Patent Application
20210364660
The technical field of the invention is the dosimetry linked to measurements and checks on beams generated by medical linear accelerators used in radiotherapy and notably the beams in stereotactic radiotherapy.
A certain number of excessive irradiations of patients have occurred, caused by an overexposure to ionizing radiations in radiotherapy treatments. These irradiations have led to serious sequelae, even deaths. The result thereof is a need to have a better quality in the predicting of the doses administered to the patients.
Currently, the dose delivered to the patient is estimated using computation code modellings, in order to best estimate the dosimetry at the level of the tumors or of the organs that are most sensitive, the measured data, on which the computation codes are based, have to be better controlled, notably for the small beams. The control of the dose delivered involves comparisons between doses that are modelled and doses that are measured in real time, on a patient, or even between doses that modelled and doses that are measured experimentally, on phantoms representative of the body of a patient.
The problem arises notably in the field of stereotactic radiotherapy, this modality successively implementing convergent irradiating beams of small size, so as to selectively irradiate a target of small volume, typically of the order of a cm3. Generally, the irradiating beam has a diameter or a greater diagonal of less than a few cm, for example 3 cm. The irradiation source can be a radio-isotope, for example 60Co, or a particle accelerator, for example a linear accelerator (LINAC), that makes it possible to obtain an X-ray or high-energy electron beam, that can potentially reach a few tens of MeV, The irradiating beams converge toward the target to be treated either by rotation, or by the geometry and the positioning of multiple sources of 60Co.
The beam emitted by the irradiation source passes generally through a collimator, notably a multi-plate collimator, composed of a plurality of dense plates whose arrangement makes it possible to obtain a spatial distribution of the dose corresponding to the geometry of the target to be treated. Each plate of the collimator can be disposed in such a way that the set of plates delimits an aperture matched to the target. In some cases, these plates are displaced during the delivery of the dose, making it possible to adapt the fluence of the photons at the target level in order to match the dose distribution as closely as possible to the lesion to be treated. This technique, called intensity-modulated conformal radiotherapy (referred to by the acronym IMCR), makes it possible to protect the healthy tissues adjacent to the target, while concentrating the dose on the target. During a rotation about the target, the configuration of the collimator can also change, so that the projection of the aperture onto the target encompasses the latter throughout the rotation.
The small stereotactic irradiation fields can also be obtained with other types of collimators, for example collimators formed with circular cones.
The irradiation beams implemented in stereotactic radiotherapy are characterized by a small size and a strong dose gradient, in particular at the periphery of the beam. Moreover, because of the small size of the irradiated zone, the condition of electronic equilibrium, in the irradiated target, may not be observed. These particular features lead to an uncertainty in the modelling of the dose integrated during an irradiation.
In order to check the dose actually delivered, experimental measurements are frequently implemented, for quality assurance purposes. Currently, the use of passive dosimeters, of radiochromic film or thermoluminescence cube type, is considered a benchmark method. These dosimeters make it possible to obtain a quantitative two-dimensional distribution of the dosimetry, complemented by punctual information when thermoluminescent cubes are used. However, implementing them is complex, takes a long time and is relatively costly, which is difficult to square with daily use. Furthermore, these dosimeters do not deliver information in real time. In addition, they are not suited to stereotactic radiotherapy guided by MRI (magnetic resonance imaging). Indeed, it has been shown that the performance of these dosimeters can be degraded by the intense magnetic fields generated by MRI.
Various alternatives to the passive dosimeters have been studied. For example, the document US2012/0292517 describes a scintillation detector, comprising scintillating optical fibers arranged parallel to one another. This detector is intended to be used for the quality control associated with radiotherapy. It can notably comprise different layers, extending parallel to one another, the fibers of one and the same layer being oriented parallel to one another, according to an orientation. However, the use of a fiber detector presents a number of drawbacks. A first limitation is linked to the size of the fibers, whose diameter is 0.5 mm, which does not make it possible to obtain an adequate spatial resolution. Furthermore, it is tedious to arrange several tens, even hundreds, of fibers alongside one another, such that the fibers are parallel to one another. Another limitation is linked to the coupling of the fibers with a photodetector, the fibers being in direct contact with the photodetector. The result thereof is a complex design, and a device that is relatively bulky and probably costly.
