Radiation Dose Determination Using Mosfet and Rpl Dosimeters in X-Ray Imaging

D 1514 OULU 2019 D 1514

UNIVERSITY OF OULU P.O. Box 8000 FI-90014 UNIVERSITY OF OULU FINLAND ACTA UNIVERSITATIS OULUENSIS ACTA UNIVERSITATIS OULUENSIS ACTA

DMEDICA Antti Kotiaho Antti Kotiaho University Lecturer Tuomo Glumoff RADIATION DOSE University Lecturer Santeri Palviainen DETERMINATION USING

Senior research fellow Jari Juuti MOSFET AND RPL DOSIMETERS IN X-RAY Professor Olli Vuolteenaho IMAGING

University Lecturer Veli-Matti Ulvinen

Planning Director Pertti Tikkanen

Professor Jari Juga

University Lecturer Anu Soikkeli

Professor Olli Vuolteenaho UNIVERSITY OF OULU GRADUATE SCHOOL; UNIVERSITY OF OULU, FACULTY OF MEDICINE; MEDICAL RESEARCH CENTER OULU; Publications Editor Kirsti Nurkkala OULU UNIVERSITY HOSPITAL

ISBN 978-952-62-2264-6 (Paperback) ISBN 978-952-62-2265-3 (PDF) ISSN 0355-3221 (Print) ISSN 1796-2234 (Online)

ACTA UNIVERSITATIS OULUENSIS D Medica 1514

ANTTI KOTIAHO

RADIATION DOSE DETERMINATION USING MOSFET AND RPL DOSIMETERS IN X-RAY IMAGING

Academic dissertation to be presented with the assent of the Doctoral Training Committee of Health and Biosciences of the University of Oulu for public defence in Auditorium 7 of Oulu University Hospital, on 24 May 2019, at 12 noon

Supervised by Professor Miika Nieminen Docent Juha Nikkinen

Reviewed by Docent Mika Kortesniemi Docent Jari Heikkinen

Opponent Docent Paula Toroi

ISBN 978-952-62-2264-6 (Paperback) ISBN 978-952-62-2265-3 (PDF)

ISSN 0355-3221 (Printed) ISSN 1796-2234 (Online)

Cover Design Raimo Ahonen

JUVENES PRINT TAMPERE 2019 Kotiaho, Antti, Radiation dose determination using MOSFET and RPL dosimeters in x-ray imaging. University of Oulu Graduate School; University of Oulu, Faculty of Medicine; Medical Research Center Oulu; Oulu University Hospital Acta Univ. Oul. D 1514, 2019 University of Oulu, P.O. Box 8000, FI-90014 University of Oulu, Finland

Abstract Medical x-ray imaging is used to visualise patients’ anatomical structures and in some cases their physiology. X-rays are ionizing radiation, thus their use needs to be optimised, as stochastic effects are assumed to increase linearly with the exposure dose. Imaging protocols need to be optimised to a radiation dose level that follows the as low as reasonably achievable principle without compromising the diagnostic value of the image. Different methods can be used to help in the optimisation process, such as simulations, radiation dose and image quality assessments with dosimeters and phantoms and utilising the latest technology in the most efficient way. The purpose of this doctoral thesis was to investigate the applicability of metal-oxide- semiconductor-field-effect-transistor (MOSFET) dosimeters for dose determinations in conventional x-ray and computed tomography (CT) examinations. Additionally, dose optimising methods were investigated in dental panoramic imaging using radiophotoluminescence (RPL) dosimeters. Anthropomorphic phantoms were used in every study to simulate patients, as their structures enable dosimeters to be positioned at locations that correspond to different organs. The MOSFET’s properties for dose determinations were evaluated against the reference dosimeter in a conventional x-ray set-up. Comparisons of absorbed and effective doses in thorax x-ray imaging were made between RPLs, MOSEFTs and Monte Carlo simulations. The effect of the organ-based tube current modulation and bismuth shields were compared against the reference imaging method in a chest CT with one scanner model. Absorbed doses and quantitative image quality were evaluated using each method. Possible dose reduction from segmented dental panoramic tomography (sDPT) imaging was compared against full DPT. Dose measurements were done using RPL dosimeters in pediatric and adult set-up using phantoms. MOSFETs are accurate enough to be used in conventional x-ray and CT, but they require a careful calibration before use as their reproducibility is limited with low doses. Bismuth shields provided the best dose reduction, but with a negative impact on quantitative image quality, especially when metal artefact removal software was used. The final study showed that the use of sDPT programmes and pediatric protocols enable a notably dose reduction compared to the full DPT adult protocol.

Keywords: computed tomography, dental panoramic imaging, optimisation, radiation dose, x-ray

Kotiaho, Antti, Säteilyannoksen määritys röntgenkuvantamisessa käyttäen MOSFET- ja RPL-dosimetreja. Oulun yliopiston tutkijakoulu; Oulun yliopisto, Lääketieteellinen tiedekunta; Medical Research Center Oulu; Oulun yliopistollinen sairaala Acta Univ. Oul. D 1514, 2019 Oulun yliopisto, PL 8000, 90014 Oulun yliopisto

Tiivistelmä Lääketieteellisessä kuvantamisessa käytetään röntgensäteilyä potilaan anatomian ja joissain tapauksissa fysiologian visualisointiin. Röntgensäteily on ionisoivaa ja stokastisten vaikutusten kasvaessa oletettavasti lineaarisesti säteilyn funktiona, tulee säteilyn olla kokonaisvaltaisesti optimoitua. Kuvauksissa käytetyn röntgensäteilyn käytön tulee noudattaa ALARA-periaatetta, minkä vuoksi kuvauksessa tulee käyttää niin vähän säteilyä kuin vain mahdollista, diagnostiikan vaarantumatta. Optimoinnin apuna voidaan käyttää esim. simulointeja, annos- ja kuvanlaatu- määrityksiä dosimetreilla ja fantomeilla, tai laitevalmistajien tuomia uusia teknologioita. Tämän väitöskirjan tarkoituksena oli tutkia metallioksidi-puolijohdekanavatransistorien (MOSFET) soveltuvuutta natiiviröntgentutkimuksissa ja tietokonetomografiassa (TT). Lisäksi työssä tutkittiin hammaskuvauksissa käytettyjä annossäästömenetelmiä radiofotoluminesenssi- dosimetreilla (RPL). Potilasvasteena työssä käytettiin antropomorfisia fantomeita, minkä ansios- ta säteilyannoksia voidaan mitata eri elimiä vastaavilta kohdilta. MOSFET annosmittarin ominaisuuksia arvioitiin natiiviröntgenasetelmassa referenssimitta- riin nähden. Absorboituneiden ja efektiivisten annosten eroa MOSFET:tien, RPL:ien ja simu- lointien kesken tutkittiin keuhkoröntgentutkimuksessa. Pintakudoksia säästävän putkivirranmo- dulointimenetelmän ja vismuttisuojien vaikuttavuutta verrattiin TT:ssä referenssimetelmää vas- ten. Vaikuttavuutta arvioitiin absorboituneiden annosten ja kvantitatiivisen kuvanlaadun avulla. Segmentoidun hammaspanoraamakuvauksen (sDPT) annossäästömahdollisuuksia verrattiin tavalliseen panoraamakuvaukseen. Annosmääritykset tehtiin käyttäen RPL dosimetreja lapsi- ja aikuisfantomeissa. MOSFET dosimetreja voidaan käyttää annosmäärityksiin natiiviröntgenkuvauksissa ja TT:ssä, mutta niiden kalibrointi ja toistettavuus matalilla annoksilla aiheuttaa kuitenkin rajoituk- sia niiden käytölle. Vismuttisuojat tuottivat parhaan annossäästön, huonontaen kuitenkin kuvan- laatua. Kuvanlaadun huonontuminen oli erityisen huomattavaa, kun metallista aiheutuvien kuva- virheiden poistamiseen suunniteltua ohjelmaa käytettiin. Viimeinen tutkimus osoitti, että sDPT ohjelmat ja lapsille suunnatut protokollat mahdollistavat huomattavan annossäästön verrattuna aikuisten kokopanoraamaan.

Asiasanat: optimointi, panoraamakuvaus, röntgen, säteilyannos, tietokonetomografia

At my age, the radiation will probably do me good. Sir Norman Wisdom

To my loved ones

8 Acknowledgements

This study was carried in the Department of Diagnostic Radiology, Oulu University Hospital and the University of Oulu during the years of 2012-2019. I owe my gratitude to my principal supervisor Professor Miika Nieminen, Ph.D., for his guidance and advices throughout this project and for giving me an opportunity to do my thesis alongside with my medical physicist residency. I’m most grateful to my second supervisor Docent Juha Nikkinen, Ph.D., who guided me to the field of Computed Tomography and has given me advices beyond count. I want to express my most sincere thanks to my colleague and co-author Ph.D.Anna-Leena Manninen for her teachings in dosimetry and giving me counsel whenever needed. I’m deeply grateful to my co-author DDS., Ph.D., Annina Sipola for her ideas, assistance and enthusiasm in our study. I wish to thank the official pre-examiners Docent Mika Kortesniemi, Ph.D., and Docent Jari Heikkinen, Ph.D., for their constructive criticism, numerous comments and suggestions to improve the quality of the thesis. I would like to thank Docent Eveliina Lammentausta, Ph.D., the chairperson of my follow-up group for her guidance during these years. I’d like to thank my colleagues/co-authors Matti Hanni, Ph.D., Arttu Peuna, M.Sc., Sakari Karhula, Ph.D., Marianne Haapea, Ph.D., Soili Kallio-Pulkkinen, DDS., Ph.D., and Essi Happo, DDS., for their efforts and guidance in these years during these studies. Finally, I want to express my deepest gratitude to my family. My better half Suvi, brother Henri, sister-in-law Päivi and my parents Tuula and Harri for their endless support, patience and love.

Oulu, April 2019 Antti Kotiaho

10 Abbreviations

ALARA As low as reasonably achievable AP Anterior-posterior CF Calibration factor CTDI Computed tomography dose index CV Coefficient of variation DAP Dose area product DICOM Digital Imaging and Communication in Medicine DLP Dose-length product DPT Dental panoramic tomography E Effective dose ESD Entrance surface dose FDD Focus-to-detector distance FOV Field of view FSD Focus-to-skin distance HU Hounsfield unit ICRP International Commission of Radiation Protection KERMA Kinetic energy released per unit mass LAT Lateral MOSFET Metal oxide semiconductor field effect transistor OBTCM Organ-based tube current modulation OEM Organ effective modulations, see OBTCM PA Posterior-anterior PMMA Polymethyl methacrylate Q Tube current-time product ROI Region of interest RPL Radiophotoluminescence sDPT Segmented dental panoramic tomography uc Combined uncertainty wT Tissue weighting factor Y Tube output

12 List of original publications

This thesis is based on the following publications, which are referred to throughout the text by their Roman numerals (I–III):

I Manninen AL*, Kotiaho A*, Nikkinen J, Nieminen MT (2015). Validation of a MOSFET dosimeter system for determining the absorbed and effective radiation doses in diagnostic radiology. Radiation Protection Dosimetry, Apr;164(3), 361–7. II Kotiaho A, Manninen AL, Nikkinen J, Nieminen MT (2018). Comparison of organ- based tube current modulation and bismuth shielding in chest CT: Effect on the image quality and the patient dose. Radiation Protection Dosimetry, Dec. Ahead of print. III Kotiaho A*, Sipola A*, Happo E, Haapea M, Nikkinen J, Kallio-Pulkkinen S, Nieminen MT (2019). Use of segmented dental panoramic tomography (sDPT) for dose reduction in comparison to full DPT. Manuscript

* Equal contribution

This thesis also contains unpublished data.

