Abstract:
The present invention relates to mixed oxide materials, methods for their preparation, detectors for ionizing radiation and CT scanners. In particular, a mixed oxide material is proposed having the formula (Y w Tb x ) 3 Al 5-y Ga y O 12 :Ce z , wherein 0.01≦w≦0.99, 0.01≦x≦0.99, 0≦y≦3.5 and 0.001≦z≦0.10 and wherein w+x+3*z=1, whereby the mixed oxide material is doped with at least 10 ppm V.

Description:
CROSS REFERENCE TO RELATED APPLICATIONS 
     This application is a national filing of PCT application Serial No. PCT/IB2014/060019, filed Mar. 21, 2014, published as WO 2014/155256 A1 on Oct. 2, 2014, which claims the benefit of U.S. provisional application Ser. No. 61/805,261 filed Mar. 26, 2013, which is incorporated herein by reference. 
    
    
     FIELD OF THE INVENTION 
     The present invention relates to mixed oxide materials, methods for their preparation, detectors for ionizing radiation and CT (Computed Tomography) scanners. 
     BACKGROUND OF THE INVENTION 
     Detectors for ionizing radiation and, in particular, solid state detectors for ionizing radiation are, e.g., widely used in CT scanners. Such solid state detectors for ionizing radiation comprise, broadly speaking, two main subunits. The first subunit comprises a fluorescent component that is usually referred to as a scintillator or a phosphor which absorbs radiation and in response emits photons in the UV, the visible or the IR region. The second subunit comprises a photodetector which can detect the photons emitted by the scintillator or phosphor and produces corresponding electrical signals. 
     With regard to the above expressions “scintillator” and “phosphor”, it needs to be noted that both are exchangeable terms and are to be understood within the scope of the invention to refer to solid state luminescent materials that, in response to a stimulation by ionizing radiation such as X-rays, β- or γ-radiation, emit radiation with photons of considerably lower energy. 
     The expression “ionizing radiation” within the scope of the invention refers to electromagnetic radiation having energy higher than that of ultraviolet radiation. 
     Detectors for ionizing radiation find broad application in X-ray-based detecting and imaging systems. One of the major medical applications for such detectors and scintillators is in CT scanners. 
     In particular for their application in CT scanners, it is preferable if those scintillators show a high light yield, so that the CT scanner can be run with as low a radiation dose for the patient as possible. Furthermore, the scintillators used in modern CT scanners should have as low an afterglow as possible, as otherwise the scanning process must be slowed down (e.g. by reducing the rotation frequency) to reduce the influence of the afterglow in subsequent images, affecting the speed of the examination. 
     Finally, it is also desirable that the scintillators are as transparent as possible to visible light, as otherwise scattering of the photons produced by the interaction between the ionizing radiation and the scintillator occurs, which results in effective background noise during the imaging process, due to optical absorption of the scintillation light in the scintillator. 
     The two materials that are at the moment commonly used as scintillators for CT scanners are scintillator materials based on Gd 2 O 2 S doped with Pr (GOS) and (Y, Gd) 2 O 3  doped with Eu. While those two materials already give reasonable results, it has been shown that GOS, due to the fact that it is not transparent to visible light but merely translucent, shows a reasonably high scattering leading to undesirable effective noise, whereas the (Y, Gd) 2 O 3 :Eu based systems show a notable afterglow which could be improved upon for the next generation of CT scanners by replacing this scintillator. 
     SUMMARY OF THE INVENTION 
     It is an object of the present invention to provide a mixed oxide material, a method for its preparation, a scintillator, a detector for ionizing radiation and a CT scanner, wherein the scintillator shows a high light yield, a very low afterglow and a high transparency. 
     In a first aspect of the present invention, a mixed oxide material having the formula (Y w Tb x ) 3 Al 5-y Ga y O 12 :Ce z  is presented, wherein 0.01≦w≦0.99, 0.01≦x≦0.99, 0≦y≦3.5 and 0.001≦z≦0.10 and wherein w+x+3*z=1, whereby the mixed oxide material is doped with at least 10 ppm V, preferably at least 25 ppm V. 
     In a further aspect of the present invention, a method for preparing a mixed oxide material as described above is presented that comprises the following steps: a) providing Y 2 O 3 , CeO 2 , Tb 4 O 7 , Al 2 O 3  and Ga 2 O 3  in proportions suitable to obtain the desired mixed oxide, b) impregnating one or several of the solids of step a) with a source of V in the desired amount, c) combining and milling the solids of step a) and step b) in the presence of a suitable dispersant, to obtain a slurry, d) drying the slurry of step c) to obtain a mixed powder, and e) sintering the mixed powder of step d) at a temperature of at least 1400° C. for at least 1 h. 
