THE UNIVERSITY OF OKLAHOMA HEALTH SCIENCES CENTER GRADUATE COLLEGE RADIATION DOSE ESTIMATION FOR DIAGNOSTIC MODALITIES

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1 THE UNIVERSITY OF OKLAHOMA HEALTH SCIENCES CENTER GRADUATE COLLEGE RADIATION DOSE ESTIMATION FOR DIAGNOSTIC MODALITIES A THESIS SUBMITTED TO THE GRADUATE FACULTY in partial fulfillment of the requirements for the degree of MASTER OF SCIENCE BY VICTOR LUIS GARCIA Oklahoma City, Oklahoma

2 RADIATION DOSE ESTIMATION FOR DIAGNOSTIC MODALITIES APPROVED BY: Weiyuan Wang, Ph.D., DABR, Chair Sulaiman D. Aldoohan, Ph.D., DABR Salahuddin Ahmad, Ph.D., DABR, FACMP,FAAPM, FACR THESIS COMMITTEE 2

3 COPYRIGHT by Victor Luis Garcia May 11,

4 ACKNOWLEDGEMENTS I would like to thank Dr. Wang for being my mentor throughout the duration of my Master s degree, whose guidance and motivation has helped me with the completion of this thesis. I would also like to express gratitude for the other members of my thesis committee, Dr. Aldoohan and Dr. Ahmad. Last, but not least, I would like to thank my mom, dad, and family, who have shown encouragement and support to me throughout the duration of this degree. 4

5 ABSTRACT The medical use of radiation in imaging has proven to be incredibly beneficial for patient diagnosis, but it does not come without the risk to the patient. In the field of medical physics, the risk can be quantified by the radiation dose the patient receives. Discussed in this thesis is the estimation of dose for the imaging modalities of radiography, fluoroscopy & interventional radiography, computed tomography, and mammography. For radiography, the skin dose and depth dose are discussed, as well as methods for calculating the depth dose (percent depth dose and tissue-air ratio methods). For fluoroscopy, the skin dose and depth dose are also discussed, in addition to the variation of the automatic brightness control and its effect on the exposure rate. In interventional radiology, the peak skin dose and the methods for calculating the peak skin dose are demonstrated. For computed tomography, methods for effective dose, size-specific dose estimate (SSDE), and fetal dose are discussed and estimated by the program. Lastly, for mammography, the calculation of average glandular dose (AGD) is discussed, as well as the variation of the AGD with breast thickness. For the radiography and fluoroscopy dose estimation, the skin dose and depth dose, both by PDD and TAR methods, are estimated for various patient parameters and imaging techniques. For interventional radiology, the peak skin dose and depth dose are estimated for different parameters. For computed tomography, the effective dose, SSDE, and fetal dose are estimated. For mammography, the AGD is estimated for various techniques and breast parameters. Caution should be made to use data from this thesis to estimate a specific patient dose. 5

6 TABLE OF CONTENTS CHAPTER I... 9 CHAPTER II...12 CHAPTER III...21 CHAPTER IV...38 CHAPTER V...45 CHAPTER VI...52 BIBLIOGRAPHY...54 PRESENTATIONS...57 APPENDIX...59 TABLE OF TABLES & FIGURES Table 1. Exposure and One-Shot HVL values measured at 40 in. source-to-chamber distance for Philips Radiographic Room OUCPB Table 2. PDD Interpolation Table for kvp and HVL Table 3. TAR Interpolation Table for HVL Table 4. Interface of the program for calculating Radiographic Dose Table 5. Technique differences for different exam types of the Philips BV Pulsera Table 6. Equivalent acrylic thickness for the Nuclear Associates Phantoms Table 7. Interface of the program for the calculation of fluoroscopic dose Table 8. Program interface for IR dose estimation Table 9. Program interface for the CT dose estimation Table 10. Percent differences between the displayed AGD and measured AGD for different units Table 11. Program interface for the mammography dose estimation Table A- 1. Exposure at 40 in. and HVL data as a function of kvp for OUCPB Table A- 2. Exposure and HVL data as a function of kvp for OUCPB Table A- 3. Exposure and HVL data as a function of kvp for OUCPB Table A- 4. Exposure and HVL data as a function of kvp for OUCPB Table A- 5. Exposure and HVL data as a function of kvp for the OUPB GE Proteus Room Table A- 6. Exposure and HVL data as a function of kvp for the OUPB GE Proteus Room Table A- 7. Tube information for the radiographic rooms mentioned in Table A-1 through A Table A- 8. Technique and exposure rate variations for the GE OEC 9900 for HLF at normal magnification Table A- 9. Variation of technique and exposure rate for the GE OEC 9900 Elite for HLF at Magnification Table A- 10. Variation of technique and exposure rate for GE OEC 9900 Elite for HLF on Magnification

7 Table A- 11. Variation of technique and exposure rate for GE OEC 9900 Elite for Standard Fluoroscopy at Normal Magnification Table A- 12. Variation of technique and exposure rate for Standard Fluoroscopy at Magnification Table A- 13. Variation of technique and exposure rate for GE OEC 9900 for Standard Fluoroscopy at Magnification Table A- 14. Variation of technique and exposure rate for the BV Pulsera for HLF with 31 cm II size Table A- 15. Variation of technique and exposure rate for the Philips BV Pulsera for HLF with 23 cm II size Table A- 16. Variation of technique and exposure rate for the PHilips BV Pulsera for HLF with 17 cm II size Table A- 17. Variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 31 cm II size Table A- 18. Variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 23 cm II size Table A- 19. Variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 17 cm II size Table A- 20. X-Ray tube information for the GE OEC 9900 Elite and Philips BV Pulsera C-Arms Table A- 21. X-Ray Tube information for CTs on campus Figure 1. Geometric illustration of the field size at the image receptor and the patient surface..15 Figure 2. Geometric illustration of the field size at the depth, d, and the image receptor Figure 3. Image of setup for measuring the exposure rate Figure 4. Image of the console display with image of the ion chamber Figure 5. Nuclear Associates Abdominal Phantom Figure 6. Nuclear Associates Chest Phantom Figure 7. Nuclear Associates Lateral Skull Phantom Figure 8. Nuclear Associates Extremity Phantom Figure 9. Nuclear Associates Lumbar Spine Phantom Figure 10. ABC Curve for the GE OEC 9900 for Standard Fluoroscopy...27 Figure 11. ABC Curve for the GE OEC 9900 for HLF...27 Figure 12. ABC Curve for Standard for the Philips BV Pulsera...28 Figure 13. ABC Curve for HLF for the Philips BV Pulsera...28 Figure 14. Acrylic thickness vs. Exposure rate for Standard mode for the GE OEC Figure 15. Acrylic thickness vs. Exposure rate for HLF for the GE OEC Figure 16. Acrylic thickness vs. Exposure rate for Standard Mode for Philips BV Pulsera...30 Figure 17. Acrylic thickness vs. Exposure rate for HLF for Philips BV Pulsera...30 Figure 18. Copper thickness and equivalent acrylic thickness for the same exposure rate at 30 cm from II Figure 19. Lateral patient setup for IR Figure 20. Frontal patient setup for IR Figure 26. Picture of the 16 cm CTDI Phantom Figure 27. Illustration showing the effective diameter versus the patient dimensions

8 Figure 29. Setup of the Rad Cal 10X6-6M ion chamber and the ACR Mammography Accreditation Phantom Figure 30. BR-12 Phantom with 8 cm thickness on a Hologic Selenia unit Figure 31. ACR Mammography Accreditation Phantom Figure 32. Phantom thickness vs. displayed AGD for 2D and 3D modes Figure 33. Phantom thickness vs. displayed AGD for 2D only system Figure 34. Phantom thickness vs. displayed AGD for SBB Figure 29. AEC Compensation Steps vs AGD for 4 cm BR-12 thickness

9 CHAPTER I INTRODUCTION Wilhelm Roentgen revolutionized the medical world in 1895 with his discovery of the X-ray [1]. Since then, X-rays have produced wonders for the advancement of medical diagnosis. X-rays allow the user to see what cannot be seen with the naked eye or through invasive methods. While this is an incredible benefit to the field of medicine, it does not come without a price. Due to the nature of ionizing radiation, its interactions with matter, including the human body, can lead to potential harm of the patient. This harm is indirectly measured in the form of radiation dose, which is analogous to the pharmaceutical dose. Effects of excessive dose to the patient can result in two types of harm. The first type is stochastic effects, which, for a certain amount of dose, give a probability for the effect to occur. There is no threshold dose for which no stochastic effects are seen. For diagnostic energies, the most important stochastic effect is the incidence of cancer. A rule of thumb for low doses, the percentage of a population of people developing a solid tumor from irradiation, for a given effective dose, is about 5% per Sv (Sv, or Sievert is a unit of effective dose) [2]. The second is deterministic effects, which have a no occurrence until a threshold dose is met. In diagnostic imaging, the most common deterministic effects are erythema (reddening of the skin), epilation (loss of hair), and radiation dermatitis (radiation burns) in the irradiated areas. The estimation of the radiation dose the patient receives for diagnostic imaging aids in the determination of deterministic or stochastic effects. While this dose is precisely calculated in the field of radiation therapy, in diagnostic imaging, these patient dose calculations are not as easily found. Therefore, the development of an easy, quick, and automatic way to estimate the radiation dose the patient receives is necessary for diagnostic imaging. 9

