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Chapter 4.
Estimation of Neutron Production fromAccelerator Head Assembly of 15 MV Medical
LINAC using FLUKA Simulations
For the production of a clinical 15 MV photon beam, the
design of accelerator head assembly has been optimized
using Monte Carlo based FLUKA code. The accelerator
head assembly consists of e- target, flattening filter,
primary collimator and an adjustable rectangular
secondary collimator. Accelerators used for medicalsecondary collimator. Accelerators used for medical
radiation therapy generate continuous energy gamma rays
called bremsstrahlung radiations by impinging high energy
electrons on high Z materials. The electron accelerators
operating above 10 MeV can result in the production of
neutrons, mainly due to photo nuclear reaction ( ,n)
induced by high energy photons in the accelerator head
materials. These neutrons contaminate the therapeuticp
beam and give a non negligible contribution to patient
dose. The gamma dose and neutron dose equivalent at the
patient plane were obtained at different field sizes and
maximum neutron dose equivalent observed near the
central axis of of 30 30 cm2.
88
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 89
4.1 Introduction
4.1.1 Importance and Objective
Mega electron voltage (MeV) bremsstrahlung produced by medical ac-
celerators are a common form of treatment modality for malignant tumors that
occur at depth below the skin surface. The linear accelerator (linac) is a primary
tool in external beam radiotherapy, which generates continuous energy gamma
rays called Bremsstrahlung radiations (BR) by impinging electrons on high Z
material (e− γ target). The clinically applicable photon beam is produced in an
e− γ target, flattened with a flattening filter, collimates in primary collimator and
beam shaping using secondary collimator. The main challenge while using such
clinical photon beam in tumor treatment is the application of high doses to the
tumorous body regions by simultaneous sparing of the healthy tissues. Dual
transmission ionization chambers are used for monitoring the photon radiation
beam output as well as the radial and transverse beam flatness. Also, this helps
to get an exact idea of dose delivered to patient.
Photon beams with energies higher than 10 MeV are preferred, if doses
should be delivered to larger depths (e.g. for the treatment of prostate cancer)
and to enhance the skin sparing. But a parasitic effect occurring is the production
of neutrons, mainly due to the photonuclear giant-dipole-resonance (GDR) reac-
tion (γ, n) induced by high energy photons in the accelerator head materials [1].
For conventional treatment techniques, the contamination is neglected for the
patient and only accounted for radiation protection. However, if precision radia-
tion treatments like intensity modulated radiation therapy (IMRT) are used, then
the leakage and neutron radiation increases, as these techniques require longer
beam-on times. It is predicted in the literature that the additional dose due to the
photoneutrons is proportional to the beam-on time [2]. The biological effective-
ness of neutrons is substantially higher than that of photons [3], therefore even
a small neutron dose will increase the risk for secondary cancer. Therefore, it
is necessary to minimize the contribution of neutrons while designing such ac-
celerator head assembly. In addition, knowledge of the energy spectrum of the
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 90
photo neutron contamination allows one to estimate the equivalent neutron dose
received by the patient, optimize the room shielding and even the use of better
energy-dependent quality factors to estimate neutron doses received by the med-
ical personnel working in and around the therapy facilities.
The requirement for the production of clinical photon beam using 15
MeV electron are that the photon beam have a spatially uniform fluence and a
well collimated in a reference plane that is perpendicular to the beam axis. Gen-
erally, this plane is defined at a depth of 10 cm in a water phantom. The surface
of water phantom is 100 cm away from the photon source i.e. surface–source dis-
tance (SSD) is 100 cm. When the required condition is met, the radiation beam
will produce a uniform dose distribution across the reference plane. The objec-
tive were met by designing a accelerator head assembly consisting of e− γ target,
primary collimator, secondary collimator (X and Y jaws) and flattening filter for
15 MeV medical linac. In addition, the neutron contamination in photon beam
has been estimated in terms of dose equivalent and energy spectra. For this work,
FLUKA simulations has been carried out to evaluate the photoneutron yield and
spectra produced through accelerator head assembly of 15 MeV medical linac as
a function of the radiation field sizes.
