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Review of Radiation Oncology Physics: A Handbook for Teachers and Students

CHAPTER 5.

ERVIN B. PODGORSAK

Department of Medical Physics

McGill University Health Centre

Montréal, Québec, Canada

5.1. INTRODUCTION

Since the inception of radiotherapy soon after the discovery of x-rays by Roentgen in 1895, the technology of x-ray production has first been aimed toward ever higher photon and electron beam energies and intensities, and more recently toward computerization and intensity-modulated beam delivery. During the first 50 years of radiotherapy, the techno-logical progress has been relatively slow and mainly based on x-ray tubes, Van de Graaff generators and betatrons.

The invention of the cobalt-60 teletherapy unit by H.E. Johns in Canada in the early 1950s provided a tremendous boost in the quest for higher photon energies, and placed the cobalt unit into the forefront of radiotherapy for a number of years. The concurrently developed medical linear accelerators (linacs), however, soon eclipsed the cobalt unit, moved through five increasingly sophisticated generations, and became the most widely used radiation source in modern radiotherapy. With its compact and efficient design, the linac offers excellent versatility for use in radiotherapy through isocentric mounting and provides either electron or megavoltage x-ray therapy with a wide range of energies.

In addition to linacs, electron and x-ray radiotherapy is also carried out with other types of accelerators, such as betatrons and microtrons. More exotic particles, such as protons, neutrons, heavy ions, and negative π mesons, all produced by special accelerators, are also sometimes used for radiotherapy; however, most of the contemporary radiotherapy is carried out with linacs or teletherapy cobalt units.

5.2. X-RAY BEAMS AND X-RAY UNITS

Clinical x-ray beams typically range in energy between 10 kVp and 50 MV, and

are produced when electrons with kinetic energies between 10 keV and 50 MeV

are decelerated in special metallic targets.

In the target, most of the electron's kinetic energy is transformed into heat and a

small fraction of the energy is emitted in the form of x-ray photons which are

divided into two groups: characteristic x-rays and bremsstrahlung x-rays.

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C hapter 5. Treatment Machines for External Beam Radiotherapy

5.2.1. Characteristic x-rays

Characteristic x-rays result from Coulomb interactions between the incident

electrons and atomic orbital electrons of the target material (collisional loss).

In a given Coulomb interaction between the incident electron and an orbital

electron, the orbital electron is ejected from its shell and the resulting orbital

vacancy is filled by an electron from a higher level shell. The energy difference

between the two shells may be emitted from the atom either in the form of a

characteristic photon (characteristic x-ray) or is transferred to an orbital electron

which is ejected from the atom as an Auger electron.

The fluorescent yield ω gives the number of fluorescent (characteristic) photons

emitted per vacancy in a shell (0≤ω≤1) and ranges from 0 for low Z atoms

through 0.5 for copper (Z = 29) to 0.96 for high Z atoms for K-shell vacancies that

are the most prominent sources of characteristic x-rays.

The photons emitted through electronic shell transitions have discrete energies

that are characteristic of the particular target atom in which the transitions have

occurred; hence the term characteristic radiation.

5.2.2. Bremsstrahlung (continuous) x-rays

Bremsstrahlung x-rays result from Coulomb interactions between the incident

electron and the nuclei of the target material.

During the Coulomb interaction between the incident electron and the nucleus, the

incident electron is decelerated and loses part of its kinetic energy in the form of

bremsstrahlung photons (radiative loss).

Photons with energies ranging from 0 to the kinetic energy of the incident electron

may be produced, resulting in a continuous bremsstrahlung spectrum.

The bremsstrahlung spectrum produced in a given x-ray target depends on the

kinetic energy of the incident electron as well as on the thickness and atomic

number Z of the target.

5.2.3. X-ray targets

In comparison with the range R of electrons of a given kinetic energy KE in the

target material, targets are divided into two main groups: thin and thick.

A thin target has a thickness much smaller than R, while the thickness of a thick

target is on the order of R.

For thin target radiation, the energy radiated is proportional to the product

(KE)×Z, where Z is the target atomic number. The intensity versus photon energy

(photon spectrum) is constant from 0 to KE, and 0 for all energies above KE.

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Review of Radiation Oncology Physics: A Handbook for Teachers and Students

A thick target may be considered as consisting of a large number of superimposed thin targets. The intensity I(hν) of a thick target spectrum is expressed as: I(hν)=CZ(KE hν) , where C hν is a proportionality constant and is the photon energy. (5.1) X-rays are used in diagnostic radiology for diagnosis of disease and in radiation oncology (radiotherapy) for treatment of disease. X-rays produced by electrons with kinetic energies between 10 keV and 100 keV

are called superficial x-rays, with electron kinetic energies between 100 keV and

500 keV orthovoltage x-rays, and with electron kinetic energies above 1 MeV

megavoltage x-rays.

Superficial and orthovoltage x-rays are produced with x-ray tubes (machines),

while megavoltage x-rays are most commonly produced with linacs and sometimes with betatrons and microtrons.

