Photon and Electron Beam in the Treatment of Cancer Patient Based on Monitor Unit Compilation

The aim of this research is to evaluate the precision of the Monitor unit (MU) calculation in radiation therapy to survive the cancerous patient, is the measurement of actual dose required to delivered. An essential key of quality assurance in radiation therapy is verifying the accuracy of monitor unit calculations. Difference between the simpler model calculation and other modifying method calculation assuming a flat water phantom must be required. The accurate measurement of field size is the most important fact since many dosimetric quantities were dependent on field size. Perhaps most commonly used method for determining monitor unit were modified which provide a new formalism.


Introduction
On the basis of clinical response data the international commission on radiation units(ICRU) and measurement states that dosimetry system must be capable of delivering dose to an accuracy of 5% or 7% [1][2]. More recently Mijnheer et al. [3,4] and Wambersie et al [3,5] proposed that the standard deviation of the uncertainty in the delivered dose should not be greater than 3.5%. For more it is need to improve on the treatment technology to get more accuracy. The delivery of therapeutic radiation is a medical procedure and as such requires independent confirmation to ensure correct and accurate delivery [6]. This confirmation is accomplished by a comprehensive calculation and chart review procedure performed both before and throughout patient treatment [6][7][8][9][10]. An evaluation of the incidence of radiotherapy errors over 10 years at a large regional cancer center concluded that treatment plan checks, including MU verification calculations, were very effective in detecting documentation and treatment planning errors [6,11]. Intensity modulated radiation therapy (IMRT) requires an enhanced quality assurance procedure. This applies in particular to the step of MU calculation verification. Because of time constraints, treatment planning systems (TPSs) normally deal only in an approximate manner with the physical processes of the interaction of ionizing radiation in the treatment head and dose deposition inside the patient. Therefore the determination of the absorbed dose needs experimental verification [12].
To get more accuracy many factor should be corrected which take an account for radiation treatment. Some of these are machine set up, patient adjustment exact tumor volume correction, field size determination, monitor unit calculation. All type fractions must included in dose calculation to get more accurate which depend on the monitor unit calculation. To ensure correct dose at any specific tumor cell it must need to know that how many monitor units is required under normalize condition for electron and photon. A small number of studies have already been published on the validity of this commercially available algorithm. The emphasis is generally on the verification of output factors and dose distributions in water [13,14]. The calculation of monitor units has developed over past several years as treatment planning for the improvement in accuracy but the clinical effect of impact parameter were unclear. The actual distribution of radiation dose accumulated in normal tissues over the complete course of radiation therapy is, in general, poorly quantified [15]. Calibrating the dose per monitor unit (DMU) for individual patients is important to deliver the prescribed dose in radiation therapy [16]. Historically, MUs were determined using a manual calculation process, where the calculations were based on water phantom data gathered at time of machine commissioning. There are a variety of radiation oncology TPSs, from widely used commercial systems to special purpose systems, with limited application to a specific delivery modality [17]. Currently, MU settings required to deliver the prescribed dose are often calculated by a computerized treatment planning system using methods and quantities different from those used in manual MU calculations. The verification calculation is not and should not be used as a check of the overall accuracy of the primary TPS; that is the function of commissioning and continual Quality assurance (QA). It is crucial that both the primary and the verification planning systems be properly and thoroughly commissioned so that they are as accurate as possible.
Monitoring the agreement between the TPS and the verification system during clinical use can aid in identifying regions where beam models or data may be improved, but such monitoring is not a substitute for the commissioning of either system. Both the TPS and the verification system should be fully tested and commissioned following accepted guidelines [6,18,19,20] prior to clinical use. In France, between 2004 and 2005, 23 patients received an overexposure of radiation resulting from the introduction of the enhanced dynamic wedge into the clinic, a review of the incident pointed to the recent elimination of an independent check of the MU calculation as a major contrib-uting factor [6,21].

Materials and Method
The experiment consists of several parts. We were used VARIAN CLINAC 2100 CD linear accelerator machine at Shaheed Ziaur Rahman Medical college & Hospital (SZMCH), Bogra which provide 6MV, 9MV, 12MV, 15MV photon and electron beam. A water phantom were used which is considered as a measurement body. Output factor (OPF), Percent depth dose (PDD), Tissue maximum ratio (TMR) were measured by using water phantom with electrometer (model Dose-1). The collected beam data were then used for beam modeling on the Pinnacle TPS (Treatment planning system). A set of cylindrical ionization chamber (farmer type FC-65P) were used to measure the radiation dose. The machine was calibrated to ensure that it works properly by comparing different measured data. To calculated monitor unit, we were used a simple model by correcting an impact factor. We were measured output factor in which included phantom scattering factor (S p ) and air scattering factor (S c ). At first we were calculated regular square field size for different irregular field size. We were used 10×10 reference field size. The water phantom were fed to the LINAC machine and measured the output for source axial distance (SAD) & source to surface distance (SSD) technique and also measured the output factor (OPF), Percent depth dose (PDD), Tissue maximum ratio(TMR), wedge factor and other attenuation factor for 6MV & 9MV photon and electron beam. For irregular field size, the output factor were obtained by interpolating neighbors field. The monitor unit was calculated for different field case and comparing with TPS (treatment planning system) calculation i.e. with manual calculation which was showed the better accuracy.

