10 Intensity-Modulated Radiation Therapy Daniel A. Low, Wei Lu, James A. Purdy, Carlos A. Perez,

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10 Intensity-Modulated Radiation Therapy

Daniel A. Low, Wei Lu, James A. Purdy, Carlos A. Perez, and Seymour H. Levitt

D. A. Low, PhD, W. Lu, PhD, C. A. Perez, MD

Department of Radiation Oncology, Mallinckrodt Institute of Radiology, Washington University School of Medicine, 4921 Parkview Place, St. Louis, MO 63110, USA

J. A. Purdy, PhD

Department of Radiation Oncology, University of California Davis Medical Center, Sacramento, CA 95817, USA

S. H. Levitt, MD

Department of Therapeutic Radiation Oncology, University of Minnesota, Minneapolis, MN 55455, USA

10.1

Introduction

The 1990s saw dramatic changes in radiation therapy treatment planning and treatment delivery, driven largely by advances in computer hardware and software. These advances inspired the development

of sophisticated three-dimensional (3D) radiation therapy treatment planning (3D RTTP) and delivery systems, making practical the implementation of 3D conformal radiation therapy (3D CRT). The purpose of 3D CRT was to conform the prescribed dose distribution to the 3D target volume (cancerous cells) shape while simultaneously minimizing the dose to neighboring normal patient structures (Meyer and Purdy 1996). Conventional 3D CRT delivery was implemented by conforming the incident-beam portal outlines to the target volume projections for a user-specified set of beam directions or during rotational beam delivery. The radiation beams had uniform intensity, or where appropriate, the intensity was varied by beam modifiers such as compensating filters or wedges. This treatment method was referred to as conventional 3D CRT.

3D CRT provides excellent conformation of the dose distribution with the tumor targets. By match- ing the radiation portal shape to the projected tumor shape, the radiation beams produces a dose distribution that encompasses the tumor shape. Because the radiation dose distribution is being delivered by relatively homogeneous fluences, the common practice is to assign the dose delivered by each single beam as the dose delivered by that beam at the linear accelerator isocenter. The isocenter point of each beam is placed at a common location within the patient’s tumor. This allows a straightforward mechanism for categorizing the dose delivered to the patient, namely the sum of the individual radiation beam isocenter doses. This process helps to make the prescription and monitor unit calculation protocols relatively straightforward.

Even though conventional radiation beams are labeled as homogeneous, the dose distribution has slight variations throughout the portal, with a rapid drop-off near the projected portal boundary.

When coupled with variations in the patient’s surface contours and internal heterogeneities (defined as variations in the patient’s electron density distribution relative to water), the total dose delivered to the target is heterogeneous. Prescription protocols,

CONTENTS

10.1 Introduction 203

10.2 Basic Physical Principles of IMRT 207 10.3 Inverse Treatment Optimization

(Cost Function and Search Algorithm) 208 10.4 IMRT Treatment Delivery Systems 210 10.4.1 Historical Review 210

10.4.2 Arc-Based Fan-Beam Dynamic MLC IMRT 211 10.4.3 Fixed Portal Cone-Beam Dynamic MLC 211 10.4.4 Fixed-Portal X-ray Compensating Filter IMRT 211 10.5 IMRT Quality Assurance

(Commissioning and Clinical) 212 10.5.1 Dosimetry System 213

10.5.2 Commissioning 214

10.5.3 System Quality Assurance 214 10.5.4 Monitor Unit Verification 215 10.5.5 Clinical Quality Assurance 215 10.5.6 Tolerance Levels 216

10.5.7 Doses of Irradiation Outside the Treatment Volume 216 10.6 Clinical Studies with IMRT 217 10.6.1 Head and Neck 218

10.6.2 Prostate 220

10.7 Discussion and Summary 224 10.8 Additional Challenges for IMRT Implementation 224 References 225

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therefore, often identify dose distribution variation limits, including the maximum acceptable dose hot spot and the minimum acceptable dose. Another criterion includes the percentage of the target encom- passed covered by the prescription (isocenter) dose.

This prescription method is often so rigid that the selection of radiation beam locations and orientations is constrained to keep the beam isocenters in a common location.

The 3D-CRT treatment planner has a few degrees- of-freedom with respect to developing the treatment plan. They can typically select from two or more beam energies, affecting the overall penetration of the beam, the depth of the build-up region, and the width of the beam-edge penumbra. They can select from a continuous spectrum of gantry angles and treatment couch angles to preferentially align the radiation beam within the patient’s anatomy.

They can select the number of beams that are used and their weights (typically expressed as the dose to isocenter delivered by each beam), and can use off-the-shelf beam-fluence modification tools such as wedges or simple custom-designed surface-compensation filters (Ellis and Miller 1944; Ellis et al. 1959). Wedges are designed to tilt the radiation beam dose distribution relative to the central axis to compensate for abnormal incidence on the patient’s surface or for the dose distribution caused by intersecting radiation beams. In many tumor sites, e.g., in the treatment of prostate cancer, class solutions are developed by clinics to improve efficiency in the treatment planning process.

In 3D CRT, while the radiation targets are able to be covered, there is no mechanism for selectively avoiding critical structures. These 3D-CRT dose distributions have convex surfaces and cannot “wrap around” critical structures. Simply shielding the projection of a critical structure by a radiation block or multileaf collimator produces heterogeneities in the resulting dose distribution that are impossible to compensate for by adjusting the other radiation portals. The relatively strict limitations on target dose distribution homogeneity precludes the partial blocking of the radiation portals. Solutions to this problem include limiting the target dose and mixing radiation beam modalities, such as photon and high-energy electron treatments. For head and neck treatments, lateral photon beams can adequately irradiate the target tissues but would over-irradiate the spinal cord. The treatment is typically divided into two courses: firstly, the spinal cord is irradiated along with the target tissues until the cord tolerance is reached. Then, the photon beam is blocked

to avoid the cord and electron beams are patched to the photon beams. The electron beam energy is selected so that the cord lays beyond the practical range of the electron beam (Fig. 10.1). Even with this level of sophistication, these treatments irradiate the salivary glands causing significant morbidity and a degradation in the quality of life.

Fig. 10.1 Setup diagram shows a hot match of photon ﬁ eld and electron ﬁ eld on skin

Electron field

Photon field Hot match

on skin

Almost immediately after the introduction of 3D CRT, the radiation therapy physics community began to understand that the radiation dose distribution shape could be significantly altered to conform to target shapes while selectively avoiding critical structures. The underlying technology required was to remove the restriction of homogeneous incident fluences and allow the fluences to vary throughout the radiation portal, with selection of the fluence intensities typically made using sophisticated algorithms. This process is sometimes referred to as the

“inverse problem” or “inverse method” of treatment planning (Fig. 10.2). While the use of fluence modulation does not remove radiation from critical structures, it does limit the dose to the structures directly within the beam’s path. The resulting cold regions within the tumor are compensated by increasing the fluence from the other radiation beams. The addition of fluence modulation leads to the term intensity-modulated radiation therapy (IMRT) as the name for this modality.

