Photon dosimetry 31 01-2022

Dose Measurement of Photon
Beam
Dr. Saira Siddique
MSc. RMO fellow
4/5/2022 Photon Dosimetry 1

Presentation Layout
• Introduction
• Phantom
• Depth Dose Distribution
• PDD
• Definition
• It’s Dependence
• TAR & BSF
• TMR & TPR

Introduction
It is seldom possible to measure dose distribution directly in patients
treated with radiation.
Data on dose distribution are almost entirely derived from measurements
in phantoms—tissue-equivalent materials, usually large enough in
volume to provide fullscatter conditions for the given beam.
These basic data are used in a dose calculation system devised to
predict dose distribution in an actual patient. Various dosimetric quantities
and methodologies have been devised to facilitate dose calculation in
clinical situations.

The methods based on quantities such as percent depth
doses (PDDs), tissue–air ratios (TARs), and scatter–air
ratios (SARs) have been used traditionally for dose
calculation involving low energy beams (up to 60Co),
which were usually calibrated in terms of exposure rate in
air or dose rate in free space.
Current methods of dose calculation use tissue–phantom
ratios (TPRs) or tissue–maximum ratios (TMRs), which
are better suited for higher energy beams and involve
measurements in phantom rather than in air.
Low energy – Exposure Rate/ Dose rate in FREE AIR  PDD, TAR, SAR
Higher energy – in phantom  TPR, TMR

TMT Various Phantom materials

Basic dose distribution data are usually
measured in a WATER PHANTOM b/c:
• it closely approximates the radiation absorption and scattering properties of
muscle and other soft tissues.
• water is universally available with reproducible radiation properties.
Ideally, for a given material to be tissue or water equivalent, it must have:
the same effective atomic number,
number of electrons per gram,
and mass density.
However, in MV photon beam (clinical range)  the most predominant mode
is Compton effect so the necessary condition for water equivalence for such
beams is to have the same electron density (number of electrons per cubic
centimeter) as that of water.

For clinical dosimetry 
anthropomorphic phantoms are
frequently used.
One such commercially available
system, known as Alderson Radiation
Therapy Phantom (formerly the
Alderson Rando Phantom),
it incorporates materials to simulate
various body tissues—muscle, bone,
lung, and air cavities.
It is shaped into a human torso and is
sectioned transversely into slices for
inserting film or other dosimeters

DEPTH DOSE DISTRIBUTION
As the beam is incident on a patient (or a phantom), the absorbed dose in the
patient varies with depth.
This variation depends on many conditions: beam energy, depth, field size,
distance from source, and beam collimation system.
Thus, the calculation of dose in a patient involves considerations in regard to these
parameters and others as they affect depth dose distribution.
An essential step in the dose calculation system is to establish depth dose variation
along the central axis of the beam.
A number of quantities, are defined for this purpose, major among these being
PDDs, TARs, TPRs, and TMRs.
These quantities are usually derived from measurements made in water phantoms
using small ionization chambers. Although other dosimetry systems such as
thermoluminescent dosimeters (TLDs), diodes, and film are occasionally used, ion
chambers are preferred because of their better precision and smaller energy
dependence.

PERCENTAGE DEPTH DOSE
One way of characterizing the central axis dose distribution is to
normalize dose at depth with respect to dose at a reference
depth.
The quantity percentage depth dose is defined as
“the quotient, expressed as a percentage, of the absorbed
dose at any depth d to the absorbed dose at a reference
depth d0 , along the central axis of the beam.”
• For orthovoltage (up to 400 kV) & lower-energy x-rays, the
reference depth is usually the surface (d0 = 0).
• For higher energies, the reference depth is usually taken at the
position of the peak absorbed dose (d0 = dmax)

PDD

PDD depends on:
• BEAM QUALITY AND DEPTH
• FIELD SIZE AND SHAPE
• SSD

PDD _ Dependence on Beam QUALITY & DEPTH
PDD (beyond the depth of maximum dose) decreases with
depth and increases with beam energy.
Higher-energy beams have greater penetrating power and
thus deliver a higher PDD at a given depth
the beam quality affects the PDD by virtue of the average
attenuation coefficient . As it decreases, the beam becomes
more penetrating, resulting in a higher PDD at any given
depth beyond the buildup region.

