WO2018117838A1

WO2018117838A1 - Bismuth silicate as detector material for tof-pet

Info

Publication number: WO2018117838A1
Application number: PCT/NL2017/050869
Authority: WO
Inventors: Dennis Robert Schaart; Stefan Enrico Karl BRUNNER
Original assignee: Technische Universiteit Delft
Priority date: 2016-12-23
Filing date: 2017-12-22
Publication date: 2018-06-28
Also published as: NL2018060B1

Abstract

The invention provides a detector system (100) comprising a detector unit (120), the detector unit (120) comprising: (a) a detector material (10) capable of generating a first radiation (11) and a second radiation (12) pursuant to interaction of a high-energy photon with the detector material (10); and a detector (20) comprising a photodetector (21), the photodetector comprising a digital silicon photomultiplier (Si-PM) photodetector capable of detecting said first radiation (11) and said second radiation (12), said detector (20) radiationally coupled with said detector material (10); wherein said detector material (10) comprises a single crystalline or ceramic A₄M₃O₁₂ material, wherein A comprise Bi and wherein M comprises one or more of Si and Ge, wherein at least part of M comprises Si, wherein the first radiation (11) has a sum of rise and a decay time at least 10 times faster than a sum of rise and decay time of the second radiation (12), and wherein the detector system (100) is configured to determine in a validation routine radiation as first radiation (11) in dependence of a predefined threshold value of a sum of first and second radiation (11, 12) in a predetermined time window t_w.

Description

BISMUTH SILICATE AS DETECTOR MATERIAL FOR TOF-PET FIELD OF THE INVENTION

The invention relates to a detector system and to a method of detecting, especially with such detector system. The invention further relates to a computer program product for executing such method. BACKGROUND OF THE INVENTION

Time-of-flight positron emission tomography systems using Cherenkov emission are known in the art. WO2010/085139, for instance, describes method of time-of-flight positron emission tomography of an object comprising a positron emitter located in a target region, said target region being spatially surrounded by an array of detector elements, the method comprising the steps of: using for the detector elements a material capable of generating a first event and a second event pursuant to a passage of a 51 1 keV photoelectron therein; detecting in the detector element the first event for timing the time-of-flight measurement; detecting in the detector element the second event for carrying out energy measurement. For the first event Cherenkov photons are used and for the second event corresponding scintillation photons are used. For the energy measurement an energy threshold is selected for suppressing undesired events, the undesired events comprise photons generated pursuant to Compton scattering. The material of the detector elements - according to WO2010/085139 - has at least 30% efficiency of detecting the full energy of an incoming 511 keV photon, has a scintillation yield of not less than 10 000 photons per event, and a Cherenkov yield for the full absorption of a 51 1 keV photon of at least 10 photon. The material of the detector elements is selected from a group consisting of: BGO, CsI(Tl), LaBr:,. SUMMARY OF THE INVENTION

It appears that there is a need to further improve current TOF-PET systems in terms of time resolution and/or sensitivity. Hence, it is an aspect of the invention to provide an alternative detector system e.g. for such TOF-PET system, which preferably further has an improved time resolution and/or sensitivity.

It was surprisingly found that with the inorganic scintillating-crystal material (61481300) a very good time response to gamma radiation of the energy of 51 1 keV can be achieved. This is unexpected as this detector material was expected to respond rather slowly to gamma radiation. This material indeed appears to provide a slow response by scintillation, but also a fast response by Cherenkov emission and/or other fast luminescent processes. With our measurements the fast response could be utilized and time resolutions of 190-320 ps FWHM could be achieved (in a non-optimized system). This fast response allows e.g. applying this material for time- of-flight positron emission tomography which is an in vivo molecular imaging tool used in medicine.

Hence, in a first aspect the invention provides a detector system comprising a detector unit, the detector unit comprising a detector material capable of generating a first radiation and a second radiation pursuant to interaction of a high-energy photon with said detector material (more especially A4M3O12 material, see below), and a detector, especially comprising a photodetector capable of detecting single (optical) photons, even more especially a digital silicon photomultiplier (Si-PM) photodetector, capable of detecting said first radiation and said second radiation, wherein said detector material especially comprises AJVLiOn material, even more especially single crystalline or ceramic A4M3O12 material, wherein A especially comprises Bi and wherein M especially comprises one or more of Si and Ge, wherein especially at least part of M comprises Si, wherein the first radiation has a sum of rise and a decay time at least 5 times faster, such as especially at least 10 times faster, like even more especially at least 20 times faster, than a sum of a rise and decay time of the second radiation, and wherein the detector system is further especially configured to determine in a validation routine radiation as first radiation in dependence of a predefined threshold value of a sum of first and second radiation in a predetermined time window t_w. Especially, the high-energy photon has an energy selected from the range of 0.1-12 MeV. For instance, the high-energy photon may have an energy selected from the range of about 250 - 750 keV. The high-energy photon is especially a 51 1 keV annihilation photon generated in a positron-electron annihilation process. Further, the detector may thus also be used for electrons/positrons from pair-production processes, (prompt) γ imaging, etc.. The high-energy photons may thus be ionizing radiation.

