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(12) United States PatentEbstein

(10) Patent No.:(45) Date of Patent:

US 9,750,958 B2Sep.5,2017

(54) DOSIMETRIC SCINTILLATING SCREENDETECTOR FOR CHARGED PARTICLERADIOTHERAPY QUALITY ASSURANCE

(56) References Cited

U.S. PATENT DOCUMENTS

(71) Applicant: Steven M. Ebstein, Newton, MA (US)

(72) Inventor: Steven M. Ebstein, Newton, MA (US)

( *) Notice: Subject to any disclaimer, the term of thispatent is extended or adjusted under 35U.S.c. 154(b) by 286 days.

4,081,673 A * 3/1978 Swindell . A61B 6/025250/237 R

5,493,121 A * 2/1996 Fitzpatrick ................ GOlT 1/20250/361 R

6,285,740 Bl * 9/2001 Seely. H05G 1/30250/367

7,412,024 Bl * 8/2008 Yun. A61B 6/484250/367

(Continued)

(21) Appl. No.: 14/528,079 OTHER PUBLICATIONS

(60) Provisional application No. 611602,301, filed on Feb.23, 2012, provisional application No. 611897,532,filed on Oct. 30, 2013.

Related U.S. Application Data

(63) Continuation-in-part of application No. 131775,325,filed on Feb. 25, 2013.

(51) Int. Cl.GOlJ 1100 (2006.01)A61N SilO (2006.01)

(52) U.S. Cl.CPC .... A61N 511075 (2013.01); A61N 200511076

(2013.01)

(58) Field of Classification SearchCPC .... G01 T 1117; GOlT 7/00; GOlT 1102; GOlT

11161; G01T 1129See application file for complete search history.

US 2015/0071408 Al (Continued)

Primary Examiner - David PortaAssistant Examiner - Mamadou Faye(74) Attorney, Agent, or Firm - Pandiscio & Pandiscio

(57) ABSTRACTAn apparatus and method are provided for performingQuality Assurance of complex beams of penetrating radia­tion inside a patient. A detector with a transverse scintillatingscreen images the radiation inside a tissue phantom withhigh spatial resolution. The scintillator is comprised of amixture of two or more scintillators emitting different spec­tra of light and having different characteristic responses as afunction of the beam LET value. The optics relaying thescintillation output have variable transmission with wave­length, further shaping the spectrum of light transmitted tothe imaging sensor which also has spectrally varying sen­sitivity. Parameters of the scintillator construction, theoptics, and the imaging sensor are chosen so the output ofthe composite detector is proportional to a characteristic ofthe input beam, for example the dose deposited as a functionof depth inside the tissue phantom.

19 Claims, 11 Drawing Sheets

Kandarakis et al. "Experimental and theoretical assessment of theperformance of Gdl02S:Tb and Lal02S:Tb phosphors andGd202S:Tb-La202S:Tb mixtures for X-ray imaging" Euro. Radiol.(2001) p. 1083-1091.*

Mar. 12, 2015

Prior Publication Data

Oct. 30, 2014(22) Filed:

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US 9,750,958 B2Page 2

References Cited

u.s. PATENT DOCUMENTS

7,515,681 B2 4/2009 Ebstein8,080,801 B2 * 12/2011 Safai GOlT 11202

250/3682011/0248188 Al 10/2011 Brusasco et al.2012/0205530 Al 8/2012 Beaulieu et al.

OTHER PUBLICATIONS

Boon, S. N., Dosimetry and Quality Control of Scanning ProtonBeams, 1998.Koba, Y. et aI., Scintillation Efficiency ofInorganic Scintillators forIntermediate-Energy Charged Particles, Progress in Nuclear Sci­ence and Technology; 2011, vol. 1, p. 218-221.Murray, R. B. et aI., Scintillation Response of Activated InorganicCrystals to Various Charged Particles, Physical Review, 1961, vol.122, No.3, p. 815-826.Safai, S. et aI., Development ofan inorganic scintillating mixture forproton beam verification dosimetry, Physics in Medicine and Biol­ogy, 2004, vol. 49, p. 4637-4655.Schippers, J. M. et aI., Applications in radiation therapy of ascintillating screen viewed by a CCD camera, Nuclear Instrumentsand Methods in Physics Research, 2002, A 477, p. 480-485.

* cited by examiner

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beam is broadened in size or angle by scattering the pencilbeam. With particle beams, the shaping can also be effectedalong the beam direction by varying the energy distributionof the charged particles, as the penetration distance and dose

5 distribution along the particles' tracks depends on theirenergy. For a monoenergetic ion beam in a homogeneousmedium, the absorbed dose versus depth describes a curveknown as a pristine Bragg peak as seen at 12 in FIG. 1.

From entrance into the body until close to the penetration10 depth, the delivered dose, which is proportional to the

average linear energy transfer (LET) given in keV/flm due toionization by the beam, rises slowly along the plateau of thecurve. Near the final depth, the LET increases rapidly to a

15 peak, then falls to zero rapidly at the end of the track of theensemble of particles in the beam, thereby increasing theradiation dose in a narrow volume. The LET value istypically a characteristic of a beam as a function of depth. Itis related to the energy loss per unit length dE/dx per unit

20 density, typically given in MeV cm2/g, of each chargedparticle. dE/dx is a function of the particle energy.

In this specification, the term beam generally refers to apenetrating radiation beam. However, it may either refer toa radiation beam produced by an accelerator, a narrow pencil

25 beam at the output of the beam delivery system, a set ofpencil beams delivered within a short time, i.e., a compositebeam, or a broad beam restricted by a metal aperture orcollimator. A treatment field generally refers to a compositeor broad beam with a set of particle energies and a defined

applications are 30 spatial distribution delivered within a short time. If either theenergies or spatial distribution or both are changed, thatgenerally refers to a new treatment field.

By irradiating with beams having different energies, or bymodulating the beam energy as it is applied, the dose

35 distribution can be shaped to deposit maximal dose in anextended volume containing the tumor while deliveringreduced dose to the surrounding healthy tissue. An example,shown at 14 in FIG. 1, uses a beam with a set of differentenergies and fluences arranged to produce a depth-dose

40 curve with a flat maximum, denoted as a spread-out Braggpeak or SOBP. The energy of particles in the beam can bereduced by inserting absorbing elements, typically plastic,upstream ofthe patient. The energy reduction depends on thethickness of the absorber, so a set or range of energies can

45 be produced with a spatially varying thickness to theabsorber or by rotating a wheel with multiple thickness stepsand temporally modulating the beam energy.