Typical thin and thick target bremsstrahlung spectra originating from 100 keV

electrons striking a thin and thick target, respectively, are shown in Fig. 5.1.

FIG. 5.1. Typical thin target (curve 1) and thick target (curves 2, 3, and 4) spectra for an x-ray tube in which 100 keV electrons strike the target. Curve (1) is for a thin target producing a constant intensity for photon energies from 0 to the kinetic energy of electrons striking the target (100 keV). Curve (2) represents unfiltered spectrum (inside the x-ray tube) for a thick target and represents a superposition of numerous thin target spectra; spectrum of curve (3) is for a beam filtered by x-ray tube window (low energy photons are filtered out); spectrum of curve (4) is for beam filtered by the x-ray tube window and additional filtration.

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C hapter 5. Treatment Machines for External Beam Radiotherapy

5.2.4. Clinical x-ray beams

A typical spectrum of a clinical x-ray beam consists of line spectra that are

characteristic of the target material and are superimposed onto the continuous

bremsstrahlung spectrum.

The bremsstrahlung spectrum originates in the x-ray target, while the

characteristic line spectra originate in the target and in any attenuators placed into

the beam.

The relative proportion of the number of characteristic photons to bremsstrahlung

photons in an x-ray beam spectrum varies with electron beam kinetic energy and

atomic number of the target. For example, x-ray beams produced in a tungsten

target by 100 keV electrons contain about 20% characteristic photons and 80%

bremsstrahlung photons, while in the megavoltage range the contribution of

characteristic photons to the total spectrum is negligible.

In the diagnostic energy range (10 to 150 kV) most photons are produced at 90°

from the direction of electron acceleration, while in the megavoltage energy range

(1 to 50 MV) most photons are produced in the direction of electron acceleration

(forward direction: 0°).

5.2.5. X-ray beam quality specifiers

Various parameters, such as photon spectrum, half-value layer, nominal accelerating potential, beam penetration into tissue-equivalent media, etc., are used as x-ray beam quality indices(ee Sections 9.8.1 and 9.8.2 for details):

Complete x-ray spectrum is very difficult to measure; however, it gives the most

rigorous description of beam quality.

Half-value layer (HVL) is practical for beam quality description in the superficial

(HVL in aluminum) and orthovoltage (HVL in copper) x-ray energy range, but not

practical in the megavoltage energy range because in this energy range the

attenuation coefficient is only a slowly varying function of beam energy .

The effective energy of a heterogeneous x-ray beam is defined as that energy of a

monoenergetic photon beam that yields the same HVL as does the heterogeneous

beam.

Nominal accelerating potential (NAP) is sometimes used for describing the

megavoltage beam quality. The NAP is determined by measuring the ionisation

ratio in a water phantom at depths of 10 and 20 cm for a 10×10 cm2 field at the

nominal source-axis distance of 100 cm.

Recent dosimetry protocols recommend the use of tissue-phantom ratios or

percentage depth doses at a depth of 10 cm in a water phantom as an indicator of

megavoltage beam effective energy (beam quality index).

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Review of Radiation Oncology Physics: A Handbook for Teachers and Students

5.2.6.

X-ray machines for radiotherapy Superficial and orthovoltage x-rays used in radiation therapy are produced with x-

ray machines. The main components of a radiotherapeutic x-ray machine are: an

x-ray tube; ceiling or floor mount for the x-ray tube; target cooling system;

control console; and an x-ray power generator. A schematic diagram of a typical

therapy x-ray tube is shown in Fig. 5.2.

The electrons producing the x-ray beams in the x-ray tube (Coolidge tube) originate in the heated filament (cathode) and are accelerated in vacuum toward

the target (anode) by an essentially constant-potential electrostatic field supplied

by the x-ray generator.

The efficiency for x-ray production in the superficial and orthovoltage energy

range is on the order of 1% or less. Most of the electron kinetic energy deposited

in the x-ray target (≈99%) is transformed into heat and must be dissipated

through an efficient target cooling system.

To maximize the x-ray yield in the superficial and orthovoltage energy range the

target material should have a high atomic number Z and a high melting point.

With x-ray tubes, the patient dose is delivered using a timer and the treatment

time must incorporate the shutter correction time (see Section 6.16) that accounts

for the time required for the power supply components to attain the steady state

operating conditions.

The x-ray tube current is controlled by hot filament emission of electrons which,

in turn, is controlled by the filament temperature (thermionic emission). For a

given filament temperature the x-ray tube current increases with the tube (anode)

voltage, first rising linearly with voltage in the space charge limited region and

saturating at higher voltages when all electrons emitted from the cathode are

pulled to the anode.

Research is currently carried out on cold field emission cathodes produced with

carbon nanotubes (CNT). The CNT-based cold cathode x-ray technology may

lead to more durable as well as miniature and portable x-ray sources for industrial

and medical applications.

FIG. 5.2. Schematic diagram of a typical therapy x-ray tube (Reprinted from Johns, H.E. and

Cunningham, J.R. with permission).

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