AAPM TG 158: Measurement and calculation of doses outside the treated volume from external-beam radiation therapy

Transcription

1 AAPM TG 158: Measurement and calculation of doses outside the treated volume from external-beam radiation therapy Stephen F. Kry a) Department of Radiation Physics, MD Anderson Cancer Center, Houston, TX 77054, USA Bryan Bednarz Department of Medical Physics, University of Wisconsin, Madison, WI 53705, USA Rebecca M. Howell Department of Radiation Physics, MD Anderson Cancer Center, Houston, TX 77054, USA Larry Dauer Departments of Medical Physics/Radiology, Memorial Sloan-Kettering Cancer Center, New York, NY 10065, USA David Followill Department of Radiation Physics, MD Anderson Cancer Center, Houston, TX 77054, USA Eric Klein Department of Radiation Oncology, Washington University, Saint Louis, MO 63110, USA Harald Paganetti Department of Radiation Oncology, Massachusetts General Hospital and Harvard Medical School, Boston, MA 02114, USA Brian Wang Department of Radiation Oncology, University of Louisville, Louisville, KY 40202, USA Cheng-Shie Wuu Department of Radiation Oncology, Columbia University, New York, NY 10032, USA X. George Xu Department of Mechanical, Aerospace, and Nuclear Engineering, Rensselaer Polytechnic Institute, Troy, NY 12180, USA (Received 28 September 2016; revised 17 May 2017; accepted for publication 25 May 2017; published 20 August 2017) The introduction of advanced techniques and technology in radiotherapy has greatly improved our ability to deliver highly conformal tumor doses while minimizing the dose to adjacent organs at risk. Despite these tremendous improvements, there remains a general concern about doses to normal tissues that are not the target of the radiation treatment; any nontarget radiation should be minimized as it offers no therapeutic benefit. As patients live longer after treatment, there is increased opportunity for late effects including second cancers and cardiac toxicity to manifest. Complicating the management of these issues, there are unique challenges with measuring, calculating, reducing, and reporting nontarget doses that many medical physicists may have limited experience with. Treatment planning systems become dramatically inaccurate outside the treatment field, necessitating a measurement or some other means of assessing the dose. However, measurements are challenging because outside the treatment field, the radiation energy spectrum, dose rate, and general shape of the dose distribution (particularly the percent depth dose) are very different and often require special consideration. Neutron dosimetry is also particularly challenging, and common errors in methodology can easily manifest as errors of several orders of magnitude. Task Group 158 was, therefore, formed to provide guidance for physicists in terms of assessing and managing nontarget doses. In particular, the report: (a) highlights major concerns with nontarget radiation; (b) provides a rough estimate of doses associated with different treatment approaches in clinical practice; (c) discusses the uses of dosimeters for measuring photon, electron, and neutron doses; (d) discusses the use of calculation techniques for dosimetric evaluations; (e) highlights techniques that may be considered for reducing nontarget doses; (f) discusses dose reporting; and (g) makes recommendations for both clinical and research practice American Association of Physicists in Medicine [ /mp.12462] Key words: late effects, neutrons, nontarget radiation, out-of-field dose e391 Med. Phys. 44 (10), October /2017/44(10)/e391/ American Association of Physicists in Medicine e391

2 e392 Kry et al.: TG-158: non-target doses e392 TABLE OF CONTENTS 1. INTRODUCTION 2. CONCERNS WITH RESPECT TO NONTARGET RADIATION 2.A. Secondary cancers 2.B. Cardiac toxicity 2.C. Implantable pacemakers and other electronic devices 2.D. Fetal dose estimates 2.E. Cataracts 2.F. Skin dose 3. DOSE ESTIMATES 3.A. Sources and description of out-of-field dose 3.A.1. Photon beams 3.A.2. Protons and heavy ions 3.B. Dose estimates outside the treated volume 3.B.1. Photon therapy 3.B.1.1. Conventional techniques 3.B.1.2. Intensity-modulated radiation therapy (IMRT) 3.B.1.3. Tomotherapy 3.B.1.4. Volumetric arc therapy (VMAT) 3.B.1.5. Stereotactic radiosurgery/stereotactic body radiotherapy (SRS/SBRT) 3.B.1.6. Flattening filter-free IMRT 3.B.2. Electron therapy 3.B.3. Proton therapy 3.B.4. Carbon ion therapy 3.B.5. Brachytherapy 3.B.6. Concomitant imaging doses 4. MEASUREMENT APPROACHES 4.A. Photon dosimetry 4.A.1. Thermoluminescent dosimeters (TLD) and optically stimulated Luminescent dosimeters (OSLD) 4.A.2. Ion chambers 4.A.3. Film 4.A.4. Diodes 4.A.5. Metal oxide silicon semiconductor field effect transistors (MOSFETs) 4.B. Neutron dosimetry 4.B.1. Thermal neutron-based detectors 4.B.1.1. Thermal neutron detectors 4.B.1.2. Thermal neutron detector moderator systems 4.B.1.3. Rem meters and extended-range Rem meters 4.B.1.4. Bonner sphere spectrometers and extended-range Bonner sphere spectrometers 4.B.2. Fast neutron detectors 4.B.2.1. Bubble detectors 4.B.2.2. Track etch detectors 4.B.2.3. Tissue-equivalent proportional counters (TEPC) 4.C. Phantoms 5. COMPUTATIONAL APPROACHES 5.A. Treatment planning systems 5.B. Analytical models for estimating out-of-field dose 5.C. Monte Carlo methods 5.C.1. Photon transport 5.C.2. Neutron transport 5.D. Phantoms 6. TECHNIQUES TO MINIMIZE NONTARGET DOSE 6.A. Reducing the target volume 6.B. Treatment modalities 6.B.1. Proton vs IMRT 6.B.2. Flattened vs FFF 6.B.3. IMRT vs 3D CRT 6.B.4. Photon vs electron 6.C. Treatment energy 6.D. Photon wedges 6.E. MLC and collimator orientation 6.F. Beam angle 6.G. Jaw tracking 6.H. Patient shielding 6.I. Accelerator shielding 6.J. Imaging dose management 7. DOSE REPORTING 7.A. Biological considerations for dose reporting 7.A.1. Relative biological effectiveness 7.A.2. Quality factor 7.A.3. Weighting factor 7.A.4. Ambient dose equivalent 7.A.5. Recommended biologically weighted dose formalism 7.B. Geometrical considerations for dose reporting 7.B.1. Point dose 7.B.2. Organ doses 7.B.3. Integral dose 7.B.4. Effective dose 7.C. Dose/MU vs dose/cgy vs dose/treatment 7.D. Plotting of neutron data 8. RECOMMENDATIONS 8.A. General/clinical care recommendations 8.B. Additional/research-focused recommendations 8.B.1. General 8.B.2. Neutron dosimetry 8.B.3. Monte Carlo simulations 8.B.4. Dose reporting REFERENCES 1. INTRODUCTION Techniques and technology for delivering radiotherapy have evolved over the past several decades. These advances have greatly improved our ability to deliver higher tumor doses while minimizing the dose to the adjacent organs at risk. However, this improved conformality has not mitigated the problem of doses to normal tissues outside the treated volume. Regardless of the type of radiotherapy being used, out-of-field doses delivered by photons, electrons, protons, or neutrons pose unique challenges to medical physicists. This report details many of those challenges and gives guidance to the medical physics community on how to best determine the out-of-field doses associated with modern radiotherapy techniques and modalities. In radiotherapy, radiation is delivered to achieve a therapeutic benefit within the clinical target volume (CTV). Uncertainties in positioning and delivery require the use of a planning

3 e393 Kry et al.: TG-158: non-target doses e393 target volume (PTV) and the corresponding irradiation of healthy tissue in this volume. Related to the target volume is the ICRU term treated volume, which is the volume of tissue planned to receive prescribed dose (which might be much larger than the PTV for conventional therapy). 1 Tissue outside of the PTV does not benefit from radiation but is unavoidably irradiated. The dose outside the PTV is referred to as nontarget dose. In external-beam radiotherapy, nontarget dose is delivered by nontarget radiation and can be subdivided into two categories. (a) in-field nontarget dose nontarget dose that is within a primary field border, such as entrance and exit dose along the beam path. (b) out-of-field nontarget dose nontarget dose that is not only outside of the PTV but also outside of any primary field edge dose deposited by stray, or secondary, radiation. In radiation therapy, any nontarget radiation should be minimized as it offers no therapeutic benefit. Nontarget dose can be classified as one of three approximate dose levels. High doses (> 30 Gy or > 50% of the prescription dose) are typically directly optimized during the planning process. Intermediate doses (3 30 Gy or 5 50% of the prescription dose) are also often addressed during optimization but are not generally the focus of the treatment plan. Low doses (< 3Gy or < 5% of the prescription dose) are not generally considered during treatment planning. However, even if these low doses were considered, it is difficult to measure, characterize, or model them in the planning system. With modern treatment planning systems, high-dose regions and areas within the primary-beam path are typically well described. However, the accuracy of dose determinations beyond a few centimeters outside the treatment field edge is usually poor. 2,3 In these cases, alternate methods are required to assess the dose to the patient. Low radiation doses outside the treated volume are a concern because they can cause deleterious effects to the patient, as described in Section 2. The risk of late effects from secondary radiation may be more evident today than in the past because the success of cancer screening and modern therapies have increased the number of cancer patients who survive and live long enough for the adverse radiation effects on healthy tissues to manifest. Over the past 30 yr, numerous published studies have detailed nontarget doses from various radiotherapy approaches (see references in Refs. 4 8 ). However, there is no current guidance to aid physicists in terms of dosimetry and computational techniques for assessing nontarget doses, as well as potential treatment and patient management options for minimizing nontarget doses. It is the physicist s responsibility to appropriately determine relevant nontarget doses and contribute to strategies to minimize these doses as needed. Once dose data are available and treatment options have been explored, it is then the treating physician s responsibility to assess the risks and benefits of the radiation treatment so as to make a sound and informed assessment of all possible treatment options. The focus of this report is on those structures that receive low doses of radiation rather than on structures generally included in plan optimization that receive high or intermediate doses. These structures are usually outside the treatment field, and although they receive low doses, they may nevertheless be of concern to the patient because of potential radiosensitivity. To this end, this report aims to address the following charges as they pertain to nontarget radiation: 1. Highlight major concerns of nontarget dose 2. Provide a rough estimate of doses associated with different treatment approaches in clinical practice 3. Discuss the uses of dosimeters and phantoms for measuring photon, electron, and neutron exposures 4. Discuss the use of calculation techniques (including Monte Carlo) for dosimetric evaluations 5. Highlight techniques that may be considered for reducing nontarget doses 6. Discuss dose reporting 7. Make recommendations for clinical and research practice 2. CONCERNS WITH RESPECT TO NONTARGET RADIATION While the risks associated with low doses of radiation are generally small, unnecessary exposures to radiation can be harmful to the patient and should be avoided. This chapter outlines the most common concerns relevant to this task group report as well as the current state of knowledge on these issues. In this section, as driven by the underlying literature, both physical doses (Gy) and biologically weighted dose in Sieverts (Sv) are reported (typically described as dose equivalent). 2.A Secondary cancers Radiation is a well-documented risk factor for cancer induction in virtually any tissue as shown through atomic bomb studies, 9,10 accidental and occupational exposures, 11,12 and medically exposed individuals. 4,13 Different organs show different susceptibilities to radiation-induced cancer per Sv of dose equivalent, and this susceptibility varies with age and, to a lesser extent, sex. 14 Children are several times more sensitive to radiation-induced cancers than adults, although the specifics of this sensitivity can vary among organs. The thyroid and female breast, for example, are very sensitive to radiation-induced cancer in young patients, but by late adulthood are relatively insensitive to radiation. 11,15 Major published epidemiologic risk estimates 4,14 show that, for many second cancer sites, females are more radiosensitive than males. However, with the exception of sex-linked organ risks (including breast), the mechanism and/or validity of this finding is under debate. 13 The relationship between radiation dose and cancer risk in vivo is complicated and not fully understood. At very low doses, less than 0.1 Sv, a linear, no-threshold relationship is typically assumed (at least in radiation protection applications) between dose and risk, but the true nature of this relationship is under active debate 13,16,17 and is unknown at this time. At doses between 0.1 and 2 Sv, the risk of developing a