14 Table of contents

Abstract Tiivistelmä Acknowledgements 9 Abbreviations 11 List of original publications 13 Table of contents 15 1 Introduction 17 2 Literature review 19 2.1 Radiation dose determinations ...... 19 2.1.1 Radiation dose quantities ...... 19 2.1.2 Organ dose determinations using phantoms ...... 22 2.1.3 Organ dose determinations using simulations ...... 22 2.2 Imaging devices ...... 23 2.2.1 Conventional x-ray and dental panoramic devices ...... 23 2.2.2 Computed tomography ...... 24 2.3 Dosimeters ...... 25 2.3.1 MOSFET dosimeters ...... 25 2.3.2 Radiophotoluminescence dosimeters ...... 28 3 Purpose of the study 31 4 Materials and methods 33 4.1 Materials ...... 33 4.1.1 Dose determinations ...... 33 4.1.2 Phantoms ...... 35 4.1.3 Imaging equipment ...... 36 4.2 Methods ...... 37 4.2.1 RPL and MOSFET in conventional radiology ...... 37 4.2.2 Organ effective modulation vs. bismuth shields ...... 38 4.2.3 Segmented dental panoramic tomography ...... 41 4.3 Measurement uncertainties...... 42 5 Results 45 5.1 RPL and MOSFET in conventional radiology ...... 45 5.1.1 MOSFET’s energy dependence of the response ...... 45 5.1.2 MOSFET’s linearity of the response ...... 45 5.1.3 MOSFET’s repeatability of the dosimeter response ...... 46 5.1.4 Absorbed and effective dose determination ...... 47

15 5.2 Organ effective modulation vs. bismuth shields ...... 50 5.3 Segmented dental panoramic tomography ...... 53 5.4 Uncertainty of dose determinations ...... 57 6 Discussion 59 6.1 RPL and MOSFET in conventional radiology ...... 59 6.1.1 MOSFET’s energy dependence of the response ...... 59 6.1.2 MOSFET’s dose linearity and repeatability ...... 60 6.1.3 Absorbed and effective dose determination ...... 60 6.2 Organ effective modulation vs. bismuth shields ...... 62 6.3 Segmented dental panoramic tomography ...... 63 7 Conclusions 65 References 67 Original publications 73

16 1 Introduction

The effects of radiation can be categorised as tissue reactions (International Commission on Radiological Protection, 2011) and stochastic effects. Tissue reaction means that there is a harmful tissue reaction due to death or a serious malfunction of the cell population, which can be characterised by a threshold dose. As the dose is increased so too is the degree of reaction. The stochastic effect of radiation means that the probability of malignant disease and heritable effects are regarded as a function of dose without a threshold level (ICRP, 2007). It should be noted that in diagnostic radiology, the tissue reactions almost never occur when the examinations and procedures are done appropriately. The linear no- threshold hypothesis that is used to estimate the stochastic effect has received criticism over the years, especially at dose levels below 100 mSv (O’Connor, 2017; Sacks, Meyerson, & Siegel, 2016), but the importance of protocol optimisation must not be overlooked. Despite the negative effects of radiation, x-rays can provide enormous benefit when used properly. The use of radiation for medical purposes needs to follow three principles: justification, optimisation and dose limits. According to justification, the benefit gained from the examination needs to be greater than the possible harm. With optimisation, the ALARA principle specifies that the dose used for medical purposes needs to be kept as low as reasonably achievable. Dose limits are set to protect the public and workers employing radiation (ICRP, 2007). To comply with ALARA in diagnostics, the optimisation of imaging devices is required. As the radiation dose and image quality are linked to each other in x-ray imaging, a fine balance between them is needed. If the radiation dose is lowered excessively, the image quality may be degraded to a point where a diagnosis cannot be made. Therefore, the optimisation process involves close teamwork between the radiologist, radiographer and medical physicist. Just as the radiation dose forms a critical part of optimisation, radiation dose measurements are one method in optimisation processes. With the help of anthropomorphic phantoms and radiation dosimeters, different imaging methods can be compared and thus optimal imaging protocol may be found. This type of measurement-based optimisation can occur when, e.g. new equipment is purchased or new software is installed. X-ray energies used in radiology can vary from roughly 10 keV to 150 keV, which poses its own challenge for dose measurements. A suitable measurement instrument depends on the intended applications, and different properties need to

17 be reviewed before use, such as radiation quality, linearity, efficiency, angular response and repeatability of the measurements. Ease of use can also play a role when selecting the instrument, as some dosimeters allow for an instant read-out of the measured quantity. As different dose parameters can be measured in radiology, e.g. skin dose, point dose, air kerma and air kerma length, the dosimeter types also vary. Some of the dosimeters require calibrations against an institution’s reference dosimeters, which in turn should be traceable back to the primary standards. (IAEA, 2007; Seco, Clasie, & Partridge, 2014)

18 2 Literature review

2.1 Radiation dose determinations

2.1.1 Radiation dose quantities

The absorbed dose is a physical quantity used in radiology and radiation biology and it is defined as follows:

∑ , (1) where d ̅ is the mean energy imparted to the examined volume, Rin is the radiation energy of all particles entering the volume and Rout is the radiation energy leaving the volume. Radiation energy consists of both charged and uncharged ionising particles. ∑Q is the sum of all changes in the rest energy of particles in the volume and dm is the mass of the volume. The unit of absorbed dose is J/kg, but it is usually referred to as gray (Gy) (IAEA, 2007). In dosimetric applications, KERMA (K, kinetic energy released per unit mass) is sometimes the preferred parameter to be used. The K is the sum of the initial energies of all charged particles (dEtr) liberated by uncharged particles in a mass dm:

. (2) KERMA has the same unit as the absorbed dose and they both have the same values when charged particle equilibrium is established. When the effects of radiation are considered, usually organ/tissue and effective doses are used. Organ/tissue dose (DT) is defined as the absorbed dose, but the volume and mass are of a given tissue (T):

. (3) To evaluate the stochastic effects of radiation, the International Commission on Radiological Protection (ICRP) has implemented the concepts of equivalent and effective dose. The equivalent dose (HT) considers the radiation type by applying a radiation weighting factor, wR, to the organ dose:

∑ , (4)

19 Different radiation weighting factors are shown in Table 1.

Table 1. Radiation weighting factors (ICRP, 2007)

Radiation type Radiation weighting factor Photons 1 Electrons and muons 1 Protons and charged pions 2 Alpha particles, fission fragments, heavy ions 20 Neutrons A continuous function of neutron energy

To evaluate the overall detriment of radiation to humans, the ICRP has introduced an effective dose in which the radiation sensitivities of different tissues are considered:

∑ , (5) where wT is the tissue weighting factor. Different tissue weighting factors are presented in Table 2.

Table 2. Tissue weighting factors based on ICRP 103 (ICRP, 2007)

Tissue wT ∑wT

Bone marrow (red), colon, lung, stomach, breast, remainder tissues1 0.12 0.72 Gonads 0.08 0.08

Bladder, oesophagus, liver, thyroid 0.04 0.16 Bone surface, brain, salivary glands, skin 0.01 0.04 Total 1 1 Remainder tissues: adrenals, extrathoracic region, gall bladder, heart, kidneys, lymphatic nodes, muscle, oral mucosa, pancreas, prostate, small intestine, spleen, thymus, uterus/cervix.

Although an effective dose is determined for evaluating the overall detriment of radiation, it has its limitations, e.g. an effective dose is not recommended for epidemiological evaluations, and therefore, organ doses are necessary if organ- specific risks for individuals or population are being estimated. (ICRP, 2007) Radiation doses from imaging devices are usually expressed as technical parameters. In conventional x-ray and dental panoramic devices, the radiation dose is typically measured using an air kerma-area product (KAP) metre, which is the revised term for the dose area product (DAP) ionisation chambers. KAP’s unit is

20 Gym2, but different prefixed units are also used. (Dowsett, Kenny, & Johnston, 2006; Lin et al., 2015) Incident air KERMA can be calculated from conventional x-ray images as follows:

, (6) where Y is the x-ray tube’s radiation output (mGy/mAs), FDD is the focus-to- detector distance, FSD is the focus-to-skin distance and Q is the x-ray tube current- time product (mAs). The entrance surface dose can be calculated from the incident air KERMA by multiplying the result with an appropriate backscatter factor (BSF). Radiation dose parameters from the computed tomography examination are presented as the computed tomography dose index (CTDI) and dose length product (DLP). The CTDI has different subscripts due to different definitions:

, (7) where D(z) is the dose profile along line z and the dose is reported as absorbed dose to air, N is the number of tomographic sections produces in one axial scan and T is the nominal section thickness selected (IEC, 2002):

, (8) where CTDI100(centre) is the CTDI100 values measured along the axis of rotation of the polymethyl methacrylate (PMMA) dosimetry phantom being used and

CTDI100(peripheral) is the average of four values of CTDI100 measured in four positions placed 10 mm interior to the surface of the PMMA phantom. (IEC, 2002)

Modern CT devices usually express CTDI values as volume CTDIw (CTDIvol), which has different definitions depending on the scanning type. For helical scanning, the CT pitch factor is required for CTDIvol calculations:

∆ , (9) where Δd is the distance patient support travels per rotation of the x-ray source and

CTDIvol is then defined as follows (IEC, 2002):

. (10)

21 The DLP from the CT examination can be defined as follows:

, (11) where L is the effective scan length (IAEA TECREPO-5, 2011).

2.1.2 Organ dose determinations using phantoms

Based on the multiple, different metrics used to present the radiation dose to a patient, the organ dose can generally be said to be one of the best (Samei, Tian, & Segars, 2014). Accurate determination of the organ dose requires an accurate knowledge of patient anatomy and the irradiation field (Samei et al., 2014). As organ dose measurements in vivo are generally unavailable, Monte Carlo-based simulations using human models or anthropomorphic phantoms are usually used (IAEA, 2007). Antropomorphic phantoms are constructed from tissue-equivalent materials and are physical representations of human anatomy. As the materials are tissue- equivalent, the radiation interactions in the phantom are nearly similar to those in humans, and they can therefore be used for dosimetric applications if there are cavities for dosimeters (Winslow, Hyer, Fisher, Tien, & Hintenlang, 2009). The number of dosimeters in a tissue of interest is usually limited for practical reasons. Thus, the organ dose is usually estimated by using the average from point doses or point-to-organ dose scaling factors (Brisse et al., 2009; Sessions et al., 2002). Dosimeters used for organ dose determinations are usually cross-calibrated against the reference dosimeter reading dose in air (Dair). To determine the energy deposited to the tissue, the mass energy absorption coefficients (µen/ρ) of the dosimeter tissue (T) and air need to be considered (Hendee & Ritenour, 2003):

∗ . (12)

2.1.3 Organ dose determinations using simulations

In diagnostic radiology, organ doses can also be obtained using Monte Carlo model simulations. The models include the radiation field (x-ray field size, direction, spectrum), photon transportation and the human body. For diagnostic radiology simulations, approximations can be made regarding the energy depositions to the point of interactions so that all the energy is locally absorbed (KERMA simulation).