10 When imaging the patient, the techniques used are important in determining the quality of the image produced. These parameters also help determine the dose the patient receives. This plays into the delicate balance of image quality and dose to the patient. There are important quantities, which may be mentioned multiple times throughout this thesis. The first parameter is kv or kvp, which refers to the electric potential placed across the X-ray tube. kvp is more specifically peak kilovoltage across the X-ray tube, which gives the maximum energy an X-ray photon can have, although the energies of photons produced is a spectrum. The assumption, unless measured exactly, is the dose is approximately proportional to the square of kv or kvp. The next quantity is ma, or the tube current of electrons that flow through from the cathode to the anode of the X-ray tube. This directly influences the rate of photons produced from the anode, which in turn directly influences the dose rate. The assumption is that the dose rate the patient receives is linearly proportional to the ma used in the imaging technique. Another quantity that may be used in place of ma is mas, which is the tube current multiplied by the time of exposure. Lastly, is the half-value layer, or HVL. HVL is defined as the amount of material, typically measured in terms of millimeter Aluminum equivalent (mm Al eq.), needed to reduce the beam intensity by half, and is used in diagnostic imaging physics as the description of beam quality. HVL indirectly determines the energy of the beam, through a quantity called effective energy, and increases the dose to the patient with decreasing HVL, due to the decrease in penetrating power of lower energy photons, which only give the patient dose with no contribution to the image. For the estimation of depth dose, Wagner et. al [3] presents methods for calculating the depth dose. In Wagner et. al [3], data for various papers are quoted [4-5] in the book used for the calculation of the dose. However, Wagner et. al was written in 1997 with the data written before However, source [6-8] have provided direct or indirect comparisons of the data presented in Wagner et. al [3]. Through comparison, there is minor deviations between the data 10

11 in [6-8] and Wagner et. al [3]. Therefore, the methods outlined in Wagner et. al [3] will be used for calculation of depth doses. 11

12 CHAPTER II RADIOGRAPHY DOSE Radiography is perhaps the oldest technique for imaging a patient through the use of X-rays. There are two types of X-ray units for imaging using radiography: radiographic/radiographic and fluoroscopy rooms (RF rooms) and portable X-ray units. When using radiography, the main concern is the occurrence of stochastic effects, mainly due to the fact the skin dose in radiography is not large enough to pass the threshold doses of the deterministic effects. Using Wagner et. al [3], there are two methods for calculating the radiation dose at a depth. The first is using the Percentage Depth-Dose (PDD) Method. The PDD method finds the relative percentage dose PDD, for a certain depth d, in the patient, based on the field size at the patient surface, the kvp, and HVLs used in the examination. Other factors to consider is the exposure A, which is defined as the charge created in air per unit mass from the ionizing radiation (in this case, X-rays), mas, and the patient thickness. Assumptions made by Harrsion [4] are: measurements are made along the central axis of the beam, measurements are made in a water tank, the distance from X-ray source to surface of water tank is 60 cm, the waveform of the X-ray tube is not applicable, and measurements are probably accurate to within +/- 10 %. Demonstrated below are the equations for calculating the dose, based off of the PDD Method: 2 D(rad or mgy t ) = Q(rad or mgy t ) PDD ( SSD(cm) d(cm) (1 + )) [Eq.1] 100 SSD+d(cm) 60 cm Q(rad or mgy t ) = A SSD (R or mgy a ) f( rad or mgy t ) B [Eq. 2] R mgy a In Eq. 1, D is defined as the Dose at depth d, Q is the entrance dose at the patient surface, PDD is the relative percentage dose at the depth d, SSD is the source-to-surface distance (or 12

13 distance to the patient surface from the X-ray focal spot). In Eq. 2, Q is, as previously mentioned, the entrance dose at the surface of the patient, A SSD is the free-in-air exposure at the patient surface, which neglects any scatter from the patient, f is the conversion factor from exposure or air kerma to dose in tissue (quoted at f = 0.93 rad/r, if using exposure, or f = 1.06 mgy t/mgy a, if using air kerma), and B is defined as the backscatter factor. For a depth of 0 cm, the percentage depth dose, P, is equal to 100. The PDD is dependent upon the HVL, kvp, depth, and field sizes. Tables for the PDD values can be found in Wagner et al. [3]. The PDD increases as a function of HVL, due to a more penetrating beam, it increases with kvp for the same reason, PDD decreases as a function of depth, and the PDD increases for larger fields, due to more scattered radiation contribution. The second method to determining dose at a depth is through the Tissue-Air Ratio (TAR) Method. The TAR method finds the ratio to convert the free-in-air exposure or air kerma at a certain point in space in the patient body is located to the dose at depth d, in the body of the patient. TAR has a dependence on the depth d, of the point of interest, the field size at depth d, and the HVL. Other factors to consider are the free in air exposure, A, at depth d, the kvp, and the mas, which both affect the exposure. Assumptions made by Säbel et. al [5] are: all measurements are made along the central axis of the beam, all measurements were made in a water tank, distance from X-ray source to point of measurement was 75 cm, and simulated three-phase, six pulse waveform were used for data. Dose at depth using the TAR method is shown as: D(rad or mgy t ) = TAR( rad or mgy t ) A R mgy SSD+d (R or mgy a ) [Eq. 3] a In Eq. 3, D is defined at dose at depth d, TAR is the tissue-air ratio, and A SSD+d is the free-in-air exposure or the air kerma. Values for TAR can be found in Wagner et. al [3] for different kvp, 13

14 field sizes, and depths. The TAR increases as a function of HVL, decreases as a function of depth, and increases with field size. In order to calculate the dose, the first quantity needed is the exposure or air kerma, kinetic energy released per unit mass. To do this, a calibrated RadCal 10X5-6 ion chamber was placed at 40 inches from the source of the X-ray tube. It should be noted that the ion chamber was free-in-air, with no potential source for scatter to add to the exposure. The technique for the X-ray tube was set at 100 mas and varying kvp, and an exposure was taken. The mas was kept constant, due to the linear relationship between dose and mas. Once each X-ray was taken, the exposure measured was then divided by the mas. The kvps chose ranged from 50 kv to 150 kv, depending on the specific unit. Since the PDD method requires the free-in-air exposure at the patient surface, and the TAR method requires the free-in-air exposure at the location of the point of interest, the exposure measured at 40 inches is used with the inverse square law to find out the exposures at either of these distances, as demonstrated in Eq. 4 and Eq in A SSD (R or mgy a ) = A 40 (R or mgy a ) ( SSD(in) )2 [Eq. 4] 40 in A SSD+d (R or mgy a ) = A 40 (R or mgy a ) ( SSD+d(in) )2 [Eq. 5] Once the exposures were collected, the HVL is measured for each of the kvps. This was done using a calibrated RadCal AGMS-DM+ Solid-State Detector. The AGMS-DM+ Solid-State Detector was used because of its feature, which takes an One-Shot HVL. This method significantly saves time when taking measurements, compared to the traditional method of taking exposures and filtering the X-ray beam with Aluminum filters until the intensity of the beam is reduced to one half of its initial value. However, there is some error from the One-Shot value and the value from the traditional method. The difference between the One-Shot value and the traditional method is <4% [9]. HVL was measured for varying kvp values and 100 mas. 14

15 The HVL measurement is necessary to find the PDD, the Backscatter Factor, and the TAR. Table 1 shows the exposures and HVL values measured across varying kvps for a Philips Radiographic Room. Table 1. Exposure and One-Shot HVL values measured at 40 in. source-to-chamber distance for Philips Radiographic Room OUCPB kvp Exposure mr/mas HVL (mm Al eq.) (mr) For calculation of PDD and backscatter factor, the field size must be determined at the patient surface. Since the field size is known at the image receptor, the field size at the patient surface can be found based off of the field size at the image receptor. The field size at the patient surface can be simply calculated using a simple geometric relationship. Figure 1 shows an illustration of the geometric relationship between the field size at the image receptor and the field size at the patient surface. Figure 1. Geometric illustration of the field size at the image receptor and the patient surface. 15