4 e394 Kry et al.: TG-158: non-target doses e394 second cancer increases linearly with dose (or in a linearquadratic manner for leukemia). 18 At doses greater than 2 Sv, there is ongoing debate about the shape of the dose response relationship. However, most epidemiologic studies 4 indicate that risk increases linearly with dose, even at doses of 40 Sv or more The notable exception is for the thyroid where the risk is dramatically decreased at high doses. 22 Radiation is a clear risk factor for second cancers, but it is only one of many etiologic bases. Roughly 10% of long-term survivors develop a second cancer 8,23,24 ; however, only a fraction of these second cancers are attributable to radiation treatment. Age, genetics, and environmental factors also contribute to the risk of developing a second cancer. 6 Recent studies evaluating this question in adults 23,25 found that of the 10% of patients who developed a second cancer, 8% of those were attributable to the actual radiation exposure from radiotherapy. 23 That is, slightly less than 1% of long-term survivors developed a second cancer from their radiotherapy, although this number may increase with longer follow-up. While radiation is expected to be, and appears to be, a larger risk factor for pediatric patients, 24 this issue requires further quantification. Epidemiologic data indicate that the majority of second cancers occur in intermediate- to high-dose regions, although low-dose regions (down to ~1 Gy) can also show significantly elevated risk of cancer induction. 26,27 Diallo et al. found that 12% of second cancers occurred within the treated volume, 66% occurred at the periphery of the treated volume (within 5 cm of the field edge), and 22% occurred more than 5 cm away from the treated volume, although this was not limited to radiation-induced second cancers. 28 A review of epidemiologic data on second cancers following radiotherapy is available from the National Council on Radiation Protection and Measurements. 4 2.B. Cardiac toxicity Cardiac toxicity is well linked with high doses of radiation, 29 but more recently, it has also been linked with much lower doses 10,30,31 including in medically exposed populations Cancer survivors have a have higher risk of several cardiac effects than the healthy population, including myocardial infarction, pericardial disease, and valvular dysfunction. 43 While radiation-induced cardiac toxicity is not completely understood, it is clear that the risk can be substantial, particularly for long-term survivors: the mortality ratio for patients with a left breast cancer compared with those with a right breast cancer increased from 1.04 at < 5 yr post-treatment to 1.53 at > 15 yr post-treatment. 36 The risk can also manifest at low doses, being evident in patients treated for testicular cancer where the heart was completely outside the treatment field. 46 In both breast cancer populations 35 and pediatric cancer populations, 43 the risk of radiation-associated cardiovascular disease was found to increase linearly with radiation dose, suggesting that a linear no-threshold relationship may be a good descriptor of the risk. It is not entirely clear which structure(s) in the heart are the most susceptible to development of heart disease. Imaging studies have shown the left ventricle and coronary arteries to be sensitive to radiation damage. Higher incidences of perfusion defects have been observed in patients having more than 5% of the left ventricle in the treatment field, 47,48 although these imaging results have not yet been linked to increased cardiac events. Radiation damage in the coronary arteries has also been linked with coronary artery stenosis. 49 However, many questions remain unanswered as the majority of dose-response data are based on mean heart dose. In additional, the relationship between risk and age of exposure remains largely unknown. These results and limitations are discussed in more detail in the recent NCRP Report. 4 2.C. Implantable pacemakers and other electronic devices Pacemakers and other implantable electronic devices (IED) are sensitive to relatively low levels of radiation as documented in AAPM Report TG Implantable electronic devices, if exposed to excessive cumulative doses, can malfunction as a result of a build-up of electrical charge (typically at doses > 2Gy). 50 A second failure mode, common in clinical practice, is a single-event upset, which can be caused by radiation with high linear energy transfer (LET), such as neutrons. 51 The risk of these singleevent effects is stochastic in nature, so there is no safe dose threshold; the risk simply increases as the number of neutrons increases. Finally, transient device interference can also occur, even at low doses, from direct irradiation of the device. 52 Additional information on these mechanisms and management of patients with implanted devices will be available in the upcoming AAPM Report TG D. Fetal dose estimates Radiotherapy treatment of pregnant patients is occasionally unavoidable. Protecting the fetus is critical because of the high radiosensitivity of developing organs and tissues. Fetal radiation exposure may cause many adverse effects at doses as low as 10 cgy, as described in AAPM Report TG Special emphasis has been placed on fetal dose assessment in the literature. This includes detailed Monte Carlo models, 54,55 although more often the dose to the fetus is estimated by using data in AAPM Report TG or similar (e.g., Peridose 56 ), and while these dose estimates are generally accurate to within 30% (on average), errors of a factor as high as 2 3 have been reported. 57 However, practical analytical models for estimating fetal dose are usually based on simple radiation fields (square, open fields) and are therefore not well suited for modern radiotherapy utilizing advanced technologies. 2.E. Cataracts Adverse effects from radiation exposure to the eye have been known for decades, 58 including cataract formation. 59,60 It is now well known that the subcapsular lens epithelium, particularly where it differentiates to lens fibers, is particularly

5 e395 Kry et al.: TG-158: non-target doses e395 susceptible to radiation damage. The development of radiation-induced cataracts is a known late effect from radiation therapy, and is dependent on radiation dose, dose rate, and age of the lens. 65 While the current guidelines for the threshold dose for cataract formation ranges from 2 to 5 Gy, recent studies indicate that the dose could be less than 0.5 Gy. 66,67 There is also strong evidence that cataracts may be better described by a linear, no-threshold model. 66 While limited data are available on cataract formation resulting from out-of-field dose, studies have indicated that very low doses can lead to cataract formation 25 yr or more after exposure F. Skin dose The skin is almost always irradiated as a nontarget structure and is relatively sensitive to low doses of radiation. Erythema, desquamation, and necrosis can all be observed in contemporary radiotherapy. 69,70 These effects are largely limited to areas of the skin within the treatment field as these skin doses are higher than outside the treatment field. 71 Correspondingly, more attention is given to the skin dose within the field. However, at even relatively low skin doses outside the treatment field, epidemiological studies have shown elevated skin cancer risk. 21 It is important to note that measurement and calculation of the skin dose require special consideration because of the high-dose gradients. This is less of a concern outside the treatment field where the gradients are more modest. This report does not address measurements or calculations of dose relevant to the skin, nor does it describe the skin doses associated with different procedures. Interested readers are directed to the recent AAPM TG-176 report DOSE ESTIMATES This chapter outlines sources and levels of secondary radiation associated with common clinical radiation therapy techniques. The sources of secondary radiation are provided to assist physicists in designing clinical treatment approaches to reduce nontarget doses. Typical dose levels are provided for each therapy approach to aid with an apriorievaluation of the dose or to verify that more detailed measurements are reasonable. Due to inter-patient and inter-plan variability, these values should not be used as a replacement for individual dose assessment. Dose equivalent is plotted throughout this section as an estimate of the biologically weighted dose relative to x rays. Studies that determined the biologically weighted dose following a different method (as discussed in Section 7, e.g., by calculating equivalent dose instead of dose equivalent) were not differentiated or modified before being plotted. 3.A. Sources and description of out-of-field dose 3.A.1. Photon beams Photon treatments that involve a linear accelerator (linac) result in three unintended sources of radiation: (a) patient scatter, scatter of the primary treatment beam outside the treatment area once it has entered the patient; (b) collimator scatter, scatter of radiation in the head of the accelerator that exits the accelerator through the treatment field opening but strikes the patient outside the treatment field; and (c) head leakage, radiation that penetrates through the accelerator head shielding and strikes the patient away from the treatment field. In close proximity to the treated volume, the dominant source of out-of-field radiation is patient scatter. At greater distances from the field edge (greater than ~20 cm, but the exact distance depends on treatment field parameters, beam modulation, and beam energy), head leakage becomes the dominant source of out-offield dose. 73 Because of the contributions of patient and collimator scatter, the energy spectrum is softer outside of the treatment field than within. While the mean energy at d max for a 6- MV beam is around 1.6 MeV (for a Varian accelerator), the average energy outside of the treatment field is much lower, typically between 0.2 and 0.6 MeV. 74,75 The single most important parameter affecting the total dose outside the treatment field is distance from the field edge. 53 The out-of-field photon dose decreases in a roughly exponential manner with distance from the field edge, although in an IMRT environment there can be some dependence on the specific treatment plan. In high gradient regions near the field edge, a small difference in plan optimization can substantially impact the absolute dose delivered to a point in the patient. This can be particularly important as these intermediate doses are associated with a large fraction of the observed second cancers in patients. 28 The dose outside the treatment field also depends on field size, increasing with increasing field size because larger irradiated volumes produce more patient scatter. This field-size dependence is most pronounced close to the field edge, where patient scattered radiation is the most important component of the out-of-field dose. Farther from the field edge, there is less field size dependence 53 because this dose is dominated by head leakage. For modulated treatments, the dose outside the treatment field also depends on beam modulation. The dose increases with increasing modulation because more head leakage and more collimator scatter is generated as the number of monitor units (MU) increases to deliver the modulated fields. 76,77 The dependence on modulation is most noticeable farther from the treatment field; near the field edge the dose is dominated by patient scatter and patient scatter depends only on the volume of tissue irradiated. For patient scatter (and the out-of-field dose near the target), modulation will only matter insofar as it changes the conformality of the treatment (smaller treated volumes yielding less out-of-field dose), although importantly, increased conformality may also change the location of the edge of the field (defined as the 50% isodose line). Out-of-field dose differs slightly among accelerator models 78 because of variations in head shielding designs. In contrast, the dose varies little with beam energy, being the same magnitude for energies ranging from 4 to 25 MV. 53 Likewise, out-of-field dose varies minimally with depth in the patient, 53,79 except near the surface of the patient. 79,80 The dose at the surface is highly elevated (typically by a

6 e396 Kry et al.: TG-158: non-target doses e396 factor of 2 5) by electrons from the accelerator head; this dose elevation is maximal at the surface and decreases with depth until the nominal beam s d max and then remains relatively constant at all depths below that. When beam energies exceed ~10 MV, neutron production in the head of the accelerator becomes relevant. Neutrons are produced primarily in the primary collimator and to a lesser extent by the target and flattening filter The jaws or multileaf collimators (MLC) can also be a major source of neutrons when they intersect the primary photon beam (such as for a closed field). 81,82 Neutron production in the patient is negligible because of the low (c,n) cross section for the low-atomic number materials that constitute tissue. 84 Neutron production in the accelerator head results in isotropic neutron emissions, exposing the patient to a relatively uniform bath of neutrons. The fluence of neutrons and associated dose equivalents increase linearly with the number of MU delivered. 85 Secondary neutron spectra measured in air in the patient plane are shown in Fig. 1 and include a fast neutron peak centered between 0.1 and 1 MeV, a maximum energy of approximately 10 MeV, and a low-energy tail that arises from neutrons being elastically scattered throughout the treatment vault. The shape of the neutron spectrum does not depend on treatment energy 82,86 or accelerator manufacturer. 86,87 However, the number of neutrons produced increases substantially with increasing photon-beam energy. For a Varian accelerator, the number of neutrons increases by ~10 times from 10 MV to 15 MV and by ~2 times from 15 MV to 18 MV (Fig. 1). 87,88 Neutron production is dependent on the accelerator manufacturer because of variation in the photon energy spectra. Varian accelerators generate approximately twice as many neutrons as Siemens or Elekta accelerators 89 because they have a higher peak photon energy for the same nominal beam energy. 90 The size of the treatment vault may also affect the neutron dose, as smaller bunkers will theoretically have increased room-scattered neutrons. Some investigators have observed this effect, 86 whereas others have found no such relationship. 88 Finally, neither the neutron spectrum nor the neutron fluence are substantially affected by distance offaxis, SSD, or field size. 73,86,91 The majority of neutron dose equivalent is deposited by fast neutrons. 89 Consequently, the depth in the patient is also a very important parameter affecting neutron dose. Once neutrons strike the patient, they are rapidly thermalized 86,91,92 and therefore show a sharp decrease in dose equivalent with increasing depth. The neutron percent depth dose equivalent (PDDE) decreases by approximately 20% per centimeter penetration in tissue (Table I), and because the spectrum does not change, these data are a good estimate of the out-of-field PDDE for any distance, field size, or SSD. 91 For high-energy photon treatments, the secondary photon dose is larger than the secondary neutron dose equivalent, except near the surface of the patient at large distances from the treated volume A.2. Protons and heavy ions Secondary dose equivalent in proton therapy is generated by neutrons, secondary protons, light charged ions ( 2 H, 3 H, 3 He, 4 He), recoil heavy ions, and photons. The vast majority of dose equivalent originates from neutrons, with secondary photons contributing up to 10% of the ambient dose equivalent Secondary neutrons are produced when primary ions undergo nuclear interactions with beam-line components (external neutrons) or the patients themselves (internal neutrons). fluence per photon Gy to isocenter (ncm -2 u -1 Gy -1 ) 4.0E E E E E E E E MV 18 MV 15 MV 0.0E+00 1E-09 1E-08 1E-07 1E-06 1E-05 1E-04 1E-03 1E-02 1E-01 1E+00 1E+01 energy (MeV) FIG. 1. Neutron fluence spectrum per unit lethargy per photon Gy to isocenter, measured in the patient plane at 40 cm from isocenter for three different x-ray beam energies from a Varian linac. Data from Howell et al. 87 See Section 7.D for an explanation of lethargy.