     In a further aspect of the present invention, a scintillator is presented that comprises the above-mentioned mixed oxide material. 
     In a further aspect of the present invention, a detector for ionizing radiation is presented that comprises the above mentioned mixed oxide material or the above-mentioned scintillator in combination with at least one photodetector. 
     In a further aspect of the present invention, a CT scanner is presented that comprises at least one detector as described above. 
     (Y W Tb x ) 3 Al 5-y Ga y O 12 :Ce z  based materials have been known for some time to be able to interact with ionizing radiation and, as a result, release photons, i.e. to have scintillator properties. The materials known so far, though, show an afterglow that is so high that they have been considered not suitable for modern CT scanners. It has now been found that by doping the above-mentioned mixed oxide materials with minute quantities of vanadium, the afterglow can be reduced significantly without sacrificing the light yields too strongly and therefore creating a scintillator material that is suitable for application in modern CT scanners. 
     In an embodiment, the mixed oxide material is doped with 10 to 250 ppm V, preferably 25 to 200 ppm V. 
     It has been shown that the addition of vanadium to the mixed oxide material in the above-mentioned quantity ranges leads to a good balance between improvements of the afterglow without significant losses in the light yield. 
     In another embodiment of the mixed oxide material, 0.1≦w≦0.9, preferably 0.2≦w≦0.8, more preferably 0.3≦w≦0.6 and even more preferably 0.35≦w≦0.5. 
     In another embodiment of the mixed oxide material, 0.1≦x≦0.9, preferably 0.2≦x≦0.8, more preferably 0.4≦x≦0.7 and even more preferably 0.5≦x≦0.65. 
     In another embodiment of the mixed oxide material, 1≦y≦3.5 preferably 2≦y≦3.5 and more preferably 2.5≦y≦3.5. 
     In still another embodiment of the mixed oxide material, 0.005≦z≦0.05, preferably 0.005≦z≦0.02 and more preferably z=0.01. 
     In a further embodiment, the mixed oxide material has the formula (Y 0.395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01  and is doped with at least 10 ppm V, preferably 10 to 250 ppm V, and, in particular, 25 to 200 ppm V. 
     It has been shown that mixed oxide materials in the above-mentioned composition ranges lead to particularly effective scintillator materials having a high light yield and a low afterglow. 
     In a further embodiment, the mixed oxide material is a single crystalline or a polycrystalline material. 
     In an embodiment of the above-mentioned method, a flux material is added in step c) when combining the solids of step a) step b). 
     By adding a flux material in step c), the diffusion of the different ions during the sintering of step d) can be improved, leading to higher quality material at lower sintering temperatures. 
     In a further embodiment, the detector further comprises a second mixed oxide material or scintillator, whereby the second mixed oxide material or second scintillator has a higher density than the above-described mixed oxide material or scintillator. 
     Through a combination of two different scintillator materials having a different density, X-rays of different energy levels can be detected, whereby the material of lower density generally detects X-rays of a lower energy, and the material of higher density generally detects X-rays of a higher energy. By creating detectors which comprise two different scintillators or scintillator materials, a detector can be created that detects two different types of X-rays which, for example, in CT scanners gives more information about the body or part of the body that is being examined. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
       These and other aspects of the invention will be apparent from and elucidated with reference to the embodiments and examples described hereinafter. In the following drawings 
         FIG. 1  shows a schematic diagram of a CT scanner according to the present invention, 
         FIG. 2  shows a schematic diagram of a first embodiment of a detector for ionizing radiation according to the present invention, 
         FIG. 3  shows a schematic diagram of a second embodiment of a detector for ionizing radiation according to the present invention, and 
         FIG. 4  shows a schematic diagram of a third embodiment of a detector for ionizing radiation according to the present invention. 
     