Calculation Formalism: For Photon Beam
Now-a-days the therapeutic machine has been improved so we need to improve our treatment technique. The equation for calculation of monitor unit (MU) in SSD technique that are used as before is The scattering factor S c , S p and ISF(inverse square factor) are included into dose rate and taken as a output factor(OPF) so the equation (1) is reduce to The equation for calculation of monitor unit in SAD technique For Electron Beam: The equation for calculation of monitor unit is

Calculation of Monitor Unit
(1

=207.68≈207
Eample-2: Calculate the MUs required to deliver 300cGy to a depth of d m 100cm SSD for a 9.4×8.6 cm 2 insert in a 10×10cm 2 applicator using 6MeV energy beam?
Answer: For the standard 100 -cm SSD, the monitor unit can be obtained buy using equation (3). Using data from table and square root rule for the output factor, the equivalent square field size is 9×9 cm 2 , the monitor units are given by

Discussions
From above result it was shown that every term in equation (1) & (2) dependent on field size so that it should be correctly measured of field size. Table-1 showed the correct measurement of several square field size from different rectangle size which provides the actual calculation of monitor unit. Table-2&3 were the measurement of output factor of different field size for 6&15MV photon beam. Table-4&5 are measurement of  PDD at different depth for different field similarly Table-6&7 were the measurement of TMR for 6&15 photon beam. Figure-1 illustrated that the variation of output factor depending on field size and it was showed that output factor increased exponentially with increasing field size. Figure-2&3 illustrated variation of PDD depending on depth and showed that maximum surface dose occurred for high energy photon beam.           Table-9 is the measurement of output factor for different field size and electron beam.    Figure: 4illustrated that as the energy of the electron beam increased the maximum dose occurred at higher depth. From above we observed that many corrections were included to get more accuracy on the monitor unit calculation. From case_1(cervix), the required monitor unit was 56 but according to TPS(treatment planning system) calculation it was 53 which was lesser than required, suppose we may assume the machine calculation is greater or less than 3 from the our calculation which means that per 100cGy the access or less monitor unit is 6, To delivered 5000cGy, it will be 300cGy so that the patient get more or less 300cGy from required which take a violent effect and the normal tissue around the tumor may be get permanently damaged if it is excess, it may not possible to damage the tumor cell if it is lesser than accurate. The result will more accurate if the ionization chamber perfect work due to use old ionization chamber.

Conclusion
Computer calculations are commonly verified using an independent manual procedure. It is difficult to calculate treatment delivery monitor units for this variant of IMRT using manual method, since manual calculations are not feasible, it is important both to understand and to verify the calculation of treatment monitor units by the planning system algorithm. A formal analysis was made of the dose calculation model and the monitor unit calculation embedded in the algorithm. Experimental verification of the dose delivered by plans computed with methodology demonstrated an agreement of better than 4% between the dose model and measurement. So it must be required to take corrected form for the calculation of monitor unit to survive the cancerous patient.

Notation and Definition.
D=The absorbed does at the point of interest from the individual field under calculation.
OP=Output or The dose rate or does per monitor unit at the point of interest.
d=Depth of the point of calculation. 1 =Water-equivalent depth of the point of calculation. 1 " =The normalization depth for photon and electron dosimetry. For photon, d o = 10cm is recommended, but not required. For each photon beam, d o is independent of field size and shall be greater than or equal to the maximum d m .
For electrons, d o is taken to be the depth of maximum dose along the central axis for the same field incident on a water phantom at the same SSD. It is field-size de-pendent. D m =The depth of maximum dose on the central axis. OAR (d,x)= Off-axis ratio (sec. 1.A.1.f). The ratio of the open field dose rate at an off-axis point to that for the same field (e.g, 10*10 cm 2 ) shifted such that the point of calculation lies on the central axis. The Primary Off-Axis Ratio, POAR, is preferred to be used for OAR (d,x). PDD (d,r,SSD)=Percent depth dose. The ratio, expressed as a percentage, of the dose rate at depth to the dose rate at d m in a water phantom for a given field size and SSD.
PDD (d,r,SSD)= Normalized percent depth dose (sec.1. A.1.b). The ratio, expressed as a percentage, of the dose rate at depth to the dose rate at the normalization depth in a water phantom for a given field for a given field size and SSD.
SAD= Source-axis distance. Distance between the x-ray physical source position and the isocenter. For most linear accelerators, this value is nominally 100cm.
SPD= Source-point distance. The distance from the x-ray physical source to the plane (perpendicular to the central axis) that contains the point of calculation.
SSD=source-surface distance. The distance along the central axis from the physical source to the patient/phantom surface.
SSD o =Standard source-surface distance. The distance along the central axis from the physical source to the patient/phantom surface under normalization conditions. SSD eff =Effective source surface distance. The distance along the central axis from the effective source to the patient /phantom surface, determined by beat fit of output versus the inverse of the distance squared.
TPR (d, r d )=Tissue phantom ratio (sec. 1.A.1.c). the ratio of the dose rate at a given depth in phantom to the dose rate at the normalization depth for a given field size.
TF=Try factor. The ratio of the central-axis dose rate for a given field with and without a blocking tray. TF is assumed independent of depth and field size in this report. This factor may be used to account for the attenuation through additional materials (e.g, special patient support devices) as needed. WF (d,r d ,x)=wedge factor (sec.1.A.1.i). the ratio of the dose rate at the point of calculation for a wedged field to that for the same field without a wedge modifier. The wedge may be a physical filter or not (i.e, dynamic or virtual). Depending on the type and angle of the WF may depend on the wedge angle, field size, depth, and off-axis distance.
R max =The maximum range (cm) is define as the depth at which maximum electron absorbed. R p =It is the practical range define as the depth at which the tangent plotted through the steepest section of the electron depth dose curve intersects with the extrapolation line of the background.