Detector, three-dimensional direct positron imaging unit, and method to estimate the differential of the radiation dose provided to cancer cells and healthy tissues during hadrotherapy

Abstract

Disclosed is a detector for a positron imaging unit, comprises a hollow body with an inner cylindrical wall and an outer wall spaced apart from the inner cylindrical wall. The hollow body includes a scintillating material, suitable to emit photons once hit by a 511 keV -ray, and one or more pairs of photo-detecting units (e.g. comprising PMTs or SiPM) for detecting photons emitted by the scintillating material; each photo-detecting unit of a pair being placed at opposite ends of the inner cylindrical wall along a radial direction. The scintillating material has scintillation decay time lower than 10 ns, an atomic number greater than 10, and a high scintillation yield greater than 8,000 photons/MeV, and comprises a mixture of xenon and argon. An imaging unit including the detector and a method to estimate the differential of the dose of radiation provided in a subject to cancer cells and to surrounding tissues in the course of hadrotherapy is also disclosed.

Claims

1. A detector for a positron imaging unit, comprising: a hollow body having an inner cylindrical wall and an outer wall spaced apart from the inner cylindrical wall, the hollow body including: a scintillating material suitable to emit photons once hit by a 511 keV -ray; one or more pairs of photo-detecting units for detecting photons emitted by the scintillating material, each photo-detecting unit of a pair comprising two photodetectors, and a first unit of the pair being placed at a surface of the inner cylindrical wall directly opposing a second unit of the pair; wherein the scintillating material has a scintillation decay time less than 10 ns, an atomic number greater than 10, and a photoelectron yield greater than 8,000 photons/MeV; and wherein the scintillating material comprises at least one of .sup.40Ar, argon depleted in .sup.39Ar, an argon-xenon mixture, or traces of xenon dissolved in argon.

2. The detector according to claim 1, wherein the scintillating material is kept in a liquid state, and the detector comprises a cryostat adapted to keep the scintillating material in the liquid state.

3. The detector according to claim 2, wherein the scintillating material further comprises at least one additional organic dopant or wavelength shifter.

4. The detector according to claim 2, further comprising at least one wavelength shifter, wherein the at least one wavelength shifter is coupled to a surface of the inner cylindrical wall and outer wall.

5. The detector of claim 1, wherein the volume inside the hollow body is divided into separate chambers by optical septa, and wherein each photo-detecting unit is delimited by at least one of the optical septa, the inner cylindrical wall, or the outer wall, and wherein each photo-detecting unit comprises scintillation material and at least one photodetector.

6. The detector according to claim 5, further comprising an anode and a cathode, each operatively connected to at least one of the two photodetectors at opposite sides, adapted to allow drift of the ionized charges and further equipped with sensing elements configured to receive an ionization signal.

7. The detector of claim 1, further comprising a shielding material located beyond the outer wall in a radial direction.

8. The detector according to claim 7, wherein the shielding material comprises at least one of deionized and demineralized water with resistivity between 17 and 18 M.Math.cm, distilled water with resistivity between 17 and 18 M.Math.cm, or organic liquid fluids with content of .sup.238U and .sup.232Th reduced below one part per trillion.

9. The detector according to claim 7, wherein the shielding material is configured either as a passive shield or as an active anticoincidence veto detecting Cherenkov photons, scintillation photons, or both.

10. The detector according to claim 1, wherein the concentration of xenon in the scintillating material is between and including 1 ppm and 1 ppt.

11. An imaging unit comprising: a first hollow body comprising an inner circular cavity suitable to house a subject; a detector housed inside the first hollow body and with an inner wall concentric with the inner circular cavity; and a processing means operatively connected to a plurality of photo-detecting units to process electrical signals output from the photo-detecting units and determine the position of emission of a positron in the subject, wherein the detector comprises: a second hollow body having an inner cylindrical wall and an outer wall spaced apart from the inner cylindrical wall, the second hollow body including a scintillating material suitable to emit photons once hit by a 511 keV -ray; one or more pairs of photo-detecting units for detecting photons emitted by the scintillating material, each photo-detecting unit of a pair comprising two photo detectors and a first unit of the pair being placed at a surface of the inner cylindrical wall directly opposing a second unit of the pair; wherein the scintillating material has a scintillation decay time less than 10 ns, an atomic number greater than 10, and a photoelectron yield greater than 8,000 photons/MeV; and wherein the scintillating material comprises at least one of .sup.40Ar, argon depleted in .sup.39Ar, an argon-xenon mixture, or traces of xenon dissolved in argon.

