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59.Thomas SR. Options for radionuclide therapy: from fixed activity to patient-specific treatment planning. Cancer Bioth Radiopharm 2002;17:71–81.

60.Maxon HR, Thomas SR, Hertzberg VS. Relation between effective radiation dose and outcome of radioiodine therapy for thyroid cancer. N Engl J Med 1983;309:937–941.

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62.Dorn R, et al. Dosimetry guided radioactive iodine treatment in patients with metastatic thyroid cancer: largest safe dose using a risk-adapted approach. J Nucl Med 2003;44:451–456

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64.Benua R, Cical N, Sonenberg M, Rawson R. The relation of radioiodine dosimetry to results and complications in the treatment of metastatic thyroid cancer. Am J Roentgenol 1962;87:171–179.

65.de Keizer B, et al. Bone marrow dosimetry and safety of high I-131 activities given after recombinant human thyroidstimulating hormone to treat metastatic differentiated thyroid cancer. J Nucl Med 2004;45:1549–1554.

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See also NUCLEAR MEDICINE, COMPUTERS IN; PHARMACOKINETICS AND PHARMACODYNAMICS.

RADIOSURGERY, STEREOTACTIC

THOMAS H. WAGNER

SANFORD L. MEEKS

M. D. Anderson Cancer Center

Orlando

Orlando, Florida

FRANK J. BOVA

University of Florida

Gainesville, Florida

INTRODUCTION

Conventional external beam radiotherapy, or teletherapy, involves the administration of radiation absorbed dose to

cure disease. The general teletherapy paradigm is to irradiate the gross lesion plus an additional volume suspected of containing microscopic disease not visible through physical examination or imaging, to a uniform dose level. External photon beams with peak photon energy in excess of 1 MeV are targeted upon the lesion site by registering external anatomy and internal radiographic anatomy to the radiation (beam) source.

Due to uncertainty and errors in positioning the patient, the radiation beam, which is directed at the lesion, may need to be enlarged to ensure that errors and uncertainty in patient positioning do not cause the radiation beam to miss some or all of the target. Unfortunately, enlarging the radiation beam results in a relatively large volume of nondiseased tissue receiving a significant radiation dose in addition to the target. For example, expansion of a 24 mm diameter spherical target volume to 26 mm to ensure that the target is fully irradiated in the presence of a 2 mm positional error will increase the irradiated volume by 60% (1). As a consequence, non-cancerous (normal) tissue in the expansion region will receive the same high dose that the target will receive. To minimize the normal tissue toxicity, the total radiation dose is delivered in many small increments (fractions), a principle first discovered by Bergonie and Tribondeau in the early twentieth century (2) and used routinely for the majority of external radiotherapy treatments.

In contrast to conventional, fractionated radiotherapy, stereotactic radiosurgery (SRS) involves the spatially precise and conformal administration of a relatively large, single dose of radiation (10–20 Gy) to a small volume of disease, thereby abandoning the advantages provided by fractionation. Hence, it is imperative to minimize the amount of normal tissue irradiated to a high dose using such an approach. Radiosurgery is commonly used to treat intracranial lesions including brain metastases, arteriovenous malformations, benign brain tumors (acoustic schwannoma, meningioma), and primary malignant brain tumors (astrocytoma, glioma, glioblastoma).

Leksel, first conceived radiosurgery for intracranial targeting in 1950 (3). His initial goal was to produce a lesion similar to one created by a radiofrequency probe but without the need to physically introduce a probe into the brain. The lesion was to be created by a very concentrated single high dose of radiation. Stereotactic targeting and arc-centered stereotaxis methods were already known to Leksell. In his initial design, Leksell mounted a therapeutic X-ray tube onto an arc-centered frame with the axis of rotation positioned in the target tissues. This approach allowed many different paths of radiation to converge on the target tissues producing a highly concentrated dose at the intersection point. While this concept did provide a concentrated dose, Leksell continued to investigate alternate radiation delivery systems in hopes of finding a better system that could produce a high concentration of radiation at the target tissues while providing more normal tissue sparing. In later designs Leksell attempted to use particle beams to take advantage of the known Bragg peak effect. The physical limitations of the particle beam delivery systems as well as the expense of the device encouraged Leksell to continue his development finally arriving at a

design based on 201 pencil thin cobalt-60 gamma radiation sources arranged on a hemisphere and focused at a single point. This device, known as the Gamma Knife (Elekta Oncology Systems), was used to treat both benign and malignant intracranial targets.

