A realistic model of photon beams generated by clinical linear accelerators has been incorporated in a convolution/superposition method to compute dose distributions in photon treatment fields. In this beam model, a primary photon source represents photons directly from the target, and an extra-focal photon source represents scattered photons from the primary collimator and the flattening filter. Monte Carlo simulation was used to study clinical linear accelerators producing photon beams. From the output of the Monte Carlo simulation, the fluence and spectral distributions of each photon component, as well as the geometrical characteristics of each photon source with respect to its distance to the isocenter and its source distribution, were analyzed. These quantities were used to reproduce realistic photon distributions in treatment fields, and thus to compute dose distributions using the convolution method. Our results showed that compared to the primary photon fluence, the extra-focal photon fluence from the primary collimator and the flattening filter was 11%-16% at the isocenter, among which 70% was contributed by the flattening filter. The variation of extra-focal photons in different treatment fields was predicted accurately by accounting for the finite size of the extra-focal source. Compared to measurements, dose distributions in photon treatment fields, including those of asymmetric jaw settings and at different SSDs were calculated accurately, particularly in the penumbral region, by using the convolution method with the new dual source photon beam model.

A realistic photon beam model based on Monte Carlo simulation of clinical linear accelerators was implemented in a convolution/superposition dose calculation algorithm. A primary and an extra-focal sources were used in this beam model to represent the direct photons from the target and the scattered photons from other head structures, respectively. The effect of the finite size of the extra-focal source was modeled by a convolution of the source fluence distribution with the collimator aperture function. Relative photon output in air (Sc) and in phantom (Scp) were computed using the convolution method with this new photon beam model. Our results showed that in a 10 MV photon beam, the Sc, Sp (phantom scatter factor), and Scp factors increased by 11%, 10%, and 22%, respectively, as the field size changed from 3 x 3 cm2 to 40 x 40 cm2. The variation of the Sc factor was contributed mostly by an increase of the extra-focal radiation with field size. The radiation backscattered into the monitor chamber inside the accelerator head affected the Sc by about 2% in the same field range. The output factors in elongated fields, asymmetric fields, and blocked fields were also investigated in this study. Our results showed that if the effect of the backscattered radiation was taken into account, output factors in these treatment fields can be predicted accurately by our convolution algorithm using the dual source photon beam model.

We have developed a convolution/superposition method to calculate dose distributions in photon treatment fields with beam modifiers such as physical wedges. The dose component due to wedge generated radiation was accounted for by using an extended phantom model, which integrated a wedge, an air gap, and a patient phantom as the calculation phantom. The inhomogeneities in the extended phantom and the effect of beam hardening by the wedge were both corrected for in the convolution dose calculation. The calculated dose was verified by Monte Carlo simulation of the same extended phantom. A new dual photon source model was also used in the convolution method to account for both primary photons from the target and extra-focal photons from the primary collimator and flattening filter. Thus, realistic photon energy fluence distributions in the extended phantom were used for the dose calculation. The calculated dose distributions and the wedge factors agreed with the measured data within 2% for a variety of treatment fields including asymmetric fields. Our results showed that the wedge-generated radiation could contribute a significant fraction of the total dose in patients. This dose component depends on a specific field configuration, thus wedge factor changes with photon energy, wedge angle, field size, depth, and patient phantom SSD. The variation of the wedge factor can be predicted accurately by our convolution approach with the extended phantom model, which allows for more accurate dose or monitor unit computation for photon fields with beam modifiers.

To account for clinical divergent and polychromatic photon beams, we have developed kernel tilting and kernel hardening correction methods for convolution dose calculation algorithms. The new correction methods were validated by Monte Carlo simulation. The accuracy and computation time of the our kernel tilting and kernel hardening correction methods were also compared to the existing approaches including terma divergence correction, dose divergence correction methods, and the effective mean kernel method with no kernel hardening correction. Treatment fields of 10 x 10-40 x 40 cm2 (field size at source to axis distance (SAD)) with source to source distances (SSDs) of 60, 80, and 100 cm, and photon energies of 6, 10, and 18 MV have been studied. Our results showed that based on the relative dose errors at a depth of 15 cm along the central axis, the terma divergence correction may be used for fields smaller than 10 x 10 cm2 with a SSD larger than 80 cm; the dose divergence correction with an additional kernel hardening correction can reduce dose error and may be more applicable than the terma divergence correction. For both these methods, the dose error increased linearly with the depth in the phantom; the 90% isodose lines at the depth of 15 cm were shifted by about 2%-5% of the field width due to significant underestimation of the penumbra dose. The kernel hardening effect was less prominent than the kernel tilting effect for clinical photon beams. The dose error by using nonhardening corrected kernel is less than 2.0% at a depth of 15 cm along the central axis, yet it increased with a smaller field size and lower photon energy. The kernel hardening correction could be more important to compute dose in the fields with beam modifiers such as wedges when beam hardening is more significant. The kernel tilting correction and kernel hardening correction increased computation time by about 3 times, and 0.5-1 times, respectively. This can be justified by more accurate dose calculations for the majority of clinical treatments.

This paper investigates measuring dose distributions for enhanced dynamic wedges (EDWs) using a commercial multichamber detector array. The technical aspects of using the chamber array, including chamber calibration, selection of measurement parameters, and use of the reference chamber, have been fully investigated. The measurement results from the chamber array were also confirmed by those from the single chamber and radiographic film measurements. The results reported here showed that proper operation of the chamber array is essential to measure dose accurately for the EDW fields; the chamber detector array can be used more efficiently than a single chamber without compromising the dose measurement accuracy.