12. A detector for a positron imaging unit, comprising: a hollow body having an inner cylindrical wall and an outer wall spaced apart from the inner cylindrical wall, the hollow body including: a scintillating material suitable to emit photons once hit by a 511 keV -ray; one or more pairs of photo-detecting units for detecting photons emitted by the scintillating material, where at least one photo-detecting unit of a pair comprises two photodetectors, and a first unit of the pair being placed at a surface of the inner cylindrical wall directly opposing a second unit of the pair; wherein the scintillating material has a scintillation decay time less than 10 ns, an atomic number greater than 10, and a photoelectron yield greater than 8,000 photons/MeV; and wherein the scintillating material comprises at least one of .sup.40Ar, argon depleted in .sup.39Ar, an argon-xenon mixture, or traces of xenon dissolved in argon.

13. An imaging unit comprising: a first hollow body comprising an inner circular cavity suitable to house a subject; a detector housed inside the first hollow body and with an inner wall concentric with the inner circular cavity; and a processing means operatively connected to a plurality of photo-detecting units to process electrical signals output from the photo-detecting units and determine the position of emission of a positron in the subject, wherein the detector comprises: a second hollow body having an inner cylindrical wall and an outer wall spaced apart from the inner cylindrical wall, the second hollow body including a scintillating material suitable to emit photons once hit by a 511 keV -ray; one or more pairs of photo-detecting units for detecting photons emitted by the scintillating material, where at least one photo-detecting unit of a pair comprises two photodetectors, and a first unit of the pair being placed at a surface of the inner cylindrical wall directly opposing a second unit of the pair; wherein the scintillating material has a scintillation decay time less than 10 ns, an atomic number greater than 10, and a photoelectron yield greater than 8,000 photons/MeV; and wherein the scintillating material comprises at least one of .sup.40Ar, argon depleted in .sup.39Ar, an argon-xenon mixture, or traces of xenon dissolved in argon.

Description

BRIEF DESCRIPTION OF DRAWINGS

(1) The invention will be described hereinafter with reference to non-limitative examples, which are provided for explanatory, non-limitative purposes in the accompanying drawings. These drawings illustrate different aspects and embodiments of this invention and, where appropriate, the structures, components, materials and/or similar elements are indicated in the different figures with similar reference numbers.

(2) FIG. 1 illustrates an imaging unit according to the present invention;

(3) FIG. 2 illustrates a transversal section of the imaging unit of FIG. 1 along the plane ;

(4) FIG. 3 illustrates a detector housed in the imaging unit of FIG. 1.

(5) FIG. 4 illustrates diagrams relating to the control of an hadron beam and to the interpretation of images taken with an imaging unit of FIG. 1 based on control of the hadron beam.

DETAILED DESCRIPTION OF THE INVENTION

(6) While the invention is susceptible to various modifications and alternative constructions, some of the illustrated embodiments are shown in the drawings and will be described below in detail.

(7) It must be understood, however, that there is no intention to limit the invention to the specific illustrated embodiments, but, on the contrary, the invention intends to cover all the modifications, alternative constructions and equivalents that fall within the scope of the invention as defined in the claims.

(8) The use of such as, etc., or indicates non-exclusive alternatives without limitations, unless otherwise indicated.

(9) The use of includes means includes, but is not limited to, unless otherwise indicated.

(10) FIG. 1 illustrates an imaging unit 1 comprising a hollow body 2 with an internal cavity 20 allowing a table 3 to slide inside out it in order to bring a subject under a detector 4 housed inside the hollow body 2.

(11) Detector 4 is adapted to detect emission of a positron emitted in the subject lying on the table 3 and to output a electrical signals to processing means, not shown in figures, adapted to determine the position of emission of a positron in the subject.

(12) Detector 4 comprises a hollow body 40 having an inner cylindrical wall 41 and an outer wall 42. In one embodiment, the inner cylindrical wall 41 has an inner diameter d of the order of 80 cm and outer diameter D bigger than 100 cm, preferably 120 cm or larger, sufficient for an optimal containment of the 511 keV -rays emitted by annihilation of positrons in the tracer injected in a subject lying on the table 3. In particular, the thickness of the cylinder (distance between walls 41 and 42) is suitable to host a layer of a scintillating material (see following description) with a thickness of four interaction lengths for 511 keV -rays or more.

