2. High-Intensity Focused Ultrasound 17 2.1 Ultrasound in Medicine

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tcartsbA desu eb nac taht euqinhcet cituepareht gnigreme na si )UFIH( dnuosartlu desucof ytisnetni-hgiH desucof dnuosartlu )UFIH( si na gnigreme cituepareht euqinhcet taht nac eb desu hgih erutarepmet a ot eussit eht gnitaeH .niks hguorht ylevisavni-non dna yllacol eussit taeh ot taeh eussit yllacol dna ylevisavni-non hguorht .niks gnitaeH eht eussit ot a erutarepmet hgih ,hguone .g.e , ot ,C°75 secudni suoenatnatsni llec htaed hguorht lamreht evitalugaoc .sisorcen deniatniam si erutarepmet eussit hcihw ni ,aimrehtrepyh rof desu eb nac UFIH ,ylevitanretlA UFIH nac eb desu rof ,aimrehtrepyh ni hcihw eussit erutarepmet si deniatniam ( emit fo doirep regnol a rof C°54–04 fo egnar a nihtiw a egnar fo C°54–04 rof a regnol doirep fo emit ( .g.e , 06 )nim ot ecnahne eht tceffe fo rehto )IRM( gnigami ecnanoser citengaM .yparehtomehc dna yparehtoidar sa hcus ,seitiladom ypareht ,seitiladom hcus sa yparehtoidar dna .yparehtomehc citengaM ecnanoser gnigami )IRM( ,tnemtaert eht gnitegrat yllaitaps rof snaem sedivorp )UFIH-RM( stnemtaert UFIH fo ecnadiug sa ecnadiug fo UFIH stnemtaert )UFIH-RM( sedivorp snaem rof yllaitaps gnitegrat eht ,tnemtaert egamad eussit gnitaulave dna ,gnitaeh gnirud emit-laer ni noitulove erutarepmet eht gnirusaem eht erutarepmet noitulove ni emit-laer gnirud ,gnitaeh dna gnitaulave eussit egamad .ypareht retfa .ypareht metsys UFIH-RM na rof noitulos aimrehtrepyh a poleved ot saw sisehT siht fo mia yramirp ehT ehT yramirp mia fo siht sisehT saw ot poleved a aimrehtrepyh noitulos rof na UFIH-RM metsys -IRM elbailer htiw niks eht morf shtped tnaveler yllacinilc ta semulov egral taeh ylefas elbane ot elbane ylefas taeh egral semulov ta yllacinilc tnaveler shtped morf eht niks htiw elbailer -IRM dnuosartlu eht fo stnemevom lacinahcem denibmoc noitulos ehT .gnippam erutarepmet desab erutarepmet .gnippam ehT noitulos denibmoc lacinahcem stnemevom fo eht dnuosartlu lortnoc ot stnemele recudsnart fo esu evitceles dna ,gnireets sucof dnuosartlu cinortcele ,recudsnart cinortcele dnuosartlu sucof ,gnireets dna evitceles esu fo recudsnart stnemele ot lortnoc erutarepmet IRM lareveS .htap maeb citsuoca eritne eht nihtiw noitubirtsid erutarepmet eht erutarepmet noitubirtsid nihtiw eht eritne citsuoca maeb .htap lareveS IRM erutarepmet aimrehtrepyh efaS .emit-laer ni gnitaeh eht rotinom dna lortnoc ot desu erew secils gnippam secils erew desu ot lortnoc dna rotinom eht gnitaeh ni .emit-laer efaS aimrehtrepyh stnemirepxe lamina ni deveihca saw ytimrofinu erutarepmet doog htiw gnitaeh htiw doog erutarepmet ytimrofinu saw deveihca ni lamina stnemirepxe ni oviv . ,eromrehtruF ni oviv lamina stnemirepxe dna tneitap gnigami yduts dewohs taht eht depoleved lacinhcet ,yllaniF .recnac latcer tnerrucer fo aimrehtrepyh rof elbisaef saw noitulos aimrehtrepyh noitulos saw elbisaef rof aimrehtrepyh fo tnerrucer latcer .recnac ,yllaniF lacinhcet .C°1 naht retteb ycarucca htiw yrtemomreht desab-IRM noitarud-gnol delbane snoitulos delbane noitarud-gnol desab-IRM yrtemomreht htiw ycarucca retteb naht .C°1 ni srecudsnart yarra-desahp esu ot syaw wen tneserp ot saw sisehT siht fo mia yradnoceS ehT ehT yradnoceS mia fo siht sisehT saw ot tneserp wen syaw ot esu yarra-desahp srecudsnart ni UFIH-RM tnereffid ni ecnamrofrep ni tnemevorpmi niag ot dlefi citsuoca dettime eht gnipahs eht dettime citsuoca dlefi ot niag tnemevorpmi ni ecnamrofrep ni tnereffid UFIH-RM eussit eht yb decneirepxe erusserp kaep eht ecuder ot elba saw gnitaeh icofitlum ,tsriF .snoitacilppa ,tsriF icofitlum gnitaeh saw elba ot ecuder eht kaep erusserp decneirepxe yb eht eussit tnemtsujda eht ,dnoceS .gnitaeh sucof-dereets elgnis ot derapmoc nehw gnitaeh aimrehtrepyh ni aimrehtrepyh gnitaeh nehw derapmoc ot elgnis sucof-dereets .gnitaeh ,dnoceS eht tnemtsujda devorpmi segami IRM no desab noitamitse thgifl-fo-emit gnizilitu sesahp tnemele-recudsnart fo tnemele-recudsnart sesahp gnizilitu thgifl-fo-emit noitamitse desab no IRM segami devorpmi ,drihT .eussit tsaerb fo snoitidnoc gnitalumis motnahp suoenegoreteh ni ssenprahs sucof eht sucof ssenprahs ni suoenegoreteh motnahp gnitalumis snoitidnoc fo tsaerb .eussit ,drihT noitcuder evitatitnauq tsaf selbane hcihw ,senob fo gnitaeh detciderp ytisnetni citsuoca detalumis citsuoca ytisnetni detciderp gnitaeh fo ,senob hcihw selbane tsaf evitatitnauq noitcuder latsocretni fo txetnoc eht ni stnemele recudsnart fo noitavitcaed hguorht gnitaeh enob fo enob gnitaeh hguorht noitavitcaed fo recudsnart stnemele ni eht txetnoc fo latsocretni .snoitacinos sdrawot UFIH-RM fo dlefi eht ecnavda sisehT siht ni detneserp snoitulos lacigolonhcet ehT ehT lacigolonhcet snoitulos detneserp ni siht sisehT ecnavda eht dlefi fo UFIH-RM sdrawot .ecitcarp lacinilc otni noitalsnart otni lacinilc .ecitcarp

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ämletsiviiT no )UFIH .st ,dnuosartlu desucof ytisnetni-hgih .lgne( inääartlu utiosukof nenitteetisnetniaekroK utiosukof inääartlu .lgne( ytisnetni-hgih desucof ,dnuosartlu .st )UFIH no itsesiviisavni-ie aj itsesillakiap atsoduk äättimmäl naadiov alloj ,akkiinket nenittueparet ävytsiely nenittueparet ,akkiinket alloj naadiov äättimmäl atsoduk itsesillakiap aj itsesiviisavni-ie nesiviitalugaok naepon aattuehia nees:C° 75 iskikremise nenimättimmäl neskoduK .ipäl nohi .ipäl neskoduK nenimättimmäl iskikremise 75 nees:C° aattuehia naepon nesiviitalugaok näätedipälly aiskoduk assoj ,naaimretrepyh söym äättyäk naadiov ätsytimmäl-UFIH .nisoorken ätsytimmäl-UFIH naadiov äättyäk söym ,naaimretrepyh assoj aiskoduk näätedipälly ,nejotoumaiparet nedium aatsohet aimretrepyH .)C° 54 – 04( assalitöpmäl assutetohok )nim 06( )nim assutetohok assalitöpmäl 04( – 54 .)C° aimretrepyH aatsohet nedium ,nejotoumaiparet citengam .lgne( neskuavukissnanoseritteengaM .aohet naiparetomek aj nodiohedäs netuk nodiohedäs aj naiparetomek .aohet neskuavukissnanoseritteengaM .lgne( citengam .st ,UFIH dediug-IRM .lgne( neeskuajho naiparet-UFIH nenimättyäk )IRM .st ,gnigami ecnanoser ,gnigami .st )IRM nenimättyäk naiparet-UFIH neeskuajho .lgne( dediug-IRM ,UFIH .st nodioh nannarues nalitöpmäl nesiakiailaaer ,neskunnedhok nodioh aatsillodham )UFIH-RM aatsillodham nodioh ,neskunnedhok nesiakiailaaer nalitöpmäl nannarues nodioh .neekläj nodioh ninnioivra nohet nodioh äkes anakia äkes nodioh nohet ninnioivra nodioh .neekläj akoj ,ämletenemaimretrepyh-UFIH-RM äättihek ilo etiovat neniajisisne najrikötsiäv nämäT nämäT najrikötsiäv neniajisisne etiovat ilo äättihek ,ämletenemaimretrepyh-UFIH-RM akoj naatnip nohi älledyyvys allitnaveler itsesiniilk neduuvalitsoduk neruus näämättimmäl yytsyp näämättimmäl neruus neduuvalitsoduk itsesiniilk allitnaveler älledyyvys nohi naatnip ämletenemsuattimalitöpmäl neniajhop-IRM avattetoul äättihek ilo sutiokrat iskäsiL .nedhän iskäsiL sutiokrat ilo äättihek avattetoul neniajhop-IRM ämletenemsuattimalitöpmäl naamiollortnok ikryp ämletenemsytimmäl yttetiheK .neesimaajho neskytimmälaimretrepyh .neesimaajho yttetiheK ämletenemsytimmäl ikryp naamiollortnok allamaajho ,itsesinaakem ätnitehälinääartlu neattukiil älläsis nätnek nesitsuka ätsimenepmäl nesitsuka nätnek älläsis neattukiil ätnitehälinääartlu ,itsesinaakem allamaajho .aamuakaj neittnemeleinääartlu netsiviitka allamiollortnok äkes itsesinortkele ättetsipsukof itsesinortkele äkes allamiollortnok netsiviitka neittnemeleinääartlu .aamuakaj .atielapiivsuavuk atiesu neätnydöyh ätsytimmäl isajho suattimalitöpmäl-IRM ,neniakiailaaeR suattimalitöpmäl-IRM isajho ätsytimmäl neätnydöyh atiesu .atielapiivsuavuk niittetuvaas ällämletenemsytimmäl ällytetiheK ällämletenemsytimmäl niittetuvaas ni oviv assiekokniäle- nenillavrut äkes neniasat nootioh nävöysnelousärep neenutuisuu suuvutlevos nämleneM .sytimmäl nenimretrepyh .sytimmäl nämleneM suuvutlevos neenutuisuu nävöysnelousärep nootioh neskytimmäL .alleskumiktutsimatnavuk ällydhet allialitop äkes alliekokniäle niittetioso alliekokniäle äkes allialitop ällydhet .alleskumiktutsimatnavuk neskytimmäL .C° 1 niuk ipmerap ilo suukkrat neskuattimalitöpmäl-IRM nytetihek neesimaajho nytetihek neskuattimalitöpmäl-IRM suukkrat ilo ipmerap niuk 1 .C° alliovat allisuu ätnitehälinääartlu auttetsiehiav äätnydöyh ilo etiovat neniot najriksötiäV najriksötiäV neniot etiovat ilo äätnydöyh auttetsiehiav ätnitehälinääartlu allisuu alliovat -UFIH-RM ire äykyksutsirous neettialotioh naamatnarap assesimaakkoum nätnek nesitsuka nätnek assesimaakkoum naamatnarap neettialotioh äykyksutsirous ire -UFIH-RM neittnemeleinääartlu ittyäk akoj ,ämletenemsytimmäl nenimretrepyh ,nikännisnE .assiskullevos ,nikännisnE nenimretrepyh ,ämletenemsytimmäl akoj ittyäk neittnemeleinääartlu iksal ,naajis nävätterriis nedhy ättetsipsukof atsiakiaäthy aatnom auttetout alleskutsiehiav auttetout aatnom atsiakiaäthy ättetsipsukof nedhy nävätterriis ,naajis iksal neättyäk sutsiehiav neittnemele ,iskesioT .atteniapimiskam aavutsidhok niiskodok itsävättikrem niiskodok aavutsidhok .atteniapimiskam ,iskesioT neittnemele sutsiehiav neättyäk ajoreakiaukluk nenääartlu aivutuehia ätsedyysineegoreteh netsoduk ,ajutteksal atsivuk-IRM ,ajutteksal netsoduk ätsedyysineegoreteh aivutuehia nenääartlu ajoreakiaukluk atneksalitteetisnetni nätnek nesitsuka ,iskennamloK .itsävles ättyyväret neskukof nesitsuka isnarap nesitsuka neskukof ättyyväret .itsävles ,iskennamloK nesitsuka nätnek atneksalitteetisnetni aneetsurep äättyäk naadiov atoj ,nesimenepmäl nuul namattuehia nenääartlu itsakrat itsunne itsakrat nenääartlu namattuehia nuul ,nesimenepmäl atoj naadiov äättyäk aneetsurep ätsimenepmäl nuul itsesiviitatitnavk äätnehäv nello niän aj neesimeklus neittnemeleinääartlu neesimeklus aj niän nello äätnehäv itsesiviitatitnavk nuul ätsimenepmäl .assodioh-UFIH-RM assavutsidhok naaskam iskikremise naaskam assavutsidhok .assodioh-UFIH-RM nakiinket-UFIH-RM naatlaso tavattua 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Preface

The work presented herein has been performed at Philips MR-Therapy during the years 2012–2016 in collaboration with University of Bordeaux 2, France; National Institutes of Health, Bethesda, MD, USA; Aarhus University Hospital, Aarhus, Denmark; Sunnybrook Health Sciences Cen- tre, Toronto, Canada; and University of Texas Southwestern Medical Cen- ter, Dallas, TX, USA. First and foremost I would like to thank my instructor Docent Heikki Nieminen for his guidance during the works of this thesis, and especially for his help with scientiﬁc writing. I would also like to express my deepest gratitude to my supervisor Prof. Risto Ilmoniemi throughout the years for his support. Pre-examiners Prof. Gail ter Haar and Docent Timo Liimatainen are greatly appreciated for their careful review and constructive criticism. I am also truly grateful to Prof. ter Haar for careful linguistic review of the thesis. It has been a privilege to work with wonderful colleagues at Philips MR- Therapy: Dr. Mika Ylihautala, Dr. Max Köhler, Dr. Julius Koskela, Dr. Ari Partanen, Dr. Robert Staruch, Dr. Charles Mougenot, Kari Rum- mukainen, Matti Lindström, Dr. Erkki Vahala, and many others, without whom this work would have not been possible. I would also like to thank all my co-authors, especially Prof. Chrit Moonen, Dr. Mario Ries, William Chu, Dr. Samuel Pichardo, and Dr. Steffen Hokland. I am grateful to my parents, Unto and Pirkko, as well as my parents- in-law, Jussi and Fia, for their support throughout the years. My deepest gratitude goes to my wife Saara and my sons Oliver and Liinus, whose love and support have kept me sane during this exciting, but long process.

Vantaa, March 21, 2017,

Matti Tillander

1 Preface

2 Contents

Preface 1

Contents 3

List of Publications 5

Author’s Contribution 7

1. Introduction 13 1.1 Background ...... 13 1.2 Aims of the Study ...... 15

2. High-Intensity Focused Ultrasound 17 2.1 Ultrasound in Medicine ...... 17 2.1.1 Ultrasound ...... 17 2.1.2 Ultrasound in Tissue ...... 17 2.2 High-Intensity Focused Ultrasound in Thermal Therapies . 19 2.2.1 Therapeutic Ultrasound ...... 19 2.2.2 Image-Guidance in HIFU ...... 19 2.2.3 Thermal Ablations with HIFU ...... 20 2.2.4 Local Hyperthermia with HIFU ...... 21 2.3 Generation of High-Intensity Focused Ultrasound Waves . . 21 2.3.1 Focusing Ultrasound Transducers ...... 21 2.3.2 Transducer Technologies ...... 23 2.3.3 Phased Array Transducers ...... 23 2.4 Simulations ...... 27 2.4.1 Acoustic Simulations ...... 27 2.4.2 Thermal Simulations ...... 28 2.5 Acoustic Field Measurements ...... 29

3 Contents

3. MRI-Guided High-Intensity Focused Ultrasound 31 3.1 Principles of MR-HIFU Therapy ...... 31 3.1.1 Magnetic Resonance Imaging (MRI) ...... 31 3.1.2 MR-HIFU Treatment Procedure ...... 32 3.2 MR-HIFU Temperature Mapping ...... 35 3.2.1 Temperature Mapping Sequence ...... 37 3.2.2 Limitations in PRF-based Thermometry ...... 38 3.3 MR-HIFU Systems ...... 40

4. Summary of The Publications 43 4.1 Publication I ...... 43 4.2 Publication II ...... 45 4.3 Publication III ...... 47 4.4 Publication IV ...... 48 4.5 Publication V ...... 50 4.6 Publication VI ...... 51

5. Conclusions 53

6. Future 55

References 57

Publications 73

4 List of Publications

This Thesis consists of an overview (chapters 1–5) and of the following Publications, which are referred to by their Roman numerals in the text.

I M. Tillander, S. Hokland, J. Koskela, H. Dam, N. P. Andersen, M. Ped- ersen, K. Tanderup, M. Ylihautala, and M. O. Köhler. High intensity focused ultrasound induced in vivo large volume hyperthermia under 3D MRI temperature control. Medical Physics, 43, 1539–1549, 2016.

II W. Chu, R. M. Staruch, S. Pichardo, M. Tillander, M. O. Köhler, Y. Huang, M. Ylihautala, M. McGufﬁn, G. Czarnota, K. Hynynen. Mag- netic resonance-guided high intensity focused ultrasound (MR-HIFU) hyperthermia for recurrent rectal cancer: MR thermometry evaluation and preclinical validation. International Journal of Radiation Oncology, Biology, Physics, 95, 1259–1267, 2016.