The use of IMRT requires that the fluences be adjustable within the constraints of the dose-delivery technology. There are two general delivery tech- nologies that can deliver IMRT dose distributions, both of which have hundreds to thousands of individual fluence elements (bixels) that require adjust-

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ment. The bixel intensities require adjustments that are too complex for individual treatment planners to conduct. The modification of single bixel intensity is coupled with the bixel intensities from other beams that intersect its path. With this level of complexity, the optimization of bixel intensities can only be conducted by computers, and the field of inverse planning is developed.

The change in radiation therapy treatments has been profound. For the first time, radiation doses could be developed that wrapped around the target volumes, selectively, avoiding critical structures. The radiation oncologists, physicists, and dosimetrists could describe their intent based on dose and the geometry of the target and organs, rather than by the delivery geometry.

Of course, the radiation beam energy and incident angles have an effect on the resulting dose distribution, but the tone of the communication between these professionals changed from a technology- to a dose-based one. This revolution in treatment planning did not come without a price.

The compensation of f luence reduction across critical structures was not perfect. This led to larger dose heterogeneities within the target than was customarily seen with 3D CRT. While treatment planners were comfortable with hot spots of a few percent, they now had to consider allowing hot spots of 1015%, depending on the complexity of the dose distribution, the beam geometry, and the treatment-planning system. Cold spots were also often more severe and treatment planners often had to compromise their previous experi- ences. Tools to selectively modify and adjust these

dose heterogeneities are now available (ActiveRx, Corvus, Cranberry Township, Pa.), but their ultimate utility has not been proven.

Two methods for delivering the optimized fluences have been developed (see later). For the pur- poses of this discussion, their principal feature is that the radiation beam is delivered using a sequence of relatively small and complex portals. Before IMRT, there was significant scientific discussion regard- ing the accuracy of dose distribution calculation protocols for such fields. For example, the calibra- tion and quality assurance of small fields used in stereotactic radiation therapy required specialized equipment and techniques to make sure that the measurements did not suffer from volume averaging due to the rapid variations of dose as a function of position. For conventional therapy, approximations were made in modeling the sources of the radiation fluences exiting the linear accelerator. For example, while most of the X-ray radiation comes from the electron beam target, the flattening filter, as well as the primary and secondary collimators, also provide radiation fluence. For larger fields, the influence of these different radiation sources is straightforward to model. For smaller fields, the relative contribution of these sources is more complex and requires more sophisticated methods for modeling the dose delivered from the linear accelerator head. In IMRT, the radiation fields can extend from relatively large to extremely small, all within the same treatment.

This stretches dose modeling and verification measurement techniques to the limit.

The two delivery mechanisms differ by their method for temporally delivering the computer-

Fig. 10.2 Conventional and intensity- modulated radiation therapy (IMRT) dose distribution to a complex target volume (hatched) and critical structure geometry.

The upper four fi gures indicate the dose distribution using conventional planning and delivery. The lower four fi gures indi- cate fi xed-beam and arc-based IMRT with different optimization criteria

Parallel opposed beam therapy Arc therapy Four field box therapy Conformation therapy Non uniform beam radiotherapy

Three field technique

Minimal mean dose outside target volume

Specified maximum dose to organ at risk

Minimal dose to organ at risk

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optimized radiation fluences. The first, termed tomotherapy (Carol et al. 1996; Grant and Cain 1998; Low 1998), uses a geometry similar to that of a computed tomography scanner. A narrow fan beam is generated using a linear accelerator mounted to the gantry such that its central axis is aligned with radii extending from the gantry rotation axis. The fan beam is oriented such that the wide extent is parallel to the plane of rotation. The fluence is modulated along the wide direction using a collimator, which is pneumatically operated. The collimation leaves move into and out of the beam in the direction parallel to the gantry axis of rotation. The delivery system moves the leaves into and out of the beam to provide the fluence modulation as the gantry is rotated. It is important to note that for any single treatment, the gantry moves at a constant rotation rate with the linear accelerator operating at a constant fluence rate. This causes a sequence that provides a constant fluence rate per angle bin. The modulation is conducted by subdividing the gantry angles into small bins (e.g., 5° bins) and leaving a leaf open a fraction of that angle bin corresponding to the requested relative fluence from that gantry angle. The first commercial application of this technique was produced by the Nomos Corporation and was called the Pea- cock system. It had the unique characteristic of being an add-on system to conventional linear accelerators.

The system included the collimator and a treatment- planning system. By rotating the collimator around the patient using the linear accelerator gantry, the system delivered an intensity-modulated dose distribution to two nearly 1-cm-thick regions (often termed “slices”). When the tumor extended beyond the limit of the collimator, it was treated using multiple passes, moving the couch precisely to abut the dose distributions within the patient.

While the Peacock system was an add-on system to conventional linear accelerators, Tomotherapy, Inc., has produced a stand-alone tomotherapy linear accelerator (Mackie et al. 1999, 2003). The accelerator (Fig. 10.3) looks like a conventional CT scanner.

The radiation beam is produced using a 6-mV linear accelerator mounted to a CT gantry. The accelerator is aimed at the rotation axis of the gantry. A pneu- matic MLC is used to modulate the radiation fluence in a method similar to the Peacock system. Unlike the Peacock system, the Tomotherapy system moves the patient couch at the same time the gantry is rotated, providing a delivery geometry that is helical in nature, rather than the serial delivery provided by the Peacock system. An added feature of the Tomotherapy system is an imaging system that is placed opposite the radiation source. This allows the Tomotherapy system to be used as a megavoltage CT system for on-board patient imaging. Because the radiation source is high energy (they detune the linear accelerator to approximately 4 mV), the radiation interaction is primarily through Compton scat- tering, so high-Z materials (contrast agents, dental fillings, and metal implants) do not cause image artifacts.

The second delivery process took advantage of the hardware on existing linear accelerators. For the “2D” and 3D conformal therapy processes, the method of conforming radiation to the tumor had been to fabricate custom-designed high-density metal collimators called blocks. Linear accelerator manufacturers marketed multileaf collimators (MLCs) to replace these custom-fabricated blocks.

The use of MLCs provided a more efficient treatment process because the therapist did not have to enter the room between each portal to change the block. The MLC leaves (typically projecting to 1 cm

Fig. 10.3 The helical tomotherapy unit installed at the University of Wisconsin. The left panel is with a cover open. The right panel is looking along the CT table into the ring gantry bore.

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at the linear accelerator isocenter plane) were computer controlled, in that they were each separately controlled and operated by the linear accelerator computer system. Implementation of IMRT required the delivery of the nonuniform optimized fluence distributions. With conventional linear accelerators, the most flexible modulation technique was to fix the gantry and collimator and deliver the fluence using a programmed set of portals such that the total fluence was equal to the desired fluence. The linear accelerators had not been designed to deliver dozens to hundreds of portals for each treatment, so during the early implementation, delivery of the fluence distribution took as long as 40 min. Current linear accelerators are capable of delivering these complex fluence distributions in 12 min per portal.