4.1.2 Literature Survey
Neutron leakage from radiotherapy accelerators has been investigated
as early as in 1951 [4]. Guidelines have been recommended regarding maximum
admissible contamination levels by the National Council on Radiation Protection
and Measurements [5]. The International Electrochemical Commission (IEC)
recommends certain limits for the neutron absorbed dose in the patient plan [6],
and the American Association of Physicists in Medicine (AAPM) has reported a
review on neutron measurement methodologies [7]. Fast neutron contamination
of high-energy bremsstrahlung beams in radiotherapy has been investigated in
previous publications [8].
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 91
Tabl
e4.
1:L
itera
ture
Surv
ey
No.
Aut
hor
J.na
me,
Ene
rgy
Met
hod
Res
ult
(yea
r),V
ol,p
p)[r
ef]
(MV
)1
Pric
eK
.W.
Med
.Phy
s.,
25ex
peri
men
tby
activ
atio
nte
chni
que
mea
sure
dfa
stan
dth
erm
al(1
978)
,5,2
85[1
]ne
utro
npr
ofile
s2
Che
nC
.C.
Med
.Phy
s.10
FLU
KA
sim
ulat
ion
and
expe
rim
entb
y–
(198
5),1
2,59
2[9
]pr
opor
tiona
lcou
nter
,neu
tron
bubb
lede
tect
or3
Para
des
L.
Rad
.Mea
s.,
18ex
peri
men
tby
CR
39tr
ack
dete
ctor
for2
00cG
y/tr
eatm
ento
fpho
ton
dose
the
neut
ron
dose
equi
vale
nt(1
999)
,31,
475
[10]
is2.
3m
Sv/tr
eatm
ent,
outs
ide
radi
atio
nfie
ldin
patie
ntpl
ane
4O
ngar
oC
.Ph
y.M
ed.B
iol,
15an
dsi
mul
atio
nby
MC
NP-
GN
and
expe
rim
ent
neut
ron
toga
mm
ado
sefo
und
(200
0)45
,L55
[11]
18by
pass
ive
neut
ron
spec
trom
eter
BD
S1
mSv
/Gy
for1
5M
eVan
d4.
8m
Sv/G
yat
18M
eV5
Loi
G.
Phy.
Med
.Bio
l,12
MeV
expe
rim
entb
ypa
ssiv
ebu
bble
dete
ctor
mea
sure
dne
utro
ndo
seeq
uiva
lent
fore
lect
ron
mod
em
achi
ne(2
006)
,51,
695
[12]
6Se
rran
oB
.R
ad.P
rot.
Dos
.,25
sim
ulat
ion
byM
CN
P,PE
NE
LO
PE,
calc
ulat
edan
dm
easu
red
dose
profi
les
(200
6),1
19,5
06[1
3]ex
peri
men
tby
ioni
zatio
nch
ambe
rat
vari
ous
field
size
s7
Gol
nik
N.
Rad
.Pro
t.D
os.,
15ex
peri
men
tby
para
llel
mea
sure
dne
utro
ndo
seeq
uiva
lent
(200
7),1
26,6
19[1
4]pl
ate
cham
ber
and
foun
d35
mSv
/Gy
at6
Gy/
min
phot
ondo
se8
Al-
Gha
mdi
H.
Rad
.Mea
s.,
18ex
peri
men
tby
CR
-39
fast
and
ther
mal
neut
ron
(200
8),4
3,S4
95[1
5]nu
clea
rtra
ckde
tect
orre
lativ
ein
tens
ity9
Esp
osito
A.
Rad
.Mea
s.18
expe
rim
entb
yac
tivat
ion
mea
sure
dam
bien
tdos
eeq
uiva
lent
(200
8),4
3,10
38[1
6]an
dT
LD
pair
sto
phot
onab
sorb
eddo
se
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 92
Various researchers have calculated and measured the neutron dose equivalent
per gamma dose using different medical linac facilities at different electron ener-
gies. There work has been cited and a brief view has been tabulated in Table 4.1.
In the present work the design of e− γ target, primary and secondary collimator
and flattening filter have been optimized for the production of clinical photon
beam. Also, the neutron contamination produced through this accelerator head
assembly has been estimated using FLUKA simulation.
