Université Bordeaux 1 Les Sciences et les Technologies au service de l’Homme et de l’environnement

N° d’ordre : 3757

THÈSE

PRÉSENTÉE A

L’UNIVERSITÉ BORDEAUX 1

ÉCOLE DOCTORALE DES SCIENCES à renseigner

Par Roel DECKERS

POUR OBTENIR LE GRADE DE

DOCTEUR

SPÉCIALITÉ : Lasers, matière, nanosciences

LE RÔLE DES ULTRASONS, DE L’IRM ET DE L’IMAGERIE OPTIQUE DANS LE CADRE DE L’ACTIVATION LOCALE

DE GÈNES ET DU DÉPÔT LOCAL DE MÉDICAMENTS

Directeur de recherche : Chrit Moonen

Soutenue le : 19 décembre 2008 Devant la commission d’examen formée de : M. TANTER, Mickael Directeur de Recherche Inserm Rapporteur M. TAVITIAN, Bertrand Chef de Laboratoire CEA (HDR) Rapporteur M. BARTELS, Wilbert Associate Professor UMC Utrecht Examinateur M. BRISSON, Alain Professeur UB1 Président de jury M. VOISIN, Pierre Maître de Conférences UB2 Rapporteur de soutenance M. MOONEN, Chrit Directeur de Recherche CNRS Examinateur

Université Bordeaux 1 Les Sciences et les Technologies au service de l’Homme et de l’environnement

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Part I. Introduction 7

Chapter 1. General introduction 9

References 11

Chapter 2. Ultrasound 13

2.1. Basics of ultrasound 13

2.2. Focused ultrasound 14

2.3. Interaction of ultrasound with tissue 15

2.4. Radiation force 20

2.5. References 21

Chapter 3. MRI 25

3.1. Introduction 25

3.2. Quantum-mechanical description 25

3.3. Classic-mechanical description 26

3.4. Relaxation and signal detection 27

3.5. Image acquisition 28

3.6. MR thermometry 29

3.7. Automatic control of temperature 31

3.8. References 32

Chapter 4. Optical imaging 35

4.1. Introduction 35

4.2. Origin of light 36

4.3. Interaction light-tissue 40

4.4. Light measurement 42

4.5. References 44

Part II. Spatio-temporal control of gene activation 47

Chapter 5. Regulatable gene expression systems 49

5.1. Introduction 49

5.2. References 51

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Chapter 6. Heat shock proteins 53

6.1. Introduction 53

6.2. Hsp promoters in gene therapy 54

6.3. References 55

Chapter 7. In vitro characterization of Hsp70 promoter 57

7.1. Introduction 57

7.2. Materials & methods 60

7.3. Results 62

7.4. Discussion 67

7.5. Conclusions 69

7.6. Reference 70

Chapter 8. In vivo characterization of the Hsp70 promoter 73

8.1. Introduction 73

8.2. Material & methods 74

8.3. Results 76

8.4. Discussion 84

8.5. Conclusion 85

8.6. References 86

Chapter 9. Local gene activation using MR guided HIFU 89

9.1. Introduction 89

9.2. Material & methods 90

9.3. Results 93

9.4. Discussion 98

9.5. Conclusions 102

9.6. References 102

Part III. Local drug delivery 105

Chapter 10. The role of ultrasound and molecular imaging in local drug delivery 107

10.1. Introduction 107

10.2. Ultrasound facilitated local drug delivery 109

10.3. Imaging of drug delivery 111

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10.4. References 117

Chapter 11. MRI monitoring of ultrasound mediated drug delivery 125

11.1. Introduction 125

11.2. Materials and methods 127

11.3. Results 130

11.4. Discussion 133

11.5. References 135

Part IV. Summaries and perspectives 139

Summary 141

Perspectives 145

Résumé 147

Perspectives 151

Part V. Word of thanks and publications 153

Word of thanks 155

List of publications 157

6

7

Part I. Introduction

8

9

Chapter 1. General introduction

Historically, image guided therapy has combined advances in imaging and therapeutic

technology to develop minimally invasive surgical and interventional techniques. Different