FIG. 1 highlights one of the differences between radio­therapy using x-rays or charged particles. For an equivalent

50 dose at the depth in the middle of the SOBP, the x-ray beamdeposits more dose in front of that depth and after the tail ofthe Bragg peak, represented by the area between the x-raydose curve 10 and the charged particle SOBP dose curve 14.This ability to limit the dose to healthy tissue outside of a

55 tumor is a major motivation for delivering radiotherapy withcharged particles.

Charged particle beams can also be scarmed across thepatient if the beam is small enough and scanning magnetsare provided to steer the beam in angles away from the

60 central axis of the system. This technique is generallytermed pencil beam scarming or PBS. PBS treatments can bedelivered without metal apertures or collimators if thescauning magnets can sufficiently define the transverseshape of the composite beam. Typically, more than one

65 radiation field will be applied from a single direction withdifferent energies or energy spectra used for the differentfields, thereby shaping the delivered dose at different depths.

FIELD OF THE INVENTION

BACKGROUND OF THE INVENTION

FEDERALLY SPONSORED RESEARCH

REFERENCE TO PENDING PRIOR PATENTAPPLICATIONS

1DOSIMETRIC SCINTILLATING SCREENDETECTOR FOR CHARGED PARTICLE

RADIOTHERAPY QUALITY ASSURANCE

The goal of radiotherapy is to irradiate unhealthy, usuallycancerous, tissue or tumors with the aim ofkilling that tissuewhile sparing, to the extent possible, the surrounding healthytissue. This requires that the patient be irradiated with one ormore radiation beams that are carefully positioned andshaped in order to precisely deliver the intended quantity ofradiation to a specific target volume. In dimensions trans­verse to the nominal direction of the beam, this shaping isoften performed with metal apertures or multi-leaf collima­tors that define the profile of the beam. The beams are oftendirected from different angles at the patient, thereby increas­ing the radiation dose in the intersecting regions. A treatmentplan is generated for each irradiation session ofa patient thatdescribes the various beam angles, shapes, energies, andquantity of radiation delivered along with the associatedmaps showing the radiation dose intended for the patient.

With high-energy X-ray beams, the technique of spatiallyshaping the radiation dose with multi-leaf collimators andmultiple beams directed from multiple angles is known asIntensity Modulated Radiation Therapy or IMRT.

Charged particles, typically protons but also heavier ions,are also used for delivering spatially complex radiotherapytreatments. The particle accelerators typically produce anarrow, or pencil, beam that can be focused, collimated, ordeflected with magnetic optics, although sometimes the

This invention was made with U.S. Government supportunder Grant 5R44CAl0361O-03 awarded by the NationalCancer Institute. The U.S. Government has certain rights inthe invention.

The three (3) above-identified patenthereby incorporated herein by reference.

This patent application:(i) is a continuation-in-part of pending prior U.S. patent

application Ser. No. 131775,325, filed Feb. 25, 2013 bySteven M. Ebstein for DOSIMETRIC SCINTILLATINGSCREEN DETECTOR FOR CHARGED PARTICLERADIOTHERAPY QA, which patent application claimsbenefit of prior U.S. Provisional Patent Application Ser. No.611602,301, filed Feb. 23, 2012 by Steven M. Ebstein forDOSIMETRIC SCINTILLATING SCREEN DETECTORFOR CHARGED PARTICLE RADIOTHERAPY QA; and

(ii) claims benefit of prior U.S. Provisional Patent Appli­cation Ser. No. 611897,532, filed Oct. 30, 2013 by Steven M.Ebstein for DOSIMETRIC SCINTILLATING SCREENDETECTOR FOR CHARGED PARTICLE RADIO­THERAPY QUALITY ASSURANCE.

The present invention pertains to a device and a methodfor monitoring the spatial characteristics of a beam ofpenetrating radiation and, more particularly, to quality assur­ance in particle therapy, namely, ensuring that the dosedelivered matches the intended treatment plan.

US 9,750,958 B23 4

validate the treatment plan. In practice, current detectors fallshort in one or more respects. Some of the detectorsemployed, like film and detectors using thin scintillatingscreens or sheets of crystal, have spatial resolution thatexceeds the requirement for beam positioning and shaping.However, their detection signal is not linear with the fluenceor LET of the radiation beam, which is required in order topredict the corresponding dose to the patient.

The light output of some scintillators will saturate when10 a clinically useful fluence is incident. Inorganic scintillators

or phosphors such as PII (ZnS:Ag), P20 (Zn,CdS:Ag), P43(Gd20 2S:Tb), P46 (Y3AIs012:Ce), and P47 (Y2SiOs:Ce)have responses that are typically linear with fluence. How­ever, scintillator light output is generally not linear with the

15 LET of the incident particles. A plot of the light outputversus depth has the same general shape as the Bragg peak.However, the ratio of the peak value to the value at a pointon the plateau differs from the ratio ofLET values calculatedusing accurate physical models or that is measured by a

20 standard detector such as an ion chamber. When the ratio oflight output is less than the ratio of LET or absorbed dose,this is termed quenching. The physical explanation is that anabundance of excited atoms provide alternate pathways forexcited species to relax by means other than photon emis-

25 slon.Scintillating screens are used, in conjunction with radia­

tion therapy, for absorbed dose measurement (as described,for example, by J. M. Schippers, S. N. Boon and P. vanLuijk, "Applications in Radiation therapy of a scintillating

30 screen viewed by a CCD camera," Nucl Instr. and Meth. A477, pp. 480-85 (2002), and S. N. Boon, thesis, "Dosimetryand quality control of scanning proton beams" (1998),available online at http://www.ub.rug.nl/eldoc/dis/science/s.n.boon!, and references therein, all of which are incorpo-

35 rated herein by reference). They are also useful for moni­toring beam delivery in real time (S. M. Ebstein, HighResolution Proton Beam Monitor, U.S. Pat. No. 7,515,681).However, as those references show, knowledge of the par­ticle beam energy spectrum and careful calibration of the

40 detector response is required to make accurate dosimetricmeasurements.

Other detectors such as ion chambers are quite linear intheir measurement of the fluence and LET, so they are quiteaccurate in predicting absorbed dose. However, due to their

45 size and complexity, they cannot be made into arrays thatcan measure the incident radiation distribution with therequired resolution.

Due to the limitations of these detectors, a complexsystem is required to perfonn and evaluate the QA measure-

50 ments. Typically, a series of representative measurementsare made with detectors that have high spatial resolution butpoor dosimetric accuracy or detectors with limited spatialresolution but good dosimetric accuracy. The response ofthedetectors to the delivered radiation beams is calculated from

55 a model of the beam delivery and detection system andcompared to the actual measurements. This approach isdescribed in C. Brusasco and B. Marchand, Device AndMethod For Particle Therapy Monitoring And Verification,U.S. Patent Application Pub. No. US 201110248188 AI.