22 The energy departed to the red bone marrow provides the exception because the secondary electrons’ range is similar to the cavities in the marrow. The human body modelling is based on mathematical (geometrical) or voxel phantoms. Mathematical phantoms are based on combinations of different geometric shapes. Voxel phantoms are based on the anatomy of individuals, usually obtained from CT or magnetic resonance images. (IAEA, 2007) The uncertainty of the results from the Monte Carlo simulations is due to the statistical nature of the simulations and systematic uncertainties. Generally, the doses in the x-ray field have less uncertainty than the doses outside the field, and the uncertainty of the results increases as the distance from the edge of the x-ray field increases and the number of transported quanta decreases. (IAEA, 2007)

2.2 Imaging devices

2.2.1 Conventional x-ray and dental panoramic devices

Conventional x-ray devices are used to take one or more projection images from the patient. The device consists of an x-ray tube, detector and any necessary electronics and mechanics. Radiation detected with the detector is usually digitally sampled, filtered, processed and windowed to have the image with the necessary details for the radiologist. (Dowsett et al., 2006) Panoramic devices use a narrow x-ray beam, which rotates horizontally around the patient’s skull base. The x-ray tube is usually behind the patient and the detector is situated in the front. As the rotation centre is located between the x-ray tube and the detector, some geometric distortion is always present in the images. The cervical spine can produce shadowing on the panoramic images if constant kVp is used. Therefore, some manufacturers use higher kVp in the spinal area to produce constant image quality. Panoramic imaging is typically used to obtain overall coverage of the dental arches and associated structures. Some dental panoramic devices enable acquisitions of only certain segments from the dental arches. (Davis, Safi, & Maddison, 2015; Langland, Langlais, & Preece, 2002)

23 2.2.2 Computed tomography

Computed tomography is based on a rotating gantry structure that includes an x- ray tube and detector inside the gantry. Multiple x-ray projections pass through the patient and attenuated x-ray projections are measured. The measured intensity projections are then converted into attenuation profiles via log transform and pre- processing. The patient is usually scanned by moving the patient table continuously (helical scan mode) or by a step-and-shoot method (axial scan mode). Before the actual scan, a survey radiograph is taken to obtain the patient’s net attenuation profile. The attenuation profile can be used to modulate the tube current along the longitudinal axis and as a function of the rotation angle. Additionally, some manufacturers have implemented an additional tube current modulation technique for radiation sensitive surface organs (Akai et al., 2016; Dixon, Loader, Stevens, & Rowles, 2016; Duan et al., 2011; Fu et al., 2017; Gandhi, Crotty, Stevens, & Schmidt, 2015; Hoang et al., 2012; Lambert & Gould, 2016; Wang et al., 2012, 2011). Organ-based tube current modulation (OBTCM) techniques are nowadays available from nearly every major CT manufacturer. The purpose of OBTCM techniques is to lower the radiation dose to radiation-sensitive organs that are located near the anterior surface. Lenses, thyroid and breasts all are located along the same anterior side, thus the tube current can be lowered only along that side regardless of the patient’s orientation. Traditionally, the radiation doses to the aforementioned organs have been reduced by placing an additional shield, usually made of bismuth or barium, on the skin’s surface at the position of the organ. Additional shields reduce the intensity of the x-ray beam passing through it, thus reducing the absorbed dose to the organ next to it. Shields also affect the x-ray spectrum, depending on the shield’s material. (McCaffrey, Shen, Downton, & Mainegra-Hing, 2007) As the x-ray tube rotates around the patient, the shield also reduces the intensity of the x-rays that have already passed through the patient, thus only reducing the image quality due to increased noise and artifacts in the reconstructed image. A survey radiograph can also be used to select the appropriate peak kilovoltage for the patient for that particular acquisition. The survey radiograph provides a patient’s attenuation profile. Furthermore, providing the acquisition’s diagnostic task, some CT devices can automatically choose the kVp level and mA profile used. As the contrast and noise change when kVp changes, the mA also needs to be adjusted. (Li, Feng, Wu, & Zhang, 2017; Siemens, 2014)

24 The geometry of the radiation beam depends on the width of the detector, which is usually from 2 to 16 cm in modern CTs. Additional patient dose reduction methods may include dynamic collimation, different bow-tie filters and different post-processing methods. (Kalender, 2011; Zacharias et al., 2013) The traditional type of image reconstruction is based on a filtered back projection (FBP). Since projections are taken from multiple angles, the linear attenuation coefficients of every voxel can be solved as an inverse problem and three-dimensional images can be obtained. Filtered back projections use reconstruction kernels to set the balance between the resolution and noise in order to modify each projection profile to obtain a suitable image for the specific purpose, e.g. soft tissue or bone images. Modern CTs usually use hybrid or model-based iterative reconstruction techniques. Iterative techniques are used to improve image quality mainly by lowering the noise level and appearance of artefacts without degrading the spatial resolution, depending on the manufacturer’s algorithm. (Geyer et al., 2015; Kalender, 2011) The calculated CT images are a map of Hounsfield Units, which can be calculated as follows:

, (13) where µi is the linear attenuation coefficient of the voxel of interest. The linear attenuation coefficient of air (µair) is sometimes excluded due to its negligible effect. In CT images, the pixel noise is determined as a standard deviation of the pixel (HU) values from a certain number of pixels within a region of interest (ROI).

2.3 Dosimeters

2.3.1 MOSFET dosimeters

Metal-oxide-semiconductor-field-effect-transistor (MOSFET) dosimeters are active dosimeters, which means that the radiation dose can be viewed immediately after the radiation event (Huang & Hsu, 2011). MOSFETs were originally intended for measuring relatively high doses in the industrial field, radiobiology and radiation therapy (Thomson, Thomas, & Berndt, 1983). Since the late 1990s, MOSFET dosimeters have been evaluated as suitable for diagnostic radiology applications (Bower & Hintenlang, 1998; Peet & Pryor, 1999). In this thesis, the

25 focus was on MOSFETs that are designed for a diagnostic radiology application, specifically high-sensitivity MOSFETs. With MOSFET dosimeters, a high voltage is applied to the polysilicon gate, which causes holes from surrounding structures to move into the oxide layer and silicon substrate adjacent to it. When the number of holes is large enough, a current between the source and drain is possible when a threshold voltage is applied. As radiation passes through the oxide layer, electron-hole pairs are formed. Holes travel to the boundary between the oxide and silicon substrate, where they become trapped. As the number of holes in the boundary between the oxide and silicon substrate increases, so too does the threshold voltage needed to create the current between the source and drain. Thus, the change in the threshold voltage is dependent on the radiation dose applied to the dosimeter. The applied threshold voltage sets a finite lifetime on the MOSFET dosimeters, which in turn adds costs when they need to be renewed. A schematic structure of a MOSFET is show in Figure 1. (Ahmed, 2014)

Fig. 1. Schematic structure of a MOSFET. Modified from (Ahmed, 2014).

When a MOSFET is used as a dosimeter, the measurement chain needs to be calibrated. The calibration set-up depends on the actual measurement set-up, but a suitable reference ionisation chamber or dosimeter with a traceable calibration in a standard laboratory is needed (Zoetelief, Julius, & Christensen, 2000). The calibration factor for the dosimeter is determined as follows:

, (14)

26 where the known radiation dose is measured with a reference dosimeter (Operator’s manual for the mobile MOSFET system, 2010). MOSFETs sensitivity (mV/mGy) has been reported to worsen over time, thus recalibration is recommended after every 3000 mV increase in the threshold voltage (Brady & Kaufman, 2012). The applicability of MOSFETs has been studied for various radiology applications, such as mammography (Benevides & Hintenlang, 2006), the entrance surface dose in diagnostic radiology (Peet & Pryor, 1999), dental CBCT (J. H. Koivisto, Wolff, Kiljunen, Schulze, & Kortesniemi, 2015) and CT (Yoshizumi et al., 2007). As the structure of MOSFETs is not homogenous, there is variation in the dosimeter’s response depending on the measurement set-up being used. The asymmetry of the dosimeter causes variation in the response as a function of the irradiation angle. The angular response of high-sensitivity MOSFETs has been reported from nearly independent of the irradiation angle (Bower & Hintenlang, 1998) to reductions of up to 63% in free-in-air measurements and up to 82% in PMMA measurements (J. Koivisto, Kiljunen, Wolff, & Kortesniemi, 2013). Certain irradiation angles have also been reported that show notable differences in sensitivity (Benevides & Hintenlang, 2006; Dong, Chu, Lan, et al., 2002). The energy dependence of dosimeters used in radiology should be carefully assessed, as the energy spectrum can vary from roughly 10 keV to 150 keV, where photoelectric and Compton effects dominate (Dowsett et al., 2006). The energy dependence of high-sensitivity MOSFETs has been reported to be both negligible (Bower & Hintenlang, 1998; J. H. Koivisto et al., 2015) and notable (Benevides & Hintenlang, 2006; Dong, Chu, Lan, et al., 2002; Peet & Pryor, 1999), depending on the measurement set-up. Since the results of energy dependence vary, the usage of MOSFETs needs to be reviewed depending on the application of intended use. Since the dose range can vary notably in different set-ups as a function of depth or due to some dosimeters measuring doses at the primary beam and others measuring only the scattered radiations, a good linearity of dosimeter response is required. It has been reported that high-sensitivity dosimeters have good linearity, but with limited reproducibility at a low dose range. Poor reproducibility at low doses sets some limits on the use of MOSFETs in diagnostic radiology applications, as overall uncertainties with a 95% confidence interval have been reported to be as high as 25% at a dose level of 1.7 mGy with a single exposure. (Bower & Hintenlang, 1998; Brady & Kaufman, 2012; J. H. Koivisto et al., 2015; Peet & Pryor, 1999; Yoshizumi et al., 2007)

27 2.3.2 Radiophotoluminescence dosimeters

Radiophotoluminescence dosimeters (RPL) can be considered passive dosimeters as the interaction of radiation with glass dosimeters needs separate read-out electronics to obtain the absorbed energy. RPL dosimeters are formed by silver- activated phosphate glass. When radiation interacts with luminescent material, so- called colour centres are formed, which are electron-hole pairs. Phosphate loses an electron from the valence band when exposed to the radiation, and a hole trap is formed: → . (15) The excited electron then combines with the silver to form an electron trap: →° . (16) The formation of the electron and hole traps are shown in Figure 2.