The main consequence of the conventional linear accelerator implementation of IMRT is that the radiation field geometry is similar to that in conventional conformal therapy in that relatively few beam directions are employed for each treatment.

The selection of those directions has changed, however. Treatment planners often decide on portal directions more to spread the fluence than to avoid specific critical structures. The assumption is that the fluence modulation will take care of the normal structure dose limits.

The introduction of IMRT has changed the radiation therapy treatment-planning process. For 2D, the physician localizes therapy using bony and cavity anatomy, implanted markers, and contrast agents. The portals are selected based on the ability to define such anatomical references. For example, anteroposterior and lateral beams are preferred because the anatomical references are straightforward to identify. Wedges and compensating filters are used to homogenize the radiation dose of intersecting beams. Confirmation of the portal positioning and shape is conducted using kilovolt- age X-ray simulator systems or using images of the treatment portals taken with the linear accelerator beam. With 3D, the beams are defined by the projections of targets in the 3D model of the tumor and normal anatomy provided, typically using CT. The negative dose distribution consequences of the interactions of these portals are still over- come with wedges and compensating filters. The connection of the portal orientation and the ability to verify the portal position began to be broken as clinics realized that the use of non-conventional beam angles provides significant improvements in some treatment sites. This led to the disconnection between portal shape verification and placement

verification. As long as a common isocenter was used, portal placement verification could be conducted using only the portals that provided useful anatomical references. Some clinics completely separated the portal and patient positioning verification, defining portals that were used only for positioning verification.

The dose prescription also changed with the introduction of the 3D treatment process. With a quantitative model of the tumor shape and size, and an accurate model of the linear accelerator, the clini- cian could be provided with a comprehensive review of the planned dose distribution and could determine if the selected beam orientations, energies, portal shapes, and relative intensities provided an adequate delivered dose distribution. The process of prescribing dose became more sophisticated, with maximum and minimum dose constraints along with the nominal prescription dose.

These trends continued with the introduction of IMRT. No longer could the portal dosimetry even be qualitatively determined by examining the portal orientation. Portals often overlapped the critical structures they were intended to protect. The physician’s prescription was now an input to the planning process, rather than simply a manual feedback for the plan quality. The use of IMRT reduced the influence that portal orientations and energies had on the resulting treatment plan. The prescription criteria, however, had a profound influence on the treatment-plan quality, and each clinic had to develop appropriate dose-prescription criteria to enter into their treatment-planning system. In fact, the prescription criteria were often adjusted not to match the ultimate desired dose, but in order to force the planning system’s optimization software to develop an adequate dose distribution.

10.2

Basic Physical Principles of IMRT

IMRT generates dose distributions that conform to complex target volume geometries in all three dimensions. The dose distributions can be considered to be produced by a superposition of individual beamlets, each consisting of a narrow incident photon beam (Fig. 10.4). Ideally, all beam entry angles and positions would be available for use, but limitations of the dose delivery devices constrain the incident radiation beam-fluence distributions to two types.

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For IMRT delivery, the fluence is modulated as a function of entry angle for each point within the target volume. The dose at any point within the patient is generated by a series of beamlets incident on that point, each with a unique entry angle. Cur- rently, treatment accelerators are capable of simultaneously delivering fluence across a broad surface area of the patient in the form of a cone beam, with a unique patient entry angle for each position within the cone beam. The radiation fluence can be modulated within the cone beam using a physical modulator, a scanning dynamic MLC, a scanning

“bremsstrahlung” photon beam, or a combination of these techniques (Convery and Rosenbloom 1992; Mackie et al. 1993; Chui et al. 1994; Grant and Bellezza 1994; Spirou and Chui 1994; Stein et al. 1994; Svensson et al. 1994; Carol 1995; Yu 1995; Yu et al. 1995; Geis et al. 1996; van Santvoort and Heijmen 1996; Hounsell and Wilkinson 1997; Richter et al. 1997; Stein et al. 1997; Webb et al. 1997); for each of these, significant time is required to deliver the incident fluence distribution. To limit the treatment to practical delivery times, the number of incident cone beams is limited at most to a dozen. The beamlet distribution at each point within the patient when cone-beam IMRT is used is therefore limited principally in number, rather than in direction. Alternatively, the accelerator can be operated using dynamic motion of one or more of the angular degrees of freedom (couch, collimator, and gantry); of these, only dynamic gantry motion is currently being investigated (Mackie et al. 1993; Yu 1995; Carol et al.

1996). When dynamic motion of the gantry is used, there is insufficient time to arbitrarily adjust the fluence at each gantry angle using standard MLC;

however, the fluence can be arbitrarily modulated along a one-dimensional array of beamlets using a rapidly actuating MLC (Woo et al. 1994; Carol et al. 1996; Webb and Oldham 1996). The collimator modulates the fluence by alternately opening and closing each leaf as the gantry rotates. This technique is termed fan-beam IMRT delivery. The fluence is modulated in time by opening and closing each of the multileaves over the course of a narrow gantry angle range. In this manner, modulated beams of 50 or more gantry angles are delivered to a roughly cylindrical volume with the cylinder thickness corresponding to the leaf opening size.

A commercial MLC is designed using two adjacent independently actuated leaf banks, so two adjacent cylindrical volumes are simultaneously irradiated (Carol 1995; Wu et al. 1996; Low 1997). After the dose is delivered to the first volumes, the patient is precisely moved and the next two cylindrical volumes treated. The couch angle can also be adjusted to allow a broad range of entry angles. The beamlet distribution for each point within the patient is limited principally by incident angular distribution (the beamlets are coplanar) rather than by number.

Standard MLCs can also be employed when using dynamic gantry motion (Yu 1995). The MLC is moved during arc rotation such that an open field of predetermined shape is delivered for each gantry angle. To modulate the beam intensity for each beamlet, multiple arcs are delivered, each with a different portal shape such that the modulated final fluence is delivered. The beamlet distribution is similar to that of the fan-beam geometry.

In traditional radiation therapy, the photon beam energy selection is used to modify shallow doses and to limit dose heterogeneities along the beam path. In megavoltage therapy, shallow and deep tumors are treated with low and high beam energies, respectively; however, when beam-intensity modulation is used, much of the dose hetero- geneity due to the depth dose can be removed by modulating nearly orthogonal beams (similar to the concept of a wedged pair); therefore, the need for higher energy beams (>10 mV), with their associated neutron contamination, may be limited. In addition, the conformation of dose within the patient relies on each beamlet independently depositing dose within the patient, which requires that the secondary electron range be small with respect to the beamlet size. Studies have so far concluded that the ideal beam energy for intensity modulation is between 4 and 10 mV.