With this invention, coincidence time resolutions of 190 ps FWHM (720 ps FWTM) have been achieved for 51 1 keV annihilation photon pairs, within a temperature range of -30°C to +20°C using a semi-optimized setup (small crystals good time resolution but bad detection efficiency). For a more realistic setup as it would be the case for a practical PET system, coincidence time resolutions well below 320 ps FWHM (2500 ps FWTM) could be achieved (long crystals high detection efficiency), though further optimization may be possible. Comparably, best state-of-the-art systems allow 300-400 ps FWHM with the much more expensive Lu_{2( 1}.x)Y_2xSiOs_:Ce (0<x<l, such as Lu₂Si0₅:Ce, possibly codoped with e.g. Ca or Mg). The present detector system may especially be useful for positron emission tomography (PET), especially time-of- flight (TOF) PET.

Radiomolecular imaging is a branch of nuclear medicine aimed at visualizing physiological processes in-vivo using radionuclides. A radioactive isotope is used to label part of the molecules of a radiopharmaceutical, or tracer, which is designed to target a particular feature of interest within the subject. By imaging the distribution of this tracer it is possible to obtain information about molecular processes non-invasively. The tracer distribution is measured by detecting the gamma rays emitted by the radionuclide. Positron emission tomography (PET) is a molecular imaging technique that makes use of tracers labeled with positron- emitting isotopes. PET is usually integrated with CT in a so-called PET/CT scanner to simultaneously obtain functional and anatomical information. Currently, the most common application of PET is to diagnose tumors and to find cancer metastasis. However, PET is also employed for treatment response monitoring, for diagnosing brain diseases, to assess cardiac viability, and as a research tool to e.g. study brain or heart function or to support drug development.

The positrons emitted by the administered radiotracer annihilate with electrons in the body almost instantaneously. The PET imaging technique is based on the coincident detection of the two 511 keV annihilation photons that are emitted in opposite directions as a result of this process. The two annihilation photons interaction points define a so-called line-of-response (LOR) on which the annihilation must have taken place (Figure 1). The combined information of many millions of LORs measured during a PET acquisition is used to produce a 3D image of the estimated tracer distribution using analytical or probabilistic image reconstruction methods. If the time difference between the two moments of interaction can be measured with sufficient precision (< 500 ps) this so called time- of-flight (TOF) information can be used to estimate the segment of the LOR on which the annihilation occurred. This helps to improve the signal-to-noise ratio of the image. Thus, PET image quality is determined largely by the performance of the detectors used to measure the position and time of interaction of the annihilation photons. Current PET detectors are based on scintillation crystals that convert the energy of the annihilation photons into tiny flashes of optical radiation and photomultiplier tubes (PMTs) that convert these optical signals into electronic pulses.

As indicated above, the invention provides a detector system. The detector system comprises a detector unit. The detector system may also comprise a plurality of detector units. In general, the detector system will also include a control system, configured to control the detector unit(s). Hence, the one or more detector units may be functionally coupled with a control system. The control system may read out the signal of the detector unit and convert this in plots, tables, etc..

The detector unit comprises a detector material ("^•scintillator") and a detector ("sensor" or "photosensor" or "photodetector"). The term "detector material" refers to the material that generates, upon receipt of the high energy photon radiation, the first and second radiation ("detector material radiation"), which first and second radiation can be detected with the detector. Photons with energies selected from the range of 0.1-12 MeV, such as selected from the range of 250-750 keV, may herein also be indicated as "high-energy photons". Optionally, the term detector material may refer to a plurality of different materials. The detector has a detecting function and the detector material has a radiation conversion function, converting the high energy radiation into radiation that can be detected by the detector. Also, the detector may have a conversion function as optical photons are converted to electrons. The terms first radiation or second radiation especially refer to "optical radiation", especially in one or more of the (V)UV, visible and IR. The detector material may herein also be indicated as "scintillator". Hence, the detector material may comprise a scintillator crystal or scintillator ceramic. As known in the art, a scintillator is a material that exhibits scintillation, i.e. upon exciting with ionizing radiation the scintillator generates luminescence. Some luminescent materials, when struck by an incoming radiation, absorb its parts or all of the energy and scintillate, (i.e. re-emit the absorbed energy in the form of visible light). Especially, the detector is radiationally coupled with said detector material. This implies that at least part of the radiation generated in the detector material can be received by the detector ("sensor"). The detector system and the detector material may thus be configured in a radiation receiving relationship. Especially, the term "radiationally coupled" means that the detector and the detector material are associated with each other so that at least part of the radiation emitted by the detector material is received by the detector (and at least partly converted into a detector signal by the detector). Instead of the term "radiationally coupled" also the term "optically coupled" may be used. The detector material may be in physical contact with the detector (with e.g. optical glue or grease in between).

