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Dosimetric Characteristics ofDosimetric Characteristics of Clinical Photon Beams
Jatinder R Palta PhDUniversity of FloridaUniversity of Florida
Department of Radiation OncologyGainesville FloridaGainesville, Florida
DisclosuresDisclosures
Research development grants from PhilipsResearch development grants from Philips Medical Systems, Elekta Oncology Systems and Sun Nuclear AssociatesSystems, and Sun Nuclear Associates.NIH research award.B kh d C l h dBankhead Coley research award.
Learning Objectivesg jUnderstanding dosimetric properties of
clinical photon beams.Understanding physical parameters that g p y p
affect dosimetric properties of clinical photon beams.pUnderstand the need for accurate
characterization of clinical photon beamscharacterization of clinical photon beams in a treatment planning system.
Photon Beam Delivery SystemsMedical Linear Accelerators: Accelerate electrons in p lses Accelerate electrons in pulses to kinetic energies from 4 to 25 MeV. Use non-conservativeUse non-conservative microwave RF fields in the frequency range from 103 MHz (L band) to 104 MHz (X band), withband) to 10 MHz (X band), with the vast majority running at 2856 MHz (S band). Some provide beams only in
S band Linear Accelerators
the low megavoltage range (4-6 MV), while others provide both photons and electrons at various
l i A i lmegavoltage energies. A typical modern high-energy linac can provide 2-3 photon energies. X band Linear Accelerators
Sources of radiation that determine dosimetric characteristics of clinical photon beams
Source
Indirect (headscatter)
Direct Radiation (Focal Radiation)
Photon radiation generated at the target that reaches
Monitor Chamber
Flattening filter
gpatient without any intermediate interactions.
Indirect Radiation (Extra-focal Radiation):
Collimator jaws
MLC
Electron Contamination Direct
focal Radiation):Photon radiation with a history of interaction/scattering in the head of the treatment unit with the flattening filter
Output radiation or Incident radiation
MLC
Charged particle contamination dose
unit with the flattening filter, collimators, or other structures in the treatment head .
Contaminant electrons/positrons
S tt d
Primary dose electrons/positrons secondary electrons and positrons released from interactions with either the treatment head or the airScatter doseSecondary
electronstreatment head or the air column .
AAPM TG74 Report
Sources of Direct and Indirect RadiationRadiation
Direct
Indirect
A Monte Carlo study (Chaney et al., Med. Phys. 21,1994) Siemens MD2, 6MV
Characterizing Dosimetric Properties of Clinical Photon Beamsof Clinical Photon Beams
Beam penetrationNormalized depth dose (NDD) or tissue phantom ratioNormalized depth dose (NDD) or tissue phantom ratio
(TPR).Beam OutputTotal output ratio: Sc,p, in-air output ratio: Sc, phantom
scatter factor: Sp.
Cross beam profileCross-beam profile Isodose distribution.
Attenuation factors for beam modifiersAttenuation factors for beam modifiers hard wedges, compensators, trays, etc.
With the ultimate goal of ensuring that computerized treatmentWith the ultimate goal of ensuring that computerized treatment plans accurately reflect the dose received by patients
Beam Penetration
dDQfdNDD f
dref
d
D,,, QfsdNDD
where d is the depth of measurement on the
f
where d is the depth of measurement on the central axis of the phantom, s is the field size at the surface of the phantom, f is the source-surface-distance, Q is the quality of the clinical d
sq yphoton beam, and Dd and Ddref are dose at depth d and dref respectively.
d
Water
sSdf 2
TPR data can be determined from measured NDD as follows:
dp
drefpd sS
sSdreffdfQfsdNDDQsdTPR
,,,,,
Normalized Depth Dose DataNormalized Depth Dose DataEnergy Dependence
B ild i
15 MV
Buildup region
6 MV
15 MV
S f i
TCPE region
6 MVSurface region
FS = 10 x 10 cm2
Normalized Depth Dose DataNormalized Depth Dose DataField Size Dependence
15 MV Photon BeamThis depth corresponds to range 15 MV Photon Beamof the highest energy contaminant charged particles
16x16
4x4
Normalized Depth Dose DataNormalized Depth Dose DataWedge/Open Comparison
FS = 10 x 10 cm2FS = 10 x 10 cm2
15 MV (W/O)
6 MV (W/O)
Normalized Depth Dose DataNormalized Depth Dose DataWedge/Open Comparison
Minima
Normalized Depth Dose Data
These data from- Siemens-- Varian. Elekta
These data from Radiological Physics Center show that all NDD. Elekta
18 MV
show that all NDD for both 6 and 18 MV photon beams at depths of 5 cm
18 MV and 15 cm for
different field sizes have a
i % f
Field sizes: 6x6, 10x10 and 20x20 cm2
6 MV
maximum %σ of 0.5% and this increases to 0.7% at a depth of 20 cm
cm2
at a depth of 20 cm.