7 e397 Kry et al.: TG-158: non-target doses e397 TABLE I. Neutron percent depth dose equivalent (PDDE) data. Photon therapy describes the PDDE characteristics for out-of-field neutrons produced by photon therapy (from Kry et al. 91 ). Proton therapy describes the PDDE characteristics for external neutrons outside of the treatment field produced by 250 MeV proton therapy (from d Errico et al. 92 based on the out-of-field spectrum measured by Howell and Burgett 94 ). Depth (cm) Photon therapy PDDE (%) Proton therapy PDDE (%) For passive scattering proton therapy, externally produced neutrons are the dominant source of out-of-field dose equivalent and comprise the majority (~85%) of the whole-body dose from neutrons. 95,98 External neutron production increases with proton energy and depends on the materials (type and dimensions) in the beam path and, hence, on the design of the proton beam line. Generally, more material intersecting the beam will result in more neutron production. Passive scattering treatment heads typically have fixed primary field sizes and thus have to block a substantial portion of the beam, thereby increasing neutron production and neutron dose as patient aperture openings decrease Because of this aperture opening dependency, and the blocking material s proximity to the patient, the field collimator used in passive scattering is often the dominant source of neutron dose equivalent in the patient. However, neutrons will also be generated in the beam scattering devices, including the modulators that are used to degrade the beam energy and patient-specific modifiers such as the aperture 104 or compensator. 101,105 Neutrons generated in the patient (internal neutrons) also contribute, particularly downstream of the target and within 10 cm from the field edge where they can contribute approximately half of the dose equivalent. 95 Internal neutron production increases with beam energy, 95 as well as with field size because the proton beam is incident on a larger volume of tissue. 101 In scanning beams, internal neutrons are the dominant source of out-of-field dose equivalent because the external neutron contribution is very low. Scanning beams, therefore, produce fewer neutrons. 106 For both passive scattering and scanning beam proton therapy, the dose due to neutrons decreases as a function of distance from the field edge, largely because the internal neutrons are attenuated. 106 The neutron dose varies with treatment site and field characteristics. 101 Importantly, there is a higher organ dose of neutrons in smaller patients, such as children, because of the closer proximity of normal tissues to the field edge and surface. 101 Both internal and external neutrons lose energy as they penetrate tissue. The percent depth dose equivalent characteristics for externally produced neutrons (average energy of 12.3 MeV from Howell and Burgett 94 ) from a 250 MeV proton beam are included in Table I based on the data of d Errico et al. 92 Because neutrons in proton therapy have a higher energy than those produced in photon therapy, the neutrons are much more penetrating. Heavy ion therapy produces neutron dose characteristics similar to those of proton therapy. While there is a higher neutron emission per ion from heavy ion therapy, this is offset by the smaller ion fluence required to deliver equivalent dose, resulting in very comparable neutron dose equivalents. 107 The major difference between secondary neutron spectra from proton and photon beams is that proton neutron spectra have two pronounced high-energy peaks rather than one (Fig. 2). Like the photon neutron spectra, the proton neutrons have a peak at a few hundred kev that is produced from evaporation processes and is isotropically generated as well as a low-energy tail that extends down to thermal energies from room-scattered neutrons. Uniquely, however, the proton neutron spectra have a second peak that starts at around 20 MeV and extends up to the maximum proton beam energy. 94, The high-energy peak contains forwarddirected neutrons from direct (nucleon nucleon) reactions. 109,110 The relative importance of each peak depends strongly on the measurement location (as they have different angular dependencies), and presence of a phantom/patient (as they will attenuate differently because of the different energy). There are substantially fewer neutrons present in data set III of Fig. 2 in part because of these geometrical and phantom considerations, but also because the proton energy was lowest for this series and because it was measured on a scanning beam, which largely eliminates the presence of external neutrons. Regardless of these differences, and as with the photon-produced neutrons, the majority of the proton-produced neutron dose equivalent is deposited via fast neutrons. 3.B. Dose estimates outside the treated volume This section summarizes available data on out-of-field dose for different treatment modalities. The data presented are combined from several similar investigations to identify the upper and lower bounds of the out-of-field doses. Other overviews of out-of-field doses are presented in recent review articles. 5,7 For modulated treatments, plan design and quality have the potential to substantially change the position and shape of the dose gradients, which can substantially impact the doses near to the treatment field. The values presented are based on a wide range of different conditions. Measurements are made at different depths in

8 e398 Kry et al.: TG-158: non-target doses e E+06 Data Set I fluence per proton Gy to isocenter (ncm -2 u -1 Gy -1 ) 1.0E E E E E+05 Data Set II Data Set III 0.0E E E E E E E E+03 energy (MeV) FIG. 2. Neutron fluence spectrum per unit lethargy per proton Gy to isocenter measured for three different proton beam lines. Data sets I 94 and II (unpublished data from R. Howell, measured at same facility and with same technique as in Howell et al. 111 ) used a 250 MeV passively scattered proton beam, with the measurement location in the patient plane 100 cm from isocenter (with no phantom). Data set III 108 used a 172 MeV scanning beam with the measurement location 1.15 m downstream from isocenter with a phantom present. different phantoms (although depth only minimally affects dose outside the treatment field 53 ). Similarly, different dosimeters were used, and several data series are computationally derived. While all series included in the below section were vetted, each has its own uncertainty based on the methods and corrections applied (see Section 4.A). While these issues contribute to some of the variability in the data presented in this section, the majority of the variability arises from differences in treatment delivery, treatment optimization, and technology employed. Correspondingly, these values can be used to estimate the out-of-field dose from a clinical treatment, but only as a rough estimate. This can be done by identifying the type of treatment from the options below, estimating the distance from the field edge for the sensitive structure of interest, and multiplying by the prescription dose. 3.B.1. Photon therapy Conventional techniques: Figures 3(a) and 3(b) summarize out-of-field doses for 6-MV and 18-MV conventional photon beams. Because the out-of-field dose varies minimally with depth, the distance from the field edge is typically considered as simply the orthogonal distance from the plane of the field edge. Approximate upper and lower bounds are provided to indicate the range of doses that might reasonably be expected. Higher out-of-field doses come from larger treatment fields. Data series with markers are from 10 cm 9 10 cm fields and highlight the impact of the MLC, whether it is absent, 53 retracted, 112 or aligned with the secondary jaws. 112 The additional shielding benefit of the MLC is clear, and this benefit is most pronounced far from the treatment field where head leakage dominates. The relative importance of neutrons from a Varian accelerator is shown for open fields based on calculations of the neutron dose equivalent in tissue. 91 Except for far from the treatment field and at shallow depths, the photon dose is larger than the neutron dose equivalent. 113,114 Neutron doses associated with high-energy photon therapy have been most commonly assessed in air as this is generally the most straightforward condition to perform these measurements. Representative ambient dose equivalent values associated with a wide range of linear accelerator manufacturers and treatment energies are shown in Table II. Additionally, neutron measurements have also been done on, as well as in, patients. It is important to note that in-air and surface measurements substantially overestimate the dose received by specific organs in the patient, often by a factor of 10 or more, 91 because of the very sharp PDDE curve. Figure 4 shows neutron dose equivalents assessed in air at different energies for a Varian accelerator and at 18 MV for a Siemens accelerator, 20 cm from the field edge. These in-air points are contrasted with two data series showing neutron doses as a function of distance from the edge of an 18-MV Varian beam at the surface of a phantom and in a phantom at typical organ depths. Intensity-modulated radiation therapy (IMRT): For IMRT treatments, the definition of field edge is unclear compared with conventional radiotherapy. This report recommends using the 50% isodose line as the field edge. Practically, the jaw edge or maximum field opening will typically be close to

9 e399 Kry et al.: TG-158: non-target doses e399 TABLE II. Neutron ambient dose equivalent in-air outside the treatment field (10 50 cm from central axis). Neutron ambient dose equivalent (outside the treatment field) Photon beams Electron beams Linac Energy H*(10) (lsv/mu) Linac Energy H*(10) (lsv/mu) Varian b Varian e a,c e a e,d a d Elekta a,c e a Elekta e b e,d Siemens b Siemens f a,c f a,c f c a Ref. [87]. b Source strength values averaged from data in Ref. [88] and converted to H*(10) by scaling with the average Q to H*(10) relationship from Howell et al. 87 c Ref. [115]. d Ref. [116]. e Ref. [117]. f Ref. [118]. FIG. 3. Out-of-field doses reported for 6-MV (a) and 18-MV (b) conventional photon beams. 53,91,112 All of the data points in the plotted data series were acquired using a 10 cm 9 10 cm field size. Upper and lower bounds are provided on the figures (solid curves) to estimate the highest and lowest out-of-field doses that would be reasonably expected. the 50% isodose line and may, therefore, be a reasonable surrogate. As with conventional treatments, the distance from the field edge is generally taken orthogonally from the plane of the 50% isodose line/field edge because of the lack of depth dependence. Compared with 3D CRT treatments, IMRT has advantages and disadvantages in terms of nontarget dose. IMRT treatments typically have lower doses near the edge of the treatment field but higher doses far from the treatment field. 119,120 Near the treatment field, IMRT provides better conformality, which constricts the field edge and thereby reduces the volume of tissue receiving high doses. 76 However, far from the treatment field, beam modulation leads to increased head leakage than can result in higher doses. The difference in conformality and modulation between IMRT and 3D CRT are driven largely by planning style and optimization, and are often in direct opposition. More conformal treatments (decreasing intermediate and high doses) are often more modulated (increasing low doses). The trade-off in importance between intermediate and high doses near the field edge and low doses further from the field may depend on the specific concern being evaluated. For example, Diallo et al. 28 reported that the majority of second cancers were in organs near the treatment field as opposed to organs far out-of-field, indicating the importance of the high-dose region. However, FIG. 4. Out-of-field neutron dose equivalent data from high-energy photon treatments showing the impact of treatment energy and depth in tissue. All data are for Varian accelerators unless otherwise stated. Data are from Kry et al. (18 MV) 91 and Howell et al. 87 for a pregnant patient, the low dose may be of utmost concern if the fetus is far from the field. Figure 5 combines reported data on IMRT treatments, including both segmental and dynamic MLC delivery and various treatment sites. All of the datasets used low-energy beams except a single 18-MV dataset 93 that includes both photon and neutron doses. Note that these dose values are normalized per dose to the target. Overlaying the data series are estimates of reasonable upper and lower bounds based on the ranges of the data series included. There is notable spread