III A. Partanen, M. Tillander, P. Yarmolenko, B. Wood, M. Dreher, M. O. Köhler. Reduction of peak acoustic pressure and shaping of heated region by use of multifoci sonications in MR-guided high-intensity focused ultrasound mediated mild hyperthermia. Medical Physics, 40, 013301, 2013.

IV C. Mougenot, M. Tillander, J. Koskela, M. O. Köhler, C. Moonen, M. Ries. High intensity focused ultrasound with large aperture transducers: A MRI based focal point correction for tissue heterogeneity. Medical Physics, 39, 1936–1945, 2012.

5 List of Publications

V C. Bing, R. Staruch, M. Tillander, M. O. Köhler, C. Mougenot, M. Ylihau- tala, T. Laetsch, R. Chopra. Drift correction for accurate PRF-shift MR thermometry during mild hyperthermia treatments with MR-HIFU. In- ternational Journal of Hyperthermia, 32, 673–687, 2016.

VI M. Tillander, M. O. Köhler, J. Koskela, M. Ylihautala. Evaluation of intensity based beam-shaping method with rib-phantom HIFU sonications. In 12th International Symposium on Therapeutic Ultrasound, Heidelberg, Germany, AIP Conference Proceedings 1503, 191–194, 2012.

6 Author’s Contribution

Publication I: “High intensity focused ultrasound induced in vivo large volume hyperthermia under 3D MRI temperature control”

The author of this Thesis had the main responsibility for designing and implementing the hyperthermia application used in this study in a com- mercial MR-HIFU system. The author had a central role in planning the animal experiments. He also operated the system in each animal experiment and collected and analyzed the data. He was the principal writer of the manuscript.

Publication II: “Magnetic resonance-guided high intensity focused ultrasound (MR-HIFU) hyperthermia for recurrent rectal cancer: MR thermometry evaluation and preclinical validation”

The author further developed the hyperthermia implementation presented in Publication I to enable the feasibility experiments for recurrent rectal cancer hyperthermia. The author participated in planning the experiments, analyzing the data, and writing the manuscript.

Publication III: “Reduction of peak acoustic pressure and shaping of heated region by use of multifoci sonications in MR-guided high-intensity focused ultrasound mediated mild hyperthermia”

The author carried out the acoustic measurements and simulations, and participated in writing the manuscript.

7 Author’s Contribution

Publication IV: “High intensity focused ultrasound with large aperture transducers: A MRI based focal point correction for tissue heterogeneity”

The author developed a hydrophone measurement system for HIFU transducers. The system was used to verify the ability of an image-based method to correct the focal phase incoherences caused by heterogeneous acoustic paths in the context of breast sonications. The author participated in writing the manuscript.

Publication V: “Drift correction for accurate PRF-shift MR thermometry during mild hyperthermia treatments with MR-HIFU”

The hyperthermia heating solution and the magnetic ﬁeld drift correction method developed by the author in Publication I were further investigated in this study. The author participated in planning the experiments and writing the manuscript.

Publication VI: “Evaluation of intensity based beam-shaping method with rib-phantom HIFU sonications”

The author constructed the measurement setup, conducted the measurements, and had the main responsibility for analyzing the results. The author was the principal author of the manuscript.

8 List of Abbreviations

2D Two-dimensional 3D Three-dimensional ADC Apparent diffusion coefficient ARFI Acoustic radiation force impulse imaging CEM Cumulative equivalent dose at 43◦C CEM43 Same as CEM CE-MRI Contrast-enhanced magnetic resonance imaging CMUT Capacitive micromachined ultrasound transducer CT Computed tomography CW Continuous wave DNA Deoxyribonucleic acid EPI Echo planar imaging, an MRI imaging method FDTD Finite-difference time-domain, simulation method FEM Finite element method, simulation method FFE Fast field echo, an MRI imaging method HIFU High-intensity focused ultrasound HIU High-intensity ultrasound MR-HIFU Magnetic resonance imaging-guided HIFU MRI Magnetic resonance imaging NMR Nuclear magnetic resonance NPV Non-perfused volume PMUT Piezoelectric micromachined ultrasound transducer PRF Proton resonance frequency PZT Lead zirconate titanate, common piezoelectric material RF Radiofrequency SENSE Sensitivity encoding, an MRI imaging method SNR Signal-to-noise ratio TOF Time-of-flight TR Repetition time, an MRI scan parameter

9 List of Abbreviations

10 List of Symbols

A Surface area of transducer

B0 Main magnetic ﬁeld

Bloc Local magnetic ﬁeld G Focusing gain I Ultrasound intensity R Transducer radius of curvature

RI Acoustic reﬂection coefﬁcient for intensity

Rp Acoustic reﬂection coefﬁcient for pressure T Temperature

T1 Longitudinal relaxation time

T2 Transverse relaxation time

Tb Temperature of blood ◦ TD43 Thermal dose in equivalent minutes at 43 C

TE Time-to-echo

Tref Reference temperature Z Acoustic impedance a Transducer aperture c Speed of sound

cmuscle Speed of sound in muscle tissue

cb Speciﬁc heat capacity of blood

ct Speciﬁc heat capacity of tissue f Ultrasound frequency h Height of the transducer rim relative to the base i Imaginary unit k Combined wavenumber and attenuation

kt Thermal conductivity

l1 Distance from element 1 to desired focus location

ln Distance from element n to desired focus location

11 List of Symbols

p Ultrasound pressure r∗ Diameter of the focus at half intensity perpendicular to the beam axis r Vector from origin to desired focus location

r1 Vector from origin to ﬁrst transducer element

rn Vector from origin to nth transducer element t Time u Magnitude of the transducer surface velocity

wb Perfusion rate of blood z∗ Diameter of the focus at half of maximum intensity along beam axis ΔT Temperature change Φ MRI image phase α Ultrasound attenuation

αs Ultrasound scattering ◦ αT Temperature dependent water chemical shift in ppm/ C

αt Ultrasound absorption γ Gyromagnetic ratio δ Chemical shift

δ0 Temperature independent chemical shift component

δT Temperature dependent chemical shift component λ Ultrasound wavelength π Pi ρ Density

ρb Density of blood σ Shielding effect

φn Phase of transducer element n

12 1. Introduction

1.1 Background

High-intensity focused ultrasound (HIFU) is an emerging medical technology that can be used to treat different benign and malignant diseases non-invasively [1]. In HIFU, ultrasound waves are focused to form a localized area of high-intensity ultrasound (HIU) in tissue, which can cause rapid temperature elevation, typically within a very confined volume near the focus. The temperature elevation can be utilized therapeu- tically. For example, heating tissue over 57◦C causes irreversible tissue damage within seconds through thermal coagulative necrosis [2, 3]. With more moderate heating, up to 45◦C, several reversible biological effects, such as disruption of the DNA repair mechanisms, can be induced and utilized to enhance the effect of radiation therapy in cancer treatment [4, 5]. The first observations of HIU waves to cause biological effects were published by Langevin in 1917, when experiments with high power un- derwater ultrasonic transmitters for submarine detection killed the fish within the sound beam [6]. Later, the biophysical effects were studied more extensively by Wood and Loomis in 1927 [7]. In 1942, Lynn et al. [8] were able to demonstrate how HIFU waves caused localized change in tissue color from red to grey indicating thermally induced coagulative necrosis in fresh bovine liver tissue samples [9]. The surrounding tissues remained intact. Despite the tremendous potential of HIFU in localized non-invasive thermal therapies, the widespread clinical acceptance of HIFU has been restricted by the lack of proper spatial guidance and temperature control ability until recently [1, 10]. Magnetic resonance imaging (MRI) was first proposed as a method for guidance of HIFU treatments in 1992 [11]. Since then, HIFU has expe-

13 Introduction rienced a renaissance in both academic research and the medical indus- try. MRI provides volumetric imaging with excellent soft tissue contrast for targeting the treatments. MRI can be also used to assess the tissue changes caused by thermal therapy [12, 13], but most importantly, MRI provides a means of accurate and precise measurement of the tissue temperature changes non-invasively during heating [14, 15]. Temperature monitoring enables protection of sensitive structures adjacent to target region from overheating, which enhances the safety profile of a treatment. If the intention of the treatment is to induce irreversible tissue damage, the technical efficacy can be enhanced by terminating the sonication after sufficient thermal exposure has been obtained in the target volume. This reduces the unnecessary overheating of the tissue. MRI-guided HIFU, i.e., MR-HIFU, has been used clinically for treatments of e.g., benign and malignant breast tumors [16–20], uterine fibroids [21, 22], prostate cancer [23], liver cancer [24], and brain tumors [25, 26], and for palliative treatment of bone tumors [27–29]. In addition, using MR-HIFU for treatments of different neurological disorders like essential tremor [30,31], Parkinson’s disease [32], and neuropathic pain [33] has emerged recently. Research is also being conducted into using MR- HIFU-mediated mild hyperthermia for controlled drug release to spatially target the effect of chemotherapy to the intended location [34–36]. The number of MR-HIFU applications has expanded due to technolog- ical advances in ultrasound transducers, heating methods, and temperature imaging. For example, modern ultrasound transducers consist of several individual transducer elements, i.e., they utilize phased array technology. This enables, e.g., enlarging the heating volume by rapid focus steering [37,38] by manipulating the phases of the signals driving individual transducer elements. Alternatively, manipulating signal amplitudes can be used to avoid rib bones when targeting the liver [39–42]. More- over, commonly used MRI-based thermometry techniques provide good thermal accuracy and precision for a short period of time for stationary organs [43, 44]. However, new methods have recently been developed for temperature imaging of moving organs [45] and for performing the reliable long-duration thermometry [46] required in various MR-HIFU hyperthermia applications. In this Thesis, the capabilities of phased array transducers are explored for (i) controlling the near-field heating during a long-duration hyperthermia sonication, (ii) avoiding the rib-bones in sonications of liver using

14 Introduction an algorithm that is fast enough for clinical use, and (iii) restoring the acoustic wave phase coherence in the focus that is destroyed due to tissue heterogeneity. In addition, methods for compensating instabilities in the main magnetic ﬁeld and the changes in the magnetic ﬁeld caused by a moving transducer to enable reliable long-duration thermometry are investigated in the context of rectal cancer MR-HIFU hyperthermia.

1.2 Aims of the Study

Publication I: To implement large-volume hyperthermia application for an MR-HIFU system The aim of this study was to develop a new large-volume, long-duration hyperthermia heating method for an MR-HIFU system that combines mechanical transducer movements with electronic focus steering and, in addition, employs selective activation of transducer elements to control the near-ﬁeld heating. With the aid of multiple MRI thermometry slices the goal was to obtain uniform temperature in the whole target volume at clinically relevant depths from the skin.

Publication II: To show the feasibility of MR-HIFU system for radio-sensitizing mild hyperthermia of recurrent rectal cancer In this study the aim was to evaluate the feasibility of performing MR- HIFU-mediated mild hyperthermia in rectal cancer patients. In particular, the accessibility of rectal tumors to a MR-HIFU system, thermometry quality near the rectal wall, and safety were evaluated in animal experiments and volunteer imaging of patients with rectal cancer to enable initiation of a human trial in the near future.

Publication III: To implement a heating method that creates multiple simultaneous foci to reduce the pressure peaks during hyperthermia. The aim of this study was to demonstrate that a sonication using multiple simultaneous focal points reduces the acoustic peak pressure experienced by the tissue compared to that obtained using electronic focal steering, and still enables homogeneous heating for mild MR-HIFU hyperthermia.

15 Introduction

Publication IV: To develop an image-based algorithm for correcting phase-aberrations caused by heterogeneous acoustic paths In this study an MRI image-based method for estimating ultrasound propagation time from individual transducer elements to the focus was developed. The aim was to enable restoration of phase coherence in the focus after its destruction by heterogeneous propagation medium in the context of sonications into breast pathologies.

Publication V: To evaluate different methods for enabling proton resonance frequency shift-based temperature mapping in long duration MR-HIFU hyperthermia heating The aim of this study is to evaluate the ability of pre-scanning procedures, dynamic center frequency stabilization, and the 3D linear ﬁtting method used in Publications I and II to correct the slow drift in the magnetic ﬁeld and to enable reliable long-duration temperature mapping for MR-HIFU hyperthermia.

Publication VI: To examine whether simulated acoustic intensity predicts the heating on bone surface in MR-HIFU In this study the ability of incident intensities obtained from acoustic simulations to predict the heating on the rib bone surface was examined. The aim was to a investigate a method for transducer element deactivation that is fast enough to obtain quantitative reduction of rib bone heating in intercostal MR-HIFU sonications to liver during clinical workﬂow.

16 2. High-Intensity Focused Ultrasound

This chapter presents the principles of high-intensity focused ultrasound. The biological effects of ultrasound are covered brieﬂy, and the necessary background information about the HIFU technology is presented. Associ- ations between HIFU technology and the publications of this Thesis are discussed.

2.1 Ultrasound in Medicine

2.1.1 Ultrasound

Ultrasound waves are acoustic pressure waves with a frequency exceeding the human hearing range (>20 kHz). An ultrasound wave travels within the propagation medium at the speed of sound causing coherent displacement of particles, i.e., molecules. If the direction of particle displacement occurs parallel to the direction of the wave propagation, the wave is called longitudinal. In the case of shear waves, the displacement and the direction of wave propagation are perpendicular to each other. Other forms of wave propagation, such as surface waves, plate waves, Stoneley waves, and Love waves, exist in the interfaces of two media. Longitudinal waves are the dominant propagating wave mode in most medical ultrasonic applications. [47]

2.1.2 Ultrasound in Tissue

When ultrasound propagates within tissue, part of the energy is deposited into tissue, leading to temperature increase. The absorption rate, and therefore, the heating rate, increases with ultrasound intensity and frequency. [6]

17 High-Intensity Focused Ultrasound

The attenuation of ultrasound wave can be described by the equation:

p(x)=p0 exp(−αx), (2.1) where p is pressure at location x, p0 is initial pressure at x =0, and α is the attenuation coefficient of the wave. Attenuation of ultrasound consists of an absorption component, αt, and a scattering component, αs so that α = αt+αs. In biological fluids and soft tissues, such as muscle and liver, the attenuation is dominated by absorption, whereas in porous-like tissues, such as lung and bone, scattering is the dominant component. Attenuation varies hugely between tissue type. For example, in human muscle tissue at 1 MHz, the attenuation coefficient is approximately 1 dB/cm. In blood, the corresponding value is 0.2 dB/cm, while in bone, the attenuation is greater than 10 dB/cm. [47] When an ultrasound wave propagates through a heterogeneous medium, such as human tissue, it can experience multiple refractions and reflections. Refraction occurs at interfaces between two tissues that have different speed of sound, c. Reflections arise from the mismatch in the acoustic impedances (Z = cρ, where ρ is density) of the two tissues. The reflection coefficient for pressure amplitudes at an interface of two materials at normal incidence can be calculated from the equation:

Z1 − Z0 Rp = , (2.2) Z1 + Z0 where Z0 and Z1 are the acoustic impedances of the intial propagation medium and the second medium, respectively. The reflection coefficient 2 for intensity is given by RI = Rp. As an example, the muscle–bone interface reflects most of the ultrasound energy at normal incidence due to the high mismatch in the acoustic impedances. Since bones have also high attenuation, they typically form a "barrier" for ultrasound propagation. [6] In addition to an oscillatory component, the molecules of the propagation medium experience a constant displacement component. In elastic materials, this causes a force, i.e., "push" that can be measured in soft tissue with, for example, acoustic radiation force impulse imaging (ARFI) [48]. In fluids, such as urine and blood, the medium is set into motion, i.e., ultrasound waves induce streaming. At high acoustic intensities, the ultrasound wave propagation becomes increasingly non-linear, and a larger proportion of the energy is transferred to higher frequency components (harmonics), compared to the fundamental frequency [6]. This further increases the attenuation, which is typically greater at higher than lower

18 High-Intensity Focused Ultrasound frequencies. At high acoustic intensities, extreme rarefactional pressures may also induce bubble mechanisms, i.e., formation of small gas or vapor cavities (cavity nucleation) or volumetric pulsation of existing bubbles due to ultrasound waves (acoustic cavitation) [49]. Cavitation can contribute to disintegration of tissue [50]. Also using cavitation to enhance heating efﬁciency in HIFU treatments has been investigated, but due to the unpredictable nature of the phenomenon, cavitation is commonly avoided intentionally [49].

2.2 High-Intensity Focused Ultrasound in Thermal Therapies

2.2.1 Therapeutic Ultrasound

In therapeutic ultrasound, the intention is to achieve biological effects favorable to the intended therapy. The large spectrum of different therapeutic ultrasound applications utilizing both thermal and non-thermal mechanisms of ultrasound is presented by Hill et al. [6] and Miller et al. [51]. As an example, ultrasound has previously been used in phys- iotherapy to accelerate the healing of soft tissues and bone fractures [6], although, the efﬁcacy of the treatments has been questioned [52]. Phys- iotherapy typically employs unfocused beams and pulsed or continuous ultrasound wave with spatial average intensity levels up to 3 W/cm2 and frequencies up to 3 MHz [6]. In HIFU, the ultrasound waves are focused and the used intensities are much higher: spatial-peak temporal-average intensities may extend up to 20 kW/cm2, which induces rapid localized temperature elevation in tissue [1]. The work of this Thesis concentrates on the use of HIFU in thermal ablations and hyperthermia.

2.2.2 Image-Guidance in HIFU

HIFU is typically conducted in the guidance of either B-mode ultrasound imaging or MRI imaging. The ultrasound guidance often utilizes the growth of hyperechoic region at HIFU focus, which is a sign of bubble ac- tivity, such as boiling and cavitation [49]. MRI on the other hand provides a means of measuring tissue temperature to guide the HIFU treatments. The works of this Thesis all utilized the guidance of MRI temperature mapping, which is described in detail in Chapter 3.