The detector material is chosen to be capable of generating a first radiation and a second radiation pursuant to an interaction of a high-energy photon with the detector material. The interaction may comprise transmission (passage) or absorption. For instance, a first step may absorption of the annihilation photon in the detector material. This may then result in a photoelectron (or Compton electron, depending on the type of interaction). The resulting hot electron may then create both the Cherenkov emission (if traveling faster than the speed of light in the material) and scintillation (by ionization). Therefore, absoprtion or inelastic scattering of an high-energy photon may lead to the first and second radiation.

Hence, outside the detector unit, positron-electron annihilations may create high energy photons, such as in a human body in a PET system. One annihilation in general creates two (or very infrequently three) annihilation photons, which are an example of high-energy photons. However, the detection system may also be used in combination with other sources of high-energy photons (that can be detected with the detection system). The photoelectron(s) generated upon photoelectric absorption of a high energy photon in the detector material thereby will create upon passage in the detector material the first radiation and the second radiation.

Especially, the detector material comprises A4M3O12 material, even more especially single crystalline or ceramic A4M3O12, wherein in embodiments A comprise Bi and wherein M comprises one or more of Si and Ge, wherein especially at least part of M comprises Si. Such material is amongst others described in E. Galenin et al., Czochralski Growth and Characterization of Mixed BGO-BSO Crystals, International Conference on Oxide Materials for Electronic Engineering OMEE-2014, which is incorporated herein by reference. Note however that the herein described single crystal is not necessarily prepared according to the Czochralski method; other methods, such as the Bridgman method, may also be possible.

It was surprisingly found that such A4M3O12 material, especially with a relatively high Si content can be used with relatively high time resolution for e.g. PET applications. Hence, in specific embodiments the single crystalline or ceramic A4M3O12 material comprises A₄(Gei_-xSix)30i2, wherein 0 < x < 1, especially wherein 0 < x < 1, such as x especially being at least 0.1, like at least 0.5, especially wherein x is at least 0.9. Therefore, in further specific embodiments, the single crystalline or ceramic A4M3O12 material comprises single crystalline or ceramic Bi₄Si 0i2.

In yet other embodiments, A may be selected from the group consisting of Sb and Bi, especially Bi. Optionally, A may in addition to Bi comprise one or more other trivalent elements. Further, in yet other embodiments, M may be selected from the group consisting of Si, Ge, Ti, and Zr, especially Si and Ge. In further specific embodiments, A M3O12 comprises Bi₄Ge30i₂. However, especially M comprises at least Si (optionally in addition to Si one or more other tetra valent elements).

Especially, A may comprise a rare earth (RE), such as dysprosium (Dy), especially in addition to bismuth. In general, the amount of the rare earth, such as Dy, is relatively low. The rare earth, such as Dy, may be used as dopant in the crystalline or ceramic Bi₄Si 0i₂. For instance, in optical processes bismuth may transfer and/or exchange energy to dysprosium. Hence, in embodiments the single crystalline or ceramic A4M3O12 material comprises wherein part of A refers to a rare earth), such as Dy. The remaining part of A may e.g. be Bi and/or Sb, especially at least Bi. Therefore, in specific embodiments the single crystalline or ceramic A4M3O12 material comprises

wherein y is selected from the range of 0-1, especially 0-0.2, more especially selected from the range of up to 0.1. The value of y may be zero, but in specific embodiments, larger than zero. RE refers to one or more rare earth elements.

RE refers to a rare earth (ion) or a combination of different rare earth ions. Especially, the rare earth includes one or more of cerium (Ce), dysprosium (Dy), erbium (Er), europium (Eu), gadolinium (Gd), holmium (Ho), lutetium (Lu), neodymium ( d), praseodymium (Pr), samarium (Sm), terbium (Tb), thulium (Tm), and ytterbium (Yb). RE may alternatively or additionally also refer to one or more of lanthanum (La), scandium (Sc) and yttrium (Y), but especially in combination with one or more of the afore-mentioned rare earth elements. Especially, the rare earth element includes one or more of praseodymium (Pr), neodymium (Nd), samarium (Sm), europium (Eu), gadolinium (Gd), terbium (Tb), dysprosium (Dy), holmium (Ho), erbium (Er), thulium (Tm), and ytterbium (Yb), even more especially at least one or more of europium and dysprosium.

In specific embodiments the single crystalline or ceramic A4M3O12 material comprises (Bii._yDy_y)4M₃0i2, wherein y is selected from the range of 0-0.2, especially selected from the range of up to 0.1. Hence, in embodiments 0 < y < 0.2, especially 0 < y < 0.2. In embodiments, y may especially be at least 0.01, such as at least 0.02. As indicated above, in embodiments M may comprise one or more of Ge and Si, especially at least Si. Hence, in embodiments the single crystalline or ceramic A4M3O12 material comprises (Bii_-yDy_y)₄Si30i2.

The fact that in the chemical formula part of A or part of M may include other ions than Bi and Si, respectively, which are used as main embodiments, does not necessarily include that the substituents (or dopants) like e.g. Dy and Ge, respectively, precisely occupy the crystallographic A and M sites. The substituents or dopants may also occupy other crystallographic places. The crystalline material may include vacancies, as known in the art. The crystalline material may in embodiments also some non-stoichiometry, as known in the art. However, with the present validation method, also other scintillator materials may be applied.

