Dosimetric Characteristics of Clinical Photon Beams
Jatinder R Palta PhD University of Florida Department of Radiation Oncology Gainesville, Florida Disclosures
Research development grants from Philips Medical Systems, Elekta Oncology Systems, and Sun Nuclear Associates. NIH research award. BkhdClBankhead Coley researc h awar d. Learninggj Objectives
Understanding dosimetric properties of clinical photon beams. Understandinggp ph ysical parameters that affect dosimetric properties of clinical photon beams. Understand the need for accurate characterization of clinical photon beams in a treatment planning system. Photon Beam Delivery Systems
Medical Linear Accelerators: Accelerate electrons in pu lses to kinetic energies from 4 to 25 MeV. Use non-conservative microwave RF fields in the frequency range from 103 MHz (L band) to 104 MHz (X band), with S band Linear Accelerators the vast majority running at 2856 MHz (S band). Some provide beams only in the low megavoltage range (4-6 MV), while others provide both photons and electrons at various megavoliAilltage energies. A typical modern high-energy linac can X band Linear Accelerators provide 2-3 photon energies. Sources of radiation that determine dosimetric characteristics of clinical photon beams
Source Direct Radiation (Focal Radiation) Indirect (headscatter) Photon radiation generated at the target that reaches patient without any Flattening filter intermediate interactions.
Monitor Chamber Indirect Radiation (Extra- focal Radiation): Collimator jaws Photon radiation with a Electron history of interaction/scattering Direct Contamination in the head of the treatment MLC unit with the flattening filter, collimators, or other structures Output radiation or in the treatment head . Charged particle Incident radiation Contaminant contamination dose Primary dose electrons/positrons secondary electrons and positrons released from interactions with either the Secondary SttdScatter dose treatment head or the air electrons column .
AAPM TG74 Report Sources of Direct and Indirect Radiation
Direct
Indirect
A Monte Carlo study (Chaney et al., Med. Phys. 21,1994) Siemens MD2, 6MV Characterizing Dosimetric Properties of Clinical Photon Beams Beam penetration Normalized depth dose (NDD) or tissue phantom ratio (TPR). Beam Output
Total output ratio: Sc,p, in-air output ratio: Sc, phantom scatter factor: Sp. Cross-beam profile Isodose distribution. Attenuation factors for beam modifiers hard wedges, compensators, trays, etc. With the ultimate goal of ensuring that computerized treatment plans accurately reflect the dose received by patients Beam Penetration
Dd NDDd, s, f ,Q f Ddref 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- s surface-distance, Q is the qqyuality of the clinical d photon beam, and Dd and Ddref are dose at depth d and dref respectively. Water
TPR data can be determined from measured NDD as follows: 2 f d S p sdref TPRd, sd ,Q NDDd, s, f ,Q f dref S p sd Normalized Depth Dose Data Energy Dependence
Bu ildup reg ion
15 MV TCPE region
6MV6 MV SfSurface regi on
FS = 10 x 10 cm2 Normalized Depth Dose Data Field Size Dependence This depth corresponds to range 15 MV Photon Beam of the highest energy contaminant charged particles
16x16 4x4 Normalized Depth Dose Data Wedge/Open Comparison
FS=10x10cm2FS = 10 x 10 cm2
15 MV (W/O)
6 MV (W/O) Normalized Depth Dose Data Wedge/Open Comparison
Minima Normalized Depth Dose Data
These data from - Siemens Radiological -- Varian Physics Center .Elekta. Elekta show that all NDD for both 6 and 18 MV photon beams at depths of 5 cm Field sizes: 18 MV and 15 cm for
6x6, 10x10 different field sizes and 20x20 have a 2 cm maxi%imum %σ of 0.5% and this 6 MV increases to 0.7% at a depth of 20 cm. Monte Carlo Calculated Photon Beam Spectra
•The spectral shapes are somewhat similar
•The differences at the high-energy end are caused by the differences in the mean incident electron energies and their spread
Sheikh-Bagheri & Rogers, Med. Phys., 29, 2002 Monte Carlo Calculated Average Energies
•The average energies for the same nominal accelerating potential are somewhat similar
•The average energies decrease at off-axis distances for all clinical beams • more pronounced difference at higggher energies
Sheikh-Bagheri & Rogers, Med. Phys., 29, 2002 Beam Penetration for Irregularly- Shaped Fields Conceppqt of Equivalent S quare: The equivalent field is defined as f that standard (square or circular) field which has the same central- axis depth dose characteristics as the given non-standard field. s d “Day’s Rule”: Water r r Sr S 1 e r e