19 High-Intensity Focused Ultrasound

2.2.3 Thermal Ablations with HIFU

In thermal ablation, tissue temperature is maintained at an elevated temperature long enough to cause irreversible tissue damage. High temperature causes acute localized bioeffects, such as denaturation and aggre- gation of proteins, loss of cell membrane integrity, mitochondrial dysfunc- tion, inhibition of DNA replication, and collapse of vasculature function, which may all lead to tissue necrosis. The delayed effects of thermal exposure (appearing within a couple of days) may contain systemic inﬂam- matory response, edema, and apoptosis of the cells in the periphal zone of the ablated region. Furthermore, damage to the vascular system may lead to delayed cell death due to ischemia. [53] The temperature threshold for irreversible damage varies with tissue type, but cell survival seems to obey a non-linear function of temperature and time. This has eventually led to the formulation of a dosimetric concept for thermal therapies: thermal dose [2, 3]. Thermal dose is used to predict the biological effects following heat exposure and to compare the results across thermal-therapy studies. Thermal dose can also be used to steer and monitor the progress of a single MR-HIFU thermal therapy event if evaluated continuously from MRI-based tissue temperature measurement during heating [54,55]. Thermal dose is deﬁned as follows:

t 43−T (r,τ) TD43(r,t)= R dτ, (2.3) 0 where R =0.25 when temperature, T (r,t), is less than 43◦C and R =0.5 at temperature 43◦C or higher. Thermal dose results are commonly reported in terms of equivalent minutes, i.e., as cumulative equivalent minutes at a reference temperature of 43◦C (annotations: CEM43 or CEM). By deﬁ- nition, keeping tissue at 43◦ for 240 minutes provides thermal dose of 240 CEM. The same effect and thermal dose is achieved in less than a second at a tissue temperature of 57◦C. The onset of thermal damage has been shown to occur in muscle tissue at 30 CEM, while 240 CEM is considered as the general threshold for ensuring complete tissue destruction for all soft tissue types [56, 57]. The thermal dose was calculated in the hyperthermia experiments presented in this Thesis to evaluate the probability of inducing thermal damage during the heating experiments.

20 High-Intensity Focused Ultrasound

2.2.4 Local Hyperthermia with HIFU

In hyperthermia, the tissue temperature is maintained above normal body temperature for a relatively long time period (typically 40–45◦C and up to 60 min). Instead of inducing direct cell death, the intention is to sen- sitize cancerous tissues to other forms of therapy [58]. Hyperthermia has been shown in several phase III trials to improve the efﬁcacy of both radiation therapy and chemotherapy [59–62]. Prolonged temperature elevation is known to increase tissue perfusion and oxygenation [63], inhibit DNA repair mechanisms [4, 5, 64], and increase vascular permeability to nanoparticles [65]. Hyperthermia is commonly divided into three categories: local, regional, and whole-body depending on the anatomical coverage the heating is applied to. In local hyperthermia, the intention is to limit the temperature elevation mainly to the targeted tumor. Traditionally local hyperthermia has been applied using external, intraluminal, or interstitial radiofrequency applicators [66]. However, HIFU provides an extraordinary technology for applying local hyperthermia since with focused ultrasound waves the heating can be very accurately localized and, importantly, in a non-invasive manner.

2.3 Generation of High-Intensity Focused Ultrasound Waves

2.3.1 Focusing Ultrasound Transducers

HIFU utilizes focusing of ultrasound waves to generate a small high- intensity focus at the desired target location. The focusing can be achieved efﬁciently by building the ultrasound transducer into a focusing shape, as shown in ﬁgure 2.1. The focusing gain, i.e., the pressure at the focus relative to the pressure at the transducer surface, is approximately G =2πh/λ, where h is the height of the rim of the transducer relative to its base and λ is the wavelength [47]. The acoustic wavelength is calculated from speed of sound c and frequency f as follows: c λ = . f (2.4)

The transducer used in the studies of this Thesis (except Publication IV) has h = 20 mm, approximately. Assuming a simple, single–element focusing transducer design and a sonication frequency of 1.2 MHz, a focus-

21 High-Intensity Focused Ultrasound

Figure 2.1. Image A illustrates how the ultrasound generated by a spherically shaped HIFU transducer propagates through the skin and intervening tissues to the targeted location. h is the height of the rim relative to the transducer base. Image B shows the components of the acoustic ﬁeld of a HIFU transducer.

ing gain of approximately 100 is obtained in muscle tissue (cmuscle = 1575 m/s [67], λ = 1.3 mm). For a simple focusing transducer the diameter of the focus at half intensity along the beam axis and perpendicular to the beam axis are z∗ = 0.886Rλ/h and r∗ =1.028Rλ/a, respectively, where R is the radius of curvature and a is the aperture [14]. With R = 140 mm and a = 120 mm, which are comparable to the transducer used in these studies, the focal point dimensions are 8 mm and 1.6 mm, respectively. The small dimensions of the focus allow localizing the heating very accurately in HIFU. The transducer aperture, radius of curvature, and the sonication frequency used in HIFU applications are designed with reference to the targeted anatomy and desired heating depth [68, 69]. The sonication frequency is typically selected based on desired target depth. Since ultrasound absorption increases with frequency, it is beneficial to use e.g., higher sonication frequency for shallow targets to optimize the heating. Hence, for transrectal or transurethral treatment of the prostate, sonication frequencies in the range of 3.0 – 9.7 MHz have been used [70, 71], while for deeper targets, such as uterine fibroids in the pelvic area, lower frequencies (1.0 – 1.5 MHz) have been employed [21, 22]. For brain sonications the use of lower sonication frequencies (0.7 MHz) help the propagation through the highly attenuating skull bone [72]. In Publications I and II the intention was to heat as deep as possible. Therefore, sonication frequencies of 800 kHz and 1.0 MHz were used, which are lower than the default frequencies (1.2 MHz and 1.45 MHz) of the Philips Sonalleve system that is designed for uterine fibroid treatments. In contrast, in Publications III and IV the intended targets were

22 High-Intensity Focused Ultrasound more shallow (rabbit muscle and a phantom mimicking human breast tissue), and therefore, a higher sonication frequency (1.45 MHz) was selected.

2.3.2 Transducer Technologies

Ultrasound waves are commonly created by applying an oscillating voltage at a radiofrequency to a piezoelectric material. The electric field through the piezoelectric material influences the piezoelectric domains in a manner that causes a stress in the material and possibly expansion and contraction in response to the applied RF voltage. Piezo electric phenom- ena were originally discovered by Jacques and Pierre Curie in 1880. PZT (lead zirconate titanate) is probably the most widely used piezoelectric material in medical ultrasound transducers. The high electromechan- ical coupling factor and machinability into various sizes and shapes make it attractive for various ultrasound applications. [6, 73] In medical ultrasound applications, a quarter-wavelength matching layer between the piezoelectric material and the propagation medium (typically water or soft tissue) can be used to compensate for the high acoustic impedance mismatch between the two materials. This improves the energy transmission from the transducer to the propagation medium. Fur- thermore, ablative and hyperthermic HIFU applications commonly use single frequency continuous-wave (CW) sonications and need to be able to deliver large amounts of energy. Using air-backing, i.e., air on the rear of the piezoelectric materiel, helps in maximizing the energy transfer in the forward direction. [6, 47] Recent developments in microelectromechanical manufacturing methods have opened opportunities for new transducer technologies, such as capacitive micromachined ultrasound transducers (CMUT) [74–76] and piezoelectric micromachined ultrasound transducers (PMUT) [77, 78]. However, these emerging technologies have not yet replaced piezoelectric materials in medical ultrasonic applications, although the potential of, e.g., CMUTs particularly in diagnostic ultrasound imaging has been considered significant [79].

2.3.3 Phased Array Transducers

An ultrasound transducer can consist of multiple transducer elements that are each driven with individual phase and amplitude. This phased

23 High-Intensity Focused Ultrasound

Figure 2.2. The left image shows a schematic illustration of the 1–3 piezocomposite struc- ture. The right image shows the cross-section of the transducer active surface. The size and shape of the phased array elements are defined by the conductor pattern on the back of the piezocomposite since only the piezo-rods that are under conducting surfaces experience and react to varying electric field. array transducer technology is used extensively in diagnostic imaging but has also become popular in therapeutic ultrasound. Phased array transducers can be built from a 1–3 piezocomposite materials where the piezoelectric material is cut into small rods, and the kerfs are filled with an electrically passive polymer (Fig. 2.2, left). Piezocom- posite provides many advantages over a transducer created from a single PZT block: piezocomposites have lower acoustic impedance, which makes creation of the matching layer easier, they are easy to cast into different shapes by using soft polymer in the kerfs, and propagation of unwanted plate waves on the transducer surface is limited. In addition, the element pattern of a phased array transducer, i.e., the number, size, and shape of the elements, can be created by engraving the pattern onto the conducting surface of the back side of the transducer (Fig. 2.2, right). The electrically passive polymer used between the piezoelectric rods also mitigates the cross-talk between elements. [80–82] The transducers used in the studies of this Thesis employed phased array technology and contained 256 individually controlled round elements.

Focal Point Steering One of the capabilities of phased array transducers is that the focus can be moved electronically by manipulating the phases of the electronic signals driving the elements. As shown in Figure 2.3, in order to obtain constructive interference at location r, the element phases are determined by l φ =2πf n , n c (2.5) where f is the sonication frequency, ln is the distance of the nth element to the desired steering location (ln = |r − rn|), and c is the speed of sound. Focal point steering enables the heating of larger volumes without mov-

24 High-Intensity Focused Ultrasound

Figure 2.3. In the left hand image all elements are driven with equal phase and the focus forms to the geometrical focus of the transducer. In the right hand image the element phases are adjusted so that the location of the constructive interference is lateral to the geometrical focus and thus the focus is displaced. l1 and ln represent the distances from the first (r1) and the nth (rn) element to the target focusing location (r), respectively. ing the transducer mechanically [37, 38]. It is also possible to split the focus into multiple simultaneous focal points in order to enlarge the heating area [83–85]. The multifoci approach reduces the applied peak pressure and thereby the probability of mechanical bioeffects in the tissue compared to a single moving focus. The use of the heating approach in long- duration hyperthermia in MRI-based temperature-mapping guidance is demonstrated in Publication III of this Thesis. The spatial distribution and size of the transducer elements on the surface of a focusing HIFU transducer has a significant effect on the transducer performance and especially on the focus steering capability. Dis- tributing the elements in a random order on the surface helps in reducing the magnitude of grating lobes, i.e., the secondary maxima within the acoustic field caused by constructive interference of side lobes of individual transducer elements [86–88]. Diminishing grating lobes is very important for achieving localized HIFU treatments without hot spots outside the target region. It is also advantageous to fill the surface with elements as densely as possible in order to maximize the output power of the transducer. Furthermore, the acoustic field of small elements is less directional, so their use improves the electronic focal steering range without significant reduction in the focus quality. However, using a large

25 High-Intensity Focused Ultrasound number of elements will require complex and expensive driving electron- ics, and so, commonly a compromise is made in the steering range of the transducer. For example, the transducer of Philips Sonalleve system for uterine ﬁbroid treatments uses 256 individual elements. This limits the deﬂection distances for electric steering to approximately to +/- 20 mm in the beam axis direction and +/- 10 mm perpendicular in the beam axis direction. In order to further expand the heating volume, which is limited by the focus steering capability of the transducer, it is possible to combine the electronic steering with mechanical transducer movements, as demonstrated in Publication I for hyperthermia sonications.

Correcting Phase Aberrations The ability of phased array transducers to control individual element phases can be used to restore the focal phase coherence that has been compromised by e.g., sub-optimal transducer manufacturing processes or heterogeneous tissue lying within the element beam paths. The latter is particularly pronounced in HIFU treatment of the brain, where varying density and thickness of the skull bone may seriously deteriorate the phase coherence of the propagating wave and cause the focus to shift, be distorted, or scatter into multiple foci [89–91]. Phase-aberration corrections are also beneﬁcial when using a wide-aperture transducer for sonications in heterogeneous tissue, such as breast, since there is more variation in the acoustic beam paths of different transducer elements than with to small aperture transducers [20, 92]. In this Thesis work, an image-based method for calculating the phase corrections in an MR-HIFU platform designed for breast sonications was developed (Publication IV).

Near Field Beam Shaping By selectively activating and deactivating transducer elements, it is possible to shape the acoustic beam of a phased array transducer. This is particularly useful if the near ﬁeld of the acoustic path contains sensitive structures, such as bones, bowel loops, or scars, and the acoustic intensity at those structures should be limited. Beam shaping has been used successfully in animal studies to reduce the heating of the rib bones during intercostal sonications to the liver [40, 93–95]. Preparations for starting a clinical trial for liver sonications that would require a beam-shaping functionality are ongoing [41, 42, 96]. In this Thesis, a new method for

26 High-Intensity Focused Ultrasound estimating the heating on the rib bones, which would be fast enough to be used clinically while the patient is on the treatment table, was examined in Publication VI. Beam shaping can also be used to even up the possible non-uniformities in the near-ﬁeld temperature pattern caused by e.g., tissue heterogeneities or non-uniformities in the transducer near-ﬁeld intensity pattern. A method for performing dynamic beam shaping during a hyperthermia sonication was developed and tested using in vivo animal experiments in Publi- cation I of this Thesis.

2.4 Simulations

2.4.1 Acoustic Simulations

Acoustic simulations can be utilized in designing new HIFU transducers and systems, in treatment planning, and even during the therapy for guiding the treatment. Acoustic simulations in HIFU often assume single harmonic linear wave propagation. Non-linear wave propagation that may appear in the focal region at high intensities and result in increase of the heating rate is often neglected to simplify the calculations [97–99]. The wave propagation of a harmonic wave in a homogeneous medium can be calculated using Rayleigh–Sommerfeld integral iρck eik|r−r |u p(r)= dA, (2.6) 2π A |r − r | where ρ is the tissue density, c is the speed of sound, k is the combined wave-number and attenuation (k =2π/λ − iα, where λ is the wavelength and α the attenuation coefficient), r is the coordinate vector of the pressure field, r is the coordinate vector of the transducer surface, u is the magnitude of the transducer surface velocity perpendicular to the surface, and A is the transducer surface area [85,100]. The Rayleigh–Sommerfeld integral can be solved very efficiently in homogeneous medium using the angular spectrum method [47]; approximative extensions to cover also inhomogeneous media have been proposed [101,102]. A simplified Rayleigh– Sommerfeld integral for focusing transducers was utilized in Publications I and VI. In Publication I, a simulation was used to forecast near-field heating to enable beam shaping to even up resultant temperature non- uniformities dynamically. In Publication VI, a simulation was used to estimate the heating of rib bones during intercostal sonications.

27 High-Intensity Focused Ultrasound

The acoustic ﬁeld in inhomogeneous media can be calculated by solving the general wave equation using, e.g., FEM or FDTD methods. However, these methods are computationally heavy in a complex geometry such as the human body. Recently, a method based on stochastic Monte Carlo method and ray tracing technique has been proposed to solve acoustic wave propagation in inhomogeneous media [103]. The method is suitable for simulations in complicated anatomies, and ﬁrst coarse estimates can be obtained very quickly, which makes it practical also for clinical practice commonly with time constraints. A ray tracer method was used in Publi- cation IV to simulate the effect of the heterogeneity in the acoustic paths to shape of the acoustic focus. In Publication VI, ray tracer simulations were used to test the correlation between incident energy density on bone surface and heating of the bone during sonications.

2.4.2 Thermal Simulations

Thermal simulations can be utilized for example in the development of HIFU heating strategies, such as the spiral trajectory heating proposed by Salomir et al. [37] and volumetric heating strategy proposed by Köhler et al. [38]. Thermal simulations for HIFU therapies can be conducted using the Pennes bioheat equation ∂T(r,t) ρc = ∇·(k ∇T (r,t)) − w c ρ (T (r,t) − T )+α I(r,t), t ∂t t b b b b t (2.7) where ρ and ct are the density and the speciﬁc heat for the tissue, respectively, and ρb and cb the corresponding values for the blood, T is the tissue temperature distribution, kt is the thermal conductivity of the tissue, wb is the blood perfusion rate, Tb is the temperature of the blood, αt is the tissue absorption, and I is the instantaneous acoustic intensity [104,105].

The heat source term, αtI(r,t), can be derived directly from the acoustic pressure ﬁeld map: |p(r)|2 I(r)= . ρc (2.8) The limitation of the bioheat model is that it does not take into account the heterogeneity of blood perfusion caused by large blood vessels. Fortu- nately, ablative HIFU sonications are relatively short so that heat conduc- tion effects are typically more dominant. Mahoney et al. [105] have, for example, investigated the agreement of the model with measurements. If the sonications are conducted in real-time MRI-based temperature- mapping guidance, the tissue temperature can be measured directly without need to rely solely on simulations.

28 High-Intensity Focused Ultrasound

The bioheat equation was used in Publication III to compare the temperature distributions obtained with multifoci heating and single steered focus heating.

2.5 Acoustic Field Measurements

Hydrophones are considered to be the gold standard method of charac- terizing HIFU transducer pressure ﬁelds. They are also used extensively when evaluating different HIFU-transducer phase-aberration correction methods [41, 90, 91, 94, 106]. A needle hydrophone measurement system was used in Publication III to characterize the pressure ﬁeld containing multiple simultaneous focal points and also in Publication IV to test the phase correction method for breast sonications. Hydrophones employ the piezoelectric effect in converting the acoustic signals to electric signals [107]. In order to improve the spatial measurement resolution and reduce the dependency of the hydrophone sensitivity on the propagation direction of the wave, the detector elements are typically made very small; hydrophones with a diameter of 40 μm have been manufactured. The hydrophones are typically either needle type or membrane type [6]. The measurements with HIFU transducers are typically performed at moderate power levels in order to avoid cavitation-induced damage [108]. Recently, optical hydrophones that are more durable with respect to cavitation have been developed [109]. These are also capable of measuring temperature and acoustic pressure simultaneously and can be placed in- terstitially. Hence, optical hydrophones could be potentially used for in situ validation of different acoustic and thermal simulation models. Needle hydrophones with a diameter of 75 μm were used in Publications III and IV of this Thesis.