The detector material comprises a single crystal of the detector material or a ceramic material. Hence, the detector material may include a single crystal or a ceramic body (further indicated as "ceramic" or "ceramic material". Therefore, in specific embodiments of the detector material are herein indicated as single crystalline or ceramic material. The single crystal or ceramic may have dimensions of e.g. at least 1 mm', such as at least 0.5 cm', such as at least 1 cm³, such as in the range of 1-500 cm'. The transmission through 1 cm crystalline or ceramic material (of the single crystal or ceramic, respectively) under perpendicular radiation with visible light (wavelength selected from the range of 380-780 run) is especially at least 30%, such as at least 50%, like especially at least 70%, such as at least 80%.

Further, in embodiments the surfaces other than the exit surface of the detector material is relatively rough and thus has a relatively high surface roughness. Hence, in embodiments the single crystalline or ceramic material, especially the crystalline A4M3O12 material, a photodetector directed surface, wherein the detector directed surface has a surface roughness Ra (arithmetical mean deviation along 1000 μιη) equal to or lower than 100 nm (but especially at least 5 11m), such as equal to or lower than 40 nm, such as in the range of 5-20 nm. The one or more other surface(s) of said single crystalline or ceramic A₄M₃O₁₂ material may have a higher surface roughness R_a, such as especially at least 2 times higher, like at least 5 times higher, such as at least 10 times higher. For instance, the surface roughness may be equal to or higher than about 100 nm, such as in the range of about 500 - 800 nm, especially however not higher than 1 μηι. Such high roughness may be obtained with not polishing the surfaces after cutting. The surface roughness may be detected with a (contact) profileometer. Especially, the roughness R_a is at least 5 nm.

Further, the time resolution can be enhanced when the surfaces, i.e. the exit surface of the detector material, and one or more of the remaining surfaces, are optimized individually. Hence, in embodiments the single crystalline or ceramic material, especially the crystalline A4M3O12 material, a photodetector directed surface, wherein the detector directed surface has a surface roughness R_a (arithmetical mean deviation measured along 1000 μηι) equal to or lower than 1 μηι, such as equal to or lower than 100 nm, such as in the range of 5-20 nm. The one or more other surface(s) of said single crystalline or ceramic A₄M₃O₁₂ material may have a surface roughness R_a, such as equal to or lower than 1 μιη, such as equal to or lower than 100 nm, such as in the range of 5-20 nm. For instance, the surface roughness may be equal to or higher than about 100 nm, such as in the range of about 500 - 800 nm, especially however not higher than 1 μιη. Especially, the roughness R_a is at least 5 nm. Such high roughness may be obtained by not polishing the surfaces after cutting.

However, with the aspect of different surface roughnesses, also other scintillator materials may be applied than A₄M₃O_H. In an embodiment a ceramic or a single crystal is applied having different faces, with one or more faces having a higher surface roughness, such as at least 2 times higher, like at least 5 times higher, such as at least 10 times higher, than one or more other faces, especially than a single surface that is used to be directed (during use of the detector system) to the detector. In other embodiments, a ceramic or a single crystal is applied having different faces, with one or more faces having the same or a different surface roughness than the others, especially than a single face that is used to be directed (during use of the detector system) to the detector.

The proposed material advantageously has a fast radiation and a slow radiation, wherein the fast radiation may have a relative low intensity, but may be used to achieve the high time resolution. The more intense slow radiation, which may still be relatively fast (in the orders of tens or hundreds of nano seconds) may be used to validate the relatively low intensity fast radiation. This will further be elucidated below.

Especially, the first radiation has a sum of rise and decay time equal to or less than 500 ps (pico seconds). For instance, the decay time may be in the range of 20-250 ps, such as in the range of 1-20 ps, or even faster. Further, especially, the second radiation has a decay time equal to or more than 5 ns (nano second).

For instance, the first radiation comprises Cherenkov emission, and the second radiation comprises scintillation. Cherenkov emission, also known as Vavilov-Cherenkov emission, is electromagnetic radiation emitted when a charged particle passes through a dielectric medium at a speed greater than the phase velocity of light in that medium. However, the fast radiation is not necessarily limited to Cherenkov emission, but may additionally or alternatively also comprise e.g. intraband luminescence, or similar fast processes.

Especially, the first radiation has a sum of rise and a decay time at least 5 times faster, even more especially at least 10 times faster, than a rise and decay time of the second radiation. Hence, the total time frame of rise and decay time of the first radiation, wherein a substantial part of the first radiation is detectable is much shorter than the total time frame of rise and decay time of the second radiation, wherein a substantial part of the second radiation is detectable. Hence, the second radiation decays essentially longer than the first radiation. This may be e.g. also be used to separate these radiations based on their different time behaviors. Further, the (different) time behavior may be also used to separate the(se) radiation(s) from possible background events, such as dark counts occurring in the photodetector. Hence, the decay time of the fast emission is at least 5 times faster, even more especially at least 10 times faster, than the decay time of the second radiation.