Monte Carlo Calculated Photon Beam SpectraBeam Spectra
•The spectral shapes•The spectral shapes are somewhat similar
•The differences at the high-energy end are caused by the differences in the mean incident electronmean incident electron energies and their spread
Sheikh-Bagheri & Rogers, Med. Phys., 29, 2002
spread
Monte Carlo Calculated Average Energies
•The average energies for the same nominalfor the same nominal accelerating potential are somewhat similarare somewhat similar
•The average energies decrease at off-axis distances for all clinical bbeams
• more pronounced difference at higher energies
Sheikh-Bagheri & Rogers, Med. Phys., 29, 2002
g g
Beam Penetration for Irregularly-Shaped FieldsShaped Fields
fConcept of Equivalent Square:
fp q q
The equivalent field is defined as that standard (square or circular) field which has the same central-
ds
field which has the same central-axis depth dose characteristics as the given non-standard field.
d
Water“Day’s Rule”:
rrSS 1 rr ereSrS 1S(r) = the central axis scatter in a field of radius r, S∞ = the central axis scatter in fi ld f i fi it di λ i li t d i di i l hfield of infinite radius, λ is a scaling parameter, and μ is a dimensionless shape parameter. They computed equivalent square fields for a complete set of rectangular fields using a value of λ=0.26 cm-1 and μ=0.5.
Equivalent square
L
Ws
Equivalent square d
Sterling Formula: g(Sterling et.al., Brit. J. Radiol. 37, 544 (1964))
PALWS /42
Assuming, λ = 0.26 cm-1., and μ = 0.5
PAWL
S /4
Assuming, λ 0.26 cm 1., and μ 0.5/2 /2
0 0( , ) 4 ( , )
L WS L W D x y dxdy 0 0
/L W( , ) / (10,10)S L W S
1 2 3 4 51.000 0.993 0.982 0.969 0.958( , ) / (10,10)S L W S 1.000 0.993 0.982 0.969 0.958
22
KLEIN- NISHINA CROSS SECTION FOR THE COMPTON INTERACTION
2
220 sin'
''
2
hh
hh
hhr
dd
Ψ
e
PHOTONS SCATTERED INTO A UNIT SOLID ANGLE, Ω
SOLID ANGLE AVAILABLE PER
sin2dd
SOLID ANGLE AVAILABLE PER UNIT ANGLE
PHOTONS SCATTERED AT AN ANGLE, Ψ
d
Based on the kinematics of Compton interaction, the average p genergy of scattered photons is less than 1Mev and is independent of the incident energy.
Measurement of Normalized Depth D dDose data
Follow AAPM TG Report # 106 recommendations:pUse 4-5 mm diameter ion chamber for depth
beyond 1cm.Use parallel plate or extrapolation chamber to
measure data near the surface.Di d d di d d t t i tDiodes and diamond detectors are appropriate
as long as data measured with these detectors is cross-referenced to data measured with an s c oss e e e ced to data easu ed t aion chamber.Prone to radiation damage and non-linear response.
Is depth ionization data depth dose?YES!!!With the caveat, TCPE exists at the point of measurement TCPE exists at the point of measurement. the energy spectrum of incident photons does not change with the
depth. fluence across the detector remains the same fluence across the detector remains the same.These conditions are met at depths beyond the range of
contaminant charged particles
However at shallow depth, The contaminants and secondary electrons have energy spectra that change rapidly with depth. Results in a variation of ~10% in restricted mass stopping power ratio data for water and air.
