoelectric conversion element 12 is a so-called semiconductor device suitable for photon counting, for example, a silicon photomultiplier (Si-PM). The silicon photomultiplier is a high-performance semiconductor photodetector that can measure photo counting and can be used for analog measurement such as scintillation detection. The silicon photomultiplier is an element in which a plurality of avalanche photodiodes (APDs) operating in a Geiger mode are connected in parallel as pixels.

[0031] The photoelectric conversion element 12 receives the photon emitted from the ceramic scintillator 11 and outputs an electric signal in a pulse shape. In a case where the intensity of the transmitted X-ray of the detected body is extremely low, the group of photons in accordance with the X-ray energy is emitted in a state of being spread in a time axis direction in the ceramic scintillator 11. At this time, the photoelectric conversion element 12 outputs a pulse signal having a wave height corresponding to the magnitude of the X-ray energy in a state of being spread in the time axis direction in a number corresponding to the dose of the X-ray having the X-ray energy. Therefore, as long as the pulse signals output within a certain time are counted by the wave height, the dose of the transmitted X-ray of the detected body can be known by the X-ray energy. In addition, as long as the pulse signals output within a certain time are counted regardless of the wave height, the dose of the transmitted X-ray of the detected body as a whole can be known.

[0032] On the photoelectric conversion element 12, there are wire patterns (not shown) formed on the ceramic substrate 4. The electrical signals from the photoelectric conversion element 12 are output to an external processing device (not shown) through these wire patterns. The electrical signals output from the photoelectric conversion element 12 can be used for projection data collection based on photon counting.

[0033] Here, the detectors used in photon counting methods generally have very high sensitivity to X-rays. Because of this high sensitivity, even with a low X-ray dose, a sufficient signal-to-noise ratio (S / N) can be obtained even with few photons emitted from the scintillator. However, if the X-ray dose is high, the pulse signals overlap, resulting in a phenomenon known as "pile-up," which prevents signal decomposition along the time axis.

[0034] Similar to indirect and direct photon counting methods, there is the issue of count rate. Count rate refers to the number of X-ray photons incident per unit area per unit time. To achieve imaging based on photon-counting X-ray detectors, detectors with high count rates are required. For example, an X-ray CT scanner requires a count rate of 10-1. 8 [cps / mm 2 This means that X-ray photons are incident at an average interval of 10 [nsec] to 1 [mm]. 2 The area of ​​[the area is missing]. Currently, scintillators capable of handling such high count rates are not yet practical. When the count rate exceeds the scintillator's capability, accumulation occurs, making it difficult to measure accurate photon energy. On the other hand, suppressing the count rate for long-term measurements leads to a decrease in throughput and temporal resolution.

[0035] The count rate of a scintillator is primarily determined by the material's response speed. Therefore, efforts are underway to develop scintillator materials that improve response speed in a way that avoids long measurement times and prevents buildup.

[0036] (Ceramic scintillators and their manufacturing methods) The main factor determining the response speed of the indirect type is the decay time constant of the scintillator. To achieve the aforementioned count rate of 10... 8 [cps / mm 2], the decay time constant of the luminescence of the scintillator is particularly important, and the decay time constant of the luminescence is preferably about 17 [nsec] or less. By making the decay time constant of the luminescence 17 [nsec] or less, as described above, it is possible to prevent pileup and achieve imaging of a practical level. Thus, experiments were performed with the aim of setting the decay time constant of the luminescence of the scintillator to about 17 [nsec] or less. As a result, it was found that the luminescence amount can be 10,000 [ph / MeV] or more. This is because, if the luminescence amount is 10,000 [ph / MeV] or more, the S / N ratio and the energy resolution are not problematic, and a detector with higher precision can be obtained.

[0037] That is, the decay time constant is preferably 17 [nsec] or less and the luminescence amount is preferably 10,000 [ph / MeV] or more. Thus, the fluorescent material of the ceramic scintillator 11 used as the light conversion element is formed of a garnet compound containing Pr. The fluorescent material of the ceramic scintillator 11 is defined as Condition 1: The absorbance at a wavelength of 320 [nm] is 0.15 or less relative to the absorbance at a wavelength of 280 [nm].

[0038] Here, the decay time constant is defined as follows: the time at which the luminescence intensity generated by irradiation of a short-time pulse becomes maximum is set to zero, and the time at which the luminescence intensity decreases to 1 / e (= 0.3679) of the maximum luminescence intensity. For example, by performing time-correlated single photon counting (TCSPC) using a pulsed X-ray tube, it is possible to record the time change in the luminescence intensity, and the decay time constant is calculated from this value. Further, the absorbance is the absorbance of the absorption spectrum measured by absorption spectrometry. The above-described condition 1 is a condition in which the absorbance at a wavelength of 320 nm, which is the peak value of the absorption spectrum of each fluorescent material as the luminescence wavelength, is expressed as a ratio relative to the absorbance at a wavelength of 280 nm, which has a relatively high absorbance.