Fig. 2. Formation of electron and hole traps. Modified from (Huang & Hsu, 2011).

RPL dosimeters require heat treatment before read-out to stabilise the colour centres (Hsu et al., 2007). When the colour centres are excited with ultra-violet light, the electrons in the electron traps are excited to higher energy levels. As electrons then descend back to the traps, 600-700 nm visible light is emitted, and the emission is detected in the photomultiplier tube. The emission is proportional to the dose absorbed in the RPL glass. (Perry, 1987) RPL dosimeters need to be annealed in 400 °C for one hour so that electrons can gain enough energy to return to the valence band (Huang & Hsu, 2011). The excitation and read-out process of the RPL dosimeters is presented in Figure 3. RPL dosimeters can be used in radiation therapy and diagnostic radiology applications. Additional tin filter can be

28 used with RPL dosimeters to compensate for the energy dependence of RPL materials when they are used to measure doses along the diagnostic radiology energy spectrum. (Perry, 1987)

Fig. 3. Schematic presentation of radiophotoluminescence dosimeter excitation and read-out process. The first phase (1) represents the formation of colour centres after radiation interacts with luminescence material. When a RPL dosimeter is placed in the read-out device (2), an ultra-violet (UV) laser is used to excite the electrons from colour centres to an excited state. After the UV excitation, electrons return to the colour centres (3) and visible light is emitted. After the annealing process, the electrons will return to the valence band from the colour centres (4). Modified from (Huang & Hsu, 2011).

Studies have shown that RPL dosimeters are suitable for radiation dose measurements in radiological applications when a tin filter is used. The energy dependence of dosimeters was within a 4% mean error when using a conventional x-ray setting at peak tube voltages of 60 to 125 kVp (Manninen, Koivula, & Nieminen, 2012). The dose linearity of the dosimeters is reliable with R2 > 0.99 at a dose range of 20 µGy to 11 mGy (Manninen et al., 2012). The vertical angular dependence of a dosimeter with a tin filter has been shown to be less than 5% with vertical angles between RPL and x-rays at 0 to 40 degrees, but if the angle is over 50 degrees, then the dose response can vary between -38% and 50% (Manninen et al., 2012). Horizontal angular dependence has been shown to be almost negligible (Hsu et al., 2007). The coefficient of variation (CV) at 20 µGy is within 12%, while

29 at 1 to 11mGy, the CV is 5% on average. It has been reported that the energy dependency of RPL dosimeter is less than 3% with an energy range of 30 to 662 keV, with a higher dependence at lower keVs. (Hsu et al., 2007; Manninen et al., 2012)

30 3 Purpose of the study

The aim of this thesis was to investigate the applicability of MOSFET dosimeters for organ dose determinations and to investigate different dose reduction methods in diagnostic and dental radiology using MOSFETs and RPLs. The specific aims of this study were to: 1. evaluate MOSFET dosimeters for absorbed and effective dose determinations in a conventional chest x-ray examination against RPL and MC results; 2. compare bismuth shielding and OBTCM against a reference method in a helical chest CT examination using MOSFET as a dosimetry method; 3. determine radiation doses from segmented and full dental panoramic procedures in pediatric and adult protocols using RPL dosimeters.

32 4 Materials and methods

This thesis consists of three studies (I–III) in which three different imaging modalities and two different dosimetry systems were used. The main materials and modalities used are summarised in Table 3.

Table 3. Summary of materials used in studies I–III

Study Modality Dosimetry system Phantoms for organ dose determination I Conventional x-ray RPL, MOSFET and MC ATOM 701 paraffin breast attachments II Computed tomography MOSFET ATOM 701 with ATOM 701-BR-02R and 701-BR-02L breast attachments III Dental panoramic RPL ATOM 701 and ATOM 705

4.1 Materials

4.1.1 Dose determinations

The radiation dose measurements in this thesis were performed using RPL and MOSFET dosimeters. A MOSFET dosimetry system (TN-RD-70-W, Best Medical Canada, Ottawa, Canada) together with high-sensitivity MOSFET dosimeters (TN- 1002RD) were used with the same manufacturer’s read-out modules (TN-RD-16) and wireless transceiver (TN-RD-48) for dose determinations in Studies I and II. The dosimetry system was connected to a laptop, where voltage changes could be converted into doses using conversion factors, as presented in equation (14). The reader module’s bias sensitivity was set at a high sensitivity level for all measurements. The high-sensitivity setting corresponds to 3.42 mV mGy-1. The used MOSFET dosimeters have a hemispherical drop-shaped epoxy bubble at the end of the dosimeter in which the active region of 0.2 x 0.2 mm is located (Figure 4).

33 Fig. 4. Structure of a TN-1002RD MOSFET. The photograph on the left and an image- processed visualisation from µCT scan (Skyscan 1176, Bruker microCT, Kontich, Belgium) on the right. The active area of the MOSFET shown in Figure 1 is marked with an arrow on the right side.

In the first study, MOSFET dosimeters were cross-calibrated against Unfors Xi R/F detector (Unfors Instruments AB, Billdal, Sweden) using a Calibration Jig TN-RD- 57-30 (Best Medical Canada, Ottawa, Canada). Calibration verifications of RPL dosimeters were previously done in the clinic against a Radcal (model 90150, Radcal Corporation, Monrovia, CA, USA) dosimeter (Manninen et al., 2012). A glass dosimeter with a tin filter (GDM-352M, Asahi Techno Glass Corporation, Chiba, Japan) was used for every RPL dose determination (Study I and III). A tin filter was used for energy compensation (Figure 5). A Doce Ace FGD- 1000 (Asahi Techno Glass Corporation, Chiba, Japan) was used for the RPL glass dosimeter read-out. Calibration of the read-out device was done according to manufacturer’s instructions using a standard calibration glass GDS-352A (Asahi Techno Glass Corporation, Chiba, Japan). For this thesis, Monte Carlo (MC) simulations were used to compare effective doses from Study I to the doses obtained via the simulations. The simulations were performed using PCXMC Dose Calculations software version 2.0, which uses mathematical phantoms (Tapiovaara & Siiskonen, 2008). The phantom used for the PCXMC simulation was an adult hermaphrodite without hands. The phantom resembles an adult with a height of 173 cm and a mass of 73 kg.

34 Fig. 5. Radiophotoluminescence dosimeter on the bottom and tin capsule on the top.

In the second study, MOSFETs were calibrated against a Piranha 657 with a CT Dose Profiler (RTI Electronics, Mölndal, Sweden) inside a custom PMMA phantom, as described by Brady and Kaufman (Brady & Kaufman, 2012).

4.1.2 Phantoms

An anthropomorphic adult male phantom (ATOM male phantom model 701-C, CIRS, Norfolk, Virginia, USA) was used in every study. Additional custom-made paraffin breast attachments were used in the first study to help determine the breast dose. The ATOM 701 phantom with paraffin breast attachments is shown in Figure 6.

Fig. 6. ATOM 701 phantom with breast attachments and MOSFET dosimeters. Photograph taken above the phantom while the phantom and dosimeters are on the table.

35 In the second study, a D-cup breast attachments (ATOM 701-BR-02R and 701-BR- 02L, Norfolk, Virginia, USA) were used for breast dose determinations. The breast attachments were modified for more suitable dose determinations by drilling holes for the dosimeters. The holes were 3.5 cm in length and 5 mm in diameter. Large 1 mm thick bismuth breast and thyroid shields (AttenuRad, F&L Medical Products) were used in the bismuth scans. Foam padding that was 1.9 cm thick was used between the breast attachments and bismuth breast shields. A five-year-old anthropomorphic phantom (ATOM 705-D, CIRS, Norfolk, Virginia, USA) was used for pediatric dose determinations in Study III (Figure 7).

Fig. 7. Five-year old anthropomorphic phantom used in pediatric dose determinations in Study III. Photograph taken above the phantom while the phantom is inside the suitcase lying on the floor.

4.1.3 Imaging equipment

Calibrations and validation of the MOSFETs for conventional x-ray dose determinations were performed using a Philips Optimus 50 (Philips Medical Systems, Eindhoven, the Netherlands). The radiation dose measurements were taken with a Fuji D-Evo (Fujifilm Medical System, Tokyo, Japan) (Study I). For the second study, every scan was performed using the same multi-detector row CT (Toshiba Aquilon One Vision Edition, Toshiba Medical Systems, Otawara, Japan) (Study II). Dental panoramic imaging was performed using an Orthopantomograph OP 200D (Instrumentarium Dental, Tuusula, Finland) (Study III). Every x-ray device is subject to an annual quality control inspection in which the proper functioning of the x-ray tube, detector, moving parts and image formation are assessed. The computed tomography device is inspected biannually

36 with respect to image quality and the displayed dose levels. Used devices also have daily and weekly quality inspections performed by radiographers.

4.2 Methods

4.2.1 RPL and MOSFET in conventional radiology

The calibrations and evaluation measurements for the MOSFETs were performed using the MOSFET’s calibration plate by placing each MOSFET in its individual slot in the plate. The evaluation measurements consisted of energy response, dose linearity and repeatability measurements. Reference dosimeter measurements were done on top of a lead apron using the same FDD of 100 cm and a field size of 10 cm x 10 cm. The lead apron was used to remove the backscatter from the table on which the calibration plate was placed. Calibrations were done by placing five dosimeters at a time on the calibration plate. Energy response was evaluated using three different dosimeters, which were placed at the centre position individually and irradiated three times at each kVp level (40, 50, 60, 70, 81, 90, 96, 109, 110 and 125). The mAs at each kVp level was set so that the absorbed dose to the dosimeter would be approximately 5 mGy. The dose linearity of the MOSFETs was evaluated by measuring the absorbed dose by three MOSFETs at seven different mAs levels (absorbed dose with reference dosimeter in mGy): 2.5 (0.14), 5 (0.29), 10 (0.58), 25 (1.45), 50 (2.90), 100 (5.79) and 200 (11.58). The peak kV of 81 was used just as with the calibrations. The repeatability of the MOSFETs was evaluated by irradiating five dosimeters at known dose levels between 0.21 and 10.29 mGy. The root-mean-square average coefficient of variation (CVRMS) was then calculated for each dose level according to equation (17). The absorbed dose and effective dose determinations from a conventional thorax examination were done using 39 and 80 RPLs and 39 MOSFETs. Different numbers of RPLs were used to demonstrate the measurement accuracy of a limited number of dosimeters. The dosimeters were positioned according to recommendations by (Scalzetti, Huda, Bhatt, & Ogden, 2008). The mAs values for different projections were set so that the ESD would be approximately 0.5 mGy, with a BSF of 1.41, according to equation (6). To reduce statistical error, the MOSFET measurements were repeated ten times and the RPL measurements were repeated four times. An energy correction factor was used for the MOSFET results

37 to compensate for the energy dependence of the dosimeters. For this thesis, equivalent doses from dose determinations were corrected using the mass energy absorption coefficients according to equation (12). Values for the mass energy absorption coefficients of muscle, bone, brain, eye lens, soft tissue and air based on the calculations were obtained at 60 keV from the National Institute of Standards and Technology database (Hubbell & Seltzer, 2004) because the mean energy of the x-ray spectrum was determined to be 66 keV by x-ray spectrum software Spektripaja (Tapiovaara & Tapiovaara, 2007). Monte Carlo simulations were performed using PCXMC 2.0 (Tapiovaara & Siiskonen, 2008). Parameters for the MC simulations were set as close to the RPL and MOSFET measurement as possible using the parameters shown in Table 4. Phantom height and mass were set at 173 cm and 73 kg, respectively, and arms were not included in the simulations. Every organ was included in the simulations. Equivalent and effective doses from the simulations were determined for AP, PA and LAT projections.