In the 1980s, several groups began to develop technology that would adapt more generally available medical linear accelerators (linacs), to deliver radiosurgical style dose distributions, thereby placing radiosurgical capabilities within the reach of many radiation therapy clinics. Betti (4) and Columbo (5) both developed linear accelerator based radiosurgery systems. Although these early linacbased systems did allow the concentration of radiation dose there existed a question as to how accurately the radiation dose could be delivered to the targeted stereotactic coordinates. Winston and Lutz addressed this issue through the use of a stereotactically positioned phantom target system (6,7). It was found that linac-based systems could maintain the accuracy of radiation beam to target coordinates to within a millimeter or two (Fig. 1 and 2). While some felt that this accuracy was adequate, it fell short of the GammaKnife claim of 0.3 mm isocentric beam accuracy. In the late 1980s Friedman and Bova (8) developed an isocentric subsystem that enabled a routine linac to achieve an

Figure 1. Modified Winston–Lutz test setup, for testing the spatial alignment of the circular radiosurgery X-ray beam with a spherical target ball. The test target sphere should be precisely aligned to the center of the X-ray field defined by the circular collimator. Several film exposures at different linac gantry rotation angles are taken.

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Figure 2. Image of a developed film from the modified Winston– Lutz test, showing the alignment between the circular radiosurgery X-ray beam with a spherical target ball. Analysis of this film image shows that in this case, the target sphere is misaligned from the center of the 20 mm diameter X-ray field by 1.06 mm in one direction, and by 0.97 mm in the other direction. Repeating this test with at least one pair of linac gantry angles gives an estimate of the spatial error inherent in the treatment delivery system.

isocentric beam accuracy of 0.2 mm, thereby matching the accuracy of the GammaKnife.

All early radiosurgery systems used a similar delivery method, namely multiple circular cross-section radiation beams converging to a common point, called the isocenter, located in the center of the target, volume with the directions chosen to minimize the overlap of beams outside the target and hence the normal tissue dose. This scheme worked well for spherical targets using a single isocenter. Non-spherical targets required the use multiple circular collimator diameters focused at multiple isocenters distributed throughout the target volume in an effort to ‘‘fill’’ the volume with dose. The ability to properly select the optimal set of sphere as well as their spacing and weighting were provided by treatment planning systems specially designed to optimize radiosurgical planning.

The next level of advance occurred in the 1990s when new computer controlled collimation devices known as multileaf collimators (MLC) were introduced that were coupled to stereotactic treatment planning software and delivery hardware designed for these new devices. While the early attempts at using MLCs had problems matching the dose conformality and steep dose gradients achieved by multiple isocentric circular collimation techniques, they nevertheless allowed complex targets to be treated more rapidly. Recently, new techniques have been developed that allow both the conformality of multiple isocenters as well as the speed of MLC delivery (9,10).

The majority of medical linear accelerators use microwave radiation in the S band to accelerate electrons and produce X rays. During the 1990s a compact medical X band linear accelerator was developed by Accuray, Inc. Adler placed this X ray source on an industrial robot gantry to create a novel stereotactic radiosurgery system called the

CyberKnife (11,12). This system based its stereotactic targeting on an integrated orthogonal X-ray system that performed real-time imaging and correction of beam orientation to compensate for patient motion during treatment. Although this method of targeting was novel in the early 1990s, targeting systems have since been introduced for use on S band medical linacs. Unlike the GammaKnife, which by design is limited to intracranial targets, the CyberKnife can be applied to targets anywhere in the body.

Although radiosurgery based on gamma and X ray sources predominate, the theoretical advantages of proton therapy beams have stimulated great interest in advancing the use of protons to treat intracranial tumors. The physics of proton beam interactions are quite different than those of a photon beam. Unlike X rays, protons have mass and charge that result in a finite range of penetration. Additionally, the density of ionization (linear energy transfer) along the track of a proton beam is greater than that of an X-ray beam, with a region of high ionization density at the end of the track known as a Bragg peak (13,14). The finite range of penetration results in zero radiation dose beyond the Bragg peak, which in theory further allows the concentration of radiation dose to a deep-seated target while sparing underlying radiosensitive structures (15–17). The theoretical advantage of the proton beam Bragg peak is somewhat tempered by practical issue that its width is usually not large enough to encompass an entire radiosurgery target, so that it becomes necessary to superimpose a multitude of proton beams of varying energies to produce a composite depth dose distribution that covers the entire target (16). Historically, proton facilities produced only stationary beams, making it very difficult to bring multiple converging beams upon the patient’s lesion. More recently rotating gantry delivery systems for protons have been introduced, which offer more flexibility in selecting beam orientations (18). Nevertheless the high cost (>$50M) of these facilities currently limits their availability to a few large metropolitan centers.