imaging modalities, mainly CT and ultrasound, but also MRI and PET/SPECT are used in the

preparation phase, during treatment and for post-treatment evaluation. The different imaging

modalities are increasingly used during minimally invasive procedures for real-time guidance

of instruments. Ultrasound is for example the primary modality for guiding radio-frequency

needles [1] and laser fibers used in tumor ablation [2]. Local injections of anti-inflammatory

drugs (steroids) for herniated disks [3] and the injection of blood cloth removal drugs are

performed with CT guidance [4]. The latest developments in image guided therapy such as

the fusion of imaging technologies with robotics and multi-modality imaging increase even

further the available information and the precision of the intervention [5,6].

MRI guided high intensity focused ultrasound (HIFU) is a completely non-invasive form of

image guided therapy and entered the clinical environment only recently for treatment of

uterine fibroids [7]. Further clinical trials are underway for treatment of cancers in breast,

liver and prostate [8-10].

New non-invasive image guided molecular therapies are evolving. These novel therapies use

molecular imaging techniques for dynamically monitoring of cellular function and molecular

processes in living animals [11]. Molecular imaging allows for (early) identification of

diseased tissue using imaging probes targeted to disease specific markers on cell membrane

[12,13] and tracking cell migration [14]. Combining molecular imaging with nanomedicine

can further improve the efficacy of molecular and cellular therapy. Monitoring of

nanomedicine such as genes, antibodies, chemotherapeutic drugs and drug carriers with

molecular imaging techniques gives insight in the local interaction of the drug with the tissue

and allows for the development of more specific nanomedicine. In this thesis two different

applications areas of image guided molecular therapy using MRI, HIFU and optical imaging

are exploited: local gene activation and local drug delivery.

Gene therapy is an experimental technique that uses genes to treat or prevent disease instead

of using drugs or surgery. With gene therapy it is possible to add a missing gene in the

genome and replace or repair a gene that malfunctions. Although gene therapy is a promising

treatment option for a number of diseases (including inherited disorders, some types of

cancer, and certain viral infections), the technique remains risky (e.g. auto-immune reponse to

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introduced gene and toxicity of viral vector) and is still being investigated to make sure that it

will be safe and effective. One part of the study is based on the delivery of the therapeutic

gene. Autoimmune reactions, toxicity and the low efficacy of gene delivery are here the major

problems to conquer. Secondly, a tight control of gene activation is necessary, because the

objective of gene therapy is to express a therapeutic gene in the region where therapy is

required and for the duration necessary to achieve a therapeutic effect and to minimize

systemic toxicity. This implies the need for spatial and temporal control of gene activation in

vivo, which will be the subject of Part II in this thesis. Part II starts with an overview of the

different approaches for spatial and/or temporal control of local gene activation. In this thesis

local hyperthermia in combination with a temperature sensitive heat shock protein (Hsp)

promoter will be used for local gene activation. In Chapter 6 a biological background on heat

shock proteins is provided, including a discussion on the possible applications of the Hsp

promoter in gene therapy. In the following chapters the influence of different heating

protocols on the promoter activity are determined in order to be able to do treatment planning.

First the characteristics of the Hsp promoter are analyzed in vitro (Chapter 7) and then in vivo

(Chapter 8). Finally, we demonstrate in Chapter 9 the use of MR guided HIFU for the local

activation of a transgene.

Not only gene therapy, but also drug delivery may take advantage of the local deposition of

thermal and/or mechanical energy by means of ultrasound and this is the subject of Part III in

this thesis. In chemotherapy anti-cancer drugs are often administrated systemically resulting

in low concentration in the tumor, hence low treatment efficacy and significant toxic side

effects. Local increase of the cytotoxic drug concentration would ameliorate the efficacy of

the therapy and reduce systemic toxicity. Ultrasound can play an important role in facilitating

local drug delivery, which is explained in Chapter 10. Another important facet of local drug

delivery is monitoring non-invasively the drug’s pharmacokinetics and pharmacodynamics,

giving real time insight in the position and concentration of the drug and its influence on the

system. An overview of the different imaging modalities and their strengths and weakness are

also given in chapter 10. In Chapter 11 we demonstrate that we can improve the deposition of

a macromolecule in the liver with a clinical echograph and that we can follow the process of

delivery dynamically with a clinical MR scanner.