Fig. 10.4 Sketch of beamlet-generated dose distribution

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10.3

Inverse Treatment Optimization (Cost Func- tion and Search Algorithm)

As indicated previously, IMRT requires a computer optimization method of determining the nonuniform beam-fluence profiles. Early attempts at automated plan optimization methods were conducted by several groups, including Hope and Orr (1965), Redpath et al. (1976), and McDonald and Rubin (1977), but with limited acceptance. With the development of 3D CRT and the corresponding increase in image and graphic data, interest in computer optimization and automation has been renewed. Inves- tigation of computer optimization has intensified because of the requirement for determining optimal nonuniform beam fluences for IMRT (Brahme 1988; Lind and Kallman 1990; Gustafsson et al.

1995).

In 1982, Brahme et al. developed the concept of determining the modulated radiation field fluence distribution necessary to produce a desired dose distribution. An analytic solution of the beam-fluence profile required to treat a circularly symmet- ric target with a critical structure at its center was developed. In 1987, Cormack (1987; Cormack and Cormack 1987) extended this approach by extending it to targets with an axis of symmetry. In 1990, Barth further developed this approach by describ- ing a mathematical solution for a target of arbitrary shape by circular component target decomposi- tion. For general mathematical solutions, the optimized incident fluence profiles contained negative fluence beamlets, an unphysically realizable result (Cormack and Quinto 1989). Application of the non-negativity constraint has been included in recent efforts to develop numerical solutions to calculate the incident fluence profiles. Reported methods include: (a) exhaustive search (Barth 1990); (b) gradient techniques (Xing and Chen 1996; Spirou and Chui 1998; Bortfeld 1999; Seco et al. 2002; Zhang et al. 2004); (c) image reconstruction approaches (Brahme 1988; Brahme et al. 1988; Bortfeld et al. 1990; Holmes et al. 1991); (d) quadratic programming (Djordjevich et al. 1990);

(e) simulated annealing (Webb 1989; Mohan et al.

1992; Djajaputra et al. 2003); (f) superposition algorithm (Scholz et al. 2003); (g) matrix inver- sion (Goldman et al. 2005); (h) genetic algorithm (Cotrutz and Xing 2003; Hou et al. 2003); and (i) adaptive control algorithm (Lof et al. 1998).

The task of generating an optimal plan can be separated into two parts: (a) specification of an opti-

mization criterion, and (b) the optimization algorithm used. The optimization criterion is expressed as a mathematical entity in the form of an objective or cost function. The objective function defines a plan’s quality and is to be maximized or minimized, as appropriate, to satisfy a set of mathematical constraints. The objective function yields a single numerical value (sometimes referred to as a score) for the plan that is used for evaluating a set of com- peting treatment plans. Often an iterative computer search is made of the fluence distributions, with guidance by the objective function. Several types of objective functions have been investigated and have been shown to be useful in some, but not all, clinical situations (Starkschall 1984; Schultheiss and Orton 1985; Langer and Leong 1987; Langer et al. 1990; Mohan et al. 1992; Niemierko 1992;

Niemierko et al. 1992; Rosen et al. 1992; Langer et al. 1993; Sodertrom and Brahme 1993; Jain et al. 1994; Djajaputra et al. 2003; Lian et al. 2003;

Yan et al. 2003; Beaulieua et al. 2004; Yang and Xing 2004; Arrans et al. 2005; Lee et al. 2005). His- torically, most objective functions use dose-based or dose-volume-based criteria, but recent efforts are examining the use of biologically based indices (e.g., tumor control probability and normal tissue com- plication probability; Niemierko 1992, 1997, 1998;

Brahme 1995; Graham et al. 1996; Deasy et al.

2002; Wu et al. 2002; Stavrev et al. 2003).

Once the optimization criteria have been selected, an algorithm is used to automatically determine a set of plan parameters that optimizes the chosen objective function. There are a number of IMRT treatment-planning optimization algorithms under investigation. If the objective function and constraints can be described by linear or quadratic functions of the plan parameters, efficient mathematical optimization techniques can be used that are guaranteed to find the best score (Langer and Leong 1987; Djordjevich et al. 1990; Langer et al.

1990; Rosen et al. 1992). If the objective function is not linear or quadratic as a function of the optimization parameters, non-linear optimization techniques must be used (Cooper 1978; Xiao et al. 2004;

Yang and Xing 2004; Olafsson et al. 2005). Other optimization approaches being used for IMRT are adopted from computed tomography (CT) image reconstruction techniques, with filtered back projection being the principal technique (Holmes et al.

1991; Gustafsson et al. 1994; Holmes and Mackie 1994). If these techniques are used in an unmodified form, the optimized fluence distribution will have negative fluences. Non-negativity constraints

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impact on the quality of the resulting distribution, so iterative techniques are often also required. One iteration method is a stochastic technique termed

“simulated annealing,” originally developed from statistical mechanics for finding the global minima of non-linear objective functions (Metropolis et al. 1953; Kirkpatric 1985). Webb (1989) applied this approach to radiation therapy plan optimization. Since then, several other researchers, as well as a commercial system, have used this method (Morrill et al. 1990, 1991; Webb 1991a,b, 1992;

Mohan et al. 1992; Mageras and Mohan 1993;

Mohan et al. 1996; Sauer et al. 1997). The gradient techniques (Xing and Chen 1996; Spirou and Chui 1998; Bortfeld 1999; Seco et al. 2002; Zhang et al. 2004), which require less computation time than simulated annealing, are also used in several commercial systems.

Optimization criteria and mathematical search algorithms are being investigated vigorously. As the capabilities of computers continue to improve, the time required to conduct an optimization search may become shorter. The biggest challenge may be the development of optimization criteria that will be appropriate for a wide variety of cancer sites and histologies and be able to account for individual physician preferences.

10.4

IMRT Treatment Delivery Systems

10.4.1

Historical Review

Early IMRT concepts were pioneered several decades ago. Seminal contributions were made by Takahashi and colleagues (1965). Their work illustrated some of the important concepts in both 3D CRT and IMRT treatment delivery. Dynamic treatments were planned and delivered by Takahashi et al.’s group using what may have been the first MLC system. The MLC used a mechanical control system to conform the beam aperture to the projected target shape as the machine was rotated around the patient.

In the late 1950s, another pioneering effort in IMRT was conducted by the group at the Massa- chusetts Institute of Technology Lahy Clinic, who independently developed an asynchronous portal- defining device similar to that of Takahashi et al.

(Wright et al. 1959; Proimos 1960, 1966; Trump et al. 1961). The Royal Northern Hospital in England

also pioneered a 3D CRT effort (Green 1965). The group developed a series of cobalt-60 teletherapy machines in which the patient was automatically positioned during rotational therapy by moving the treatment couch and gantry during the radiation delivery using electromechanical systems. This was called the “Tracking Cobalt Project” because the planning and delivery system attempted to track around the path of disease spread and subsequently to conform the dose distribution. The work was extended in the 1970s and 1980s by Davy and Brace at the Royal Free Hospital in London (Davy et al. 1975;

Davy and Brace 1979; Brace et al. 1981a,b; Brace 1982, 1985; Davy 1985), The 1970s saw important advances in IMRT from the Harvard Medical School by Bjarngard, Kijewski, and others (Bjarngard and Kijewski 1976; Bjarngard et al. 1977; Kijewski et al. 1978; Chin et al. 1981, 1983). Unfortunately, computer technology had not yet advanced to the capac- ity required for practical implementation of traditional 3D CRT and IMRT. This work and that of the previous researchers led the way for modern IMRT.