Early stereotactic targeting systems relied upon orthogonal radiographs for target localization, however, stereotactic procedures did not gain wide acceptance until the late 1980s with the development of three-dimensional (3D) treatment planning based on the use of computerized tomography (CT) imaging to obtain a 3D model of the patient. By the 1990s CT based target definition was augmented with magnetic resonance (MR) imaging that provided superior anatomical definition of the central nervous system. Because of difficulties related to MR incompatibility of stereotactic head fixation systems, MR imaging is performed without such hardware and the image set aligned or fused with CT images obtained with head fixation.

Initially, all intracranial stereotactic procedures used a rigid stereotactic head ring, or frame, screwed into the patient’s skull to achieve a rigid, reproducible geometry for CT imaging and treatment. While frame-based procedures are still the most precise radiosurgery method they obviously are invasive to the patient and place a time limit on the completion of the procedure within hours of the frame placement. Noninvasive frameless head fixation systems were subsequently developed to address these

issues. While less precise then ring-based approaches, they can be used in certain radiosurgery procedures where extreme accuracy is not required, such as treatment of brain metastases that are not located near critical structures like the brain stem or optical apparatus. Some of these systems are based upon the fitting of patients with thermoplastic face masks (19), while other systems separate the fixation and localization processes through the used of biteplates and thermoplastic masks (20,21).

Extracranial stereotactic targeting of lesions outside of the skull has been made possible through the development of a number of new technologies. One of the first was the use of ultrasound to allow the clinician to obtain a twodimensional (2D) or 3D image set of the patient in position for radiosurgery (22). The ultrasound probe is tracked during image acquisition and the image voxels are mapped to a reference that allows precise targeting of the radiation beam. To enhance the ability to target tissues these scans are often registered to pretreatment CT and MR scans. Other methods involving fixed stereotactic X-ray tubes with image intensifiers have been developed by Accuray (23) and BrainLab (24). These systems function by obtaining either orthogonal or stereoscopic radiographs that are registered to the projection of a previously obtained 3D CT dataset. While these planar X-ray localization methods work well for bony anatomy the poorer contrast of soft tissues makes them less useful for localizing targets that are not rigidly fixed to bone. More recently the development of large format amorphous silicon detectors has facilitated the development and integration of cone beam CT scanning systems onto medical linear accelerators allowing stereotactic localization and registration of soft tissue anatomy to the linear accelerator’s reference coordinate system. These new units have promise of providing unprecedented targeting accuracy to extra cranial targets.

STEREOTACTIC IMAGING AND LOCALIZATION

The uncertainty inherent to the imaging modality used can be the largest source of uncertainty the radiosurgery process. Poor imaging techniques increase this uncertainty and nullify the efforts to improve accuracy in treatment planning and delivery. Therefore, it is important to understand stereotactic imaging techniques, the increased quality assurance demands that are placed on the diagnostic imaging apparatus used, and the inherent limitations associated with each modality. Following are brief explanations of the three stereotactic imaging techniques used in radiosurgery: computed tomography, magnetic resonance imaging, and angiography.

Computed tomography is the primary modality used for radiosurgery treatment planning due to its spatial accuracy and electron density information that are both useful for accurate dose calculation and targeting. Stereotactic CT images can be obtained with a CT-compatible localizer attached to the stereotactic head ring such as the Brown– Roberts–Wells (BRW) (Fig. 3) or other commercially available designs (25,26). Since the geometry of the localizer is known relative to the head ring, stereotactic coordinates of any point in a volumetric CT image set may be

Figure 3. Computed tomography localizer attached to frame.

accurately calculated, using the localizer fiducial markers in each axial image (Fig. 4). For example, the characteristic N-shaped rods of the BRW localizer allow the x–y–z coordinates of any point in space to be mathematically determined relative to the head ring rather than relying on the CT coordinates. This method provides more accurate spatial localization, and minimizes the CT scanner quality assurance requirements. In order to minimize the inaccuracies associated with the stereotactic imaging, it is important to obtain all imaging studies with the best available spatial resolution. Typically this means reducing the uncertainty to < 1 mm by acquiring CT images at an

Figure 4. Axial CT image of patient with BRW stereotactic headframe and CT localizer attached. There are three sets of N- shaped rods; the location of any point in a CT image that contains all nine CT localizer rods, can be computed using the interrod distances to precisely define the plane of the image with respect to the BRW headframe and its coordinate system.