(13) The axial length of detector 4 may cover any size from a very short (few cm) cylinder to 200 cm in case of a full-body scanner.

(14) The hollow body 40 includes a scintillating material 43 suitable to emit photons once hit by a particle (e.g., the 511 keV -rays emitted during annihilation of a positron) and photodetectors (e.g., SiPM or PMTs) suitable to detect photons emitted by the scintillating material. The scintillating material is selected to have very low scintillation decay time , preferably lower than 10 ns, a high atomic number Z, preferably greater than 10, and a high scintillation yield, preferably greater than 8,000 photons/MeV. Requirements may be met by organic scintillators, but best performance is typically guaranteed by noble elements kept in liquid state inside a cryostat, not shown in the figures, and kept by a cryocooler at a suitable and constant temperature. The cryostat contains the full volume of the scintillating material as well as the photodetectors.

(15) The scintillating material may be argon, and argon-xenon mixtures in any proportion, including pure xenon, with or without additional organic dopants, and/or wavelength shifters (e.g., TetraPhenylButatliene, TPB), and/or quantum splitting wavelength shifters. The addition of dopants or wavelength shifters, including the possible use of traces of xenon in argon, is preferable to convert all excited states responsible for the emission of scintillation light into fast decaying states, such as to obtain a very fast scintillator emitting photons within a few nanoseconds. In one embodiment, organic wavelength shifters are in solution in the noble element in the liquid state. In other embodiments, they could be placed on the walls of the detector facing the scintillating material.

(16) Argon commercially available is extracted from the atmosphere and has an activity of the radioactive .sup.39Ar of 1 Bq/kg. The tail end of spectrum of .sup.39Ar (Q=565 keV, t.sub.1/2=269 yr) overlaps the Region Of Interest (ROI) for the detection of the full energy deposition peak of the 511 keV -rays. In one embodiment, devised for clinical studies that may require the use of a very low radioactive dose, the scintillating material comprises argon depleted in .sup.39Ar, so as to increase the Signal-to-Noise Ratio (SNR) and, consequently the dose of radioactive tracer to be inoculated into the subject. Argon depleted in .sup.39Ar can be obtained by use of known techniques, e.g., by extracting argon from underground sources [3,4], where the .sup.39Ar is naturally reduced by the suppression of cosmic rays, or by active isotopic depletion, such as it may be achieved by the use of gas centrifuges or through cryogenic distillation.

(17) In one embodiment, detector 4 is surrounded by a passive shield of ultrapure material (e.g., deionized and demineralized or distilled water with resistivity between 17 and 18 M.Math.cm, or organic liquid fluids with content of .sup.238U and .sup.232Th reduced below one part per trillion), surrounding all sides of the positron detector with the exception of one or two of the ends of the cylinder, suitable to suppress signals induced by environmental radioactivity to increase the SNR. The shielding material can be configured either as a passive shield or as an active anticoincidence veto detecting Cherenkov and/or scintillation photons. In one embodiment, that can be used in isolation or in combination with the passive or active shield, the detector is deployed in an underground laboratory, suppressing the background from cosmic rays and increasing the SNR.

(18) In one embodiment, that can be used in isolation or in combination with the passive or active shield and/or the deployment in an underground laboratory, the imaging unit comprises means to force fresh airpreferably processed to reduce its Radon (.sup.222Rn, .sup.220Rn) concentrationinside the cavity 20 where the subjects reside. This solution further suppresses background noise and increases the SNR.

(19) In order to detect photons emitted by the scintillating material, detector 4 comprises a plurality of pairs of photo-detecting units 45, each pair of photo-detecting units 45 being placed at opposite ends of the inner cylindrical wall 41 along a radial direction. For sake of clarity, only some radial directions are indicated by dashed lines in FIG. 2.

(20) In the embodiment of FIGS. 2 and 3, the cylindrical volume of detector 3 is subdivided by optical septa in separate photo-detecting units 45 of footprint from a few cm.sup.2 to a fraction of a cm.sup.2. Photodetectors 44, in particular photomultipliers like SiPMs or PMTs, are placed at each radial extremity of the photo-detecting unit 45, while scintillating material 43 is comprised between the photodetectors thanks to the optical septa.

(21) Photo-detecting units 45 output electrical signals once the photodetectors detect 511 keV -rays emitted by the annihilation of positrons in the subject; the output electrical signals are then processed by the processing means which determine the position of emission of a positron in the subject and reconstruct an image of the subject.