16 The calculation of field size at the patient surface is simply just a proportional to the ratio of the SSD to the SID, as demonstrated in Eq. 6. x SSD (cm) = x SID (cm) SSD(cm) SID(cm) [Eq. 6] For the TAR method, the field size must be found at the patient depth. Similar to the surface field size calculation, the field size at the patient depth d, is simply proportional to the two distances. Figure 2 shows an illustration for the field size at depth in the patient. Figure 2. Geometric illustration of the field size at the depth d, and the image receptor. The field size at the depth d, is calculated using the proportion between SSD+d and SID, as demonstrated in Eq. 7. x SSD+d (cm) = x SID (cm) SSD+d(cm) SID (cm) [Eq. 7] Once the field size and HVL are determined, the backscatter factor can be determined when using the PDD method. The backscatter factor is a function of the field size and HVL from Wagner et. al [3]. Note, the backscatter factors are only listed for HVLs from 1.0 to 4.0 mm Al and for 10x10 cm 2 to 30x30 cm 2 field sizes. For other field sizes and HVL values, the data was extrapolated from the data found from the source. While this extrapolation is being used, these backscatter factors can be measured in the future to correct for potential error from extrapolation with an ion chamber and an acrylic phantom. 16

17 Similar to the backscatter factors, the PDD values are limited to certain values of kvp, from 60 kv to 100 kv, and HVL values, from 1 mm Al eq. to 4 mm Al eq., in Wagner et. al [3]. These kvp and HVL values don t encompass every possible value that could be potentially used in the clinic. Therefore, to get the PDD values for these missing values, this new data must be extrapolated from the data present. Table 2 shows a screenshot of the Dose Estimation Program for PDD interpolation. First, the interpolation is done between depth and field size to display the table presented in Table 2. The highlighted data shows these values. Once the interpolation is done between depth and field size, the data is then extrapolated from the highlighted data to get the PDD values for kvp and HVL that are missing. After this data is calculated, then an interpolation between this new data for kvp and HVL is done to get the final PDD that will be used to calculate dose using the PDD method. Table 2. PDD Interpolation Table for kvp and HVL. Similar to the PPD interpolation, the TAR method has a similar table generated to use for interpolation of the TAR value. First, the TAR is found for each depth to generate the table. Once the table is generated, TAR is interpolated between HVL to find the TAR value to calculate the dose. The generated TAR table is shown in Table 3. Note, the values in Wagner et. al [3] are limited to certain HVL values, so the values highlighted shown these TARs, and the other values are extrapolated from this data. 17

18 Table 3. TAR Interpolation Table for HVL. The TAR values shown are in terms of rad/r. When using air kerma, kinetic energy released per unit mass, in the TAR method, the following relationship is used to find the conversion from air kerma to dose in tissue: TAR ( mgy t ) = 1.14 TAR ( rad ) [Eq. 8] mgy a R The program is able to calculate dose based on various parameters, such as the HVL, kvp, mas, patient thickness, depth, and field size. Table 4 shows the program interface for the Radiographic Dose Estimation. In the section, labeled Radiographic Unit, there is the option to select a specific radiographic unit for dose calculation. Note, all the parameters that must be input are denoted in blue, and the dose that is calculated is shown in yellow. The dose calculated using the PDD method and the dose calculated using the TAR method are very similar to each other. 18

19 Table 4. Interface of the program for calculating Radiographic Dose. While the program provides a quick and automatic way to calculate dose, there are some limitations to the program. One of the assumptions made is that the point of interest is on the central axis of the beam. This means there is no off-axis correction or any out-of-field calculation for the dose. The next limitation stems from the data for PDD, backscatter factor, and TAR. Due to the limited values across HVL, kvp, and field sizes, the extrapolated data may cause additional error from the extrapolation of the data. Lastly, the program does not account for any spectral changes for the X-ray spectrum in more modern tubes since the publication of the data in Wagner et. al [3]. For future consideration, the removal or reduction of error in these 19

20 limitations can be accomplished by measurement of the missing values. These measurements can be made physically or by computational methods. 20

21 CHAPTER III FLUOROSCOPY DOSE Fluoroscopy is an imaging modality which allows the user to see the live image of the patient. This modality can be divided into three main subsections for imaging: C-Arms (portable units), fluoroscopy rooms, and interventional radiology (IR). Prolonged fluoroscopy can lead to high doses, which can lead to occurrences of the deterministic effects of radiation, while also increasing the risk for the occurrence of the stochastic effects. When calculating dose, however, the dose for the C-Arms and fluoroscopy rooms use the same estimation program with the same parameters, and the dose for the IR cases uses a different estimation program with similar parameters. When calculating skin dose and depth dose for the C-Arms and fluoroscopy rooms, similar equations and parameters are used as in the case of radiography dose. Due to this, the same assumptions do apply to both the PDD and TAR methods when calculating the depth dose. An additional assumption for the C-Arms is that there is no rotational movement of the C-Arm itself. To calculate dose in this case, the following technique and patient parameters are needed: kvp, ma, duration of exposure, field size at the image intensifier, patient thickness, depth, exposure rate, and the HVL. The calculation of skin dose for C-Arms and fluoroscopy is given by Eq. 9: Q(rad or mgy t ) = A SSD( R or mgy a min min ) t(min) f(rad R or mgy t mgy a ) B [Eq. 9] In Eq. 9, Q is the surface dose, A SSD is the exposure rate at the patient surface, t is the duration of exposure, f is the conversion factor from exposure or air kerma to dose in tissue, and B is the backscatter factor. To calculate the depth dose using the PDD method, this new definition of surface dose is used in Eq. 1, rather than the surface dose as defined in Eq

22 Similarly, to calculate depth dose using the TAR method, the definition must use the exposure rate and duration of exposure. The dose using the TAR method for the C-Arm and fluoroscopy rooms can be found with Eq. 10: D(rad or mgy t ) = TAR( rad R or mgy t mgy a ) A SSD+d( R or mgy a min min ) t (min) [Eq. 10] In Eq. 10, D is the dose in tissue, TAR is the tissue-air ratio, A SSD+d is the exposure rate at the depth of interest, and t is the duration of exposure. The dose is determined indirectly by the Automatic Brightness Control (ABC), which chooses the kvp and ma to maintain image quality. These parameters will change based on the patient thickness. It is important to check the variation of the fluoroscopy technique based of the patient thickness and find the change in exposure rate to the patient. By comparing the acrylic thickness to the patient thickness, an approximation of the technique and exposure rate can be determined for dose estimation. Using GE OEC 9900 and Philips BV Pulsera C-Arms, a calibrated RadCal 10X5-6 ionchamber was placed at 30 cm from the image intensifier (II). An Acrylic phantom was placed on the image intensifier, ranging from thicknesses of 1 inch to 12 inches, as demonstrated in Figure 3. For each thickness of Acrylic, the kvp and ma were chosen by the ABC and the exposure rate were measured by the ion-chamber. The kvp, ma, and exposure rate were recorded for various types of fluoroscopy examinations (e.g. Thorax, Abdominal, etc.), Standard and High- Level fluoroscopy (HLF), and for the various magnification modes. The technique is shown on the image on the console display screen as shown in Figure 4. It should be noted that the ion chamber is also centered in the image in Figure 4. The techniques were compared with those using Copper and various Nuclear-Associates phantoms that are used in the clinic (Figures 5-9). 22

23 X-Ray Tube Image Intensifier (II) Figure 3. Image of setup for measuring the exposure rate. Figure 4. Image of the console display with image of the ion chamber. 23

24 The abdominal phantom is constructed by 8 in. of acrylic (Figure 5). The chest phantom is constructed using 4 in. of acrylic, a 1 mm thick Aluminum plate, and a 2 mm Aluminum plate, in the order of 1 in. acyrlic+2 mm Al+1 in. acrylic+2 in. Air+1 in. acrylic+1 mm Al+1 in. acrylic (Figure 6). The lateral skull phantom is constructed with 6 in. of acrylic, a 1 mm thick Aluminum plate, and a 2 mm Al plate, in the order of 1 in. acyrlic+2 mm Al+4 in. acrylic+1 mm Al+1 in. acrylic (Figure 7). The Extremity phantom is constructed using 2 in. of acrylic and the 2 mm Aluminum plate, in the order of 1 in. acrylic+2 mm Al+1 in. acrylic (Figure 8). The Lumbar Spine phantom is similar to the Abdominal phantom. It is constructed using 8 in. of acrylic with a narrow 5 mm thick Aluminum plate (Figure 9). Figure 5. Nuclear Associates Abdominal Phantom. 24

25 Figure 6. Nuclear Associates Chest Phantom. Figure 7. Nuclear Associates Lateral Skull Phantom. 25

26 Figure 8. Nuclear Associates Extremity Phantom. Figure 9. Nuclear Associates Lumbar Spine Phantom. The ABC for the GE OEC 9900 adjusts the kvp and ma in a linear fashion until the ma saturates, at 6.00 ma for Standard and at ma for HLF, and only the kvp increases, as the thickness of the Acrylic increases. As the magnification increased, the kvp and ma increased at a much faster rate than the lower magnifications for the GE OEC Figure 10 and Figure 11 show the ABC curves for the 9900 for Standard and HLF, respectively. 26

27 ma ma 8 6 ABC for Varying Acrylic Thickness for Standard Fluoroscopy Normal Mag 1 Mag kvp Figure 10. ABC Curve for the GE OEC 9900 for Standard Fluoroscopy. 20 ABC for Varying Acrylic Thickness for High Level Fluoroscopy Normal Mag 1 Mag kvp Figure 11. ABC Curve for the GE OEC 9900 for HLF. The ABC curves for the Philips BV Pulsera follow the same curve across magnification modes. For the same acrylic thickness, for higher magnification modes, the points were further along the curve. For the Standard (Figure 12) versus HLF (Figure 13), the HLF mode uses a higher ma for the same acrylic thickness. Note, for the Pulsera, only up to 8 in. of acrylic were used. 27