578 RADIOSURGERY, STEREOTACTIC

image resolution and slice spacing less than this amount. For example, a minimum 34.5 cm field of view is just large enough to image all of the stereotactic fiducials of a BRW localizer. This FOV corresponds to a pixel size of 0.67 mm for a 512 512 image matrix. In addition, current multislice diagnostic helical CT scanners can obtain CT images at 0.5–1 mm splice spacing.

Magnetic resonance (MR) imaging often provides superior tumor visualization, but spatial distortion inherent in the MR images due to magnetic field non-uniformities and patient-specific artifacts, and secondarily the lack of electron density information makes the use of MR images less desirable than CT images for radiotherapy dose calculations. Introducing a stereotactic frame and localizer into an MR imager will perturb the magnetic field producing image distortions on the order of 0.7–4 mm in each orthogonal plane (axial, sagittal, coronal) of a stereotactic MRI (27,28). Furthermore the size of the stereotactic head frame may be incompatible with the geometry of standard MRI head coil necessitating the use of a larger MRI coil, such as the standard body coil, with a consequent degradation in image quality due to a reduced signal/noise ratio. These problems are overcome by eliminating the head frame during the MR imaging procedure and using image fusion techniques to register the MR image volume to the CT image volume of the patient in the head frame. For frame-based radiosurgery, the 3D volumetric MR scan is acquired prior to head ring placement using a pulse sequence that allows a fast image acquisition to minimize image distortion due to patient movement. All currently available image correlation routines consider the MR images as rigid bodies, and do not remove local image distortions that can exist in the MR data, hence careful review of the coregistered MRI and CT image sets is essential. This comparison should focus on internal anatomy, such as the ventricles, tentorium, sulci. and avoid the external contour since it can be shifted 3–4 mm due to the fat shift (a distortion resulting in the difference in the resonant frequency of protons in fat relative to their resonant frequency in water).

The third imaging modality important to radiosurgery, angiography, is used for diagnosis and anatomic characterization of cerebral arteriovenous malformations (AVMs). Unlike volumetric CT and MR tomography, stereotactic planar angiography utilizes a set of orthogonal radiographs of a special localizer attached to the stereotactic head ring bearing radio-opaque fiducials. The stereotactic coordinates of any point within the localizer may be calculated very accurately since the geometry of the fiducials is known relative to the head ring. The orthogonal film pair is obtained with contrast injected rapidly at the location of the AVM nidus allowing excellent visualization of fine vasculature and fiducials.

The use of orthogonal images as the sole localization method for treatment planning is inadequate for accurately determining the shape, size and location of an arbitrarily shaped AVM nidus (29–31). Furthermore, overlapping structures, such as feeding or draining blood vessels, may obscure the view of the AVM nidus and will result in unnecessary irradiation of normal tissue if these blood vessels are included in the targeted volume. Because of

these issues a volumetric CT angiography image dataset (1 mm slice thickness; intravenous contrast infused at a rate of 1 cm3 s 1) is always acquired in addition to, or in replacement of, stereotactic angiography. The resultant CT images provide an accurate 3D description of the AVM nidus, along with the feeding and draining vessels.

RADIOSURGERY DELIVERY TECHNIQUES

Numerous radiosurgery techniques have been devised based noncoplanar configurations static beams or arcs. The majority of radiosurgery treatments use circular collimators to create spherical regions of high dose. The classic example of a static beam delivery system is the GammaKnife unit, which consists of 201 narrow-beam cobalt-60 sources arranged on a hemisphere. A collimation helmet containing 201 circular collimators, each of the same diameter, is placed between the hemisphere of sources and the patient’s head with the collimator’s focal point centered on the intracranial target. This produces a spherical dose distribution or shot in GammaKnife parlance. An irregular volume is treated with multiple shots whose diameters are selected based on the available helmet collimators (32). The CyberKnife robotic radiosurgery unit is also used in a similar manner to deliver treatments from fixed beam orientations using a circular collimator.