(22) This solution allows a use of detector 4 both as scintillation-only detector or as a scintillation and ionization detector. The ionization detector could be added by providing, through a drift cage, a strong electric field capable of separating the electron-ions pairs resulting from the ionizing events in the scintillating material and drifting the electrons towards the photodetectors 44, where the electrons charge would be detected by specially devised sensors.

(23) Some considerations are reported here below to show the performances of the above described detector.

(24) With liquid argon doped to reduce the scintillation decay to 4 ns and use of high-performance SiPM at 87 K, a photoelectron yield in excess of 5,000 photoelectrons/MeV is practical and achievable [5,6]. This would result in the detection of 2,500 photelectrons for a fully contained 511 keV -ray, which in turn guarantees detection of arrival of the first photons on the surface of the photo detectors in 4 ns/2,5002 ps. Taking into account practical limitation on the time required for detection and processing in the photodetectors, a 25 resolution of t30 ps is possible with the use of SiPM at cryogenic temperature or special PMTs. This would allow obtaining a position resolution in the radial direction for each T event of c t/24 mm. At the same time, the dimensions of the voxels in the axial and transaxial directions would allow retaining the 4-5 mm resolution in those directions already achieved in commercial PET and TOF-PET units. Finally, combining the reconstructed position in three-dimensions of all events recorded, with the additional aid of advanced computer-aided algorithms, an increase in the resolution of the position of tumors to the length scale of 1 mm is expected.

(25) Both compton- and photo-electrons can be recorded, 5 processed and analyzed as signal. The photoelectron yield indicated above for the imaging unit would result in a 2-3% energy resolution at 511 keV, to be compared with the 10-15% of current commercial units.

(26) Considering that:

(27) 1) Monte Carlo simulations indicate that 16% of .sup.18F decays results in single triggers and a 310.sup.4 fraction of decays results in T events;

(28) 2) the typical fraction of 18F-FDG consumed by tumor metabolism is of 1%;

(29) 3) the uptake of sugar-hungry tumors such as carcinoma and breast cancer results in a 80:1 uptake relative to nearby healthy cells; and

(30) 4) detection of a tumor of size 1/10 cm.sup.3 can be established, at the spatial sensitivity of 1 mm introduced above, for a conservative statistics of 1,000 localized T events. A procedure with the goal of determining size and position of a tumor of 1/10 cm.sup.3 would require 310.sup.8 total .sup.18F decays, corresponding to an activity of 150 kBq or 4 Ci and to a singles trigger rate of 30 kBq.

(31) Usage of the 10 ns coincidence window typical of traditional PET and TOF-PET units would result in a very strong suppression of the dominant background sources, R events, whose rate RR would become completely negligible. S events would be strongly suppressed by the improvement in energy resolution.

(32) From the above description it is clear that several modification can be made by a person skilled in the art of positron imaging units and detectors for positron imaging units without departing from the scope of protection of the present invention as defined by the claims.

(33) As an example, the cylindrical volume of detector 4 can be a single, contiguous optical volume housing the scintillation material 43 and all the photodetectors 44 instead of being divided into a plurality of volumes by optical secta like in the above described example. Photodetecting units therefore can consists of single photodetectors immersed into a scintillating material.

(34) In one embodiment, the detectors is provided with only a pair of detectors (or voxels) placed along a radial direction at two opposite side to detect back-to-back -rays. In this embodiment, the positron imaging unit can be provided with means to rotate the detector 3 around the longitudinal axis.