28 ma ma ABC Curve for Standard Fluoroscopy 31 cm Mag 23 cm Mag 17 cm Mag kvp Figure 12. ABC Curve for Standard for the Philips BV Pulsera. ABC Curve for HLF 31 cm Mag 23 cm Mag 17 cm Mag kvp Figure 13. ABC Curve for HLF for the Philips BV Pulsera. For both the GE OEC 9900 (Figure 14-15) and the Philips BV Pulsera (Figure 16-17), the exposure rate increased exponentially with the increased thickness of acrylic for both Standard and HLF modes. For increasing magnification, the exposure rate was higher for the same acrylic thickness. 28

29 Exposure Rate (R/min) Exposure Rate (R/min) 10 Acrylic Thickness vs. Exposure Rate Standard Normal Mag 1 Mag Acrylic Thickness (in) Figure 14. Acrylic thickness vs. Exposure rate for Standard mode for the GE OEC Acrylic Thickness vs. Exposure Rate HLF Normal Mag 1 Mag Acrylic Thickness (in) Figure 15. Acrylic thickness vs. Exposure rate for HLF for the GE OEC

30 Exposure Rate (R/min) Exposure Rate 10 Acrylic Thickness vs. Exposure Rate Standard 31 cm Mag 23 cm Mag 17 cm Mag Acrylic Thickness (in) Figure 16. Acrylic thickness vs. Exposure rate for Standard Mode for Philips BV Pulsera. 10 Acrylic Thickness vs. Exposure Rate for HLF 31 cm Mag 23 cm Mag 17 cm Mag Acrylic Thickness (in) Figure 17. Acrylic thickness vs. Exposure rate for HLF for Philips BV Pulsera. When comparing the different examination types, the abdominal phantom was used. On the Philips BV Pulsera C-Arm, the four exam types were Abdominal, Orthopaedics, HQ Orthopaedics, and Head/Spine. The 31 cm II setting size was used for the different exam types 30

31 with 8 in. of acrylic. The variation between exam types is shown in Table 5. The exposure rate was measured at 30 cm from the surface of the II. Table 5. Technique differences for different exam types of the Philips BV Pulsera. Exam Type Mode kvp ma Exposure Rate (R/min) Abdominal Standard Orthopaedics Standard HQ Orthopaedics Standard Head/Spine Standard As seen in the different exam types, the fluoroscopy technique changes with the exam type. These variations in the technique should be investigated for a better application of the program. The phantoms mentioned above (Figures 5-9) were compared to thicknesses of acrylic. To calculate the equivalent acrylic thickness, the exposure rate measured at 30 cm from the II was compared to the exposure rates from the varying thicknesses. The equivalent thickness is defined as the acrylic thickness that gives the same exposure rate as the phantom used. Table 6 shows the equivalent acrylic thicknesses for the various phantoms, as well as the exposure rate measured. Table 6. Equivalent acrylic thickness for the Nuclear Associates Phantoms. Phantom Phantom Materials Calculated Equivalent Acyrlic Thickness (in) Exposure Rate (R/min) Extremity 1"Acrylic+2 mm Al+1"Acrylic Chest Lateral Skull 1"Acyrlic+2mm Al+1"Acrylic+2" Air+1"Acrylic+1mm Al+1"Acrylic 1"Acrylic+2mm Al+4"Acrylic+1 mm Al+1"Acrylic Abdominal 8"Acrylic Lumbar Spine 8"Acrylic+5mm Narrow Al Plate

32 Acrylic Thickness (inch) As noted above the abdominal phantom is made with 8 in. of acrylic. When comparing the abdominal phantom exposure rate to the varying thickness for 8 in. of acrylic exposure rate, the results are very similar when concerning the equivalent acrylic thickness. For the different phantoms, the equivalent acrylic thickness was found to be higher than the amount of acrylic present in the phantom itself. This was to expected due to the aluminum plates in the phantom. The last comparison was made between copper and acrylic. The copper was placed on the II with ion chamber still at 30 cm from the II. When doing image quality assessments, 2 mm of copper filtration are used to simulate an adult abdomen, and 1 mm of copper filtration is used to simulate an extremity. The thickness of copper was compared to the thickness of acrylic resulting in the same exposure rate, as shown in Figure Copper Thickness vs Acrylic Thickness for Same Exp Rate Copper Thickness (mm) Figure 18. Copper thickness and equivalent acrylic thickness for the same exposure rate at 30 cm from II. Comparing the copper thickness to the equivalent acrylic thickness, the 2 mm of copper corresponds to an equivalent acrylic thickness of approximately 6.2 in. of acrylic and the 1 mm of copper has an equivalent acrylic thickness of 4.58 in. The calculated copper equivalent acylic thickness values are different from the assumption that 2 mm of copper simulates an abdomen 32

33 and the 1 mm of copper for the extremity, which have equivalent acrylic thicknesses of 7.84 in. and 2.47 in. in Table 3, but the measurements were made with copper placed on the II to drive the technique, whereas the image quality measurements are made with the copper on the exit of the X-ray tube. Similar to the radiography interface, the fluoroscopy dose can be calculated on a machine-to-machine basis. The skin dose and depth dose can be calculated as a function of exam type, magnification, patient thickness, depth, kvp, ma, and duration of exposure. Table 7 shows the program interface for fluoroscopy dose (C-arms and Fluoroscopy rooms). Table 7. Interface of the program for the calculation of fluoroscopic dose. (mm Al eq.) The limitations in the fluoroscopy program are similar to those of the radiography. An assumption made is the PDD, TAR, and backscatter factors are made on the central axis of the beam, meaning there is no off-axis correction or any out-of-field calculation for the dose. The next limitation stems from the data for PDD, backscatter factor, and TAR. Due to the limited values across HVL, kvp, and field sizes presented in Wagner et. al [3], the extrapolated data 33

34 may cause additional error from the extrapolation of the data. Lastly, similar to the radiography dose, the program does not take into account any changes in the X-ray spectrum from the publication of data from Wagner et. al [3]. The next portion concerns the patient dose estimation from interventional radiology (IR) procedures. In the field of diagnostic imaging physics, the doses given in the IR can be incredibly high and have the potential to be the largest of the imaging modalities. This high dose can cause permanent damage to the patient s skin, or damage needing surgery (i.e. skin graft). It is critical to monitor this dose to ensure the patient has not received an excess amount of dose in the IR procedure. A typical IR suite setup is two C-Arms set perpendicular to each other to image the patient in a frontal direction and lateral direction. The C-Arms are allowed to rotate around the patient, as well. The program for IR dose estimate is a correction from the displayed doses given on the monitors in the suite. There are three displayed doses that are used: frontal, lateral, and rotational digital angiography (DA). The displayed doses are given at the Interventional Reference Point (IRP). The IRP is defined on the central axis of the X-ray beam at 15 cm from isocenter towards the X-ray tube [10]. A typical patient setup is shown in Figures where the patient is centered at isocenter. Figure 19. Lateral patient setup for IR. 34

35 Figure 20. Frontal patient setup for IR. It is important to note in the frontal setup (Figure 20), the table and pad are in the beam, thus their attenuation should be accounted for in the dose estimation. To estimate the dose, it is necessary to know the HVLs, kv, SIDs, patient thickness, depth, field size at the II, and the lateral, frontal, and rotational displayed doses. Eq. 11 shows the skin dose estimation for the lateral exposure. D lat (Gy) = D dis(mgy) B 1.06 mgy t mgya ( IRP(cm) 1000 SSD lat (cm) )2 [Eq. 11] In Eq. 11, D lat is the lateral skin dose, D dis is the displayed dose for the lateral tube, B is the backscatter factor, IRP is the source-to-interventional reference point distance, and SSD lat is the source-to-surface distance for the lateral tube. Eq. 12 shows the skin dose estimation for the frontal exposure. D frontal (Gy) = D dis(mgy) B 1.06 mgy t mgya 1000 C att ( 2 IRP(cm) ) SSD frontal (cm) [Eq. 12] In Eq. 12, D frontal is the frontal skin dose, D dis is the displayed dose for the frontal tube, C att is the correction factor for the table-pad attenuation, and SSD frontal is the source-to-surface distance for the frontal tube. Lastly, Eq. 13 shows the skin dose estimation for the rotational dose. 35

