Radio-sensitizing effect of MEK inhibition in glioblastoma in vitro and in vivo.

INTRODUCTION: Glioblastoma (GBM) is an incurable cancer type. New therapeutic options are investigated, including targeting the mitogen-activated protein kinase (MAPK) pathway using MEK inhibitors as radio-sensitizers. In this study, we investigated whether MEK inhibition via PD0325901 leads to radio-sensitization in experimental in vitro and in vivo models of GBM. MATERIALS AND METHODS: In vitro, GBM8 multicellular spheroids were irradiated with 3 fractions of 2 Gy, during 5 consecutive days of incubation with either PD0325901 or MEK-162. In vivo, we combined PD0325901 with radiotherapy in the GBM8 orthotopic mouse model, tumor growth was measured weekly by bioluminescence imaging and overall survival and toxicity were assessed. RESULTS: Regrowth and viability of spheroids monitored until day 18, showed that both MEK inhibitors had an in vitro radio-sensitizing effect. In vivo, PD0325901 concentrations were relatively constant throughout multiple brain areas and temporal PD0325901-related adverse events such as dermatitis were observed in 4 out of 14 mice (29%). Mice that were treated with radiation alone or combined with PD0325901 had significantly better survival compared to vehicle (both P < 0.005), however, no significant interaction between PD0325901 MEK inhibition and irradiation was observed. CONCLUSION: The difference between the radiotherapy-enhancing effect of PD0325901 in vitro and in vivo urges further pharmacodynamic/pharmacokinetic investigation of PD0325901 and possibly other candidate MEK inhibitors.

How flatbed scanners upset accurate film dosimetry.

Film is an excellent dosimeter for verification of dose distributions due to its high spatial resolution. Irradiated film can be digitized with low-cost, transmission, flatbed scanners. However, a disadvantage is their lateral scan effect (LSE): a scanner readout change over its lateral scan axis. Although anisotropic light scattering was presented as the origin of the LSE, this paper presents an alternative cause. Hereto, LSE for two flatbed scanners (Epson 1680 Expression Pro and Epson 10000XL), and Gafchromic film (EBT, EBT2, EBT3) was investigated, focused on three effects: cross talk, optical path length and polarization. Cross talk was examined using triangular sheets of various optical densities. The optical path length effect was studied using absorptive and reflective neutral density filters with well-defined optical characteristics (OD range 0.2-2.0). Linear polarizer sheets were used to investigate light polarization on the CCD signal in absence and presence of (un)irradiated Gafchromic film. Film dose values ranged between 0.2 to 9 Gy, i.e. an optical density range between 0.25 to 1.1. Measurements were performed in the scanner's transmission mode, with red-green-blue channels. LSE was found to depend on scanner construction and film type. Its magnitude depends on dose: for 9 Gy increasing up to 14% at maximum lateral position. Cross talk was only significant in high contrast regions, up to 2% for very small fields. The optical path length effect introduced by film on the scanner causes 3% for pixels in the extreme lateral position. Light polarization due to film and the scanner's optical mirror system is the main contributor, different in magnitude for the red, green and blue channel. We concluded that any Gafchromic EBT type film scanned with a flatbed scanner will face these optical effects. Accurate dosimetry requires correction of LSE, therefore, determination of the LSE per color channel and dose delivered to the film.

Accurate dosimetry with GafChromic EBT film of a 6 MV photon beam in water: what level is achievable?