However, this approach cannot give complete confidencethat the delivered beam matches the plan due to reduceddimensionality of the measurements. In addition, the workrequired to construct the models and perfonn the calcula­tions of the expected measurements adds to the cost and

65 complexity of performing QA.There exists a need for improved detectors, especially for

charged particle therapy QA. The beam delivery systems are

Several different fonns of PBS are used to deliver radio­therapy. A technique known as Intensity Modulated ProtonTherapy or IMPT allows the beam intensity to vary as afunction of lateral position by modulating the beam currentor modulating the scauning rate. In some scenarios, thebeam position is continually scanned and in others it istranslated to a series of fixed positions where the beam isturned on for a period of time, a technique called spotscanning. For this approach, the magnetic scanning of thebeam defines the shape of the irradiated volume transverseto the beam direction. An approach, which is tenned eitherwobbling or uniform scanning, has the beam scan a givenarea, typically a regular region like a circle or rectangle, anddeliver a uniform dose over that area. A fixed aperture ormulti-leaf collimator is positioned between the beam sourceand the patient to restrict the transverse distribution ofradiation.

In all of these various fonns of radiotherapy, the standardfor positioning and shaping the beam relative to the patientis typically less than one millimeter across a field of regardthat can range up to 40 em in diameter. This requires that theradiation pattern be delivered with precise shaping andpositioning of the beam(s) over those scales, as well as inprecise amounts so the dose absorbed by the patient matchesthe prescription of the treatment plan. This requires carefulcalibration and measurement of the delivery system and thetreatment plan, i.e., the beam(s) which will be delivered tothe patient, a process referred to as quality assurance or QA.QA measurements are routinely perfonned that reassure thepractitioners that the correct radiation dose will be deliveredin the correct amount over the correct spatial distribution tothe patient. These measurements are performed with avariety of different radiation detectors.

With the complex series ofradiation fields that are used inradiotherapy, especially in charged particle therapy, it iscritical to make measurements which predict the absorbeddose inside of a patient, not merely in a plane just outside ofthe patient. One way this is accomplished is by placingabsorbing material, a tissue phantom, in between the beamsource and the detector so the radiation incident on thedetector corresponds to the radiation incident on the corre­sponding location(s) inside of the patient. The absorbingmaterial is typically water (since the body is mostly made upof water), sheets of plastic with water-equivalent absor­bance, or more complex phantoms with multiple materials.

There exist several different detectors used for perfonningQA of these radiotherapy beams. Film has been used sincethe beginnings ofradiotherapy to view the beam as deliveredto the patient. Various other technologies have been used,including scintillating screen based portal imagers, flat paneldevices consisting of a sheet of scintillating crystals atop anarray of semiconductor detectors, ionization chambers assingle devices, stacked along the beam direction, or arrayedtransverse to the beam direction, multi-wire proportionalchambers, gaseous electron multiplier (GEM) detectors, andother detectors, etc. Typically, these detectors are one part ofthe overall system and procedure for perfonning QA.

Requirements for these detectors include stability, robust­ness, linearity, accuracy, and precision, i.e., signal to noiseratio (SNR), spatial resolution, and dosimetric accuracy. The 60

detector output should be translatable into a measure of theeffective dose delivered to the patient. In addition, market­ability factors such as the cost and usability of the detectoroutput are important for the usefulness of a detector for QAapplications.

An ideal QA detector would measure the delivered doseinside a tissue phantom with sufficient spatial resolution to

US 9,750,958 B25 6

is proportional to a characteristic of the radiation beamincident on the scintillator at each measurement positionaccessible with the tissue phantom.

In another preferred form of the present invention, thereis provided a method for determining the spatial distributionand intensity of penetrating radiation beams characterizedgenerally by a propagation direction inside a body, themethod comprising:

providing a tissue phantom between the radiation source10 and a radiation detector;

providing a scintillating screen behind the tissue phantomfor emitting light in response to the radiation comprising amixture of at least two scintillators wherein each scintillator

15 has a different characteristic response and a different spectraloutput;

providing a means of optical communication of the scin­tillator output to at least one imaging sensor wherein themeans of optical communication has a nonuniform spectral

20 transmission; andproviding at least one imaging sensor in optical commu­

nication with the scintillating screen for providing a highresolution imaging sensor output indicative of the spatialdistribution and intensity of the radiation beam wherein the

25 imaging sensor has a nonnniform spectral sensitivity;wherein the composition of the scintillator, the means of

optical communication, and imaging sensor are selected soas to comprise a system wherein the imaging sensor outputis proportional to a characteristic of the radiation beam

30 incident on the scintillator at each measurement positionaccessible with the tissue phantom.

In another preferred form of the present invention, thereis provided a method for determining the spatial distribution

35 and intensity of penetrating radiation beams characterizedgenerally by a propagation direction inside a body, themethod comprising:

providing a tissue phantom between the radiation sourceand a radiation detector;

providing a scintillating screen behind the tissue phantomfor emitting light in response to the radiation comprising amixture of at least two scintillators wherein each scintillatorhas a different characteristic response and a different spectraloutput;

providing a means of optical communication of the scin­tillator output to an imaging sensor wherein the means ofoptical communication has a nonuniform spectral transmis­sion; and

providing an imaging sensor in optical communication50 with the scintillating screen for providing a high resolution

imaging sensor output indicative of the spatial distributionand intensity of the radiation beam wherein the imagingsensor has a nonnniform spectral sensitivity;

wherein a first composition of the scintillator, a first55 means of optical communication, and a first imaging sensor

are selected so as to comprise a first system wherein theimaging sensor output is proportional to a characteristic ofthe radiation beam incident on the scintillator at eachmeasurement position accessible with the tissue phantom;

60 andperforming characteristic measurements with a penetrat­

ing radiation beam and the first system;wherein one or more of (a) the composition of the

scintillator, (b) the means of optical commnnication, or (c)65 the imaging sensor is adjusted so as to comprise a second

system wherein the imaging sensor output is proportional toa characteristic of the radiation beam incident on the scin-

SUMMARY OF THE INVENTION

complex instruments and are required to produce a series ofbeams with complex shapes and different energy spectra.Current detectors and systems cannot directly measure thedelivered dose as a function ofdepth with sufficient accuracyand spatial resolution to quickly and easily perform the QAtask. A detector which combines the high spatial resolutionof either film or a scintillating screen detector and thedosimetric accuracy of an ion chamber would be very usefulin performing QA of complex radiotherapy treatment plans,especially for IMPT.

