Estimation of Neutron Production from Accelerator Head Assembly...

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


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].


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


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.


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


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


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


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].


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)


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).


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


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.


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


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


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


0

20

40

60

80

100

-30 -20 -10 0 10 20 30

Rel

ativ

e D

ose

(%

)



(a)

0

20

40

60

80

100

-30 -20 -10 0 10 20 30

Rel

ativ

e D

ose

(%

)



(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


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)


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


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)


Figure 4.11: The neutron fluence spectra at iso-center for optimized accelerator headassembly.


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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