29 High-Intensity Focused Ultrasound

30 3. MRI-Guided High-Intensity Focused Ultrasound

This chapter gives a brief introduction to magnetic resonance imaging and how it is used to guide HIFU treatments. The most commonly used temperature mapping method in HIFU, proton resonance frequency shift method (PRF), is discussed in detail. In addition, the association of the method with the publications of this Thesis and its limitations is described.

3.1 Principles of MR-HIFU Therapy

3.1.1 Magnetic Resonance Imaging (MRI)

MRI is a technique used extensively in diagnostic imaging. It is based on the phenomenon of nuclear magnetic resonance (NMR) where certain atomic nuclei in magnetic field absorb and re-emit radiofrequency energy. The nuclei utilized predominantly in medical applications are hydrogen nuclei, i.e., single protons, which are especially abundant in soft tissues and in fat. Conventional MRI devices use strong, static magnetic fields to augment the NMR phenomenon. Despite of its quantum-mechanical nature, NMR phenomenon in medical imaging can usually be described with a net magnetization vector, which originates from the atomic nuclei in a magnetic field. The magnetization vector can be manipulated for example by tilting it from its equilibrium state using RF pulses at Larmor frequency. Magnetization in non-equilibrium state can be detected using sensitive RF coils. The speed of exponential return of the magnetization back to the equilibrium state is defined by the relaxation time constant T1. In addition, the decay of mea- surable magnetization due to dephasing of the protons in non-equilibrium state is described by the exponential relaxation time constant T2. Both T1

31 MRI-Guided High-Intensity Focused Ultrasound and T2 are tissue speciﬁc and are used extensively in differentiating between pathological and normal tissue. MRI is known for its versatility and excellent soft tissue contrast. In its simplest form, MRI measures the density of protons. However, by varying the scanning parameters, contrast between tissues that have different

T1 and T2 can be obtained. Furthermore, it is possible to e.g., suppress fat content from images to help detect edemas, follow diffusion of water molecules in tissue, and examine tissue perfusion. Combined with excellent spatial resolution and the possibility of acquiring images in any plane orientation makes MRI an excellent imaging modality for differentiating pathological tissue from healthy tissue. Many of the MRI parameters that define the tissue contrast are also temperature-dependent, which enables the use of MRI imaging for spatial and temporal monitoring of temperature during treatments. Hence, the capabilities of MRI make it suitable for guiding thermal therapy treatments. MRI does not use ionizing radiation in contrast to computed tomography (CT), which utilizes X-rays, and it is considered a safe imaging modality. Adverse events are typically related to ferromagnetic objects, such as stents and pacemakers, which are influenced by the strong magnetic field, or keys and tools, which may become projectiles when brought too close to the MRI scanner. However, MRI imaging is usually slower and more expensive than for example CT imaging. A detailed descriptions of MRI technologies and imaging sequences for obtaining different contrasts are provided in e.g., Magnetic Resonance Imaging: Physical Principles and Sequence Design by Brown et al. [110].

3.1.2 MR-HIFU Treatment Procedure

An MR-HIFU treatment procedure typically comprises three phases that are performed in the MRI scanner: planning, treatment, and output veri- ﬁcation.

Planning The planning of MR-HIFU therapy can be performed immediately before starting the treatment; the patient is in treatment position on the table of the therapy device when MRI images of the target region are acquired

[22, 111]. T1 and T2-weighted images are typically used to discern the diseased tissue from healthy tissue and thereby target the treatment to right location [17, 28, 112]. MR-HIFU systems may have speciﬁc software

32 MRI-Guided High-Intensity Focused Ultrasound

Figure 3.1. Example of treatment planning in Philips Sonalleve MR-HIFU system for ablations of uterine fibroids. Planning is done by placing green ellipsoidal graphical objects on top of planning images. Each object is an individual sonication and the size represents the predicted treatment volume. These objects are called "cells" in the Sonalleve framework, but should not be mis- interpreted to mean living tissue cells. The overlay for the acoustic field of the transducer can be used to estimate the risk to sensitive structures, such as the bowels. The planning can utilize multiple imaging slices (left: sagittal, right: coronal). tools for guiding the planning, e.g., as shown in figure 3.1. Alignment of the coordinate system of the transducer with the MR images helps in spatially targeting the heating to the correct locations.

Treatment The ablative MR-HIFU treatment consists typically of a series of short (e.g., 20 s [113]) heating events, i.e., sonications, whose targeting is based on the acquired planning images (Fig. 3.1). In each sonication, the tissue is heated to a sufﬁciently high temperature (e.g. above 57 ◦C) to ensure coagulative necrosis of the tissue. Each sonication is followed by a cooling period, this is necessary to protect fat layer and skin from overheating [113–116]. The targeted volume is spatially populated with sonications in order to cover the entire target sufﬁciently. The population of sonications can be planned e.g., using a "one-layer"-strategy proposed by Kim et al. [117]. This means that all sonications are placed in one plane normal to the beam axis. Prior to each sonication, it is a common practice to conduct a "test shot". This is a low-power sonication that does not induce irreversible damage to the tissue but can be used to verify the correct heating location [112,118]. The test shot can help correct small displacement errors in the heating due to e.g., heterogeneity in the acoustic path. In addition, the thermal response of the test shot can be used to predict the required power level for

33 MRI-Guided High-Intensity Focused Ultrasound subsequent therapy sonications [119]. This depends, e.g., on the material properties within the acoustic path and blood perfusion around the target, which acts as a heat sink. In contrast to ablative thermal therapy, the treatment procedure for MR-HIFU hyperthermia has not yet been established. This will probably be delivered using a single, or a few, long-duration sonications, as demonstrated in Publications I, II, III, and V of this Thesis.

Outcome Veriﬁcation MRI provides various options for assessing the outcome of MR-HIFU thermal therapies. The most common method in ablative therapy of uterine ﬁbroids is contrast–agent enhanced imaging (CE-MRI), which is typically acquired immediately after the desired treatment volume has been covered with sonications [21, 112, 120, 121]. High temperatures cause occlusion of the blood vessels in the target area and thereby reduce uptake of the contrast agent [53, 122–126]. The contrast agent decreases the T1 relaxation of the tissue and regions with reduced uptake are seen as dark in T1-weighted images. The dark regions, i.e., non-perfused volumes (NPVs), are known to correlate well with histology-proven coagulative necrosis [125,127–129]. NPV right after treatment overestimates the necrosed volume slightly compared to histology, but an excellent match has been reported after 7 days [130]. Destruction or temporary occlusion of a large blood vessel during treatment can also extend the NPV outside the treated volume via ischemia [53,131, 132]. The disadvantage of CE-MRI is that it can be performed only after treatment; the contrast agent typically used, gadolinium, biases PRF thermometry and it may also dissolve to toxic compounds at high temperatures [133, 134].

Both T1 and T2-weighted images have also been used to assess the MR- HIFU outcome, although with inconsistent results. E.g., the treated region has been reported to become hypointense, hyperintense, or to show heterogeneous intensity change in T1-weighed images, depending on the study. T1 and T2 are probably sensitive to multiple counteracting bioeffects induced by high-intensity ultrasound waves, which complicates their use in outcome assessment. [53] Diffusion imaging can potentially detect changes in cellular and nuclear membranes caused by HIFU heating [53]. However, studies with post- treatment diffusion imaging of uterine ﬁbroids have again given inconsis-

34 MRI-Guided High-Intensity Focused Ultrasound tent results [132]. Hector et al. [53] speculate that the outcome is influenced by the aggressiveness of the heating method: treatments employ- ing very high intensities cause both thermal and mechanical effects and disrupt the cell membrane. This further causes an increase in apparent diffusion coefficient (ADC) of the tissue. However, in more conservative heating, without mechanical effects, the membrane stays intact and a decrease in ADC is observed, possibly due to formation of edema. A more recent method of assessing the tissue changes caused by thermal ablations is MRI-based elastography [135–137]. Slowly propagating shear waves are introduced into the tissue and imaged using a motion- sensitive MRI sequence to deduce variations in tissue stiffness indirectly. However, although increase in tissue stiffness was observed in necrotic ex vivo tissue after thermal exposure, in vivo experiments showed decrease in the stiffness [138]. The conclusion was that tissue stiffness was more sensitive to the presence of edema than coagulative necrosis. Measuring the biological changes caused by hyperthermia can be chal- lenging. The hyperthermia-induced effects are intended to be reversible, and the biological changes that enhance radio or chemotherapy may be so small that they cannot be detected with MRI. For example, the effect of mild hyperthermia on dynamic contrast-agent imaging (DCE-MRI) and intra-voxel incoherent motion led to inconclusive outcome [139]. Monitor- ing of hyperthermia-mediated drug release has been investigated by en- capsulating MR contrast agent (gadolinium) in temperature sensitive liposomes, together with the chemotherapy drug. When the contrast agent is within the liposome, the effect on tissue relaxation parameters is in- significant. However, when the carrier liposome disrupts at certain temperature, releasing the drug and also the contrast agent, a clear difference is seen in the relaxation parameters and thereby in the image contrast. [140,141]

3.2 MR-HIFU Temperature Mapping

Many of the MRI parameters, such as the proton density, the T1 and T2 relaxation coefﬁcients, the diffusion coefﬁcient, magnetization transfer, and the proton resonance frequency (PRF) are sensitive to temperature changes [44]. The most common method for temperature mapping in MR- HIFU applications is PRF, which was also the method used in all publications of this Thesis. PRF is advantageous over other methods since it

35 MRI-Guided High-Intensity Focused Ultrasound has rather strong linear dependency on temperature in the thermal range relevant to HIFU heating, it works for aqueous soft tissues and it is relatively independent of tissue type and also unaffected by coagulation, and finally, it can be imaged with fast gradient echo sequences enabling control of rapid temperature changes induced by HIFU [44, 142, 143]. One disadvantage of the PRF method is that it does not work for adipose tissues since the PRF of fat protons is almost independent of temperature [144]. To avoid temperature-mapping errors in tissues that contain both water and fat, the fat signal can be suppressed without a major penalty to the imaging speed for example by using frequency selective slice excitations [145]. In the publications of this Thesis fat suppression was done using 121-binomial water-selective excitation where the excitation is split into three RF pulses. The known resonance frequency difference between water and fat protons is utilized in designing the pulse sequence so that fat protons are left mostly unaffected by the excitation. Since tissue susceptibility also changes with temperature, it can intro- duce errors to PRF temperature mapping. In water-rich tissue, such as muscle, the errors due to susceptibility changes remain within 10% and are therefore usually neglected [44]. However, in fat-rich breast tissue the errors might be significant [146]. Due to the limitations of the PRF method, alternative temperature mapping methods for fat have been developed [44,147–149] The temperature dependence of PRF was originally observed by Hind- mann in 1966 [150]. Use of the PRF phenomenon for MRI-based temperature mapping was proposed by Ishihara et al. in 1995 [15]. PRF temperature mapping is based on the effect a temperature change has on electronic shielding of protons. Increasing tissue temperature also increases the shielding effect, σ, thereby decreasing the local magnetic field experienced by the protons. The local magnetic field can be expressed as follows:

Bloc =(1− σ(T ))B0 =(1+δ(T ))B0, (3.1) where B0 is the strength of the main magnetic ﬁeld and δ(T ) describes the change in resonance frequency, i.e., chemical shift. [43]

The chemical shift has a temperature dependent component, δT (T ), and a temperature independent component, δ0, caused by B0 ﬁeld inhomo- geneities. The temperature dependence has been found to be linear in the temperature range relevant to MR-HIFU heating, so the chemical shift

36 MRI-Guided High-Intensity Focused Ultrasound can be written as

δ(T )=δ0 + δT (T )=δ0 + αT T, (3.2)

◦ where αT is the temperature dependency coefﬁcient in ppm/ C [44]. Most of the in vivo calibrations of the coefﬁcient have given values between −0.1 and −0.09 ppm/◦C in aqueous soft tissues [143]. During MRI, the chemical shift affects the phase of the protons accord- ing to equation

Φ(T )=γδ(T )TEB0, (3.3) where γ is the gyromagnetic ratio (42.58 MHz/T) and TE is the echo time of the imaging sequence. The effect of the chemical shift can be seen in the gradient echo-based phase images. In order to remove the temperature independent chemical shift component from the phase images, a subtraction method is commonly used. A reference phase map acquired at known temperature, Tref , is subtracted from the phase map acquired at temperature T to obtain the temperature change:

Φ(T ) − Φ(Tref ) ΔT = T − Tref = . (3.4) γαT TEB0

Hence, the PRF temperature mapping method is capable of showing relative but not absolute temperatures. In MR-HIFU treatments using the PRF method, the body temperature is typically measured before starting the therapy and added to ΔT to obtain absolute tissue temperature [15]. More information on the PRF and other MRI-based temperature mapping techniques can be found in comprehensive reviews by Quesson et al. [43], de Senneville et al. [151], and Rieke et al. [44].

3.2.1 Temperature Mapping Sequence

The temperature mapping sequence used for PRF-based monitoring of MR-HIFU heating needs to have adequate spatial and temporal resolution, spatial coverage, and SNR. Fast gradient echo-based imaging methods, such as FFE (fast ﬁeld echo), have been shown to be suitable for monitoring MR-HIFU treatments [16, 21, 152]. To increase the imaging speed FFE can also be combined with parallel imaging (e.g., SENSE) or with echo-planar (EPI), which utilizes fast reversal of readout gradient to acquire multiple echoes on a single excitation [44]. RF spoiling is commonly used to suppress the stimulated echoes caused by very short RF pulse repetition time (TR) [43, 145, 153]. In order to increase the coverage of the temperature mapping, a multi-slice monitoring approach can

37 MRI-Guided High-Intensity Focused Ultrasound be used [38]. This also improves the safety of the treatment, as some of the slices can be used to monitor, e.g., the temperature-sensitive regions in the near ﬁeld. The voxel size of temperature mapping sequences will always be ﬁnite, and therefore, MRI-based temperature mapping methods suffer from par- tial volume effect. This might be problematic for example in the presence of steep spatial temperature gradients. Moreover, comparison to inserted thermocouple or optical probe, which are typically very small, might show discrepancies. In all publications of this Thesis, the temperature mapping was performed using the PRF method and an FFE-EPI, multi-slice, RF-spoiled imaging sequence. The imaging parameters varied between publications. For example, in Publications I and II, where good spatial coverage was important, 6 temperature mapping slices were acquired with spatial resolution of 2.5 mm × 2.5 mm × 7.0 mm and update interval per slice of 0.5 s. On the other hand, in Publications III, IV, and VI, obtaining small voxel size, such as 1.0 mm × 1.0 mm × 4.0 mm in Publication IV, for a plane at focal depth was more important and only 1 or 2 slices were acquired.

3.2.2 Limitations in PRF-based Thermometry

Heating History The reference images for the PRF temperature mapping have to be acquired at a stable baseline temperature, otherwise the previously heated regions would cool down during the sonication and would show a drop in temperature below the body temperature. In ablative sonications, the cool-down period after each sonication protects the fat layer and the skin from overheating and gives time also for the target region to return back close to the baseline temperature [114–116].

Drift of The Magnetic Field

The B0 magnetic ﬁeld of all modern MRI scanners drifts slowly over time due to heating of the scanner components. This particularly occurs with imaging sequences that use gradients intensively, such as EPI sequences [46]. The PRF depends linearly on the magnetic ﬁeld and therefore the drift directly affects the temperature-mapping accuracy. In hyperthermia sonications, where the desired temperature range is narrow (40–45◦C) and the sonication time is long (up to 60 min), this drift must be corrected

38 MRI-Guided High-Intensity Focused Ultrasound since it may be tens of degrees per hour [46]. A common approach to drift correction is to select reference temperature regions in the images that are not heated, and are assumed to stay at constant temperature throughout the treatment. The reference regions can be, e.g., tissues outside the target area [36] or external reference tubes filled with aquaous gel or oil [46, 144, 154], which stay at a constant temperature during the hyperthermia procedure. In the reference regions the apparent temperature change is assumed to be due solely to the drift in the magnetic field. This information can be used to remove the drift component from the temperature maps of the heated region. The magnetic field drift may have also a spatially varying component, which makes the magnetic field gradually more inhomogeneous. An improved correction can be made, e.g., by making a polynomial fit to the data in unheated regions instead of a global correction. A linear fitting method for drift correction was implemented for Publication I and used in Publications II and V. The implementation used data from several temperature-mapping slices simultaneously to do the correction in 3D. In Publication III, a zeroth- order correction (spatially invariant) was conducted separately for each slice acquired. In Publication V, different approaches for making drift corrections in MR-HIFU applications were examined. E.g., dynamic stabilization of the center frequency of the MRI scanner was examined for also correcting image shifts that may appear with strong drifts.

Moving Organs Organ motion may lead to severe errors in PRF-based temperature mapping. The images acquired may lose alignment with the reference images and lead to impairment of the voxel-by-voxel phase-image subtraction. If the movement is periodic, e.g., due to respiratory motion, gating of the imaging using MRI navigator echoes can be used to eliminate motion ar- tifacts [155]. Alternatively, a multi-baseline approach can be used where a collection of PRF reference images, i.e., an atlas, is acquired to represent the different phases of the respiratory motion [156]. An atlas-based approach has been used also in PRF thermometry in the breast [157]. The organ itself does not move, but the changing volume of the lungs induces magnetic ﬁeld changes that extend also to the breast. PRF-based thermometry is problematic in organs with non-periodic motion, such as the bowel. Temperature mapping methods that do not need a separate reference image, but use the tissue around the heating area

39 MRI-Guided High-Intensity Focused Ultrasound to estimate the reference phase, have been developed. However, the magnetic field in the surrounding regions need to be homogeneous enough and therefore using these methods in the presence of organ interfaces near the heated volume may be difficult. [151] In Publication II, rectal filling was used to stabilize the peristaltic motion of the rectum in volunteers with rectal cancer. The purpose was to examine whether motion-free PRF-based thermometry can be carried out near the rectal wall despite the tendency of the rectum to transient movements. The procedure was used, in addition to long duration hyperthermia and temperature mapping methods developed in Publications I and II, in the first clinical MR-HIFU hyperthermia treatment of recurrent rectal cancer [158].