The recent introduction of commercially available MLCs (Boyer et al. 1992; Webb 1993; Klein 1994; Mohan et al. 1996) and the simultaneous development of medical accelerator computer control systems (Brahme 1987; Fraass 1994; Boyer 1995; Fraass et al. 1995; Mohan et al. 1996) assisted the rapid and widespread development of practical IMRT. Until recently, medical accelerators were produced with relatively primitive beam collimation systems consisting of opposed jaw pairs that were constrained to move symmetrically to provide square or rectangular beam apertures. A low melting point lead alloy was used to fabricate a customized beam block for each radiation portal to provide static conformal field shaping (Powers et al. 1973). This system was relatively inefficient, in that the radiation therapist was required to enter the treatment room and manually change the aperture block for each treatment portal. A limited improvement came when the collimator jaws were allowed to move independently, allowing the fabrication of lighter blocks.

A rudimentary form of dynamic intensity modulation, the dynamic wedge, was developed using independent jaws (Kijewski et al. 1978; Leavitt et al.

1990, 2000; Klein et al. 1995, 1998; Leavitt et al.

1997; Edlund et al. 1999; Kubo and Wang 2000).

The MLC replaced (or supplemented, depending on the accelerator manufacturer) the rectangular jaw system with a set of independently adjustable, relatively narrow, tungsten leaves (typically projecting to 1 or 1.25 cm wide at isocenter). The leaves were

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placed under computer control and were therefore able to create custom-shaped beam apertures and allow the change of portal shapes without the requirement of room entry by the radiation therapist. The computer control system was often used to automatically rotate the gantry (and collimator), position the treatment table, and control dose rate.

The IMRT delivery is made practical by using a MLC in a dynamic mode similar to dynamic wedge (i.e., moving each pair of MLC leaves with the beam on while at a fixed gantry angle or during a gantry rotation). This method, and three others, for delivering IMRT are discussed in the following sections.

10.4.2

Arc-Based Fan-Beam Dynamic MLC IMRT

A commercial implementation of a fan-beam approach to IMRT (tomotherapy) uses a mini-MLC system that is mounted to an unmodified linear accelerator, and treatment is delivered to a narrow slice of the patient during arc rotation (Carol 1992, 1995a,b; Carol et al. 1996; Low 1997; Verellen et al. 1997; Low 1998a- c; Al-Ghazi et al. 2001; Saw et al. 2001a,b). The beam is collimated to a narrow slit and beamlets are turned on and off by driving the mini-MLC leaves out and in the beam path, respectively, as the gantry rotates around the patient. A complete treatment is accom- plished by sequential delivery to adjoining slices.

As previously indicated, this type of IMRT system (Peacock, Nomos Corp., Pittsburgh, Pa.) was the first broadly implemented IMRT planning and delivery system (Woo et al. 1994).

The helical version of tomotherapy was developed by Mackie et al. and is commercially available from Tomotherapy, Inc. (Mackie et al. 1993a,b, 1997, 1999; Welsh et al. 2002; Jeraj et al. 2004). The dose is delivered using a narrow MLC and small linear accelerator mounted in a modified CT scanner gantry. The radiation is delivered while the patient is moved through the gantry in the same way that a helical CT study is conducted (Fang et al. 1997;

Mackie et al. 1997).

10.4.3

Fixed Portal Cone-Beam Dynamic MLC

Other IMRT researchers have concentrated on using the MLC to provide a full-field or cone-beam modulation technique (Kallman et al. 1988; Convery and Rosenbloom 1992; Galvin et al. 1993; Bortfeld et

al. 1994; Svensson et al. 1994; van Santvoort and Heijmen 1996; Webb et al. 1997). For IMRT, the MLC is operated in a dynamic mode where the gap formed by each opposing leaf sweeps under computer control across the target volume to produce the desired fluence profile. The technique for setting the gap opening and its speed for each MLC leaf pair was first determined by a technique introduced by Convery and Rosenbloom (1992) and extended by Bortfeld et al. (1994) and Spirou and Chui (1994).

The Memorial Sloan Kettering Cancer Center was one of the first centers to adopt a variation of this technique (Ling et al. 1996). Because of its implementation on conventional linear accelerators, this technique currently treats the greatest number of IMRT patients worldwide (Burman et al. 1997; de Neve et al. 1999; Ma et al. 1999; Mott et al. 1999;

Wu et al. 2000; Budgell et al. 2001; Chui et al. 2001;

Low et al. 2001).

10.4.4

Fixed-Portal X-ray Compensating Filter IMRT

The original and common use of compensating filters is to account for missing tissue deficits while maintaining the skin-sparing benefit of megavoltage photon beams. Filters designed using 3D RTTP are capable of compensating for not only the missing tissue deficit, but also for internal tissue heterogeneities. Filters can be designed by calculat- ing a thickness along a ray line using an effective attenuation coefficient for the filter material and dose-ratio parameters for effective depths may consider scattered radiation from the filter. The filter construction process can be automated using numerically controlled milling machines, and can be extended for use when inverse treatment planning is employed, fabricating the compensation filter that generates the desired IMRT fluence profile when the filter is placed in the radiation beam.

A carousel or stack device could be supplied to hold multiple filters that would be automatically moved into the beam under computer control. Being relatively simple, this method of delivering IMRT may prove to be the easiest to clinically implement. Com- parisons of IMRT dose distributions delivered using physical modulators and MLC-based approaches are under investigation, with indications that the dose distributions provided by these modalities are similar (Stein et al. 1997). Stein et al. (1997) have developed a method for producing physical modulators by stacking layered filters up to 1.6 cm thick.

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The filters are fabricated using low-melting-point alloy poured into foam molds, which are cut using a computer-controlled cutter. Other authors have also investigated the use of physical filters for IMRT (Low et al. 1996a,b).

10.5

IMRT Quality Assurance (Commissioning and Clinical)

As with all treatment modalities, verification and characterization measurements are required before clinically implementing of IMRT. Dosimetric verification traditionally includes a comparison of measured and calculated dose distributions for selected test treatment plans. Acquisition of benchmark quality dose distributions for static-beam therapy relies on ionization chambers, which measure doses at individual points. Acquisition of isodose distributions relies on the assumption that the dose distribution being delivered during the measurements is constant in time (at least to within an overall constant dose rate, monitored by a reference ionization chamber), and that the chamber can be scanned throughout the dose distribution during the measurement. Dynamic IMRT dose distributions are delivered using a tempo- ral sequence of incident fluences, so this assumption does not hold. As a consequence, the measurement of dose at a single point using an ionization chamber requires that the entire fluence delivery sequence be delivered, requiring measurement times as long as 20 min. Similar difficulties are seen with all electronic point dosimeters (e.g., diodes). Stand-alone integrating point dosimeters, such as thermolumi- nescent dosimeters (TLDs), allow the simultaneous measurement of dose at numerous locations, but with the additional workload required for off-line readout. Practical application of real-time point dosimeters provide, at most, dozens of measurement points within the treatment volume.