The publication, by Goulet M. entitled “High resolution 2D device based on a few long scintillating fibers and tomographic reconstruction”, Med. Phys., 39 (8) August 2012, describes the use of a detector comprising optical fibers, of 1 mm diameter, extending parallel to one another, on a plane. The detector is rotationally movable. Upon an exposure of the detector to an irradiation beam the detector is successively turned according to different orientations. The measurements performed on each orientation are used in a tomography algorithm to obtain a spatial distribution of the irradiation beam. Such a method presents limitations linked to the use of the optical fibers. Furthermore, it requires a sequential rotation of the detector, the latter having to be accurate if the aim is to obtain a good quality tomographic reconstruction. The sequential acquisition, according to different orientations, is affected by possible temporal variations of the irradiation beam. The method is therefore relatively complex to implement, because of the presence of a means for rotating the detector. A method based on the use of optical fibers is also described in US20140217295.
Another detector, targeting the same type of application, is described in US2009/0236510. The device comprises fibers of which one spot end is scintillating, to generate a light signal representative of a dose. The scintillating end is linked to a non-scintillating fiber whose function is to guide the light signal to an image sensor, of CCD or CMOS type. Such a detector presents the same drawbacks as those cited concerning US2012/0292517, namely a complex setup, reflected by a high cost, and a certain bulk, because of the presence of fibers extending to the detector. Moreover, only the end of the fibers is scintillating, the scintillating volume being less than 2 mm3. Such a detector is suitable in the case of spot measurements, but is not suited to the performance of a measurement of the spatial distribution of the dose in an irradiation beam whose diagonal is of the order of 2 or 3 cm.
The inventors have developed a detector that is simple, inexpensive and easy to implement for experimentally measuring a dose delivered by an ionizing radiation beam extending along a diagonal of a few centimeters. The detector makes it possible to simply evaluate a two-dimensional spatial distribution of the irradiation beam.
A first subject of the invention is a multilayer scintillation detector, comprising at least three layers superposed on top of one another, and each extending parallel to a plane, called detection plane, the detector being such that:
each layer comprises a first material, called scintillation material, that can interact with an ionizing radiation and form, following the interaction, a scintillation, light in a scintillation spectral band;
The scintillation spectral band can be situated in the visible or in the near ultraviolet. It generally lies within the 200 nm-800 nm spectral range.
The number of layers is preferably between 3 and 20.
The plate, in which the channels are formed, can comprise a non-scintillating bottom part, such that, after the formation of the channels, the light guides rest on the bottom part. The latter keeps the light guides parallel to one another.
The channels can extend to all or part of a thickness of the plate, and preferably to 90% of the thickness of the plate, the thickness being defined at right angles to the detection plane.
The light guides of each layer are kept, by the plate, secured to one another,
According to one embodiment, each light guide of one and the same layer extends, according to the detection plane, to a face of the detector, called detection face, the detection face being disposed transversely to the detection plane, and preferably at right angles thereto, such that the scintillation light generated in the light guide is propagated to the detection face. The device can comprise several detection faces that are different from one another, each detection face comprising ends of light guides formed in one and the same layer. A detection face can comprise ends of light guides formed in different layers.
The detector can have, in the detection plane, a polygonal section.
The height of at least one light guide, at right angles to the detection plane, preferably lies between 100 μm and 1 mm. The width of a light guide, in the detection plane, at right angles to the axis of orientation according to which the light guide extends, preferably lies between 100 μm and 500 μm.
The second material can be air. The first material can be an organic scintillator.
At least one layer can be separated from another layer which is superposed on it by a thickness of a third material, of a third optical index, lower than the first optical index and/or opaque and/or reflecting.