For instance, assuming a first radiation having a rise time of 1 ps or less, and a decay time of 10 ps, the sum of the rise and decay time is about 10-1 1 ps. Assuming a second radiation having a rise time of 1 ps or less, and a decay time of about 5 ns, then the sum of the rise and decay time is about 5 ns. The second radiation is thus much slower than the first radiation.

Especially, the single crystalline or ceramic A₄M₃O₁₂ material is configured to generate equal to or more than 100 second radiation photons and equal to or less than 30 first radiation photons pursuant to said passage of said photoelectron created by absorption of said high-energy photon therein. Hence, the present materials seem to be much more sensitive than some prior art materials. The single crystalline or ceramic A4M3O12 material may generate less than about 2000, such as less than about 1000 second radiation photons.

The detector may absorb (annihilation created) high energy photons, such as e.g. photons of 51 1 keV. An electron of the detector material, especially of the single crystalline or ceramic material (more especially of an atom comprised by the material), is released and is quasi-free in the material, propagating with a kinetic energy given by 51 1 keV with the binding energy of the electron subtracted (typically less than 116 keV). This "hot" electron (with a typical energy of about 450 keV) is fast enough to generate Cherenkov emission during its slowing down in the material (slowing down, e.g., by excitation and ionization, which may result in the generation of secondary electrons, where a few of the secondary electrons even may have sufficient energy to generate Cherenkov emission themselves). After thermalization of the secondary electrons (there may be thousands), electrons and holes may recombine via different channels. For some of those channels they are releasing scintillation photons. The Cherenkov emission photons and the scintillation photons can be measured with the detector.

The detector, which may also be indicated as "photodetector" or "sensor" is especially a photodetector capable of detecting single photons in the visible wavelength range, and especially also in the VUV, UV and/or the infrared region. Further, especially the photodetector is capable of detecting single photons with a high time resolution. Especially, the detector comprises a silicon photomultiplier (Si-PM) photodetector, even more especially a digital silicon photomultiplier (Si-PM) photodetector. Such photodetector is capable of detecting said first radiation and said second radiation. Especially, it appears that with such digital photodetector now the fast radiation can be detected, whereas with an analogous detector, the fast radiation may be less easily retrievable from the background signal.

Especially, the detector may have a single photon time resolution (SPTR) < 500 ps, especially < 200 ps, even more especially < 50 ps, like at least 0.01 ps, such as at least 0.1 ps.

In specific embodiments, the detector comprises a kl *ll tile of independent digital silicon photomultiplier (Si-PM) photodetectors, which may also be called "dies", wherein each photodetector (or "die") comprises k2*k2 pixels, and wherein each pixel comprises k3 *13 single photon avalanche diodes (SPADs), wherein kl,l l,k2,12,k3, and 13 are each independently at least 1, and wherein at least two of kl,ll,k2,12,k3, and 13 are at least 4, 16, 1600, 3200, etc.. Especially, kl and 11 are each independently at least 1. Hence, kl and 11 may each independently be selected from the range of 1-24, such as 2-4. Especially, kl=ll . Especially, k2 and 12 are each independently at least 2. Hence, k2 and 12 may each independently be selected from the range of 1-24, such as 2-4. Especially, k2=12. For instance, each die may comprise four pixels. Especially, k3 and 13 are each independently at least 2. Hence, k3 and 13 may each independently be selected from the range of 4-200, such as 20-80. Especially, k3=13, though this is not necessarily the case. For instance, each pixel may include 3200 single photon avalanche diodes (SPADs). Especially, the detector comprises at least 200 single photon avalanche diodes, such as at least 400 single photon avalanche diodes, like at least 20.000. A large number of single photon avalanche diodes may allow a high energy resolution.

Optionally, the detector may comprise a PMT-like system, such as combination of a photocathode and a vacuum-electron multiplier. Hence, the detector may e.g. comprise a SiPM or vacuum-based single photon detector. In yet other embodiments, an electron emission membrane may be applied, such as described by Y. Bilevych, S. E. Brunner, H. W. Chan, E. Charbon, H. van der Graaf, C. W. Hagen, G. Nutzel, S. D. Pinto, V. Prodanovic, D. Rotman, F. Santagata, L. Sarro, D. R. Schaart, J. Sinsheimer, J. Smedley, S. Tao, and A. M. M. G. Theulings, "Potential applications of electron emission membranes in medicine," Nucl. Instruments Methods Phys. Res. Sect. A Accel. Spectrometers, Detect. Assoc. Equip., vol. 809, pp. 171-174, 2016. By placing, in vacuum, a stack of transmission dynodes (tynodes) on top of a CMOS pixel chip, a single free electron detector could be made with outstanding performance in terms of spatial and time resolution. The essential object is the tynode: an ultra-thin membrane, which emits, at the impact of an energetic electron on one side, a multiple of electrons at the other side.