Characteristics of IMRT dose distributions include complex 3D treatment volume geometries and steep spatial dose gradients. The precise characterization of these distributions using only point dosimeters is not practical. To obtain the large quantity of dose measurements necessary for the description of the dose distributions, at least planar dosimeters are necessary. Currently, the only practical planar dosimeter is film, either radiographic or radiochromic. Benchmark-quality dose distributions for megavoltage beams cannot be acquired using radio-

graphic film due to its energy response and relative dosimetric characteristics (Williamson et al. 1981;

Hale et al. 1994; van Battum and Heijmen 1995;

Mayer et al. 1997; Dogan et al. 2002; Zhu et al.

2002). Radiochromic film uses a tissue-equivalent dye that changes color when irradiated (Stevens et al. 1996; Zhu et al. 1997, 2002; Niroomand-Rad et al. 1998; Dogan et al. 2002; Esthappan et al. 2002).

Early implementations of radiochromic film had significant difficulties with response homogeneity and poor dose sensitivity. Current radiochromic film has homogeneous response (better than 5%) and sufficient sensitivity to use for single-fraction IMRT quality assurance. The cost of the film was also significantly reduced such that the cost of the film may be less than the cost of radiographic film when the processor purchase and maintenance is considered.

The new radiochromic film is sufficiently novel that its full impact on IMRT quality assurance has not been determined.

Even considering the dosimetric limitations of radiographic film, it is at least capable of accurately localizing regions of steep spatial dose gradients. For example, in a region of a 5% mm-1 dose gradient, an error of 5% in the relative dose measurement results in a spatial error in the localization of that gradient region of only 1 mm. Radiographic film therefore provides an opportunity for accurately localizing the edges of a complex dose distribution within the 2D measured area.

To determine the spatial accuracy of the treatment planning and delivery system, the spatial location of the measured and calculated doses must be accurately and independently determined. Most treatment-planning systems use a Cartesian coordinate system to localize the calculated dose distribution. For example, many planning systems tie the coordinate system directly to treatment delivery. Before acquisition of the volumetric CT scan of noninvasively immobilized patients, radiopaque markers are placed on the immobilization system (e.g., thermoplastic masks), corresponding to the lateral and anterior (for supine patients) projection of a selected point, typically the treatment isocenter.

During treatment planning, the isocenter location is identified on the CT-scan data set. Verification measurements need to be conducted having a quantitative alignment with the treatment plan isocenter.

The treatment-planning system may provide instructions to offset the patient position in a specified direction and distance to place the accelerator isocenter at the optimal location for treatment. Veri- fication measurements should be made in the same

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way. Localization wires are placed on the dosimetry phantom identifying the projected origin. The wires are visualized on the CT-scan data set to determine the isocenter location within the phantom. The dosimeter position locations are determined relative to the origin using physical measurements or machine drawings of the phantom.

10.5.1

Dosimetry System

There is no one dosimetry system that conveniently measures all of the dose information necessary for quality assurance. Systems that balance thorough dosimetric measurements with labor and complexity have been developed in many academic centers (Klein et al. 1998; Chao 2002; Olch 2002; Higgins et al. 2003; Agazaryan et al. 2004; Letourneau et al. 2004; Moran et al. 2005; Wiezorek et al. 2005;

Winkler et al. 2005; Yan et al. 2005); these exploit the optimal characteristics of point dosimeters (e.g., ionization chambers) and planar dosimeters. Radio- graphic film is used to measure the spatial position of the dose distribution (by aligning the steep-dose gradient regions) and ionization chambers are used to measure the absolute dose in relatively shallow dose-gradient regions.

Careful dosimetry phantom design is important to commissioning and quality assurance procedures. Anthropomorphic phantoms offer the advantage that they are shaped and sized similarly to patients; however, if the treatment-planning system does not accurately consider internal heterogeneities, determination of the cause of discrepan- cies between measured and calculated doses may be difficult, especially within and near bony anatomy.

While TLDs provide convenient single-point measurements, the spacing between locations for TLD chips may be larger than desired. Film preparation is made more difficult with anthropomorphic phantoms due to irregular external contours. Radiogrphic film preparation may require careful cutting in the darkroom to conform the film shape to the phantom outline. Homogeneous cuboid dosimetry phantoms do not share these limitations.

Multipoint planar dosimeters, such as electronic portal imaging devices (EPIDs) and diode system (Sun Nuclear), have been developed that can greatly assist in the quality assurance process (McCurdy et al. 2001; van Esch et al. 2001, 2004; Chang et al. 2003, 2004; Letourneau et al. 2004; Baker et al. 2005;

Wiezorek et al. 2005). Rather than measure dose

within a phantom, the dosimeter system is placed at isocenter and irradiated using the delivered beams.

A prediction of the incident dose from each beam is made by the treatment planning system and com- pared against the measurement. Typically, the com- parisons are made port by port, but some planning systems are also capable of computing the summed dose, i.e., the dose delivered from all beams using a common gantry angle. While this does not reflect the delivered dose, gross delivery or planning errors would be detected by such a system. One of the greatest challenges of using these dosimeters is the selection of acceptance criteria. Because the dose delivered to these planar dosimeters does not reflect the dose to the patient, developing passfail criteria based on the total delivered dose is not possible. The user can only compare the measured planar dose with the corresponding calculated planning dose distribution, for the same beam and depth. For this purpose, two main analysis tools are used: overlay comparison of calculated and measured dose profiles; and 2D plot of measurement points that fail both measures of a two component user-selected criteria consisting of distance to agreement (DTA) and dose difference (percentage of dose difference; Low 1998; Low and Dempsey 2003). Commonly, the passing criteria are set at 3 mm for DTA and 3% for dose difference. The DTA criteria would then fail any measured dose point for which a corresponding point could not be found in the calculated distribution, with the same dose, within a 3-mm radius. A point measurement would fail the percentage of dose-difference criteria if the difference in dose between the measured point and the equivalent point in the calculated distribution was greater than 3%.

An ideal dosimeter would be capable of providing dose measurements in a 3D volume, with an energy response similar to that of water or muscle.

BANG (bis, acrylamide, nitrogen, and gelatin) and other polyacrylamide gels have been investigated for use in dynamic IMRT dose-distribution measurements (Maryanski et al. 1994, 1996a,b, 1997; Gore et al. 1996; Low et al. 2000; Islam et al. 2003). The dosimeters operate on the principal that monomers can be made that cross-link when irradiated by ionizing radiation. There is a subsequent increase in the solvent proton relaxation in the presence of the polymer. The increased proton relaxation rate (R₂=1/T₂) can be imaged using mag- netic resonance imaging (MRI). The gel is irradiated using IMRT delivery and subsequently imaged using a clinical MRI unit. T1- and T2-weighted scans are obtained for each gel and the volumetric

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distribution of R₂ is determined using these scans.