4.2 Structure of the Linear Accelerator Head
Clinical photon beams emanating from a medical linac are produced
in the e− γ target, flattened with a flattening filter and collimates in collimators.
The components of accelerator head assembly is shown in Figure 4.1(a). The
Photon collimation and beam shaping is achieved with primary collimator and
movable secondary collimator. The beam shaping using primary and secondary
collimator is shown in Figure 4.1(b).
4.2.1 e – γ target
Production of bremsstrahlung beam begins at the e− γ target. An elec-
tron beam from the electron gun is accelerated to a very high speed in the linac
waveguide. This electron beam is steered by the 270° magnet to hit the surface
of e− γ target perpendicularly. As the electron beam penetrated the target ma-
terial, the Coulomb interactions between the electron beam, atomic electrons in
the target and the protons in the nuclei of the target material occur. The Coulomb
interactions result in production of photon beam, is called as a Bremsstrahlung
photon, detail mechanism is discussed in Chapter 3. The bremsstrahlung yield
depends on the electron energy and the atomic number of the target element.
Therefore, high atomic number elements are used as e− γ target. The thickness
of the e− γ target was optimized such that all the electron incident on the target
get absorbed in target itself. Therefore, the target thickness was optimized more
than the range of 15 MeV electron in e− γ target.
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 93
Electron beam
E- target
Primary
Collimator
Forward peak
X-ray beam
Ion chamber
Flattening Filter
Secondary
Collimator X-Jaws
Secondary
Collimator
Y-Jaws
Flattened X-ray beam
Patient Plane at iso-center
(a)
Electron beam e- target
Primary
CollimatorCollimator
+ shielding
Flattening
FilterIonization
Chamber
Upper (X)
Jaws of
secondaryLower (Y)
Chamber
secondary
collimatorJaws of
secondary
collimator
X-ray field
at 1 m from
target
(b)
Figure 4.1: (a)The schematic of various components of accelerator head assembly and(b) The beam shaping using primary and secondary collimator (Not to the scale).
4.2.2 Primary Collimator
Immediately after e− γ target is the primary collimator. It is designed
to absorb all unwanted sections of the X-ray field. The circular conical hole
also defines the maximum divergence of the beam and therefore the maximum
circular field size. For obtaining maximum field of 51 cm diameter at patient
plane, the primary conical collimator comprises a 28° cone bored in a metal
block. The axis of the conical hole defines the beam axis and passes through
the center of the source of radiation. The sides of the conical opening projecting
on to edges of the target on one end of the block and on to the flattening filter
on the other end. The distance between the e− γ target and primary collimator
is optimized such that the system must produce the minimum penumbra. The
thickness of the shielding block is usually designed to attenuate the mean primary
X-ray beam intensity to be less than 0.1% of the initial value (three tenth-value
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 94
layers (TVLs)). According to International Electrotechnical Commission (IEC)
recommendations, the maximum leakage should not exceed 0.2% of the open
beam value. High Z element is often used for this component because of high
attenuation coefficient.
4.2.3 Flattening Filter
The photon beam exiting from the primary collimator does not have
uniform spatial intensity. It has an angular distribution that is strongly peaked in
forward direction with respect to he initial electron beam [17]. A more uniform
angular distribution of the photon beam can be achieved by passing it through a
flattening filter. The filter is designed differentially to absorb the radiation and
reduce the dose rate at the beam center. The flattening filter is like of Gaussian
shaped. The dose distribution is very sensitive to the position of the flattening
filter. A small misalignment of the flattening filter within a few millimeters in the
linac head would cause large variations in the dose distribution [18].
The flattening filter also has another effect on the beam called beam
hardening [19]. At high photon energies the flattening filter either increases or
decreases the variation of the effective photon energy across the beam. Conse-
quently, the depth of dose maximum and the penetration of the beam increases
or decreases. The former problem is due to significant beam hardening in the
forward direction when using low atomic number flattening filters, whereas a
significant spectral degradation is obtained with high atomic number filters due
to their preferential absorption of high energy photons. The flattening filter not
only hardens the beam as a whole, but further enhances the relative hardness near
the center.