This paper focuses on the accuracy, in absolute dose measurements, with GafChromicTM EBT film achievable in water for a 6 MV photon beam up to a dose of 2.3 Gy. Motivation is to get an absolute dose detection system to measure up dose distributions in a (water) phantom, to check dose calculations. An Epson 1680 color (red green blue) transmission flatbed scanner has been used as film scanning system, where the response in the red color channel has been extracted and used for the analyses. The influence of the flatbed film scanner on the film based dose detection process was investigated. The scan procedure has been optimized; i.e. for instance a lateral correction curve was derived to correct the scan value, up to 10%, as a function of optical density and lateral position. Sensitometric curves of different film batches were evaluated in portrait and landscape scan mode. Between various batches important variations in sensitometric curve were observed. Energy dependence of the film is negligible, while a slight variation in dose response is observed for very large angles between film surface and incident photon beam. Improved accuracy in absolute dose detection can be obtained by repetition of a film measurement to tackle at least the inherent presence of film inhomogeneous construction. We state that the overall uncertainty is random in absolute EBT film dose detection and of the order of 1.3% (1 SD) under the condition that the film is scanned in a limited centered area on the scanner and at least two films have been applied. At last we advise to check a new film batch on its characteristics compared to available information, before using that batch for absolute dose measurements.

Formalisms for MU calculations, ESTRO booklet 3 versus NCS report 12.

Although the relevance and importance of quality assurance and quality control in radiotherapy is generally accepted, only recently, methods for monitor unit (MU) calculation and verification have been addressed in recognized recommendations, published by the European Society of Therapeutic Radiation Oncology (ESTRO) and by the Netherlands Commission on Radiation Dosimetry (Dutreix A, Bjärngard BE, Bridier A, Mijnheer B, Shaw JE, Svensson H. Monitor unit calculation for high-energy photon beams. Physics for clinical radiotherapy. ESTRO Booklet No. 3. Leuven: Garant, 1997; Netherlands Commission on Radiation Dosimetry (NCS). Determination and use of scatter correction factors of megavoltage photon beams. NCS report 12. Deift: NCS, 1998). Both documents are based on the same principles: (i) the separation of the output factor into a head and a volume (or phantom) scatter component; (ii) the use of a so-called mini-phantom to measure and verify the head scatter component; and (iii) the recommendation to use a single reference depth of 10 cm for all photon beam qualities. However, there are substantial differences between the approach developed in the IAEA-ESTRO task group and the NCS approach for MU calculations, which might lead to confusion and/or misinterpretation if both reports are used simultaneously or if data from the NCS report is applied in the algorithms of the ESTRO report without careful consideration. The aim of the present paper is to discuss and to clearly point out these differences (e.g. field size definitions, phantom scatter parameters, etc.). Additionally, corresponding quantities in the two reports are related where possible and several aspects concerning the use of a mini-phantom (e.g. size, detector position, composition) are addressed.

Dependence of the tray transmission factor on collimator setting and source-surface distance.

When blocks are placed on a tray in megavoltage x-ray beams, generally a single correction factor for the attenuation by the tray is applied for each photon beam quality. In this approach, the tray transmission factor is assumed to be independent of field size and source-surface distance (SSD). Analysis of a set of measurements performed in beams of 13 different linear accelerators demonstrates that there is, however, a slight variation of the tray transmission factor with field size and SSD. The tray factor changes about 1.5% for collimator settings varying between 4x4 cm and 40 x 40 cm for a 1 cm thick PMMA tray and approximately 3% for a 2 cm thick PMMA tray. The variation with field size is smaller if the source-surface distance is increased. The dependence on the collimator setting is not different, within the experimental uncertainty of about 0.5% (1 s.d.), for the nominal accelerating potentials and accelerator types applied in this study. It is shown that the variation of the tray transmission factor with field size and source-surface distance can easily be taken into account in the dose calculation by considering the volume of the irradiated tray material and the position of the tray in the beam. A relation is presented which can be used to calculate the numerical value of the tray transmission factor directly. These calculated values can be checked with only a few measurements using a cylindrical beam coaxial miniphantom.

Effect of electron contamination on scatter correction factors for photon beam dosimetry.

