The term Grenz-ray therapy is used to describe treatment with beams of very soft (low-energy) x-rays produced at potentials below 20 kV. Because of the very low depth of penetration (Fig. 4.1, line a), such radiations are no longer used in radiation therapy.

A contact therapy or endocavitary machine operates at potentials of 40 to 50 kV and facilitates irradiation of accessible lesions at very short source (focal spot) to surface distances (SSD). The machine operates typically at a tube current of 2 mA. Applicators available with such machines can provide an SSD of 2.0 cm or less. A filter of 0.5- to 1.0-mm thick aluminum is usually interposed in the beam to absorb the very soft component of the energy spectrum.

Because of very short SSD and low voltage, the contact therapy beam produces a very rapidly decreasing depth dose¹ in tissue. For that reason, if the beam is incident on a patient, the skin surface is maximally irradiated but the underlying tissues are spared to an increasing degree with depth. The dose versus depth curve or simply the depth-dose curve of a typical contact therapy beam is shown in Figure 4.1, line b. It is readily seen that this quality of radiation is useful for tumors not deeper than 1 to 2 mm. The beam is almost completely absorbed with 2 cm of soft tissue. Endocavitary x-ray machines have been used in the treatment of superficial rectal cancers.

The term superficial therapy applies to treatment with x-rays produced at potentials ranging from 50 to 150 kV. Varying thicknesses of filtration (usually 1- to 6-mm aluminum) are added to harden the beam to a desired degree. As mentioned in Section 3.6, the degree of hardening or beam quality can be expressed as the half-value layer (HVL). The HVL is defined as the thickness of a specified material that, when introduced into the path of the beam, reduces the exposure rate by one-half. Typical HVLs used in the superficial range are 1.0- to 8.0-mm Al.

The superficial treatments are usually given with the help of applicators or cones attachable to the diaphragm of the machine. The SSD typically ranges between 15 and 20 cm. The machine is usually operated at a tube current of 5 to 8 mA.

As seen in Figure 4.1, line c, a superficial beam of the quality shown is useful for irradiating tumors confined to about 5-mm depth (˜90% depth dose). Beyond this depth, the dose dropoff is too severe to deliver adequate depth dose without considerable overdosing of the skin surface.

The term orthovoltage therapy, or deep therapy, is used to describe treatment with x-rays produced at potentials ranging from 150 to 500 kV. Most orthovoltage equipment is operated at 200 to 300 kV and 10 to 20 mA. Various filters have been designed to achieve HVLs between 1 and 4 mm Cu. An orthovoltage machine is shown in Figure 4.2.

Although cones can be used to collimate the beam into a desired size, a movable diaphragm, consisting of lead plates, permits a continuously adjustable field size. The SSD is usually set at 50 cm.

Figure 4.1, line d, shows a depth-dose curve for a moderately filtered orthovoltage beam. Although the actual depth-dose distribution would depend on many conditions such as kilovoltage, HVL, SSD, and field size, some generalizations can be made from this curve about the orthovoltage beam characteristics. The maximum dose occurs close to the skin surface, with 90% of that value occurring at a depth of about 2 cm. Thus, in a single field treatment, adequate dose cannot be delivered to a tumor beyond this depth. However, by increasing beam filtration or HVL and combining two or more beams directed at the tumor from different directions, a higher dose to deeper tumors is delivered. As will be discussed in further detail in Chapter 11, there are severe limitations to the use of orthovoltage beam in treating lesions deeper than 2 to 3 cm. The greatest limitation is the skin dose, which becomes prohibitively large when adequate doses are to be delivered to deep-seated tumors. In the early days of radiation therapy, when orthovoltage was the highest energy available, treatments were given until radiation tolerance of the skin was reached. Although methods were developed to use multiple beams and other techniques to keep the skin dose under tolerance limits, the problem of high skin dose remained an overriding concern in the orthovoltage era. With the availability of cobalt teletherapy, the skin-sparing properties of higher-energy radiation (Fig. 4.1, line e) became the major reason for the modern trend to megavoltage beams.

Although skin dose and depth-dose distribution have been presented here as two examples of the limitations posed by low-energy beams, there are other properties such as increased absorbed dose in bone and increased scattering that make orthovoltage beams unsuitable for the treatment of tumors behind bone.