The thesis will begin with an introduction of the most important techniques (ultrasound, MRI

and optical imaging) used for the research performed in this thesis.

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References

[1] Solbiati L, Ierace T, Goldberg SN, Sironi S, Livraghi T, Fiocca R, Servadio G,

Rizzatto G, Mueller PR, Del Maschio A, Gazelle GS. Percutaneous US-guided radio-

frequency tissue ablation of liver metastases: treatment and follow-up in 16 patients.

Radiology 1997;202(1):195-203.

[2] Nolsoe CP, Torp-Pedersen S, Burcharth F, Horn T, Pedersen S, Christensen NE,

Olldag ES, Andersen PH, Karstrup S, Lorentzen T, et al. Interstitial hyperthermia of

colorectal liver metastases with a US-guided Nd-YAG laser with a diffuser tip: a pilot

clinical study. Radiology 1993;187(2):333-337.

[3] Braun J, Bollow M, Seyrekbasan F, Haberle HJ, Eggens U, Mertz A, Distler A, Sieper

J. Computed tomography guided corticosteroid injection of the sacroiliac joint in

patients with spondyloarthropathy with sacroiliitis: clinical outcome and followup by

dynamic magnetic resonance imaging. J Rheumatol 1996;23(4):659-664.

[4] Montes JM, Wong JH, Fayad PB, Awad IA. Stereotactic computed tomographic-

guided aspiration and thrombolysis of intracerebral hematoma : protocol and

preliminary experience. Stroke 2000;31(4):834-840.

[5] Rasmus M, Huegli RW, Bilecen D, Jacob AL. Robotically assisted CT-based

procedures. Minim Invasive Ther Allied Technol 2007;16(4):212-216.

[6] Yap JT, Carney JP, Hall NC, Townsend DW. Image-guided cancer therapy using

PET/CT. Cancer J 2004;10(4):221-233.

[7] Tempany CM, Stewart EA, McDannold N, Quade BJ, Jolesz FA, Hynynen K. MR

imaging-guided focused ultrasound surgery of uterine leiomyomas: a feasibility study.

Radiology 2003;226(3):897-905.

[8] Gelet A, Chapelon JY, Bouvier R, Rouviere O, Lasne Y, Lyonnet D, Dubernard JM.

Transrectal high-intensity focused ultrasound: minimally invasive therapy of localized

prostate cancer. J Endourol 2000;14(6):519-528.

[9] Hynynen K, Pomeroy O, Smith DN, Huber PE, McDannold NJ, Kettenbach J, Baum

J, Singer S, Jolesz FA. MR imaging-guided focused ultrasound surgery of

fibroadenomas in the breast: a feasibility study. Radiology 2001;219(1):176-185.

[10] Kennedy JE, Wu F, ter Haar GR, Gleeson FV, Phillips RR, Middleton MR, Cranston

D. High-intensity focused ultrasound for the treatment of liver tumours. Ultrasonics

2004;42(1-9):931-935.

[11] Weissleder R, Mahmood U. Molecular imaging. Radiology 2001;219(2):316-333.

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[12] Cai W, Chen X. Multimodality molecular imaging of tumor angiogenesis. J Nucl Med

2008;49 Suppl 2:113S-128S.

[13] Dunphy MP, Strauss HW. Molecular imaging of atherosclerosis. Curr Cardiol Rep

2008;10(2):121-127.

[14] Bulte JW, Kraitchman DL. Monitoring cell therapy using iron oxide MR contrast

agents. Curr Pharm Biotechnol 2004;5(6):567-584.