As indicated above, in specific embodiments the detector system is configured to determine in a validation routine radiation as first radiation in dependence of a predefined threshold value of a sum of first and second radiation in a predetermined time window t_w. As the number of first radiation photons is relatively low, the validation routine is used to confirm whether a detected fast photon is indeed a first radiation photon or noise. To this end, (mainly) the second radiation, which provides much more photons, is applied. Hence, during a short time frame after detecting a possible first photon, the sum of "^•fast" and "^•slow" photons are counted. When a threshold value is reached, it can be confirmed that the high energy photoelectron has indeed entered the detector material and generated first radiation (and second radiation). The sum of the first and second radiation here refers especially to the sum of first radiation photons and second radiation photons.

In specific embodiments, the validation routine comprises identifying a possible first radiation photon at t_¾ measuring during said time window t_w starting at t₀ the number i¾ of first and second radiation photons, and determining said possible first radiation photon as first radiation photon when said number n₂ of first and second radiation photons is larger than said predefined threshold value of first and second radiation, wherein 0.05 ns < t_w < 50 ns, especially 0.1 ns < t_w < 50 ns, such as 1 ns < tw < 50 ns, like 5 ns < t_w < 50 ns. Further, especially the threshold value of first and second radiation is defined as a number of first and second radiation photons selected from the range of 1-100, such as in the range of 1-30, such as in the range of 4-30, such as at least 4 photons, like selected from the range of 5- 20 photons. Hence, after counting e.g. a total of 7 first and second radiation photons a detected fast signal can indeed be ascribed to a first radiation photon.

Hence, the second radiation may e.g. be used for energy determination. Further, as the number of photons detected from the first radiation may be too low for robust and reliable validation, the second radiation may also be used to validate the first radiation. The first radiation especially provides the basis for a high time resolution.

The present detector system may especially be useful for positron emission tomography (PET). Hence, in a further aspect the invention provides a positron emission tomography (PET) detector system comprising said detector system as defined herein. Further, especially time-of-flight measurements may be done with such detector system. Therefore, in specific embodiments the PET detector system is a time-of-flight (TOF) PET detector system. Such (time-of-flight) positron emission tomography (PET) detector system may especially comprise a plurality of such detector units, for instance configured in a detector ring. However, in general the invention can be used to improve the time resolution in any radiation detector based on scintillators which detect photons in the energy range of 0.1-12 MeV, such as especially 0.1-1 MeV. In yet a further aspect, the invention also provides a method of detecting a high energy photon, the method comprising detecting a first radiation and a second radiation with especially a detector system as defined herein. As indicated above, the second radiation comprises photons generated by said high-energy photon. Yet further, as indicated above especially the first radiation has a sum of rise and decay time at least 5 times, such as at least 10 times faster than (a sum of rise and decay time of) the second radiation.

Further, in specific embodiments the detector system is configured to determine in a validation routine radiation as first radiation in dependence of a predefined threshold value of a sum of first and second radiation in a predetermined time window t_w, and wherein said detector material comprises a single crystalline or ceramic A4M3O12 material, wherein A may especially comprise Bi and wherein M may especially comprise one or more of Si and Ge, and wherein at least part of M comprises Si (see further also above for alternative embodiments). Even more especially, the method is a time-of-flight positron emission tomography method, wherein the method comprises time-of-flight measurement of two positron-electron annihilation created photons, generated by a single annihilation.

In yet further aspects, the invention also provides a computer program product, when running on a computer (which computer) is functionally coupled to detector system as defined herein, is capable of bringing about the method as described herein.

BRIEF DESCRIPTION OF THE DRAWINGS

Embodiments of the invention will now be described, by way of example only, with reference to the accompanying schematic drawings in which corresponding reference symbols indicate corresponding parts, and in which:

Figs, la-lb schematically depicts an imaging principle of PET; and

Figs. 2a-2c schematically depict some aspects of the detector system.

The schematic drawings are not necessarily on scale. DETAILED DESCRIPTION OF THE EMBODIMENTS

Fig. la schematically depicts an imaging principle of PET: (a) after annihilation of a positron and an electron, two 511 keV annihilation photons, indicated with reference P, are emitted in (almost) opposite directions. Reference E indicates e.g. a positron emitting nucleus; (b) when two interactions are simultaneously detected within a ring of detectors surrounding the patient, it is assumed that an annihilation, indicated with reference A, occurred on the so-called line-of-response (LOR) connecting the two interactions. By recording many LORs the activity distribution can be topographically reconstructed. Reference 100 indicates a detector system, and reference 1000 indicates a positron emission tomography detector system. References 1 1, 12 indicate a (fast) first radiation and a (slower) second radiation, respectively. Reference DR indicates a detector ring.