36 D rot (Gy) = 1 3 D dis (mgy) 1000 [Eq. 13] In Eq. 13, D rot is the rotational skin dose and D dis is the displayed rotational skin dose. The 1/3 in Eq. 13 comes from the assumption the same portion of skin is in the X-ray beam about 1/3 of the time. To find the peak skin dose, it is simply just the sum of D lat, D frontal, and D rot, as shown in Eq. 14, where D peak is the peak skin dose. This method assumes the overlap of the frontal and lateral fields on the patient skin. D peak (Gy) = D lat (Gy) + D frontal (Gy) + D rot (Gy) [Eq. 14] To calculate dose at a depth, the PDD and TAR methods can be employed similarly as above. The IR dose estimation program can calculate skin dose and depth dose for varying HVLs, kv, SIDs, patient thickness, depth, field size at the II, and the lateral, frontal, and rotational displayed doses. Similarly, the dose can be estimated for different machines. Table 8 shows the IR dose estimation interface. Table 8. Program interface for IR dose estimation. (mm Al eq.) 36

37 There are some limitations to the program for IR dose estimation. According to the FDA 21CFR (k)(6), the maximum deviation between the displayed and measured doses should be less than +/-35% [11]. This means if the displayed dose is used by itself the uncertainty should be taken to be +/-35%. However, this deviation is measured in the annual inspections of the IR suites, and the uncertainty can be corrected for from the deviation in the annual reports. Similar to the interfaces for the radiography and fluoroscopy doses, the measurements for the dosimetric parameters is made on the central-axis of the X-ray beam. Also, concerning the PDD, backscatter factor, and TAR, the extrapolation of data outside of what is presented in Wagner et. al [3] can cause additional error. Again, for the IR dose, the program does not take into account the changes in the X-ray spectrum since the publication of Wagner et. al [3]. 37

38 CHAPTER IV COMPUTED TOMOGRAPHY DOSE Computed tomography (CT) is perhaps the most well-known of all the imaging modalities. CT uses a rotating X-ray tube to take many X-ray projections of the patient to create a series of cross-sectional images, or slices. While the CT can be used to produce sectional images, CT scans can give patients a large amount of dose, in terms of diagnostic imaging, which can be the largest when compared to other imaging modalities (the other comparable imaging modality being fluoroscopy, namely IR). The concern from CT is mainly stochastic effects, but there have been instances in the United States of occurrences of erythema and epilation from excessive dose. To calculate dose in CT, two parameters, volume computed tomography dose index (CTDI vol) and dose length product (DLP), are typically used. CTDI vol was developed to provide a standardized method to compare the radiation output between different scanners or different protocols through the use of a reference phantom, made of polymethyl methacrylate (PMMA), and the DLP is related to the total imparted energy to the reference phantom [12]. It should be noted that any CTDI mentioned is not the patient dose, but it is related to patient dose. CTDI vol and DLP are dependent on the kv, ma, gantry rotation time, pitch, and bowtie filter, but is independent of patient size [12]. Current CT scanners typically display the CTDI vol and DLP before and after each scan is taken [12]. The PMMA cylindrical phantom has two diameters, one 16 cm and the other 32 cm. The 16 cm phantom fits inside the 32 cm phantom, each have four holes bored at the 12, 3, 6, and 9 o clock positions, and the 16 cm has a hole bored in the center, as shown in Figure 21. These 38

39 holes can either hold the 100 mm long pencil ion chamber or have PMMA rods inside them when making CTDI vol measurements. Typically, when a head scan is performed the scanner choses the CTDI vol based off of the 16 cm diameter phantom. When a torso or abdomen of an adult is scanned, the 32 cm phantom is used as the reference phantom. Lastly, for pediatric body scans the manufacturer can choose to use the 16 cm or 32 cm diameter phantom for calculation of CTDI vol and DLP [12]. Figure 21. Picture of the 16 cm CTDI Phantom. The CTDI vol is defined through the following equations [13]: CTDI 100 (mgy) = 1 z=+50 mm D(z) dz nt z= 50 mm [Eq. 14] In Eq. 14, CTDI 100 is the CTDI for the 100 mm pencil chamber, n is the number of tomographic sections imaged simultaneously in a single rotation of the x-ray tube, T is the width of one tomographic section along the longitudinal z axis (nt is the nominal beam width), D(z) is the dose, and z is the position in the z direction (defined as direction perpendicular to the bore). The quantity measured for the CTDI 100 is air kerma, or kinetic energy released per unit mass, not the dose in tissue or dose in PMMA, and it is measured in terms of mgy. The CTDI 100 is measured 39

40 at the center of the phantom and at the periphery of the phantom. The weighted CDTI (CTDI w) is defined as: CTDI w (mgy) = 1 CTDI center (mgy) + 2 CTDI periphery (mgy) [Eq. 15] Lastly, the CTDI vol is defined as: CTDI vol (mgy) = CTDI w(mgy) P [Eq. 16] In Eq. 16, CTDI w is defined in Eq. 15 and P is the helical pitch, which is defined in Eq. 17 as: P = I (mm) nt(mm) [Eq. 17] where I is the table feed per 360 rotation of the gantry, and nt is the nominal beam width defined in Eq. 14. From the CTDI vol, the DLP is defined as: DLP(mGy cm) = CTDI vol (mgy) L(cm) [Eq. 18] where DLP is the dose-length product, CTDI vol is defined in Eq. 16, and L is the scan length. Once the CTDI vol and DLP are found, the dose can be investigated. There are many doses that the program can calculate. The first that is investigated is the effective dose. Effective dose is defined by the ICRP [2] as the tissue-weighted sum of the equivalent doses in all specified tissue and organs. This effective dose is found from the DLP and k-factor, which converts the DLP to effective dose (in msv) [14]. The k-factor values for different exam types for adults, 0 year old, 1 year old, 5 year old, and 10 year old can be found in AAPM Report 96 [14]. It should be noted the adult k-factors are for an adult of standard physique. The k-factors noted for adult head and neck and pediatric patients are for the 16 cm diameter phantom, and the other k-factors are for the 32 cm diameter CTDI phantom [14]. The k-factor for the different exam types decreases with age. To calculate the effective dose using the DLP, the following equation is used: 40

41 E (msv) = k( msv mgy cm ) DLP(mGy cm) [Eq. 19] In Eq. 19, E is the effective dose, k is the k-factor mentioned above, and the DLP is the doselength product. It should be noted the effective dose is not used to assess an individual s risk for cancer, rather give a comparison of the risk of irradiation relative to other procedures, to other radiation risks when concerning the patient, or to risks of everyday life, according to AAPM Report 96 [14]. The next dose quantity that is estimated in the program is the Size Specific Dose Estimate (SSDE). In AAPM Task Group (TG) 204 [12], the report provides conversion factors to be applied to CTDI vol, which is independent of patient size, to allow the practitioner to estimate patient dose. These factors take into account the patient size and estimate the patient dose based off of the CTDI vol and the factors associated with the size of the patient [12]. The SSDE estimation uses the definition of an effective diameter to determine the conversion factor for the SSDE. The effective diameter is based off of the patient s lateral dimension and anterior-posterior (AP) dimension. In AAPM TG 204 [12], the effective diameter is defined as: d eff (cm) = d lat (cm) d AP (cm) [Eq. 20] In Eq. 20, d eff is the effective diameter, d lat is the patient lateral dimension, and d AP is the patient AP dimension. Figure 22 shows an illustration of the effective diameter compared to a patient. It should be noted that a patient abdomen is ellipsoidal. Figure 22. Illustration showing the effective diameter versus the patient dimensions. 41

42 The size dependent conversion factors can be applied to both small and large patients. To calculate the SSDE, the following definitions are used [12]. 32 SSDE (mgy) = f cm 32 size CTDI cm vol (mgy) [Eq. 21] 16 SSDE (mgy) = f cm 16 size CTDI cm vol (mgy) [Eq. 22] Eq. 21 uses the conversion factors based off of the effective diameter for the 32 cm PMMA phantom and the CTDI vol calculated for the 32 cm PMMA phantom from the CT scanner. Eq. 22 uses the conversion factor for the effective diameter based off of the 16 cm phantom and the CTDI vol calculated for the 16 cm phantom displayed on the CT scanner. The conversion factors, as a function of effective diameter, for the 32 cm and 16 cm PMMA phantoms are from the report [12]. The conversion factors decrease with an increase in effective diameter and are higher for the 32 cm phantom for the same effective diameter. In AAPM TG 204 [12], it is suggested that these SSDE is used to estimate the patient dose from the output, CTDI vol, with a consideration based off of the patient size. There is a clear statement in TG 204 [12] that the SSDE should not be used to calculate a DLP, nor calculate an effective dose E, by using k-factors. Lastly, the last dose that is calculated in the dose estimation program is fetal dose. The fetal dose calculation is based off of the CTDI vol of the CT scan. This method is found in Wagner et. al [3]. The equations for calculating fetal dose in a CT scan are shown in Eqs D Primary (mgy) = D Ref (mgy) F(z = 0 mm) [Eq. 23] D Scatter (mgy) = ( T i (mm) 10 mm ) DRef (mgy) F(z i (mm)) i [Eq. 24] D Total (mgy) = D Primary (mgy) + D Scatter (mgy) [Eq. 25] 42