A monotonic relationship exists between R₂ and absorbed dose, but because the gel radiation sensitivity is batch dependent, some gels are irradiated to known doses and scanned to obtain a dose cali- bration curve. A measurement of the full 3D dose distribution is provided when this medium is used, with measurement voxels as small as 1×1×2 mm³. The requirement of MRI for readout has limited the use of this detector medium to a few dosimetry studies. In addition to its MRI properties, the gel is normally transparent and becomes opaque on irradiation. Optical absorption imaging has also been investigated as a method for extracting the dose (Maryanski et al. 1996a,b; Kelly et al. 1998;

Wolodzko et al. 1999), but gels have not enjoyed widespread clinical implementation.

10.5.2

Commissioning

The core of a treatment-planning system commissioning procedure is the comparison of calculated and measured dose distributions. Traditional commissioning procedures investigate beams incident on a water phantom for a variety of field sizes, blocking geometries, beam modulators (wedges), and incident angles. However, IMRT dose distributions are generated using a superposition of complex fluence distributions; therefore, commissioning of these systems relies on developing dose distributions and fluence-delivery sequences for a series of selected target and critical structure volume geometries.

The volumes are selected based on expected sizes and shapes of clinical target volumes. Neighboring critical structures can also be defined to provide complex target geometries. Treatment plans used for commissioning should include a variety of target sizes, shapes, and locations within the test phantom, as well as at least one test using phantoms of different sizes. The test target volume can be cylindrical to simplify preparation and subsequent description of the target geometry. Both diameter and length of the targets should span the range of sizes used for clinical cases. Experiments should also include a broad range of optimization parameters.

Each commissioning experiment will include both film and absolute dose measurements (ionization chamber and/or TLDs, and possibly the use of radiochromic film). The results should be tabulated to show the dosimetric and spatial accuracy of the system.

10.5.3

System Quality Assurance

The accurate delivery of IMRT dose distributions depends on thorough accelerator- and delivery- system QA programs. A description of all accelerator QA procedures is beyond the scope of this chapter.

A thorough review was published by the American Association of Physicists in Medicine (Kutcher et al. 1994). Some of the specialized procedures specific to IMRT are mentioned in this chapter.

The accurate localization of the accelerator isocenter relative to the patient alignment fiducial markers is important for noninvasively immobilized patients. The isocenter position within the patient is aligned to the accelerator using the positioning lasers. As in all external beam therapy, if the lasers are not correctly aligned, the localization of the dose distribution within the patient will suffer.

For IMRT treatment planning and delivery, dose distributions are generated using fluence distributions incident on the patient from a series of directions. The superposition of the fluence distributions generates the planned dose distribution, and the planning system assumes that the orientation and position of each beam angle is accurate. Dose-delivery errors can occur due to excessive gantry sag and gantry and collimator angle misalignments. Qual- ity assurance tests have been developed that check beam and isocenter alignment.

For indexed sequential arc therapy, accurate patient positioning between arcs is extremely important. Carol et al. (1996) determined that an incorrect placement of the patient between successive arc treatments will cause a 10% mm^–1 dose heterogene- ity in the abutment region. Consequently, the accuracy of the patient immobilization and placement system is critical to accurate dose delivery. Periodic testing of patient indexing system accuracy should be conducted. One method is to place a film at the plane of isocenter and irradiate it using the open collimated portal.

The film is then moved using the same appara- tus as used in patient treatments, and the process is repeated six to ten times. Ideally, the film will have a homogeneous rectangular irradiated region, with no evidence of high- or low-dose regions. Overlaps or underlaps can serve to identify problems with either the positioning system or the treatment couch support structure or bearings. Tests are also being developed for the QA of dynamic MLC delivery (Chui et al. 1996) and helical tomotherapy (Yan et al. 2005).

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The implications of incorrect treatment setup are being investigated. Studies (Convery and Rosenbloom 1995; Low 1997; Ahmad et al. 2005;

Hong et al. 2005; Siebers et al. 2005) have shown that variations in the delivered-dose distribution can arise when the gantry, collimator, or couch angle are incorrectly set.

10.5.4

Monitor Unit Verification

A manual or simple computer calculation has been found useful to double check the monitor unit of a treatment in 3D CRT. In IMRT the treatment-planning system provides the MUs for each field. The MUs are verified either by direct measurement (Low 1998; Xing et al. 1999; Li et al. 2001) or by an independent calculation system (Boyer et al. 1999; Chen et al. 2000; Kung et al. 2000; Ma et al. 2000, 2003;

Xing et al. 2000; Pawlicki and Ma 2001; Chen et al.

2002; Yang et al. 2005). Independent computation is more efficient and less manpower intensive than direct measurement. The most accurate computation methods proposed for IMRT verification are based on Monte Carlo simulation (Ma et al. 2000, 2003;

Pawlicki and Ma 2001; Yang et al. 2005). Other computation methods include convolution of pencil beams (Boyer et al. 1999), self-consistent monitor unit and isocenter point-dose calculation (Chen et al. 2002), scatter-summation algorithm (Xing et al.

2000), ray tracing (Chen et al. 2000), and modified Clarkson integration (Kung et al. 2000).

10.5.5

Clinical Quality Assurance

The delivery of IMRT is similar in most respects to the delivery of standard conformal radiation therapy.

The principal difference lies in the requirement to keep the patient stationary during the entire course of treatment. For traditional conformal therapy, the motion of the patient during treatment will result in minor dose variations within the portal outlines and a major dose variation near the portal bound- aries; however, margins are typically applied to the tumor volumes to account for an expected amount of motion so that the targets will generally remain within the portal boundary. The exception to this rule is when compensating filters or wedges are used.

In these cases, patient motion during treatment will alter the doses within the beams an amount propor-

tional to the lateral fluence gradient. For example, the dose gradient for a 45q wedge is typically 4

5% cm-1, so movement within the portal boundary of 5 mm will result in a dose-delivery error of only 23%, respectively.

In IMRT, the dose distribution is due to a delivery sequence of incident fluences, each with a poten- tially large lateral fluence gradient. Yu et al. (1997) showed that the fluence delivery error can be as high as a factor of two when dynamic-MLC delivery is coupled with breathing motion; however, Bortfeld et al. (2002, 2004) have shown that, under certain assumptions, the dose-delivery errors in realistic

Fig. 10.5 a A digitally reconstructed radiograph of lateral pro- jection depicts a representative arc through nasopharynx and sphenoid sinus. b A double-exposure portal ﬁ lm taken at the same geometry from the linear accelerator corresponds to the digitally reconstructed radiograph.

aa

b b

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treatments is significantly less and averages out over multiple fractions such that the overall effect of tumor motion is to blur the resulting dose distribution. Clearly these two results appear contradictory, and the truth in actual clinical cases probably lies between the two. More research is necessary using more sophisticated simulations before the actual effects are understood using four-dimensional techniques. Until then, clinicians are cautioned to avoid using IMRT for tumors that move due to breathing, and to provide good immobilization techniques for all patients (Yu et al. 1997; Bortfeld et al. 2002, 2004; Engelsman et al. 2005).