S(r) = the central axis scatter in a field of radius r, S∞ = the central axis scatter in field o f in finit e ra dius, λ ilitdis a scaling parameter, and μ idiilis a dimensionless s hape 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. L s W Equivalent square d
Sterling Formula: (Sterling et.al., Brit. J. Radiol. 37, 544 (1964)) 2LW S 4A/ P L W Assuming, λ = 0.26 cm -1., and μ = 0.5
LW/2 /2 SLW(, ) 4 Dxydxdy (,) 00 L /W 12345 SLW( , ) / S (10,10) 1.000 0.993 0.982 0.969 0.958 KLEIN- NISHINA CROSS SECTION FOR THE COMPTON INTERACTION 2 d r 2 h ' h h ' e 0 sin 2 dΨ 2 h h ' h
PHOTONS SCATTERED INTO A UNIT SOLID ANGLE, Ω
SOLID ANGLE AVAILABLE PER UNIT ANGLE d 2 sin d
PHOTONS SCATTERED AT AN ANGLE, Ψ Based on the kinematics of Comppgton interaction, the average energy of scattered photons is less than 1Mev and is independent of the incident energy. Measurement of Normalized Depth DdDose data Follow AAPM TG Report # 106 recommendations: Use 4-5 mm diameter ion chamber for depth beyond 1cm. Use parallel plate or extrapolation chamber to measure data near the surface. Diod es and diamon d de tec tors are appropr ia te as long as data measured with these detectors isscoss cross-ref er en ced to data m easur ed wi th an ion 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. the energy spectrum of incident photons does not change with the depth. 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
f f
Sc Sc,p
10 cm c c Water S s c, p (Derived) S p s Sc c In-air output Ratio Elekta: 4 -18 MV clinical photon beams. Monte Carlo Calculations of In- Air O ut put R ati o (BEAMnrc code)
In-Air Output Ratio
1.05
1.00 o O
0.95 6 MV measured 6 MV calcu lated 18 MV measured 18 MV calculated Simulation Geometry 0.90 (Varian 2100EX) 0 5 10 15 20 25 30 35 40 45 Side of square field /cm
/tex/rof/clxyro 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 MeV In-air output Ratio e: Elekta,,, s: Siemens, and v: Varian (for clinical photon beams ranging from 6-25 MV. Monitor Back Scatter
Machine MBS Publication
Flattening Filter Varian Clinac 1800 1-5% Kubo, Med. Phys. Monitor Chamber Beam Modifier 16, 295 (1987) (internal wedge) Upper Collimator Therac 20 7.5% Hounsell, P.M.B.
Lower Collimator 43, 445 (1998)
Tertiary Collimator Beam Modifier (Cerrobend Block Elekta SL15 <1% Yu et.al. P.M.B. (external wedge) or Varian MLC) (()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 MtfIMeasurement of In-Air O ut put R ati os • Mini phantom – Water-equivalent materials. – 4g/cm2 diameter and 10g/cm2 depth to maintain lateral CPE 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 TG 74 recommendations 2 Cross Beam Characteristics Affected by the radially symmetric conical high Z- material flatteningg, filter, which 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 points. hardens the beam D D Cross beam flatness is defined as: F 100 max min Dmax Dmin One flattening filter for each clinical photon beam results in a compromise of beam flatness characteristics of small and large fields. Fla tten ing filters are des igne d to g ive a gra dua lly increas ing ra dia l in tens ity. 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. 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 Profile 6 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 Profile 6 MV Photon Beam,,,p,,,, Field Size of 10.4x10.4 cm2, Depths of 1.5, 5.0, 10.0, 15.0, 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 Steering on Beam Flatness
Symmetric Tilted Displaced Effect of a Dippgole Magnet on Exit Beam
Energy Spread Radial Displacement Radial Divergence Cross Beam Symmetry
Area Area S 100 left right Arealeft Arearight
Dosimetry and beam steering system Isodose Distribution
30 cm X 30 cm 18 MV X-ray beam Isodose Distributions (20 X 20 Cm 2)
6 MV 18 MV
Note contaminant electrons contribute to dose outside the field at s ha llo w dept hs . The mag nitude an d ex ten t of dose outsi de the geometric edge of a field at shallow depths increases with beam energy. Isodose Distributions (20 X 20 Cm 2, 18 MV)
Note Contaminant electrons contribute to dose outside the field at shallow deppgths. The magnitude and extent of dose outside the geometric edge of a field at shallow depths increases even more in the presence of beam modifiers. Cross Beam Measurements
Wha t is the aff ect of detector size? Incorrect measurement of penumbra region
Diode CC04 CC13 0.8x0.8 Diameter 4 mm 6 mm mm2 Penumbra 4.0 mm 6.1 mm 7.2 mm 20%~80% Detector Size Effect on TPS Commissioning
Impact of Treatment Planning detector size System effect on dose Commissioning dist rib uti on???