43 In Eq. 23, D Primary is the primary beam dose, D Ref is the reference dose (i.e. CTDI vol), F(0) is the phantom reference factor for z=0, which is the slice location of the fetus. If the fetus is not in the CT scan, the D Primary is equal to zero. In Eq. 24, D Scatter is the scatter beam dose, T i is the slice thickness of the i-th slice, D Ref is the reference dose (i.e. CTDI vol), F(z i) is the phantom reference factor for the distance z, of the i-th slice from the location of the fetus. In Eq. 25, D Total is the total fetal dose from primary and scatter dose, D Primary and D Scatter are the primary dose and scatter dose, defined in Eq. 23 and Eq. 24, respectively. The phantom reference factors are derived from Felmlee et. al [15]. With an increase in distance from the fetal slice, the phantom reference factor decreases until it is at a 180 mm distance. For slice thicknesses smaller than the 10 mm, since the table in Wagner et. al [3] goes in steps of 10 mm, the phantom reference factors can be linearly interpolated to obtain the factors for these smaller slices. Unfortunately, the error from the fetal dose calculation is unknown. The dose estimation program can calculate the effective dose for a CT scan as a function of the CTDI vol and the DLP, which are dependent on the kv, ma, gantry rotation time, pitch, and bowtie filter, as well as the age of the patient. As noted before, the effective dose is not meant to evaluate a specific patient risk from the CT scan. The program can also calculate the SSDE, as a function of the CTDI vol, which is again dependent on the kv, ma, gantry rotation time, pitch, and bowtie filter, as well as the patient size. The dose estimation program also can calculate fetal dose, as a function of the CTDI vol, slice thickness, and distance from the slice. Table 9 shows the interface of the dose estimation program for CT examinations. Lastly, the effective dose, SSDE, and fetal doses can be calculated on a machine-to-machine basis. While the program can estimate the dose to the patient, it does not come without limitations. First, as mentioned before, when using the calculation of the effective dose, the effective dose cannot be used to give a patient specific risk of cancer. Second, the SSDE only adds the consideration for patient size to find a patient dose, but should not be used to calculate 43

44 the effective dose through DLP and k-factors. For the fetal dose, the calculation does not take into account a change in depth, since the method outlined in Wagner et. al [3] does not provide the F(z) values as a function of depth for any value other than z=0 mm. When using the CTDI vol to calculate any of the mentioned doses, there is error when using the displayed CTDI vol. If the displayed CTDI vol is used by itself, the assumption of the difference of +/-20% from the measured value should be taken into account, since this would be the worst-case assumption for the unit to being used to scan within regulation, as defined in the ACR 2017 CT QC Manual [16]. However, this difference is measured for the annual CT QC, and can be easily corrected for. The program does not take into account any spectral data from the CT in the calculation of the doses. Lastly, these results are scanner dependent. Table 9. Program interface for the CT dose estimation. 44

45 CHAPTER V MAMMOGRAPHY DOSE Mammography is a very beneficial image modality to use. Use of the modality has yielded the early detection of breast cancers, which in turn reduces the mortality and morbidity rates of breast cancers [17]. Similar to the other imaging modalities, the use of mammography can present potential risk to the patient. However, the dose in mammography relative to the other imaging modalities is low, but there is still a risk of the development of secondary cancers, no matter how small the dose, from the linear no-threshold model. Since the dose is low in mammography, the main concern is the stochastic effects of radiation exposure. The typical dose referred to in mammography is the average glandular dose (AGD). In mammography, the breast is assumed to be composed of glandular tissue and adipose (fatty) tissue. Typical assumptions of breast composition are 50% glandular tissue and 50% adipose tissue, or what is known as 50% glandularity. The AGD is calculated from the entrance skin exposure (ESE) or entrance skin dose (ESD) and the use of dose conversion tables. The ESE or ESD can be found on the patient images or can be measured with an ion chamber (Figure 23). For the BR-12 phantom (Figure 24), the ESD and AGD were taken from the images. When measuring the ESE or ESD, the ACR Mammography Accreditation Phantom (Figure 25), which has a thickness of 4.5 cm is used. When measuring the AGD, a RadCal 10X6-6M ion chamber is used. The ion chamber must be calibrated every two years, and this ion chamber was calibrated on Feb. 16, The ESE and ESD have a dependence on the technique, such as the kvp and mas, used to image the patient, which depends on the compressed breast thickness and the composition. According to the RadCal calibration report, the measured dose and dose rate measurements have an uncertainty of +/-5% at the 95% confidence level. 45

46 Figure 23. Setup of the Rad Cal 10X6-6M ion chamber and the ACR Mammography Accreditation Phantom. Figure 24. BR-12 Phantom with 8 cm thickness on a Hologic Selenia unit. 46

47 Figure 25. ACR Mammography Accreditation Phantom. The AGD is calculated using the following equation from the ACR 2016 Digital Mammography QC Manual [18]. AGD (mgy) = K(R) g c s 8.76 mgy R [Eq. 26] In Eq. 26, AGD is the average glandular dose, K is the entrance skin exposure, g is the g-factor for breast simulated with acrylic or BR-12, c is the c-factor for breast simulated with acrylic or BR-12, and s is the s-factor for clinically used spectra. K is the exposure with the absence of backscatter at the entrance surface of the breast. The g-factor corresponds to a glandularity of 50%, the c-factor corrects for any deviation in breast composition from 50% glandularity, and the s-factor corrects for differences, due to choice in X-ray spectrum (e.g. target/filter combination) [18]. The s-factors are all greater than 1 (except for Mo/Mo), due to the fact the s- factor is relative to Mo/Mo. This method only applies to the 2D mammography dose modes. To estimate the patient dose, it is important to note how the AGD changes with different phantom thicknesses. The AGD variation should also be investigated between 2D and 3D modes. Another important aspect is to note the differences between the displayed and the measured AGDs to ensure an accurate estimation of patient dose. 47

48 Displayed AGD (mgy) Testing was done using Hologic Selenia Dimensions 3D, Hologic Selenia 2D Mammography Systems, and Hologic MultiCare Platinum Stereotactic Breast Biopsy (SBB). Variations of the technique were recorded for different BR-12 phantom thicknesses ranging from 2 cm to 8 cm, and for different compensation steps, using the 4 cm thickness. A compensation step of 0 was used for all measurements, except for the compensation step measurements. The displayed AGD was also noted for each thickness of phantom used, as well as the target and filter used. Measurements of the AGD were made using a RadCal 10X6-6M chamber and the ACR Mammographic Accreditation Phantom. When applicable, 2D and 3D displayed and measured AGD values were compared for the same setups and phantom thickness. For the Dimensions with 3D capability, the displayed AGD for both the 2D and 3D modes change exponentially for differing BR-12 phantom thicknesses, as shown in Figure 26. For the 2 cm and 4 cm thicknesses, the 3D displayed AGD was higher than that of the 2D displayed AGD. For the 6 cm and 8 cm BR-12 thicknesses, the 3D displayed AGD was lower than the 2D displayed AGD. 6 Phantom Thickness vs. AGD Hologic Selenia Dimensions D 3D y = e x R² = y = e 0.239x R² = BR-12 Phantom Thickness (cm) Figure 26. Phantom thickness vs. displayed AGD for 2D and 3D modes. For the Hologic Selenia with only 2D mode, the displayed AGD was linear as a function of the BR-12 phantom thickness, as shown in Figure

49 Displayed AGD (mgy) Displayed AGD (mgy) Phantom Thickness vs. AGD Hologic Selenia 2D y = x R² = BR-12 Phantom Thickness (cm) Figure 27. Phantom thickness vs. displayed AGD for 2D only system. For the SBB unit, the displayed AGD changed exponentially with BR-12 phantom thickness, as shown in Figure 28. Phantom Thickness vs. AGD Hologic MultiCare Platinum y = e x R² = BR-12 Phantom Thickness (cm) Figure 28. Phantom thickness vs. displayed AGD for SBB. Comparing the displayed AGD to the measured AGD, the absolute value of the percent differences ranged from 0.87% to 29.97%, shown in Table 10. These percent differences are measured across the different units for different modes (2D, 3D, & combo) and different target/filter combinations. For the 95% confidence level, the measured AGD values have an uncertainty of +/- 5% from the RadCal calibration report. 49

50 AGD (mgy) Table 10. Percent differences between the displayed AGD and measured AGD for different units. Mammography Unit Mode Target/Filter Displayed AGD Measured AGD Percent (mgy) (mgy) Difference 2D W/Rh / % 2D W/Ag / % Hologic Selenia /- 3D W/Al 1.58 Dimensions # % C3D W/Al / % C2D W/Rh / % /- Hologic Selenia 2D W/Ag % Dimensions #2 2D W/Rh / % Hologic Selenia 2D 2D W/Rh / % Hologic MutiCare Platinum 2D Mo/Mo / % Lastly, the AEC compensation steps were tested using a constant 4 cm BR-12 phantom thickness. The change in AGD appears to be linear with compensation step, as demonstrated in Figure 29. It should be noted that the 6 cm BR-12 phantom thickness, with a compensation step of 0, had an AGD of 2.48 mgy on the Hologic Selenia (Figure 29), which is lower than the 4 cm BR-12 phantom thickness with a compensation step of 4. AEC Compensation Steps vs. AGD y = x R² = cm BR AEC Compensation Steps Figure 29. AEC Compensation Steps vs AGD for 4 cm BR-12 thickness. 50