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Chapter 2. Ultrasound

2.1. Basics of ultrasound

Ultrasound is a pressure wave that propagates within a medium, inducing mechanical

vibrations of particles at a frequency above 20 kHz. In most cases, the oscillatory

displacement of particles is in the direction of wave propagation (i.e. longitudinal wave),

creating regions with high pressure (compressions) and low pressure (rarefaction) (Figure

2.1). The wavelength (λ) of the ultrasound wave, which is distance between two rarefactions,

depends on both frequency (f) and propagation speed (c) of ultrasound:

f

c=λ 2-1

The typical wavelength in soft tissue is about 1 mm at a frequency of 1.5 MHz. The

propagation speed of ultrasound is about 1550 ms-1 for soft tissue, independent of the

ultrasound frequency. In fatty tissue the average speed is only slightly lower (1480 ms-1),

whereas in air spaces a value of 343 ms-1 is found. In bone, the speed is much higher (between

1800 and 3700 ms-1). When ultrasound meets the interface between two media, it may be

partially reflected and partially transmitted, depending on the incident angle and the

difference in acoustic impedance of the two media. The large difference in acoustic

impedance (product of propagation speed in medium and density of medium) between air and

soft tissue causes an almost complete reflection of the ultrasound wave at tissue/air interfaces.

At the bone/soft tissue interface there is also a complete reflection at incident angles larger

than 30º.

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Figure 2.1 Longitudinal and transverse waves. In longitudinal waves the oscillatory

displacement of particles is in the direction of wave propagation, creating regions with high

pressure (compressions) and low pressure (rarefaction). In contrast, the particle motion in

transverse waves is perpendicular to the wave motion. The distance between two succeeding

rarefaction (or compressions) equals the wave length (λ).

2.2. Focused ultrasound

In general, focused ultrasound is generated using a curved resonant piezo-electric element.

The principle is based on interference of ultrasound waves. At the focal point, waves

originating from different points on the ultrasound transducer are in phase, resulting in

constructive inference and thus a maximal ultrasound intensity. The shape of the focal point

depends on the architecture of the transducer and the frequency. The geometry of a transducer

is often described by an F-number, which is the ratio between the focal depth (l) and the

diameter of aperture (d) of the transducer (F-number = l/d). A smaller F-number leads to

stronger focusing with a shorter focal point length along the beam axis (i.e. axial resolution).

The minimum width of the focal point is half λ (i.e. lateral resolution), thus by increasing the

frequency, the size of the focal point will be reduced. An alternative method to focus

ultrasound is the combination of multiple small ultrasound elements, a so-called phased array.

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2.3. Interaction of ultrasound with tissue

It is well known that the interaction of ultrasound with tissue can produce a wide variety of

biological effects [1,2]. The underlying physical mechanisms include acoustic local heating,

radiation pressure and cavitation. Below the physical principles associated with these three

mechanisms and their bioeffects are discussed in more detail.

2.3.1. Heat

The intensity of an ultrasonic wave travelling through a medium may be attenuated. In

experimental studies, attenuation has been found to be dominated by absorption. During

absorption, the mechanical energy of the acoustic wave (micrometer displacements with a

frequency in the kHz-MHz range) is predominantly converted into heat (atomic vibrations

with sub-nanometer displacements with frequencies well above 1GHz). There are several

mechanisms by which absorption can occur. However, the exact mechanisms by which

ultrasound is absorbed by biological materials are rather complicated. It has been observed

that, within the frequency range used for medical ultrasonic imaging (2-15 MHz), most

tissues have an absorption coefficient that is linearly proportional to the frequency [3]. The

intensity Ix at depth x with respect to that at the original position (I0) for plane wave in tissue

may be described as

100 10

x

x II⋅−

⋅=α

2-2

where the symbol α represents the absorption coefficient of the wave amplitude per unit path

length and usually lies in the range of 50 to 350 dB m-1 MHz-1 in soft tissue [4]. For example,

using 1.5 MHz ultrasound, the intensity in muscle will drop to about 50% at 50 mm

penetration. For a given absorption coefficient, the rate of heat generation is proportional to

the local intensity, and thus diminishes also in an exponential way with increasing depth.

However, by using focused ultrasound instead of planar waves much higher intensities can be

obtained at the focal point, resulting in heating of deep lying tissue with relatively low

increase of the temperature of tissue between the target region and the transducer [5].