Table 4. Parameters used in Monte Carlo simulations (only changes in the lateral projection are shown)

Parameter AP and PA1 LAT2

Focus-skin-distance (cm) 200 Image width (cm) 35 24

Image height (cm) 34 Phantom exit-image distance (cm) 5.7 0.0 Xref 0.0

Yref 10.0 0.0 Zref 51.0 Max energy (keV) 130

Number of photons 20 000 1 Anterior-posterior and posterior-anterior projection. 2 Lateral projection.

4.2.2 Organ effective modulation vs. bismuth shields

In the second study, organ doses were determined from three different methods in a chest CT scan. An organ effective modulation (OEM), which is meant to reduce the surface dose, was used as the OBTCM method. OEM and bismuth shield scans were compared against the reference method. The reference scan used only

38 longitudinal and angular tube current modulation. Thirteen MOSFETs were used to measure the absorbed doses from different measurement locations (Figure 8). For image quality determination, HU and noise values were obtained from Digital Imaging and Communication in Medicine (DICOM) images using ImageJ (ImageJ 1.48v, National Institutes of Health, USA). The MOSFETs were removed from the phantom before the image quality scans due to the artifacts occurring from metal parts of the dosimeters. Metal artefact removal software was initially used with the intention of removing artifacts from the dosimeters, but as the bismuth shields caused additional artifacts, the MOSFETs had to be removed before doing the image quality scans. Rectangular 18 pixel x 18 pixel (13.86 mm x 18.86 mm) ROIs were used from three consecutive slices at the same location as in the dosimetry scans. The imaging and image reconstruction parameters used in Study II are listed in Tables 5 and 6.

Table 5. Imaging parameters used in Study II (table modified from Study II)

Parameter Value Scanning parameters Peak kilovoltage 100 Beam width/collimation (mm x detector rows) 0.5 * 80 Rotation time (s) 0.275 Pitch 0.813 Scan field of view (mm) Large (400) SureExposure3D parameters1 SD 11 SureIQ Body:STD. Axial Image thickness (mm) 5 Recon kernel FC18 Dose reduction AIDR3D STD2 1 Adaptive iterative dose reduction 3D-algorithm with Standard setting (Irwan, Nakanishi, & Blum, 2011). 2 Image quality parameters used in CT devices by Toshiba. SD defines the standard deviation of noise (Angel, 2012).

39 Table 6. Dose values and image reconstruction parameters used in Study II (table modified from Study II)

Parameter Reference and bismuth scans OEM scan1 Dose report on dosimetry scan 2 CTDIVOL (mGy) 6.2 6.2 DLP3 (mGycm) 210.1 208.4 Dose report on image quality scan

CTDIVOL 6.3 6.3 DLP 210.9 211.3

Image reconstructions for every scan Slice thickness 3mm 3mm Recon kernel FC08 (soft tissue) FC51 (lung) AIDR3D setting Enhanced Enhanced 1 Organ effective modulation scan. 2 Volume Computed Tomography Dose Index. 3 Dose Length Product.

Fig. 8. Organ dose measurement locations and mean tube currents of the different scans used in Study II (figure obtained from Study II).

40 4.2.3 Segmented dental panoramic tomography

In the third study, 4323 segmented dental panoramic tomography (sDPT) images were retrospectively studied under permission from the Oulu University Hospital review board (No. 20078/2016). Indications and the segments used for every image were reviewed for further analysis. Organ dose determinations for different sDPTs and full DPT protocols were performed using adult and pediatric phantoms and protocols. Absorbed doses were determined using RPL dosimeters. Every dose measurement was repeated ten times to accumulate a higher dose for the dosimeters. A fixed tube current of 10 mA was used for every measurement, and the measured pediatric doses were adjusted to the dose level that would have been obtained by using clinical protocols (6.3 mA). The organ doses obtained from the absorbed doses were evaluated by multiplying the measured doses by the mass energy absorption coefficient of the tissue at 60 keV, divided by the mass energy attenuation coefficient of air at 60 keV, according to the equation (12). Effective doses were determined from the organ doses, as shown in equation (5). Dosimeter locations in the adult phantom were based on a study by Granlund et al. (Granlund, Thilander-Klang, Ylhan, Lofthag-Hansen, & Ekestubbe, 2016); for the pediatric phantom, the locations are listed in Table 7. Different segments of DPT are presented in Figure 9. A full DPT consists of every segment, the temporomandibular joint (TMJ) (4,5), right TMJ (5), dentition (2,3,4), anterior region (3), half panoramic (3,4,5), molar region (2) and third molars (1,2,4,5).

Table 7. Number of dosimeters used per organ in measurement set-ups and dosimeter location in the pediatric phantom (table obtained from Study III)

Tissue Number of RPLs in Number of RPLs in 5 year Dosimeter locations as adult phantom old phantom presented in the 5 year phantom Thyroid 4 5* 22, 23, 24, 25, * Submandibular gland 6 Parotis 12 2 17, 19 Cerebellum 1 Pituitary 1 Brain/Cranium 2 5 8, 11, 12, 13, 14 Lens 2 2 * Oral mucosa † 1 18 * Dosimeters were taped in front of the lens and thyroid, † dose to the oral mucosa was estimated from six dosimeters shared with submandibular and parotid glands. The dosimetry locations of the adult phantom were similar to work done by Granlund et al., as cited in the article.

Fig. 9. Different segments in a segmented dental panoramic tomography (figure obtained from Study III).

4.3 Measurement uncertainties

Study I Repeatability of the MOSFETs was investigated by calculating the root-mean- square average coefficient of variation (CVRMS) as follows:

∑ , (17) where CVi is the coefficient of variation of dosimeter i and n is the number of dosimeters. CV was calculated as follows:

, (18) where σ is the standard deviation of the measured dose and µ is the mean of the doses.

Study II Uncertainty of dose measurements in Study II was evaluated as standard uncertainty (standard deviation of the mean). Since dosimeters were calibrated for

42 the used energy and the radiation angle varied throughout the examination, only the standard deviation of the dose measurements was used.

Study I and III For Studies I and III, the combined uncertainty of the RPL and MOSFET dosimeter measurements was calculated based on (Zoetelief et al., 2000):

. , (19)

where uD represents the standard uncertainty of the measured dose (Cmeas), energy (KE), angle (θ) and source (s).

Study III Uncertainty of the effective doses was the determined as follows:

∑ ∑ ∑ , (21) where wT is the tissue weighting factor, n is the number of dosimeters, fT is the fraction of tissue that was irradiated and σ(dt) is the combined uncertainty of type A and type B standard uncertainties (equation 19). For red bone marrow, bone surface, lymphatic nodes and muscle, the fT was based on work done by Davis et al. and Granlund et al. (Davis et al., 2015; Granlund et al., 2016). The standard uncertainty of the fraction of tissue irradiated σ(fT) was estimated as 25% based on work done by Koivisto et al. (J. Koivisto, Kiljunen, Kadesjö, Shi, & Wolff, 2015).

A simplification fraction of the irradiated tissue (fT) for other tissues located in the x-ray field was estimated as 100% with 0% uncertainty.

44 5 Results

5.1 RPL and MOSFET in conventional radiology

5.1.1 MOSFET’s energy dependence of the response

We evaluated the MOSFET’s dose response at different peak kilovoltages ranging between 40 kVp and 125 kVp. The dose response had a less than 10% error between 50 and 110 kVp. The dose response was more than 10% smaller than the reference dose at kVps under 40 and over 110 (Figure 10).

Fig. 10. Correction factors for the MOSFET dosimeter in typical diagnostic peak kilovoltages calculated against the reference doses. Error bars represent the coefficient of variation of three measurements (figure obtained from Study I).

5.1.2 MOSFET’s linearity of the response

The MOSFET’s dose correlated strongly with the reference dose. Pearson correlation coefficient was 0.999 at a dose range of 0.14 to 11.58mGy, as shown in Figure 11.

45 Fig. 11. Dose linearity of the MOSFET for doses between 0.14 and 11.58 mGy. The error bars represent the standard deviation of the dose from three dosimeters at each dose level (figure obtained from Study I).

5.1.3 MOSFET’s repeatability of the dosimeter response

We determined the repeatability of MOSFET’s response at dose levels of between

0.21 and 10.29 mGy. The mean CVRMS was between 49 and 5% (Table 8). The

CVRMS was below 10% at dose levels of 5.21 mGy and 10.29 mGy.

Table 8. Root-mean-square average coefficient of variation for applied doses in the repeatability measurements (table obtained from the Study 1).

Absorbed dose (mGy) Mean CVRMS 0.21 0.485 0.41 0.369 0.64 0.256 0.81 0.209 1.16 0.207 1.64 0.185 2.59 0.135 5.21 0.059 10.29 0.053

46 5.1.4 Absorbed and effective dose determination

We compared the mean equivalent doses in Study I using an AP projection radiograph of the phantom. Equivalent doses were mostly similar for the different groups, as shown in Table 9. The greatest relative differences in equivalent doses between MOSFETs and RPLs were in the active bone marrow, thyroid, bone surface, brain and urinary bladder (Table 9). The results for doses determined using 39 and 80 RPL dosimeters were similar. The dose for the brain was the only result with a relative difference over 15% when comparing the two methods. The brain’s equivalent dose was 43% higher with 39 dosimeters. The mass energy absorption coefficient-corrected equivalent and effective doses from the MOSFET and RPL dose determinations are presented in Tables 8, 9 and 10 along with the results from the MC simulations. Active bone marrow, lung and remainder organs were the only organ doses derived from the measurements that were less than 50% different from the doses obtained from the MC simulations in every set-up. The greatest differences between the dosimeter and MC results were in the equivalent doses for the brain and urinary bladder, where the relative difference varied between 548% and 8298% in the AP, PA and LAT projections. Equivalent doses from almost every organ location were consistent for each measurement method (MOSFET, RPL39 and RPL80) in the AP radiographs. The equivalent dose for the bone surface showed the greatest difference between the

MOSFET and RPL80, with a 0.16 mSv difference. When comparing the equivalent doses from the measurements with the MC results, the difference was notable. Oesophagus, breasts, lungs, liver and remainder tissues were the only organ doses within a 30% difference when comparing the different methods. The effective doses from the AP projection were 0.17 mSv (MOSFET), 0.17 mSv (RPL39), 0.16 mSv

(RPL80) and 0.14 mSv (MC). The determined effective doses from every set-up in the AP projection were below a 15% difference.

