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The issue of neutron leakage has recently been raised in connection with dedicated electron-only linear accelerators used for intraoperative radiation therapy (IORT). In particular, concern has been expressed about the degree of neutron production at energies of 10 MeV and higher due to the need for additional, perhaps permanent, shielding in the room in which the device is operated. In particular, three mobile linear accelerators available commercially offer electron energies at or above the neutron threshold, one at 9 MeV, one at 10 MeV, and the third at 12 MeV. To investigate this problem, neutron leakage has been measured around the head of two types of electron accelerators at a distance of 1 m from the target at azimuthal angles of 0 degrees, 45 degrees, 90 degrees, 135 degrees, and 180 degrees. The first is a dedicated electron-only (nonmobile) machine with electron energies of 6 (not used here), 9, 12, 15, and 18 MeV and the second a conventional machine with electron energies of 6 (also not used here), 9, 12, 16, and 20 MeV. Measurements were made using neutron bubble detectors and track-etch detectors. For electron beams from a conventional accelerator, the neutron leakage in the forward direction in Sv/Gy is 2.1 x 10(-5) at 12 MeV, 1.3 x 10(-4) at 16 MeV, and 4.2 x 10(-4) at 20 MeV, assuming a quality factor (RBE) of 10. For azimuthal angles > 0 degrees, the leakage is almost angle independent [2 x 10(-6) at 12 MeV; (0.7-1.6) x 10(-5) at 16 MeV, and (1.6-2.9) x 10(-5) at 20 MeV]. For the electron-only machine, the neutron leakage was lower than for the conventional linac, but also independent of azimuthal angle for angles > 0 degrees: {[0 degrees: 7.7 x 10(-6) at 12 MeV; 3.0 x 10(-5) at 15 MeV; 1.0 x 10(-4) at 18 MeV]; [other angles: (2.6-5.9) x 10(-7) at 12 MeV; (1.4-2.2) x 10(-6) at 15 MeV; (2.7-4.7) x 10(-6) at 18 MeV]}. Using the upper limit of 6 x 10(-7) Sv/Gy at 12 MeV for the IORT machine for azimuthal angles > 0 degrees and assuming a workload of 200 Gy/wk and an inverse square factor of 10, the neutron dose equivalent is calculated to be 0.012 mSv/wk. For the primary beam at 12 MeV (0 degrees), the 10 x higher dose would be compensated by the attenuation of a primary beam stopper in a mobile linear accelerator. These neutron radiation levels are below regulatory values (National Council on Radiation Protection and Measurements, "Limitation of exposure to ionizing radiation," NCRP Report No. 116, NCRP Bethesda, MD, 1993).

The x-ray leakage from the housing of a therapy x-ray source is regulated to be <0.1% of the useful beam exposure at 1 m from the source. It is to be expected that the machine leakage in the backward direction would be less because the gantry and stand contain significant amounts of additional metal to attenuate the x rays. A reduction in head leakage in this direction will have a direct effect on the thickness of the shielding wall behind the linear accelerator. However, no reports have been published to date on measurements in this area. The x-ray leakage in the backward direction has been measured from linacs having energies of 4, 6, 10, 15, and 18 MV using a 100 cm ionization chamber and Al2O3 dosimeters. The leakage was measured at nine different positions over the rear wall using a 3 x 3 matrix with a 1-m separation between adjacent horizontal and vertical points with either the leftmost or rightmost column aligned with the target and isocenter. In general, the leakage is less than the canonical value, but the exact value depends on energy, gantry angle, and measurement position. There is significantly greater attenuation directly behind the gantry stand for all energies. Leakage at 10 MV for some positions exceeded 0.1%. Additionally, neutron leakage measurements were made for 10, 15, and 18 MV x-ray beams using track-etch detectors. The average neutron leakage was less than 0.1% except for 18 MV, where neutron leakage was more than 0.1% of the useful beam at some positions.