According to one embodiment, at least one layer comprises a so-called auxiliary detector, disposed in a channel, called measurement channel, the auxiliary detector being able to induce an optical or electronic signal when it is exposed to the ionizing radiation. The auxiliary detector can be formed by a solid material, linked to an optical or electrical connection, the connection extending in the measurement channel. The auxiliary detector is preferably a point detector, the auxiliary detector having a detection volume less than 1 mm3, and preferably less than or equal to 0.5 mm3. It can be a scintillation detector, notably of gallium nitride (GaN), linked to an optical fiber, the latter forming the optical connection.
According to one embodiment, the detector comprises marks, formed on at least one layer, using a material forming a contrast agent in an examination by magnetic resonance imaging, such that the marks form reference points that are visible when the detector is examined by magnetic resonance imaging.
A second subject of the invention is a device for detecting an ionizing radiation, comprising:
The device can comprise at least one optical coupling system, such that each pixel is optically coupled to a light guide by the optical coupling system. The optical coupling system can comprise optical fibers or one or more lenses.
According to one embodiment, at least one layer of the multilayer scintillation detector comprises an auxiliary detector, disposed in a channel, called measurement channel, the auxiliary detector being able to induce an optical or electronic signal when it is exposed to the ionizing radiation, the detection device comprising a measurement unit, linked to the auxiliary detector, and configured to measure a level of irradiation detected by the auxiliary detector. The auxiliary detector can be a scintillator of GaN type, linked to an optical fiber, the optical fiber extending in the measurement channel.
A third subject of the invention is a method for reconstructing a two-dimensional spatial distribution of an irradiation beam emitted by an irradiation source, using the detection device according to the second subject of the invention, the method comprising the following steps:
According to one embodiment,
The steps a) to c) can be performed by arranging a multilayer scintillation detector at different distances from the irradiation source, so as to obtain, for each distance, a two-dimensional spatial distribution of the irradiation.
Other advantages and features will emerge more clearly from the following description of particular embodiments of the invention, given as nonlimiting examples, and represented in the figures listed below.
Unless explicitly stated otherwise, the term “a” should be interpreted to mean “at least one”. Moreover, the arrangement is in an orthogonal reference frame defined by the axes X, Y and Z, the Z axis corresponding to the vertical axis.
The energy of the ionizing radiation can be between 500 keV and 22 MeV when the irradiation source comprises a particle accelerator. When it is an isotopic source, the energy of the radiation corresponds primarily to the emission lines of the source. In the case of 60Co, the energies of emission of these lines are equal to 1173 keV and 1332 keV.
The patient undergoing the radiotherapy treatment is generally positioned on a table 14. In this example, a phantom 2 is represented, simulating the body of a patient. The phantom can be produced according to different materials exhibiting an attenuation comparable to the body of a patient (atomic number and density both very close to those of tissues), for example an organic material of PMMA (polymethylmethacrylate) type. Such a phantom is called “tissue equivalent”.
A detection device 1 is represented that makes it possible to estimate the dose generated by the irradiation beam Ω, and to estimate a two-dimensional spatial distribution thereof according to a plane. The detection device 1 comprises a multilayer scintillator 20 and at least one pixelated photodetector 30. The multilayer scintillator 20 extends essentially according to a plane, called detection plane P. Under the effect of the irradiation beam, the multilayer scintillator generates a scintillation light, the latter being guided to the photodetector 30. The scintillation light is generated in a scintillation spectral band, the latter depending on the scintillation material used. The scintillation spectral band is generally the visible or near ultraviolet range, therefore lying between 100 nm and 800 nm.
The photodetector 30 is situated at a distance from the multilayer scintillator 20. The detection device 1 also comprises a processor 3, for example a microprocessor, linked to a memory 4, comprising instructions for implementing the detection and reconstruction methods described hereinbelow. The processor 3 can be linked to a screen 5. The detection device 1, and more particularly the multilayer scintillator 20, form an important aspect of the invention, described more broadly in
In the example of
The spatial extent of the irradiation beam Ω, at right angles to the irradiation axis ZΩ, is determined by the multi-plate collimator 12. Such a collimator is represented in
As indicated in the prior art, the diameter or the greater diagonal of a section of the irradiation beam Ω, at right angles to the axis of propagation ZΩ, can be less than 5 cm, even than 3 cm. Generally, at the target tissue, the surface of the irradiation beam Ω, at right angles to the irradiation axis ZΩ, is less than 50 mm2.