Figs. 2a-2c schematically depicts some aspects of the detector system 100. The detector system 100 comprises the detector unit 120. The detector unit 120 comprises a detector material 10 capable of generating first radiation 1 1 and second radiation 12 pursuant to the absorption of an annihilation photon therein (which photon is created by a positron-electron annihilation). The detector 20 comprises for instance a digital silicon photomultiplier Si-PM photodetector 21 capable of detecting said first radiation 1 1 and said second radiation 12, said detector 20 radiationally coupled with said detector material 10. The detector 20 has a detector face 24 directed to the detector material 10, especially its photodetector directed surface, which is indicated with reference 15. Reference 14 indicates other surfaces, which may receive high energy photons. Such one or more surfaces may also be indicated as high-energy photon receipt surfaces. Especially, on the photodetector directed surface 15 has a surface roughness, whereas one or more of the others, especially all other surfaces have the same or a different surface roughness.

Fig. 2b very schematically depicts the probability density distribution as function of time after a photoelectron generates radiations 11 and 12. The first radiation 1 1 is fast and the second radiation 12 is slow. The intensities of the radiations are not to scale. As the fast radiation is substantially instantaneous, the max of the curve may be indicated with to. The rise time of the fast radiation / first radiation 11 may be about 1-100 ps. The rise time of the second radiation 12 may be about 1-1000 ps. As can be seen from the graph, the fast radiation has a much faster decay time than the second radiation. Further, the rise time of the first radiation may be equal to or even faster than of the second radiation. Hence, the sum of the rise time and decay time of the first radiation is substantially smaller than the sum of the rise time and decay time of the second radiation.

Especially, the detector system 100 is configured to determine in a validation routine radiation as first radiation 1 1 in dependence of a predefined threshold value of a sum of first and second radiation 12 in a predetermined time window t_w. Reference t_w is also indicated in fig. 2b. Fig. 2c schematically depicts that in an example the detector 20 may comprises a kl *ll tile of independent digital silicon photomultiplier Si-PM photodetectors 21. Each photodetector 21 comprises k2*k2 pixels 22. Further, each pixel 22 may comprise k3*13 single photon avalanche diodes 23 SPADs. The single crystal or ceramic may radiationally be coupled with a single pixel 22.

EXPERIMENTAL

The following pairs of single crystals (3 mm x 3 mm cross section) were measured coupled to pixels of Philips Digital Photon Counters (type 3200).

From these data, it appears that not polishing the surfaces other than the exit surface 15 is advantageous. Further, it appears that the Si containing crystals (here BSO) provide a higher time resolution.

The term "substantially" herein, such as in "substantially all radiation" or in "substantially consists", will be understood by the person skilled in the art. The term "substantially" may also include embodiments with "entirely", "completely", "all", etc. Hence, in embodiments the adjective substantially may also be removed. Where applicable, the term "substantially" may also relate to 90% or higher, such as 95% or higher, especially 99% or higher, even more especially 99.5% or higher, including 100%. The term "comprise" includes also embodiments wherein the term "comprises" means "consists of. The term "and/or" especially relates to one or more of the items mentioned before and after "and/or". For instance, a phrase "item 1 and/or item 2" and similar phrases may relate to one or more of item 1 and item 2. The term "comprising" may in an embodiment refer to "consisting of but may in another embodiment also refer to "containing at least the defined species and optionally one or more other species".

Furthermore, the terms first, second, third and the like in the description and in the claims, are used for distinguishing between similar elements and not necessarily for describing a sequential or chronological order. It is to be understood that the terms so used are interchangeable under appropriate circumstances and that the embodiments of the invention described herein are capable of operation in other sequences than described or illustrated herein.

The devices herein are amongst others described during operation. As will be clear to the person skilled in the art, the invention is not limited to methods of operation or devices in operation.

It should be noted that the above-mentioned embodiments illustrate rather than limit the invention, and that those skilled in the art will be able to design many alternative embodiments without departing from the scope of the appended claims. In the claims, any reference signs placed between parentheses shall not be construed as limiting the claim. Use of the verb "to comprise" and its conjugations does not exclude the presence of elements or steps other than those stated in a claim. Unless the context clearly requires otherwise, throughout the description and the claims, the words "comprise", "comprising', and the like are to be construed in an inclusive sense as opposed to an exclusive or exhaustive sense; that is to say, in the sense of "including, but not limited to". The article "a" or "an" preceding an element does not exclude the presence of a plurality of such elements. The invention may be implemented by means of hardware comprising several distinct elements, and by means of a suitably programmed computer. In the device claim enumerating several means, several of these means may be embodied by one and the same item of hardware. 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.

The invention further applies to a device comprising one or more of the characterizing features described in the description and/or shown in the attached drawings. The invention further pertains to a method or process comprising one or more of the characterizing features described in the description and/or shown in the attached drawings.

The various aspects discussed in this patent can be combined in order to provide additional advantages. Further, the person skilled in the art will understand that embodiments can be combined, and that also more than two embodiments can be combined. Furthermore, some of the features can form the basis for one or more divisional applications.