10 e400 Kry et al.: TG-158: non-target doses e400 in the IMRT data because of differences in target volumes and in modulation. Larger target volumes (such as head and neck) will cause more patient scatter and therefore higher out-of-field doses, particularly near the treatment field. More modulation (i.e., more MU/dose) will cause more head leakage and therefore higher out-of-field doses far from the treatment field. Among other influences, modality affects the degree of modulation. For example, sliding window delivery typically requires 20 30% more MU than step-and-shoot IMRT Tomotherapy: The out-of-field doses associated with various forms of tomotherapy, including serial tomotherapy 128,129 and more modern helical systems, 127,130 have been evaluated. While the dose varies with the complexity of the target volume and the total number of MUs delivered, in general, the magnitude and distribution of the out-of-field doses from tomotherapy are similar to those of IMRT. 128 This has included data with slightly higher doses (observed in the tomotherapy data included in Fig. 5), 127 as well as slightly lower doses 130 than conventional IMRT. Volumetric arc therapy (VMAT): VMAT is a form of IMRT delivery and thus shows out-of-field dose characteristics similar to those of other IMRT delivery techniques. However, VMAT will produce a slightly different nontarget dose distribution in the plane of the treatment delivery compared with IMRT because VMAT spreads out the dose delivery over more angles. Consequently, while the average dose will be similar between VMAT and IMRT, more tissue will be irradiated to a lower dose with VMAT. As with other IMRT techniques, out-of-field doses are primarily dependent on field size and number of MU. While field size is restricted by the size of the target, MU depends FIG. 5. Distance and dose data from various IMRT treatments representing different treatment sites, treatment energies, and delivery systems. 89, Dose equivalents (msv/gy) are plotted as a fraction of the dose to the target. Values with an asterisk 89 have relatively high modulation. The upper and lower bounds shown (solid curves) describe the empirical range of doses associated with different IMRT techniques. These bounds are included on subsequent figures. very strongly on the number of arcs employed. Single arc therapy typically has fewer MUs than traditional IMRT. Multiple arc treatment will have more MU than single arc therapy, and can have comparable or even more MU than traditional IMRT. For prostate treatments, MUs have been reported to be 16 39% lower for VMAT than for step-andshoot IMRT 124,131,132 and 42 68% lower than for sliding window IMRT. 124,133,134 In cases where fewer MU are required (particularly single arc), leakage radiation is lower for VMAT, and consequently out-of-field doses are also slightly lower for VMAT than for step-and-shoot IMRT 124,135 as shown in Fig. 5. Near the treatment field, where patient scatter dominates, there is no difference between VMAT and other IMRT modalities except as the difference achieved in plan optimization affects the dose fall-off (eg, single arcs are less conformal and will therefore usually have more tissue inside the 50% isodose cloud. 136 Stereotactic radiosurgery/stereotactic body radiotherapy (SRS/SBRT): In addition to traditional linacs using MLCs or cones, stereotactic radiation therapy can be delivered with specialty delivery devices such as CyberKnife and Gamma Knife. These different devices and different beam modifiers can create unique nontarget dose distributions. Importantly, the flexibility in beam direction during many SRS/SBRT treatments increases the reliance on noncoplanar beams as compared with traditional therapies. Highly noncoplanar, particularly vertex fields, can substantially increase the dose throughout the patient 137 as primary beam is delivered down the axis of the patient. Data for several stereotactic treatments using different delivery systems are shown in Fig. 6, which also includes the IMRT upper and lower dose limits from Fig. 5. Original Cyberknife units were found to have very high out-of-field doses relative to other modalities. 126 The CyberKnife system was subsequently upgraded with an additional tungsten collimator, 138,139 which reduced nontarget doses within 30 cm of the field edge to levels comparable to those of other modalities. For distances greater than 30 cm from the field edge, CyberKnife nontarget doses remain high relative to other treatment modalities. 130 Furthermore, some CyberKnife treatments require more MUs than other modalities, 126 and the robotic placement of the accelerator may place the source of leakage radiation within close proximity to the patient, all contributing to higher nontarget doses. 126 Because Gamma Knife units have a relatively low-energy spectrum, doses very close to the treatment field are often higher than those of CyberKnife or linac systems. 126,139 For most distances from the field edge, Gammaknife treatments have been shown to produce intermediate doses higher than linac-based therapy but lower than those achieved by CyberKnife. Linac-based stereotactic procedures have been studied recently 125,126,140 and found to produce relatively low doses compared to other stereotactic modalities. 126 They also

11 e401 Kry et al.: TG-158: non-target doses e401 FIG. 6. Published out-of-field dose data for different stereotactic procedures for the brain 126,138 and for the lung (using an MLC) 125 on several different SRS/SBRT platforms (CK, Cyberknife; GK, Gammaknife). Dose equivalents (msv/gy) are plotted as a fraction of the dose to the target. The Petti and Chuang cyberknife doses are for an original Cyberknife unit and an updated unit with additional head shielding, respectively. The upper and lower bounds produced for the IMRT data in Fig. 5 are also included for reference. produce low out-of-field doses compared with traditional IMRT treatments, as stereotactic procedures use relatively small fields and usually have low modulation, minimizing both patient scatter and head leakage. Total out-of-field doses may also be lower because the prescription dose is usually lower in stereotactic procedures. Flattening filter-free IMRT: In modern radiotherapy, treatment delivery is feasible without the flattening filter for IMRT or stereotactic radiosurgery/radiotherapy. 141 In addition to drastically increasing the dose rate, 142,143 the flattening filter-free (FFF) modality also reduces the outof-field dose 125, as highlighted in Fig. 7. It reduces head leakage because less target current is required. Similarly, collimator scatter is reduced because the flattening filter is no longer a source of scatter. However, patient scatter may be increased in FFF modes. Increased patient scatter has been documented in the Varian implementation of FFF (in which the beam energy is not changed in moving to the FFF mode) because of the softer primary-beam spectrum. 146 Regardless of the FFF implementation, for small-field treatments (such as used in SRS/SBRT) and IMRT treatments, doses are consistently lower for FFF deliveries. FFF beams produce comparable or slightly reduced doses near the treated volume 125,143,146 and substantially reduced doses further from the treated volume. 125,143,144,146 Most current FFF beams are operated at energies of 10 MV or lower, so neutron production is not generally a concern. However, even for higher energy photon beams, neutron production in FFF mode is reduced by as much as 70% over conventional flattened beams because of the more efficient delivery and elimination of the flattening filter as a source of neutrons FIG. 7. Published out-of-field dose data comparing FF vs FFF treatments. 125,146 Dose equivalents (msv/gy) are plotted as a fraction of the dose to the target. The upper and lower bounds produced for the IMRT data in Fig. 5 are also included for reference. 3.B.2. Electron therapy Nontarget doses associated with electron radiotherapy have received some recent attention. The characteristics of the nontarget dose from electron therapy are starkly different than those from photon therapy, and the magnitude of the nontarget dose from electron therapy is highly dependent on accelerator manufacturer, as shown in Fig. 8, and even based on specific applicator type. 150 While out-of-field doses generally decrease with distance from the edge for electron therapy, this is less pronounced than for photon therapy (particularly for Elekta and Siemens accelerators), and the dose may show regions of increase, particularly around 20 cm from the field edge. The majority of the out-of-field dose is from scattered electrons, particularly scattered primary electrons or Compton scattered electrons originating in the lower trimmer. 152 The increase in dose at ~20 cm from the field edge arises from scattered electrons originating from the rounded surface of the MLC. 152 The difference in dose between the Infinity (MLCi2 head) and Versa HD (Agility head) from Elekta (particularly at 20-cm distance) is likely the result of a change in head design: the MLCi2 has a divergent jaw, whereas the jaw and MLC of the Agility head are rounded. The rounded surface provides more electron scatter outside of the treatment head. Variability is also seen between Varian models; the 2100C accelerator produces notably more dose than the TrueBeam (multiple Varian 21iX machines were found to produce comparable doses to the TrueBeam). 153 While the out-of-field doses from electron therapy can be substantial, they are also superficial. Because the dose is deposited by primary electrons, they only penetrate a few centimeters. The dose decreases rapidly, such that by ~2 6cm depth, there is only residual bremsstrahlung dose (which is about ~10% of the superficial electron dose). 154 Particularly for the Siemens accelerator and the Elekta Versa HD, out-of-field doses from electron therapy are relatively high compared with doses from photon therapy (even IMRT). In general, the out-of-field dose from electron

12 e402 Kry et al.: TG-158: non-target doses e402 therapy increases with treatment energy. It also increases with obliquity (i.e., with the gantry tilted toward the measurement point) but varies little with applicator size. 154,155 Neutron production has been measured for electron treatments. 118,156 Neutron dose equivalents associated with different energies and accelerator manufacturers are shown in Table II. The neutron production from electron therapy is typically ~5% of the neutron production associated with photon therapy of the same nominal energy. 3.B.3. Proton therapy The main contribution to the out-of-field dose in proton therapy is due to neutrons. Even for scanning beams, the contribution of secondary photons to the ambient dose equivalent is typically only ~10% of the neutron contribution. 97 The neutron yield and out-of-field dose during proton treatments depends on geometrical and physical parameters. Figure 9 presents combined data from several investigations on out-offield neutron dose equivalent from proton therapy. 103, Marked variability has been observed in these doses. This spread in dose data reflects different delivery techniques (scanning and scattering), different treatments and treatment energies, and different beam lines/manufacturers, and is also affected by the challenges inherent to conducting high-energy neutron measurements. In passive scattering, the parameters determining the neutron contamination of the primary proton beam include the characteristics of the beam entering the treatment head, the material in the double-scattering system and range modulator, the position of the treatment head relative to the patient, and the field size upstream of the final patient-specific aperture. 101,109,159,163 These characteristics vary among machines, largely preventing definition of a typical neutron dose in proton therapy. 99,100 The additional dependence of the neutron dose on the passive scattering technique, beam-line materials, and specific field settings in the treatment nozzle implies that even for the same facility there can be substantial variation in fields. Nevertheless, relative conclusions can be drawn. For passive scattering therapy, higher neutron doses are observed for higher treatment energies and smaller field sizes. For scanning beam therapy, higher neutron doses are observed for higher energies and larger field sizes). Although proton therapy produces more neutron dose equivalent than high-energy photon therapy, 164 overall proton therapy generally offers a substantial benefit in nontarget dose. This is largely due to the in-field nontarget sparing associated with the distal fall-off beyond the Bragg peak. However, out-of-field dose equivalents resulting from proton therapy are also typically lower than the out-of-field doses resulting from photon therapy (Fig. 9). Near the treatment field, the neutron dose equivalent from proton therapy is typically much lower than the scattered dose from photon therapy. At greater distances, the doses are often comparable and may be even larger for proton therapy. 95,165 There is generally an overall advantage for proton therapy vs photon therapy in the treatment of a given disease site This advantage is more pronounced for lower energy proton treatments 139 and for scanning beam therapy B.4. Carbon ion therapy Treatment with carbon ions is a relatively novel approach. 169 As in proton therapy, nontarget doses are generally lower than those from photon therapy primarily because of the reduced exit dose associated with the Bragg peak and minimal lateral scatter of the primary particles. Relative to protons, carbon ion therapy has a dosimetric advantage in the entrance region because the ratio of the Bragg peak dose to the dose in the entrance region is larger for heavy ions. However, carbon ion performance is poorer than protons distal to the Bragg peak because the nuclear interactions along the FIG. 8. Published out-of-field dose data for different electron beams. 112, Doses were measured at a superficial depth (typically ~1 cm). Dose equivalents (msv/gy) are plotted as a fraction of the maximum dose on central axis. The upper and lower bounds produced for the IMRT data in Fig. 5 are also included for reference. FIG. 9. Published data from several investigations on out-of-field neutron dose from proton therapy 103, (passive scatter delivery unless otherwise stated in the legend) and carbon therapy. 163 Dose equivalents (msv/gy) are plotted as a fraction of the dose to the target. The upper and lower bounds produced for the IMRT data in Fig. 5 are also included for reference.