The first layer 21 is formed by a support plate 21s, also referred to by the term “scintillating sheet”, produced according to a first material, called scintillation material. The support plate 21s extends according to the detection plane P. As previously described, the first material is a scintillation material, capable of generating a scintillation light when it is exposed to an irradiation. The scintillation material is for example an organic scintillator. Indeed, such a scintillator is called “tissue equivalent”; when it is exposed to an irradiation beam, it generates a scintillation light whose intensity is proportional to an instantaneous dose which would be delivered to a biological tissue. The organic scintillators are materials commonly used in the field of nuclear measurement. They can be available in different sizes, at a reasonable cost. Their response time is rapid and they exhibit a generally low remanence, making them particularly appropriate to repeated exposures to intense irradiation beams. Furthermore, they can easily be structured by simple microstructuring methods. In the present case, the material used is the reference BC408 from the manufacturer Saint Gobain Crystals. It can emit a scintillation light according to a scintillation spectral band centered on the 425 nm wavelength. Other organic scintillation materials known to a person skilled in the art can be used, and for example the reference BC412 (Saint Gobain Crystals), or even the references SCSF-78, SCSF-81, SCSF-3HF (Kuraray), or the reference EJ200 (Eljen Technologies). It is also possible to use a scintillating resin, for example the reference BC490 (Saint Gobain Crystals).
The scintillation material has a first optical refractive index n1. Generally, the first refractive index is, at a wavelength of 450 nm, greater than, or equal to 1.3, even 1.5. In the case of BC408, the refractive index is equal to 1.58 at this wavelength.
The support plate 21c has been structured, so as to form hollow channels 21c extending along the plate, according to a length l. Over all or part of the length l, the hollow channels 21, extend parallel to one another, being oriented according to a first axis of orientation Δ1. The first axis of orientation Δ1 is coplanar to the detection plane P according to which the first support plate 21 extends. The thickness of the support plate 21s, before the formation of the channels 21c, can lie between 100 μm and 5 mm. It is preferably less than 2 mm, and more preferably less than 1 mm. The channels are formed to all or part of a thickness of the support plate 21s, the thickness of the plate extending at right angles to the detection plane.
The structuring of the channels 21c makes it possible to delimit light guides 21g, each light guide extending between two adjacent channels, parallel thereto. It is important for the light guides to extent parallel to one another, according to the first axis of orientation Δ1, in at least a central part of the layer 21s, intended to be exposed to the irradiation beam Ω. In the example represented, the channels 21c and the light guides 21g extend parallel to one another, according to the first axis of orientation Δ1, over all their length.
After the structuring of the support plate 21s, each channel 21c is filled with a second material, different from the first material forming the support plate 21s. The second material has a refractive index n2 lower than that of the first material. In the examples described, the second material is air. The second material is, preferably, non-scintillating.
According to one embodiment, the support plate comprises a bottom part, corresponding to the thinner part, formed by a support material different from the scintillation material. The support material is preferably non-scintillating, and its refractive index is advantageously lower than the refractive index of the scintillation material. The support material is for example a plastic material. After the formation of the channels, the light guides 21g are held by the bottom part of the support plate 21s, the latter serving as non-scintillating support. The support material can, preferably, have a refractive index lower than the refractive index of the scintillation material.
Between two adjacent channels 21c there extends a light guide 21g. The height of the light guide, at right angles to the detection plane P, corresponds to the depth of the adjacent channels. The higher it is, the greater the detection sensitivity. Like the channels 12c, each light guide 21g extends according to the axis of orientation Δ1 of the first layer 21. The width of a light guide, at right angles to the axis of orientation Δ1, preferably lies between 100 μm and 500 μm, for example between 200 μm and 300 μm. This width conditions the spatial resolution of the detection, as is understood from the experimental tests described hereinbelow.