Yan G et. al., Med. Phys (35)., 2008 Extraction of True Profile IMRT QA results: DTA 2%/2 mm
CC13 CC04 Deconvolved Measurement of Attenuation FtFactors f fBor Beam M Mdifiodifiers 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 trayy,s, accessories 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 contamination The attenuation devices that are in contact with the patient skin (immobilization apparatus, table top, etc.) require additional considerations. These devices not only attenuate the incident beam but they introduce scatter radiation that increase the scatter to primary ratio within the patient. It is best to include such attenuation devices as a part of the patient in 3DRTPS Measurement of Wedge Factors
The WF is defined as the ratio of the dose rate atthft the reference d dthfepth for a wed ddfildtthtged field to that for the same field without a wedge modifier . The field size dependency of the WF originates from a wedge-induced increase in head scatter. the field size deppyendence of the WF is correctly
accounted for by in-air output ratios (Sc)wedge specifically measured for wedged fields These data should be measured with the chamber axis perpendicular to the gradient direction of the wedge Two sets of measurements should be made with the wedge in opposite orientations to ensure the correct placement of the chamber Characterizing Clinical Photon Bi3DRTPSBeams in 3DRTPS
Ahnesjo 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 phantom configurations corrections to water data
““CorrectionCorrection” based ““ModelModel” based methods methods
Figure 8.9,The Modern Technology of Radiation Oncology; J. Van Dyk Correction vs. Model Based Methods
Correction Based Model Based Measured data used as basis for Measured data used to setup Dose Computation. 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 Require more time for tuning of functions used to reduce size of model parameters. data set for convenient clinical use (i.e. less storage space).
Patient dose distribution obtained by Patient dose distribution obtained by first computing Dose in water from computing beam and beam transport generating function, then correcting (i.e. beam interactions in treatment for tissue heterogeneity, patient head and in patient) directly. contdbdifitour, and beam modifiers. Accuracy Goal in Dose Calculations
•Reqqyuired accuracy (overall treatment < 5%):
Ahnesjo et al., PMB 1999 Characterizing Clinical Photon Beams in 3DRTPS
MUST model the following features realistically: Finite size of source (& penumbra) EtExtra-focal radi ati on ( pri mary colli ma tor, fla tten ing filter ) Beam spectrum (& change in spectrum with position) Beam intensity variation across field (e.g., beam horns) Transmission through secondary collimators Sca tter out sid e field (re la te d to ex tra-flditi)focal radiation) MLC, blocks, block tray Dynamic wedge, fixed wedge, compensators (beam hardening) Characterizing Clinical Photon Beams in 3DRTPS
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 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 source functions, modeling of off-axis softening with a simple parametric fit, source size, etc. are best effort estimates of physical processes Characterizing Clinical Photon Beams 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 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 (Deve lope d un der NIH in itia tive )
A collaborative effort involving Sun Nuclear Associates; the contttractor, and consult lttfants from: th thUiitfFlide University of Florida; the RPC at M.D. Anderson Cancer Center; the University of Iowa; and the Vassar Brothers Hospital. Already measured a complete set of data on the new generation of Elekta (Synergy), Siemens (Oncor) and Varian (Trilogy) linear accelerators Measured 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 TPS companies to compare the accuracy of their dose modeling for treatment delivery Summary
The dosimetric properties of a clinical photon beam are characterized by: Its ability to penetrate a tissue-like medium (water) its changgpe in dose output with field size Its cross beam behavior Its attenuation through modifying devices (e.g., wedge, compensator etc .) The dosimetric properties of clinical photon beams 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 it interacts. Summary It is quite evident that all modern clinical linear acce lera tors (linacs ) o f a par ticu lar commercial make produce beams of very similar characteristics High quality benchmark data have already been acquired 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 to examine the accuracy of dose-calculation algorithms.