13 e403 Kry et al.: TG-158: non-target doses e403 TABLE III. Nominal radiation dose ranges to patients from representative standard clinical concomitant imaging procedures. Note that actual values will depend heavily upon protocol and mode selection as well as patient size and image quality requirements. 173,174, Imaging method Portal imaging (MV) Planar imaging (kv) Cone-beam computed tomography (kv) (Elekta XVI â, Varian OBI â ) Parameter assumptions or body area Solid-state flat-panel detectors; 100 cm source/isocenter distance Gantry-mounted kv system or ceiling/floor-mounted kv sources/flat-panel detectors Estimated radiation dose per image (cgy) < 1 5 ~ Imaging site and doses to structures: Head/neck Soft tissue ~0.2 2 Red bone marrow ~4 5 Bone surfaces ~9 20 Chest Soft tissue ~ Red bone marrow ~1 4 Bone surfaces ~2 10 Pelvis Soft tissue ~0.5 5 Red bone marrow ~2 3 Bone surfaces ~5 8 beam path produce a build-up of light fragments that can deposit energy distally to the spread-out Bragg peak (i.e., a fragmentation tail). As in proton therapy, the entire body of the patient is exposed to nuclear secondaries (primarily neutrons), which are produced primarily in beam-line components of the beam delivery system, such as the scattering foil, but are also produced in the patient. Neutrons produced during these interactions are highly forward-directed, unlike neutrons from proton therapy, which are more isotropic. The total neutron dose equivalent is fairly comparable between proton and carbon ion therapies. 163 Although the cross section for nuclear interactions is higher for carbon ion beams than for proton beams, fewer carbon ions are required to deliver the therapeutic dose. Consequently, the net neutron production during carbon therapy is similar to that during proton therapy (Fig. 9). source (> 30 cm) because of the low-energy spectrum emitted from current commercially available electronic brachytherapy sources (i.e., kvp). Comparisons of nontarget doses between brachytherapy and external-beam therapy are complicated by the external-beam treatment having a field edge while the radioactive source does not. Figure 10, which aligns the field edge and source location, shows notably that nontarget doses from high-energy brachytherapy sources are actually reasonably close to doses from external-beam therapy. In comparing brachytherapy ( 60 Co and 192 Ir) to external-beam therapy for treatment of the prostate, Candela-Juan et al. 170 found similar or even higher doses from brachytherapy as far as 35 cm from the prostate. At distances beyond 35 cm from the target, the nontarget dose from brachytherapy became substantially less than that from external-beam therapy. 3.B.6. Concomitant imaging doses Although imaging dose is not generally considered part of the stray radiotherapy dose, its presence is highly relevant to overall out-of-field dose and typical doses from standard imaging procedures are shown in Table III. Patients undergoing radiotherapy are subjected to two types of imaging procedures: patient simulation before treatment and daily localization during treatment. A variety of imaging solutions are available for these purposes, and specific overviews of image-guidance techniques and their associated radiation dose levels are available in AAPM Task Group Report and Purdy. 174 As such, careful evaluation, measurement, and optimization of concomitant imaging doses over the whole course of planning and treatment are necessary (for additional information, see AAPM Reports 104, , 176 and the upcoming TG-180 report, as well as International Commission on Radiological Protection (ICRP) Publication ). The relative contribution of nontarget imaging doses compared with nontarget therapeutic doses was evaluated by Halg et al. 178 Considering all nontarget structures, the 3.B.5. Brachytherapy Although not external beam, brachytherapy provides an alternative modality that can be compared in terms of nontarget dose. Available data for doses away from the target during brachytherapy are shown in Fig. 10. For the highenergy sources, the dose depends only minimally on the type of brachytherapy source. While no data have been published on low-energy brachytherapy sources, such as 125 Ior 103 Pd, these sources would be expected to produce much lower doses than those presented in Fig. 10. Nontarget doses from electronic brachytherapy sources are lower than those from their radioactive counterparts at greater distances from the FIG. 10. Published nontarget doses from various brachytherapy sources The upper and lower bounds produced for the IMRT data in Fig. 5 are also included for reference based on distance to the edge of the treatment field.

14 e404 Kry et al.: TG-158: non-target doses e404 effective dose was found to be dominated by the therapeutic dose; imaging dose (assuming imaging once per fraction) contributed up to 30% of the nontarget effective dose, but was typically much lower depending on the type of imaging used. However, the spatial distribution of the imaging dose may require additional consideration because the imaging field encompasses more of the patient than the therapy dose. The total accumulated imaging dose is not relevant inside the treated volume where the therapeutic dose will dominate, and it is unlikely to be relevant outside the imaging field where it will be only a small fraction of the therapeutic out-of-field dose. However, because imaging fields are typically much larger than radiotherapy fields, there will be tissue well outside the treated volume but still inside a daily imaging field. In these tissues, daily imaging doses could be comparable to the stray dose from the radiotherapy procedure, particularly for high-dose imaging procedures. The need for image-guidance procedures should be justified, particularly in sensitive populations such as children. It should be noted that for image-guidance techniques, there is often interplay between increased imaging and improved therapeutic dose conformity (i.e., reduced PTV volume) that suggests the possibility of optimizing, rather than simply minimizing, the imaging dose. 4. MEASUREMENT APPROACHES 4.A. Photon dosimetry Measurement of nontarget photon (or electron) doses (particularly outside of the treatment field) poses many unique challenges because the radiation field is different outside of the treatment field than within it. The characteristics of the radiation field outside the treated volume are discussed in Section 3.A. Specific considerations include: 1. Dose at the surface: Outside the treatment field, the dose at the surface is increased by stray electrons, so there is a build-down effect instead of a build-up effect near the surface. The dose is elevated by a factor of as much as 5 at the patient surface, and decreases to a depth of approximately d max, below which the dose becomes approximately constant with depth. 79,80 Therefore, if a dosimeter is placed on the patient surface, it will overestimate the dose (by a factor of as much as 5) unless a very superficial dose estimate is desired. If a superficial dose estimate is not desired, the dosimeter should be covered by bolus of a thickness of approximately d max. If measurements are made inside a phantom, this depth consideration generally is not an issue. 2. Energy spectrum: The average energy is much lower outside the treatment field (mean photon energy as low as 0.20 MeV). If the dosimeter s calibration is done in the harder, primary beam (mean photon energy ~1.5 MeV for a 6 MV beam), a dosimeter with a higher effective atomic number than tissue will overrespond to this softer radiation. While this means that measurements will tend to be conservative overestimates, this effect must at least be known, and can be sizeable to the point of unacceptable accuracy unless it is accounted for. 3. Dosimeter dynamic range: Because dose levels outside the treatment field are low, the MU for phantom measurements often must be scaled up to achieve an appropriate reading by the dosimeter. This is clearly not possible for in vivo measurements. 4. Presence of other particles: This section outlines measurement of photon or electron doses. It is important to know and consider if measurements are being made in a mixed field (e.g., photon/neutron), as dosimeters can respond very differently to different types of radiation. 4.A.1. Thermoluminescent dosimeters (TLD) and optically stimulated luminescent dosimeters (OSLD) TLD and OSLD perform well for out-of-field measurements. 184 These detectors are small, safe, and unobtrusive for in vivo measurements. Many phantoms readily accommodate these detectors, especially TLD. Special issues that require consideration include: Energy dependence: If the TLD/OSLD is calibrated in a primary MV beam, the TLD/OSLD signal per delivered dose outside of the treatment field will be greater than expected because of the dosimeter s overresponse to the softer spectrum outside the field. This overresponse is up to 12% for LiF-based TLD (e.g., TLD-100); 74 however, TLD based on other materials (such as Ca) may show substantially greater overresponse. The overresponse is also greater with Al 2 O 3 :C OSLD (e.g., nanodots), ranging from 5% to 31%. 185 Dose level: Calibration (generation of a calibration curve) for TLD/OSLD is done either at a specific dose or over a specific dose range because these dosimeters have nonlinearities in their response. For low doses associated with out-of-field applications, the dose response is quite linear as nonlinearities are most pronounced for higher doses. As nonlinearities are dependent on the relative calibration dose vs the measured dose, attention should be paid to these dose levels to ensure that nonlinearities are either not relevant or corrected for. AAPM Report TG-191 discusses this point further. Contaminating neutrons: Standard TLD-100 overresponds to neutrons. Although this dosimeter is accurate within the treatment field of high-energy photon and particle beams, it is grossly inaccurate outside of the treatment field, overresponding by a factor of as much as In such situations, TLD-100 should not be used. Rather, a neutroninsensitive dosimeter (such as TLD-700) should be used to measure photon doses, and separate neutron dosimetry should be conducted to determine the neutron dose (see Section 4.B). Al 2 O 3 :C OSLD also has a very low neutron cross section and will essentially show only the photon component.

15 e405 Kry et al.: TG-158: non-target doses e405 4.A.2. Ion chambers Ion chambers have many good traits and have occasionally been used to measure nontarget doses. 53,187 One of the main challenges for using ion chambers is that phantoms are typically designed to house other detectors (such as TLD), which makes the use of ion chambers challenging. Application of these devices is further impeded because most ion chambers are also relatively bulky for surface measurements on a phantom or patient, and such surface dosimetry is particularly challenging. Beyond these logistical issues, the following issues should be noted: Energy dependence: Ion chambers typically exhibit favorably small energy dependence, but the energy dependence is highly dependent on the materials of the chamber, particularly the central electrode Farmer chambers and scanning chambers constructed of low-atomic number (Z 13) materials show a nearly flat energy response, either over- or under-responding, but typically by less than 5% down to effective energies of 40 kev (compared to 60 Co). However, microchambers containing high-z collecting electrodes show a dramatic overresponse that typically exceeds 50% at energies around 100 kev (and increases at lower energies). 191 Low readings: To overcome the problem of low readings, it is common to run beam for an extended period of time and at lower charge settings. Potential electrometer drift over this time period should be monitored. High voltage: Ion chambers present a problem for in vivo measurements in that they necessitate the use of biased electronics on or near the patient and management of wires. 4.A.3. Film EBT Gafchromic film has been used to measure doses outside of the treatment field. 114,192,193 Although this provides planar information, it presents challenges. Energy dependence: Gafchromic films have an effective atomic number comparable to that of tissue, but it can be slightly higher or lower so they can over- or under-respond at low energy. This energy dependence (at out-of-field energies compared with a primary 6-MV beam) is small but present. EBT film, for example, was found to underrespond by up to 8% (for EBT), 194 while EBT2 and 3 were found to overrespond by 5 10%. 195,196 These effects depend slightly on dose, color channel, and other scan parameters. 196 Dose level: EBT2 film, in particular, is only indicated for doses > 1 cgy (ISP users guide: Ashland/Static/Documents/ASI/Advanced%20Materials/ebt2. pdf). Doses outside the treated volume may easily be less than this, requiring either up-scaling of the delivered MU or limiting the measurement location to relatively small distances from the field edge. Dynamic range: Depending on the range of measurement locations, doses can span many orders of magnitude. Creation of a calibration curve that can span this range can be challenging 193 although van der Heuvel recently reported on extending the dynamic range of film to low doses A.4. Diodes Diodes have been used to measure many in vivo out-offield doses. 50,198,199 Use of diodes should include the following considerations: Energy dependence: The energy dependence is caused by overresponse from high-atomic-number materials such as silicon, electrode attachments, the protective housing, and build-up. 200,201 The dependence varies substantially between different types of diodes, but is often very pronounced. 201 Just 1 cm outside of the treatment field the energy response was found to be up to 70% different than in-field for an EDD-5 diode. 201 Specific diode dosimeters are typically provided for a certain energy range. It is, thus, important to choose appropriately low-energy diodes for out-of-field dosimetry. Even so, large corrections or uncertainties are likely. Instantaneous dose-rate dependence: The instantaneous dose-rate dependence impacts the diode s sensitivity with its distance from the radiation source and thus is commonly referred to as SSD dependence. This dependence is in the range of 2 7% for different types of diodes, 202 which adds another challenge for out-of-field dosimetry because there are no universal correction factors that can be applied at various distances from the source. Angular dependence: The angular dependence is mainly caused by the detector construction. Cylindrical diodes have relatively small angular dependence of approximately 2% in the axial direction, while the angular dependence of conical or hemispherical diodes can exceed 5%. 198,203 Temperature dependence: The change in sensitivity of commercial diodes with temperature is between 0.1% and 0.54% per C. 204 An appropriate temperature correction factor should generally be applied unless a system with an automatic correction feature is used A.5. Metal oxide silicon semiconductor field effect transistors (MOSFETs) Advantages of MOSFETs include their small size, immediate readout, ease of use, and linear response for large dose ranges. 205,206 However, the following considerations are warranted: Energy dependence: MOSFET dosimeters have energy dependence because of the intrinsic design using non-tissueequivalent silicon materials. The energy dependence is relatively low, around 3% in primary MV beam energy ranges, 205 but results in an overresponse between 50% and 600% to low-energy kv photons It should be noted that MOS- FETs with different sensitivities may have different energy dependencies because of variations in the percentage of nontissue-equivalent material used in the intrinsic design. Dose calibration for the photon spectrum at the point of measurement outside the treatment field can be crucial.