Alternatively, a conventional medical linear accelerator can be outfitted with a circular collimator and multiple (5–9) noncoplanar arc delivery used to achieve a spherical dose distribution. When used with linear accelerators, the circularly collimated beam is rotated around the target at isocenter by moving the gantry in arc mode while the patient and treatment couch are stationary, producing a parasagittal beam path around the target. Betti and Derichinsky developed their linac radiosurgery system with a special chair, the Betti chair, which moved the patient in a side to side arc motion under a stationary linac beam, and which produced a set of para-coronal arcs (4). With modern, computer controlled linear accelerators, more complex motions other than these simple arcs are possible. The Montreal technique, which involves synchronized motion of the patient couch and the gantry while the radiation beam is on, is an example of this, producing a baseball seam type of beam path (33). The rationale of using arcs with circular collimators is to concentrate radiation dose upon the target, while spreading the beam entrance and exit doses over a larger volume of nontarget tissue, theoretically reducing the overall dose and toxicity to nontarget tissue.

Radiosurgery based on circular collimators produces a spherical region of high dose with steep falloff, or gradient, that is adequate for spherical targets. Irregular target volumes require the use of multiple spheres, or isocenters, abutted together to conform the dose more closely to the shape of the target so as to minimize nontarget tissue dose (34). A consequence of the multiisocenter approach is that the shape of the total dose distribution is very sensitive to the abutment of the spherical dose distributions due to their steep falloff. For this reason, it is common practice to accept 30% or greater dose variation over the irregular target volumes using circular collimation.

The linear accelerator offers additional flexibility in that tertiary computer-controlled multileaf collimators (MLCs) may be used to produce noncircular beams and beams with nonuniform intensity profiles that conform dose distributions more closely to irregular target volumes with less dose non uniformity. The most common type of MLC consists of two banks of opposed high density metal plates, or leaves that can be moved in a plane perpendicular to the beam’s direction. The MLC can be rotated with the treatment machine’s collimator in order to align the leaves for the best fit to the target’s projected shape. The simplest use of an MLC is simply as a functional replacement for custom made beam shaping blocks, in which the rectangular MLC edges are used to approximate a continuous target outline shape (Fig. 5) (35). This field shaping can be used for either static field treatments or for dynamic arcs in which the MLC shape is continually changed to match the beam’s-eye-view projection of the target volume. Moss investigated the efficacy of performing radiosurgery treatments with a dynamically conforming MLC in arc mode, and concluded that dynamic arc MLC treatments offered target coverage and normal tissue sparing comparable to that offered by single and multiple isocenter radiosurgery (36). Nedzi (37) showed that even crude beam shaping devices offered some conformal benefit over single isocenter treatments with circular collimators. Since the midto late-1990s, the use of miniature multileaf collimators (MLCs with a leaf width projected to isocenter of 5 mm or less) has become increasingly common.

However, the MLC may be used in a more sophisticated fashion to form many different beam shapes of arbitrary size and intensity (by varying the amount of radiation applied through each beam aperture). In this manner, radiation fields with a similar dose profile as a shaped, wedged field may be delivered using only the computercontrolled MLC, shall can also deliver intensity modulated dose profiles similar to those achievable using custom beam compensators, but without the disadvantages of fabrication time or of needing to manually change a physically mounted beam filter between each treatment field (38).

Thus, a computer-controlled MLC and treatment machine offer the potential to deliver more sophisticated radiation treatments to each patient with the same time and cost resources available.

The MLC-based solutions are available for both static multiple beams and arc-based delivery. Radionics initially introduced the use of a mini-MLC for defining static beam shapes that conformed to the projected shape of the target volume in the beam’s eye view (35,39,40). The device consisted of multiple thin plates, or MLC leaves, that were mechanically clamped together to form an irregular beam shape defined by a plastic template corresponding to the projected shape of the target. Subsequent developments by other vendors added computer-controlled motorization to the leaves so that treatments could be carried out more efficiently. While most mini-micro-MLC implementations were based on static delivery of few fixed beams, NOMOS, Inc. and 3D Line, Inc. developed specialized arc-based intensity-modulated radiosurgery (IMRS) systems. Most, if not all, tertiary MLC vendors have now developed integrated treatment planning systems designed specifically for their MLCs and treatment applications, including IMRS.