In accordance with preferred embodiments of the presentinvention, an apparatus and method are provided for per­forming QA of complex beams of penetrating radiation thatproduce a spatially varying dose distribution inside of apatient. The apparatus provides for inserting tissue phantommaterial in between the beam source and a detector. Thedetector incorporates a scintillator disposed transverse to thenominal direction of the beam and one or more high reso­lution imaging sensors in optical commnnication with thescintillator for generating an image of the beam distributionacross the scintillator. The scintillator is comprised of amixture of two or more scintillators emitting different spec­tra oflight and having different characteristic responses as afnnction of the beam LET value. The means of opticalcommnnication of the scintillation output, i.e., the optics,have variable transmission with wavelength, further shapingthe spectrum of light transmitted to the imaging sensor. Theimaging sensor has sensitivity that varies across the spec­trum of light input. Parameters of the scintillator construc­tion, the optics, and the number and types of imaging sensorsare chosen so the output of the composite detector isproportional to a characteristic of the input beam. In oneembodiment, the characteristic is the dose deposited as afnnction of depth inside the tissue phantom. In anotherembodiment, the scintillator is disposed in a thin, planarscreen that is oriented perpendicular to the central beamaxis. In another embodiment, the scintillator output is com- 40

municated to multiple sensors whose outputs are combinednumerically to produce an output proportional to the desiredcharacteristic.

In one preferred form of the present invention, there isprovided an apparatus for determining the spatial distribu- 45

tion and intensity of penetrating radiation beams character­ized generally by a propagation direction inside a body, theapparatus comprising:

a tissue phantom disposed between the radiation sourceand a radiation detector;

a scintillating screen disposed behind the tissue phantomfor emitting light in response to the radiation comprising amixture of at least two scintillators wherein each scintillatorhas a different characteristic response and a different spectraloutput;

a means of optical commnnication of the scintillatoroutput to at least one imaging sensor wherein the means ofoptical communication has a nonnniform spectral transmis­sion; and

at least one imaging sensor in optical commnnication withthe scintillating screen for providing a high resolution imag­ing sensor output indicative of the spatial distribution andintensity of the radiation beam wherein the imaging sensorhas a nonnniform spectral sensitivity;

wherein the composition of the scintillator, the means ofoptical communication, and imaging sensors are selected soas to comprise a system wherein the imaging sensor output

7US 9,750,958 B2

8tillator at each measurement position accessible with thetissue phantom with greater accuracy than with the firstsystem.

BRIEF DESCRIPTION OF THE DRAWINGS

The foregoing features of the invention will be morereadily understood by reference to the following detaileddescription, taken with reference to the accompanying draw­ings, in which:

FIG. 1 depicts the relative radiation dose from a photonbeam, a pristine proton beam (Bragg peak), and a spread outBragg peak (SOBP) using protons composed of multiplepristine proton Bragg peaks of different energies. The areabetween the curves is the excess dose from a photon beam.

FIG. 2 shows a schematic depiction of a system fordelivering a spatially complex radiation beam.

FIG. 3 shows a schematic depiction of the QA apparatusincluding tissue phantom material, a scintillating screen, andoptics relaying the image to an imaging sensor.

FIG. 4 shows the luminous efficiency of a scintillatorversus dE/dx.

FIG. 5 shows the spectral content of scintillator output forvarious scintillators.

FIG. 6 shows exemplary spectral transmission of com­pound lenses.

FIG. 7 shows exemplary spectral transmission of coloredglass filters.

FIG. 8 shows exemplary spectral response of an imagingsensor.

FIG. 9 shows depth dose measurements with a Markus ionchamber and several scintillators.

FIG. 10 shows exemplary Relative Biological Effective­ness (RBE) versus Linear Energy Transfer (LET).

FIG. 11 shows an apparatus which records multipleimages of the scintillator.

DETAILED DESCRIPTION OF THEPREFERRED EMBODIMENTS

In accordance with preferred embodiments of the presentinvention, the spatial distribution of a beam of penetratingradiation is imaged with a scintillator. The term "penetratingradiation," as used herein, and in any appended claims,refers both to particles with mass, such as protons, as wellas to photons, i.e., to electromagnetic radiation such asx-rays or gamma rays. Moreover, in the case of massiveparticles, the particles are typically charged, such as protonsor heavier atomic ions, however, neutrons or other electri­cally neutral particles may also be detected, and their beamsimaged within the scope of the present invention.

Referring to FIG. 2, a system for delivering a spatiallycomplex radiation beam is designated generally by 20. Thesystem consists of a radiation source, 21, and a means ofsetting the energy of the radiation, 22, setting the fluence,and means for shaping the transverse spatial distribution ofthe radiation, 23. In the case of charged particle beams, theshaping means can comprise a scatterer which scatters aunidirectional beam into a beam with a wide angular spreadin combination with a fixed or variable collimator or aper­ture which defines the shape of the beam. Alternatively, theshaping means can comprise two orthogonal scanning mag­nets which can steer the angle of a pencil beam, 24, from avirtual point located near the magnets across a field ofregard,25.

Referring to FIG. 3, a QA apparatus consists of a tissuephantom 31 disposed between the beam shaping means and

the detector 40 which images the radiation beam. There isprovision for varying the effective depth within the phantom,e.g., by varying the thickness of the phantom material,where the detector is located. For example, the tissuephantom may consist of a series of sheets of a polymermaterial, e.g., a water-equivalent plastic. Sheets may beadded or removed so as to vary the effective depth within thephantom where the detector images the beam. In order tosimulate measurement at depth within a body, the detector

10 and tissue phantom are typically translated away from thebeam source so as to keep the position of the initial surfaceof the phantom at the same distance from the beam source.

Alternatively, the tissue phantom may comprise a watertank. With the beam directed downwards, as is possible with

15 many radiotherapy delivery systems which comprise a rotat­ing gantry, the water level in the tank can be continuallyvaried. As water is added to the tank, the water tank anddetector can be translated downward, away from the beamsource, so as to preserve the distance from the beam source

20 to the top of the water layer, thereby simulating measure­ment at depth within a body. If the water tank comprises abellows-like apparatus, the thickness of the water layer canbe varied with a beam that is not vertical.

The detector comprises a thin scintillating screen 41,25 denoted the scintillator, at least one imaging sensor 44, and

a means of optical communication of the scintillation lightfrom the screen to the sensor, 42 and 43. In order for thedetector output to match a characteristic such as theabsorbed dose, the scintillator comprises a mixture of two or

30 more scintillating materials comprising a thin layer bound toa supporting substrate, typically a thin plastic film such asMylar. Since the output ofa single material will not typicallymatch the desired characteristic as a function of depth withinthe tissue phantom, the scintillator composition comprises a

35 mixture of materials whose combined output, in combina­tion with the rest of the apparatus, can match the desiredcharacteristic.

