Dosimetric Characteristics of
Clinical Photon Beams
Jatinder R Palta PhD
University of Florida
Department of Radiation Oncology
Gainesville Florida
Gainesville,
Disclosures
Research development grants from Philips
Medical Systems, Elekta Oncology
Systems and Sun Nuclear Associates
Systems,
Associates.
NIH research award.
B kh d C
Bankhead
Coley
l research
h award.
d
Learning
g Objectives
j
Understanding dosimetric properties of
clinical photon beams.
Understanding
gp
physical
y
p
parameters that
affect dosimetric properties of clinical
photon beams.
p
Understand the need for accurate
characterization of clinical photon beams
in a treatment planning system.
Photon Beam Delivery Systems
Medical Linear
Accelerators:
S band Linear Accelerators
X band Linear Accelerators
 Accelerate electrons in p
pulses
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
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
l
energies.
i
A typical
i l
modern high-energy linac can
provide 2-3 photon energies.
Sources of radiation that determine dosimetric
characteristics of clinical photon beams
Direct Radiation (Focal
Radiation)
Source
Photon radiation generated
at the target
g that reaches
patient without any
intermediate interactions.
Indirect (headscatter)
Flattening filter
Indirect Radiation (Extrafocal Radiation):
Monitor Chamber
Collimator jaws
Electron
Contamination
Direct
MLC
Charged particle
contamination dose
Output radiation or
Incident radiation
Primary dose
Secondary
electrons
S tt d
Scatter
dose
Photon radiation with a
history of interaction/scattering
in the head of the treatment
unit with the flattening filter
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
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
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
NDD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 q
quality
y of the clinical
photon beam, and Dd and Ddref are dose at
depth d and dref respectively.
f
s
d
Water
TPR data can be determined from measured
NDD as follows:
  f  d   S p s dref 
TPR d , s d , Q   NDD d , s, f , Q   


S p s d 
 f  dref 
2
Normalized Depth Dose Data
Energy Dependence
B ild region
Buildup
i
15 MV
S f
Surface
region
i
6 MV
FS = 10 x 10 cm2
TCPE region
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
Normalized Depth Dose Data
Wedge/Open Comparison
FS = 10 x 10 cm2
15 MV (W/O)
6 MV (W/O)
Normalized Depth Dose Data
Wedge/Open Comparison
Minima
Normalized Depth Dose Data
- Siemens
-- Varian
. Elekta
18 MV
Field sizes:
6x6, 10x10
and 20x20
cm2
6 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
i
%
%σ off
0.5% and this
increases to 0.7%
at a depth of 20 cm
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
g
energies
g
at higher
Sheikh-Bagheri & Rogers, Med. Phys.,
29, 2002
Beam Penetration for IrregularlyShaped Fields
Concept
p of Equivalent
q
Square:
q
f
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
“Day’s Rule”:

S r   S  1  e
 r
   r e
 r

Water
S(r) = the central axis scatter in a field of radius r, S∞ = the central axis scatter in
fi ld off infinite
field
i fi it radius,
di
λ iis a scaling
li parameter,
t and
d μ iis a di
dimensionless
i l
shape
h
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.
d
L
Equivalent square
W
d
Sterling
g Formula:
(Sterling et.al., Brit. J. Radiol. 37, 544 (1964))
2 LW
S
 4A/ P
L W
Assuming, λ = 0.26 cm
cm-1.,
1., and μ = 0.5
S ( L, W )  4 
L /2
0
L /W

W /2
0
D( x, y )dxdy
1
2
S ( L, W ) / S (10,10) 1.000 0.993
3
4
5
0.982 0.969 0.958
s
KLEIN- NISHINA CROSS SECTION
FOR THE COMPTON INTERACTION
d e r0  h '   h h '