There are several definitions of photon field flatness in the literature.
The IEC [20, 21] recommends that for square fields larger than 30 cm × 30 cm
the flattened area is 3 cm out from the field edge along the major axes, 6 cm from
the corners along the diagonal. These dimensions are presumably for the plane
containing the isocenter. The ratio of the maximum to minimum dose within the
flattened area should not exceed 1.10. The American Association of Physicists in
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 95
Medicine (AAPM) [22] defines flatness as the difference between the maximum
and minimum doses in the flattened area divided by the sum of the maximum and
minimum doses multiplied by 100, with the flattened region defined as 80% of
the profile width.
The flattening filter in this simulation was represented as a series of
eight truncated right circular cylinder of various thickness and of increasing radii
constructed by iron slabs. The 3 mm thick aluminum plate to which the filter is
attached was also included in the simulation.
4.2.4 Monitoring System and Mirror
Following the flattening filter are the monitoring system and the mirror.
The monitoring system consists of four quadrants of ion chambers which are
fixed in the beam direction. The ion chambers measure the radiation beam dose
output in terms of Monitor Units (MU) and the radial and transverse symmetry
of the radiation beam. The monitor chamber system also has a mechanism to
provide feedback to disable the linac from beaming due to the lack of symmetry
and or when the required number of MU is delivered.
Mirror is used to project light from the optical source to replicate the
shape of the radiation field. The angle and position of the mirror and the light
source are carefully aligned, so that the light field is coincident with the radiation
field. These components are designed so that they have minimal effect on the
radiation beam.
4.2.5 Secondary Collimator
Below the mirror are two sets of jaws which constitute a secondary
collimator. The maximum circular field defined by primary collimator are trun-
cated with an adjustable rectangular collimator which consists of upper and lower
independent movable jaws for producing rectangular and square fields with a
maximum dimension of 40 × 40 cm2 at the linac iso-centre. These blocks have
sufficient thickness to shield out unwanted radiations. The upper jaws move in
the in-plane direction and the lower jaws move in the cross-plane direction. The
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 96
collimating face of each jaw moves in such a way that it always lies along the
direction of propagation of the radiation, i.e. along a radius from the source.
These jaws are designed to move in an arc shape to account for the divergence
of the photon beam. This is not a concern for modeling because the jaws are
static in simulation. For each field size using different positions of jaws on the
arc simulation were carried out. The IEC recommends that the transmission of
the primary photon beam through the rectangular collimator should not exceed
2% of the open beam value. The material and dimension of collimators were
optimized such that the neutron contamination in the gamma beam was below
the allowed limit.
There are other components such as Multi-leaf collimators (MLC) and
Wedges which can also be mounted on the accelerator head assembly.
4.3 Neutron production through photonuclear reactions
When high energy electron interacts with high Z target, it generates a
cascade shower of bremsstrahlung radiations. The radiations produced in this
way, get absorbed in the material. If the absorbed photon has energy greater
than the binding energy of the neutron to the material, then neutron is emitted.
The photo neutron production threshold energy is varying in general from 8 to
19 MeV for light nuclei (A < 40) and from 6 to 8 MeV for heavy nuclei. The
exceptions are Eth = 2.23 MeV for deuterium and 1.67 MeV for beryllium [5].
The Giant Dipole Resonance (GDR) neutrons are produced by photons with en-
ergies from threshold energy to 30 MeV. In the GDR process, the electric field of
the photon transfers its energy to the nucleus by inducing an oscillation in which
the protons as a group move in opposite direction to the neutrons as a group.
The GDR neutron yield is proportional to the product of the length l of the ma-
terial traversed by photons of each energy (i.e the photon track-length) and the
GDR photoneutron cross section. The dependence of the photon track-length on
the photon energy k is expressed as the differential photon track length dl/dk,
representing the total track-length of all photons with energy in file interval k,
k + dk [10].