The first plate 21s, structured thus, forms a first layer 21 of the multilayer scintillator 20. Under the effect of an exposure to an irradiation beam Ω, a scintillation light is generated by the first layer 21, in particular within each light guide (or waveguide) 21g, the volume of the first layer 21 being essentially composed of the light guides 21g. Because of the difference in refractive index between each light guide 21g and the channels 21c that are adjacent to it, the scintillation light generated within each light guide 21g is propagated therein, according to the axis of orientation Δ1.
The structuring of the first layer can be performed by a method combining an etching method and lithography, for example photolithography, or by thermoforming of “hot embossing” type, or by molding or even by micro-machining. It makes it possible to simultaneously obtain a large number of waveguides, of small width, within one and the same layer, this number exceeding 100, even 1000. That makes it possible to perform measurements by benefiting from a high spatial resolution. When photolithography is implemented, it can be UV photolithography, for example at 375 nm, through a chrome on glass mask. A structured scintillator is thus obtained.
The structuring of the support plate makes it possible to simultaneously obtain waveguides of small width, separated from one another by a few tens of microns, and secured to one another. The waveguides are fixed to one another, being held by the thinner part of the support plate. A layer produced in this way is easily manipulable.
The second layer 22 and the third layer 23 have a structure similar to the first layer 21. They extend according to the same detection plane P. Thus, the second layer comprises light guides 22g, delimited by channels 22c, and extending, over at a least a part of their length, parallel to a second axis of orientation Δ2. The second axis of orientation is parallel to the detection plane P, but not parallel to the first axis of orientation Δ1. Thus, when the second, layer is exposed to an irradiation beam, a scintillation light is generated in each light guide 22g, and is propagated in each of them, according to the second axis of orientation Δ2.
Similarly, the third layer comprises light guides 23g, delimited by channels 23c, and extending, over at least a part of their length, parallel to a third axis of orientation Δ3. The third axis of orientation is parallel to the detection plane P, but not parallel to the first axis of orientation Δ1, or to the second axis of orientation Δ2.
Two layers can be directly superposed on one another. Alternatively, a third material, of a third refractive index n3, can extend between two adjacent layers. The third material can be identical to the second material, for example air. In this case, spacers are used to space apart two superposed layers. When the two adjacent layers are not in contact with one another, the distance separating them is preferably as small as possible, for example between 10 μm and 100 μm. The third refractive index n3 is preferably lower than the first refractive index n1, in the scintillation spectral band, notably when the third material is transparent in the scintillation spectral band. That allows for a better containment of the light in the light guides. Alternatively, the third material can be an opaque material, so as to optically isolate the two superposed layers that it separates. The third material can also be reflecting. When the support plate comprises a bottom part formed by a non-scintillation material, as previously described, the non-scintillation material can be the third material.
The thickness ε of the structured multilayer scintillator 20, at right angles to the detection plane P, is as small as possible, such that the layers can be considered to be exposed to one and the same irradiation beam, according to one and the same plane. The thickness ε of the multilayer scintillator must however allow each layer to have a sufficient thickness for the detection sensitivity to be acceptable. The thickness ε of the multilayer scintillator varies according to the number of layers, but it is preferable for it to be less than 2 cm or 1 cm.
Preferably, at each position of the pixelated photodetector 30, the relationship between the pixels collecting the scintillation light is bijective, such that a light guide is optically coupled to a pixel, or a group of pixels, the pixels that are optically coupled to one light guide being different from the pixels that are optically coupled to another light guide. According to a variant, several light guides are optically coupled to one and the same pixel.
It is possible to provide a coupling of the light guides of a layer, to the pixels of a photodetector, by optical fibers. However, it is, preferable for this coupling to be effected by an optical system 35, which is less complex to implement. That also makes it possible to keep each photodetector 30 at a distance from the multilayer scintillator 20. Alternatively, the light guides can be directly coupled to the pixels of the photodetector. In such a case, it is preferable for an optical coupling fluid, of coupling oil or gel type, to be disposed at the interface between the pixels and the light guides, so as to obtain an index matching between the light guides and the pixels.