Moving Transducer A moving transducer can cause similar effects to the magnetic field as the changing lung volume in breast sonications, due to the transducer susceptibility. This renders the electronic focus steering and multifoci approaches advantageous over mechanical transducer scanning in enlarging of the sonication volume when PRF-based temperature control is used. In Publication I, the heating volume was extended beyond the focus-steering capabilities of the phased array transducer by adding mechanical transducer movements. The effect of the transducer movement in the magnetic field was eliminated by first obtaining a reference image atlas for each transducer position. Prior to starting heating, the transducer is moved through all the mechanical positions to be used during heating, and reference images are acquired at each position.

3.3 MR-HIFU Systems

The two commercially available MR-HIFU systems for treatment of uterine ﬁbroids are the Insightec Exablate 2000 and the Philips Sonalleve. They both have a phased array transducer inside an MR-compatible treatment table. The transducers are attached to motorized translation systems to extend the treatment reach without the need to move the patient during the therapy. Both systems utilize PRF-based temperature mapping for monitoring the temperature elevation [21, 22]. The volumetric heating approach of the Sonalleve system provides heating volumes lateral to the beam axis that are twice as large as those from the Exab-

40 MRI-Guided High-Intensity Focused Ultrasound late, while in the beam-axis direction the heated volumes are comparable [38,159].

41 MRI-Guided High-Intensity Focused Ultrasound

42 4. Summary of The Publications

This chapter brieﬂy summarizes the main results of the publications of this Thesis.

4.1 Publication I

Publication I describes implementation of a sonication method for long- duration (60 min) hyperthermia in the Philips Sonalleve MR-HIFU system that combined electronic focal-point steering with mechanical transducer movement to enlarge the heating volume (Fig. 4.1 (a)). This increases the accessible treatment volume substantially (three-fold increase in diameter lateral to beam axis) compared to prior art, e.g., heating with either mechanical movement only [35] or with electronic steering only [36]. A similar combined heating approach for MR-HIFU hyperthermia was recently proposed by Yarmolenko et al. [160]. However, the key advantage of the method proposed in Publication I is that the heating is controlled throughout the entire beam path. In contrast, in the study by Yarmolenko et al. the heating was controlled using only one temperature- mapping slice in the focal plane (Ari Partanen, personal communication, Nov. 2016), which increased the risk of overheating in the near field being unnoticed. Therefore, Publication I advances the state of the art by enhancing safety. In Publication I, the hyperthermic heating was primarily controlled using two temperature-mapping slices in MRI (slice thickness: 7 mm). The first slice was located at the depth of the focus in a plane parallel to the beam axis. The second slice was in the near field, parallel to the first slice. Electronic focal point steering was used to control heating in the focal plane and selective use of transducer elements, i.e., beam shaping, in the near field. In addition, three parallel temperature-mapping slices

43 Summary of The Publications

Figure 4.1. (a) A schematic representation of multiposition sonication cells in a plane parallel with the beam axis. The sonication is divided into seven transducer positions arranged on a hexagonal grid with one position in the middle. In each focal position, steering is used. (b) An example of measured focal-plane temperature map in in-vivo porcine muscle tissue during heating (sonication cell grid overlayed in magenta). (c) Beam axis view of one hyperthermia sonication. The contours represent the minutes the area was within hyperthermia temperature range (41–45◦C) during sonication. (Modified from Publi- cation I) at different depths were used for secondary temperature control; the sonication was temporarily terminated if too high temperature (45◦C) was achieved or exceeded in any of the slices. Temperature maps were up- dated every 3.2 seconds and provided feedback for the heating algorithm to decide (i) the mechanical transducer position, (ii) the electronic steering points, and (iii) which transducer elements are active. This thermal control strategy is unique in the field of MR-HIFU. Moreover, the study implemented a safe hyperthermia heating method for clinical use that is capable of monitoring and controlling the temperature over the entire beam path, not just in the focal plane. Reference images for PRF temperature-mapping were collected for each transducer position prior to heating. The drift in the magnetic field was compensated for using a 3D linear fit to the temperature-map voxels outside the beam path. The voxels were assumed to remain at constant temperature, and the apparent temperature change was estimated to be caused merely by the drift. The heating method was tested in 11 sonications of porcine thigh muscle in vivo. The sonications gave excellent temperature control for the target area in the focal plane; spatial and temporal average temperature, over all 11 sonications, was 42.0±0.6◦C (mean±SD) when the target temperature was 42.5◦C. In the near field the average temperature was lower, 39.3±0.8◦C, which indicates that even more energy could have been applied for heating, e.g., larger volumes or deeper targets while not compro-

44 Summary of The Publications mising safety. Alternatively, using a higher sonication frequency would have deposited more energy in the near field. The hyperthermia sonication frequencies used were 800 kHz and 1 MHz, which are lower than those normally used in ablative treatments with the Sonalleve (1.2 MHz and 1.4 MHz). The lower frequencies were selected in order to test the maximum attainable heating depth. The shape of the heating volume represented the shape of the focusing acoustic field (Fig. 4.1 (c)). The greatest heating volume achieved was over 100 ml, obtained with a 44 mm diameter heating cell. One experiment with a 58 mm heating cell failed due to overheating of a bone in the edge of the beam path in the near field. This exemplifies how large heating cells require a wide unobstructed acoustic window in the near field, which may become a lim- iting factor for the achievable hyperthermia heating volume. The developed time-in-range (TIR) concept was used to quantify the volume of the tissue exposed to hyperthermia. TIR is the cumulative time each voxel stays within hyperthermia range 41–45◦C. Successful hyperthermic heating was performed with the center of the heating cell set at 80 mm below the skin. The temperature mapping slices used to monitor and control the heating did not show high temperatures that could have led to thermal damage. Furthermore, CE-MRI images taken after sonication experiments showed no signs of thermal damage. Hence, the large volume hyperthermia heating was considered safe throughout the entire acoustic beam path, as is essential for clinical applications. In literature, most MR-HIFU hyperthermia experiments have been conducted in small animals, such as rabbits and mice. However, in this study a large animal model (pig) was used, and the MR-HIFU hyperthermia was extended to heating volumes and depths relevant to clinical use. This study therefore advances the field towards the clinical use of MR-HIFU hyperthermia.

4.2 Publication II

In Publication II, the feasibility of the MR-HIFU mild hyperthermia implementation of Publication I was evaluated for the treatment of recurrent rectal cancer. In order to investigate the temperature proﬁle, safety, and accessible heating depth in muscle tissue and near the rectum, a series of in vivo animal experiments was conducted. The accessibility of rectal

45 Summary of The Publications

Figure 4.2. Coronal (left) and sagittal (center) temperature maps for a hyperthermia sonication near the porcine rectal wall. The plot on the right shows the temperature evolution of spatial average, spatial T10, and spatial T90 temperatures for the target area at focal depth. (Modified from Publication II) cancer targets to MR-HIFU heating was evaluated in an imaging study with human volunteers with pre-existing rectal cancer. In addition, the ability of rectal filling to stabilize peristaltic motion and to enable precise PRF-based thermometry near rectum was examined in both animal subjects and in human volunteers. The animal experiments demonstrated that hyperthermia in temperature range of 41–43◦C can be safely achieved in conditions that represent recurrent rectal cancer. CE-MRI images and autopsy showed no signs of thermal damage when the maximum temperature within the acoustic path was limited to 43◦C. Heating was performed with the center of the cell set at a maximum depth of 80 mm from the skin. The imaging experiments with human volunteers demonstrated that an 80–mm reach from the skin is adequate for accessing rectal cancer targets. The heating volume will be limited by the available acoustic window, i.e., the space between coccyx and ischium bones. Rectal filling stabilized peristaltic motion and enabled precise temperature mapping (approx. 1◦C) in both animal experiments and volunteer imaging. Without rectal filling, the temperature mapping precision was not on an acceptable level (7.8 ± 7.8 ◦C) in volunteer imaging. The temperature mapping accuracy (approx. 1◦C) was verified against inserted optical thermosensor in animal experiments. Publication II demonstrated the feasibility of MR-HIFU hyperthermia in recurrent rectal cancer targets: the target is accessible for HIFU heating, and reliable temperature mapping can be performed. These enable safe and effective heating. To the authors’ knowledge, no similar feasibility study for MR-HIFU hyperthermia has been performed in any previous clinical application. This study enabled initiation of a clinical trial and treatment of the first recurrent rectal cancer patient with MR-HIFU hy-

46 Summary of The Publications perthermia, outside the remit of this Thesis [158].

4.3 Publication III

Figure 4.3. Simulated acoustic pressure ﬁeld for multifocal heating (right) and single steered focus heating (left) in a plane parallel with the beam axis, i.e., coronal plane. (Modiﬁed from Publication III)

In Publication III, a MR-HIFU hyperthermia heating method utilizing multiple simultaneous focal points was implemented and compared to heating with a single focus that is moved using electronic steering (Fig. 4.3). The multifoci approach is a well-known technique for enlarging the heating region in HIFU [83, 161], but in ablative therapies it has been replaced by, e.g., the more versatile volumetric heating method [38, 117]. However, quantitative comparison between multifoci and single steered focal heating approaches in hyperthermia has not been investigated sys- tematically in earlier literature. The motivation in this study was to reduce the acoustic pressure peaks (positive and negative) experienced by the tissue so that the risk of mechanical damage, such as that caused by cavitation, during hyperthermia heating could be mitigated. Control- ling the biological effects is essential in translating the medical technologies developed to clinical use. Also, a comparison between multifoci and steered-focus heating with regard to heating-region shape and temperature uniformity was carried out. The pressure ﬁeld from a multifocal sonication was characterized with hydrophone measurements. The results were compared against acoustic simulations using the Rayleigh–Sommerfeld integral. Heating experiments were conducted in phantoms and animals in vivo. The maximum acoustic pressure amplitude (at the fundamental soni-

47 Summary of The Publications cation frequency) was reduced by approximately 70% in the multifocal approach compared to that for the single electronically steered focus, in acoustic simulations. Excellent agreement was obtained between simulations and hydrophone measurement. In heating experiments in VX2 tumor, the diameter of the heated area in focal plane, i.e., lateral to the beam axis, was comparable for multifocal and steered focus approaches (difference: 1.25%). In the beam-axis direction the length of the heating volume was 35% shorter with the multifoci heating approach. No thermal or mechanical tissue damage was visible in the dissected tissue, suggest- ing the heating was thermally safe. The multifocal hyperthermia heating approach introduced in Publica- tion III advances the ﬁeld of MR-HIFU hyperthermia by providing a heating method with comparable heating capabilities to a steered focus while signiﬁcantly reducing maximum acoustic pressure, which is a risk factor for deleterious cavitation.

4.4 Publication IV

In Publication IV, a method for correcting transducer-element phase incoherence in the focus was described in the context of sonications for breast pathologies. In breast tissue, the phase incoherence is due to differences in the acoustic paths of individual transducer elements. Each path contains a unique combination of fat, glandular, and ﬁbrous tissue, which each have a different speed of sound. Therefore the time-of-ﬂight (TOF), i.e., the propagation time of ultrasound from transducer element to focus, varies between elements. Large phase incoherence may make the focus less localized, or may shift or split the focus [89–91]. The correction method was developed for a specialized Sonalleve breast MR-HIFU system. In the correction method developed, the main tissue components are segmented from the MRI images, and each component is given a speed of sound value in order to calculate the TOFs from each element to the target, in order that each wave component can have the same phase at the focus. Refraction is neglected. A heterogeneous two-component phantom (Fig. 4.4) was constructed to test the method. The two components simulated the speed of sound of glandular tissue and fat. Corrections were compared to hydrophone measurements, which are considered the gold standard for phase aberration

48 Summary of The Publications

Figure 4.4. The left hand image shows an MRI image of the heterogeneous breast phantom. The bright areas are polymer material (c = 1528 m/s) and grey areas are agar gel (c = 1485 m/s). The right hand image shows the location of transducer elements relative to the phantom. The system used was a dedicated breast MR-HIFU system [92] that used a wide aperture transducer to protect thoracic cage from heating during sonications. (Modified from Publication IV) measurements in HIFU [41, 90, 91, 94, 106]. The effect of hydrophone based and MRI-based TOF corrections on the maximum temperature and heating-region shape was examined in phantom sonications. The hydrophone measurements were performed with a needle hydrophone attached to a positioning system consisting of three linear stages. The hydrophone was moved along a predefined rectangular measurement grid. The movement was continuous, and while passing each grid point, generation of a new pulse from the transducer was triggered. The continuous motion enabled gathering pressure field maps quickly while keeping the related positioning errors negligible (less than 1 μm). In the heating experiments, the increase in focal temperature was on average 27.7 ± 1.1◦C (standard deviation) without any phase corrections. With MRI and hydrophone-based corrections, the temperatures were 34.0 ± 0.9◦C and 39.2 ± 2.7◦C, respectively. The drive signal amplitudes to the elements were kept constant during the experiments. The full-width half-maximum of the focal temperature was widest for the uncorrected sonications, indicating that phase incoherence spreads the heating in the focal region, making it less localized. The image-based correction therefore made the focus sharper and increased the maximum focal temperature substantially. The correction method presented was successful in a simple heteroge- nous phantom, but clinical use brings additional difficulties that need to be considered. For example, the heterogeneity of breast tissue is far more fine-grained, and segmenting tissue components may be more challeng-

49 Summary of The Publications

Figure 4.5. Comparison between optical thermometer reading and MRI thermometry with different drift correction in a phantom sonication. Images A and B refer to optical thermosensors 1 and 2. DS refers to dynamic stabilization only, DC+3D to first order polynomial fit correction method used in Publications I and II, and DS+DC+3D to combination of these two. 10 minute preheating was used in all experiments. (Modified from Publication V) ing. Perhaps MRI methods, such as that of Dixon [162], which is able to differentiate between water and fat, could be used in the process. Also, inter-patient speed of sound variation of different tissue components is currently not known. Finally, the speed of sound of the tumor itself was not considered in the experiments. This could affect the correction. The phase incoherence correction method was implemented into a MR- HIFU system designed for breast treatments. The same MR-HIFU system was used successfully in a phase I clinical trial for the treatment of breast cancer [20], although this correction method was not used during the trial. The breast tissue heterogeneity did not prevent the formation of the focus and the production of localized heating. This indicates that some degree of phase incoherence is acceptable for breast sonications. On the contrary, in neurological applications, the sensitivity of adjacent structures sets higher safety requirements for the localization accuracy, and therefore phase coherence is a necessity.

4.5 Publication V

In Publication V, different methods for improving the accuracy and stability of long-duration PRF-based temperature mapping for MR-HIFU applications were investigated. The aim was to correct the slow temporal drift of the magnetic ﬁeld strength that arises from heating of the MRI scanner components [46]. First, a 10-minute pre-scanning period was used to produce a temperature steady state to reduce the drift during a subse-

50 Summary of The Publications

Figure 4.6. Simulated energy density on bone surface with (right) and without (left) selective elements shutdown, obtained from ray-tracer simulation. quent imaging. Second, a dynamic stabilization was used prior to image acquisition to compensate for the change in the MRI scanner center resonance frequency, which changes with the magnetic field drift. Finally, these two methods were combined with the drift correction methods presented in Publication I and used in Publication II. The accuracy of the corrections was examined in phantom sonications and compared against readings from optical thermosensors. Validation of the methods was conducted in in vivo animal experiments. The pre-scan period reduced the variation in the magnetic field drift in the beginning of a temperature mapping scan. Combining the pre-scan period with dynamic stabilization and the correction method used in Pub- lications I and II gave a temperature error less than 0.6◦C when compared to readings from thermosensors. This accuracy can be considered adequate for MR-HIFU hyperthermia. Quantitative assessment of thermometry quality is essential for MR- HIFU hyperthermia applications where the desired temperature range is narrow (40–45◦C). This study provided valuable information on the be- havior of Philips MRI scanners during long-duration PRF temperature mapping. Moreover, this study introduced feasible correction methods for the magnetic field drift to enable safe MR-HIFU hyperthermia.

4.6 Publication VI

In Publication VI, the ability of simulated incident ultrasound intensity on the bone surface to predict the heating of the bone despite the complex interactions (reﬂections, attenuation, and mode conversions) ultrasound can have with bone, was discussed. The purpose was to beam-shape by deactivating transducer elements so that the exposure of the bones to acoustic waves, and thereby, the bone heating, reduces. The intensity- based heating estimate could be used to reduce heating of the bones to

51 Summary of The Publications a safe level without being overly conservative. The target application in this study was intercostal sonications of liver. The functionality of the Philips Sonalleve system, originally developed for avoiding scars in uterine fibroid treatments, estimated the ultrasound intensity field through the simplified Rayleigh–Sommerfeld integral and allowed selected ultrasound elements to be disabled. The calculation takes only few seconds in the Sonalleve operating computer, which is fast enough for clinical use. The element deactivation method was evaluated in 33 sonication experiments in an agar silica sonication phantom with embedded porcine rib bone and two optical thermosensors. The thermosensors were introduced via small through-holes to the surface of the bone facing the transducer. Some sonications were performed with all elements active and some with transducer elements deactivated based on the simulation. In addition, a full-wave simulation using an acoustic ray tracer [103] was applied off- line to get a more accurate estimate of the intensity at the bone surface. The simulated intensity was converted to energy density by multiplying by sonication duration. The correlation between energy density and thermosensor readings at the end of sonication was examined. The element deactivation method, based on incident intensity on the bone surface, was able to significantly reduce the heating of the bone. The energy density at the bone surface obtained from full-wave acoustic simulations calculated off-line was also reduced (Fig. 4.6). A linear dependence (R2: 0.80, slope: 40.7 Kmm2/J) was found between measured bone heating and simulated energy density. This indicates that the intensity simulation was able to predict the heating reduction at bone surface during the sonications. A possible use could be for acquiring first a priori knowledge on bone heating through a test sonication and then identifying, in silico, the intensity to reduce the in vivo bone heating to an acceptable level. Current methods for disabling transducer elements to spare ribs require either thermal simulations [93], which may be complicated in a heteroge- nous environment, or use simplified acoustic "shadowing" of the ribs without assessing their heating [40,41]. The method presented in Publication VI for element switch-off is fast and provides a heating estimate for the bones, which makes it an attractive approach for sparing ribs from dam- aging thermal exposure during intercostal sonications.