Translates into a spatial uncertainty of less than 1.5 mm in dose in the build up region
Beam Output
ff ffSc Sc,p
10 cm
cc
Water
c
cSsS
sSc
pcp
, (Derived)
In-air output Ratio Elekta: 4-18 MV clinical photon beamsElekta: 4-18 MV clinical photon beams.
Monte Carlo Calculations of In-Ai O t t R tiAir Output Ratio
(BEAMnrc code)
In-Air Output Ratio
1.05
0.95
1.00
6 MV measured6 MV calc lated
Oo
0 5 10 15 20 25 30 35 40 45
0.90
6 MV calculated 18 MV measured 18 MV calculatedSimulation Geometry
(Varian 2100EX)0 5 10 15 20 25 30 35 40 45
/tex/rof/clxyro
Side of square field /cm
Energy spectrum of head scattered photons
Mean Energy:0.5 MeV
(Varian 2100C.)
Energy spectrum of head scattered photons
(Varian 2100C.)
Mean Energy:0 5 MeVMean Energy:0.5 MeV
In-air output Ratio e: Elekta, s: Siemens, and v: Varian , ,
(for clinical photon beams ranging from 6-25 MV.
Monitor Back ScatterMachine MBS Publication
Flattening FilterMonitor Chamber
Beam Modifier(internal wedge)
Upper Collimator
Varian Clinac 1800 1-5% Kubo, Med. Phys.16, 295 (1987)
Therac 20 7.5% Hounsell, P.M.B.Lower Collimator
Tertiary Collimator(Cerrobend Block
or Varian MLC)Beam Modifier(external wedge)
,43, 445 (1998)
Elekta SL15 <1% Yu et.al. P.M.B.(with 3 mm AL) 41, ( )
1107(1996)5%
(without Al)
Varian 600c/2100C 2-5% Lam et. al. Med.Varian 2100C Phys. 25, 334
(1998)
The differences in In-Air Output Ratio for the same field size on different machines is primarily attributed to the difference in monitor back scatter
M t f I Ai O t t R tiMeasurement of In-Air Output Ratios• Mini phantomp
– Water-equivalent materials.– 4g/cm2 diameter and 10g/cm2 depth to maintain lateral
CPE and eliminate contaminant electronCPE and eliminate contaminant electron.• For small segment fields (c<4cm), high Z material
(Brass etc.) should be used.– Corrections for energy absorption coefficients and energy
spectra change are needed.r1
h
1
2TG 74 recommendations
Cross Beam CharacteristicsCross Beam Characteristics Affected by the radially symmetric conical high Z-
material flattening filter, whichg , Flattens the beam by differentially absorbing more photons in the
center and less in the periphery unwanted consequence of flattening the beam is the differential
change in beam quality at off-axis pointschange in beam quality at off-axis points. hardens the beam
Cross beam flatness is defined as:minmax
minmax100DDDD
F
One flattening filter for each clinical photon beam results in a compromise of beam flatness characteristics of small and large fields. Fl tt i filt d i d t i d ll i i di l i t it Flattening filters are designed to give a gradually increasing radial intensity.
This is referred to as “horns” on a cross-beam profile
Cross beam profiles may not be radially symmetric due to non circular focal spot.p Therefore, cross-beam data is characterized by a set of two
orthogonal dose profiles measured perpendicular to the beam’s central axis at a given depth in a phantom
Cross Beam Profile6 MV Photon Beam, Depth of 5.0 cm, Field size of 4x4, 10.4x10.4, and 21x21 cm2.
The flatness of photon beams is extremely sensitive to change in energy of the incident beam. A small change in the penetrative quality of a photon beam results in very large change in beam flatness.
Cross Beam Profile6 MV Photon Beam, Field Size of 10.4x10.4 cm2, Depths of 1.5, 5.0, 10.0, 15.0, , , p , , , ,
and 25.0 cm.