Physical quantities for use in megavoltage photon beam dose calculations which are defined at the depth of maximum absorbed dose are sensitive to electron contamination and are difficult to measure and to calculate. Recently, formalisms have therefore been presented to assess the dose using collimator and phantom scatter correction factors, Sc and Sp, defined at a reference depth of 10 cm. The data can be obtained from measurements at that depth in a miniphantom and in a full scatter phantom. Equations are presented that show the relation between these quantities and corresponding quantities obtained from measurements at the depth of the dose maximum. It is shown that conversion of Sc and Sp determined at a 10 cm depth to quantities defined at the dose maximum such as (normalized) peak scatter factor, (normalized) tissue-air ratio, and vice versa is not possible without quantitative knowledge of the electron contamination. The difference in Sc at dmax resulting from this electron contamination compared with Sc values obtained at a depth of 10 cm in a miniphantom has been determined as a multiplication factor, Scel, for a number of photon beams of different accelerator types. It is shown that Scel may vary up to 5%. Because in the new formalisms output factors are defined at a reference depth of 10 cm, they do not require Scel data. The use of Sc and Sp values, defined at a 10 cm depth, combined with relative depth-dose data or tissue-phantom ratios is therefore recommended. For a transition period the use of the equations provided in this article and Scel data might be required, for instance, if treatment planning systems apply Sc data normalized at d(max).

A consistent formalism for the application of phantom and collimator scatter factors.

A coherent system for the use of scatter correction factors, determined at 10 cm depth, is described for dose calculations on the central axis of arbitrarily shaped photon beams. The system is suitable for application in both the fixed source-surface distance (SSD) and in the isocentric treatment set-up. This is in contrast to some other proposals where only one of these approaches forms the basis of the calculation system or where distinct quantities and data sets are needed. In order to derive the relations in the formalism, we introduced a separation of the phenomena related to the energy fluence in air and to the phantom scatter contribution to the dose. Both are used relative to quantities defined for the reference irradiation set-up. It is shown that dose calculations can be performed with only one set of basic beam data, obtained at a reference depth of 10 cm. These data consist for each photon beam quality of measured collimator and phantom scatter correction factors, in combination with a set of (percentage/relative) depth-dose or tissue-phantom ratio values measured along the central axis of the beam. Problems related to measurements performed at the depth of maximum absorbed dose, due to the electron contamination of the beam, are avoided in this way. Collimator scatter correction factors are obtained by using a mini-phantom, while phantom scatter correction factors are derived from measurements in a full scatter phantom in combination with the results of the mini-phantom measurements. For practical reasons the fixed SSD system was chosen to determine the data. Then, dose calculations in a fixed SSD treatment set-up itself are straightforward. Application in the isocentric treatment set-up needs simple conversion steps, while the inverse approach, from isocentric to fixed SSD, is described as well. Differences between the two approaches are discussed and the equations for the conversions are given.

A three-Gaussian fit of phantom scatter correction data.

The phantom scatter correction factor Sp of megavoltage photon beams can be accurately described using a three-Gaussian fit. The model leads to six parameters, with which Sp(r) is described as a smooth function of the field radius r for beam qualities in the range from 60Co up to 25 MV. The parameters allow Sp values to be calculated at intermediate beam energies and for any field shape. Calculated Sp(X, Y) values for rectangular fields (X, Y) can be subsequently used as reference values to compare with measured Sp(X, Y) values, for example when appraising a new beam.

Comparison of parametrization methods of the collimator scatter correction factor for open rectangular fields of 6-25 MV photon beams.

PURPOSE: To facilitate the use of the collimator scatter correction factor, Sc, parametrization methods that relate Sc to the field size by fitting were investigated. MATERIALS AND METHODS: Sc was measured with a mini-phantom for five types of dual photon energy accelerators with energies varying between 6 and 25 MV. Using these Sc-data six methods of parametrizing Sc for square fields were compared, including a third-order polynomial of the natural logarithm of the field size normalized to the field size of 10 cm2. Also five methods of determining Sc for rectangular fields were considered, including one which determines the equivalent field size by extending Sterling's method. RESULTS: The deviations between measured and calculated Sc-values were determined for all photon beams and methods investigated in this study. The resulting deviations of the most accurate method varied between 0.07 and 0.42% for square fields and between 0.26 and 0.79% for rectangular fields. A recommendation is given as to how to limit the number of fields for which Sc should be measured in order to be able to accurately predict it for an arbitrary field size.