47 Table 9. Equivalent and effective doses from the MOSFET, RPL and MC dose determinations in the AP projection (the difference between the dose determinations derived from the measurements and the MC simulations, in percentages, are presented in parentheses)

1 1 Organ MOSFET (mSv) RPLD39 (mSv) RPLD80 (mSv) MC (mSv) Active bone marrow 0.06 (-14%) 0.09 (32%) 0.09 (39%) 0.07 Thyroid 0.11 (93%) 0.17 (197%) 0.16 (180%) 0.06 Oesophagus 0.21 (29%) 0.21 (29%) 0.21 (29%) 0.16 Skin 0.08 (64%) 0.09 (89%) 0.09 (89%) 0.05 Bone surface 0.26 (203%) 0.4 (362%) 0.42 (387%) 0.09 Salivary glands 0.02 (111%) 0.02 (129%) 0.02 (162%) 0.01 Breasts 0.49 (9%) 0.46 (2%) 0.42 (-6%) 0.45 Large intestine (colon) 0.04 (642%) 0.04 (650%) 0.03 (557%) 0.00 Lung 0.23 (-17%) 0.20 (-26%) 0.20 (-28%) 0.27 Stomach 0.25 (67%) 0.26 (74%) 0.26 (73%) 0.15 Liver 0.19 (15%) 0.18 (6%) 0.16 (-7%) 0.17 Brain 0.02 (1469%) 0.01 (297%) <0.00 (176%) <0.00 Urinary bladder 0.02 (2684%) <0.00 (346%) <0.00 (346%) <0.00 Gonads 0.01 (458%) 0.01 (322%) 0.01 (322%) <0.00 Remainder 0.11 (-5%) 0.11 (-7%) 0.11 (-8%) 0.11

Effective dose 0.17 (15%) 0.17 (15%) 0.16 (11%) 0.14 1 Radiophotoluminescence dosimeter. 2 Monte Carlo.

The equivalent doses obtained from the dosimeters at the PA projection radiograph were all within 0.04 mSv when compared at the same measurement locations (Table 10). The effective doses were all 0.09 mSv, while the MC simulations yielded 0.11 mSv. Equivalent doses for the breasts, lungs, stomach, liver and remainder tissues were all within a 30% difference.

48 Table 10. Equivalent and effective doses from the MOSFET, RPL and MC dose determinations in the PA projection (the difference between the dose determinations from the measurements and the MC simulations, in percentages, are presented in parentheses)

1 2 Organ MOSFET (mSv) RPLD39 (mSv) RPLD80 (mSv) MC (mSv) Active bone marrow 0.07 (-45%) 0.08 (-38%) 0.07 (-43%) 0.12 Thyroid 0.08 (110%) 0.06 (75%) 0.06 (57%) 0.04 Oesophagus 0.10 (-54%) 0.12 (-44%) 0.12 (-44%) 0.22 Skin 0.09 (83%) 0.10 (122%) 0.10 (122%) 0.05 Bone surface 0.31 (96%) 0.35 (120%) 0.32 (102%) 0.16 Salivary glands 0.02 (190%) 0.01 (73%) 0.01 (71%) 0.01 Breasts 0.09 (-6%) 0.10 (3%) 0.07 (-26%) 0.10 Large intestine 0.03 (406%) 0.02 (243%) 0.01 (152%) 0.01 (colon) Lung 0.24 (-29%) 0.25 (-27%) 0.25 (-28%) 0.34 Stomach 0.08 (15%) 0.07 (2%) 0.09 (22%) 0.07 Liver 0.13 (-2%) 0.10 (-26%) 0.12 (-7%) 0.13 Brain 0.01 (548%) <0.00 (72%) <0.00 (84%) 0.00 Urinary bladder 0.01 (2250%) <0.00 (-27%) <0.00 (-27%) 0.00 Gonads 0.03 (801%) 0.05 (1500%) 0.05 (1500%) 0.00 Remainder 0.08 (-15%) 0.07 (-27%) 0.08 (-19%) 0.09

Effective dose 0.09 (-15%) 0.09 (-15%) 0.09 (-17%) 0.11 1 Radiophotoluminescence dosimeter. 2 Monte Carlo.

Thyroid’s equivalent dose was the only location from the MOSFET and RPL dose determinations in the LAT projection that had more than a 0.05 mGy difference when comparing the different set-ups (Table 11). The greatest differences between the dosimeter and MC results were with the thyroid dose, where the RPL39 showed a 0.18 mGy larger dose. The salivary glands equivalent dose determined with the dosimeters was generally higher than the equivalent dose determined via MC simulations. The effective doses obtained with different methods were consistent:

0.06 mSv (MOSFET), 0.07 mSv (RPL39), 0.06 mSv (RPL80) and 0.07 mSv (MC).

49 Table 11. Equivalent and effective doses from the MOSFET, RPL and MC dose determinations in the LAT projection (the difference between dose determinations from the measurements and the MC simulations, in percentages, are presented in parentheses)

1 2 Organ MOSFET (mSv) RPLD39 (mSv) RPLD80 (mSv) MC (mSv) Active bone marrow 0.03 (-44%) 0.05 (-15%) 0.04 (-26%) 0.05 Thyroid 0.04 (119%) 0.20 (1056%) 0.09 (442%) 0.02 Oesophagus 0.09 (-21%) 0.14 (18%) 0.14 (18%) 0.12 Skin 0.08 (122%) 0.01 (-71%) <0.00 (-89%) 0.04 Bone surface 0.14 (91%) 0.21 (189%) 0.19 (151%) 0.07 Salivary glands 0.01 (96%) 0.01 (147%) 0.02 (267%) <0.00 Breasts 0.05 (-60%) 0.07 (-43%) 0.06 (-51%) 0.13 Large intestine 0.02 (710%) 0.01 (379%) 0.01 (352%) <0.00 (colon) Lung 0.12 (-29%) 0.13 (-20%) 0.11 (-34%) 0.17 Stomach 0.11 (22%) 0.11 (17%) 0.13 (46%) 0.09 Liver 0.06 (237%) 0.05 (162%) 0.04 (114%) 0.02 Brain 0.02 (2032%) <0.00 (300%) <0.00 (158%) <0.00 Urinary bladder 0.02 (8298%) 0 (-100%) 0 (-100%) <0.00 Gonads <0.00 (273%) 0 (-100%) 0 (-100%) <0.00 Remainder 0.06 (4%) 0.06 (11%) 0.06 (14%) 0.05

Effective dose 0.06 (-14%) 0.07 (4%) 0.06 (-5%) 0.07 1 Radiophotoluminescence dosimeter. 2 Monte Carlo.

5.2 Organ effective modulation vs. bismuth shields

The use of a bismuth in the chest CT examination provided the greatest dose reduction at every measurement location (Table 12). The dose reduction at different measurement locations compared to the reference dose was between 12 and 39% in the bismuth scan. For the OEM scan, the dose reduction was between 1 and 17%. In the OEM set-up, seven measurement locations showed dose reductions greater than 5% and only two locations showed a dose reduction greater than 10%. For the breasts, the dose reduction with OEM was 6% (right breast) and 9% (left breast), while with the bismuth shields it was 37% and 38%, respectively.

50 Table 12. Absorbed organ doses (mGy) ± 1 SD in a helical chest CT scan series, with the difference between used dose reduction method and the reference scan shown as a percentage in parentheses (table obtained from Study II)

Measurement point Reference OEM1 Bismuth Thyroid (1) 2.4 ± 0.3 2.3 ± 0.1 (-4.3) 1.5 ± 0.1 (-39.2) Sternum (3) 3.5 ± 0.4 2.9 ± 0.1 (-16.7) 2.9 ± 0.4 (-18.9) Liver (13) 3.8 ± 0.4 3.3 ± 0.1 (-12.5) 3.1 ± 0.2 (-19.1) Thoracic spine (9) 2.7 ± 0.2 2.6 ± 0.2 (-3.4) 2.2 ± 0.2 (-21.0) Breast, right (11) 4.7 ± 0.3 4.4 ± 0.2 (-6.3) 3.0 ± 0.1 (-37.3) Breast, left (12) 4.6 ± 0.4 4.2 ± 0.2 (-8.9) 2.9 ± 0.2 (-38.1) Heart (4) 3.5 ± 0.3 3.4 ± 0.1 (-2.8) 2.6 ± 0.1 (-26.2) Heart (8) 3.7 ± 0.3 3.7 ± 0.2 (-0.7) 2.8 ± 0.2 (-25.7) Lung, right (2) 2.7 ± 0.3 2.6 ± 0.1 (-6.4) 2.2 ± 0.2 (-18.1) Lung, left anterior (7) 3.6 ± 0.3 3.4 ± 0.3 (-5.8) 2.4 ± 0.1 (-31.4) Lung, left posterior (6) 3.2 ± 0.2 3.1 ± 0.2 (-1.6) 2.8 ± 0.1 (-11.7) Lung, right (5) 3.5 ± 0.4 3.7 ± 0.3 (5.6) 3.0 ± 0.1 (-13.0) Lung, right (10) 4.4 ± 0.3 4.3 ± 0.3 (-3.2) 3.7 ± 0.3 (-15.3) 1 Organ Effective Modulation.

The HU values obtained from the bismuth scan images were generally different compared to the OEM and reference scan images, as shown in Table 13. HU values from the OEM scan differed from the reference scan by -0.1% on average. Similarly, HU values from the bismuth scan differed from those from the reference scan by 2.9% on average. The greatest increases in noise were also observed in the bismuth scan, where noise values varied from -1.9 to 338.1% compared to the reference scan. The OEM scan showed noise variation between -10.6 and 9.6% compared to the reference scan. The average noise difference when comparing the reference scan and the OEM scan, excluding the thyroid’s location, was 0.6% and for the bismuth scan 7.4%. The bismuth shields caused additional image artefacts when metal artefact removal software was used, as presented in Figure 12.

52 52

Table 13. HU and noise values from different helical chest CT scans, with differences from the reference scan shown as a percentage in parentheses (table obtained from Study II)

Measurement point Reference HU 1 OEM2 HU Bismuth HU Reference noise OEM noise Bismuth noise Thyroid (1) 20.6 23.8 (15.1) 139.2 (574.9) 10.0 9.6 (-3.8) 43.9 (338.1) Sternum (3) 950.3 949.4 (-0.1) 949 (-0.1) 25.3 27.7 (9.5) 29.9 (18.2) Liver (13) 11.3 10.4 (-8.4) 16 (41.3) 14.6 15 (2.4) 15.5 (6.1) Thoracic spine (9) 962.5 965.3 (0.3) 940.4 (-2.3) 21.6 21.5 (-0.6) 21.2 (-1.9) Breast, right (11) -45.8 -48.9 (6.7) -27.9 (-39.1) 14.6 14.5 (-0.5) 14.1 (-3.4) Breast, left (12) -43.8 -43.3 (-1.1) -24.5 (-43.9) 11.6 12.4 (7.3) 11.9 (2.6) Heart (4) 23.2 20.5 (-11.5) 30.9 (33.2) 13.8 15.1 (9.6) 14.2 (3.2) Heart (8) 17.5 17 (-2.6) 26.8 (53.2) 14.5 13 (-10.6) 14.7 (1.4) Lung, right (2) -776.5 -772.4 (-0.5) -774.6 (-0.3) 26.3 24.7 (-6.3) 25.9 (-1.6) Lung, left anterior (7) -790.0 -793.7 (0.5) -767.3 (-2.9) 19.9 18.4 (-7.8) 22.8 (14.5) Lung, left posterior (6) -775.4 -776.4 (0.1) -767.7 (-1) 19.5 19.9 (2.4) 22.4 (15) Lung, right (5) -778.9 -778.1 (-0.1) -767.7 (-1.4) 18.8 18.6 (-0.7) 22.6 (20.7) Lung, right (10) -768.6 -772.7 (0.5) -758.1 (-1.4) 22.3 22.9 (3.0) 25.3 (13.8) 1 Hounsfield unit. 2Organ Effective Modulation.