52 5. Conclusions

This Thesis has presented several technical advances in MR-HIFU hyperthermia and ablative therapies of recurrent rectal cancer, breast cancer, and liver cancer. In particular, the capabilities of phased array transducers were utilized and extended into new applications. Furthermore, the performance and reliability of long-duration temperature mapping in MR-HIFU hyperthermia was studied. Publication I described implemention of an MR-HIFU application for long-duration, large-volume hyperthermia that utilized electronic steering of the focus and selective use of transducer elements to control the heating. The work was continued in Publication II, in which the feasibility of the developed method for hyperthermia of recurrent rectal cancer was shown. The work led to the treatment of the ﬁrst MR-HIFU hyperthermia patient with recurrent rectal cancer in a clinical trial [158]. The long-duration MRI based temperature mapping methods were further described in Publication V. In Publication III, the capability of phased array transducers to produce multiple simultaneous focal points by manipulating the element phases for MR-HIFU hyperthermia was described. Publication IV discussed the use of element phases to make MRI image-based focal phase incoherence corrections in the context of breast sonications. Finally, a fast simulation method for reducing rib-bone heating by disabling elements in liver sonications was discussed in Publication VI.

The main conclusions of this Thesis are:

1. Combining mechanical transducer movement with focus steering and utilizing multiple temperature-mapping slices enables safe and long duration MR-HIFU hyperthermia for a large tissue volume with real-time

53 Conclusions

temperature control over the entire beam path. (Publication I)

2. MR-HIFU hyperthermia for recurrent rectal cancer, when rectal ﬁlling is used, has been demonstrated to be feasible and safe for the ﬁrst time. (Publication II)

3. Long-duration MR-HIFU hyperthermia using multiple simultaneous foci produced heating regions and temperature uniformity comparable to those for a single steered focus while achieving substantial reduction in acoustic instantaneous peak pressures. (Publication III)

4. Compensating the variation in ultrasound propagation times from element to focus based on MRI images improved focal sharpness and heating efﬁciency in wide-aperture transducer in a dedicated breast MR- HIFU system. (Publication IV)

5. Adequate accuracy in long-duration PRF temperature mapping was obtained when the drift of the magnetic field was corrected by (i) using a pre-scan period to stabilize the temperature of scanner components, (ii) using dynamic stabilization of the scanner center resonance frequency, and (iii) utilizing a retrospective first-order polynomial fit on temperature maps. (Publication V)

6. Simulated acoustic intensity on bone surface correlated with bone heating. This enables the use of fast intensity simulation methods for quantitative reduction of bone heating in intercostal sonications through selective use of phased array transducer elements. (Publication VI)

54 6. Future

The technical advances presented in this Thesis have in part promoted the translation of MR-HIFU applications into clinical practice. For example, the implemented large volume hyperthermia application later enabled the start of a clinical trial and radiotherapy treatment of the first recurrent rectal cancer patient with neoadjuvant MR-HIFU hyperthemia. However, there is still ample of room for technical improvements, for in- stance, in the field of transducer design and technology and in the MRI- based temperature mapping and treatment outcome assessment. In addition, the field would greatly benefit from, e.g. standardized treatment methods and dosimetry, automated treatment planning solutions, and better screening methods to find patients that are suitable for HIFU. It is of course important to continue making technical improvements to enable effective HIFU treatments, but lacking technical solutions are not necessarily the biggest hurdles in gaining widespread acceptance for HIFU. Even more important is to find clinical applications in which HIFU would be particularly useful and gather evidence on treatment efficacy and clinical value through high-quality trials. Based on solid evidence it is possible to build general acceptance, adoption, and standardization of HIFU therapies. MR-HIFU is a growing, disruptive technology that also needs to compete with existing technologies that already have received clinical acceptance and established standardized treatment procedures. On a time of skyrocketing healthcare costs, this is a challenge for HIFU. In addition, new medical devices need to pass regulatory requirements, which are probably more stringent currently than in the past, when for example radiation therapy was first started. MR-HIFU has great potential, but substantial efforts are still required before it can be translated into both clinical and economic success.

55 Future

56 References

[1] G. ter Haar and C. Coussios, “High intensity focused ultrasound: physical principles and devices,” International Journal of Hyperthermia, vol. 23, no. 2, pp. 89–104, 2007.

[2] S. A. Sapareto and W. C. Dewey, “Thermal dose determination in cancer therapy,” International Journal of Radiation Oncology, Biology, Physics, vol. 10, pp. 787–800, June 1984.

[3] M. W. Dewhirst, B. L. Viglianti, M. Lora-Michiels, M. Hanson, and P. J. Hoopes, “Basic principles of thermal dosimetry and thermal thresholds for tissue damage from hyperthermia,” International Journal of Hyperther- mia, vol. 19, no. 3, pp. 267–294, 2003.

[4] P. M. Krawczyk, B. Eppink, J. Essers, J. Stap, H. Rodermond, H. Odijk, A. Zelensky, C. van Bree, L. J. Stalpers, M. R. Buist, T. Soullié, J. Rens, H. J. M. Verhagen, M. J. O’Connor, N. A. P. Franken, T. L. M. Ten Ha- gen, R. Kanaar, and J. A. Aten, “Mild hyperthermia inhibits homologous recombination, induces BRCA2 degradation, and sensitizes cancer cells to poly (ADP-ribose) polymerase-1 inhibition,” Proceedings of the National Academy of Sciences of the United States of America, vol. 108, pp. 9851– 9856, June 2011.

[5] A. L. Oei, L. E. M. Vriend, J. Crezee, N. A. P. Franken, and P. M. Krawczyk, “Effects of hyperthermia on DNA repair pathways: one treatment to inhibit them all,” Radiation Oncology (London, England), vol. 10, Aug. 2015.

[6] C. R. Hill, J. C. Bamber, and G. ter Haar, Physical Principles of Medical Ultrasonics, vol. 2. Wiley Online Library, 2004.

[7] R. W. Wood and A. L. Loomis, “XXXVIII. The physical and biological effects of high-frequency sound-waves of great intensity,” The London, Edinburgh, And Dublin Philosophical Magazine And Journal of Science, vol. 4, no. 22, pp. 417–436, 1927.

[8] J. G. Lynn, R. L. Zwemer, A. J. Chick, and A. E. Miller, “A new method for the generation and use of focused ultrasound in experimental biology,” The journal of general physiology, vol. 26, no. 2, p. 179, 1942.

[9] S. L. Thomsen, “Mapping of thermal injury in biologic tissues using quantitative pathologic techniques,” in BiOS’99 International Biomedical Op- tics Symposium, pp. 82–95, International Society for Optics and Photonics, 1999.

57 References

[10] K. Hynynen, A. Darkazanli, E. Unger, J. F. Schenck, et al., “MRI-guided noninvasive ultrasound surgery.,” Medical Physics, vol. 20, no. 1, p. 107, 1993.

[11] H. E. Cline, J. F. Schenck, K. Hynynen, R. D. Watkins, S. P. Souza, and F. A. Jolesz, “MR-guided focused ultrasound surgery.,” Journal of Computer As- sisted Tomography, vol. 16, no. 6, pp. 956–965, 1992.

[12] K. Hynynen, A. Darkazanli, C. A. Damianou, E. Unger, and J. F. Schenck, “The usefulness of a contrast agent and gradient-recalled acquisition in a steady-state imaging sequence for magnetic resonance imaging-guided noninvasive ultrasound surgery.,” Investigative Radiology, vol. 29, no. 10, pp. 897–903, 1994.

[13] S. J. Graham, G. J. Stanisz, A. Kecojevic, M. J. Bronskill, and R. M. Henkel- man, “Analysis of changes in MR properties of tissues after heat treatment,” Magnetic Resonance in Medicine, vol. 42, no. 6, pp. 1061–1071, 1999.

[14] H. E. Cline, K. Hynynen, C. J. Hardy, R. D. Watkins, J. F. Schenck, and F. A. Jolesz, “MR temperature mapping of focused ultrasound surgery,” Magnetic Resonance in Medicine, vol. 31, no. 6, pp. 628–636, 1994.

[15] Y. Ishihara, A. Calderon, H. Watanabe, K. Okamoto, Y. Suzuki, K. Kuroda, and Y. Suzuki, “A precise and fast temperature mapping using water proton chemical shift,” Magnetic Resonance in Medicine, vol. 34, no. 6, pp. 814–823, 1995.

[16] K. Hynynen, O. Pomeroy, D. N. Smith, P. E. Huber, N. J. McDannold, J. Ket- tenbach, J. Baum, S. Singer, and F. A. Jolesz, “MR imaging-guided focused ultrasound surgery of ﬁbroadenomas in the breast: A feasibility study,” Radiology, vol. 219, no. 1, pp. 176–185, 2001.

[17] P. E. Huber, J. W. Jenne, R. Rastert, I. Simiantonakis, H.-P. Sinn, H.-J. Strittmatter, D. von Fournier, M. F. Wannenmacher, and J. Debus, “A new noninvasive approach in breast cancer therapy using magnetic resonance imaging-guided focused ultrasound surgery,” Cancer Research, vol. 61, no. 23, pp. 8441–8447, 2001.

[18] D. Gianfelice, A. Khiat, M. Amara, A. Belblidia, and Y. Boulanger, “MR imaging-guided focused ultrasound surgery of breast cancer: correlation of dynamic contrast-enhanced MRI with histopathologic ﬁndings,” Breast cancer research and treatment, vol. 82, no. 2, pp. 93–101, 2003.

[19] H. Furusawa, K. Namba, H. Nakahara, C. Tanaka, Y. Yasuda, E. Hirabara, M. Imahariyama, and K. Komaki, “The evolving non-surgical ablation of breast cancer: MR guided focused ultrasound (MRgFUS),” Breast Cancer, vol. 14, no. 1, pp. 55–58, 2007.

[20] R. Deckers, L. G. Merckel, B. D. de Senneville, G. Schubert, F. M. Knuttel, W. T. M. Mali, C. T. W. Moonen, M. van den Bosch, L. W. Bartels, and others, “Performance analysis of a dedicated breast MR-HIFU system for tumor ablation in breast cancer patients,” Physics in Medicine and Biology, vol. 60, no. 14, p. 5527, 2015.

58 References

[21] C. M. C. Tempany, E. A. Stewart, N. McDannold, B. J. Quade, F. A. Jolesz, and K. Hynynen, “MR Imaging–guided Focused Ultrasound Surgery of Uterine Leiomyomas: A Feasibility Study,” Radiology, vol. 226, pp. 897– 905, Mar. 2003.

[22] M. J. Voogt, H. Trillaud, Y. S. Kim, W. P. T. M. Mali, J. Barkhausen, L. W. Bartels, R. Deckers, N. Frulio, H. Rhim, H. K. Lim, T. Eckey, H. J. Niemi- nen, C. Mougenot, B. Keserci, J. Soini, T. Vaara, M. O. Köhler, S. D. Sokka, and M. A. A. J. van den Bosch, “Volumetric feedback ablation of uterine ﬁ- broids using magnetic resonance-guided high intensity focused ultrasound therapy,” European Radiology, vol. 22, pp. 411–417, Feb. 2012.

[23] R. Chopra, A. Colquhoun, M. Burtnyk, W. A. N’djin, I. Kobelevskiy, A. Boyes, K. Siddiqui, H. Foster, L. Sugar, M. A. Haider, and others, “MR imaging–controlled transurethral ultrasound therapy for conformal treatment of prostate tissue: initial feasibility in humans,” Radiology, vol. 265, no. 1, pp. 303–313, 2012.

[24] A. Okada, T. Murakami, K. Mikami, H. Onishi, N. Tanigawa, T. Marukawa, and H. Nakamura, “A case of hepatocellular carcinoma treated by MR- guided focused ultrasound ablation with respiratory gating,” Magnetic Resonance in Medical Sciences, vol. 5, no. 3, pp. 167–171, 2006.

[25] Z. Ram, Z. R. Cohen, S. Harnof, S. Tal, M. Faibel, D. Nass, S. E. Maier, M. Hadani, and Y. Mardor, “Magnetic Resonance Imaging-guided, High- Intensity Focused Ultrasound For Brain Tumore Therapy,” Neurosurgery, vol. 59, no. 5, pp. 949–956, 2006.

[26] N. McDannold, G. Clement, P. Black, F. Jolesz, and K. Hynynen, “Tran- scranial MRI-guided focused ultrasound surgery of brain tumors: Initial ﬁndings in three patients,” Neurosurgery, vol. 66, no. 2, p. 323, 2010.

[27] R. Catane, A. Beck, Y. Inbar, T. Rabin, N. Shabshin, S. Hengst, R. M. Pf- effer, A. Hanannel, O. Dogadkin, B. Liberman, and Kopelman, D., “MR- guided focused ultrasound surgery (MRgFUS) for the palliation of pain in patients with bone metastases—preliminary clinical experience,” Annals of Oncology, vol. 18, no. 1, pp. 163–167, 2007.

[28] D. Gianfelice, C. Gupta, W. Kucharczyk, P. Bret, D. Havill, and M. Clemons, “Palliative Treatment of Painful Bone Metastases with MR Imaging–guided Focused Ultrasound 1,” Radiology, vol. 249, no. 1, pp. 355–363, 2008.

[29] B. Liberman, D. Gianfelice, Y. Inbar, A. Beck, T. Rabin, N. Shabshin, G. Chander, S. Hengst, R. Pfeffer, A. Chechick, and others, “Pain palliation in patients with bone metastases using MR-guided focused ultrasound surgery: a multicenter study,” Annals of Surgical Oncology, vol. 16, no. 1, pp. 140–146, 2009.

[30] N. Lipsman, M. L. Schwartz, Y. Huang, L. Lee, T. Sankar, M. Chapman, K. Hynynen, and A. M. Lozano, “MR-guided focused ultrasound thalamotomy for essential tremor: a proof-of-concept study,” The Lancet Neurology, vol. 12, no. 5, pp. 462–468, 2013.

59 References

[31] W. J. Elias, N. Lipsman, W. G. Ondo, P. Ghanouni, Y. G. Kim, W. Lee, M. Schwartz, K. Hynynen, A. M. Lozano, B. B. Shah, D. Huss, R. Dal- lapiazza, R. Gwinn, J. Witt, S. Ro, H. Eisenberg, P. Fishman, D. Gandhi, C. Halpern, R. Chuang, K. Butts Pauly, T. Tierney, M. Hayes, G. R. Cos- grove, T. Yamaguchi, K. Abe, T. Takaomi, and J. W. Chang, “A randomized trial of focused ultrasound thalamotomy for essential tremor,” New Eng- land Journal of Medicine, vol. 375, no. 8, pp. 730–739, 2016.

[32] A. Magara, R. Bühler, D. Moser, M. Kowalski, P. Pourtehrani, and D. Jean- monod, “First experience with MR-guided focused ultrasound in the treatment of Parkinson’s disease,” Journal of Therapeutic Ultrasound, vol. 2, no. 1, p. 1, 2014.

[33] E. Martin, D. Jeanmonod, A. Morel, E. Zadicario, and B. Werner, “High- intensity focused ultrasound for noninvasive functional neurosurgery,” An- nals of Neurology, vol. 66, no. 6, pp. 858–861, 2009.

[34] S. Dromi, V. Frenkel, A. Luk, B. Traughber, M. Angstadt, M. Bur, J. Poff, J. Xie, S. K. Libutti, K. C. Li, and B. J. Wood, “Pulsed-high intensity focused ultrasound and low temperature–sensitive liposomes for enhanced targeted drug delivery and antitumor effect,” Clinical Cancer Research, vol. 13, no. 9, pp. 2722–2727, 2007.

[35] R. Staruch, R. Chopra, and K. Hynynen, “Localised drug release using MRI-controlled focused ultrasound hyperthermia,” International Journal of Hyperthermia, vol. 27, no. 2, pp. 156–171, 2011.

[36] A. Partanen, P. S. Yarmolenko, A. Viitala, S. Appanaboyina, D. Haem- merich, A. Ranjan, G. Jacobs, D. Woods, J. Enholm, B. J. Wood, and M. R. Dreher, “Mild hyperthermia with magnetic resonance-guided high- intensity focused ultrasound for applications in drug delivery,” Interna- tional Journal of Hyperthermia, vol. 28, no. 4, pp. 320–336, 2012.

[37] R. Salomir, J. Palussière, F. C. Vimeux, J. A. de Zwart, B. Quesson, M. Gauchet, P. Lelong, J. Pergrale, N. Grenier, and C. T. W. Moonen, “Lo- cal hyperthermia with MR-guided focused ultrasound: Spiral trajectory of the focal point optimized for temperature uniformity in the target region,” Journal of Magnetic Resonance Imaging, vol. 12, no. 4, pp. 571–583, 2000.

[38] M. O. Köhler, C. Mougenot, B. Quesson, J. Enholm, B. Le Bail, C. Laurent, C. T. W. Moonen, and G. J. Ehnholm, “Volumetric HIFU ablation under 3d guidance of rapid MRI thermometry,” Medical Physics, vol. 36, no. 8, pp. 3521–3535, 2009.

[39] Y. Y. Botros, E. S. Ebbini, and J. L. Volakis, “Optimal synthesis of phased array ﬁeld patterns in the presence of the ribcage,” in Ultrasonics Sympo- sium, 1997. Proceedings., 1997 IEEE, vol. 2, pp. 1751–1754, IEEE, 1997.

[40] B. Quesson, M. Merle, M. O. Köhler, C. Mougenot, S. Roujol, B. D. de Sen- neville, and C. T. Moonen, “A method for MRI guidance of intercostal high intensity focused ultrasound ablation in the liver,” Medical Physics, vol. 37, no. 6, p. 2533, 2010.

[41] F. Marquet, J. F. Aubry, M. Pernot, M. Fink, and M. Tanter, “Optimal transcostal high-intensity focused ultrasound with combined real-time

60 References

3d movement tracking and correction,” Physics in Medicine and Biology, vol. 56, no. 22, p. 7061, 2011.