The potential for improvement presented by some of these newer and more sophisticated treatment delivery methods has spurred interest in their evaluation relative to the more traditional linac SRS methods of multiple intersecting arcs and circular collimators. These studies are usually conducted by those who have had difficulty achieving the conformality routinely published by those experienced in multiisocenter planning. These comparisons generally demonstrate that for small to medium (up to 20 cm3) intracranial targets multiple static beams offer conformity with ratios of normal tissue to target tissue treatments in the range of 1.5–2.0, with target dose homogeneity on the order of 10–20%, while offering a more standard radiation therapy treatment planning interface and process (39,41–43). These studies go on to show that static beam IMRT techniques generally performed comparable to or better than static beam plans, usually increasing the dose homogeneity and possibly conformality (44,45).

Figure 6. Ideal target (a) and nontarget volume (b) direct DVHs. Note that in the ideal direct DVH of the nontarget volume (right side), the plot is empty, since there is no nontarget volume receiving any dose in the ideal case.

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A potential problem with these comparison studies is that they may not equitably compare the full potential of multiple isocenter radiosurgery with circular collimators. A qualitative inspection of the multiple isocenter dosimetric results shown in these comparisons leads one to suspect that in many cases, suboptimal multiple isocenter plans are being compared with reasonably optimized static beam and dynamic MLC arcs–IMRT plans. Although the multiple isocenter treatment plans in these comparisons in the literature may represent a level of plan quality achievable by an average or unfamiliar user, they do not always represent the experience of expert users. Some expert users have reported on the use of multiple isocenter– circular collimator radiosurgery systems to plan and deliver tightly conformal dose distributions to irregularly shaped targets near radiosensitive structures, while maintaining a sharp dose gradient away from the target toward radiosensitive structures (34,46–48).

TOOLS FOR EVALUATING RADIOSURGERY TREATMENT PLANS

The clinical objective of radiosurgery is to deliver a tumorcidal radiation dose to a target volume while minimizing the dose to surrounding tissues. The following tools are available to the treatment planner to evaluate a 3D dose distribution in order to quantify the degree to which this objective is achieved: (1) 2D isodose curves and 3D isodose surfaces, (2) dose–volume histograms, and (3) physical dose–volume figures of merit. The following sections explain the use of each of these tools in radiation therapy and radiosurgery treatment planning.

It is possible to display 3D semitransparent surface renderings of constant dose levels overlaid on 3D renderings of the target volume to determine if the target adequately covered, but these can be difficult to analyze quantitatively. For this reason, 2D cross-sections of the 3D dose distribution are evaluated making it easier to quantitative assess target coverage. The 2D dose crosssections are displayed as isocontour plots (isodose plots) overlaid on the patient’s CT and MR images to allow visual assessment of dose coverage to an accuracy of within one image pixel. Although this implies submillimeter precision, the 1 pixel uncertainty in isodose position can result in a large uncertainty in dose coverage for small intracranial targets. In the case of a 0.67 mm pixel, a 20 mm sphere,

equal to 4.2 cm3, would apparently be equally well covered by an isodose surface ranging in volume from 3.8 to 4.6 cm3 corresponding to a 10% uncertainty in volume. Hence, although visual inspection of isodose plots on multiple images is commonly performed, it is cumbersome and there is a large uncertainty in assessing the dose coverage that is associated representing small targets using finite size pixels.

One commonly used solution to this problem is to use dose–volume histograms (DVHs). DVHs are a method of condensing 3D dose information into a more manageable form for analysis. The simplest type of DVH is a differential histogram of volume versus dose (49). This is simply a histogram showing the number of occurrences of each dose value within a 3D volume. A second more common representation is the cumulative DVH, which is the integral of the differential DVH as a function of dose. Unfortunately, in either type of DVH, the spatial information of which specific volumes are exposed to each dose level is lost in the process of constructing a DVH. For this reason, DVHs are generally used clinically in conjunction with the evaluation of multiple isodose plots as mentioned earlier.

The ideal treatment planning situation is one in which the target volume receives a uniform dose equal to the maximum dose, and the nontarget volume receives zero dose. This would correspond to ideal differential DVHs for target and nontarget volumes as shown in Fig. 6. Clinically realistic differential DVHs for target and nontarget volumes for a more typical (non-ideal) radiosurgery dose distribution are shown in Fig. 7. Figure 8 shows two differential DVHs from competing radiosurgery plans plotted on a common axes to allow a direct comparison of the plans. Note that it can difficult to evaluate competing plans using such differential histograms (50), as demonstrated in Fig. 8. Above 40 units of dose, both plans appear to be identical, but the two plans expose differing volumes of brainstem at doses less than 40 units. For this reason cumulative DVH analysis is more commonplace. Transforming the differential DVHs into cumulative DVHs, by plotting the volume receiving at least a certain dose versus dose, makes it simpler to evaluate the differences in the dose distributions, as shown in Fig. 9.