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 97
The GDR neutron yields are calculated by integrating, over the photon
energy spectra generated by electrons, the product of the differential photon track
length and the published GDR photoneutron cross sections. The GDR neutron
yield per incident electron can be determined analytically for each photoneutron
reaction using [10]:
YGDR =6.023 × 10−4ρ f Nn
AE0
∫ Emax
Eth
σGDR(k)(
dldk
)dk (4.1)
where YGDR = GDR neutron yield (neutron-MeV−1/electron), ρ = density of target
(g-cm−3), f = isotope fractional abundance, Nn = numbers of neutrons produced
per photoneutron reaction, A = atomic weight (g mol−1), E0 = electron energy
(MeV), σGDR(k) = photoneutron cross section (mb), dl/dk = differential photon
track length (cm MeV−1), k = photon energy (MeV), Eth = threshold energy of
the reaction (MeV), Emax = upper energy limit of the reaction or electron energy
when upper energy limit of the reaction is larger than the electron energy (MeV).
Neutrons generated in accelerator through photon induced GDR reac-
tion can be classified in two groups: the first has a Maxwellian energy distribution
and are called evaporation neutrons. The second are direct neutrons, which are
produced through direct interaction between the photon and neutron in the nu-
cleus of the target atom. The direct neutrons are, approximately, 15% of the total
produced by the (γ,n) reactions and their energy are greater than the evaporation
neutrons [23]. The angular distribution of the direct neutrons is assumed to be
in the form of 1 + Csin2θ, where θ is the angle between the incoming photon
and the emitted neutron and C is a constant dependent on neutron energy and the
media isotope [24, 25]. For the evaporation neutrons with energies < 2.5 MeV,
the emissions are assumed to be isotropic, i.e., C = 0. The evaporation neutrons
dominate at low neutron energies (< 1 – 2 MeV), and direct neutrons dominate
at high energies (> 2 MeV) [11]. Neutron spectra of evaporation neutrons is
described by [26],dNdEn
=En
T 2 exp(−En
T
)(4.2)
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 98
where En is the neutron energy in MeV and T is the nuclear temperature (in
MeV) to a particular nucleus. For instance, the corresponding temperature for
the production of neutrons in tungsten is 0.5 MeV [23].
4.4 Results
Firstly, to optimize the e− γ target the results discussed in Chapter 3 are
referred. Amongst the materials studied, tungsten was found to be best suitable
as e − γ target because of its physical properties like melting point, heat conduc-
tivity and highest bremsstrahlung yield. The 0.42 cm thick tungsten has been
optimized as an e− γ target for 15 MV medical LINAC since it absorbs almost
all the incident electrons. The bremsstrahlung spectrum estimated on the target
surface and collimator incident face is shown in Figure 4.2. It is observed from
figure that the bremsstrahlung spectra have peak at 0.5 MeV energy and have
continuous energy spectrum upto the incident electron energy. The mean energy
of the bremsstrahlung spectrum is around 2.033 MeV. The FLUKA simulations
were performed to find out the tenth value layer (TVL) thickness in various ma-
terials for 15 MeV electron beam generated bremsstrahlung spectrum. For the
1e-06
1e-05
0.0001
0.001
0.01
0.1
1
0 2 4 6 8 10 12 14 16
Bre
mss
trah
lung F
luen
ce (
(photo
n-M
eV-1
-cm
-2) /e
_)
Bremsstrahlung Energy (MeV)
calculated at 3 cm from the target
calculated on front surface of the target
Figure 4.2: Bremsstrahlung spectra at e− γ target and collimator surface (at 3 cm).
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 99
simulation, a 15 MeV electron beam is incident on the e− γ target, the generated
bremsstrahlung radiations were allowed to fall on the material of which TVL
thickness has to be determined. Different input files were prepared for varying
thicknesses of Iron, Lead, Tungsten, Bismuth, Tungsten+copper, Tantalum mate-
rials. The incident photon fluence on the material and transmitted photon fluence
beyond the material was estimated using FLUKA code for each run. The rela-
tive transmission of photon fluence with thickness of material for all the element
is shown in Figure 4.3. From the figure TVL values for Iron, Lead, Tungsten,
Bismuth, Tungsten+copper, Tantalum have been estimated and they are 7.64,
3.87, 2.93, 4.40, 3.37 and 3.30 cm respectively. The neutron fluence produced
0
0.1
0.2
0.3
0.4
0.5
0.6
0.7
0.8
0.9
0 2 4 6 8 10 12 14 16 18 20
Rel
ativ
e B
rem
stra
hlu
ng F
luen
ce
Material thickness cm
BiFePbTaW
W-Cu
Figure 4.3: Relative transmission of photon fluence with respect to thickness of mate-rial.