One consequence of IMRT is the lack of a convenient open portal for image-based patient position and treatment verification. In IMRT, the two concepts must be separated: the patient position is verified separately from the treatment delivery. Patient positioning verification can be conducted using orthogonal portal films using fixed open fields and comparing against digitally reconstructed radio- graphs (DRRs) or simulator films. Figure 10.5 shows examples of the DRRs and portal films for a head and neck patient (Chao et al. 2000). Determination of patient positioning accuracy is made by manually marking bony landmarks and overlaying the portal films and DRRs on a light box. Treatment verification can be conducted using measurement phantoms. The use of in vivo dosimetry may be limited to intracavitary dosimetry. Engler et al. (1997) used in vivo TLD dosimetry placed beneath bolus on the patient’s skin to verify the delivered dose to arc- based fan-beam treatments; however, the acceptable dose tolerances were by necessity large due to the high spatial dose gradient near the patient surface.

10.5.6

Tolerance Levels

The consensus in the radiation therapy community is that the dose delivered to the tumor volume should be within 5% of the prescribed dose (Dische et al.

1993). The more complex IMRT systems have dif- ficulty in meeting the 2% in relative dose accuracy in shallow-dose gradients or 2-mm spatial accuracy in regions with steep-dose gradients recommended by the International Commission of Radiation Units and Measurements (van Dyk et al. 1993). Palta et al. (2003) proposed a set of more appropriate criteria for IMRT plan validation. The over-uncertainty in delivered dose should be less than 5%. The proposed confidence limits for relative dose difference

in treatment planning are 3% for regions with high dose and shallow-dose gradient, 10% or 2 mm DTA (distance to agreement) for regions with high dose and steep-dose gradient, 4% for regions with low dose and shallow-dose gradient, and 2 mm DTA for dose fall-off regions. The proposed tolerance limits for delivery systems with MLC are: 1 for SMLC and 0.5 mm for DMLC for leaf position accuracy; 0.2 mm for leaf position reproducibility and for gap width reproducibility; 0.75 mm for isocenter; 2% for SMLC and 3% for DMLC for low MU (<2 MU); and 2% for low MU symmetry.

10.5.7

Doses of Irradiation Outside the Treatment Volume

Followill et al. (1997), based on the observation that with IMRT some form of X-ray attenuation to modulate intensity is required, suggested that the number of MUs used to deliver a given treatment is increased over that used for conventional radiation therapy, resulting in the likelihood of increased whole-body dose to the patient by leakage and scat- tering of X-rays and, for higher energies, leakage of neutrons through the collimation and treatment head assemblies of the linear accelerator. These authors made estimates of the leakage neutron dose equivalent or photon dose at isocenter in the patient plane at a point 5 cm away from the central axis of a pelvic field, and estimated the whole-body X- ray dose equivalent contribution for 50 cm from the center of a 20×20-cm treatment field based on previously measured data (Stovall et al. 1995). They also calculated the MUs required to deliver a treatment with conventional X-rays (wedged or unwedged) and with beam-intensity modulated irradiation from two different systems (Varian MLC or Nomos Pea- cock MLC). Average total MUs derived from treatment plans for six patients are shown in Table 10.1.

The unwedged 6-mV beam required 20% more MUs than the 18-mV beam. With wedges, the 6-mV-beam MUs increased by 60% over that for 18 mV. With the Nomos Peacock system the total number of MUs per fraction increased with the number of arcs used. The total whole-body dose equivalent to a point 50 cm from the center of the pelvic field, when 70 Gy was delivered, was significantly higher with the intensity-modulated plans (Table 10.2). The authors also calculated the estimated risk of any fatal secondary cancer associated with scattered dose from the prescribed treatment as a percentage increase in likeli-

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hood in non-irradiated populations. The smallest increased risk was 0.4% for the 6 mV unwedged conventional technique, and the greatest risk was 24.4% with the tomotherapy technique using 25 MV X-rays. The authors caution against the use of 25 MV X-rays with intensity-modulated techniques, which may carry a seriously high risk of secondary can- cers, and recommend that the use of X-ray energies greater than 10 MV for beam-intensity modulated conformal therapy should not be attempted until the X-ray and neutron leakage dose and risks associated with the treatment have been carefully evaluated in long-term studies. At our institution Mutic and Low (1997) also measured whole-body irradiation doses from arc-based IMRT to the head and neck region using a water-equivalent plastic block whole- body phantom and a polystyrene phantom, placing TLDs in multiple locations arranged into orthogonal linear arrays. To assess the leakage dose component, the MLC leaves remained closed during treatment delivery. The total midplane whole-body dose from internal scatter and leakage was approximately 2.5%

of the total target dose, decreasing to 0.5% at 30 cm

from the target. The whole-body dose was entirely due to head irradiation leakage. The internal scatter dose was significant near the target but became negligible beyond 15 cm from the target in relation to the leakage dose; therefore, the total-body dose was proportional to the total MUs used to deliver a given treatment.

10.6

Clinical Studies with IMRT

Recently, Guerrero Urbano and Nutting (2004) presented two reviews on clinical use of IMRT in tumors of the head and neck region, central ner- vous system, lung, and in prostate, gynecological, breast, and gastrointestinal malignancies, as well as in other issues related to the clinical use of this new technique. The main indications at every site have been dose escalation and acceptable or reduced normal tissue toxicity. In general, a better dose distribution by IMRT has been well documented,

Table 10.1 The MU/cGy and total MU to deliver conventional and modulated beam-intensity radiotherapy. MLC multileaf collimator

Beam energy (mV)

Conventional Beam-intensity modulated

Unwedged Wedged Varian MLC

modulated

Nomos MLC tomotherapy

(MU/cGy) (MU/cGy) (MU/cGy) (MU/cGy)

MU per dose^a 6 1.2 2.4 3.4 9.7

18 1.0 1.5 2.8 8.1

25 1.0 1.5 2.8 8.1

(MU) (MU) (MU) (MU)

Total MU^b 6 8,400 16,800 23,800 67,900

18 7,000 10,500 19,600 56,700

2.5 7,000 10,500 19,600 56,700

aThe number of MU needed for the indicated technique to deliver 1 cGy to isocenter

bThe total MU needed, using the indicated technique and energy, to deliver 70 Gy to isocenter

Table 10.2 The estimated total whole-body dose equivalent (mSv) from a total delivered dose of 70 Gy at isocenter

6 mV 18 mV 25 mV

No wedges Wedges No wedges Wedges No wedges Wedges

Conventional 67 134 326 488 602 903

MLC modulated 190 -- 911 -- 1686 --

Tomotherapy 543 -- 2637 -- 4876