Optimal cumulative DVH curves for target structures will be as far toward the upper right hand corner of the plot as possible, while the those for nontarget structures will be as close as possible to the lower left hand corner of the

Dose (% max)

Figure 8. Direct DVHs for a radiosensitive nontarget structure in two hypothetical treatment plans.

plot as shown in Fig. 10. Considering the two completing plans shown in Fig. 9, the better plan will have its target cumulative DVH further to the upper right corner and its nontarget cumulative DVH further to the lower left corner than the poorer plan. Plan 1 is the preferred plan, since its nontarget DVH for the brainstem lies below and to the left of that for Plan 2. The relative ease of this comparison underscores the general utility of cumulative DVHs over differential DVHs (49,51). Unfortunately, it is rare for the cumulative DVHs of rival treatment plans to separate themselves from one another so cleanly. Typically, the DVHs cross one another, perhaps more than once as shown in Fig. 11. The simple rules for evaluating DVHs cannot resolve this situation, in which case other means must be used to evaluate the treatment plans by applying a score to each plan derived from a clinically relevant figure of merit.

Dose (% max)

Figure 9. Cumulative DVH plot of the direct DVH data shown in Fig. 8.

0.0

Dose (percent of maximum)

Figure 10. Ideal cumulative DVH curve for target and nontarget volumes.

The three properties of radiosurgery dose distributions that have been correlated with clinical outcome and that lend themselves to clinical figures of merit are (1) dose conformity, (2) dose gradient, and (3) dose homogeneity (34). The conformity of the dose distribution to the target volume may be simply expressed as the ratio of the prescription isodose volume to the target volume, frequently referred to as the PITV ratio (52).

PITV ¼ Prescription isodose volume=target volume (1)

Perfect conformity of a dose distribution to the target, that is, PITV ¼ 1.00, is typically not achievable, and some

Dose (relative units)

Figure 11. Crossing cumulative DVH curves.

582 RADIOSURGERY, STEREOTACTIC

Figure 12. Transaxial, sagittal, and coronal isodose distributions for five arcs of 1008 each delivered with a 30 mm collimator. Isodose lines in each plane increase from 10 to 90% in 10% increments, as indicated. The isocenter is marked with cross-hairs.

volume of nontarget tissue must be irradiated to the same dose level as the target, resulting in PITV ratios greater than unity. The most conformal treatment plans are those with the lowest PITVs, if all of the plans under comparison provide equivalent target coverage. This stipulation is necessary because the definition of PITV does not specify how the prescription isodose is determined. It is possible (but undesirable) to lower, and thus improve, the PITV by selecting an isodose level that incompletely covers the target as the prescription isodose, and therefore reduces the numerator of Eq. 1. Many investigators report isodose shells that cover in the neighborhood of at least 95% of the target volume or 99% of the target volume (17,34,46,48,53–57). This ensures a more consistent basis of comparisons for all treatment plans.

A sharp dose gradient (fall off in dose with respect to distance away from the target volume) is an important characteristic of radiosurgery and stereotactic radiotherapy dose distributions. Dose gradient may be characterized by the distance required for the dose to decrease from a therapeutic (prescription) dose level to one at which no ill effects are expected (half prescription dose). For illustrative purposes, a typical dose distribution in a hemispherical water phantom for a single isocenter delivered with five converging arcs and a 30 mm collimator is depicted in Fig. 12. A quantitative measure of gradient is obtained from examining the dose profiles along orthogonal directions in the principal anatomical planes (transaxial, sagittal, and coronal), as shown by cross-plots in Fig. 13. As in this example the steepest dose gradient (4.6 mm) occurs between the 80% and 40% isodose shells, and for this reason single isocenter dose distributions are prescribed to the 80% isodose shell (34). Table 1 lists dose gradient information between the 80 and 40% isodose shells for single isocenter spherically symmetric dose distributions with 10–50 mm diameter collimators.