16 e406 Kry et al.: TG-158: non-target doses e406 Angular dependence: MOSFET dosimeters have angular dependence because of their asymmetric design. The angular dependence in the radial direction is the most studied and is in the range of 2 6%. 205,208 The data on angular dependence in the longitudinal direction are sparse, and one study shows 16% variation. 211 Limited lifetime: MOSFET dosimeters have a limited lifetime with a maximum accumulated dose. This is not a major concern because out-of-field measurements typically involve low doses. Sensitivity: Even a high-sensitivity MOSFET dosimeter in a high-sensitivity voltage setting can measure dose only down to the order of 1 mgy. Appropriate dose levels are necessary. 4.B. Neutron dosimetry Neutron dosimetry is a particularly challenging pursuit. It is very easy to make mistakes that result in errors of several orders of magnitude. In order to perform neutron dosimetry correctly, it is essential to understand the properties of not only the detector but also the neutron field being measured. Most importantly, neutron detectors are sensitive to neutrons only over a particular energy range. Unlike with photon dosimetry, the neutrons that are detected are routinely of vastly different energy than the neutrons that contribute most of the dose to the patient. Because detector responses to different neutron energies often vary by several orders of magnitude, it is critical to know the detector response function and how it relates to the neutron spectrum being measured. This issue becomes extremely important in the presence of a patient or phantom, which strongly influences the energy spectrum of the neutrons (Fig. 11). The calibration of the detector (which is typically done in air) is usually incorrect if it is used in vivo or even placed on the surface of a phantom or patient. Such applications are difficult and require careful consideration of the energy of the neutrons at the measurement location compared with the energy of the neutrons during calibration. Neutron dosimeters with location-specific calibration factors have been proposed as a solution to this very difficult problem. 86,212 While neutron dosimetry is often conducted in air to provide the most accurate measurements of neutron spectrum or dose, the dose equivalent in air (or at the surface of a patient) substantially overestimates the dose to organs within the patient because of the very sharp depth-dose dependency of neutrons. An estimate of the dose in the patient can be made by propagating the dose in air/at the surface to dose in a phantom or patient based on neutron percent depth doseequivalent curves. 91,92 As an additional consideration, because primary proton and photon beams are pulsed, passive detectors are generally preferred for measurements. Active detectors, such as 6 LiI detectors and BF 3 proportional counters, are subject to pulse pile-up effects. This section describes the most common types of neutron detectors. Broadly, this section is divided into thermal neutron detectors (such as activation foils and TLD) and accompanying thermalizing systems (Section 4.B.1) and fast neutron detectors, such as bubble detectors and track etch detectors (Section 4.B.2). 4.B.1. Thermal neutron-based detectors Thermal neutron detectors: Most neutron detectors have the highest interaction cross section to thermal neutrons, where absorption is often followed by emission of various charged particles. Unlike the neutrons, these emitted charged particles can be easily measured to generate signal. The most common active thermal neutron detectors use boron, lithium, or helium gas as the principal absorbing material to exploit the 10 B(n,a), 6 Li(n,a), or 3 He(n,p) reaction, respectively. The count rate then indicates the number of interactions, that is, the quantity of thermal neutrons. Other active detectors exist, including those based on electronic upset events where the frequency of electronic upsets is proportional to the neutron flux. 213,214 The most common passive thermal neutron detectors are TLDs and activation foils. The TLDs are used in pairs, one sensitive to both photons and thermal neutrons (TLD-600) and the other sensitive only to photons (TLD-700). The neutron component is determined by the difference in the response in the two types of materials. Activation foils are thin metal discs that are activated by thermal neutron absorption. Gold foils, 197 Au, form 198 Au (T1/2 = 2.7 days), and the resultant 411 KeV gammas can be measured. Indium foils, 115 In, foils have a higher (n,c) cross section, but are not used as frequently for medical physics applications because of the short (54 min) half-life of the excited 116m In. It is essential to recognize that these detectors respond primarily to thermal neutrons, and they must be used in conjunction with a neutron moderating system that thermalizes fast neutrons. Moreover, this moderating system must be incorporated into the calibration (ie, the calibration is specific to a neutron spectrum and moderating system). For a detector such as a TLD-600/700 pair, it is not possible to get meaningful neutron dosimetric information by simply placing them, eg, in an anthropomorphic phantom and using a single calibration factor. Even though the phantom will moderate the fast neutrons, the thermal neutron fluence varies drastically at different depths in the phantom (Fig. 11). The differences in thermal neutron fluence will lead to substantial changes in the detector response that do not correspond in any meaningful way to dose to the patient (which is dominated by fast neutrons). Therefore, such measurements yield the relative thermal neutron fluence and not the dose distribution. The magnitude of this error can be crudely estimated by assuming that the entire signal is from thermal neutrons and the entire dose is uniformly from fast neutrons. Based on the curves in Fig. 11, the error in the dose estimate associated with using an inair calibration for in-phantom measurements can be substantial. For example, the dose is comparable between the

17 e407 Kry et al.: TG-158: non-target doses e E-09 Neutron energy spectra at different depths in tissue (cm) 3.0E E-09 Air 2.5E-09 Fluence (A.U.) 2.0E E E E E E E E E+00 1.E-08 1.E-06 1.E-04 1.E-02 1.E+00 E (MeV) 0.0E+00 FIG. 11. Variations in the neutron spectrum from an 18 MV photon beam as a function of depth in tissue. Included is an in-air location, along with the same location at the surface (0.1 cm depth) of a phantom. Other depths in the phantom are also included up to 19.5 cm depth. Of note, the fast neutron peak (which deposits the dose) decreases sharply with depth, while the thermal neutron peak (which is often responsible for depositing signal in the detector) shows a very different dependence on depth, increasing up to 4.5 cm depth and still being notably present at 19.5 cm depth. Data are compiled from Kry et al. 91 surface location and the in-air location (similar number of fast neutrons), but the signal (number of thermal neutrons) is more than three times as large at the surface location, so the signal based on an in-air calibration would overestimate the dose at the surface by approximately a factor of 3. The error would be an overestimation by more than a factor of 25 at 4.5 cm depth and a factor of 250 at 19.5 cm depth. Of note, these differences are not changing uniformly with depth (the thermal peak changes independently and differently than the fast peak). Therefore, relative measurements in a phantom do not describe the relative dose. It is possible to avoid these substantially erroneous results by requiring that the detectors are separately calibrated for each specific measurement location in the phantom (accounting for changes in the neutron spectrum and the response of TLD-600). Such location-specific calibrations have been successfully employed. 86,212 Thermal neutron detector moderator systems: Thermal neutron detectors only detect low-energy neutrons. However, dose deposition is dominated by high-energy neutrons, which generally makes these fast neutrons the neutrons of interest. Therefore, thermal neutron detectors must be used in combination with moderators to determine information about the higher energy neutrons. Neutron moderators are composed of low-atomic number material (frequently hydrogenous polyethylene) that surround the thermal detector. Incident neutrons undergo elastic scattering, primarily off of hydrogen atoms, until the energy is sufficiently lowered so that it can be captured by the thermal neutron detector. This method of detection is referred to as moderate and capture. The volume of moderating material determines the energy range of neutrons that will be thermalized and thereby detected. Larger volumes of moderating material will moderate higher energy neutrons, allowing for their detection. However, such moderating systems are generally useful only for neutron energies of up to ~10 MeV; at greater energies, additional moderation poorly differentiates neutrons of different energies. Therefore, to measure higher energy neutrons, a high-z moderator detector system is required, in which the thermal neutron detector is surrounded by layers of both high-z and low-z materials. The incident neutrons interact in the high-z material, which has a large (n,xn) cross section and functions as a neutron multiplier, increasing the number of neutrons. The low-z layer serves to decrease the energies so that the neutrons can be captured by the thermal neutron detector. The following two subsections describe types of thermal neutron detector moderator systems. Rem meters and extended-range Rem meters: Neutron Rem meters, commonly referred to as neutron survey meters, are used for real-time measurement of neutron dose equivalent. 215,216 They include an active thermal neutron detector surrounded by a moderator. Because Rem meters have active detectors, pulse pile-up is a concern and these detectors are generally not well suited to measurements in or near the primary beam. Commercially available Rem meters [including

18 e408 Kry et al.: TG-158: non-target doses e408 the Victoreen neutron survey meter (Fluke Biomedical, Everett, WA, USA) and the ThermoElectron NRD neutron survey meter (Thermo Fisher Scientific, Inc., Waltham, MA, USA)] often use BF 3 proportional counters to detect thermal neutrons. The response per unit dose equivalent is shown in Fig. 12(a). 215 At 1 MeV, the response is close to 1, and these detectors do not overrespond or underrespond by more than 50% between 0.1 and 10 MeV. At energies greater than 10 MeV, however, the response functions for both detectors rapidly decrease with increasing energy. This underresponse is by a factor of ~6 when the energy reaches 100 MeV. In the context of photoneutron surveys, traditional Rem meters are a reasonable choice for measurements, but for neutron measurements in proton or other ion beams, this detector would greatly underestimate the high-energy neutrons and are not recommended. Because of the poor response of traditional Rem meters to energies greater than 10 MeV, high-energy Rem meters have been developed: the PRESCILA (Proton Recoil Scintillator Los Alamos) and the WENDI-II (also referred to as the Smart WENDI or SWENDI) use lead and tungsten, respectively, to extend the high-energy response 217 [Fig. 12(b)]. These detectors have a more accurate response than traditional Rem meters at energies greater than 10 MeV. However, both of these detectors still demonstrate substantial variability in their response to neutrons of different energies. Unless careful corrections based on the known spectrum are applied, measured doses will likely be less accurate than 50%. Bonner sphere spectrometers and extended-range bonner sphere spectrometers: Bonner spheres use a thermal neutron detector in a series of polyethylene moderating spheres, or Bonner spheres, ranging in diameter from 2 to 12 inches (2, 3, 5, 8, 10, and 12 inches). 218 Each detector moderator combination has a different response to a given neutron energy or energy spectrum (Fig. 13). These systems can be used to measure neutron spectra ranging from thermal up to approximately 20 MeV. At energies greater than 10 MeV, the responses rapidly decrease and become nondifferentiating with increasing energy, making this detector unsuitable for neutron measurements around proton facilities. Measurements must be repeated for each detector sphere combination. Then, from these data, a neutron spectrum is determined by mathematical deconvolution ( unfolding ); additional details are available in the literature Bonner spheres were originally designed for use with 6 LiI detectors but can be used with any thermal neutron detector, including BF 3, gold and indium activation foils, and TLD pairs. High-Z shells are typically employed to increase the sensitivity of Bonner spheres to high-energy neutrons (> 20 MeV), such as those encountered around proton or carbon ion beams that extend up to hundreds of MeV Use of these shells creates an extended-range Bonner sphere spectrometer, which can be used in conjunction with any of the thermal neutron detectors described above to measure secondary neutrons from proton therapy. 94,108,110, Similar to standard Bonner sphere spectrometers, response functions for each detector sphere combination must be determined, accounting for the high-z shell. During a neutron measurement, the irradiation must be repeated with each detector sphere combination, and data must be unfolded to determine a neutron spectrum. 223,227, A related but streamlined detector was recently developed by Bedroni et al. 234,235 whose Spherical Spectrometer (SP2) condenses the functionality of extended range Bonner sphere systems into a single moderator comprised of a polyethylene sphere, a lead shell, and multiple thermal neutron detectors 236 The response functions of this detector indicate it would be suitable for measurements in particle therapy. 236 Bonner spheres are generally considered the standard for neutron detection, and measurements around high-energy photon linacs can be made with uncertainties on the order of 3% in fluence and 7% in dose. 87 However, these systems are difficult and inefficient to use, require substantial ancillary FIG. 12. (a) Response functions for two common neutron Rem meters to different neutron energies, showing the large variability in response, particularly for high neutron energies. Figure from personal communication, Richard Olsher, Los Alamos National Laboratory (retired), as shown in Ref. [215]. (b) Response function for high-energy neutron Rem meters. Reproduced with permission from Ref. [217].