X-ray therapy in the range of 500 to 1,000 kV has been designated as high-voltage therapy or supervoltage therapy. In a quest for higher-energy x-ray beams, considerable progress was made in the 1950s and 1960s toward developing higher-voltage machines. The major problem at that time was insulating the high-voltage transformer. It soon became apparent that conventional transformer systems were not suitable for producing potential much above 300 kVp. However, with the rapidly advancing technology of the times, new approaches to the design of high-energy machines were found. One of these machines is the resonant transformer, in which the voltage is stepped up in a very efficient manner.

X-ray beams of energy 1 MV or greater can be classified as megavoltage beams. Although the term strictly applies to the x-ray beams, the x-ray beams produced by radionuclides are also commonly included in this category if their energy is 1 MeV or greater. Examples of clinical megavoltage machines are accelerators such as Van de Graaff generator, linear accelerator, betatron and microtron, and teletherapy γ-ray units such as cobalt-60.

The magnetron is a device that produces microwaves. It functions as a high-power oscillator, generating microwave pulses of several microseconds’ duration and with a repetition rate of several hundred pulses per second. The frequency of the microwaves within each pulse is about 3,000 MHz.

The magnetron has a cylindrical construction, having a central cathode and an outer anode with resonant cavities machined out of a solid piece of copper (Fig. 4.7). The space between the cathode and the anode is evacuated. The cathode is heated by an inner filament and the electrons are generated by thermionic emission. A static magnetic field is applied perpendicular to the plane of the cross section of the cavities and a pulsed DC electric field is applied between the cathode and the anode. The electrons emitted from the cathode are accelerated toward the anode by the action of the pulsed DC electric field. Under the simultaneous influence of the magnetic
field, the electrons move in complex spirals toward the resonant cavities, radiating energy in the form of microwaves. The generated microwave pulses are led to the accelerator structure via the waveguide.

Typically, magnetrons operate at a 2-MW peak power output to power low-energy linacs (6 MV or less). Although most higher-energy linacs use klystrons, accelerators of energy as high as 25 MeV have been designed to use magnetrons of about 5 MW power.

The klystron is not a generator of microwaves but rather a microwave amplifier. It needs to be driven by a low-power microwave oscillator.

Figure 4.8 shows a cross-sectional drawing of an elementary two-cavity klystron. The electrons produced by the cathode are accelerated by a negative pulse of voltage into the first cavity, called the buncher cavity, which is energized by low-power microwaves. The microwaves set up an alternating electric field across the cavity. The velocity of the electrons is altered by the action of this electric field to a varying degree by a process known as velocity modulation. Some electrons are speeded up while others are slowed down and some are unaffected. This results in bunching of electrons as the velocity-modulated beam passes through a field-free space in the drift tube.

As the electron bunches arrive at the catcher cavity (Fig. 4.8), they induce charges on the ends of the cavity and thereby generate a retarding electric field. The electrons suffer deceleration, and by the principle of conservation of energy, the kinetic energy of electrons is converted into high-power microwaves.

Bremsstrahlung x-rays are produced when the electrons are incident on a target of a high-Z material such as tungsten. The target is water cooled, and it is thick enough to absorb the incident electrons. As a result of bremsstrahlung-type interactions (Section 3.4.A), the electron energy is converted into a spectrum of x-ray energies with maximum energy equal to the incident electron energy. The average photon energy of the beam is approximately one-third of the maximum energy.

As mentioned previously, the electron beam, as it exits the window of the accelerator tube, is a narrow pencil about 3 mm in diameter. In the electron mode of linac operation, this beam, instead of striking the target, is made to strike an electron scattering foil to spread the beam as well as get a uniform electron fluence across the treatment field. The scattering foil consists of a thin high-Z metallic foil (e.g., lead, tantalum). The thickness of the foil is such that most of the electrons are scattered instead of suffering bremsstrahlung. However, a small fraction of the total energy is still converted into bremsstrahlung and appears as x-ray contamination of the electron beam. Most systems also employ a secondary low-Z foil of variable thickness to flatten the electron beam. The low atomic number material is chosen to minimize additional brehmsstrahlung radiation produced in the beam.