The pixelated photodetector 30 can be an image sensor, of CCD or CMOS type. The different layers of the multilayer scintillator 20 being offset from one another according to the irradiation axis ZΩ, one and the same photodetector 30 can simultaneously address several layers, the pixels that are optically coupled to one layer being different from the pixels that are optically coupled to another layer. It is also possible to use linear sensors, comprising a strip of pixels extending along a row. Such sensors can be applied directly against the light guides emerging from a layer, even from each layer. An example of a linear sensor is the reference S11865-128 (Hamamatsu). The direct coupling of a sensor against the light guides of a layer enhances the compactness and can be produced from inexpensive and widely used linear sensors.
At least one face, and preferably each face, can comprise an opaque mask covering the face, apart from the light guides emerging from said face. The opaque mask can be obtained by the application of an opaque coating on the face. It can for example be an opaque paint or an absorbent sheet, applied to the face. That prevents a scintillation light, not guided by a light guide, from emanating from a face of the multilayer scintillator, by emerging notably from a thinner part of a plate. The addition of the opaque mask on one or more faces can affect all the embodiments. The opaque mask can be reflecting.
The central zone ZC of the scintillator 20 encompasses a projection of the aperture 13, defined by the collimator 12, according to the axis ZΩ of the irradiation beam Ω, on the detection plane P according to which each layer extends. The projection of the irradiation beam in the detection plane P is designated by the term “irradiation field”.
Outside of the central zone ZC, the waveguides of one and the same layer are directed toward one and the same detection face F1, the latter being common to several layers, and in this particular case to all the layers. The fact that the light guides of several layers emerge from one and the same detection face makes it possible to collect the scintillation light generated in each light guide with one and the same pixelated photodetector 30, coupled to the optical system 35. In
Regardless of the embodiment, when a light guide extends between an end coupled to a photodetector, and an end that is not coupled to a photodetector, the latter can be coated with a reflecting material, so as to return the scintillation light to the end of the light guide that is coupled to the photodetector. That makes it possible to increase the quantity of light collected by the photodetector which enhances the measurement sensitivity.
Other scintillation materials can be implemented to form the auxiliary detector 28. Preference will be given to detectors that are compact, of weak remanence and compatible with strong irradiation levels, and insensitive to the incidence of the irradiation beam. Other scintillation materials capable of forming the auxiliary detector that can be cited include, non-exhaustively, BGO (bismuth germanate), CsI(TI) (thallium-doped cesium iodide), LSO (lutetium oxyorthosilicate), LYSO (scintillating crystal based on cerium-doped lutetium), GSO (gadolinium orthosilicate), or LaBr3 (lanthanum bromide).
Several auxiliary detectors can thus be disposed in one and the same layer, even in different layers. Preferably, at least one auxiliary detector is disposed at the isocenter of the irradiation beam Ω. It is recalled that, in the case of stereotactic radiotherapy, the isocenter corresponds to the intersection of the successive irradiation axes ZΩ during the rotation of the irradiation source.
The auxiliary detector 28 allows for an accurate estimation of a dose at a point. This information, accurate but isolated, can advantageously be combined with the estimation of the spatial distribution of the irradiation beam, in the detection plane, described hereinbelow. Spatial information is then combined with a spot quantitative measurement.
Whatever the embodiment, the multilayer scintillator can comprise marks forming reference points, visible by MRI. These marks can be symbols of dot, cross or line type, produced in a material forming an agent of contrast in MRI, for example gadolinium. It will thus be possible to delimit an outline of the multilayer scintillator or identify noteworthy points, for example a center of the scintillator in the detection plane P. It is specified that the multilayer scintillator is preferably amagnetic, which makes it compatible with use in the strong magnetic fields produced in examination by MRI.
Experimental trials were carried out by implementing a single-layer scintillator of square section, comprising a single layer, similar to the first layer 21 of the scintillator 20 described in association with
In order to check the capacity of the light guides to propagate the scintillation light, the multilayer scintillator was first of all exposed to a UV irradiation (375 nm) at right angles to the plane XY. The irradiation beam forms, in the detection plane, a rectangle of 8 mm (according to the axis X) by 50 mm (according to the axis Y). An optical system 35 and a photodetector 30 of CMOS sensor type (Andor Zyla 5.5 CMOS camera optically coupled to a Navitar 7000 macro zoom) was disposed opposite the face of the scintillator, extending according to a plane XZ. The face of the scintillator, an image of which is formed by the CMOS sensor, is designated “detection face”.