The choice of the ultrasound frequency to be used for a given application is a compromise

between tissue penetration, spot size and heat generation and changes for different sonication

depths. As mentioned before, the attenuation of the ultrasound by the tissue in the path of the

beam will increase with frequency. However, the increased absorption coefficient due to the

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higher ultrasound frequency also results in better heating. This is illustrated in Figure 2.2

where the energy density and temperature are plotted for a simulation of acoustic heating at

two different depths in tissue. Generally, the ultrasound frequencies for HIFU heating are in

the lower megahertz range. Hyperthermia finds its applications in gene therapy as well as in

local drug delivery. Both applications will be discussed in more detail in Part III.

Figure 2.2 Energy density and temperature as function of penetration depth in tissue. The

absorption coefficient is linearly proportional to the ultrasound frequency. Therefore, the

energy density after a certain distance of tissue penetration is always lower for the higher

frequency (a and b). However, the heat generation is linearly proportional to the energy

density and the absorption coefficient. Therefore, the optimal frequency for heating depends

on the penetration depth in tissue (c and d).

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2.3.2. Cavitation

Acoustic cavitation is defined as the formation and/or activity of gas-filled bubbles in a

medium exposed to ultrasound [6]. These gas-filled bubbles can be of natural origin or

artificially made (i.e. microbubbles). The sustained growth of cavitation bubbles and their

oscillations over several acoustic cycles is known as stable or non inertial cavitation [7]. In

contrast, when a cavitation bubble grows violently and collapses in less then a cycle this is

called transient or inertial cavitation [7]. In general, the likelihood and intensity of inertial

cavitation increases at higher pressures (p in MPa) and lower frequencies (f in MHz) [8].

These two exposure parameters have been combined in a single mechanical index (MI)

defined as [9]:

f

pMI = 2-3

The pressure (p) of an ultrasound wave is related to the acoustic intensity (I) as follows:

c

pI

ρ2

2

= 2-4

where ρ and c are the density and ultrasound propagation speed of the medium, respectively.

This equation indicates that high intensities obtained with HIFU also results in high local

pressures.

The cavitation process will also be affected by the number and size of the cavitation bubbles,

the available spaces for the bubbles to oscillate and their physical properties such as the type

of gas in its interior and the composition of its shell [7]. These parameters, among others (i.e.

pulse length, pulse repetition frequency and duration of exposure), were recently investigated

for ultrasound/microbubble-mediated gene delivery into cultured adherent cells by Rahim et

al. [10]. The authors found an approximately linear-dependence between the gene delivery

efficiency and acoustic pressure over the range 0.1-0.5 MPa. The optimized parameters for

high levels of gene delivery and cell viability using 1 MHz US were found to be a pulse

pressure amplitude of 0.25 MPa (peak-negative), a pulse repetition frequency of 1kHz, 40

cycles pulse length and 10 s exposure. Other studies on the subject have been published (e.g.

[11]). Since many parameters are involved (including tissue parameters), some variability in

optimal parameters may not be surprising.

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Stable Cavitation

Stable cavitation occurs at low pressure waves and is associated with the physical phenomena

of rectified diffusion, microstreaming, and coalescence of small bubbles to form larger ones

[12]. Rectified diffusion describes the slow growth of an oscillating bubble related to a net

inflow of gas into the bubble over successive cycles, until it reaches a resonant size [13]. At

resonant size the bubble will show stable, low amplitude oscillations. The resonance

frequency (fr) of this linear, undamped system is given by [14]:

ργ

πP

Rf r

3

2

1

0

= 2-5

where R0 is the initial radius of the bubble, γ is the heat capacity ratio (γ equals 1.4 for a

diatomic gas such as oxygen), P is the ambient pressure and ρ is the density of the

surrounding medium. Therefore, a reasonable estimation of resonance frequency (in MHz) of

a free spherical gas bubble (with an initial radius R0 in µm) in water at 20° C and atmospheric

pressure is given by:

0

3,3

Rf r ≈ 2-6

The low amplitude oscillations induce inhomogeneous cyclic pressure fields around the

bubbles, which cause small flows in the bubble surrounding fluid, a process known as

microstreaming [15]. Marmottant et al. showed that weak oscillations related to stable

cavitation can be sufficient to rupture single cells. The microstreaming creates high shear

stresses (10-2 N m-1) near the bubble surface that may increase strain in the membrane of

surrounding cells and even exceed its critical value for membrane rupture [16]. This may

result in cell wall permeation, a phenomenon advantageous for drug delivery across the cell

membrane. This effect was clearly visualized by Van Wamel et al., showing the uptake of

propium iodide (PI) by a cell that was deformed by microbubble oscillations whereas an un-

deformed reference cell did not show uptake of PI [17].