    
    
     DETAILED DESCRIPTION OF THE INVENTION 
     Examples 
     Examples 1 to 4 
     Stoichiometric amounts of Y 2 O 3  (Rhodia), CeO 2  (Neo Materials), Tb 4 O 7  (Guangdong and Neo Materials), Al 2 O 3  (Baikowski) were weighed in, in proportions to create mixed oxide materials having the formula (Y 0.395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01 . In order to dope these materials with 25 ppm, 50 ppm, 100 ppm and 200 ppm, respectively of V, a corresponding amount of NH 4 VO 3  was dissolved in ethanol, mixed with the Al 2 O 3 , precipitated and dried on a rotary evaporator. The modified Al 2 O 3  obtained this way was then employed in the solid state synthesis of the desired mixed oxides. The solid starting materials were mixed and milled with heptane in agate pots. After the mixing process, the samples were dried in a tube oven to remove the mixing liquid, and the samples were sintered in a horizontal tube furnace (Entech 01820 series) at 1550° C. in an aluminium crucible for 4 hours in a H 2 /N 2  flow in order to reduce Ce 4+  to Ce 3+  and Tb 4+  to Tb 3+ . 
     The obtained samples were tested for photoluminescence, whereby the photoluminescence emission spectra were recorded at room temperature using a xenon lamp with an Edinburgh instruments FLSP920 spectrometer featuring double monochromators to improve resolution and to reduce stray light. Afterglow measurements were performed using X-ray excitation and a photodiode. The light yield was measured by determining the area under the emission curve and expressed as a percentage yield compared to Comparative Example 1 in each of the tables. 
     The results of the measurements compared to (Y 0.395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01  without vanadium are shown below in Tables 1 and 2. 
     Table 1, light yields for Examples 1 to 4 compared to (Y 0.395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01  without the addition of vanadium. Percentage values are relative to Comparative Example 1. 
                                                                                                                 280 nm Tb 3+     345 nm Ce 3+     378.5 nm Tb 3+             band excitation   band excitation     7 F 6 - 5 D 3  excitation                        % Light        % Light       % Light           V Content    Area   Yield   Area   Yield   Area   Yield                    Comparative    0 ppm   2.26E+08   100.0   8.33E+07   100.0   7.28E+07   100.0       Example 1                                   Example 1    25 ppm   2.14E+08   94.9   8.75E+07   105.0   8.24E+07   113.2       Example 2    50 ppm   1.97E+08   87.1   7.95E+07   95.4   7.37E+07   101.2       Example 3   100 ppm   1.87E+08   82.6   7.41E+07   88.9   6.43E+07   88.3       Example 4   200 ppm   1.59E+08   70.5   7.02E+07   84.3   5.99E+07   82.3                    
Table 2, afterglow for Examples 1 to 4 compared to (Y 0.395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01  without the addition of vanadium. Percentage values are relative to Comparative Example 1. Ppm values are relative to initial intensity of the corresponding material.
 