Intraoperative radiation therapy (IORT) has been customarily performed either in a shielded operating suite located in the operating room (OR) or in a shielded treatment room located within the Department of Radiation Oncology. In both cases, this cancer treatment modality uses stationary linear accelerators. With the development of new technology, mobile linear accelerators have recently become available for IORT. Mobility offers flexibility in treatment location and is leading to a renewed interest in IORT. These mobile accelerator units, which can be transported any day of use to almost any location within a hospital setting, are assembled in a nondedicated environment and used to deliver IORT. Numerous aspects of the design of these new units differ from that of conventional linear accelerators. The scope of this Task Group (TG-72) will focus on items that particularly apply to mobile IORT electron systems. More specifically, the charges to this Task Group are to (i) identify the key differences between stationary and mobile electron linear accelerators used for IORT, (ii) describe and recommend the implementation of an IORT program within the OR environment, (iii) present and discuss radiation protection issues and consequences of working within a nondedicated radiotherapy environment, (iv) describe and recommend the acceptance and machine commissioning of items that are specific to mobile electron linear accelerators, and (v) design and recommend an efficient quality assurance program for mobile systems.

The calibration and monthly QA of an electron-only linear accelerator dedicated to intra-operative radiation therapy has been reviewed. Since this machine is calibrated prior to every procedure, there was no necessity to adjust the output calibration at any time except, perhaps, when the magnetron is changed, provided the machine output is reasonably stable. This gives a unique opportunity to study the dose output of the machine per monitor unit, variation in the timer error, flatness and symmetry of the beam and the energy check as a function of time. The results show that, although the dose per monitor unit varied within +/- 2%, the timer error within +/- 0.005 MU and the asymmetry within 1-2%, none of these parameters showed any systematic change with time. On the other hand, the energy check showed a linear drift with time for 6, 9, and 12 MeV (2.1, 3.5, and 2.5%, respectively, over 5 years), while at 15 and 18 MeV, the energy check was relatively constant. It is further shown that based on annual calibrations and RPC TLD checks, the energy of each beam is constant and that therefore the energy check is an exquisitely sensitive one. The consistency of the independent checks is demonstrated.

Wedge transmission factors have been measured for two sets of physical wedges for the 4 MV beam of a new dual energy linear accelerator as a function of field size and depth. The field size dependence of these factors has been compared with the 4 MV beam from another, single-energy linear accelerator and a difference of a factor of approximately 2 has been observed in the relative wedge transmission factors between the two machines.

A significant timer error has been found on a linear accelerator used for intraoperative radiation therapy. Since typical treatment doses range up to 20 Gy, calibrations that do not account for this effect can incur a large dosimetry error. The effect is dose rate dependent, but energy independent. For a dose rate of 900 cGy min(-1), the error in dose calibration varies between 2 and 3% for an average of 2.5%. At dose rates of 600 and 300 cGy min(-1), the error was 1.7% and 0.8% respectively. The average timer error at 900 cGy min(-1) was found to be 1.2 +/- 0.1 monitor units. It is argued that even a small timer error has serious dosimetry consequences for the implementation of intensity modulated radiation therapy.

Ensuring safe operation for a medical accelerator is a difficult task. Users must assume more responsibility in using contemporary equipment. Additionally, users must work closely with manufacturers in promoting the safe and effective use of such complex equipment. Complex treatment techniques and treatment modality changeover procedures merit detailed, unambiguous written procedural instruction at the control console. A thorough "hands on" training period after receiving instructions, and before assuming treatment responsibilities, is essential for all technologists. Unambiguous written instructions must also be provided to guide technologists in safe response when equipment malfunctions or exhibits unexpected behavior or after any component has been changed or readjusted. Technologists should be given a written list of the appropriate individuals to consult when unexpected machine behavior occurs. They should be assisted in identifying aberrant behavior of equipment. Many centers already provide this instruction, but others may not. Practiced response and discussion with technologists should be a part of an ongoing quality assurance program. An important aspect of a safety program is the need for continuous vigilance. Table III gives a summary of a comprehensive safety program for medical accelerators. Table IV gives a list of summary recommendations as an example of how one might mitigate the consequences of an equipment failure and improve procedures and operator response in the context of the environment described. Most of these recommendations can be implemented almost immediately at any individual treatment center.