A method has been proposed that uses easily obtainable DVH information to generate a numerical measure, or score, of the overall dose gradient for evaluating the dose conformality of radiosurgery dose distributions. The Conformity–Gradient Index score, or CGIg, has been proposed as a metric for quantifying dose gradient of a stereotactic treatment plan (58,59). From treatment planning experience at the University of Florida, it has been observed that it is possible to achieve a dose distribution that decreases from the prescription dose level to half of prescription dose in a distance of 3–4 mm away from the target. Taking this as a guide, a gradient score CGIg

Figure 13. Dose cross-plots through the isocenter, corresponding to the isodose distributions shown in Fig. 12. The sharpest dose fall-off, from dose D to half-dose 0.5D, occurs between dose D of 80% to 0.5D ¼ 40%, which occurs in a distance of 4.6 mm. The D to 0.5D fall-off distance is larger for 90–45% (5.1 mm) and for 70–35% 4.9 mm) doses.

may be computed as

CGIg ¼ 100 f100 ½ðReff;50%Rx Reff;RxÞ 0:3 cm&g (2)

where Reff50%Rx is the effective radius of the half-prescription isodose volume, and ReffRx is the effective radius of

the prescription isodose volume. The effective radius of a

Table 1. Single Isocenter (Five Converging Arcs) Dose– Volume and Gradient Information for 10–50 mm Diameter Circular Collimators

volume is the radius of a sphere of the same volume, so that Reff for a volume V is given by

The volumes of the prescription isodose shell and the half prescription isodose shell are obtained from a DVH of the total volume (or a sizeable volume that completely encompasses the target volume and a volume that includes all of the half prescription isodose shell) within the patient image dataset. The CGIg score is a dimensionless number that exceeds 100 for dose gradients < 3 mm (steeper falloff from prescription to half-prescription dose level), and which decreases < 100 as a linear function of the effective distance between the prescription and half-prescription isodose shells.

Dose conformity is another important characteristic of a radiosurgery treatment plan that should be considered in plan evaluation. The Conformity-Gradient Index (conformal), or CGIc, is defined as (58):

The CGIc converts PITV into a numerical score expressing the degree of conformity of a dose distribution to the target volume. The CGIg score increases as the dose gradient improves, and the CGIc score increases as dose conformity improves. Perfect conformity (assuming the target is adequately covered) of the prescription isodose volume to the target is indicated by a PITV ¼ 1.00 and a CGIc ¼ 100.

As dose gradient and dose conformity are both important parameters in judging a stereotactic radiosurgery or radiotherapy plan, an overall figure of merit for judging radiosurgery plans should incorporate both of these characteristics. Since clinical data to indicate the relative importance of conformity versus gradient is currently lacking, an index, the Conformity-Gradient Index (CGI) is proposed that assigns equal importance to both of these factors. The overall Conformity-Gradient Index score, or CGI, for a radiosurgery or radiotherapy plan is the average of the CGIc and CGIg scores:

A final measure of plan quality considered by some to be an important factor in evaluating treatment plans is dose homogeneity. While is a homogenous dose is desirable for conventional, fractionated radiotherapy (60), its role is less clear in radiosurgery. Several studies have associated large radiosurgical dose heterogeneity (maximum dose to peripheral dose ratio, or MDPD, > 2.0) with an increased risk of complications (61,62). However, some radiosurgeons have hypothesized that the statistically significant correlation between large dose inhomogeneities and complication risk may be associated with the relatively nonconformal multiple isocenter dose distributions with which some patients in these studies were treated, and not with dose inhomogeneity alone. One theory is that the extreme hot spots associated with large dose heterogenities may be acceptable, if the dose distribution is very conformal to the target volume and the hot spot is contained within the target volume. Nonconformal dose distributions could easily cause the hot spots to occur outside of the

target, greatly increasing the risk of a treatment complication. The extensive successful experience of gamma unit treatments administered worldwide (almost all treatments with MDPD 2.0) lends support to this hypothesis (63). Therefore, as a general principle, one strives for a homogeoneous radiosurgery dose distribution, but this is likely not as important a factor as conformity of the high dose region to the target volume, or the dose gradient outside of the target.

SUMMARY

While the use of radiosurgery is now in its fifth decade the basic principles of dose prescription and delivery have changed very little from those first conceived by Leksell. The primary, and still most effective method to treat relatively small target tissues with a high dose and to maintain a very steep dose gradient is through the use of many beams that all converge on the target tissue and diverge along independent paths while approaching and leaving the target region. Other dose targeting and restriction techniques, such as intensity modulation, provide the ability to position beams that geometrically avoid tissue and potentially provide a more powerful tool for dose optimization. These techniques are often combined to allow for the best of both optimized dose planning and efficient dose delivery.

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