[42] J. W. Wijlemans, M. de Greef, G. Schubert, L. W. Bartels, C. T. W. Moonen, M. A. van den Bosch, and M. Ries, “A clinically feasible treatment protocol for magnetic resonance-guided high-intensity focused ultrasound ablation in the liver,” Investigative radiology, vol. 50, no. 1, pp. 24–31, 2015.

[43] B. Quesson, J. A. de Zwart, and C. T. W. Moonen, “Magnetic resonance temperature imaging for guidance of thermotherapy,” Journal of Magnetic Resonance Imaging, vol. 12, no. 4, pp. 525–533, 2000.

[44] V. Rieke and K. Butts Pauly, “MR thermometry,” Journal of Magnetic Res- onance Imaging, vol. 27, no. 2, pp. 376–390, 2008.

[45] B. D. de Senneville, S. Roujol, C. T. W. Moonen, and M. Ries, “Motion correction in MR thermometry of abdominal organs: a comparison of the refer- enceless vs. the multibaseline approach,” Magnetic Resonance in Medicine, vol. 64, no. 5, pp. 1373–1381, 2010.

[46] A. M. El-Sharkawy, M. Schär, P. A. Bottomley, and E. Atalar, “Monitor- ing and correcting spatio-temporal variations of the MR scanner’s static magnetic ﬁeld,” Magnetic Resonance Materials in Physics, Biology and Medicine, vol. 19, no. 5, pp. 223–236, 2006.

[47] R. S. Cobbold, Foundations of Biomedical Ultrasound. Oxford University Press, 2006.

[48] K. Nightingale, M. S. Soo, R. Nightingale, and G. Trahey, “Acoustic radiation force impulse imaging: in vivo demonstration of clinical feasibility,” Ultrasound in Medicine & Biology, vol. 28, no. 2, pp. 227–235, 2002.

[49] >. C. Coussios, C. H. Farny, G. Ter Haar, and R. A. Roy, “Role of acoustic cavitation in the delivery and monitoring of cancer treatment by high- intensity focused ultrasound (HIFU),” International Journal of Hyperther- mia, vol. 23, pp. 105–120, Jan. 2007.

[50] W. W. Roberts, “Development and translation of histotripsy: current status and future directions,” Current Opinion in Urology, vol. 24, no. 1, p. 104, 2014.

[51] D. L. Miller, N. B. Smith, M. R. Bailey, G. J. Czarnota, K. Hynynen, I. R. S. Makin, B. C. o. t. A. I. o. U. i. Medicine, and others, “Overview of therapeutic ultrasound applications and safety considerations,” Journal of Ul- trasound in Medicine, vol. 31, no. 4, pp. 623–634, 2012.

[52] V. J. Robertson and K. G. Baker, “A review of therapeutic ultrasound: ef- fectiveness studies,” Physical Therapy, vol. 81, no. 7, pp. 1339–1350, 2001.

[53] S. J. Hectors, I. Jacobs, C. T. W. Moonen, G. J. Strijkers, and K. Nicolay, “MRI methods for the evaluation of high intensity focused ultrasound tumor treatment: current status and future needs,” Magnetic Resonance in Medicine, vol. 75, no. 1, pp. 302–317, 2016.

[54] D. Arora, D. Cooley, T. Perry, M. Skliar, and R. B. Roemer, “Direct thermal dose control of constrained focused ultrasound treatments: phantom and in vivo evaluation,” Physics in medicine and biology, vol. 50, p. 1919, 2005.

61 References

[55] J. K. Enholm, M. O. Kohler, B. Quesson, C. Mougenot, C. T. W. Moonen, and S. D. Sokka, “Improved volumetric MR-HIFU ablation by robust binary feedback control,” IEEE Transactions on Biomedical Engineering, vol. 57, no. 1, pp. 103–113, 2010.

[56] C. Damianou and K. Hynynen, “The effect of various physical parameters on the size and shape of necrosed tissue volume during ultrasound surgery,” The Journal of the Acoustical Society of America, vol. 95, pp. 1641–1649, Mar. 1994.

[57] N. J. McDannold, R. L. King, F. A. Jolesz, and K. Hynynen, “Usefulness of MR imaging-derived thermometry and dosimetry in determining the threshold for tissue damage induced by thermal surgery in rabbits,” Radi- ology, vol. 216, pp. 517–523, Aug. 2000.

[58] B. L. Viglianti, P. Stauffer, E. Repasky, E. Jones, Z. Vujaskovic, and M. De- whirst, “Hyperthermia,” Holland-Frei Cancer Medicine, ed, vol. 8, pp. 528– 540, 2010.

[59] C. C. Vernon, J. W. Hand, S. B. Field, D. Machin, J. B. Whaley, J. van der Zee, W. L. van Putten, G. C. van Rhoon, J. D. van Dijk, D. González González, F. F. Liu, P. Goodman, and M. Sherar, “Radiother- apy with or without hyperthermia in the treatment of superﬁcial localized breast cancer: results from ﬁve randomized controlled trials. International Collaborative Hyperthermia Group,” International Journal of Radiation Oncology, Biology, Physics, vol. 35, pp. 731–744, July 1996.

[60] J. van der Zee, D. González González, G. C. van Rhoon, J. D. van Dijk, W. L. van Putten, and A. A. Hart, “Comparison of radiotherapy alone with radiotherapy plus hyperthermia in locally advanced pelvic tumours: a prospective, randomised, multicentre trial. Dutch Deep Hyperthermia Group,” Lancet (London, England), vol. 355, pp. 1119–1125, Apr. 2000.

[61] M. R. Horsman and J. Overgaard, “Hyperthermia: a potent enhancer of radiotherapy,” Clinical Oncology (Royal College of Radiologists (Great Britain)), vol. 19, pp. 418–426, Aug. 2007.

[62] R. D. Issels, L. H. Lindner, J. Verweij, P. Wust, P. Reichardt, B.-C. Schem, S. Abdel-Rahman, S. Daugaard, C. Salat, C.-M. Wendtner, Z. Vujaskovic, R. Wessalowski, K.-W. Jauch, H. R. Dürr, F. Ploner, A. Baur-Melnyk, U. Mansmann, W. Hiddemann, J.-Y. Blay, P. Hohenberger, European Or- ganisation for Research and Treatment of Cancer Soft Tissue and Bone Sarcoma Group (EORTC-STBSG), and European Society for Hyperther- mic Oncology (ESHO), “Neo-adjuvant chemotherapy alone or with regional hyperthermia for localised high-risk soft-tissue sarcoma: a randomised phase 3 multicentre study,” The Lancet. Oncology, vol. 11, pp. 561–570, June 2010.

[63] B. Hildebrandt, P. Wust, O. Ahlers, A. Dieing, G. Sreenivasa, T. Kerner, R. Felix, and H. Riess, “The cellular and molecular basis of hyperthermia,” Critical Reviews in Oncology/Hematology, vol. 43, pp. 33–56, July 2002.

[64] K. F. Chu and D. E. Dupuy, “Thermal ablation of tumours: biological mechanisms and advances in therapy,” Nature Reviews Cancer, vol. 14, no. 3, pp. 199–208, 2014.

62 References

[65] G. Kong, R. D. Braun, and M. W. Dewhirst, “Hyperthermia enables tumor- speciﬁc nanoparticle delivery: effect of particle size,” Cancer Research, vol. 60, pp. 4440–4445, Aug. 2000.

[66] J. van der Zee, “Heating the patient: a promising approach?,” Annals of Oncology, vol. 13, no. 8, pp. 1173–1184, 2002.

[67] H. Azhari, “Ultrasonic Imaging Using the Pulse-Echo Technique,” Basics of Biomedical Ultrasound for Engineers, pp. 191–231, 2010.

[68] C. R. Hill, “Optimum acoustic frequency for focused ultrasound surgery,” Ultrasound in Medicine & Biology, vol. 20, no. 3, pp. 271–277, 1994.

[69] I. H. Rivens, R. L. Clarke, and G. ter Haar, “Design of focused ultrasound surgery transducers,” IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 43, no. 6, pp. 1023–1031, 1996.

[70] A. Gelet, J. Y. Chapelon, R. Bouvier, O. Rouviere, Y. Lasne, D. Lyonnet, and J. M. Dubernard, “Transrectal high-intensity focused ultrasound: mini- mally invasive therapy of localized prostate cancer,” Journal of Endourol- ogy, vol. 14, no. 6, pp. 519–528, 2000.

[71] R. Chopra, M. Burtnyk, M. A. Haider, and M. J. Bronskill, “Method for MRI-guided conformal thermal therapy of prostate with planar transurethral ultrasound heating applicators,” Physics in Medicine and Biology, vol. 50, no. 21, p. 4957, 2005.

[72] G. T. Clement, J. Sun, T. Giesecke, and K. Hynynen, “A hemisphere array for non-invasive ultrasound brain therapy and surgery,” Physics in Medicine and Biology, vol. 45, pp. 3707–3719, Dec. 2000.

[73] F. A. Jolesz and K. Hynynen, MRI-Guided Focused Ultrasound Surgery. CRC Press, Sept. 2007.

[74] H. T. Soh, I. Ladabaum, A. Atalar, C. F. Quate, and B. T. Khuri-Yakub, “Sil- icon micromachined ultrasonic immersion transducers,” Applied Physics Letters, vol. 69, no. 24, pp. 3674–3676, 1996.

[75] O. Oralkan, A. S. Ergun, J. A. Johnson, M. Karaman, U. Demirci, K. Ka- viani, T. H. Lee, and B. T. Khuri-Yakub, “Capacitive micromachined ultrasonic transducers: Next-generation arrays for acoustic imaging?,” IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 49, no. 11, pp. 1596–1610, 2002.

[76] B. T. Khuri-Yakub and O. Oralkan, “Capacitive micromachined ultrasonic transducers for medical imaging and therapy,” Journal of Micromechan- ics and Microengineering : Structures, Devices, and Systems, vol. 21, pp. 054004–054014, May 2011.

[77] J. J. Bernstein, S. L. Finberg, K. Houston, L. C. Niles, H. D. Chen, L. E. Cross, K. K. Li, and K. Udayakumar, “Micromachined high frequency fer- roelectric sonar transducers,” IEEE Transactions on Ultrasonics, Ferro- electrics, and Frequency Control, vol. 44, pp. 960–969, Sept. 1997.

[78] Y. Qiu, J. V. Gigliotti, M. Wallace, F. Griggio, C. E. Demore, S. Cochran, and S. Trolier-McKinstry, “Piezoelectric micromachined ultrasound transducer (PMUT) arrays for integrated sensing, actuation and imaging,” Sensors, vol. 15, no. 4, pp. 8020–8041, 2015.

63 References

[79] A. S. Savoia, G. Caliano, and M. Pappalardo, “A CMUT probe for medical ultrasonography: from microfabrication to system integration,” IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 59, no. 6, pp. 1127–1138, 2012.

[80] W. A. Smith, “The role of piezocomposites in ultrasonic transducers,” in Ul- trasonics Symposium, 1989. Proceedings., IEEE 1989, pp. 755–766, 1989.

[81] D. Cathignol, O. A. Sapozhnikov, and Y. Theillere, “Comparison of acoustic ﬁelds radiated from piezoceramic and piezocomposite focused radiators,” The Journal of the Acoustical Society of America, vol. 105, no. 5, pp. 2612– 2617, 1999.

[82] G. Fleury, D. M. de Lima, K. Hynynen, R. Berriet, O. Le Baron, and B. Huguenin, “New piezocomposite transducers for therapeutic ultrasound,” in Biomedical Optics 2003, pp. 227–236, International Society for Optics and Photonics, 2003.

[83] E. S. Ebbini and C. A. Cain, “Multiple-focus ultrasound phased-array pattern synthesis: optimal driving-signal distributions for hyperthermia,” Ul- trasonics, Ferroelectrics and Frequency Control, IEEE Transactions on, vol. 36, no. 5, pp. 540–548, 1989.

[84] D. R. Daum and K. Hynynen, “Thermal dose optimization via temporal switching in ultrasound surgery,” IEEE Transactions on Ultrasonics, Fer- roelectrics, and Frequency Control, vol. 45, no. 1, pp. 208–215, 1998.

[85] D. R. Daum and K. Hynynen, “A 256-element ultrasonic phased array system for the treatment of large volumes of deep seated tissue,” IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 46, pp. 1254–68, 1999. 5.

[86] S. A. Goss, L. A. Frizzell, J. T. Kouzmanoff, J. M. Barich, and J. M. Yang, “Sparse random ultrasound phased array for focal surgery,” IEEE Trans- actions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 43, no. 6, pp. 1111–1121, 1996.

[87] F. Dupenloup, J. Y. Chapelon, D. J. Cathignol, and O. A. Sapozhnikov, “Re- duction of the grating lobes of annular arrays used in focused ultrasound surgery,” IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 43, no. 6, pp. 991–998, 1996.

[88] L. R. Gavrilov and J. W. Hand, “A theoretical assessment of the relative performance of spherical phased arrays for ultrasound surgery,” IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control, vol. 47, no. 1, pp. 125–139, 2000.

[89] J. L. Thomas and M. A. Fink, “Ultrasonic beam focusing through tissue in- homogeneities with a time reversal mirror: application to transskull therapy,” Ultrasonics, Ferroelectrics and Frequency Control, IEEE Transactions on, vol. 43, no. 6, pp. 1122–1129, 1996.

[90] K. Hynynen and F. A. Jolesz, “Demonstration of potential noninvasive ultrasound brain therapy through an intact skull,” Ultrasound in medicine & biology, vol. 24, no. 2, pp. 275–283, 1998.

64 References

[91] G. T. Clement and K. Hynynen, “A non-invasive method for focusing ultrasound through the human skull,” Physics in Medicine and Biology, vol. 47, no. 8, p. 1219, 2002.

[92] L. G. Merckel, L. W. Bartels, M. O. Köhler, H. D. van den Bongard, R. Deck- ers, P. T. M. Willem, C. A. Binkert, C. T. W. Moonen, K. G. Gilhuijs, and M. A. van den Bosch, “MR-guided high-intensity focused ultrasound ablation of breast cancer with a dedicated breast platform,” Cardiovascular and Interventional Radiology, vol. 36, no. 2, pp. 292–301, 2013.

[93] H.-L. Liu, H. Chang, W.-S. Chen, T.-C. Shih, J.-K. Hsiao, and W.-L. Lin, “Feasibility of transrib focused ultrasound thermal ablation for liver tumors using a spherically curved 2d array: A numerical study,” Medical Physics, vol. 34, no. 9, pp. 3436–3448, 2007.

[94] E. Cochard, C. Prada, J. F. Aubry, and M. Fink, “Ultrasonic focusing through the ribs using the DORT method,” Medical Physics, vol. 36, no. 8, p. 3495, 2009.

[95] M. de Greef, G. Schubert, J. Wijlemans, J. Koskela, L. W. Bartels, C. T. W. Moonen, and M. Ries, “Feasibility of intercostal high intensity focused ultrasound ablation of clinically relevant volumes under the application of beam shaping,” Journal of Therapeutic Ultrasound, vol. 3, no. 1, p. 1, 2015.

[96] J.-F. Aubry, K. B. Pauly, C. T. W. Moonen, G. ter Haar, M. Ries, R. Salomir, S. D. Sokka, K. M. Sekins, Y. Shapira, F. Ye, and others, “The road to clinical use of high-intensity focused ultrasound for liver cancer: technical and clinical consensus,” Journal of Therapeutic Ultrasound, vol. 1, no. 1, p. 1, 2013.

[97] K. Hynynen, “Demonstration of enhanced temperature elevation due to nonlinear propagation of focussed ultrasound in dog’s thigh in vivo,” Ul- trasound in Medicine & Biology, vol. 13, no. 2, pp. 85–91, 1987.

[98] E. A. Filonenko and V. A. Khokhlova, “Effect of acoustic nonlinearity on heating of biological tissue by high-intensity focused ultrasound,” Acousti- cal Physics, vol. 47, no. 4, pp. 468–475, 2001.

[99] M. S. Canney, V. A. Khokhlova, O. V. Bessonova, M. R. Bailey, and L. A. Crum, “Shock-induced heating and millisecond boiling in gels and tissue due to high intensity focused ultrasound,” Ultrasound in medicine & biology, vol. 36, no. 2, pp. 250–267, 2010.

[100] H.-L. Liu, N. McDannold, and K. Hynynen, “Focal beam distortion and treatment planning in abdominal focused ultrasound surgery,” Medical Physics, vol. 32, no. 5, pp. 1270–1280, 2005.

[101] U. Vyas and D. Christensen, “Ultrasound beam propagation using the hy- brid angular spectrum method,” in 2008 30th Annual International Confer- ence of the IEEE Engineering in Medicine and Biology Society, pp. 2526– 2529, IEEE, 2008.

[102] J. Georgii, C. von Dresky, S. Meier, D. Demedts, C. Schumann, T. Preusser, J. Bender, K. Erleben, and E. Galin, “Focused Ultrasound-Efﬁcient GPU Simulation Methods for Therapy Planning.,” in VRIPHYS, pp. 119–128, 2011.

65 References

[103] J. Koskela, E. Vahala, M. d. Greef, L. P. Laﬁtte, and M. Ries, “Stochastic ray tracing for simulation of high intensity focal ultrasound therapya),” The Journal of the Acoustical Society of America, vol. 136, pp. 1430–1440, Sept. 2014.

[104] H. H. Pennes, “Analysis of tissue and arterial blood temperatures in the resting human forearm,” Journal of Applied Physiology, vol. 1, no. 2, pp. 93–122, 1948.

[105] K. Mahoney, T. Fjield, N. McDannold, G. Clement, and K. Hynynen, “Com- parison of modelled and observed in vivo temperature elevations induced by focused ultrasound: implications for treatment planning,” Physics in Medicine and Biology, vol. 46, no. 7, p. 1785, 2001.

[106] M. Pernot, J.-F. Aubry, M. Tanter, J.-L. Thomas, and M. Fink, “High power transcranial beam steering for ultrasonic brain therapy,” Physics in Medicine and Biology, vol. 48, no. 16, p. 2577, 2003.