19 e409 Kry et al.: TG-158: non-target doses e409 equipment and neutron calibration capabilities, require expertise in spectra unfolding, and can be used only for in-air measurements. Thus, they are best suited for characterizing neutron spectra at new therapy facilities and for benchmarking Monte Carlo models for neutron simulations. 4.B.2. Fast neutron detectors Bubble detectors: A bubble detector is a small sealed tube with tiny superheated droplets of liquid dispersed throughout a polymer (Fig. 14), and have been recently reviewed. 237 When neutrons strike a droplet, secondary charged particles are emitted that vaporize the droplet, producing a bubble, which remains fixed in the polymer. 215 Neutron dose can then be determined by counting the bubbles and applying a calibration factor. The advantages of bubble detectors over other neutron detectors are that they are very easy to use, are reusable, and can be read instantaneously. Disadvantages of bubble detectors include energy dependence, loss of linearity at high doses, and the potential for spurious bubbles. These detectors should not be used in the primary photon beam because spurious signal has been observed within the polymer, overestimating the true result. After irradiation, the number of bubbles is typically scored manually and dose determined on the basis of a calibration factor provided by the manufacturer. Irradiations should be balanced such that a reasonable number of bubbles are produced. As statistical accuracy is generally lower for these detectors than other neutron detectors, multiple readings at each location may be appropriate (either with multiple detectors or repeat irradiations). Bubble Technology Industries (Chalk River, Ontario, Canada) recommends exposing the detectors to a neutron dose sufficient to achieve approximately 100 bubbles to maintain the product specified accuracy of 20%. The stated accuracy must be considered very carefully. Generally, it is treated as a best case scenario, as accuracy may be highly dependent on differences between the calibration and experimental neutron spectra. Commercially available bubble detectors, e.g., BD-PND (BTI Bubble Technology Industries, Ontario, Canada) have only a somewhat uniform response (per unit dose equivalent) to neutrons over the energy range of approximately 100 kev 20 MeV 238 [Fig. 14(b)]. And at energies less than and greater than this range, the response decreases rapidly. While this energy range encompasses the majority of neutrons produced during photon therapy, room-scattered neutrons (which have low energy) are underestimated by bubble detectors because they are largely unresponsive to these low-energy neutrons. As such, they are of little use in a vault maze or outside a treatment room, where the average neutron energy is low. Caution should also be used before using bubble detectors in a phantom. Inside a phantom, neutrons lose energy rapidly and bubble detectors will, therefore, under-respond to them (Fig. 11). Unless the calibration of the detector is able to account for the spatial changes in the spectrum, dosimetric uncertainty will increase, potentially substantially. Similarly, bubble detectors are not well suited for measuring neutrons from proton therapy because their response to high-energy neutrons (> 20 MeV) is only about 1/3 of their response to neutrons with energy around 1 MeV. 239 The details described above are largely specific to BD- PND (Bubble Technology Industries), the most common type of bubble detector, which is recommended for personnel monitoring and neutron surveys around high-energy photon linacs. Bubble detectors designed for thermal neutron detection are also available, as well as sets of bubble detectors with FIG. 13. Response function for a Bonner sphere and gold-foil system. Curves show the response of the foil to neutrons of different energies incident on a moderating sphere. Different curves are for moderating spheres of different diameters. Based on data from Ref. [229].

20 e410 Kry et al.: TG-158: non-target doses e410 FIG. 14. (a) Bubble detectors for measuring neutron dose equivalent (from reproduced with permission). The top detector is unexposed while the bottom detector shows the bubbles formed in response to neutrons. (b) The response (per fluence) of BD-PND to neutrons of different energies, including a drop in response to neutrons of low or high energy. 238 Reproduced with permission. neutron reaction thresholds that vary between 0.1 and 10 MeV, which can be used to determine neutron spectra over this energy range. The threshold bubble detectors require data unfolding to determine neutron spectra, and generally produce cruder results than Bonner sphere systems. 237 Track etch detectors: Solid-state nuclear track etch detectors are typically made of polymers that include mainly hydrogen, carbon, and oxygen (often CR-39 polymer). When a neutron interacts in these materials, the recoil nuclei from the constituent materials leave behind microscopic damage trails or tracks. Chemical processing (etching) is used to enlarge the tracks. The track density can then be scored and neutron dose determined by using a calibration factor. Multiple polymers sensitive to different energy ranges may be used; for example, the Neutrak 144 (Landauer, Inc., Glenwood, IL, USA) uses two different radiators to increase sensitivity to fast and thermal neutrons. While the Neutrak 144 can detect fast neutrons from 40 kev to over 35 MeV, its response to fast neutrons decreases with increasing energy and rapidly falls off at energies greater than 10 MeV (Fig. 15a). Most track etch detectors are not suitable for measurement around proton (or carbon) beams because of the lack of response to high-energy neutrons. PN3 detectors (Thermo Fisher Scientific, Waltham MA, USA) have been successfully used in proton beams (Fig. 15b), but characterization and depth dependence in a phantom showed unique characteristics compared to the neutron fields encountered in highenergy photon therapy. 212 Furthermore, track etch detectors have additional uncertainty because photon- or protoninduced nuclear reactions can result in additional particle tracks that will artificially inflate the number of tracks and should only be used with caution inside the primary photon or proton beam. 240 Track etch detectors are very sensitive and are, therefore, suitable for use outside the primary beam in high-energy photon therapy. Track etch systems such as Neutrak 144 are also suitable for dose or personnel monitoring outside the treatment vault for both photon and proton therapy, as these detectors retain sensitivity to thermal neutrons. Because these detectors are intended for personnel monitoring, they are typically calibrated on the surface of a phantom. Therefore, using them in air or at depth in phantom will increase the measurement uncertainty (potentially substantially), as the fluence and spectrum are both substantially altered by the absence of scattering material or additional build-up. Calibration of these detectors for in vivo measurements is possible through application of multiple detector calibration factors that vary as a function of position in the phantom/patient, thereby accounting for the changes in the neutron spectrum. The variation in neutron spectrum with depth in a phantom was found to make a 30 40% difference in the calibration factor for PN3 track etch detectors. 86,212 Tissue-equivalent proportional counters (TEPC): The relative photon and neutron components of a mixed radiation field can be distinguished on the basis of the LET of the charged particles generated by the incident radiation with a low-pressure proportional counter. Unique to tissue-equivalent proportional counters (TEPCs), the output from these systems is primarily microdosimetric quantities (e.g., the LET of the incident radiation). TEPCs have been applied to measure single-event distributions of many types of ionizing radiation, including photons, neutrons, and heavy charged particles. These detectors can and have been used to conduct measurements around proton and carbon therapy beams. 163,241,242 The event-size spectra obtained during measurements can range from around 0.1 kev/lm to > 1000 kev/lm. The absorbed dose contributions from various types of secondary particles can be differentiated via the microdosimetric spectrum because different secondary particles have different lineal energies in the active volume. TEPC can also be used for determination of absorbed dose as well as calculation of radiation quality factors, kerma factors, and equivalent doses. The wall of this type of detector

21 e411 Kry et al.: TG-158: non-target doses e411 FIG. 15. Response of (a) the Neutrak 144 and (b) the PN3 track etch detectors to neutrons of different energies. The individual data points are underlying experimental data. Note that the x-axes in (a) and (b) cover very different spans of neutron energy. Panel a is courtesy of Landauer Inc.(Glenwood Illinois), while panel b is reproduced with permission from Ref. [212]. can be made of various materials, for example, for the determination of kerma factors TEPC detectors are usually calibrated with an internal alpha particle source, 244 Cm, where the energy deposited in the cavity, expressed in kev/ lm, can be obtained from the stopping power of the alphas and the counter diameter. 242, This calibration is generally the principal source of uncertainty in this type of measurement, suffering from a poor resolution of the calibration peak as well as uncertainties associated with the stopping power value and mean chord length. Given that the signal from this type of detector depends on the measurement of each energy deposition event, the flux is a major concern when using this detector in radiotherapy environments because of pulse pile-up and dead time effects. Trying to manage these effects is a major challenge in radiotherapy as instantaneous fluence rates are often high, even outside the treatment volume. 4.C. Phantoms Out-of-field doses have been measured in large scanning tanks, 53 plastic slab phantoms put together in an anthropomorphic shape, 112,129,187,249 and adult and pediatric anthropomorphic phantoms. 2, A more realistic humanoid phantom is conceptually more appealing and better suited for study of specific treatment plans as it captures the complex contours of human anatomy and heterogeneities such as the lungs and bony anatomy. Such phantoms also facilitate definition of organ locations, including appropriate distances between the treatment site and organs of interest. The location of such organs relative to the field edge can substantially influence the dose received by such organs. 251 However, while a more realistic phantom seems to be more appropriate, the impact, if any, on dose accuracy has not been quantified by a direct comparison between measurements using a plastic slab phantom, or even a water tank, and an anthropomorphic phantom. Because the anatomical structures in children are closer in proximity to the treated volume, the nontarget dose in the same tissue will be higher in children than in adults. 187 Phantoms are usually incompatible with neutron dosimetry. Although there can be obvious difficulties with using large neutron dosimeters (e.g., Bonner spheres) inside a phantom, the more serious problem is the impact of the phantom on the neutron spectrum and subsequent response of the dosimeter (see Section 4.B). The phantom presence may be further complicating because the material composition of the phantom is relevant, particularly if neutron production is important (i.e., proton or ion therapy). The phantom must have reasonable neutron production cross sections in order to mimic the production of internal neutrons; otherwise errors up to 30 50% are possible. 255, COMPUTATIONAL APPROACHES Calculations can be less laborious and cumbersome than measurements. However, sufficient care is needed to ensure that the calculations being performed are accurate and have been benchmarked correctly. 5.A. Treatment planning systems Radiotherapy treatment planning systems (TPS) are not commissioned or designed for out-of-field dose calculations. 257 Even though the TPS displays doses at points far from the treatment field, it should not generally be used for estimating out-of-field dose. Large discrepancies have been reported between measured doses and those calculated from photon-beam planning systems, 2,3,78,258 even for modern convolution/superposition and analytic anisotropy algorithms. These differences can exceed 30% of the local dose as close as 3 cm from the field edge, and differences increase dramatically at greater distances. These errors have been found across planning systems and across delivery systems.