51 The AGD varies with phantom thickness, target/filter combination, AEC compensation steps, and 2D/3D modes. However, it should be noted that it is unknown how the mammography system calculates the AGD. Importing this data found into a dose estimation program can give a quick, automatic way to estimate the patient breast dose. The program calculates AGD as a function of kvp, mas, HVL, breast composition, compression thickness, and target/filter combination. The program can also calculate the AGD on a machine-to-machine basis. Table 11 shows the program interface for mammography dose estimation. Table 11. Program interface for the mammography dose estimation. (mm Al eq.) While the program will estimate the dose, it does not come without limitations. The first is the program calculates the AGD based on a breast with 50% glandularity, which does not apply to specific patients, as there can be many combinations of breast composition. Second, whether using the BR-12 phantom or the ACR Mammography Accreditation Phantom, the AGD is calculated to the phantoms, not to the breast of the patient. While they maybe similar, they are not the same material. Third, if the displayed AGD is used by itself, there is a difference from the measured AGD, as shown in Table 10. However, this can be corrected if the percent difference is accounted for. 51

52 CHAPTER VI CONCLUSION While ionizing radiation can provide wonderful benefits, it can also do harm to the patient. While in the chapters above the risk has been discussed, there has been little mention of how to quantify the risk to the patient. There was a mention about stochastic risks in Chapter 1, and the risk of a population developing a secondary cancer from some amount of dose is approximately 5% per Sv [2]. For estimating the risk of radiation to the fetus, the main source for the evaluation is NCRP 54 [19], which states the risk of abnormality is negligible if the dose is less than 50 mgy and significantly increased for doses higher than 150 mgy for the fetus. When concerning the dose to the skin, there is different effects discussed in Balter et. al [20]. In the paper, the tissue reactions for the skin are discussed, as well as a general time frame before effects are seen for different doses, from 0 Gy to >15 Gy. Similarly, for fluoroscopy skin dose, the Joint Commission has defined a Sentinel Event of 15 Gy, over a period of 6 months to a year for fluoroscopy guided interventional procedures [21]. If a Sentinel Event occurs, the facility may report the occurrence to the Joint Commission. According to the Mammography Quality Standards Act (MQSA), the average glandular dose per view for a 4.2 cm compressed breast with 50% glandularity must be less than 3.0 mgy [22]. While the dose estimation program does have its limitations, it still provides a quick, easy, and automatic way of estimating patient dose. For determining the fetal dose risk, the program can be applied prospectively and retrospectively to give a risk estimate based on the procedure. Similarly, when evaluating skin dose in radiography and fluoroscopy, the program can provide the estimate of dose before and/or after a procedure has occurred. When concerning IR, however, due to the method by which the peak skin dose is estimated, the program is much better applied retrospectively. However, the program can be applied before, 52

53 with a ballpark of the displayed doses. For the CT portion, the program works well in a prospective and retrospective manner, since the CTDI vol and DLP are displayed before and after a scan, for modern scanners. Lastly, for mammography, the program can be applied in both manners, however, should be taken with some caveat, due to assumptions about the glandularity and the phantom material. When concerning the Joint Commission, having a program that quickly and automatically estimates the dose the patient receives, can provide a benefit to determining the occurrence of a Sentinel Event. For future work, there are a few things that can be investigated further. First, is the investigation of uncertainty analysis and error propagation. Throughout the thesis, some of the uncertainties have been discussed, but not the propagation to give a final uncertainty in the patient dose estimation. Second, is the investigation into the differences in the energy spectrum for X-ray production, since the publication of Wagner et. al [3]. The differences in the X-ray spectrum may have an effect on dosimetric parameters (i.e. PDD, Backscatter factor, TAR, etc.), which will have an effect on the dose estimation. Lastly, further investigation into the PDD, backscatter factors, and TAR values mentioned in Wagner et. al [3] is needed. They are limited by HVL, kvp, and field sizes. Investigating the PDDs, backscatter factors, and TARs will improve the quality of the program, as well as remove the need for the extrapolation of the data found in Wagner et. al [3], thus removing potential error from extrapolation. Methods for investigation of the PDDs, backscatter factors, and TARs can be through measurement or Monte Carlo simulations. 53

54 BIBLIOGRAPHY [1] Chodos, A., & Ouellette, J. (Eds.). November 8, 1895: Roentgen's Discovery of X- Rays. Retrieved February 16, 2018, from [2] ICRP, The 2007 Recommendations of the International Commission on Radiological Protection. ICRP Publication 103. Ann. ICRP 37 (2-4). [3] Wagner, L.K., Lester, R.G., Saldana, L.R. Exposure of the Pregnant Patient to Diagnostic Radiation. Madison: Medical Physics Publishing; [4] Harrison R.M. Central Axis Depth-Dose Data for Diagnostic Radiology. Phys Med Biol 26: 657, [5] Säbel M., Bednar W., Weishaar J. Investigation of the exposure to radiation of the embryo/fetus in the course of radiographic examinations during pregnancy. First communication: Tissue-Air Ratios for X Rays with Tube Voltages between 60 kv and 120 kv. Strahlentherapie 156: 502, [6] Fetterly, K.A., Gerbi, B.J., Parham, A., Geise, R.A. Measurement of the dose deposition characteristics of x-ray fluoroscopy beams in water. Medical Physics 28(2): [7] Dauer, L.T., et al. Radiation Management for Interventions Using Fluoroscopic or Computed Tomographic Guidance during Pregnancy: A Joint Guideline of the Society of Interventional Radiology and the Cardiovascular and Interventional Radiological Society of Europe with Endorsement by the Canadian Interventional Radiology Association. Journal of Vascular and Interventional Radiology 23: [8] Reynoso, F.J., Tailor, R., Wang, C., Cho, S.Y. Comparison of filtered x-ray spectra and depth doses derived from a hybrid Monte Carlo model of an orthovoltage x-ray unit with experimental measurements. Biomedical Physics & Engineering Express 2: [9] Layfield, C., Garcia, V.L., Wang, W. Radcal Solid State Multisensor 1-Shot Half Value Layer Evaluation. American Association of Physicists in Medicine Annual Meeting,

55 [10] Balter, S. (2006). Methods for measuring fluoroscopic skin dose. Pediatric Radiology, 36(Suppl 2), [11] Food and Drug Administration. Performance Standards for Ionizing Radiation Emitting Products. 21 CFR [12] American Association of Physicists in Medicine. Size-specific dose estimates (SSDE) in pediatric and adult body CT Examinations: report of AAPM Task Group 204. College Park, Md: American Association of Physicists in Medicine, [13] Bauhs, J.A., Vrieze, T.J., Primak, A.N., Bruesewitz, M.R., McCollough, C.H. CT Dosimetry: Comparison of Measurement Techniques and Devices. RSNA RadioGraphics, Vol. 28, Issue 1, [14] American Association of Physicists in Medicine. The Measurement, Reporting, and Management of Radiation Dose in CT: report of AAPM Task Group 23 of the Diagnostic Imaging Council Committee. College Park, Md: American Association of Physicists in Medicine, [15] Felmlee, J.P., Gray, J.E., Leetzow, M.L., Price, J.C. Estimated fetal radiation dose from multislice CT studies. AJR , [16] American College of Radiology Computed Tomography Quality Control Manual. ACR [17] American Cancer Society. American Cancer Society Recommendations for the Early Detection of Breast Cancer. ACS Retrieved from [18] Berns E.A., Baker J.A., Barke L.D., et al. Digital Mammography Quality Control Manual. Reston, Va: American College of Radiology [19] National Council on Radiation Protection and Measurements. NCRP Report 54: Medical Radiation Exposure of Pregnant and Potentially Pregnant Women. Washington: National Council on Radiation Protection and Measurements; 1-31, [20] Balter, S., Hopewell, J.W., Miller, D.L., Wagner, L.K., Zelefsky, M.J. Fluoroscopy Guided Interventional Procedures: A Review of Radiation Effects on Patients Skin and Hair. Radiology, 254(2),

56 [21] The Joint Commission. Radiation Overdose as a Reviewable Sentinel Event. Oakbrook Terrance: The Joint Commission. 1-2, Retrieved from [22] Food and Drug Administration. Mammography Quality Standards Act Regulations. MQSA Subpart B Sec Retrieved from EmittingProducts/MammographyQualityStandardsActandProgram/Regulations/ucm htm 56