For example, it is known that many scintillators such asGSO, LYSO, Y3A150l2:Ce (P46), Gd202S:Tb (P43), and

40 Y2Si05:Ce (P47) exhibit quenching with increasing dE/dxor LET. That is, their light output per deposited energy,dUdE, decreases monotonically with increasing dE/dx suchas occurs within the Bragg peak. However, there are scin­tillators such as NaI(Tl), CsI(Tl), ZnS:Ag (Pll) and (Zn,

45 Cd)S:Ag (P20) which show anti-quenching behavior, i.e.,dUdE is not monotonic and increases with dE/dx for somevalues. Examples of both quenching and anti-quenchingbehavior are shown in R. B. Murray and A. Meyer, Scintil­lation Response of Activated Inorganic Crystals to Various

50 Charged Particles, Phys. Rev. Vol 122, No.3, pp. 815-826(1961), Y. Koba, H. Iwamoto, K. Kiyohara, T. Nagasaki, G.Wakabayashi, Y. Uozumi, N. Matsufuji, "Scintillation Effi­ciency of Inorganic Scintillators for Intermediate-EnergyCharged Particles," Progress in Nuclear Science And Tech-

55 nology, Vol. 1, p. 218-221 (2011), S. Safai, S. Lin, and E.Pedroni, "Development of an inorganic scintillating mixturefor proton beam verification dosimetry," Phys Med BioI.2004 Oct. 7; 49(19):4637-55 and S. Safai, Inorganic Scin­tillating Mixture And A Sensor Assembly For Charged

60 Particle Dosimetry, U.S. Pat. No. 8,080,801. FIG. 4 showsthe luminous efficiency of several scintillators as a functionof dE/dx.

One can generate an output signal that is proportional toa desired characteristic by combining the outputs from

65 different scintillators in ways detailed below. It is apparentthat for these common scintillators, different characteristicresponses such as quenching and anti-quenching behavior

US 9,750,958 B29

may also be accompanied by different spectral outputs.Therefore, the detection and combination ofthe signals fromeach scintillator may depend on their varying spectral out­puts. In order to generate an output signal that is propor­tional to a desired characteristic, several parameters must bechosen so that they work in combination to produce a signalthat is linear in the desired characteristic. These parametersinclude (a) the scintillator composition; (b) the spectraltransmission ofthe optics; and (c) the number ofand spectralsensitivity of imaging sensors which detect the scintillated 10

light.In a preferred embodiment of the invention, the scintil­

lating screen comprises a mixture of a quenching and ananti-quenching inorganic scintillator comprising a substrate 15

comprised of a thin sheet of supporting material, e.g.,polyester sheet, and a thin layer of the scintillating materialranging from 10-1000 mg/cm2 in thickness. In addition tothe scintillating material, a polymer material is typicallyincluded as a binding agent to dimensionally stabilize the 20

scintillator mixture and adhere the scintillating layer to thesubstrate that supports it. The substrate, the scintillatormixture and any binding agents will affect the characteristicsof the scintillation light output. The substrate may be clearor partially reflective. The inorganic scintillators are typi- 25

cally opaque and will scatter or absorb some or all of thescintillation light output from the portions ofthe scintillatinglayer proximal to the source of the radiation beam. A varietyof binding agents are known to the art, including a varietyof single component and dual component optical cements 30

and epoxies which cure by heat, UV light, or chemically.These binding agents are selected because of one or morecharacteristics including their cost, ease of use, their abilityto form a uniform scintillating layer, their optical clarity, and 35

their ability to withstand radiation without darkening, yel­lowing, or otherwise suffering radiation damage.

In a preferred embodiment of the invention, the light fromthe scintillator is communicated to the imaging sensor by amirror and a compound lens close to the imaging sensor. The 40

mirror, designated the fold mirror, directs light from thescintillator to the side so the imaging sensor can be locatedoutside of the radiation beam. The lens forms a highresolution image of the scintillating screen on the imagingsensor so there is an accurate, high resolution sampling of 45

the radiation beam as it is incident on the scintillator.Typically, both the fold mirror and the lens will have anonuniform spectral effect on the scintillation light trans­mitted to the sensor. The reflectance of the fold mirror,especially when used at a 45 degree angle, varies with 50

wavelength over the range of visible scintillators, and lenstransmission varies with wavelength, as well.

In a preferred embodiment of the invention, the imagingsensor comprises a thermoelectrically cooled CCD detector.

Safai et al. have disclosed a mixture scintillator which has 55

one component with a quenching characteristic and one withan anti-quenching characteristic, which in combination pro­duces a signal that has a quenching characteristic differentthan either component alone. QA of the treatment would besimpler if the signal is proportional to the deposited dose. 60

However, the spectral output of each scintillator componentis typically not identical. Thus, using the typical measure ofscintillator efficiency, output power per input power (W/W),is not sufficient to predict the detection signal. The signalwill depend on the interaction of the spectral content of the 65

scintillator output with the spectral transmission of theoptics and the spectral sensitivity of detectors which receive

10the light from the mixture scintillator. A signal proportionalto deposited dose may result with certain optics and detec­tors but not with others.

The spectral content and quantity of the scintillation lightis a function of several multiplicative factors within thelinear range of the scintillator and the sensor. First, eachscintillating component has a characteristic spectrum Sn(A)as well as an areal density Pn' Most important, its luminousefficiency dL/dE=En(LET, F), i.e., the unit light output is afunction of LET and the overall fluence F. The output isattenuated by a factor An(A) due to the self-absorption andscattering by the scintillation layer which depends on thescintillator density and the characteristics of the binder suchas the binder areal density and transmission. The imagingsensor will typically have a sensitivity Q(A) that reflects theeffective quantum efficiency of the sensor, as a function ofwavelength, in addition to a gain factor that depends on theelectronic amplification and digitization of the photodetec­tion signal. Finally, the spectral transmission T(A) of theoptics transmitting the scintillation light to the sensor alsodetermines the detected signal.

Thus, the detection signal D(LET, F) per deposited energyequals

For this equation, we have assumed that the characteris­tics of the scintillating screen are uniform across its area.While that is only true with a given accuracy-scintillatingscreen thickness can typically vary by several percent acrossa screen tens of cm in size-the variation is generallyindependent of LET and multiplies D(LET) uniformly as afunction of position or angle which can be calibrated by aflat field correction. In addition, evidence from Boon is thatthe spectral content of the scintillation output does not varywith LET for some common scintillators.

In order for the detection signal to match a characteristicof the radiation beam, the densities of the components of themixture must be engineered so that the detection signalmatches that characteristic, for a given spectral transmissionT(A) and spectral sensitivity Q(A). For a dosimetric detector,the desired characteristic is that the detection output isproportional to the energy deposited, i.e., D(LET, F) isessentially independent of the LET value and the fluence.For a two component scintillator, that can be accomplishedif one component shows decreasing output with LET(quenching), the other shows increasing output with LET(anti-quenching), and the proportions ofthe two componentsare such that the combination has uniform output with LET.Within a range of densities, the proportions should reduce toa ratio of the density of the two components.