Fig. 12. The artefacts caused by bismuth shields when used with metal artefact correction software. The artefacts shown on the left image resemble beam hardening and streaking artefacts (figure obtained from Study II).

5.3 Segmented dental panoramic tomography

In Study III, pediatric protocols were selected in 93% of all sDPT examinations for children between 2 and 6 years of age and in 69% of all examinations for those between 7 and 12 years of age (Table 14). The DAP reduction when using pediatric collimation and dedicated pediatric programmes was, on average, 50% for children between 7 and 12 years of age. The highest effective doses with both phantoms were determined in full DPT examinations (Table 15 and 16). Salivary glands, oral mucosa, oesophagus and extrathoracic regions were determined to receive the highest organ doses. The dose reduction in sDPT compared to the full DPT was between 13 and 78% in the pediatric set-up and between 23 and 86% in the adult set-up.

Table 14. Mean dose area product (DAP) values and number of examinations (in brackets) in the youngest age groups from the pediatric and adult programmes together with the percentage of how often a pediatric programme was chosen (third molars and implants are excluded from the table, as none of the examinations of 2-to-12-year-old patients were due for them)

Indication Pediatric protocol Adult protocol %1

Orthodontic reason 2-–6 years 22.7 (9) 26.9 (1) 90 7–12 years 23.1 (918) 45.8 (411) 69

TMJ 2–6 years - - - 7–12 years 7.5 (2) 23.3 (4) 33

Lesions 2–6 years - - - 7–12 years 16.2 (7) 43.3 (3) 70

Trauma 2–6 years 26 (1) - 100 7–12 years 19.7 (4) 32.6 (4) 50

Other 2–6 years 23.4 (3) - 100 7–12 years 21.3 (9) 41.1 (5) 64

Total 2–6 years 23.1 (13) 26.9 (1) 93 7–12 years 22.9 (940) 45.4 (427) 69 1 Percentage of examinations done using a pediatric protocol.

Table 15. Equivalent and effective doses from pediatric phantom measurements in µSv (effective dose uncertainty in µSv is presented in parentheses)

Organ DPT 1 TMJ2 Right TMJ Dentition Anterior region Right side of DPT Molar region Third molars Thyroid 25.9 8.3 4.9 18.8 8.6 15.4 6.6 21.6 Salivary glands 188.1 112.9 73.0 135.5 28.2 135.6 92.0 187.6 Brain 58.8 28.6 13.3 46.6 20.5 42.7 13.9 34.8 Lens 10.4 2.0 1.1 9.5 4.9 8.4 3.7 7.3 Red bone marrow 16.9 8.2 5.4 12.6 3.8 12.5 7.8 14.6 Oesophagus 11.2 6.2 4.1 8.0 2.0 7.8 5.0 10.9 Skin 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 Remainder tissues Bone surface 71.5 34.5 22.9 53.3 16.1 52.8 32.9 61.6 Lymphatic nodes 10.1 6.1 3.9 7.3 1.5 7.3 4.9 10.1 Oral mucosa 61.6 11.8 4.3 59.7 45.6 53.2 13.7 38.9 Muscle 10.1 6.0 3.9 7.3 1.5 7.3 4.9 10.1 Extrathoracic region 188.1 112.9 73.0 135.5 28.2 135.6 92.0 187.6 Effective dose 14.6 (4.1) 7.1 (2.4) 4.6 (1.8) 10.8 (3.0) 3.3 (0.7) 10.6 (3.3) 6.4 (2.3) 12.7 (4.1) 1 Dental panoramic tomography. 2 Temporomandibular joint.

56 56 Table 16. Equivalent and effective doses from adult phantom measurements in µSv (effective dose uncertainty in µSv is presented in parentheses)

Organ DPT 1 TMJ2 Right TMJ Dentition Anterior region Right side of DPT Molar region Third molars Thyroid 45.8 5.9 14.5 29.6 13.5 23.2 12.9 29.8 Salivary glands 864.0 227.4 329.4 508.8 170.1 407.0 244.2 688.2 Brain 85.6 8.7 19.6 57.0 49.7 55.9 5.6 31.0 Lens 12.1 0.7 1.0 6.3 5.3 4.5 2.3 6.0 Red bone marrow 7.1 1.6 3.4 4.2 1.7 3.5 2.2 5.2 Oesophagus 82.7 30.0 61.9 51.0 17.7 52.0 20.4 63.6 Skin 0.0 0.0 0.0 0.0 0.0 0.0 0.0 0.0 Remainder tissues Bone surface 30.1 6.8 14.5 17.8 7.2 14.9 9.2 22.0 Lymphatic nodes 45.1 12.8 12.3 25.2 7.4 19.1 13.6 37.2 Oral mucosa 1711.5 457.7 948.3 825.5 87.2 743.6 397.0 1237.7 Muscle 45.0 12.7 12.2 25.1 7.4 19.1 13.6 37.2 Extrathoracic region 928.4 303.3 286.1 522.5 117.5 391.8 293.8 821.0 Effective dose 39.1 (12.6) 10.8 (5.5) 17.9 (8.4) 21.5 (7.6) 5.6 (3.1) 18.2 (8) 10.3 (4.5) 30.0 (10.1) 1 Dental panoramic tomography. 2 Temporomandibular joint.

5.4 Uncertainty of dose determinations

Study I

In the first study, we estimated the positioning uncertainty to be 5% and the source uncertainty obtained from the imaging device’s annual quality control measurements as 1.3%.

For the MOSFETs, we estimated the uD(Cmeas) to be 21.1% based on the average uncertainty of the MOSFET response at the used dose range, while the uD(KE) was 9.01% based on the energy dependence measurements. We estimated angle uncertainties uD(θ) for the axis rotation to be 5%, normal-to-axial uncertainties as 8% and tangent-to-axial uncertainties as 7% based on the PMMA measurements done by Koivisto et al. (J. Koivisto et al., 2013). The average uc for the MOSFET measurements was therefore 26.3% according to equation (19).

For the RPL dosimeters, we estimated the uD(Cmeas) to be 6.0% based on the uncertainty of the RPL in the dose range, while the uD(KE) was 2.9% and the uD(θ) for the vertical angle was 1.7% (Manninen et al., 2012). Therefore, the uc was 8.7% according to equation (19). Statistical organ dose uncertainty from the MC simulation of the AP projection was, on average, 10.9% and the statistical effective dose uncertainty was 0.6%. Similarly, the average organ dose uncertainties for the PA and the LAT projections were 12.4% and 12.7%, respectively, and for the effective doses 0.6% and 0.6%.

Study II

In the second study, we calibrated the MOSFETs individually for the used energy; thus, the uncertainty varied between 2.4% and 13.6%, with a mean uncertainty of 7.3%.

Study III

We obtained the energy and mean vertical angle uncertainties of the RPLDs from Manninen et al. (Manninen et al., 2012). We estimated the positioning uncertainty to be 5%, while the source uncertainty obtained from the imaging device’s annual quality control measurements was 3.2%. In the third study, the uD(KE) was 4.0%,

the uD(Cmeas) for the adult set-up was on average 3.8% and the uc was on average

6.6%. For the pediatric set-up, the uD(Cmeas) was on average 3.5% and uc was, according to equation (19), 6.4% on average. We estimated effective dose uncertainties for Study III using equation (21), and they are shown in Tables 14 and 15

6 Discussion

6.1 RPL and MOSFET in conventional radiology

We evaluated the properties of the MOSFET dosimeters using a diagnostic x-ray spectrum, while we determined radiation doses from the conventional thorax x-ray projections using MOSFET and RPL dosimeters.

6.1.1 MOSFET’s energy dependence of the response

The energy dependence of the MOSFET dosimeters have been evaluated using different set-ups. The results vary because the methods used to determine the energy dependence are not identical. Peet et al. reported similar behaviour regarding MOSFET’s energy response along the lower part of energy spectrum. The sensitivity of dosimeter was lower in the lower kVp range than at the calibration kVp (Peet & Pryor, 1999). Koivisto et al. used two different additional filtrations, 2.5mmAl and 2.5mmAl+0.5mmCu, with a CBCT x-ray unit (J. Koivisto et al., 2015). Their results showed that the sensitivity of the MOSFETs in different kVp set-ups were nearly independent of the energy. With a 2.5mmAl additional filtration, the coefficient of determination (R2) was lower than 0.01, but with a higher filtration R2 was more than 0.25. Bower et al. presented results similar to those of Koivisto et al., where the sensitivity of the dosimeters did not seem to change with different kVps (Bower & Hintenlang, 1998). Dong et al. showed that the energy response of MOSFETs was nearly within a 10% range of variation close to the calibration energy, which is similar to our results (Dong, Chu, Lee, et al., 2002). Variation in the energy dependence results from the different MOSFET studies can mostly be explained by different backscatter mediums. Peet et al. measured energy dependency in free-in-air set-up that was similar to our lead apron set-up. Dong et al. did not report the thickness of their solid water phantom, while Koivisto et al. used a 15 cm x 15 cm x 10 cm PMMA block and Bower et al. used some type of acrylic plate. With a thicker backscatter medium, more radiation is scattered towards the dosimeter, which might reduce the effect of angular dependence of the dosimeter. Bower et al. reported a difference in the dosimeter’s response whether the dosimeter is used with the epoxy bubble toward x-ray source or with the flat side of the dosimeter toward the x-ray source. It has been suggested that phantom measurements along the surface or near the surface should be made

by facing the flat side of the MOSFET toward the x-ray source, which would seem to be more accurate compared to the Monte Carlo-simulated results (Jones, Pazik, Hintenlang, & Bolch, 2005). The coefficient of variation of the averaged measurements in our study was relatively large and could have been reduced by increasing the dose or the number of averages.