57 PRESENTATIONS *=denotes presenting author A Program for Fetal Dose Estimation in Radiographic Examinations V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center Oral Presentation Graduate Research Education and Technology Symposium 2017 University of Oklahoma Health Sciences Center Oklahoma City, Oklahoma Patient Radiation Dose Estimation Program V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center Oral Presentation Graduate Research Education and Technology Symposium 2017 University of Oklahoma Health Sciences Center Oklahoma City, Oklahoma A Program for Fetal and Peak Skin Dose Estimations for Radiographic and Fluoroscopic Examinations V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center eposter Presentation AAPM Annual Meeting 2017 The American Association of Physicists in Medicine Denver, Colorado Radcal Solid State Multisensor 1-Shot Half Value Layer Evaluation C. Layfield*, V. Garcia, W. Wang, University of Oklahoma Health Sciences eposter Presentation AAPM Annual Meeting 2017 The American Association of Physicists in Medicine Denver, Colorado Evaluation of the Role of the Physicist When Imaging the Pregnant Patient V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center Poster Presentation Education Week 2017 University of Oklahoma Health Sciences Center Oklahoma City, Oklahoma Variation of Fluoroscopy Technique & Exposure Rate for Different Phantoms for Dose Estimation Program V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center Poster Presentation AAPM Spring Clinical Meeting 2018 The American Association of Physicists in Medicine Las Vegas, NV Patient Radiation Dose Estimation Program for Radiography, Fluoroscopy, Computed Tomography, and Mammography Examinations V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center Oral Presentation Graduate Research Education and Technology Symposium 2018 University of Oklahoma Health Sciences Center Oklahoma City, Oklahoma 57

58 A Program for Mammography Radiation Dose Estimation V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center Oral Presentation Graduate Research Education and Technology Symposium 2018 University of Oklahoma Health Sciences Center Oklahoma City, Oklahoma Mammography, Tomosynthesis, and SBB AEC Performance and Displayed AGD Accuracy for Dose Estimation Program V. Garcia*, W. Wang, University of Oklahoma Health Sciences Center eposter Presentation AAPM Annual Meeting 2018 The American Association of Physicists in Medicine Nashville, Tennessee 58

59 APPENDIX This appendix is added to show extra data collected that was not presented in the main body of the thesis, as well as extra information that is relevant. While only presented once in the main body, many radiographic rooms were tested to find the exposure at 40 in. and the HVL data as a function of kvp. These were tested with the RadCal10X5-6 ion chamber placed 40 in. from the focal spot of the X-ray tube to measure exposure and the AGMS-DM+ Solid-State Detector for the One-Shot HVL data. Table A-1 shows the exposure and HVL data collected for Philips Radiographic Room in OUCPB Table A- 1. Exposure at 40 in. and HVL data as a function of kvp for OUCPB Table A-2 shows the exposure data and HVL for the OUCPB 3512 Philips Radiographic Room. 59

60 Table A- 2. Exposure and HVL data as a function of kvp for OUCPB Table A-3 shows the exposure and HVL data for OUCPB 3513 Philips Radiographic Room. Table A- 3. Exposure and HVL data as a function of kvp for OUCPB

61 Table A-4 shows the exposure and HVL data for the Philips Radiographic Room in OUCPB Table A- 4. Exposure and HVL data as a function of kvp for OUCPB Table A-5 shows the exposure and HVL data for the GE Proteus Radiographic Room in OUPB Rm 3. Table A- 5. Exposure and HVL data as a function of kvp for the OUPB GE Proteus Room 3. 61

62 Table A-6 shows the exposure and HVL data for the GE Proteus Radiographic Room in OUPB Room 4. Table A- 6. Exposure and HVL data as a function of kvp for the OUPB GE Proteus Room 4. While this is not all the radiographic rooms on OUHSC campus, further experimentation is required for each radiographic room. Testing for each room gives the machine-to-machine basis that was presented in the main body of the thesis. More important information to consider is the X-ray tube information. The recording of the X-ray tube information provides a way to retrace the X-ray tube used for testing. While this may not necessarily be important for rooms that have house X-ray tubes, this is incredibly important for portable units. Table A-7 shows the X-ray tube information for the radiographic rooms tested for the exposure and HVL data. 62

63 Table A- 7. Tube information for the radiographic rooms mentioned in Table A-1 through A-6. For the Variation of Fluoroscopy Technique and Exposure Rate for Varying Acrylic Thickness project, the data was collected for GE OEC 9900 Elite and Philips BV Pulsera C-Arms. The RadCal 10X5-6 ion chamber was placed at 30 cm from the image intensifier (II) for both C-Arms. The data was presented for the ABC curves in Figures 10-13, and the measured exposure rates for varying acrylic thickness in Figures The figures were adapted from the following tables to create the ABC curves and the exposure rate data. Table A-8 shows the technique and exposure rate data for the high-level fluoroscopy on normal magnification mode for the GE OEC 9900 Elite. Table A- 8. Technique and exposure rate variations for the GE OEC 9900 for HLF at normal magnification. 63

64 Table A-9 shows the variation of technique and exposure rate for the GE OEC 9900 Elite for HLF on Magnification 1 for varying acrylic thickness. Table A- 9. Variation of technique and exposure rate for the GE OEC 9900 Elite for HLF at Magnification 1. Table A-10 shows the variation of the technique and exposure rate for the GE OEC 9900 Elite for HLF on Magnification 2 for varying acrylic thickness. Table A- 10. Variation of technique and exposure rate for GE OEC 9900 Elite for HLF on Magnification 2. 64

65 Table A-11 shows the variation of technique and exposure rate for the GE OEC 9900 Elite for Standard Fluoroscopy for Normal Magnification. Table A- 11. Variation of technique and exposure rate for GE OEC 9900 Elite for Standard Fluoroscopy at Normal Magnification. Table A-12 shows the variation of technique and exposure rate for the GE OEC 9900 Elite for Standard Fluoroscopy for Magnification 1. Table A- 12. Variation of technique and exposure rate for GE OEC 9900 Elite Standard Fluoroscopy at Magnification 1. 65

66 Table A-13 shows the variation of technique and exposure rate for the GE OEC 9900 Elite for Standard Fluoroscopy at Magnification 2. Table A- 13. Variation of technique and exposure rate for GE OEC 9900 for Standard Fluoroscopy at Magnification 2. Next, is the data for the Philips BV Pulsera C-Arm. When testing the variation, only up to 8 in. of acrylic were used, as that was what we were able to bring to testing. Table A-14 shows the variation of technique and exposure rate for the Philips BV Pulsera for HLF with a 31 cm II size. Table A- 14. Variation of technique and exposure rate for the BV Pulsera for HLF with 31 cm II size. 66

67 Table A-15 shows the variation of technique and exposure rate for the Philips BV Pulsera for HLF with 23 cm II size. Table A- 15. Variation of technique and exposure rate for the Philips BV Pulsera for HLF with 23 cm II size. Table A-16 shows the variation of technique and exposure rate for the Philips BV Pulsera for HLF with 17 cm II size. Table A- 16. Variation of technique and exposure rate for the PHilips BV Pulsera for HLF with 17 cm II size. Table A-17 shows the variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 31 cm II size. 67

68 Table A- 17. Variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 31 cm II size. Table A-18 shows the variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 23 cm II size. Table A- 18. Variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 23 cm II size. Table A-19 shows the variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 17 cm II size. 68

69 Table A- 19. Variation of technique and exposure rate for the Philips BV Pulsera for standard fluoroscopy with 17 cm II size. As with the radiography X-ray tubes, the same was recorded for the C-Arm X-ray tubes. While this does not encompass every tube on the OUHSC Campus, these X-ray tubes were the ones the most detailed data was collected for. Table A-20 shows the X-ray tube information for the GE OEC 9900 Elite and the Philips BV Pulsera C-Arms. Table A- 20. X-Ray tube information for the GE OEC 9900 Elite and Philips BV Pulsera C-Arms. For computed tomography, data will be collected in the future. Table A-21 shows the CT X-ray tube information for these X-ray tubes. Table A- 21. CT X-Ray Tube information. 69

70 The BR-12 phantom thickness vs. displayed AGD shown in Figures are adapted from the Tables A-22 through A-24. Table A-22 shows the technique, ESD, and AGD information for varying BR-12 phantom thickness for the Hologic Selenia Dimensions from Figure 26. Table A- 22. Technique and dose information from the Hologic Selenia Dimensions unit. Table A-23 shows the technique, ESD, and AGD for varying BR-12 phantom thickness for the Hologic Selenia with 2D only capability from Figure 27. Table A- 23. Technique and dose information for the Hologic Selenia 2D. Table A-24 shows the technique, ESD, and AGD for varying BR-12 phantom thickness for the Hologic MultiCare Platinum Stereotactic Breast Biopsy (SBB) from Figure

71 Table A- 24. Technique and Dose information for the Hologic MultiCare Platinum SBB unit. For the compensation steps vs. AGD shown in Figure 29, the figure was adapted from the data presented than Table A-25. This was found on the Hologic Selenia 2D. Table A- 25. AEC compensation steps vs. mas and displayed AGD for Hologic Selenia 2D. Lastly, similar to the radiography, fluoroscopy, and CT units, the X-ray tube information of mammography systems was recorded in Table A-26. Table A- 26. Mammography X-Ray tube information. 71