[107] C. R. Hill, I. Rivens, M. G. Vaughan, and G. ter Haar, “Lesion development in focused ultrasound surgery: a general model,” Ultrasound in Medicine & Biology, vol. 20, no. 3, pp. 259–269, 1994.

[108] Y. Zhou, L. Zhai, R. Simmons, and P. Zhong, “Measurement of high intensity focused ultrasound ﬁelds by a ﬁber optic probe hydrophone,” The Journal of the Acoustical Society of America, vol. 120, no. 2, p. 676, 2006.

[109] P. Morris, A. Hurrell, A. Shaw, E. Zhang, and P. Beard, “A Fabry–Pérot ﬁber-optic ultrasonic hydrophone for the simultaneous measurement of temperature and acoustic pressure,” The Journal of the Acoustical Soci- ety of America, vol. 125, no. 6, pp. 3611–3622, 2009.

[110] R. W. Brown, Y.-C. N. Cheng, E. M. Haacke, M. R. Thompson, and R. Venkatesan, Magnetic Resonance Imaging: Physical Principles And Se- quence Design. John Wiley & Sons, 2014.

[111] J. Hindley, W. M. Gedroyc, L. Regan, E. Stewart, C. Tempany, K. Hyn- nen, N. Macdanold, Y. Inbar, Y. Itzchak, J. Rabinovici, and others, “MRI guidance of focused ultrasound therapy of uterine ﬁbroids: early results,” American Journal of Roentgenology, vol. 183, no. 6, pp. 1713–1719, 2004.

[112] E. A. Stewart, W. M. Gedroyc, C. M. Tempany, B. J. Quade, Y. Inbar, T. Ehrenstein, A. Shushan, J. T. Hindley, R. D. Goldin, M. David, and others, “Focused ultrasound treatment of uterine ﬁbroid tumors: safety and feasibility of a noninvasive thermoablative technique,” American Journal of Obstetrics and Gynecology, vol. 189, no. 1, pp. 48–54, 2003.

[113] E. A. Stewart, J. Rabinovici, C. M. Tempany, Y. Inbar, L. Regan, B. Gastout, G. Hesley, H. S. Kim, S. Hengst, and W. M. Gedroye, “Clinical outcomes of focused ultrasound surgery for the treatment of uterine ﬁbroids,” Fertility and Sterility, vol. 85, no. 1, pp. 22–29, 2006.

[114] C. Damianou and K. Hynynen, “Focal spacing and near-ﬁeld heating during pulsed high temperature ultrasound therapy,” Ultrasound in Medicine & Biology, vol. 19, no. 9, pp. 777–787, 1993.

66 References

[115] X. Fan and K. Hynynen, “Ultrasound surgery using multiple sonications—treatment time considerations,” Ultrasound in Medicine & Biology, vol. 22, no. 4, pp. 471–482, 1996.

[116] N. J. McDannold, F. A. Jolesz, and K. H. Hynynen, “Determination of the Optimal Delay between Sonications during Focused Ultrasound Surgery in Rabbits by Using MR Imaging to Monitor Thermal Buildup in Vivo1,” Radiology, vol. 211, no. 2, pp. 419–426, 1999.

[117] Y.-s. Kim, J.-H. Kim, H. Rhim, H. K. Lim, B. Keserci, D.-S. Bae, B.-G. Kim, J.-W. Lee, T.-J. Kim, and C. H. Choi, “Volumetric MR-guided high-intensity focused ultrasound ablation with a one-layer strategy to treat large uterine ﬁbroids: initial clinical outcomes,” Radiology, vol. 263, no. 2, pp. 600–609, 2012.

[118] K. Hynynen, W. R. Freund, H. E. Cline, A. H. Chung, R. D. Watkins, J. P. Vetro, and F. A. Jolesz, “A clinical, noninvasive, MR imaging-monitored ultrasound surgery method,” Radiographics, vol. 16, no. 1, pp. 185–195, 1996.

[119] Y.-s. Kim, H. K. Lim, J.-H. Kim, H. Rhim, B. K. Park, B. Keserci, M. O. Köhler, D.-S. Bae, B.-G. Kim, J.-W. Lee, and others, “Dynamic contrast- enhanced magnetic resonance imaging predicts immediate therapeutic response of magnetic resonance-guided high-intensity focused ultrasound ablation of symptomatic uterine ﬁbroids,” Investigative Radiology, vol. 46, no. 10, pp. 639–647, 2011.

[120] F. M. Fennessy and C. M. Tempany, “MRI-guided focused ultrasound surgery of uterine leiomyomas 1,” Academic Radiology, vol. 12, no. 9, pp. 1158–1166, 2005.

[121] A. M. Venkatesan, A. Partanen, T. K. Pulanic, M. R. Dreher, J. Fischer, R. K. Zurawin, R. Muthupillai, S. D. Sokka, H. J. Nieminen, N. Sinaii, and others, “Magnetic resonance imaging–guided volumetric ablation of symptomatic leiomyomata: Correlation of imaging with histology,” Journal of Vascular and Interventional Radiology, vol. 23, no. 6, pp. 786–794, 2012.

[122] I. J. Rowland, I. Rivens, L. Chen, C. H. Lebozer, D. J. Collins, G. ter Haar, and M. O. Leach, “MRI study of hepatic tumours following high intensity focused ultrasound surgery.,” The British Journal of Radiology, vol. 70, no. 830, pp. 144–153, 1997.

[123] L. Chen, G. ter Haar, D. Robertson, J. P. Bensted, and C. R. Hill, “Histo- logical study of normal and tumor-bearing liver treated with focused ultrasound,” Ultrasound in Medicine & Biology, vol. 25, no. 5, pp. 847–856, 1999.

[124] F. Wu, W.-Z. Chen, J. Bai, J.-Z. Zou, Z.-L. Wang, H. Zhu, and Z.-B. Wang, “Tumor vessel destruction resulting from high-intensity focused ultrasound in patients with solid malignancies,” Ultrasound in Medicine & Bi- ology, vol. 28, no. 4, pp. 535–542, 2002.

[125] H.-L. M. Cheng, C. M. Purcell, J. M. Bilbao, and D. B. Plewes, “Prediction of subtle thermal histopathological change using a novel analysis of Gd- DTPA kinetics,” Journal of Magnetic Resonance Imaging, vol. 18, no. 5, pp. 585–598, 2003.

67 References

[126] O. Noterdaeme, T. A. Leslie, J. E. Kennedy, R. R. Phillips, and M. Brady, “The use of time to maximum enhancement to indicate areas of ablation following the treatment of liver tumours with high-intensity focused ultrasound.,” The British Journal of Radiology, vol. 82, no. 977, pp. 412–420, 2009.

[127] A. H. Chung, F. A. Jolesz, and K. Hynynen, “Thermal dosimetry of a focused ultrasound beam in vivo by magnetic resonance imaging,” Medical Physics, vol. 26, no. 9, pp. 2017–2026, 1999.

[128] J. D. Hazle, R. J. Stafford, and R. E. Price, “Magnetic resonance imaging- guided focused ultrasound thermal therapy in experimental animal models: Correlation of ablation volumes with pathology in rabbit muscle and VX2 tumors,” Journal of Magnetic Resonance Imaging, vol. 15, no. 2, pp. 185–194, 2002.

[129] J. D. Hazle, C. J. Diederich, M. Kangasniemi, R. E. Price, L. E. Olsson, and R. J. Stafford, “MRI-guided thermal therapy of transplanted tumors in the canine prostate using a directional transurethral ultrasound applicator,” Journal of Magnetic Resonance Imaging, vol. 15, no. 4, pp. 409–417, 2002.

[130] J. W. Wijlemans, R. Deckers, M. A. van den Bosch, B. A. Seinstra, M. van Stralen, P. J. van Diest, C. T. W. Moonen, and L. W. Bartels, “Evolution of the ablation region after magnetic resonance–guided high-intensity focused ultrasound ablation in a Vx2 tumor model,” Investigative Radiology, vol. 48, no. 6, pp. 381–386, 2013.

[131] N. McDannold, C. M. Tempany, F. M. Fennessy, M. J. So, F. J. Rybicki, E. A. Stewart, F. A. Jolesz, and K. Hynynen, “Uterine leiomyomas: MR imaging–based thermometry and thermal dosimetry during focused ultrasound thermal ablation,” Radiology, vol. 240, no. 1, pp. 263–272, 2006.

[132] M. C. Pilatou, E. A. Stewart, S. E. Maier, F. M. Fennessy, K. Hyny- nen, C. Tempany, and N. McDannold, “MRI-based thermal dosimetry and diffusion-weighted imaging of MRI-guided focused ultrasound thermal ablation of uterine ﬁbroids,” Journal of Magnetic Resonance Imaging, vol. 29, no. 2, pp. 404–411, 2009.

[133] N. M. Hijnen, A. Elevelt, and H. Grüll, “Stability and trapping of magnetic resonance imaging contrast agents during high-intensity focused ultrasound ablation therapy,” Investigative Radiology, vol. 48, no. 7, pp. 517– 524, 2013.

[134] N. M. Hijnen, A. Elevelt, J. Pikkemaat, C. Bos, L. W. Bartels, and H. Grüll, “The magnetic susceptibility effect of gadolinium-based contrast agents on PRFS-based MR thermometry during thermal interventions,” Journal of Therapeutic Ultrasound, vol. 1, no. 1, p. 1, 2013.

[135] T. Wu, J. P. Felmlee, J. F. Greenleaf, S. J. Riederer, and R. L. Ehman, “As- sessment of thermal tissue ablation with MR elastography,” Magnetic resonance in medicine, vol. 45, no. 1, pp. 80–87, 2001.

[136] R. Souchon, O. Rouvière, A. Gelet, V. Detti, S. Srinivasan, J. Ophir, and J. Y. Chapelon, “Visualisation of HIFU lesions using elastography of the human prostate< i> in vivo: preliminary results,” Ultrasound in medicine & biology, vol. 29, no. 7, pp. 1007–1015, 2003.

68 References

[137] R. Souchon, R. Salomir, O. Beuf, L. Milot, D. Grenier, D. Lyonnet, J.-Y. Chapelon, and O. Rouvière, “Transient MR elastography (t-MRE) using ultrasound radiation force: Theory, safety, and initial experiments in vitro,” Magnetic Resonance in Medicine, vol. 60, pp. 871–881, Oct. 2008.

[138] B. Larrat, M. Pernot, J.-F. Aubry, E. Dervishi, R. Sinkus, D. Seilhean, Y. Marie, A.-L. Boch, M. Fink, and M. Tanter, “MR-guided transcranial brain HIFU in small animal models,” Physics in medicine and biology, vol. 55, no. 2, p. 365, 2009.

[139] M. K. Lam, C. Oerlemans, M. Froeling, R. Deckers, A. D. Barten-Van Ri- jbroek, M. A. Viergever, C. T. W. Moonen, C. Bos, and L. W. Bartels, “DCE- MRI and IVIM-MRI of rabbit Vx2 tumors treated with MR-HIFU-induced mild hyperthermia,” Journal of Therapeutic Ultrasound, vol. 4, no. 1, p. 1, 2016.

[140] M. de Smet, E. Heijman, S. Langereis, N. M. Hijnen, and H. Grüll, “Mag- netic resonance imaging of high intensity focused ultrasound mediated drug delivery from temperature-sensitive liposomes: an in vivo proof-of- concept study,” Journal of Controlled Release, vol. 150, no. 1, pp. 102–110, 2011.

[141] A. H. Negussie, P. S. Yarmolenko, A. Partanen, A. Ranjan, G. Jacobs, D. Woods, H. Bryant, D. Thomasson, M. W. Dewhirst, B. J. Wood, and others, “Formulation and characterisation of magnetic resonance image- able thermally sensitive liposomes for use with magnetic resonance-guided high intensity focused ultrasound,” International Journal of Hyperther- mia, vol. 27, no. 2, pp. 140–155, 2011.

[142] R. D. Peters, R. S. Hinks, and R. M. Henkelman, “Ex vivo tissue-type inde- pendence in proton-resonance frequency shift MR thermometry,” Magnetic Resonance in Medicine, vol. 40, no. 3, pp. 454–459, 1998.

[143] N. McDannold, “Quantitative MRI-based temperature mapping based on the proton resonant frequency shift: review of validation studies,” Interna- tional Journal of Hyperthermia, vol. 21, no. 6, pp. 533–546, 2005.

[144] J. D. Poorter, “Noninvasive MRI thermometry with the proton resonance frequency method: study of susceptibility effects,” Magnetic resonance in medicine, vol. 34, no. 3, pp. 359–367, 1995.

[145] J. A. de Zwart, F. C. Vimeux, C. Delalande, P. Canioni, and C. T. W. Moo- nen, “Fast lipid-suppressed MR temperature mapping with echo-shifted gradient-echo imaging and spectral-spatial excitation,” Magnetic resonance in medicine, vol. 42, no. 1, pp. 53–59, 1999.

[146] S. Sprinkhuizen, M. Konings, M. van der Bom, M. Viergever, C. Bakker, and L. W. Bartels, “Temperature-induced tissue susceptibility changes lead to signiﬁcant temperature errors in PRFS-based MR thermometry during thermal interventions,” Magnetic Resonance in Medicine, vol. 64, no. 5, pp. 1360–1372, 2010.

[147] K. Hynynen, N. McDannold, R. V. Mulkern, and F. A. Jolesz, “Tempera- ture monitoring in fat with MRI,” Magnetic Resonance in Medicine, vol. 43, no. 6, pp. 901–904, 2000.

69 References

[148] P. Baron, M. Ries, R. Deckers, M. Greef, J. Tanttu, M. Köhler, M. A. Viergever, C. T. W. Moonen, and L. W. Bartels, “In vivo T2-based MR thermometry in adipose tissue layers for high-intensity focused ultrasound near-ﬁeld monitoring,” Magnetic Resonance in Medicine, vol. 72, no. 4, pp. 1057–1064, 2014.

[149] M. Parmala, M. Eriksson, M. Rytioja, J. Tanttu, and M. Köhler, “Temper- ature measurement in human fat with T2 imaging,” Journal of Magnetic Resonance Imaging, 2015.

[150] J. C. Hindman, “Proton resonance shift of water in the gas and liquid states,” The Journal of Chemical Physics, vol. 44, no. 12, pp. 4582–4592, 1966.

[151] B. D. de Senneville, B. Quesson, and C. T. W. Moonen, “Magnetic resonance temperature imaging,” International Journal of Hyperthermia, vol. 21, no. 6, pp. 515–531, 2005.

[152] A. H. Chung, K. Hynynen, V. Colucci, K. Oshio, H. E. Cline, and F. A. Jolesz, “Optimization of spoiled gradient-echo phase imaging for in vivo localization of a focused ultrasound beam,” Magnetic Resonance in Medicine, vol. 36, no. 5, pp. 745–752, 1996.

[153] Y.-C. Chung and J. L. Duerk, “Signal formation in echo-shifted sequences,” Magnetic Resonance in Medicine, vol. 42, no. 5, pp. 864–875, 1999.

[154] C. Wyatt, B. Soher, P. Maccarini, H. C. Charles, P. Stauffer, and J. Macfall, “Hyperthermia MRI temperature measurement: Evaluation of measurement stabilisation strategies for extremity and breast tumours,” Interna- tional Journal of Hyperthermia, vol. 25, no. 6, pp. 422–433, 2009.

[155] K. K. Vigen, B. L. Daniel, J. M. Pauly, and K. Butts, “Triggered, navigated, multi-baseline method for proton resonance frequency temperature mapping with respiratory motion,” Magnetic Resonance in Medicine, vol. 50, no. 5, pp. 1003–1010, 2003.

[156] B. D. de Senneville, B. Quesson, P. Desbarats, R. Salomir, J. Palussiere, and C. T. W. Moonen, “Atlas-based motion correction for online MR temperature mapping,” in Image Processing, 2004. ICIP’04. 2004 International Conference on, vol. 4, pp. 2571–2574, IEEE, 2004.

[157] S. Hey, G. Maclair, B. D. de Senneville, M. Lepetit-Coiffe, Y. Berber, M. O. Köhler, B. Quesson, C. T. W. Moonen, and M. Ries, “Online correction of respiratory-induced ﬁeld disturbances for continuous MR-thermometry in the breast,” Magnetic Resonance in Medicine, vol. 61, no. 6, pp. 1494–1499, 2009.

[158] W. Chu, R. Staruch, S. Pichardo, Y. Huang, C. Mougenot, M. Tillander, M. O. Köhler, M. Ylihautala, M. McGufﬁn, G. J. Czarnota, and K. Hyny- nen, “MR-HIFU mild hyperthermia for sensitization of radiation and chemotherapy for recurrent rectal cancer: First phase I clinical trial results,” (Singapore), ISMRM 24th Annual Meeting, May 2016.

[159] K. R. Gorny, N. J. Hangiandreou, G. K. Hesley, B. S. Gostout, K. P. McGee, and J. P. Felmlee, “MR guided focused ultrasound: technical acceptance measures for a clinical system,” Physics in Medicine and Biology, vol. 51, no. 12, p. 3155, 2006.

70 References

[160] P. S. Yarmolenko, E. Callardo, A. Partanen, A. Ranjan, D. Woods, A. H. Negussie, C. Burke, L. W. Bartels, M. W. Dewhirst, B. J. Wood, and M. R. Dreher, “MR-HIFU drug paintbrush: Large volume, conformal mild hyperthermia with MR-HIFU used to trigger and monitor release from image- able, temperature sensitive liposomes,” in 3rd International Symposium on Current and Future Applications of MR-Guided Focused Ultrasound, (Bethesda, Washington DC), pp. 102–103, 2012.

[161] X. Fan and K. Hynynen, “Control of the necrosed tissue volume during noninvasive ultrasound surgery using a 16-element phased array,” Medi- cal Physics, vol. 22, no. 3, pp. 297–306, 1995.

[162] W. T. Dixon, “Simple proton spectroscopic imaging.,” Radiology, vol. 153, no. 1, pp. 189–194, 1984.

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