Particle Therapy in Non-Small Cell Lung Cancer

Zhongxing Liao; Charles B. Simone II

Disclosures

Transl Lung Cancer Res. 2018;7(2) 

In This Article

Dosimetry and Radiobiology of Charged Particle Therapy

The radiobiology and dosimetric characteristics of charged particle therapy have been reviewed in depth elsewhere.[7] Briefly, charged particle radiotherapy involves the use of charged particles such as protons or carbon ions to treat cancer. The depth dose characteristics of charged particles are well understood and described elsewhere.[8] When a "fast" charged particle moves through matter, it interacts with the electrons within atoms and causes ionizations, which deposit energy and dose along its path. The energy loss per unit path length is relatively constant until it reaches a peak (the so-called Bragg peak), where energy deposition occurs at a depth that is a function of the energy and nature of the charged particle. Beyond that Bragg peak, very little dose remains. In passively scattering proton therapy (PSPT), the Bragg peak is spread both longitudinally and laterally to create a spread-out Bragg peak (SOBP), which provides a uniform dose to cover the entire volume of a target. Conformal coverage of the tumor is achieved by using range modulation wheels, compensators, and beam apertures. Pencil beam scanning proton therapy, on the other hand, uses magnetic scanning of thin beamlets of protons of a sequence of energies, delivered from different directions, to produce the desired pattern of dose distribution. The tumor is "scanned" layer by layer, with one layer per energy, until the entire target is covered. This technique provides greater flexibility and control for ideal dose distributions and allows delivery of intensity-modulated proton therapy (IMPT), to date the most advanced form of proton therapy.[7] Many treatment-planning comparison studies have demonstrated dosimetric advantages of IMPT over intensity-modulated photon radiation therapy (IMRT).[9,10]

The biological interaction of ionizing radiation with matter (i.e., tissues) is related to the amount of energy transferred to the matter over a specified path length [known as linear energy transfer (LET)]. For particles such as protons and helium, the LET is thought to be nearly equivalent to that of photons and, therefore, the relative biological effectiveness (RBE) is also nearly equivalent (the RBE for protons:photons is approximately 1.1).[11,12] For heavier charged particles such as carbon ions, the density of ionization is greater at the end of their range, which causes greater damage to the DNA within cells at the end of that range and results in carbon ions having a higher RBE (1.5–3). However, it is becoming increasingly evident that RBE is a complex, variable function of radiation dose per fraction, total dose, LET, cell and tissue type, choice of endpoint, and other factors.[13,14] Thus, the RBE may be less than 1.1 at the entrance, may increase with depth, and may be highest at the distal edge of the beam.

These two physical properties of protons, i.e., having a finite range in tissue and having a higher RBE at the distal edge of the beam, make proton therapy both appealing and potentially problematic. Proton therapy is exquisitely sensitive to changes in tumor position and density and to differences in tissue composition; this sensitivity is particularly problematic for thoracic tumors, because the tumors move with lung ventilation and diaphragm motion, and because the tissues along the beam path are quite heterogeneous in structure and density. In PSPT, extreme care must be taken to consider the need to compensate for tumor motion, changes in lung density due to respiration, and uncertainties in proton range with regard to respiration-induced tumor motion and lung density changes (Figure 1). These variables should be assessed separately for each beam direction, and some amount of dosimetric uncertainty should be built into the planning of each beam.[15] Although lung motion and density uncertainties can be accounted for during the treatment planning process by adding generous internal and smearing margins, practical issues regarding inconsistencies in patient setup and positioning and changes in tumor volume between treatment sessions must also be accounted for during the course of treatment.[16]

Figure 1.

Proton dose distribution is vulnerable to changes in anatomy. First row, dose distribution of the initial plan for 4-field passively scattering proton therapy (PSPT) in axial and sagittal views from the (A) left lateral (A1, axial view, A2, coronal view) and (B) left anterior oblique fields (B1, axial view, B2, coronal view). Second row, verification plan in week 4 (C1, lateral field, axial view, C2, lateral field, coronal view) (D1, left anterior oblique field, axial view, D2, left anterior oblique field, coronal view). Note the overshoot of protons to the spinal cord due to tumor cavitation. Third row, dose-volume histogram of initial (square) and verification (triangle) plan showing increased dose to the heart, lung, esophagus, and the spinal cord. Red arrows indicate the dose difference in spinal cord due to tumor shrinkage.

In IMPT, conformity of the proximal and lateral field is achieved by limiting the position of the spots to within the target region only. Dynamic apertures that can change shape layer by layer are being developed to address the issues of large spots appearing in pencil beam scanning proton therapy. In treatment planning, the position and intensities for a matrix of spots within the target volume for each scanned beam are calculated automatically by the treatment planning system to achieve the desired dose distribution. The IMPT dose distribution is more sensitive to uncertainties in set-up and motion than is the PSPT dose distribution. To address this heightened sensitivity, "robust optimization" techniques that simultaneously consider multiple uncertainty scenarios and optimize intensities in the face of all those scenarios are being actively investigated.[17,18]

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