[0039] In addition, the composition formula of the garnet compound of the ceramic scintillator 11 is represented by R3M5O 12 Further preferably, the following condition 2 or 3 is satisfied.

[0040] Condition 2: The element R in the composition contains Pr and at least one or more selected from Y, La, Ce, Pr, Nd, Sm, Eu, Gd, Tb, Dy, Ho, Er, Tm, Yb, and Lu, The element M in the composition contains Al and Ga.

[0041] Condition 3: The element R in the composition contains Pr, The element M in the composition includes Al and Ga and at least one or more selected from Sc, Si, Ge, and Sn.

[0042] Next, the results of investigating the light emission amount and the decay time constant of the light emission of the phosphor of the embodiment having the composition satisfying the above condition 1 and the phosphor of the comparative example not satisfying the above condition 1 are described. The phosphor of the embodiment and the phosphor of the comparative example are produced through the following processes.

[0043] First, in the first process, a powder of a mixture of the oxide of the element R and the oxide of the element M (oxide powder) is filled in an alumina container and mixed. In the second process, the mixed powder is fired at a temperature of 1300 [°C] or higher. In the third process, the product obtained by the second process is filled in an alumina container and fired at a temperature of 1200 [°C] or higher in a nitrogen-hydrogen mixed atmosphere. In the fourth process, the product obtained by the third process is shaped. In the fifth process, a ceramic scintillator is produced by sintering the product obtained by the fourth process.

[0044] Sintering is performed once. By performing sintering once, a ceramic scintillator having a small composition deviation and good crystallinity can be obtained, and a scintillator having a small decay time constant can be obtained because Ga volatilization can be suppressed, unlike when sintering is performed two or more times.

[0045] The phosphor is formed of the elements contained in the above composition, and contains no element other than the above elements except for inevitable impurities. This is because the light emission decay time constant varies depending on the contained elements, and thus when a large amount of impurities are contained, the light emission decay time constant sometimes increases. Further, as long as the required light emission amount and / or the required light emission decay time constant are satisfied, impurities can be contained in an amount of 100 ppm or more. Here, when Si functions as a sintering aid, it remains as an impurity in the grain boundary of the finished ceramic scintillator. Therefore, when Si remains in the grain boundary, it can be determined that it is added as a sintering aid, and when it is present in the base material, it can be determined that it is added as a raw material. In the ceramic scintillator of the embodiment, except for the case where several tens of ppm of Si is present as an impurity, it means that a compound containing Sn, Si, Ge, or at least one of them is not present in the grain boundary, and a garnet compound containing Si, Ge, or Sn is contained in the base material.

[0046] (Above conditions 1 to 3) Use Figures 3-5This section explains the relationship between the absorptivity at 320 nm and the absorptivity at 280 nm (320 nm absorptivity / 280 nm absorptivity), corresponding to the types of elements R and M in the composition of the garnet compound, and the relative luminescence intensity relative to the relative decay time constant (relative luminescence intensity / relative decay time constant). Here, relative luminescence intensity represents the percentage of luminescence intensity when the luminescence intensity in Example 1 is used as a baseline (100). Furthermore, relative decay time constant represents the percentage of decay time constant when the decay time constant in Example 1 is used as a baseline (100).

[0047] right Figure 3 Please provide an explanation. Figure 3 Examples 1-16 represent elements R and M in the composition that satisfy condition 2 above, and Examples 17-20 represent elements M that satisfy condition 3 above. For example, Example 1 represents a garnet compound containing Lu and Pr as element R in a mass ratio of 0.998:0.002 and Al and Ga as element M in a mass ratio of 0.9:0.1. In Examples 1-20, condition 1 is satisfied, that is, the 320nm absorbance / 280nm absorbance is 0.15 or less. Generally, the decay time constant of luminescence and the luminescence amount are in a trade-off relationship, so the performance is considered based on the relative luminescence amount (relative luminescence amount / relative decay time constant) relative to the relative decay time constant. In Examples 1-20, the relative luminescence amount / relative decay time constant is 0.9 or more. In this case, the decay time constant is about 17 [nsec] or less, and the luminescence amount is 10000 [ph / MeV] or more, which is a preferred range of performance. It should be noted that, comparing Examples 17-20, the case where element M contains a tetravalent element (Examples 18-20) has a lower 320nm / 280nm absorbance and a larger relative luminescence / relative decay time constant compared to the case where element M does not contain a tetravalent element (Example 17). Therefore, element M preferably contains a tetravalent element.