through photonuclear reaction in these materials are given in Figure 4.4. The ma-
terial having less TVL thickness and low neutron production is the best material
to be used for primary collimator. Therefore, the W–Cu material has been opti-
mized to design the primary collimator. To attenuate the bremsstrahlung beam
intensity to less than 0.1% of initial value, W–Cu thickness has been optimized to
more than three TVL thickness. The simulations were carried out using a block
of W–Cu having conical opening of 28° for the calculation of bremsstrahlung
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 100
0
2e-08
4e-08
6e-08
8e-08
1e-07
1.2e-07
1.4e-07
1.6e-07
0 2 4 6 8 10 12 14 16 18 20
Neu
tron F
luen
ce (
(neu
tron-c
m-2
)/e_
)
Material thickness cm
BiFePbTaW
W-Cu
Figure 4.4: Variation in neutron fluence as a function of material thickness.
radiations at iso-center and leakage radiation. The iso-center is defined at 100
cm from the source to surface distance(SSD). The bremsstrahlung fluence profile
at iso-center for different thickness of primary collimator is shown in Figure 4.5.
It is observed from the figure that there is almost same bremsstrahlung fluence
1.0e-5
1.0e-4
0 5 10 15 20 25 30
Bre
mss
trah
lung F
luen
ce (
(photo
n-M
eV-1
-cm
-2) /e
_)
Distance from beam axis (cm)
thickness = 8 cmthickness =10 cmthickness =12 cmthickness =14 cm
Figure 4.5: Bremsstrahlung fluence profile at iso-center for different thickness of pri-mary collimator.
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 101
for different thickness of primary collimator in beam area of 25.5 cm radius. The
increase in thickness of primary collimator decreases the penumbra. In addition,
the leakage radiations were calculated offside at a distance of 1 m from the beam
center and found to be less than 0.2% of beam value (recommended by IEC) for
the thickness more than 8 cm of primary collimator. Therefore, it was optimized
to use 10 cm thickness of W–Cu for primary collimator. The neutron fluence
calculated at iso-center is 3.94 ×10−9 n−cm−2/e−
0
1
2
3
4
5
-5 -4 -3 -2 -1 0 1 2 3 4 5
Hei
ght
(cm
)
Radius (cm)
(a) (b)
0
0.05
0.1
0.15
0.2
0.25
0.3
0.35
-30 -20 -10 0 10 20 30
Abs
orbe
d do
se (
Gy-
min
-1-µ
A-1
)
Off axis distance (cm)
(c)
Figure 4.6: (a)Dimension of Gaussian shaped flattening filter.(b) 3-D drawing of flat-tening filter (c) Flattened dose profile due to optimized flattening filter for 15 MV
LINAC.
The unflattened absorbed dose distribution can be modeled as Gaussian
along a plane transversal to the beam axis. Therefore, a Gaussian-shaped filter
may reflect a smoothly increased attenuation towards the central beam axis. A
Gaussian shaped filter was divided in eight truncated right angle cone (TRC) as
shown in Figure 4.6(a). For different values of the height, base radius and top
radius of the each TRC, the FLUKA simulations were carried out to obtain ab-
sorbed dose in water phantom at SSD = 100 cm. The beam profile was a key
parameter for the design of a flattening filter. As seen from Figure 4.4, the less
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 102
number of neutron are produced from iron material as compared to lead material,
therefore, iron has been used as a filter. The dimensions of optimized flatten-
ing filter made of eight TRC’s is shown in Figure 4.6(a). The 3D drawing of
the optimized Iron flattering filter is shown in Figure 4.6(b) which is plotted in
Simplegeo 4.2 [27], the flattened dose estimated in water phantom is shown in
Figure 4.6(c) and it gives flattened dose for 40 × 40 cm2 field size.