Is there a need for a revised table of equivalent square fields for the determination of phantom scatter correction factors?

The use of the British Journal of Radiology (BJR) (supplement 17) tables of equivalent square fields for dose calculations is widespread. A revised version of the supplement was published recently, with a more elaborate discussion, but without changes in data given in these tables (Br. J. Radiol. suppl 25). The tables were generated for use in dose calculations, with relative beam data such as PDD, BSF, PSF, all with d(max) as the reference depth. However, the current philosophy in dose calculational methods is based on quantities defined at a reference depth, d(ref) = 10 cm, on a separation of phantom and head scatter, and on the use of the relative depth-dose or tissue-phantom ratios normalized at d(ref). By using these quantities as a starting point, problems at shallow depths related to the influence of contaminating electrons in the beam can be eliminated. Recently, a comprehensive set of phantom scatter factor data with d(ref) = 10 cm has been published for a set of square field sizes and a wide range of photon beam energies, showing that phantom scatter is a smoothly varying function of field size and quality index. It is not a priori evident that the conventional concept of equivalent squares for rectangular fields is also fully applicable for phantom scatter factors and phantom scatter related quantities at a depth of 10 cm. It was questioned whether or not new tables of equivalent square fields are needed for this purpose. In this paper, new tables have been constructed for four photon beam energies in the range of Co-60 to 25 MV (quality index from 0.572 to 0.783). The small differences between the outcome of these new tables allowed the construction of one averaged table of equivalent square fields. Phantom scatter factors were calculated for rectangular fields based on the use of the BJR table and on the use of the newly constructed tables and the differences were quantified. For Co-60 no improvements could be shown when using the new averaged table, but for beam energies of 6 to 10 MV small improvements of the order of 0.5 to 1.0% were found. For a higher beam energy of 25 MV the improvement is smaller. Deviations resulting from the BJR table are within the limits of accuracy as stated by the authors. Therefore, for clinical use, the continued use of the BJR table of equivalent squares for phantom scatter factors and phantom scatter related quantities of rectangular fields is justified, irrespective of photon beam energy.

Differences in wedge factor determination in air using a PMMA mini-phantom or a brass build-up cap.

The head scatter dose contribution to the output of a treatment machine has been determined for an open and wedged 60Co gamma-ray beam and for open and wedged x-ray beams of 4, 8, and 16 MV. From those data wedge factor values "in air" have been deduced, expressed as the ratio of the dose to water, measured in air, for the situation with and without wedge, for the same number of monitor units (or treatment time for 60Co). The measurements have been performed using a polymethyl-metacrylate (PMMA) and a graphite-walled ionization chamber inserted in a brass build-up cap and in a PMMA mini-phantom, respectively. Absolute wedge factor values deduced with both detector systems and based on the ratio of ionization chamber readings, differ for the investigated photon beams, up to 3.5% for the 4 MV x-ray beam. The deviations results from the difference in composition between the detector materials and water and can be taken into account by conversion of the ionization chamber readings for both the open and wedged photon beams to the absorbed dose to water. For the brass build-up cap detector system the ratio of the conversion factors for the wedged and open beam changes the ratio of the ionization chamber readings up to about 3.6% for the 4 MV x-ray beam. For the mini-phantom the conversion factors for the wedged and open beam are almost equal for all photon beams. Consequently, for that system wedge factors based on ionization chamber readings or dose values are the same. With respect to the wedge factor variation with field size a somewhat larger increase has been determined for the 60Co and 4 MV photon beam using the brass build-up cap: about 1% for field sizes varying between 5 cm x 5 cm and 15 cm x 15 cm. This effect has to be related to an apparent more pronounced variation of the head scatter dose contribution with field size for the wedged photon beams if the brass build-up cap detection system is used. It can be concluded that determination of wedge factors "in air" under reference irradiation conditions, performed with both the mini-phantom and brass build-up cap yields within 0.5% the same result if the wedge factors are based on a dose to water ratio. However, by using high-Z build-up materials the determination is more complicated because appropriate conversion factors are then required, while similar conversion factors can be ignored if more water equivalent build-up materials such as PMMA are applied.