However, it may be difficult to design that ratio inadvance. Factors such as An(A) factor are not well under­stood and hard to control, as it can be difficult in practice toprecisely control the absolute value of Pn deposited on thesubstrate and its uniformity. Factors such as Q(A) are typi­cally average values for sensors tested with a specific inputbeam of light. A particular sensor with different conditionsof the input beam may yield a measurably different effi­ciency as a function of wavelength due to interferenceeffects resulting from the thickness of a semiconductorsensor, for example.

It is possible to tailor the specifications of the optics tocompensate for such variations. FIG. 5 shows the consider-

US 9,750,958 B211

able variation in the spectral content of the light output ofdifferent inorganic scintillators. Thus, the relative contribu­tion of different components of a scintillator mixture to thedetection signal can be adjusted by varying the spectralthroughput of the optics. The optics transmission functionT(A) can be modified by including an element in the opticswith variable transmission or reflection as a function ofwavelength. In the case ofthe fold mirror, different reflectivecoatings such as aluminum, protected aluminum, and vari­ous dielectric coatings have non-uniform spectral reflectiv­ity as a function of wavelength. In the case of the lens,different optical glasses have non-uniform spectral transmis­sion. In addition, coatings applied to the glass can havenon-uniform spectral transmission. FIG. 6 shows exemplarytransmission curves for compound lenses. Furthermore, acolor filter can be used with a lens to enhance or diminishvarious portions of the spectrum.

Colored filters such as Wratten filters, Schott coloredglass, and Hoya colored glass can be used for this purpose.FIG. 7 shows exemplary transmission curves for coloredglass filters. In addition, numerous filters used for colortemperature adjustment, color correction, and light balanc­ing are available. These latter filters are typically designed tocorrect the effective color (black body) temperature fromartificial (tungsten or fluorescent) light to that of sunlight, orvice versa. They often come in a series which can applysmaller or larger color corrections depending on the inputlighting or desired output.

It is also possible to employ variable optical componentsin order to modifY the spectral response and hence thedegree to which the detection signal matches the desiredcharacteristic. For example, variable optical filters are avail­able with changeable passbands or transmission propertiesthat are modified by changing the temperature of the filter,the voltage across a liquid crystal transmissive element, orby some other means.

Another consideration is that properties of the systemcomponents may vary over time due to gradual modificationof their properties by aging, slow chemical reactions, orradiation damage. For example, radiation exposure can coloror darken optical glasses and optical cements such as thepolymer binder in the scintillator, or it can introduce crystaldefects into the scintillating material that change the effi­ciency of the scintillator. The performance of optical coat­ings can change over time, e.g., if they absorb moisture.Absorbing dyes in colored filters can change, chemically,over time to alter their transmission properties. Any of thesechanges will affect the elements of the right-hand side of theequation describing the detection signal given above.

Some of these variations in output over time do not affectthe basic ability of the system to measure the appropriatecharacteristic. A spectrally uniform reduction of the overallsignal, e.g., due to a colorless (gray) darkening of an opticalelement, effectively multiplies the detection signal equationby a constant factor independent of the wavelength A. Thismultiplicative or scale factor can be calibrated by measuringthe detection response to a radiation beam whose fluenceand energy spectrum are known from measurements withother detectors.

Other variations can affect the ability of the system tomeasure the appropriate characteristic. Variations thatchange the relative contributions of the various componentsin the scintillation mixture to the detection signal can changewhether the system linearly measures the appropriate char­acteristic of the radiation or not. Those variations maycomprise a differential change in (a) the output of thescintillator components or (b) a change in how that scintil-

12lator output is transmitted or detected. For example, one ofthe optical elements could undergo a colored darkening of amaterial due to radiation damage which would reduce thecontribution from a first scintillator component more thanfrom a second scintillator component.

A variation in one element of the equation can be com­pensated for by changing one or more of the other elementsof the system. For example, the colored darkening from theprevious example could be compensated for by adding a

10 filter which reduces transmission from the second compo­nent more than from the first component. Alternatively, ifradiation damage reduces output from a first component ofthe scintillator more rapidly than from a second component

15 of the scintillator, a filter could be added that preferentiallytransmits the spectral output of the first component relativeto the second component so as to compensate.

It may be difficult to predict the elements ofthe right-handside of the detection equation with sufficient accuracy. As

20 noted, the scintillator output can be quite sensitive to thescintillator mixture, including the binder, and the layerthickness. Moreover, the values of some elements of theequation may have significant uncertainties due to processvariation. There may also be practical considerations where

25 technical specifications for some components are unavail­able and it may be impractical to measure those specifica­tions of the components. In those cases, some assumptionsabout the specifications of those elements can be made inorder to generate an initial design of the system. The design

30 would include selection of the mixture proportions and thevarious components.

The initial design can be tested by performing calibrationmeasurements with a well-characterized radiation beam,

35 preferably monoenergetic, and observing the detection sig­nal as a function of depth inside the tissue phantom. Due tothe uncertainties, the signal may not match the desiredcharacteristic to a sufficient accuracy. In that case, the designcan be modified by changing either the scintillator compo-

40 sition, the optical system, or the imaging sensor. Forexample, if dose is the desired characteristic and it is foundthat the detection signal using a two-component scintillatorexhibits a quenching characteristic, another scintillatorcould be fabricated with the fraction of the quenching

45 phosphor reduced relative to the initial design.Another approach involves adding or changing an optical

filter that enhances the contribution of one component of thescintillator. Referring to the previous example, a filter couldbe added that has higher transmission for the spectral

50 content of the anti-quenching component than for the spec­tral content of the quenching component. A filter with thenecessary transmission ratio for the two components canadjust the output of the system so it matches the desiredcharacteristic response. In the case of dosimetry, the

55 adjusted system would not exhibit quenching or anti­quenching.

The images from the imaging sensor must undergo someprocessing in order to render the image value at each pixelproportional to the light incident on the sensor. It is known

60 to the art that two point correction consisting of subtractionof a bias value and multiplication of the result by a gainfactor is generally required to linearize the output of animaging detector. Moreover, these two point correctionstypically vary across the imaging sensor and from one copy

65 of a sensor to another so these correction factors are typi­cally specific to a particular detector and take on differentvalues for each pixel of the sensor. We thus speak of the bias

US 9,750,958 B213

and gain images or arrays that are used to process the rawsensor output, since the correction factors typically vary foreach pixel.

