Patient-specific radiation dose and cancer risk estimation in CT: part I. development and validation of a Monte Carlo program

Abstract

Purpose: Radiation-dose awareness and optimization in CT can greatly benefit from a dose-reporting system that provides dose and risk estimates specific to each patient and each CT examination. As the first step toward patient-specific dose and risk estimation, this article aimed to develop a method for accurately assessing radiation dose from CT examinations.

Methods: A Monte Carlo program was developed to model a CT system (LightSpeed VCT, GE Healthcare). The geometry of the system, the energy spectra of the x-ray source, the three-dimensional geometry of the bowtie filters, and the trajectories of source motions during axial and helical scans were explicitly modeled. To validate the accuracy of the program, a cylindrical phantom was built to enable dose measurements at seven different radial distances from its central axis. Simulated radial dose distributions in the cylindrical phantom were validated against ion chamber measurements for single axial scans at all combinations of tube potential and bowtie filter settings. The accuracy of the program was further validated using two anthropomorphic phantoms (a pediatric one-year-old phantom and an adult female phantom). Computer models of the two phantoms were created based on their CT data and were voxelized for input into the Monte Carlo program. Simulated dose at various organ locations was compared against measurements made with thermoluminescent dosimetry chips for both single axial and helical scans.

Results: For the cylindrical phantom, simulations differed from measurements by -4.8% to 2.2%. For the two anthropomorphic phantoms, the discrepancies between simulations and measurements ranged between (-8.1%, 8.1%) and (-17.2%, 13.0%) for the single axial scans and the helical scans, respectively.

Conclusions: The authors developed an accurate Monte Carlo program for assessing radiation dose from CT examinations. When combined with computer models of actual patients, the program can provide accurate dose estimates for specific patients.

Figure 1

(a) A point source and an effective beam width were used in the simulations to account for the dose delivered by both the umbra and the penumbra regions of the beam. (b) The fan beam was chosen to be just broad enough to cover the imaging object. A pseudoimpact detector was added below the bowtie filter at the level of the tungsten cam collimators to register the information of each incident particle in a phase-space file for use in the subsequent simulations of axial and helical scans.

Figure 2

(a) Custom-designed CT dose phantom for measuring center-to-periphery dose distributions. The locations of the seven drill holes are numbered. (b) Measurements of dose distribution in the custom-designed phantom. The phantom was attached to one end of the CT table and positioned so that its long axis matched the axis of gantry rotation.

Figure 3

Dose measurements in (a) pediatric one-year-old CIRS phantom and (b) adult female CIRS phantom. (c) The phantoms were composed of axially sliced, 25 mm thick, contiguous sections. Each section contained one or more 5 mm diameter through-holes at various organ locations.

Figure 4

Model of the ion chamber (6.6 mm long, 0.18 cm³ active volume, model 10×5-0.18∕9015, Radcal Corporation, Monrovia, CA) used in the Monte Carlo simulations. “C552” here refers to C552 air-equivalent plastic. Material data file for polyoxymethylene was used for the polyacetal cap.

Figure 5

Simulated x-ray energy spectra at the exit of the x-ray tube and before filtration by the bowtie filter (prebowtie spectra).

Figure 6

Measured and simulated dose distributions in the custom-designed cylindrical phantom for single axial scans at four kVp values using the (a) small, (b) medium, and (c) large bowtie filters and a 40 mm beam collimation. Simulated dose values are shown by lines. Measured dose values are shown by symbols and their error bars reflect one standard deviation. Most error bars are too small to appreciate.

Figure 7

Measured and simulated dose from a single axial scan in (a) the pediatric and (b) the adult female phantoms. Error bars reflect one standard deviation. Percent discrepancies between simulation and measurement are labeled on the bottom. The high dose values reported here were results of the high-dose scan parameters (Table 1) used in the measurements and simulations and were not typical of clinical dose values.

Figure 8

Measured and simulated dose from (a) a full-body helical scan in the pediatric phantom and (b) a chest scan in the adult female phantom. Error bars reflect one standard deviation. The degrees are x-ray tube start angles relative to 12 o’clock. Percent discrepancies between simulation (averaged over tube starting angles) and measurement are labeled on the bottom. At four organ locations in the adult female phantom, one of the two TLD chips was cracked during the experiment; dose uncertainty could not be assessed for these four locations. The high dose values reported here were results of the high-dose scan parameters (Table 1) used in the measurements and simulations and were not typical of clinical dose values.