Initial Dose Buildup
PDD decreases with depth beyond the depth of maximum dose. However, there
is an initial buildup of dose that becomes more and more pronounced as the
energy is increased. In the case of the orthovoltage or lower-energy x-rays, the
dose builds up to a maximum on or very close to the surface. But for higher-
energy beams, the point of maximum dose lies deeper into the tissue or
phantom.
“The region between the surface and the point of maximum dose is called
the dose buildup region.”
dose buildup effect (higher-energy)  the skin-sparing effect.
For MV beams (cobalt-60 & higher energies), the surface dose is much smaller
than the Dmax . This offers a distinct advantage over the lower-energy beams
for which the Dmax occurs at or very close to the skin surface.
Thus, in MV beams, higher doses can be delivered to deep-seated tumors
without exceeding the tolerance of the skin. This, of course, is possible because
of both the higher PDD at the tumor and the lower surface dose at the skin.

PDD _ EFFECT OF FIELD SIZE AND SHAPE
Field size may be specified either
geometrically
dosimetrically.
The geometric field size: “the projection, on a plane
perpendicular to the beam axis, of the distal end of the
collimator as seen from the front center of the source”
The dosimetric, or the physical, field size is the distance
intercepted by a given isodose curve (usually 50% isodose)
on a plane perpendicular to the beam axis at a stated
distance from the source.

PDD _ EFFECT OF FIELD SIZE AND SHAPE
For a sufficiently small field, the depth dose at a point is primarily the result
of the primary radiation (the photons that have traversed the overlying
medium without interacting).
The contribution of the scattered photons to the depth dose in this case is
negligibly small and may be assumed to be 0.
But as the field size is increased, the contribution of the scattered
photons to the absorbed dose increases because the volume that can
scatter radiation gets larger with field size. Also, this increase in scattered
dose will be greater at greater depths.
The increase in PDD caused by increase in field size depends on beam
quality. Because the scattering probability or scatter cross section decreases
with increase in energy and the higher-energy photons are scattered more
predominantly in the forward direction, the field size dependence of PDD is
less pronounced for the higher-energy than for the lower-energy beams.

PDD _ DEPENDENCE ON SSD
Photon fluence emitted by a point source of radiation varies inversely as a
square of the distance from the source. The source is considered as a point
source (at large SSDs – 80cm). Thus, the exposure rate or “dose rate in free
space” from such a source varies inversely as the square of the distance
(assuming it as a Primary beam – no scatter).
PDD increases with SSD as a result of the inverse square law.
Although the actual dose rate at a point decreases with increase in distance
from the source, the PDD, which is a relative dose with respect to a
reference point, increases with SSD.

Relative dose rate from a point
source is plotted as a function of
distance from the source, following
the inverse square law.
The plot shows that:
the drop in dose rate
between two points is much
greater at smaller distances
from the source than at large
distances.
This means that the PDD, which
represents depth dose relative to a
reference point, decreases more
rapidly near the source than far away
from the source.

TISSUE–AIR RATIO
TAR may be defined as:
“the ratio of the dose at a given point in the phantom (Dd) to the
dose in free space at the same point (Dfs).”
For a given quality beam, TAR depends on
 depth d
 field size at that depth.
Because the PDD depends on the SSD, the SSD correction to the
PDD will have to be applied to correct for the varying SSD—a
cumbersome procedure to apply routinely in clinical practice. TAR—
has been defined to remove the SSD dependence.
concept of TAR has been refined to facilitate calculations not only for
rotation therapy but also for stationary isocentric techniques as well
as irregular fields. (specifically for rotation therapy calculation in past)

TAR
EFFECT OF DISTANCE:
The most important properties attributed to TAR is that it is
independent of the distance from the source. TAR for the primary
beam is only a function of depth, not of SSD.
VARIATION WITH ENERGY, DEPTH, AND FIELD SIZE
TAR varies with energy, depth, and field size very much like the PDD.
For the megavoltage beams, the TAR builds up to a maximum at the
depth of maximum dose (dm ) and then decreases with depth.

BSF
The term backscatter factor or peak scatter factor (PSF) is simply
the TAR at the reference depth of maximum dose on central axis
of the beam. It may be defined as:
“the ratio of the dose on central axis at the reference depth of
maximum dose to the dose at the same point in free space.”
The BSF, like the TAR, is independent of distance from the source
and depends only on the beam quality and field size.

TMR & TPR

TMR & TPR
TPR (Tissue Phantom Ratio) is defined as the ratio of the dose
rate at a given depth in phantom to the dose rate at the
same source-point distance, but at a reference depth.
If dmax is adopted as a fixed reference depth, the quantity TPR
gives rise to the TMR.
Thus, TMR is a special case of TPR
TMR (Tissue Maximum Ratio) is defined as the ratio of the
dose rate at a given point in phantom to the dose rate at the
same source-point distance and at the reference depth of
maximum dose.

THANK YOU!

Photon dosimetry 31 01-2022

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