Wedge factor constituents of high-energy photon beams: head and phantom scatter dose components.

The head and phantom scatter contribution to the output of a treatment machine have been determined for open and wedged 60Co gamma-ray beams and 4, 8, 16 and 25 MV X-ray beams, using an extended and a small-sized phantom. The wedge factor variation with field size and phantom depth have been analysed as a function of both scatter components. For the wedged beams a stronger increase of the head scatter contribution with field size, i.e. 4-9% for field sizes increasing from 5 cm x 5 cm to 20 cm x 20 cm, has been observed compared with open beams. This result indicates that the wedge factor variation with field size is related to a change of the primary photon fluence. Our study shows that the ratio of the head and phantom scatter contribution for the wedged and open beams remains unchanged for all beams except the 4 and 25 MV X-ray beam. This implies that, except for these latter energies, the variation of the wedge factor with phantom depth is determined by the wedge-induced change of the primary photon energy fluence. For the 4 and 25 MV X-ray beam it is shown that the wedge factor is also influenced by a change of the phantom scatter contribution. The wedge factor for the 25 MV X-ray beam is strongly influenced by the electron contamination for phantom depths up to 6 cm. For the 60Co and the 4 MV photon beam it is shown that the wedge factor decreases slightly with increasing source-to-skin distance due to a reduced contribution to the total dose from photons scattered in the wedge. For clinical use, an algorithm is given to calculate the wedge factor variation with field size and phantom depth.

Wedge factor constituents of high energy photon beams: field size and depth dependence.

Wedge factors have been determined as a function of field size and phantom depth for a 60Co gamma-ray beam and X-ray beams in the range from 4 MV to 25 MV. The results show an increase of the wedge factor with field size, up to 9.1% for the 25 MV X-ray beam. The magnitude of this increase is a linear function of the product of that part of the irradiated wedge volume that can be observed from the point of measurement, its mass energy-absorption coefficient and mass density. This relationship is independent of the photon beam energy, the type of wedge material and the wedge angle. Differences in variation of the amount of backscatter to the monitor with field size for the open and wedged photon beam yielded only a minor influence, up to 0.7%. For the 4-16 MV X-ray beams the wedge factor increases linearly with phantom depth, almost independently of field size. For the 60Co gamma-ray beam and the 25 MV X-ray beam the wedge factor variation is a more complicated function of phantom depth for a particular field size.

Quality assurance of the simultaneous boost technique for prostatic cancer: dosimetric aspects.

Quality assurance (QA) in radiotherapy is of particular importance if a new irradiation technique is introduced. The dosimetric aspects of such a QA program concern the check of the dose calculation procedure, i.e. the prediction of the relative dose distribution, as well as the verification of the absolute value of the target absorbed dose specified at a particular point. In our institution a QA program has been developed for a new conformal irradiation technique of prostatic cancer: the simultaneous boost technique. With this technique the dose of the boost field and the large field are given simultaneously, using customized 10 mm thick Roses-metal plates in which the boost field has been cut out. The computation of the dose distribution, using a procedure adapted from a commercially available 2D treatment planning system, has been compared with isodose distributions measured in a water phantom. Good agreement, better than 3% or 3 mm, was observed for both open and wedged 8 MV X-ray beams. In vivo dose measurements have been performed on individual patients to check the dose delivery at the specification point. An agreement better than +/- 2% with the calculated dose value was required. The average ratio for 18 patients of the actual and expected dose value amounted to 1.005 +/- 0.017 (1 SD) after a correction of the number of monitor units for 2 patients during the treatment. Quality control of the dose transmission factor of the Roses-metal plates has been performed.(ABSTRACT TRUNCATED AT 250 WORDS)

In vivo dosimetry during pelvic treatment.