     
       
         
               
               
               
               
               
               
               
             
               
               
               
               
               
             
               
               
               
               
               
               
               
               
             
               
               
               
               
               
               
               
               
             
           
               
                   
               
               
                   
                 % afterglow  
                   
                   
                   
                   
                   
               
               
                   
                 after 5 ms 
                   
                   
                   
                   
                   
               
             
          
           
               
                   
                 (relative to  
                 PD 5 ms 
                 PD 500 ms 
                 PD 2100 ms 
               
               
                   
                 Comparative  
                 [ppm] 
                 [ppm] 
                 [ppm] 
               
             
          
           
               
                   
                 Example 1) 
                 μ 
                 σ 
                 μ 
                 σ 
                 μ 
                 σ 
               
               
                   
               
             
          
           
               
                 Comparative 
                 100.0 
                 4628 
                 15 
                 122 
                 1 
                 43 
                 1 
               
               
                 Example 1 
                   
                   
                   
                   
                   
                   
                   
               
               
                 Example 1 
                 5.7 
                 263 
                 9 
                 &lt;10 
                   
                 &lt;10 
                   
               
               
                 Example 2 
                 8.2 
                 380 
                 65 
                 31 
                 1 
                 &lt;10 
                   
               
               
                 Example 3 
                 10.3 
                 476 
                 42 
                 10 
                 2 
                 12 
                 3 
               
               
                 Example 4 
                 2.3 
                 108 
                 7 
                 &lt;10 
                   
                 &lt;10 
               
               
                   
               
             
          
         
       
     
     Comparative Examples 1 to 4 
     As comparative examples, (Y 0.0395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01  doped with Ti, Cr and Mn as well as (Y 0.395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01  without the addition of any dopant was prepared. 
     The synthesis of the Comparative Examples took place in analogy to the synthesis of Examples 1 to 4 with the NH 4 VO 3  being omitted or replaced with Ti-n-butoxide, Cr(NO 3 ) 3 .9H 2 O, and Mn(NO 3 ) 2 .4H 2 O respectively. A dopant level of 50 ppm was used for all Comparative Examples comprising a dopant. 
     The analysis of the materials obtained according to Comparative Examples 1 to 4 took place according to the conditions described above for Examples 1 to 4 and the results are shown below in Tables 3 and 4. 
     Table 3, light yields for Comparative Examples 1 to 4. Percentage values are relative to Comparative Example 1. 
                                                                                             280 nm Tb3+               band excitation                            % Light               Dopant   Area   Yield                            Comparative   none   6429590    100.0           Example 1                       Comparative    Ti   5734990    89.2           Example 2                       Comparative    Cr   5907660   91.9           Example 3                       Comparative   Mn   6055820    94.2           Example 4                        
Table 4, afterglow for Comparative Examples 1 to 4. Percentage values are relative to Comparative Example 1. Ppm values are relative to initial intensity of the corresponding material.
 
     
       
         
               
               
               
               
               
             
               
               
               
               
               
               
               
               
             
               
               
               
               
               
               
               
               
             
           
               
                   
               
               
                   
                 % afterglow after  
                   
                   
                   
               
               
                   
                 5 ms (relative to  
                 PD 5 ms 
                 PD 500 ms 
                 PD 2100 ms 
               
               
                   
                 Comparative  
                 [ppm] 
                 [ppm] 
                 [ppm] 
               
             
          
           
               
                   
                 Example 1) 
                 μ 
                 σ 
                 μ 
                 σ 
                 μ 
                 σ 
               
               
                   
               
             
          
           
               
                 Comparative 
                 100.0 
                 4628  
                 15 
                 122  
                 1 
                 43 
                 1 
               
               
                 Example 1 
                   
                   
                   
                   
                   
                   
                   
               
               
                 Comparative 
                 49.4 
                 2285  
                 17 
                 147  
                 1 
                 52 
                 1 
               
               
                 Example 2 
                   
                   
                   
                   
                   
                   
                   
               
               
                 Comparative 
                 94.1 
                 4355  
                 40 
                 115  
                 1 
                 39 
                 1 
               
               
                 Example 3 
                   
                   
                   
                   
                   
                   
                   
               
               
                 Comparative 
                 268.1 
                 12406 
                 12 
                 392  
                 2 
                 85 
                 1 
               
               
                 Example 4 
                   
                   
                   
                   
                   
                   
                   
               
               
                   
               
             
          
         
       