During a second experimental trial, the UV irradiation was formed by two beams of 200 μm width (according to the axis X) and of respective lengths (according to the axis Y) equal to 47.5 mm and 40 mm.
During another series of tests, the single-layer scintillator was exposed to an irradiation beam Ω of X photons from an accelerator raised to the 6 MV potential, the dose rate rising to 14 Gy/minute. The length of the beam, according to the axis Y, rose to 10 cm. The width of the beam, according to the axis X, was successively set at 3 cm, 2 cm, 1 cm, 0.5 cm and 0.1 cm.
During another series of tests, the use of a scintillator of trapezoidal section, similar to the example described in association with
Step 100: arrangement of the multilayer scintillator 20 in an irradiation beam Ω, the scintillator extending in a detection plane P. In this example, the detection plane P is orthogonal to the axis of irradiation PΩ, but this condition is not essential.
Step 110: parameterizing of the tomography. This involves performing a modelling so as to obtain a transfer matrix M. The detection plane is discretized into a number of individual meshes and a transfer matrix is calculated, in which each term M(i,j) corresponds to a contribution to the light intensity measured at the output of a light guide i of the scintillator when a mesh j is exposed to a given irradiation level. Establishing such a transfer matrix is a conventional step in tomography. The dimension of the matrix is I×J, in which denotes the number of waveguides and J denotes the number of meshes.
Step 120: acquisition, by one or more pixelated photodetectors, of images representative of the quantity of light emanating from the light guides respectively formed in each detection layer. A projection of the irradiation beam is then obtained according to the orientation respectively associated with each layer. When there is a sufficient number of pixels, for example by using an imager, the acquisition of the images can be simultaneous, so as to obtain information as to the extent and the intensity of the irradiation beam. The quantity of each signal can form a measurement vector V, of which each term V(i) is representative of a signal quantity collected by each vector. The dimension I of the measurement vector corresponds to the numbers of waveguides taken into account.
Step 130: inversion. This involves determining an irradiation vector W, each term W(j) of which corresponds to an irradiation quantity detected in a mesh j. The dimension of the irradiation vector corresponds to the numbers J of meshes taken into account. The measurement vector V, the transfer matrix M and the irradiation vector are linked by the equation: V=M×W. The inversion allows for an estimation of the vector W that best satisfies this relationship. It is performed according to different methods known to a person skilled in the art.
Step 140: obtaining of a two-dimensional spatial distribution of the irradiation beam Ω, from the irradiation vector W estimated in the preceding step.
Another example of a tomography algorithm is also described in the publication by Goulet M., “High resolution 2D measurement device based on a few long scintillating fibers and tomographic reconstruction”, cited in the prior art.
When the multilayer scintillator comprises an auxiliary detector, the latter can be used to perform a realignment of the two-dimensional spatial distribution obtained from an exposure value measured by the auxiliary detector. Spatial information is then combined with one or more spot quantitative measurements.
The multilayer scintillator 20 described above will be able to be used to predict the dosimetry prior to radiotherapy interventions, in particular in stereotactic radiotherapy. It will for example be possible to arrange several multilayer scintillators 20, parallel to one another, in a phantom 2, as represented in
It is also possible to envisage arranging different multilayer scintillators respectively in different planes, so, as to obtain a two-dimensional spatial distribution of the irradiation beam respectively in the different planes.
Whatever the disposition of the multilayer scintillator or scintillators, the phantom 2 can comprise a point detector, for example a GaN scintillator of small volume, typically less than 1 mm3, at the isocenter.
| Number | Date | Country | Kind |
|---|---|---|---|
| 18 52480 | Mar 2018 | FR | national |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/FR2019/050637 | 3/20/2019 | WO | 00 |