High precision in vivo entrance and exit dose measurements have been performed with p-type diodes on patients during 8 MV X-ray irradiation of the pelvis, to investigate the accuracy of dose calculations in this region. Based on phantom measurements the accuracy of the p-type diode measuring system itself, i.e. the agreement with ionisation chamber dose measurements, was shown to be better than 0.7% while the reproducibility in the dose determination was 1.1%, 1.5% and 1.6% (1 S.D.) at the entrance point, isocentre and exit point, respectively, for the wedged lateral fields. Patient movement and the uncertainty in the diode position increased these values to 1.7%, 1.5% and 3.1% (1 S.D.) for dose determinations on patients. From the entrance and exit in vivo dose values the dose actually delivered to the isocentre was determined. For the anterior-posterior beams a good correspondence for most patients was observed at the entrance and exit point and at the isocentre between the in vivo and calculated dose values. For the wedged lateral beams a systematic deviation of about 3% was observed. In addition to the in vivo dose measurements phantom dose measurements have been performed to quantify the accuracy of the dose calculation algorithms including the computation of the number of monitor units. These measurements also served to quantify the effects of the actual patient on the dose delivery. The measurements showed that accurate calculation of the dose requires a separation of the head and phantom scatter contribution of the output of the treatment machine. The dependence of the wedge factor on field size, depth and source-skin-distance has also to be considered for accurate dose calculations. The effect of the patient on the dose calculation is mainly related to the actual electron densities of fat and bone structures compared to water: neglecting these densities in the dose computation could yield deviations up to 8.5% for the exit point in wedged beams. Based on these results, improvements in the dose calculation algorithms and monitor unit calculation including the use of the actual electron densities will be implemented in the treatment planning procedure.

Should inhomogeneity corrections be applied during treatment planning of tangential breast irradiation?

Due to the inclusion of lung tissue in the treatment volume, some parts of the breast will get a higher dose during tangential breast irradiation because of the lower lung density. Data on the accuracy of dose calculation algorithms, investigated by phantom measurements, determinations of the geometry and density of the actual lung in the patient and the results of in vivo dose measurements, are presented. From this information it can be concluded that a lung correction varying between about 3% and 7% is needed but its magnitude is slightly overpredicted in a number of commercial treatment planning systems. Because this increase in dose is already in a high dose region, it is recommended that inhomogeneity corrections should be applied during tangential breast irradiation.

In vivo dosimetry during tangential breast treatment.

The 3-dimensional (3-D) dose distribution as calculated in clinical practice for tangential breast treatment was verified by means of in vivo dosimetry. Clinical practice in our institution implies the use of 8 MV X-ray beams, a 2-D treatment planning system, collimator rotation and a limited set of patient data for dose calculations. By positioning diodes at the central beam axes as well as in the periphery of the breast the magnitude of the dose values at the isocentre and in points situated in the high-dose regions behind the lung could be assessed. The position of the diodes was verified by means of an on-line portal imaging device. The reproducibility of these in vivo dose measurements was better than 2% (1 SD). Our study showed that on the average the dose delivery at the isocentre is 2% less and at the points behind the lung, 5.7% higher with respect to the calculated dose values. Detailed analysis of these in vivo dosimetry results, based on dose measurements performed with a breast shaped phantom, yielded the magnitudes of the errors in the predicted dose due to several limitations in the dose calculation algorithms and dose calculation procedure. These limitations are each introducing an error of several percent but are compensating each other for the dose calculation at the isocentre. We concluded that the dose distribution in a patient for our treatment technique and dose calculation procedure can be predicted with a 2-D treatment planning system in an acceptable way. A more accurate prediction of the dose distribution can be performed but requires an estimation of the lack of scatter due to missing tissue, the change in the dose distribution due to oblique incident beams and the incorporation of the actual output of the treatment machine in the assessment of the number of monitor units.