The field flatness changes with depth. This is attributed to an increase in scatter to primary dose ratio with increasing depth and decreasing incident photon energy off axis
Effect of Electron Steeringon Beam Flatnesson Beam Flatness
Symmetric Tilted Displaced
Effect of a Dipole Magnet on Exit p gBeam
Radial DivergenceRadial DisplacementEnergy Spread
Cross Beam SymmetryCross Beam Symmetry
rightleft AreaAreaS
100
rightleft AreaArea
Dosimetry and beam steering system
Isodose DistributionIsodose Distribution
30 cm X 30 cm18 MV X-ray beam
Isodose Distributions(20 X 20 Cm2)(20 X 20 Cm2)
6 MV 18 MV
Note contaminant electrons contribute to dose outside the field at shallow depths. The magnitude and extent of dose outside at s a o dept s e ag tude a d e te t o dose outs dethe geometric edge of a field at shallow depths increases with beam energy.
Isodose Distributions(20 X 20 Cm2 18 MV)(20 X 20 Cm2, 18 MV)
Note Contaminant electrons contribute to dose outside the field at shallow depths. The magnitude and extent of dose p goutside the geometric edge of a field at shallow depths increases even more in the presence of beam modifiers.
Cross Beam Measurements
Wh t i th ff t fWhat is the affect of detector size?Incorrect measurement of penumbra region
Diode CC04 CC13
Diameter 0.8x0.8 mm2 4 mm 6 mm
Penumbra20%~80% 4.0 mm 6.1 mm 7.2 mm
Detector Size Effect on TPS Commissioning
Impact ofTreatment Planning
SystemCommissioning
Impact of detector size
effect on dose di t ib ti ???Commissioning distribution???
Yan G et. al., Med. Phys (35)., 2008
Extraction of True Profile
IMRT QA results: DTA 2%/2 mm
CC13CC13CC04Deconvolved
Measurement of Attenuation F t f B M difiFactors for Beam Modifiers
The attenuation factor for a beam modifier is defined as the ratio of the dose rate at the point of calculation for a given field with and without the modifier in place. Attenuation factors for devices such as block trays, accessories y ,
etc. are often assumed to be independent of field size, depth and SSD. These factors should be measured at a depth well beyond the
maximum range of electron contaminationmaximum range of electron contamination
The attenuation devices that are in contact with the patient skin (immobilization apparatus, table top, etc.) req ire additional considerationsrequire additional considerations. These devices not only attenuate the incident beam but they
introduce scatter radiation that increase the scatter to primary ratio within the patientratio within the patient. It is best to include such attenuation devices as a part of the patient
in 3DRTPS
Measurement of Wedge FactorsMeasurement of Wedge FactorsThe WF is defined as the ratio of the dose rate
t th f d th f d d fi ld t th tat the reference depth for a wedged field to that for the same field without a wedge modifier .The field size dependency of the WF originates fromThe field size dependency of the WF originates from
a wedge-induced increase in head scatter. the field size dependence of the WF is correctly p y
accounted for by in-air output ratios (Sc)wedgespecifically measured for wedged fieldsThese data should be measured with the chamber axisThese data should be measured with the chamber axis
perpendicular to the gradient direction of the wedgeTwo sets of measurements should be made with the wedge
in opposite orientations to ensure the correct placement ofin opposite orientations to ensure the correct placement of the chamber
Characterizing Clinical Photon B i 3DRTPSBeams in 3DRTPS
Ahnesjo et al PMB 1999Ahnesjo et al., PMB 1999
Approaches to Dose Computation Algorithms
Data measured in water and in air
Parameterize water data
Reconstitute water data Calculate dose directly based on beam and
Calculate inhomogeneity corrections to water data
based on beam and phantom configurations
““CorrectionCorrection” based ” based methodsmethods
““CorrectionCorrection” based ” based methodsmethods
““ModelModel” based ” based methodsmethods
““ModelModel” based ” based methodsmethodsmethodsmethodsmethodsmethods methodsmethodsmethodsmethods
Figure 8.9,The Modern Technology of Radiation Oncology; J. Van Dyk
Correction vs. Model Based Methods
Correction Based Model BasedMeasured data used as basis for Dose Computation.
Measured data used to setup description of treatment beam.
Require measurements with buildup Require a parameter to estimate size cap in air or in a mini-phantom. of photon source at target.
Require lots of data. Generating functions used to reduce size of
Require more time for tuning of model parameters.
data set for convenient clinical use (i.e. less storage space).
p
Patient dose distribution obtained by Patient dose distribution obtained byPatient dose distribution obtained by first computing Dose in water from generating function, then correcting for tissue heterogeneity, patient
t d b difi
Patient dose distribution obtained by computing beam and beam transport (i.e. beam interactions in treatment head and in patient) directly.
contour, and beam modifiers.