The jaws of secondary collimator was positioned such that the rotation
of respective X-jaws and Y-jaws forms square field size. The thickness of sec-
ondary collimator was optimized such that the transmission of the primary X-ray
beam should not exceed 2% of the open beam value. Therefore, the thickness of
the secondary collimator was optimized to 8 cm. Using the optimized value of
e− γ target, primary collimator, filter and secondary collimator, the structure was
modeled in FLUKA as shown in Figure 4.7. Using the trigonometry, the position
of the X and Y jaws of secondary collimator has been calculated for different
field sizes.
e- target
Secondary
Collimator
(X Jaws)
Secondary
Collimator
(Y Jaws)
Lead Shielding
Primary
Collimator
Iron Shielding
Lead Shielding
Water
Phantom at
100cm SSD
Movement in
Y direction
Movement in
X direction
Z
XY
Iron Flatt-
ening filter
Beam
line
Figure 4.7: Schematic of accelerator head assembly modeled in FLUKA for 15 MVmedical LINAC (for 0 × 0 cm2 field size)(not to the scale).
The rotation of X and Y jaws of secondary collimator along the arc
changes the radiation field size area from 0 × 0 to 40 × 40 cm2. Using cal-
culated positions of jaws for each field size, the accelerator head assembly has
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 103
30
40
50
60
70
80
90
100
0 5 10 15 20 25 30
Rel
ativ
e D
ose
(%
)
Depth in water (cm)
Field 10 X 10Field 20 X 20Field 30 X 30Field 40 X 40
Figure 4.8: Relative photon depth dose distribution for various field sizes at SSD of100 cm.
been modeled in FLUKA to estimate the gamma absorbed dose and neutron dose
equivalent. In addition, the neutron fluence and respective spectra for different
field sizes have also been estimated.
Figure 4.8 shows the relative photon depth dose distribution, for 10 ×
10 cm2, 20 × 20 cm2, 30 × 30 cm2, 40 × 40 cm2 field sizes at an SSD of 100 cm
using the optimized accelerator head geometry. The distance at which maximum
dose delivered in water is 2 cm. The Figure 4.9(a) and 4.9(b) shows the flattened
and unflattened dose profiles in water phantom for different field sizes.
The Table 4.2 shows that the bremsstrahlung fluence, current, ratio of
Table 4.2: Bremsstrahlung fluence, current, mean energy and dose at iso-center for theoptimized accelerator head geometry.
Field Size Fluence Current Fluence/Current Mean Max. Dose(cm2 × cm2) (photon−cm−2/e−) (Ratio) Energy (Gy−min−1-
× 10−5 (MeV) µA−1)10 × 10 7.216 7.211 1.0007 4.141 0.29320 × 20 8.051 8.028 1.0029 3.735 0.32330 × 30 8.652 8.596 1.0065 3.431 0.34240 × 40 8.974 8.875 1.0112 3.180 0.335
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 104
0
20
40
60
80
100
-30 -20 -10 0 10 20 30
Rel
ativ
e D
ose
(%
)
Off axis distance (cm)
Field 10 X 10Field 20 X 20Field 30 X 30Field 40 X 40
(a)
0
20
40
60
80
100
-30 -20 -10 0 10 20 30
Rel
ativ
e D
ose
(%
)
Off axis distance (cm)
Field 10 X 10Field 20 X 20Field 30 X 30Field 40 X 40
(b)
Figure 4.9: (a) Flattened and (b) Unflattened dose profile in water phantom for differ-ent field sizes.
fluence to current, mean energy of bremsstrahlung spectrum and maximum dose
delivered in water at iso-center. It is observed from the results that maximum the
field size more the dose delivered. However, in case of 40 × 40 cm2 field size
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 105
the dose observed to be reduced as compared to 30 × 30 cm2 field size because
of the effective area offered by the conical beam is less than the square field
size. The fluence to current ratio increases with field size which implies that the
sharpness of the beam decreases. (If the ratio is 1 then all particle coming par-
allel to each other and exactly perpendicular to detector face). The mean energy
of bremsstrahlung spectrum observed at center is almost one third of the initial
electron energy.
