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.