(35) In one embodiment, the imaging unit 1 is operated in combination with a hadrotherapy machine. The imaging unit is used as an in-beam three dimensional direct positron imaging unit, capable of determining the differential dose delivered by the hadrons beam to cancer cells relative to that delivered to surrounding healthy tissues. To this end, prior to the start of the procedure subjects are inoculated special tracers targeting cancer cells. These tracers differ from the most commonly used [.sup.18F]FDG (Fluorodeoxyglucose, 2-Deoxy-2-[.sup.18F]fluoroglucose, C.sub.6H.sub.11.sup.18FO.sub.5), a carrier of .sup.18F, an unstable nuclide emitting positrons. The tracers for the in-beam three dimensional direct positron imaging are chosen among those containing nitrogen and are loaded in the stable isotope .sup.15N (natural isotopic abundance: 0.37%) in lieu of the 20 naturally predominant .sup.14N (natural isotopic abundance: 99.63%). For example, for a treatment of brain cancer, a suitable tracer containing nitrogen FDOPA (Fluorodopa, or 2-fluoro-5-hydroxy-L-tyrosine, C.sub.9H.sub.10FNO.sub.4), whose efficiency as a tracer was demonstrated to be superior to more common [8]. FDOPA would be loaded in the .sup.15N isotope, produced, for example, by cryogenic distillation, and, once administered to subjects in form of [.sup.15N]FDOPA (Fluorodopa, or 2-fluoro-5-hydroxy-L-tyrosine, C.sub.9.sup.15FO.sub.4) would preferentially concentrate in cancer cells. A possible alternative to [.sup.15N]FDOPA is [.sup.15N]NH.sub.3. Preferably, in order to accumulate a sufficient concentration of .sup.15N in the cancer cells of subjects, subjects will need to be inoculated sufficient quantities of tracers, possibly also in form of proteins and amino acids targeting the metabolism of cancer cells. The technique known as Stable Isotope LAbeling of Mammals tissues [9,10] results in the substitution of .sup.14N with .sup.15N with high percentages (>80%).

(36) After administration of tracers loaded in the stable isotope .sup.15N in the subject, the subjects' tissues are irradiated with a proton beam. Positron emitting nuclides are produced in the healthy tissues through reactions such as .sup.16O(p,pn).sup.15O, .sup.16O(p,) .sup.13N, .sup.13N(p,pn).sup.13N, .sup.12C(p,pn).sup.11C, .sup.13N(p,).sup.11C, .sup.16O(p,pn).sup.11C. Production of these nuclides is typically family abundant, resulting in a planar integral density of positron emitting nuclide of a few per mm for planar integral doses at the Bragg peak of a few mGy.Math.mm.sup.2 [11], which roughly translates in a positron emitting nuclide every 50-100 protons [11] and overall positron decay rates of several hundred kBq [12] at termination of a typical hadrotherapy session. In cancer cells, positron productions would be boosted by the additional channel .sup.15N(p,n).sup.15O, which has a cross section significantly larger than .sup.16O(p,pn).sup.15O.

(37) In order to discriminate the dose administered to cancer cell with respect to that administered to healthy tissues, the following procedure (described with reference to FIG. 4) is applied.

(38) The hadrons beam is cyclically switched on and off (see upper diagram of FIG. 4) with a period T, while the imaging unit 1 continuously takes images of subject's tissues to create three-dimensional maps of positron annihilation vertices (see lower diagram of FIG. 4).

(39) Since .sup.15O has a half-life of 2.0 min, much shorter than that of .sup.13N (10.0 min) and .sup.11C (20.4 min), the proton beam is switched on and off with a frequency comprised between 1 and 10 minutes and preferably of 4.0 minute while the positron annihilation events are continuously recorded with the imaging unit 1.

(40) While the hadron beam is switched on and after it has been switched on for a time , preferably equal or bigger than the half-life of .sup.15O, and therefore from time from to T with reference to FIG. 4, the imaging unit 1 takes images of positron annihilation vertices due to .sup.15O, .sup.13N and .sup.11C.

(41) While the hadron beam is switched off and after the hadron beam has been switched off for the same time T, and therefore from time T+ to time 2T in FIG. 4, having .sup.15O decayed much faster than .sup.13N and .sup.11C, the images taken after T have a strong and progressive reduction in the positron annihilation events due to .sup.15O and therefore provide a proxy of the positron annihilation vertices due to .sup.13N and .sup.11C.

(42) By subtraction of the two images, recorded in time intervals from to T and from T+ to 2T, we can reconstruct the maps of the positron annihilation vertices from .sup.15O and from .sup.13N and .sup.11C.

(43) The map of the positron annihilation vertices from .sup.15O, activated both from .sup.15N nuclides concentrated in cancer cells and from .sup.16O nuclides distributed in all tissues, is a proxy of the dose delivered both to cancer cells and all tissues. The map of positron annihilation vertices due to .sup.13N and .sup.11C, activated from .sup.16O, .sup.14N, and .sup.12C, is a proxy of the dose delivered to all tissues.

(44) From further manipulation of the two maps we can extract a proxy indicator of the differential of the dose provided to cancer cells and surrounding healthy tissues, correlated with their positions reconstructed in three dimensions.

REFERENCES

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