0
0.1
0.2
0.3
0.4
0.5
0.6
0.7
0.8
-30 -20 -10 0 10 20 30
Rat
io o
f neu
tron D
E t
o p
hoto
n d
ose
(m
Sv/G
y)
Off axis distance (cm)
0x0 cm
2
10x10 cm2
20x20 cm2
30x30 cm2
40x40 cm2
Figure 4.10: The ratio of neutron dose equivalent to central axis photon absorbed doseat patient plane for different field sizes.
The ratio of neutron dose equivalent to central axis photon absorbed
dose along the longitudinal axis at patient plane for different field sizes is shown
in Figure 4.10. The maximum neutron dose equivalent observed near the cen-
tral axis of 30 × 30 cm2 field size. This is 0.71% of the central axis photon
dose rate of 0.3 Gy/min at 1 µA electron beam current. The values of neutron
dose equivalent estimated are consistent with the results of other measurements
reported in literature [14] and fall within the allowed limit by International Elec-
trotechnical Commission (IEC). In addition, the ratio of neutron dose equivalent
to central axis photon dose was maintained below the allowed limit set by IEC
(< 1 mSv/Gy) inside and 0.5 mSv/Gy outside of photon field. The Table 4.3
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 106
shows the neuron fluence, current, ratio of fluence to current, mean energy of
neutron spectrum and thermal neutron percentage calculated in neutron spec-
trum at iso-center for the optimized accelerator head geometry. It is observed
that the neutron fluence increases with field size. The neutron fluence spectrum
calculated for different field sizes using FLUKA is shown in Figure 4.11.
Table 4.3: Neutron fluence, current, mean energy and percentage of thermal neutron atiso-center for the optimized accelerator head geometry.
Field Size Fluence Current Fluence/Current Mean Thermal neutron(cm2 × cm2) (neutron−cm−2/e−) (Ratio) Energy percentage
× 10−9 (keV) (%)10 × 10 2.452 2.413 1.0158 0.404 11.6020 × 20 3.408 3.612 1.0139 0.886 11.4630 × 30 3.612 3.538 1.0210 0.511 10.6640 × 40 3.518 3.437 1.0235 0.163 10.97
1e-13
1e-12
1e-11
1e-10
1e-09
0.001 0.01 0.1 1 10
Neu
tron F
luen
ce (
(neu
tron-c
m-2
) /e_)
Neutron Energy (MeV)
Field 10 X 10Field 20 X 20Field 30 X 30Field 40 X 40
Figure 4.11: The neutron fluence spectra at iso-center for optimized accelerator headassembly.
Chapter 4. Accelerator head assembly for 15 MeV gamma ray therapy 107
4.5 Conclusion
For the production of 15 MeV photon beam in clinical applications, the
design of accelerator head assembly has been proposed and optimized. Using the
optimized design, the flattened dose calculated at 100 cm SSD is 0.34 Gy/min at
1 µA for 30 × 30 cm2 field size. The maximum square field size can be produced
by the collimator is 30 × 30 cm2. In addition, the neutron produced in accelerator
head assembly has been estimated and the the ratio of neutron dose equivalent
to gamma dose is found below the allowed limit recommended by IEC i.e. < 1
mSv/Gy.
4.6 Future Scope
Neutron characterization around medical accelerators has been studied
extensively in this chapter. However, heavy particles such as neutron, proton and
alphas are also produced by photonuclear processes in the patient body. Allen
and Chaudhari [28] have calculated the photonuclear absorbed dose to be 0.094%
of the photon absorbed dose for a 24 MV photon beam. They also estimated that
24% of the absorbed dose due to photonuclear reactions could be attributed to
(γ, n) reactions and that the (γ,P)and (γ, α) processes give rise to 69% and 7%,
respectively. Therefore, to estimate the heavy particle dose relative to the photon
absorbed dose and equivalent dose in tissue due to photonuclear processes in the
patient body, further simulations can be carried out. The results will show the
exact analysis of neutron dose which will be received by patient.
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