6.1.2 MOSFET’s dose linearity and repeatability

Our results from Study I showed a good dose linearity with R2 over 0.9985, which is in accordance with the previously reported R2=0.99987 at dose levels ranging from 0.24 to 17.5 mGy and R2 over 0.999 at doses ranging from 0 to 20 mGy. (Dong, Chu, Lan, et al., 2002; J. Koivisto et al., 2015). Koivisto et al. reported dose responses ranging from 0.24 to 17.5 mGy with standard deviations between 13 and 2% (J. Koivisto et al., 2015). Our results showed standard deviations in percentages that varied between 47 and 5% in the dose range between 0.21 and 10.29 mGy. It should be noted that the x-ray spectrum was not the same, and the field size and thickness of the backscatter material have a notable effect on the sensitivity of the MOSFETs (Brady & Kaufman, 2012). Poor repeatability of MOSFETs with doses lower than 3 mGy have also been reported previously (Dong, Chu, Lan, et al., 2002; Peet & Pryor, 1999). With poor repeatability at low doses, the use of MOSFETs should be limited to applications where the absorbed doses are sufficient for accurate measurements or else multiple exposures should be used to increase the net exposure level.

6.1.3 Absorbed and effective dose determination

The equivalent doses for organ locations that were in the primary beam (oesophagus, breasts, lungs and liver) showed the greatest similarity between the different dose determination methods. Dose differences for organ locations receiving only scattered radiation showed the greatest difference between different methods in the AP projection. As there are a limited number of dosimeters with a limited dose and energy range, the results from the MC simulations differed more in the areas that received only scattered radiation and where organs had a higher volume/surface, e.g. bone surface and stomach. It should be noted that differences between the experimental set-up and the MC simulations can arise from, among other things, differences in the anthropomorphic models, organ definitions and dose sampling schemes. Low statistical uncertainty from the MC simulations can be

misleading, as a systematic error can arise from all input parameters used in the simulations, e.g. misplacement of the x-ray field, organ compositions and density (Schmidt, Dance, Skinner, Smith, & McNeill, 2000). When comparing AP and PA projections, radiation-sensitive tissues such as thyroid, breasts, liver and stomach received notably less radiation due to the different irradiation angle. Since many of the radiation-sensitive organs receive less radiation, the decrease in the effective dose is understandable. Thus, a PA projection should be the first choice for thorax examinations if possible, as is generally recommended (Carmichael, J. H. E; Moores, B. M. et al; Maccia, 1996). The LAT projection showed similar results between the different dose determination methods. The largest absolute differences between the measured doses and simulated doses were in the thyroid, salivary glands, active bone marrow and skin doses. Thyroid and salivary glands are focal organs, and as they are located close to the edge of the primary beam, a small difference in the primary beam setting or model differences can explain the differences. The phantom set-up in measurements also had a larger breast phantom compared to the MC simulation phantom, which can increase scatter to the superior organs. Skin and active bone marrow differences are likely to occur because their dose is estimated only from a few points in the measurements, whereas simulations consider entire organs, including both the parts receiving primary and scattered radiation. The LAT projection had smaller effective doses than the PA and AP projections in our measurement set-up. This might not be the case in a clinical set-up, where automatic exposure control is usually used, whereas we used the same entrance skin dose for every projection. The phantom used in the MC simulation is not an exact copy of the phantom used in the dose measurements, and thus there are differences in the size and location of the different organs. Furthermore, the breasts used in the dosimetry set- ups differ from the breasts used in the MC simulations in shape, size and material composition. Despite the phantom differences, the MC simulations provided another method for obtaining dose information. The relatively high average organ dose uncertainties in the MC simulations are due to the high uncertainties from organs that lie outside of the primary x-ray photons, thus they receive dose only from scattered radiation. The high uncertainties could be decreased by using a higher number of photons in the simulations, thus lowering the statistical uncertainty. It should be noted that organs located outside the primary beam have lower equivalent doses, and therefore, their level of uncertainty does not greatly affect the uncertainty of the effective dose.

6.2 Organ effective modulation vs. bismuth shields

Organ-based tube current modulation techniques have been developed to reduce doses to radiation-sensitive organs located near the surface without causing image- quality degradation. As the method for dose reduction varies for each CT manufacturer, every method should be validated before generalisations regarding their use can be made. Details on different OBTCM solutions are presented on p. 1 of Study II. Most of the current studies on OBTCM methods have been performed on X-Care by Siemens and on ODM by GE, while the number of studies on OEM are still few. ODM and OEM work based on similar techniques, where the tube current is reduced along the anterior side of the patient without compensation on the other sides. The most notable difference between ODM and OEM has to do with how much the tube current is reduced on the side of the radiation-sensitive organs and whether the dose reduction is applied to the entire scan. With ODM, the user can select the length in the z-direction to which ODM is applied, whereas with OEM the dose reduction is applied to the entire scan. X-Care, on the other hand, increases the tube current on the contralateral side and on the lateral sides to keep the image quality constant throughout the entire scan. With X-Care, the pitch value is limited to 0.6 (Siemens, 2014). As the degree of tube current reduction can vary with different devices, we evaluated the performance of OEM and the bismuth shields against the reference method. The dose reduction with the bismuth shields was shown to be more effective compared to the OEM in our set-up. Bismuth shields in general have been shown to be good for dose reduction (Hoang et al., 2012; Nikupaavo, Kaasalainen, Reijonen, Ahonen, & Kortesniemi, 2015; Raissaki, Perisinakis, Damilakis, & Gourtsoyiannis, 2010; Samei, 2014). The problems with bismuth shields have been associated with changes in the HU values (Wang et al., 2011), an increase of noise (Lambert & Gould, 2016), additional artifacts (Wang et al., 2011) and interference with AEC (Samei, 2014). Thus, their use cannot always be recommended. Our study shows there is variation in the HU numbers and an increase of noise at many measurement locations, which causes bismuth shields to not always be applicable with the studied device, especially if metal artefact removal software is needed. OEM reduced the absorbed doses at all but one lung location, although the dose reduction was quite modest in other organs. The noise and HU difference were small, as has been shown with other OBTCM methods (Duan et al., 2011; Hoang et al., 2012; Kim, Sung, Choi, Kim, & Kim, 2013; Lambert & Gould, 2016; Wang et al., 2012, 2011).

Since the dose measurements were performed on imperfect dosimeters, the uncertainty of the measurements needs to be considered. The standard deviation of the mean was between 0.08 mGy and 0.39 mGy, which corresponds to CV variations between 2% and 14%. As the recommendation for the overall standard uncertainty for dose measurements in diagnostic radiology is less than 12.5% (Zoetelief et al., 2000), the applicability of MOSFETs for this type of dose measurement is at its limit. Due to the image-quality problems with bismuth shields, the OBTCM methods provide an alternative, although less effective, method for dose reduction with this specific CT scanner model and setting.

6.3 Segmented dental panoramic tomography

The effective doses from dental x-ray examinations are low compared to conventional x-ray examinations, CT examinations and interventional x-ray examinations. Although the radiation doses in dental x-ray examinations are low. the number of examinations are high, thus contributing to the cumulative effective dose of the population (Medical Radiation Exposure of the European Population Part 1/2, 2015). Different manufacturers have brought in new technologies to lower the dose, such as segmented DPT and pediatric collimation (Benchimol, Koivisto, Kadesjö, & Shi, 2018; Davis et al., 2015). Our measurements are in line with previously reported effective doses regarding DPT and sDPT (Benchimol et al., 2018; Granlund et al., 2016). The uncertainties regarding effective doses were between 20 and 38% for the pediatric set-up and between 32 and 56% for the adult set-up. The large degree of uncertainty regarding effective doses is mainly due to the high uncertainty value of the fraction of tissue irradiated for red bone marrow, bone surface, lymphatic nodes and muscle.

7 Conclusions

The main finding from this thesis can be summarised as follows: 1. The MOSFET system can be used for radiation dose determinations with different devices in radiology if its inherent properties, such as poor dose reproducibility at dose levels in general lower than 5 mGy, energy dependence, angular dependency and calibration procedures are carefully assessed and considered. 2. In a helical chest CT scan with the studied scanner model, protocol setting and phantom model, the bismuth shield provided larger dose savings compared to the OEM, but with a negative impact on quantitative image quality. OEM provided a small dose reduction without any notable alteration in image quality 3. In dental panoramic tomography, the use of segmented DPT and application of a pediatric protocol, when properly indicated, can reduce the radiation dose.

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Original publications

I Manninen AL*, Kotiaho A*, Nikkinen J, Nieminen MT (2015). Validation of a MOSFET dosimeter system for determining the absorbed and effective radiation doses in diagnostic radiology. Radiation Protection Dosimetry, Apr;164(3), 361-7. II Kotiaho A, Manninen AL, Nikkinen J, Nieminen MT (2018). Comparison of organ- based tube current modulation and bismuth shielding in chest CT: Effect on the image quality and the patient dose. Radiation Protection Dosimetry, Dec. Ahead of print. III Kotiaho A*, Sipola A*, Happo E, Haapea M, Nikkinen J, Kallio-Pulkkinen S, Nieminen MT (2019). Use of segmented dental panoramic tomography (sDPT) for dose reduction in comparison to full DPT. Manuscript Reprinted with permission from Oxford Journals (I and II).

Original publications are not included in the electronic version of the doctoral thesis.

74 ACTA UNIVERSITATIS OULUENSIS SERIES D MEDICA

1499. Piispala, Johanna (2019) Atypical electrical brain activity related to attention and inhibitory control in children who stutter 1500. Leppänen, Joni (2019) The role of hypoxia, innate immunity receptors and stromal response in pancreatic cancer 1501. Lahtinen, Sanna (2019) Complications, quality of life and outcome after free flap surgery for cancer of the head and neck 1502. Rajavaara, Päivi (2019) Children’s dental general anaesthesia : reasons and associated factors 1503. Varpuluoma, Outi (2019) Drugs, dermatitis herpetiformis and celiac disease as risk factors for bullous pemphigoid in Finland 1504. Mäkelä-Kaikkonen, Johanna (2019) Robotic-assisted and laparoscopic ventral rectopexy in the treatment of posterior pelvic floor procidentia 1505. Lahtinen, Antti (2019) Rehabilitation after hip fracture : Comparison of physical, geriatric and conventional treatment 1506. Alakärppä, Antti (2019) Primary sinonasal surgery and health-related quality of life in adults 1507. Komulainen-Ebrahim, Jonna (2019) Genetic aetiologies and phenotypic variations of childhood-onset epileptic encephalopathies and movement disorders 1508. Jussila, Päivi (2019) Prevalence and associated risk factors of temporomandibular disorders (TMD) in the Northern Finland Birth Cohort (NFBC) 1966 1509. Määttä, Jenni (2019) Effects of the hypoxia response on metabolism in atherosclerosis and pregnancy 1510. Taka-Eilola, Tiina (2019) Mental Health Problems in the Adult Offspring of Antenatally Depressed mothers in the Northern Finland 1966 Birth Cohort; Relationship with Parental Severe Mental Disorder 1511. Korhonen, Tommi (2019) Bone flap survival and resorption after autologous cranioplasty 1512. Sliz, Eeva (2019) Genetics and molecular epidemiology of metabolic syndrome- related traits : Focus on metabolic profiling of lipid-lowering therapies and fatty liver, and the role of genetic factors in inflammatory load 1513. Ollila, Meri-Maija (2019) The role of polycystic ovary syndrome (PCOS) and overweight/obesity in women’s metabolic and cardiovascular risk factors and related morbidities

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