Accuracy Goal in Dose CalculationsAccuracy Goal in Dose Calculations
• Required accuracy (overall treatment < 5%):q y ( )
Ahnesjo et al., PMB 1999
Characterizing Clinical Photon Beams in 3DRTPSBeams in 3DRTPS
MUST model the following features realistically:MUST model the following features realistically:Finite size of source (& penumbra)E t f l di ti ( i lli t fl tt i filt )Extra-focal radiation (primary collimator, flattening filter)
Beam spectrum (& change in spectrum with position)
Beam intensity variation across field (e.g., beam horns)Transmission through secondary collimatorsS tt t id fi ld ( l t d t t f l di ti )Scatter outside field (related to extra-focal radiation)MLC, blocks, block trayDynamic wedge fixed wedge compensators (beamDynamic wedge, fixed wedge, compensators (beam
hardening)
Characterizing Clinical Photon Beams in 3DRTPSBeams in 3DRTPS
Caveats:Caveats: Almost all photon dose computation with convolution models
assumes kernel invariance, which requires the photon dose kernel to be constant with spatial locations in the calculation phantombe constant with spatial locations in the calculation phantom. However, in clinical treatments, patient inhomogeneities, as well as beam
divergence and polychromaticity, cause kernel variation in various ways.
Modeling of charged particle contaminants is at best an approximation of real clinical situation
Modeling of indirect radiation as a single or multiple analytical Modeling of indirect radiation as a single or multiple analytical source functions, modeling of off-axis softening with a simple parametric fit, source size, etc. are best effort estimates of physical processesprocesses
Characterizing Clinical Photon Beams in 3DRTPSBeams in 3DRTPS
Caveats (continued):One can always use a set of beam modeling parameters
to get the best agreement between the computed and measured beam data in a phantommeasured beam data in a phantom. . However, that would not be a sufficient condition for robust and
accurate beam modeling .
The value or function used to describe a parameter should have some physical meaning.each parameter used in the dose calculation algorithm should
model the physical reality it represents even if there is less than perfect agreement between measure and computed data. The observed differences often reflect limitations of the dose
computation algorithm
Benchmark Dataset(D l d d NIH i iti ti )(Developed under NIH initiative)
A collaborative effort involving Sun Nuclear Associates; the t t d lt t f th U i it f Fl idcontractor, and consultants from: the University of Florida;
the RPC at M.D. Anderson Cancer Center; the University of Iowa; and the Vassar Brothers Hospital. p
Already measured a complete set of data on the new generation of Elekta (Synergy), Siemens (Oncor) and Varian (Trilogy) linear acceleratorsVarian (Trilogy) linear acceleratorsMeasured data are comprehensive in beam geometries to
validate dose computation for any clinical situation.data are sufficient in spatial resolution and were validated by
independent measurements
This benchmark datasets will be sufficient for the TPSThis benchmark datasets will be sufficient for the TPS companies to compare the accuracy of their dose modeling for treatment delivery
SummaryyThe dosimetric properties of a clinical photon
beam are characterized by:beam are characterized by: Its ability to penetrate a tissue-like medium (water) its change in dose output with field sizeg p Its cross beam behavior Its attenuation through modifying devices (e.g., wedge,
compensator etc )compensator etc.) The dosimetric properties of clinical photon
beams from linacs depend on the photon energybeams from linacs depend on the photon energy fluence distribution emanating from the treatment head, on the geometry of the linac, and on the radiological properties of the medium with which itradiological properties of the medium with which it interacts.
SummaryyIt is quite evident that all modern clinical
li l t (li ) f ti llinear accelerators (linacs) of a particular commercial make produce beams of very similar characteristicssimilar characteristics High quality benchmark data have already been
acquired by comprehensively characterizingacquired by comprehensively characterizing single linacs of each make. These benchmark data thoroughly describe the
characteristics of photon beams so that treatment-planning companies and clinics throughout the United States can use it tothroughout the United States can use it to examine the accuracy of dose-calculation algorithms.