     
     The data from the Examples and Comparative Examples show clearly that the addition of vanadium to the (Y 0.395 Tb 0.595 ) 3 Al 5 O 12 :Ce 0.01  mixed oxide material significantly reduces the afterglow without impacting the light yield excessively. Furthermore, the Comparative Examples show that this seems to be a very specific effect for vanadium, as related d-group metals such as titanium, chrome or manganese do not show such an effect. 
     In  FIG. 1 , a CT scanner in its entirety is denoted with reference numeral  10 . The CT scanner  10  comprises a rotating gantry  12  on which on opposing sides an X-ray source  14  and a detector array  16  are arranged. The detector array  16  consists of a number of individual X-ray detectors one of which is for exemplary purposes denoted here with the reference numeral  18 . The rotating gantry  12  is arranged such that the X-ray source  14  and the detector array  16  are on opposing sides of an examination area  20  into which a patient  22  is inserted. In use, the X-ray source emits a wedge-shaped, cone-shaped or otherwise shaped X-ray beam directed into the examination area  20 , in the instant case in the direction of a patient  22 . The patient  22  can be linearly moved in a z direction (perpendicular to the plane of drawing), while the X-ray source  14  and, correspondingly, the detector array  16  rotate around the z axis. In general, the rotating gantry  12  rotates simultaneously with the linear advancement of the patient  22  leading to a generally helical trajectory of the X-ray source  14  and, correspondingly, the detector array  16  around the examination area  20 . However, other imaging modes can also be employed, such as a single- or multi-slice imaging mode in which the gantry rotates as the subject support remains stationary, to produce a generally circular direction of the X-ray source  14  and, correspondingly, the detector array  16  over which an axial image is acquired. 
     As can be seen in the picture, the detector array  16  is arranged on the gantry  12  on the opposing side of the X-ray source  14 , so that in use the X-rays emitted by the X-ray source  14  pass through e.g. a patient  22  and are then detected by the detector array  16 . The detector array  16  generally comprises a multitude of detectors  18 , whereby the detector array  16  can be a single line of detectors  18  or two-dimensional array of detectors  18 . A more detailed explanation of the function of the detectors  18  within the detector arrays  16  is given below in respect to various embodiments of the detectors shown in  FIGS. 2 to 4 . 
     In  FIG. 2 , a first embodiment of a detector for ionizing radiation is denoted in its entirety with reference numeral  30 . The detector  30  comprises two subunits, namely the scintillator  32  and the photodetector  34 . The photodetector  34  comprises a photodiode  36  which is arranged such that the active area of the photodiode  36  is facing the scintillator  32 . 
     In use, the detector is arranged such that the scintillator  32  points towards the source of potential source of radiation to be detected. The scintillator  32  thereby e.g. consists of the material described above under Example 1. If ionizing radiation, for example X-rays, now impinges on the scintillator  32 , the scintillator  32  interacts with those X-rays and, in response, releases one or multiple photons which are emitted from the scintillator  32  and can be detected by the photodiode  34  generating an electric signal indicating the presence of X-rays. In order to improve the yield of photons detected by the diode  34 , the scintillator  32  can be covered on one or several sides not facing the photodetector with a material reflective to the emitted photons. 
     In  FIG. 3 , a second embodiment of a detector for ionizing radiation is designated in its entirety with reference numeral  40 . Again, this detector  40  comprises two subunits, namely a scintillator  42  and a photodetector  44 . In contrast to the embodiment of  FIG. 1 , in this case the scintillator  42  consists of two different scintillator materials, a first scintillator material  46  and a second scintillator material  48 . The first scintillator material  46 , in the instant case, is the material of the above-mentioned Example 2, and the second scintillator material  48  is thereby a scintillator material having a higher density than the first scintillator material  46 . In the instant case, the second scintillator material  48  is a Gd 2 O 2 S doped with Pr. 
     Corresponding to the first scintillator material  46  and the second scintillator material  48 , the photodetector  42  comprises two photodiodes, a first photodiode  50  and a second photodiode  52 . In use, X-rays with different energies impinge on the detector  40  from top, i.e. from the direction of the first scintillator material  46 . Due to its lower density, the first scintillator material  46  absorbs X-rays of lower energy and in response thereto emits photons of a first frequency. After passing through the first scintillator material  46 , the X-rays strike the second scintillator material  48 , whereby through the interaction with the second scintillator material  48 , photons of a second wavelength are emitted. 
