Clinical Photon Beams

Transcription

1 Dosimetric Characteristics of Clinical Photon Beams Jatinder R Palta PhD University of Florida Department of Radiation Oncology Gainesville, Florida

2 Disclosures Research development grants from Philips Medical Systems, Elekta Oncology Systems, and Sun Nuclear Associates. NIH research award. Bankhead d Coley research award.

3 Learning Objectives Understanding dosimetric properties of clinical photon beams. Understanding gphysical parameters that affect dosimetric properties of clinical photon beams. Understand the need for accurate characterization of clinical photon beams in a treatment planning system.

4 Photon Beam Delivery Systems S band Linear Accelerators X band Linear Accelerators Medical Linear Accelerators: Accelerate electrons in pulses to kinetic energies from 4 to 25 MeV. Use non-conservative microwave RF fields in the frequency range from 10 3 MHz (L band) to 10 4 MHz (X band), with 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 megavoltage energies. A typical modern high-energy linac can provide 2-3 photon energies.

5 Sources of radiation that determine dosimetric characteristics of clinical photon beams Indirect (headscatter) Flattening filter Monitor Chamber Collimator jaws Electron Contamination MLC Charged particle contamination dose Secondary electrons Source Direct Output radiation or Incident radiation Primary dose Scatter dose Direct Radiation (Focal Radiation) Photon radiation generated at the target that reaches patient without any intermediate interactions. Indirect Radiation (Extra- focal Radiation): Photon radiation with a history of interaction/scattering in the head of the treatment unit with the flattening filter, collimators, or other structures in the treatment head. Contaminant electrons/positrons secondary electrons and positrons released from interactions with either the treatment head or the air column. AAPM TG74 Report

6 Sources of Direct and Indirect Radiation Direct Indirect A Monte Carlo study (Chaney et al., Med. Phys. 21,1994) Siemens MD2, 6MV

7 Characterizing Dosimetric Properties of Clinical Photon Beams Beam penetration Normalized depth dose (NDD) or tissue phantom ratio (TPR). Beam Output Total output ratio: S c,p, in-air output ratio: S c, phantom scatter factor: S p. Cross-beam profile Isodose distribution. Attenuation factors for beam modifiers hard wedges, compensators, trays, etc. With the ultimate goal of ensuring that computerized treatment With the ultimate goal of ensuring that computerized treatment plans accurately reflect the dose received by patients

8 Beam Penetration NDD d D d d, s, f, Q D dref 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 sourcesurface-distance, Q is the quality of the clinical photon beam, and D d and D dref are dose at depth d and dref respectively. Water s f d TPR data can be determined from measured NDD as follows: TPR d, s, Q NDD d, s, f, Q d 2 f d S s f dref p S p dref s d

9 Normalized Depth Dose Data Energy Dependence Buildup region 15 MV TCPE region Surface region 6MV FS = 10 x 10 cm 2

10 Normalized Depth Dose Data Field Size Dependence This depth corresponds to range of the highest energy contaminant charged particles 15 MV Photon Beam 16x16 4x4

11 Normalized Depth Dose Data Wedge/Open Comparison FS=10x10cm2 15 MV (W/O) 6 MV (W/O)

12 Normalized Depth Dose Data Wedge/Open Comparison Minima

13 Normalized Depth Dose Data Field sizes: 6x6, 10x10 and 20x20 cm 2 6 MV - Siemens -- Varian.Elekta 18 MV These data from Radiological Physics Center show that all NDD for both 6 and 18 MV photon beams at depths of 5 cm and 15 cm for different field sizes have a maximum %σ of 0.5% and this increases to 0.7% at a depth of 20 cm.

14 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

15 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 higher energies Sheikh-Bagheri & Rogers, Med. Phys., 29, 2002

16 Beam Penetration for Irregularly- Shaped Fields Concept of Equivalent Square: The equivalent field is defined as that standard (square or circular) field which has the same centralaxis depth dose characteristics as the given non-standard field. s f d Day s Rule : S r r r S 1 e r e Water S(r) = the central axis scatter in a field of radius r, S = the central axis scatter in field of infinite it radius, λ is a scaling parameter, and μ is a dimensionless i 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.

17 L Equivalent square d W s Sterling Formula: (Sterling et.al., Brit. J. Radiol. 37, 544 (1964)) LW S 2 4 A L W / P Assuming, λ = 0.26 cm-1., and μ = 0.5 L/2 W/2 SLW (, ) 4 Dxydxdy (, ) 0 0 L / W SLW (, ) / S (10,10)

18 KLEIN- NISHINA CROSS SECTION FOR THE COMPTON INTERACTION d e d Ψ 2 2 r0 h ' h h ' sin 2 h h ' h 2 PHOTONS SCATTERED AT AN ANGLE, Ψ PHOTONS SCATTERED INTO A UNIT SOLID ANGLE, Ω SOLID ANGLE AVAILABLE PER UNIT ANGLE d d 2 sin Based on the kinematics of Compton interaction, the average energy of scattered photons is less than 1Mev and is independent of the incident energy.

19 Measurement of Normalized Depth Dose 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. Diodes and diamond d detectors t are appropriate as long as data measured with these detectors is scoss cross-referenced e e ced to data measured ed with an ion chamber. Prone to radiation damage and non-linear response.