Claims

CLAIMS:

A detector system ( 100) comprising a detector unit (120), the detector unit (120) comprising:

a. a detector material (10) capable of generating a first radiation (11) and a second radiation (12) pursuant to interaction of a high-energy photon with said detector material (10); and

b. a detector (20) comprising a photodetector (21), the photodetector comprising a digital silicon photomultiplier photodetector capable of detecting said first radiation (1 1) and said second radiation (12), said detector (20) radiationally coupled with said detector material (10); wherein said detector material (10) comprises a single crystalline or ceramic A₄M₃O₁₂ material, wherein A comprise Bi and wherein M comprises one or more of Si and Ge, wherein at least part of M comprises Si, wherein the first radiation (1 1) has a sum of rise and a decay time at least 10 times faster than a sum of rise and decay time of the second radiation (12), and wherein the detector system ( 100) is configured to determine in a validation routine radiation as first radiation (11) in dependence of a predefined threshold value of a sum of first radiation and second radiation (1 1, 12) in a predetermined time window t_w. The detector system (100) according to claim 1, wherein the single crystalline or ceramic A4M3O12 material comprises A4(Gei_-xSi_x)30i2, wherein 0.1<x<l, especially wherein x is at least 0.9.

The detector system (100) according to any of the preceding claims, wherein the single crystalline or ceramic A₄M₃O₁₂ material comprises (B _ yREy)₄M30₁₂, wherein y is selected from the range of 0-0.2, and wherein RE refers to one or more rare earth elements.

The detector system (100) according to any of the preceding claims, wherein the single crystalline or ceramic A₄M₃O₁₂ material comprises (B _ yDy_y)4M₃0₁₂, wherein y is selected from the range of 0-0.1.

5. The detector system (100) according to any one of the preceding claims, wherein the single crystalline or ceramic A4M3O12 material has a photodetector directed surface (15), wherein the photodetector directed surface ( 15) has a surface roughness R_a equal to or lower than 1 μιη along 1000 μιη, and wherein one or more other surface(s) of said single crystalline or ceramic A4M3O12 material have a the same or a different surface roughness, equal to or lower than 1 μηι along 1000 μιη.

6. The detector system (100) according to any one of the preceding claims, wherein the first radiation (11) has a sum of rise and decay time equal to or less than 500 ps, and wherein the second radiation (12) has a sum of rise and decay time equal to or more than 5 ns, wherein the first radiation (1 1) comprises Cherenkov emission, and wherein the second radiation (12) comprises scintillation.

7. The detector system (100) according to any one of the preceding claims, wherein the detector (20) comprises a kl *ll tile of independent digital silicon photomultiplier (Si-PM) photodetectors (21), wherein each photodetector (21) comprises k2*k2 pixels (22), and wherein each pixel (22) comprises k3 *13 single photon avalanche diodes (23) (SPADs), wherein kl,ll,k2,12,k3, and 13 are each independently at least 1, and wherein the detector (20) comprises at least 200 single photon avalanche diodes (23).

8. The detector system (100) according to any one of the preceding claims, wherein the validation routine comprises identifying a possible first radiation photon at to, measuring during said time window t_w starting at to the number n₂ being the sum of first and second radiation photons, and determining said possible first radiation photon as first radiation photon when said number n₂ of first and second radiation photons is larger then said predefined threshold value of first and second radiation, wherein 0. 1 ns < t_w < 50 ns.

9. The detector system (100) according to any one of the preceding claims, wherein said threshold value of first and second radiation (11, 12) is defined as a number of second radiation photons selected from the range of 1-30.

10. The detector system (100) according to any one of the preceding claims, wherein the single crystalline or ceramic A4M3O12 material is configured to generate equal to or more than 100 second radiation photons and equal to or less than 30 first radiation photons pursuant to said interaction of the high- energy photon.

11. The detector system (100) according to any one of the preceding claims, wherein the high-energy photon has an energy selected from the range of 0.1- 12 MeV.

12. A positron emission tomography (PET) detector system (1000) comprising said detector system (100) according to any one of the preceding claims 1-10.

13. The PET detector system (1000) according to claim 12, wherein the PET detector system (1000) is a time-of-flight (TOF) PET detector system (1000).

14. A method of detecting a high-energy photon, the method comprising detecting a first radiation (1 1) and a second radiation (12) with a detector system (100) according to any one of the preceding claims 1-0, wherein the first radiation (1 1) has a sum of rise and a decay time at least 10 times faster than a sum of rise an decay time of the second radiation (12), wherein the detector system (100) is configured to determine in a validation routine radiation as first radiation (11) in dependence of a predefined threshold value of a sum of first and second radiation (11, 12) in a predetermined time window t_w, and wherein said detector material (10) comprises a single crystalline or ceramic A4M3O12 material, wherein A comprise Bi and wherein M comprises one or more of Si and Ge, and wherein at least part of M comprises Si.

15. The method according to claim 14, wherein the method is a time-of-flight positron emission tomography method, wherein the method comprises time- of-flight measurement of two positron-electron annihilation created high- energy photons, generated by a single annihilation.

16. A computer program product, when running on a computer which is functionally coupled to detector system as defined in any one of the preceding claims 1-11, is capable of bringing about the method as described in any one of the preceding claims 14-15.