[0048] In condition 3 above, element M in the composition sometimes contains one or more of the tetravalent elements Si, Ge, and Sn. Of the luminescent element Pr, Pr is the one that emits light. 3+ ,Pr 4+ It does not emit light but absorbs visible light. Therefore, Pr 4+ The presence of Pr is the main reason for the reduction in the luminescence intensity and decay time constant of the scintillator. 3+ Therefore, although most of them are Pr 3+ However, Pr exists 4+ The issue of trace residues. Here, if a divalent element is added, Pr easily becomes tetravalent to maintain charge balance, thus increasing the absorption rate at 320 nm.Figure 4 The Ca shown 2+ 、Sr 2+ Mg 2+ The luminescence intensity and decay time constant of the scintillator are further reduced. On the other hand, when a tetravalent element is added, Pr easily becomes trivalent in order to maintain charge balance, thus reducing the absorption rate at wavelength 320 nm. Figure 4 The Si shown 4+ Ge 4+ Sn 4+ Similarly, it can inhibit Pr 4+ The residue makes Pr 3+ Stabilization improves properties. However, the element M in the composition is not limited to containing tetravalent elements.

[0049] right Figure 5 Please provide an explanation. Figure 5 Comparative Examples 1-15 indicate that elements R and M in the composition do not satisfy either condition 2 or condition 3 above (in Comparative Examples 1-15, element M includes elements other than Al and Ga, therefore condition 2 is not satisfied). In Comparative Examples 1-15, the 320nm absorbance / 280nm absorbance is 0.17 or higher. In Comparative Examples 1-15, condition 1 is also not satisfied, i.e., the relative luminescence / relative decay time constant is less than 0.90. In this case, the characteristics are unlikely to be within the preferred range.

[0050] The ceramic scintillator formed from a garnet compound containing Pr described above provides a fluorescent material with a small decay time constant for luminescence. Furthermore, by using this fluorescent material as a ceramic scintillator for medical X-ray detectors, it is possible to achieve the required reactivity.

[0051] According to at least one embodiment described above, a ceramic scintillator that corresponds to a high count rate, a photon counting type X-ray detector having a ceramic scintillator, and a method for manufacturing a ceramic scintillator can be provided.

[0052] Further, the ceramic scintillator 11 is not limited to the case of being used for the photon counting type X-ray detector provided with a silicon photomultiplier in the X-ray CT apparatus. For example, the ceramic scintillator 11 can also be used for an X-ray detector provided with a photodiode in the X-ray CT apparatus. In addition, the ceramic scintillator 11 can also be used for a flat panel detector (FPD) provided with a CMOS (Complementary Metal Oxide Semiconductor) in the X-ray diagnostic apparatus. Further, the ceramic scintillator 11 can also be used for a photon counting type detector provided with a silicon photomultiplier in a PET (Positron Emission Tomography) apparatus. In addition, the ceramic scintillator 11 can also be used for imaging for industrial use such as baggage inspection and non-destructive inspection.

[0053] The embodiments of the present application have been described, but these embodiments are presented as examples and are not intended to limit the scope of the application. These novel embodiments can be implemented in other various ways, and various omissions, substitutions, and changes can be made within the scope of the gist of the application. These embodiments and modifications thereof are included in the scope and gist of the application, and are also included in the scope of the application and equivalents thereof recited in the claims.

Claims

1. A ceramic scintillator which is a ceramic scintillator formed of a garnet compound containing Pr, wherein 320 [nm] is 0.15 or less relative to the absorbance at a wavelength of 280 [nm].

2. The ceramic scintillator of claim 1, wherein, The composition formula of the garnet compound is represented by R3M5O 12 wherein R represents a rare earth element, M represents a trivalent metal element, and O represents oxygen. The element R in the composition of the garnet compound contains Pr and at least one or more selected from the group consisting of Y, La, Ce, Pr, Nd, Sm, Eu, Gd, Tb, Dy, Ho, Er, Tm, Yb, and Lu, The element M in the composition contains Al and Ga.

3. The ceramic scintillator of claim 1, wherein, The composition formula of the garnet compound is represented by R3M5O 12 wherein R represents a rare earth element, M represents a trivalent metal element, and O represents oxygen. The element R in the composition of the garnet compound contains Pr, The element M in the composition contains Al and Ga and at least one or more selected from the group consisting of Sc, Si, Ge, and Sn.

4. A photon counting type X-ray detector comprising: the ceramic scintillator according to any one of claims 1 to 3, and a silicon photomultiplier.

5. A method of manufacturing a ceramic scintillator according to any one of claims 1 to 3, comprising the following steps: First step, a mixture of oxide powders of elements R and M in the composition formula R3M5O 12 of the garnet compound is filled into an alumina container and mixed; a second step of performing sintering of the mixture after mixing at a temperature of 1300°C or higher; a third step of filling a product obtained by the second step into an alumina container and performing sintering at a temperature of 1200°C or higher in a nitrogen-hydrogen mixed atmosphere; a fourth step of performing molding of a product obtained by the third step; and a fifth step of performing sintering of a product obtained by the fourth step to manufacture the ceramic scintillator.

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