Sperduto_Khan's Treatment Planning in Radiation Oncology, 5e

438 SECTION III Treatment Planning: Physics and Dosimetric Principles

automatically removed and, in some cases, replaced with a treatment couch by the planning system. Also relevant are temporary contrast agents that can produce a CT num ber that mimic a higher density material within the body. Usually, the contrast agent is used to aid in the tissue seg mentation, and so only the additional step of providing a more realistic CT number in the segmented region is required to correct for the presence of the contrast agent. The spatial reliability of CT scanners is typically within 2%, which leads to dose uncertainties of ∼ 1%. 3 Other imaging modalities provide information that will aid in the location and delineation of structures, but is of less value in the calculation of dose. For example, the advent of cone-beam CT within the treatment room provides invaluable information regarding patient alignment. How ever, the scatter contained within the images makes accu rate determination of density difficult. Although magnetic resonance imaging (MRI) is often able to provide superior tissue contrast, the information in MRI is not strongly related to electron density. Furthermore, MRI images are more prone to artifacts during image formation, which will degrade the quality of the calculated dose distributions. In addition to electron density, it is also necessary to determine the tissue composition for more modern calcula tion algorithms. In convolution/superposition algorithms, fluence attenuation tables are typically computed using mass-attenuation coefficient data, which are somewhat weakly dependent on material. Often these coefficients are determined for each voxel by linearly interpolating between published results of two different materials (e.g., water and bone) based on the density assigned to the voxel. For both Monte Carlo (MC) and Boltzmann transport calculations, a full material assignment must be made to allow for accu rate cross-sectional determination of both photon and elec tron transport throughout the patient volume. Ideally, the size of the voxels in the treatment planning CT should be close to the dose grid resolution used for calculation. A CT volume set typically consists of 50 to 200 images with a voxel matrix dimension of 512 × 512 for each image. For a 50-cm field of view, this corresponds to a voxel size of ∼ 1 mm in the transverse direction. The longitudinal voxel size depends on the slice thickness, but is typically from 2 to 5 mm. In many planning sys tems, the CT slice thickness is chosen as the voxel size of the dose grid. For these systems, it may be appropriate to downsample the CT image set to 256 × 256. This makes the transverse resolution more closely matched to that of the longitudinal direction, with only a minor degrada tion in the image. Degrading the resolution further from 256 × 256 may result in an unacceptable loss of detail. BASIC RADIATION PHYSICS FOR PHOTON BEAM DOSE CALCULATION Here, we present an introduction to the important aspects of X-ray production and interaction to understand the

capabilities and limitations of model-based photon treat ment planning algorithms.

Megavoltage Photon Production Figure 20.1 displays a cross-sectional view of a linear accelerator treatment head, which consists of a high-­ density shielding material such as lead, tungsten, or a lead-tungsten alloy. It consists of an X-ray target, flatten ing filter, ion chamber, and a primary and movable colli mator. High-energy electrons are accelerated in the linac’s accelerating structure and impinge on the X-ray target. The production of Bremsstrahlung, or braking radiation, occurs when the high-energy electrons strike a tungsten target located in the head of the accelerator. The size of the focal spot of the electrons on the target is on the order of a few millimeters. 4 This finite size contributes to the penum bra or the blurring of the beam near the edges of the field. A primary collimator, fabricated from a tungsten alloy, defines the maximum field diameter that can be used for treatment. At megavoltage energies, Bremsstrahlung is directed primarily in the forward direction. In most conventional C-arm accelerators, to make the beam intensity more uniform, a conical filter positioned in the beam prefer entially absorbs the photon fluence along the central axis. The presence of the field-flattening filter alters the energy spectrum, since the beam passing through the thicker central part of the filter has a higher proportion of low energy photons absorbed by the filter. This may not be necessary for modern treatment deliveries where modu lation is used to vary the intensity of the beam. Indeed, many treatment units now have the option of removing the filter for these treatments (e.g., Varian TrueBeam, Palo Alto, CA; Elekta Versa HD, Atlanta, GA), or have removed the flattening filter entirely (e.g., Accuray, Inc. TomoHD, Sunnyvale, CA) when a uniform field is not needed. Compton Scatter Photons can inelastically scatter via three main processes: photoelectric absorption, incoherent (Compton) scattering with atomic electrons, and pair production in the nuclear or electron electromagnetic field. In the energy range used for radiation therapy, most interactions are Compton scat tering events, which are discussed in more detail here. Compton-scattered photons may originate in either the accelerator treatment head or the patient (or phan tom). Most of the scatter dose generated by the accelera tor head is produced within the primary collimator and the field-flattening filter. These scattered photons and electrons are sometimes referred to as “extrafocal radia tion” which may be added to the primary photon beam emitted from the source. As the collimator jaws open, more scattered radiation is allowed to leave the treatment head, which results in an increase in the machine output with field size. This effect is known as collimator scatter , 5

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