A series of calibrations are employed to measure the twopoint correction factors, i.e., recording the sensor responsewhen no radiation is incident (a dark exposure) or theresponse to a uniform radiation pattern (flat field). In somecases, the sensor output is non-linear with the underlyingcharacteristic, e.g., the intensity or fluence of the incidentradiation. In these cases, a non-linear transformation of thesensor signal is performed after its response has beencarefully calibrated. In what follows, the resulting signalsdescribed are those after the imaging sensor output has beenprocessed to produce a signal proportional to the incidentlight.

In accordance with a preferred embodiment of the inven­tion, the scintillator comprises a mixture oftwo components,by weight, of 25% Pll and 75% P43 in a layer with arealdensity 120 mg/cm2 on a thin polyester sheet with a polymerbinder. The resulting screen has transverse dimensions sev­eral centimeters or greater in order to simultaneously mea­sure an area with clinical significance, i.e., an area thatsurrounds a tumor. The optics comprise a 45 degree firstsurface mirror of enhanced aluminum coated on glass and aCanon EF camera lens with 50 mm focal length. Theimaging sensor comprises a Kodak KAF two-phase CCDemploying a transparent gate. FIG. 8 shows the typicalsensitivity of this CCD as a function of wavelength. Withthis composition, the image output is proportional to theradiation dose as measured by a Markus ion chamber.

Note that this composition of the scintillator comprises ananti-quenching blue phosphor, PI I , and a quenching yellow­green phosphor, P43. This choice of a scintillator mixturewith significantly different color outputs enables the detec­tion properties to be easily tuned with spectrally-varyingcomponents that enhance blue versus yellow and green orvIce versa.

A favorable aspect of this composition is that it does notuse toxic materials. Recently, the desire to reduce the use ofhazardous materials such as lead in solder has led to regu­latory requirements for electronic and other instrumentswhich do not use such materials. This trend is exemplified bydesignation of ROBS (reduction of hazardous substances)qualified components, which are now required in somejurisdictions. Other scintillators such as P20 or (Zn, Cd)S:Ag include cadmium, a hazardous material.

FIG. 9 shows the results of measurements with a Markusion chamber, a quenching scintillator (P46), and two mixturescintillators irradiated by a narrow energy spectrum protonbeam with varying thicknesses of a plastic absorber. The"Mixl" scintillator was designed with two scintillator com­ponents whose proportions and thickness, together with theoptics and imaging sensor, produce an output that matchesthe dose as measured by the Markus chamber. The Mixlscintillator is comprised, by weight, of 25% Pll and 75%P43. The "Mix2" scintillator was designed with a differentproportion of the same two scintillator components to besomewhat quenching. The Mix2 scintillator is comprised, byweight, of 12.5% Pll and 87.5% P43.

A scaling factor was applied to the P46, Mixl and Mix2measurements so the plateau region (starting at 0 em depth)of all 4 data sets lie on the same curve. As the curvesapproach the Bragg peak, they separate due to whether theyexhibit quenching or not, relative to the ion chamberresponse. As is evident from FIG. 9, the Mixl output

14matches the Markus chamber output very well and the Mix2output exhibits some quenching, though not as much as theP46 scintillator.

Measurements such as those shown in FIG. 9 indicatehow well the detection measurements match a characteristic,in this case defined by the ion chamber response. Althoughthe detection measurements with the Mix2 scintillator do notmatch the characteristic, the method described above couldbe used to adjust the detection response, namely, a filter

10 could be added with higher transmission for blue light thanyellow-green light. That would have the effect of enhancingthe anti-quenching component of the signal.

In addition to dosimetry, other characteristics of the15 deposited radiation may be of interest. For example, the

efficacy of radiation therapy is a function of the relativebiological efficiency (RBE) of a particular radiation beam.Different types of radiation-photons or protons or heavierions-have different RBE values and the RBE may be a

20 function of the LET value. FIG. 7 shows the RBE values asa function of LET for various probabilities of cell survival.So, it may be desirable to measure a detection signal that isappropriately weighted to match the RBE of the radiationbeam for a given probability of cell survival. If the scintil-

25 lator components have a different light output per unitenergy depending on LET, then scintillator mixtures can beproduced with light outputs that match the appropriate RBEof the beam for a given probability of cell survival.

A properly designed system has the detection signal30 proportional to the desired characteristic. Especially as treat­

ment modes incorporate very narrow pencil beams, theincident fluence may increase to the point where the lumi­nous efficiency (light output) En(LET, F) of one or morescintillator components is not linear with the fluence F but

35 saturates with increasing fluence. If there is significantdifferential non-linearity of the scintillator components, thedetection signal may only be linear with the characteristic tothe required accuracy for a limited range of fluence. Sincepencil radiation beams typically have a Gaussian intensity

40 profile, the fluence varies as a function of position, and thelinear measurement may only be achievable over a portionof the beam spot.

The system design can be tuned so the linear rangecorresponds with the clinically relevant range of fluence. In

45 this fashion, measurement of a characteristic of the beamthat is limited to a range offluence can be linearized with themultiple component scintillator. Furthermore, a changeableoptical filter, either by replacement or use of a variable filter,can be used to dynamically tune the optical system. Thus one

50 filter (or setting ofa variable filter) can be used for one rangeof fluence and another filter (or setting of a variable filter)can be used for a different range of fluence.

The discussion above shows how a custom spectral filtercan compensate for the detailed characteristics of a particu-

55 lar mixture scintillator. Filter selection can also compensatefor other non-idealities of the detector system or enhancecertain performance characteristics. For example, compo­nents of a scintillator may have different linear ranges as afunction of the incident fluence. A spectral filter can be

60 selected to enhance the signal contribution from a scintilla­tor component with a large linear range as a function of thefluence. Changing the filter can also compensate for someaging effects of the detector or the scintillator. For example,if radiation damage to the scintillator or a lens results in

65 relative changes to their spectral transmission, installing afilter with the opposite relative changes can make thedetector output insensitive to these aging effects.