2


 sin  
 

2  h   h ' h
dΨ

2
2
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 Compton
p
interaction, the average
g
energy of scattered photons is less than 1Mev and is independent
of the incident energy.
Measurement of Normalized Depth
D
Dose
d
data
Follow AAPM TG Report
p # 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
Di d and
d di
diamond
dd
detectors
t t
are appropriate
i t
as long as data measured with these detectors
is
s cross-referenced
c oss e e e ced to data measured
easu ed with
t a
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
measurement.
 the energy spectrum of incident photons does not change with the
depth.
 fluence across the detector remains the same
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
S p s  
S c , p s 
S c c 
Water
(Derived)
In-air output Ratio
Elekta: 4-18 MV clinical photon beams.
Monte Carlo Calculations of InAi Output
Air
O t t Ratio
R ti
(BEAMnrc code)
In-Air Output Ratio
1.05
Oo
1.00
6 MV
6 MV
18 MV
18 MV
0.95
Simulation Geometry
measured
calc
calculated
lated
measured
calculated
0.90
(Varian 2100EX)
0
5
10
15
20
25
30
Side of square field /cm
/tex/rof/clxyro
35
40
45
Energy spectrum of head scattered photons
Mean Energy:0.5 MeV
(Varian 2100C.)
Energy spectrum of head scattered photons
(Varian 2100C.)
Mean Energy:0
Energy:0.5
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
Flattening Filter
Machine
MBS
Publication
Varian Clinac 1800
1-5%
Kubo, Med. Phys.
16, 295 (1987)
Therac 20
7.5%
Hounsell,, P.M.B.
43, 445 (1998)
Elekta SL15
<1%
Yu et.al. P.M.B.
((with 3 mm AL)) 41,
Monitor Chamber
Beam Modifier
(internal wedge)
Upper Collimator
Lower Collimator
Beam Modifier
(external wedge)
Tertiary Collimator
(Cerrobend Block
or Varian MLC)
1107(1996)
5%
(without Al)
Varian 600c/2100C
Varian 2100C
2-5%
Lam et. al. Med.
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
Measurement
t off In-Air
I Ai Output
O t t Ratios
R ti
• Mini phantom
p
– Water-equivalent materials.
– 4g/cm2 diameter and 10g/cm2 depth to maintain lateral
CPE and eliminate contaminant electron
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
TG 74 recommendations
1
2
Cross Beam Characteristics
 Affected by the radially symmetric conical high Zmaterial flattening
g 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
points.
 hardens the beam
 Cross beam flatness is defined as:
Dmax  Dmin
F  100 
Dmax  Dmin
 One flattening filter for each clinical photon beam results in a
compromise of beam flatness characteristics of small and large
fields.
 Fl
Flattening
tt i filt
filters are d
designed
i
d tto give
i a gradually
d ll iincreasing
i radial
di l iintensity.
t
it
This is referred to as “horns” on a cross-beam profile
 Cross beam profiles may not be radially symmetric due
p
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,, Field Size of 10.4x10.4 cm2,, Depths
p
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 Dipole
p
Magnet
g
on Exit
Beam
Energy Spread
Radial Displacement
Radial Divergence
Cross Beam Symmetry
S  100 
Area left  Area right
Area left  Area right
Dosimetry and beam
steering system
Isodose Distribution
30 cm X 30 cm
18 MV X-ray beam
Isodose Distributions
(20 X 20 Cm2)
6 MV
18 MV
Note contaminant electrons contribute to dose outside the field
at shallow
s a o depths.
dept s The
e magnitude
ag tude a
and
de
extent
te t o
of dose outs
outside
de
the geometric edge of a field at shallow depths increases with
beam energy.
Isodose Distributions
(20 X 20 Cm2, 18 MV)
Note Contaminant electrons contribute to dose outside the
p
The magnitude
g
and extent of dose
field at shallow depths.
outside the geometric edge of a field at shallow depths
increases even more in the presence of beam modifiers.
Cross Beam Measurements
Whatt iis th
Wh
the affect
ff t off
detector size?
Incorrect
measurement of
penumbra region
Diameter
Diode
CC04
CC13
0.8x0.8
mm2
4 mm
6 mm
6.1 mm
7.2 mm
Penumbra
4.0 mm
20%~80%
Detector Size Effect on TPS
Commissioning
Treatment Planning
System
Commissioning
Impact of
detector size
effect on dose
di t ib ti ???
distribution???
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
F t
Factors
for
f Beam
B
Modifiers
M difi
 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,
y , 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.)
req ire additional considerations
require
considerations.
 These devices not only attenuate the incident beam but they
introduce scatter radiation that increase the scatter to primary
ratio within the patient
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
att the
th reference
f
depth
d th for
f a wedged
d d field
fi ld tto th
thatt
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
p
of the WF is correctly
y
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
B
Beams
iin 3DRTPS
Ahnesjo et al
al., PMB 1999
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
Correction vs. Model Based Methods
Correction Based
Model Based
Measured data used as basis for
Dose Computation.
Measured data used to setup
description of treatment beam.
Require measurements with buildup
cap in air or in a mini-phantom.
Require a parameter to estimate size
of photon source at target.
Require lots of data. Generating
functions used to reduce size of
data set for convenient clinical use
(i.e. less storage space).
Require more time for tuning of
model p
parameters.
Patient dose distribution obtained by
first computing Dose in water from
generating function, then correcting
for tissue heterogeneity, patient
contour,
t
and
d beam
b
modifiers.
difi
Patient dose distribution obtained by
computing beam and beam transport
(i.e. beam interactions in treatment
head and in patient) directly.
Accuracy Goal in Dose Calculations
• Required
q
accuracy
y (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)
 Extra-focal
E t f
l radiation
di ti (primary
( i
collimator,
lli t flflattening
tt i filt
filter))
 Beam spectrum (& change in spectrum with position)
 Beam intensity variation across field (e.g., beam horns)
 Transmission through secondary collimators
 Scatter
S tt outside
t id fi
field
ld ((related
l t d tto extra-focal
t f
l radiation)
di ti )
 MLC, blocks, block tray
 Dynamic wedge
wedge, fixed wedge
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
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
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
(Developed
l
d under
d NIH iinitiative)
iti ti )
A collaborative effort involving Sun Nuclear Associates; the
contractor,
t t and
d consultants
lt t from:
f
the
th University
U i
it off Florida;
Fl id
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 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
y
 The dosimetric properties of a clinical photon
beam are characterized by:
 Its ability to penetrate a tissue-like medium (water)
g in dose output
p with field size
 its change
 Its cross beam behavior
 Its attenuation through modifying devices (e.g., wedge,
compensator etc.)
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
y
It is quite evident that all modern clinical
li
linear
accelerators
l t
(li
(linacs)) off a particular
ti l
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.