22 e412 Kry et al.: TG-158: non-target doses e412 Most commonly, photon-beam planning systems have been reported to underestimate the true out-of-field dose, 2,3,114,258 although overestimations have been found. 78,140 Large and comparable errors have been reported for both simple conformal fields 2,78,258 and IMRT fields. 3 The error in the planning system calculation is attributed to underestimation of patient scatter and head leakage and substantial underestimation of collimator scatter. 3 Given these fundamental limitations in the planning system calculations, it is unclear if dose calculation accuracy could be improved by additional modeling efforts (e.g., extending the beam profile measurements used for commissioning beyond 5 cm outside of the field edge). For proton therapy, close to the field edge modeling is difficult for analytical dose calculation methods for beam scanning because of the contribution of the nuclear halo. 259 Farther from the field edge, for proton therapy (as with photon therapy), neutron doses are not considered in any available commercial treatment planning software, substantially limiting the ability of proton planning systems to predict nontarget dose. TPS dose calculations are often used for volumetric assessments (mean organ dose). In many cases, organs span the edge of the treatment field, being partially in-field and partially out-of-field. If the majority of the organ is within the 5% isodose line, particularly if parts of the organ receive a large fraction of the prescribed dose, then the mean dose reported by the TPS may be reasonably accurate. However, if the majority of the organ is outside the 5% isodose line, the mean dose reported by the TPS is likely in error. 258,260 5.B. Analytical models for estimating out-of-field dose Analytical models have been historically available for the calculation of out-of-field dose from photon therapies. For example, data from AAPM Report TG have been used in combination with mathematical phantoms to determine the out-of-field dose to different organs in a patient. 261 A software program called Peridose was developed to calculate out-of-field dose with a user-friendly interface. 56 Other, similar models also exist (e.g., as described in Refs. [89,253,254]). In general, these analytical tools are accurate within approximately 30%, 56,262 although much larger errors are possible. Recent studies have developed more refined analytical models for predicting out-of-field dose. These models have increased complexity (i.e., more parameters), but provide improved accuracy in dose calculation and can estimate doses from IMRT treatments. 263,264 Neutrons have not been included in any of these analytical models to date. Neutrons have been included in analytical models for the calculation of out-of-field neutron doses from proton therapy. This has included a model of neutron equivalent dose in air, 161 and in a phantom, 265 and was able to agree with Monte Carlo results within 17% on average. 5.C. Monte Carlo methods Several different Monte Carlo codes have been used to calculate out-of-field dose, including MCNP5, MCNPX, 266,267 Geant4, 268 EGSnrc, 269 and FLUKA. While any code may be a viable option, all models must be validated (i.e., benchmarked) against appropriate measurements. Uncertainty in Monte Carlo calculations should be reported with adequate consideration for both systematic and random uncertainties. Random uncertainties can be minimized through variance reduction techniques, whereas systematic uncertainties are more likely to be a function of the modeling and therefore harder to reduce. For out-of-field dose calculations, both sources of error may be relatively large compared with in-field calculations. When reporting uncertainties, both the random and systematic uncertainties should be quantified and quoted unless one source is clearly dominant. Monte Carlo models have been used for out-of-field dose calculation for various treatment techniques and geometries. Detailed models of the treatment head include head shielding components and structural components, which are needed to ensure that leakage radiation and head-scattered radiation, as well as neutron production and scattering during high-energy photon or proton therapy, are appropriately modeled. Several such detailed models have been developed for photon (including neutron) transport 79,93, as well as proton transport. 101,105,160, However, development of detailed models is difficult because manufacturer blueprints are often not readily available, and coding and validation are laborious. More simplified models have been considered, such as incorporating only beam-line components, but such models are less accurate and may be inappropriate, as described in the next subsection. 5.C.1. Photon transport Detailed Monte Carlo models (which include head shielding and structural components) have shown good agreement with measurements even beyond 50 cm from the field edge. 79,93,270 Much simpler Monte Carlo models that include only beam-line components have also been used. 78,276 While simple models are capable of accounting for patient scatter outside the treatment field, they do not allow for proper modeling of head leakage and, potentially, collimator scatter. Correspondingly, good accuracy in beam-line models can be achieved out to ~15 cm from the field edge, but at greater distances, there is increasingly poor accuracy. Any Monte Carlo model used for dose evaluation outside of the treatment field should be explicitly validated against detailed out-of-field measurements. In-field PDDs and profiles do not constitute validation of out-of-field accuracy; as such parameters depend overwhelmingly on the primary radiation beam, which does not guarantee that secondary components (head scatter and leakage) are adequately modeled. A further validation consideration is relevant near the phantom surface (at depths shallower than the beam s d max ), as the

23 e413 Kry et al.: TG-158: non-target doses e413 out-of-field dose near the surface is substantially elevated by electrons. 79,276 Accuracy of calculated dose near the phantom surface should also be confirmed against measurements when relevant. Doses associated with photons are optimally calculated with a dose scoring tally (electron energy deposition). However, the computation time needed to achieve good statistics for such tallies can be prohibitively long. Photon kerma can be calculated instead, and it has been found to offer accuracy within 3% over most of the out-of-field region. 79,93 Exceptions to this agreement are regions without electronic equilibrium, including the surface of the patient, the field edge, and tissue interfaces. If the dose is required in these regions, electron energy deposition must be used. 5.C.2. Neutron transport For high-energy photon beams, most neutron simulations have included detailed linac head models. 93, ,277 Such models have shown agreement with fluence measurements within 20% on average 93, and with measured doseequivalent metrics within 10% on average. 93 As a simplification, it is also possible to represent the linac as a beam-line structure (or neutron source) encased in heavy metal shells or an otherwise simplified shielding structure. 272,278,279 Such an approach makes the modeling much easier, but as it is a simplification, it results in poorer agreement with measurement. Fluence calculations and dose equivalents have been found to agree within ~40% as compared to complete models and measurements. 272,279 Finally, some simulations have included only the beam-line components of the accelerator and neglected all head shielding and structural components entirely. However, this is an even larger simplification that fails to account for many important physical processes: the neutron production in the shielding and structural components are not modeled, the degradation of the fast neutron spectrum that occurs in the shielding components is not modeled, and the focusing of scattered neutrons through the collimator opening and toward the isocenter is not modeled. 277 Consequently, errors in fluence of a factor of 2 3, and errors in the simulated spectra, have been observed with this approach as compared to complete models. 277 This report, therefore, recommends including head shielding (at least in a simplified form) in any model of an accelerator used to simulate neutrons. Neutron production in proton and ion beams has also been simulated by Monte Carlo methods. 101,105,160,161,273,274 Some of these studies included detailed models of treatment heads. As with photon simulations, it is reasonable to expect that the inclusion of more detail will increase the accuracy of the simulation. Accurate neutron modeling in a proton environment is more challenging than in photon therapy because measurements suitable for validating the Monte Carlo model are more difficult and uncertain in proton therapy (see Section 4). Whether photon or proton beams are being considered, the majority of neutrons are generally produced in the beamline components. Because neutrons may experience a large number of scattering processes before being captured, scattering from the treatment vault walls and possibly even ancillary equipment (such as the treatment couch) will typically affect the neutron fluence. Kase et al. 272 found that approximately 30% of the neutrons in the patient plane were scattered neutrons, supporting the need to model room-scattering structures. Notably, these neutrons are of lower energy and may, therefore, contribute only around 5% of the neutron dose to the patient. 89 Therefore, inclusion of a description of vault walls may be appropriate (particularly for model validation), but this can generally be done in a simplified manner. Any developed Monte Carlo model should be validated against measurements, and neutron Monte Carlo simulations require explicit validation against neutron measurements. For high-energy photon beams, good agreement with photon measurements does not guarantee good agreement with neutron measurements. Practically, the optimal source description for modeling neutrons may be somewhat different than for photons. Often, a compromise must be struck between photon and neutron agreement. For any neutron-generating beam line, the model is best validated against measurements in air. However, because such measurements require specialized equipment and are prone to error, a reasonable approach is to validate a Monte Carlo model against neutron measurements in the literature (e.g., fluence and spectra in Ref. [87]) if in-house measurements are not feasible. Once the model is validated against in-air measurements, it can be used to calculate neutron doses throughout a patient or phantom geometry. 5.D. Phantoms Analytical dose calculation models have often been paired with simple stylized phantoms, that is, phantoms made of simple geometrical shapes. In these cases, dose is typically overlaid on top of a geometrically defined phantom [e.g., in Ref. [254] and Fig. 16(a)]. Often, more complex phantoms are desired, often in an effort to better define organs, and to more closely mimic actual human geometry. Stylized phantoms [e.g., Fig. 16(b)] 95,257,259,261,273, are geometrically defined but could still have a large number of structures. 93,246,264 and can represent males, females, and different ages. 281,282 These models often allow great flexibility in terms of scaling of patient size and/or weight, although they are limited by being based on simple geometrical shapes, rather than on the more complex human anatomy, which results in errors in calculated doses. 283,284 A second class of phantoms are voxelbased models which provide a more realistic representation of the human body as each voxel corresponds to a unique tissue type and organ identification. This includes VIP-Man [Fig. 16(c)], RPI-P pregnant female models [Fig. 16(d)], and age-specific pediatric models [Fig. 16(e)]. 101,268,273,285 While voxel phantoms surpass stylized phantoms in terms of anatomic accuracy, they are less flexible in terms of permitting changes to body posture, size, weight, or internal organ shape, size, position, or location. Consequently, voxelized

24 e414 Kry et al.: TG-158: non-target doses e414 phantoms still struggle to model a specific patient or patient population. The trade-off between the stylized and voxelized phantoms can lead to notable differences in calculated doses; differences of as much as 150% have been observed in some organs. 284 Recently, a third class of phantoms, hybrid phantoms, has been developed to preserve anatomical accuracy while overcoming the shortcomings of voxelized phantoms. Hybrid phantoms combine voxel data with surface equations to produce a more realistic anatomy. In these models, the boundary of an organ can be adjusted to a desired shape and volume using patient-specific images and deformable image registration, such as the NURBS (nonuniform B-spline fits) method. 286,287 Out-of-field organ doses have been investigated successfully using hybrid models for photon therapy 288 and proton therapy. 289 While these models are the most flexible, they are also the most complex to manage. When a model is used to represent a patient or patient population, the phantom should match the patient in size and other characteristics to the extent reasonable. Major differences between a reference phantom and a specific patient may include the size and position of the organs within the patient, the girth of the patient, and the height of the patient. In photon therapy, girth matters only by a few percent, 290 while organ position and size only matters in terms of distance from the organ to the field edge. 251 Height matching between the patient and the phantom is, therefore, important, while weight matching is much less so. Similar results have been found in proton therapy. 101,291 In general, this issue would benefit from further study for both proton and photon therapy; however, patient height (age) appears to be the most important parameter to consider when matching a phantom to a patient; patient girth is secondary, and organ size and location appears to be only a tertiary consideration. When using a CT dataset in a Monte Carlo dose calculation, the conversion of CT number to material composition may be relevant in proton therapy. Material definition will affect neutron production cross sections and could affect neutron transport, which would be most relevant for scanning beam techniques where patient-produced neutrons are most important. This effect can reach 30 50% depending on material definition. 255, TECHNIQUES TO MINIMIZE NONTARGET DOSE 6.A. Reducing the target volume Reducing the size of the CTV or PTV can be one of the most potent options for reducing the dose to nontarget structures. This directly reduces the volume of tissue receiving a high dose of radiation. The reduced high-dose volume is particularly relevant when second cancers are of concern because of the high incidence of second cancers in high-dose regions. Although the primary benefit is likely from reducing the volume receiving a high dose, reducing the target volume has additional benefit in that the field edge is moved further away from nontarget structures, thereby decreasing the dose they receive. Reducing the target size also reduces the field size and thereby reduces the amount of scattered radiation in the patient. All of these factors reduce nontarget dose. Clinically, the substantial challenge, of course, is to be able to safely reduce the target volume. In clinical practice, the CTV has been reduced in some clinical situations. In Hodgkin s lymphoma, for example, the field is often reduced from an entire mantle to the involved regions only. 294 Similarly, testicular treatments now tend to involve smaller targets than historical treatments. There has also recently been investigation into total-marrow irradiation as a replacement for total-body irradiation, with a substantial reduction in the dose to healthy tissue. In general, however, reducing the volume of the CTV is a very challenging clinical decision. More achievable is reducing the PTV volume by reducing the size of the PTV margin. This could be achieved by, for example, improved localization and/or immobilization. FIG. 16. Various computational phantom types. (a) Stylized phantom used for late-effect studies as in Ref. [254]. (b) Stylized MIRD phantom as described in Ref. [282]. (c) Voxelized VIP-Man described in Ref. [292]. (d) The RPI-P9 voxelized model for a 9-month pregnant female 293 and (e) UF voxelized pediatric phantoms. 284