     The first photodiode  50  is now equipped with a first filter  54  which filters out photons of the second wavelength, ensuring that only the photons of the first wavelength, i.e. the photons generated by the first scintillator material  46 , are detected by the first photodiode  50 . 
     Correspondingly, the second photodiode  52  is equipped with a second filter  56  which blocks photons of the first wavelength, ensuring that only photons of the second wavelength, i.e. photons generated by the second scintillator material  48 , reach the second photodiode  52  and are detected thereby. 
     Through the above-mentioned set-up, it is possible with the detector  40  to detect and differentiate X-rays of two different energy levels and create corresponding signals increasing the amount of information available in the CT scan. 
     In  FIG. 4 , a third embodiment of a detector for ionizing radiation is designated in its entirety with reference numeral  60 . The detector for ionizing radiation  60  is similar in function to the detector  40  of  FIG. 3 , but shows a different design. 
     Again, the detector  60  consists of a scintillator  62  and a photodetector  64 . In this case again the scintillator  62  consists of a first scintillator material  66  and a second scintillator material  68 . The first scintillator material  66  is, thereby, for example the material of Example 3, whereby the second scintillator material  68  again is Pr doped Gd 2 O 2 S, i.e. a material of higher density than the first scintillator material  66 . In contrast to the embodiment of  FIG. 3 , in  FIG. 4  the photodetector  64  is not arranged underneath the scintillator  62  but on the side of it, whereby a first photodiode  70  is arranged on the side of the first scintillator material  66  and a second photodiode  72  is arranged on the side of the second scintillator material  68 , when seen in the direction of the incoming ionizing radiation to be detected, as indicated by arrow  74 . 
     Both scintillator materials  66  and  68  are covered on those sides which do not face the first photodiode  70  and second photodiode  72 , respectively, with a coating that is reflective to photons in the wavelength range emitted by the first and second scintillator material  66  and  68 , respectively, yet transparent to ionizing radiation. 
     In use, the ionizing radiation travels in the direction indicated by arrow  74  towards the first scintillator material  66 , whereby, due to the lower density, the lower energy part of the ionizing radiation interacts with the first scintillator material  66  and stimulates the emission of one or several photons. Due to the reflective coating, on the outside of the first scintillator material  66 , the photons can only exit the first scintillator material  66  towards the first photodiode  70  and are detected thereby. After traveling through the first scintillator material  66 , the ionizing radiation travels to the second scintillator material  68 , whereby, due to the fact that the density of the second scintillator material  68  is higher than the first scintillator material  66 , higher energy radiation is absorbed and, as a result, a second set of photons is generated. Again, due to the fact that the second scintillator  68  material is covered by reflective material on those sides not facing the second photodiode, the photons can only exit the second scintillator material  68  towards the second photodiode  72  and are detected thereby. 
     Again, due to the fact that each scintillator material interacts with radiation of a specific energy level and in response emits photons which are directed to specific photodiodes and detected thereby, different X-rays can be detected with the detector for ionizing radiation creating more information, for example about a body to be investigated in a CT scanner. 
     While the invention has been illustrated and described in detail in the drawings and foregoing description, such illustration and description are to be considered illustrative or exemplary and not restrictive; the invention is not limited to the disclosed embodiments. Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure, and the appended claims. 
     In the claims, the word “comprising” does not exclude other elements or steps, and the indefinite article “a” or “an” does not exclude a plurality. A single element or other unit may fulfill the functions of several items recited in the claims. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage. 
     Any reference signs in the claims should not be construed as limiting the scope.