20 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

21 Beam Output f f S c S c,p 10 cm c c S p s S c, p S c s c (Derived) Water

22 In-air output Ratio Elekta: 4-18 MV clinical photon beams.

23 Monte Carlo Calculations of In- Air Output t Ratio (BEAMnrc code) In-Air Output Ratio O o Simulation Geometry (Varian 2100EX) MV measured 6 MV calculated lated 18 MV measured 18 MV calculated Side of square field /cm /tex/rof/clxyro

24 Energy spectrum of head scattered photons Mean Energy:0.5 MeV (Varian 2100C.)

25 Energy spectrum of head scattered photons (Varian 2100C.) Mean Energy:0.5 MeV

26 In-air output Ratio e: Elekta, s: Siemens, and v: Varian (for clinical photon beams ranging from 6-25 MV.

27 Monitor Back Scatter Machine MBS Publication Flattening Filter Beam Modifier (internal wedge) Monitor Chamber Varian Clinac % Kubo, Med. Phys. 16, 295 (1987) Lower Collimator Upper Collimator Therac % Hounsell, P.M.B. 43, 445 (1998) Beam Modifier (external wedge) Tertiary Collimator (Cerrobend Block or Varian MLC) 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

28 Measurement of In-Air Output t Ratios Mini phantom Water-equivalent materials. 4g/cm 2 diameter and 10g/cm 2 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. r 1 h TG 74 recommendations 1 2

29 Cross Beam Characteristics Affected by the radially symmetric conical high Z- material flattening 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 Cross beam flatness is defined as: D F 100 D max max D One flattening filter for each clinical photon beam results in a compromise of beam flatness characteristics of small and large fields. Flattening filters are designed d to give a gradually increasing i radial intensity. it 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 D min min

30 Cross Beam Profile 6 MV Photon Beam, Depth of 5.0 cm, Field size of 4x4, 10.4x10.4, and 21x21 cm 2. 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.

31 Cross Beam Profile 6 MV Photon Beam, 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

32 Effect of Electron Steering on Beam Flatness Symmetric Tilted Displaced

33 Effect of a Dipole Magnet on Exit Beam Energy Spread Radial Displacement Radial Divergence

34 Cross Beam Symmetry S 100 Area left Area right Arealeft Area right Dosimetry and beam steering system

35 Isodose Distribution 30 cm X 30 cm 18 MV X-ray beam

36 Isodose Distributions (20 X 20 Cm 2 ) 6 MV 18 MV Note contaminant electrons contribute to dose outside the field at shallow depths. The magnitude and extent t of dose outside the geometric edge of a field at shallow depths increases with beam energy.

37 Isodose Distributions (20 X 20 Cm 2, 18 MV) Note Contaminant electrons contribute to dose outside the field at shallow depths. 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.

38 Cross Beam Measurements What is the affect of detector size? Incorrect measurement of penumbra region Diameter Penumbra 20%~80% Diode CC04 CC13 0.8x0.8 mm 2 4 mm 6 mm 4.0 mm 6.1 mm 7.2 mm

39 Detector Size Effect on TPS Commissioning Treatment Planning System Commissioning Impact of detector size effect on dose distribution??? ib ti Yan G et. al., Med. Phys (35)., 2008

40 Extraction of True Profile

41 IMRT QA results: DTA 2%/2 mm CC13 CC04 Deconvolved

42 Measurement of Attenuation Factors 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 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

43 Measurement of Wedge Factors The WF is defined as the ratio of the dose rate at the reference depth for a wedged d field to that t 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 dependence of the WF is correctly accounted for by in-air output ratios (S c ) 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

44 Characterizing Clinical Photon Beams in 3DRTPS Ahnesjo et al., PMB 1999

45 Approaches to Dose Computation Algorithms Data measured in water and in air Parameterize water data Reconstitute water data Calculate inhomogeneity corrections to water data Correction Correction based methods Calculate dose directly based on beam and phantom configurations Model Model based methods Figure 8.9,The Modern Technology of Radiation Oncology; J. Van Dyk

46 Correction vs. Model Based Methods Correction Based Measured data used as basis for Dose Computation. Require measurements with buildup cap in air or in a mini-phantom. Require lots of data. Generating functions used to reduce size of data set for convenient clinical use (i.e. less storage space). Patient dose distribution obtained by first computing Dose in water from generating function, then correcting for tissue heterogeneity, patient contour, and beam modifiers. Model Based Measured data used to setup description of treatment beam. Require a parameter to estimate size of photon source at target. Require more time for tuning of model parameters. Patient dose distribution obtained by computing beam and beam transport (i.e. beam interactions in treatment head and in patient) directly.

47 Accuracy Goal in Dose Calculations Required accuracy (overall( treatment < 5%): Ahnesjo et al., PMB 1999

48 Characterizing Clinical Photon Beams in 3DRTPS MUST model the following features realistically: Finite size of source (& penumbra) 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 collimators Scatter outside field (related to extra-focal radiation) MLC, blocks, block tray Dynamic wedge, fixed wedge, compensators (beam hardening)

49 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

50 Characterizing Clinical Photon Caveats (continued): Beams in 3DRTPS 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

51 Benchmark Dataset (Developed under NIH initiative) iti A collaborative effort involving Sun Nuclear Associates; the contractor, t and consultants t from: the University it 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 This benchmark datasets will be sufficient for the TPS companies to compare the accuracy of their dose modeling for treatment delivery

52 Summary The dosimetric properties of a clinical photon beam are characterized by: Its ability to penetrate a tissue-like medium (water) its change 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.

53 Summary It is quite evident that all modern clinical linear accelerators (linacs) of a particular 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.