US 9,750,958 B215 16

Dn(lLT, F)in = .z.: D;(LET, F)

a scintillating screen disposed behind the tissue phantomfor emitting light in response to the radiation receivedby the scintillating screen, the scintillating screen com­prising a mixture of at least two scintillators whereineach scintillator has a different characteristic responseand a different spectral output;

a means of optical communication between the output ofthe scintillating screen and at least one imaging sensor,wherein the means of optical communication has anonuniform spectral transmission; and

at least one imaging sensor in optical communication withthe scintillating screen for providing a high resolutionimaging sensor output indicative of the spatial distri­bution and intensity of the radiation beam, wherein theimaging sensor has a nonuniform spectral sensitivity;

wherein the composition of the scintillating screen, themeans of optical communication, and the at least oneimaging sensor form a system wherein the relativecontribution of the at least two scintillators of themixture produce an imaging sensor output which isproportional to a characteristic of the radiation beamincident on the scintillating screen at each measurementposition accessible with the tissue phantom;

wherein each scintillator of the mixture contributes adetection signal per deposited energy Dn(LET,F)=fdAPnSn(A)En(LET,F)An(A)Q(A)T(A) to the total imag­ing sensor output and a fractional contribution

65

to the total imaging sensor output;wherein n is the index ofthe nth scintillator in the mixture,

Dn(LET,F) is the detection signal per deposited energyfor the nth scintillator as a function of the linear energytransfer (LET) and fluence (F), Pn is the areal density ofthe nth scintillator, Sn(A) is the characteristic spectrumof the nth scintillator, En(LET,F) is the luminous effi-ciency of the nth scintillator as a function of the linearenergy transfer (LET) and fluence (F), An(A) is theattenuation factor for the nth scintillator due to self­absorption and scattering by the scintillation layer,Q(A) is the spectral sensitivity of the imaging sensor,and T(A) is the spectral transmission of the means ofoptical communication;

and wherein the fractional contribution, fm of at least onescintillator of the mixture to the sensor output varies byat least 3% from what it would be with a means ofoptical communication having uniform spectral trans­mission T(A)=I and with the at least one imagingsensor having uniform spectral sensitivity Q(A)=I for abeam with LET and fluence that is clinically relevant.

2. The apparatus as set forth in claim 1, wherein thecharacteristic of the radiation beam is the dose.

3. The apparatus as set forth in claim 1, wherein the60 characteristic of the radiation beam is the response of an ion

chamber.4. The apparatus as set forth in claim 1, wherein the

characteristic of the radiation beam is the relative biologicalefficiency.

5. The apparatus as set forth in claim 1, wherein thescintillating screen comprises a mixture of at least twoinorganic scintillators.

50

10

MODIFICATIONS

What is claimed is:1. An apparatus for determining the spatial distribution

and intensity of a penetrating radiation beam characterizedgenerally by a propagation direction inside a body, theapparatus comprising:

a tissue phantom disposed between a radiation source anda radiation detector;

In a preferred embodiment of the system, a multi-colorimage of the scintillator is formed. The multi-color imagecontains color and intensity information, i.e., two or morespectral component images with significantly non-overlap­ping spectral content. A detection signal can be produced byan arbitrary linear combination of the spectrally distinctcomponent images. In particular, an output proportional tothe desired characteristic can be formed by combining twoor more of the spectral component images with the appro­priate weighting factors.

The multi-color image of the scintillator can be formed ina variety of ways known to the art. Examples include videocameras with a patterned array of color filters overlaid on asingle focal plane sensor, typically a CCD or CMOS array.These cameras record intensity at each pixel, but color 15

information has lower resolution since the intensity ofseveral neighboring pixels with different color filters arecompared to derive the local color of the image. An examplewith comparable color and intensity resolution known to theart is a video cameras with a color separation prism and 3 20

CCDs to separately record red, green, and blue (RGB)component images.

An optimal detector for a two component scintillator mayonly record two color components whose spectral content ismatched to the scintillator components. In a preferred 25

embodiment, shown in FIG. 11, a dichroic mirror 45 sepa­rates the blue output of PII and the yellow-green output ofP43 and dual images are recorded by using two separateimaging sensors 44 and 47. Alternatively, the dual imagescould be spatially multiplexed onto a single imaging sensor. 30

The two images can be spatially registered and co-addedwith a weighting factor that produces the best fit to linearizethe output with respect to the desired characteristic, e.g., thewater-equivalent dose. Before co-adding the images, stan­dard two-point correction (spatially variable bias and gain) 35

or a more complex correction is applied to each spectrallydistinct image so it is a linear representation of the incidentlight in the corresponding spectral band.

More complicated detection signals can be synthesizedfrom the two (or more) spectral images than simply co- 40

adding the images with a single weighting factor. If one ormore of the scintillator components exhibits saturation withincreasing fluence, a non-linear function of the spectralimages can be constructed that linearizes the output withrespect to fluence, as well as the desired characteristic. This 45

can be implemented by calibrating the response and mod­eling the output signal at each pixel as a polynomial functionof the corresponding pixel in each spectrally distinct image.Alternatively, a lookup table that is indexed by the corre­sponding value from each separate image can be used.

The described embodiments of the invention are intendedto be merely exemplary and numerous variations and modi- 55

fications will be apparent to those skilled in the art. All suchvariations and modifications are intended to be within thescope of the present invention as defined in the appendedclaims.

US 9,750,958 B217 18

***

PI

PI +P2

P2

PI +P2

**

is in the range 10-40% and the fractional density ofP43,

is in the range 60-90%.19. The apparatus as set forth in claim 13 wherein the

fractional contribution, fm of at least one scintillator of themixture varies by at least 3% from what it would be with afilter having uniform spectral transmission.

forms two spatially distinct images on the at least oneimaging sensor, and further wherein the two spatially dis­tinct images are linearly combined to produce the highresolution imaging sensor output.

17. The apparatus as set forth in claim 1 wherein thecharacteristic of the radiation beam includes the fluence.

18. The apparatus as set forth in claim 7 wherein PII hasareal density PI' P43 has areal density P2' and the total arealdensity, Pl+P2' is in the range 40-400 mg/cm2,

10 and wherein the fractional density of Pll,

6. The apparatus as set forth in claim 1, wherein the atleast two scintillators are chosen from the group consistingof Pll, P20, P43, P46, and P47.

7. The apparatus as set forth in claim 1, wherein the atleast two scintillators comprise PII and P43.

8. The apparatus as set forth in claim 1, wherein thescintillating screen comprises a planar screen disposed per­pendicular to the propagation direction of the radiationbeam.

9. The apparatus as set forth in claim 1, wherein the meansof optical communication comprise a fold mirror and acompound lens.

10. The apparatus as set forth in claim 1, wherein the atleast one imaging sensor comprises a CCD.

11. The apparatus as set forth in claim 1, wherein the 15

tissue phantom comprises polymer sheets.12. The apparatus as set forth in claim 1 wherein the tissue

phantom comprises water.13. The apparatus as set forth in claim 1 wherein the

means of optical communication include at least one filter 20

with nonuniform spectral transmission.14. The apparatus as set forth in claim 1 wherein the

means of optical communication include at least one filterwith variable nonuniform spectral transmission.

15. The apparatus as set forth in claim 1 wherein the at 25

least one imaging sensor forms a multi-color image.16. The apparatus as set forth in claim 1 wherein the

means of optical communication includes a dichroic mirrorwhich separates the light emitted by the scintillating screeninto two significantly non-overlapping spectral bands and