A detailed Monte Carlo approach has been employed to investigate the sources of electron contamination for the 25-MV photon beam generated by Varian's Clinac-35. Three sources of contamination were examined: (a) the flattening filter and beam monitor chamber, (b) the fixed primary collimators downstream from the monitor chamber and the adjustable photon jaws, and (c) the air volume separating the treatment head from the observation point. Five source-to-surface distances (SSDs) were considered for a single field size, 28 cm in diameter at 80 cm SSD. It was found that for small SSDs (80-100 cm), the dominant sources of electron contamination were the flattening filter and the beam monitor chamber which accounted for 70% of the unwanted electrons. Thirteen percent of the remaining electrons originated in the downstream primary collimators and the photon jaws, and 17% were produced in air. At larger SSDs, the fraction of unwanted electrons originating in air increased. At 400 cm SSD, 61% of the contaminating electrons present in the beam were produced in air, 34% originated in the flattening filter and beam monitor chamber, and 5% were due to interactions in the fixed collimators downstream from the monitor chamber and the adjustable photon jaws. These calculated results are substantiated by recent experiments.

A new version of a manufacturer's couch no longer has rails that span the length of each side of the couch. Modification of the original couch to include a spinal section and a stretched mylar top have split this rail into three sections. The gaps between these sections have been fitted with interlocking bars that can readily be removed. By this means, the usefulness of the earlier model couch in terms of attaching treatment devices and aids to any part of the side has been restored, while at the same time not interfering with the demands of the new, multipurpose couch.

We have made a device that improves the dose distributions of X ray fields from a 25 MV accelerator by increasing the width of the 95% dose relative to the 50%. This device, called a beam width improving device, is a specially designed "compensator," consisting of two lead sheets taped onto a plastic blocking tray and is tailored for each field so that it can accommodate square fields, rectangular fields and fields with corner cuts. Comparison of profiles with and without this compensator indicates an increase in the width of the 95% dose of approximately 8 mm for small fields to approximately 14 mm for large fields.

An attachment to an existing linear accelerator couch is described that extends the range of SSD's at which patients can be treated. This couch attachment, which is mounted on castors for ease of mobility, is attached to and locks into the rail of the accelerator couch such that the axis of the new device is normal to the existing couch. All motions of this assembly are then under the control of the existing couch. The overall size of the couch attachment is 215 cm x 90 cm, sufficiently large to treat all clinically used fields.

With an increasing number of high energy accelerators in operation, intra-oral electron radiotherapy is likely to become a more widely-used modality in the treatment of lesions of the oral cavity. However, there is one potential problem associated with this modality which concerns patient safety. There must never be any uncontrolled movement of the couch or gantry while the cone is in the patient's mouth, otherwise serious consequences could occur. In an effort to overcome this problem, a set of cones previously constructed by us for use in intra-oral electron radiotherapy has been modified to include a breakaway safety feature. This modification consists of separating the plate into which each of the treatment cones screws, into two pieces, the dividing line being in the shape of a cone frustum flaring out in the upward direction. A set of four spring plungers projecting into a notch on the inner ring prevents the cone from being pushed upwards without the application of an upward force. The amount of force required to push the cone upwards can be changed by adjusting the pressure of the spring plungers.

A method is described for computing the effective center of an x-ray spoke shot pattern and determining its separation from a given reference point, such as the mechanical isocenter. This method is as quick as, but more accurate than, the standard technique of using a ruler.

The quality of portal films taken at megavoltage energies is frequently poor, despite the use of various techniques such as lead screens to improve the image. When critical structures, such as the spinal cord, are to be blocked out of the treatment field, it is frequently difficult to determine whether the block is correctly placed. To solve this problem, a diagnostic X ray source has been mounted on the side of a 10 MV linear accelerator to provide accurate verification of patient positioning and location of shielding blocks. The coincidence between the mechanical, 10 MV and diagnostic X ray isocenters is about 1 mm. The system has been designed so that procedures to monitor this coincidence, and adjustment procedures to maintain it, are easily performed.