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Figure 2.3 Oscillation of microbubble causes uptake of propium iodide (PI) by mechanically

deformed cell. The frames are recorded by an ultra fast camera (Brandaris) (A) First frame of

a Brandaris recording in which contours (membranes) of the cells are drawn (dashed lines).

Two cells can be distinguished as well as the intercellular space and the microbubbles. (B)

Brandaris recording: 6 selected frames out of a total of 128 frames. The pushing and pulling

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behavior of the vibrating microbubbles nearby the cells is shown. Cell I was not deformed,

cell II was deformed (arrows in frame 0009). (C) Fluorescence images of PI uptake of the two

cells are shown. The first frame is before US exposure and the two circles indicate the

position of the microbubbles. The second frame is taken 0.25 min after US was turned off, in

this image the deformed cell II shows PI uptake, whereas the un-deformed cell I shows no PI

uptake. The deformed cell contained PI only for a short period of time. The last frame is a

bright field image of the cells 3 min after US exposure. Figure adapted from [17].

Inertial cavitation

Inertial cavitation occurs at higher acoustic pressure and is associated with several physical

phenomena [18]. During the rapid collapse of the gas bubble, the inward moving wall of fluid

has sufficient inertia so that it cannot reverse direction when the acoustic pressure reverses

direction, but continues to compress the gas in the bubble to a very small volume. During this

process the pressure and temperature can reach thousands of bars and degrees Kelvin possibly

even leading to emission of light (sonoluminescence) [19] and production of free radicals

(sonochemistry) [20]. The collapse of the bubble also generates shock waves that spherically

diverge in the surrounding environment of the bubble. When collapsing bubbles are in the

vicinity of solid boundaries (e.g. cell membranes), the collapse will be asymmetrical and can

result in the formation of high speed, fluid microjets [21]. The preceding physical phenomena

may facilitate drug delivery in two ways. First, they can disrupt the shell of drug delivery

systems thereby releasing the enclosed drugs. The fragmentation thresholds for several

ultrasound contrast agents (i.e. microbubbles) were investigated by Chen et al. [22] and Shi et

al. [23]. Secondly, these phenomena can cause cell membrane permeabilization and capillary

rupture. Lokhandwalla and Sturtevant performed a theoretical analysis of how shock waves,

bubble wall motion (i.e. bubble expansion and collapse) and microjets may affect membrane

permeability. They show that these mechanisms can cause the membrane to deform beyond

the threshold strain for rupture [24]. Sundaram et al. performed experimental and theoretical

analyses confirming the influence of inertial cavitation in cell membrane permeabilization

[25].

2.4. Radiation force

Although heat deposition and cavitation are the most widely investigated ultrasound-related

mechanisms producing biological effects in tissues, there is an increasing amount of evidence

that radiation forces may be used for enhanced delivery during high intensity focused

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ultrasound (HIFU) exposures [26,27]. Radiation forces are produced by HIFU pulses, in

which the absorption and/or reflection of acoustic energy causes a transfer of momentum

from the ultrasound wave to the medium [28]. The primary radiation force acts in the

direction of acoustic wave propagation and a secondary radiation force acts between

individual bubbles [29]. Radiation force-induced displacements may cause shear forces

between displaced and non-displaced tissue. The strain resulting from these stresses may

induce gaps between endothelial cells [30,31] and widening intracellular spaces in epithelial

tissue [32,33]. The former effect will increase the drug extravasation from the vasculature and

the latter effect will increase the intracellular drug diffusion. Although more in depth

investigations are necessary to show that this proposed mechanism is conclusive, there are

several preclinical studies that show the feasibility of this mechanism. The combination of

pulsed-HIFU targeted exposure with systemically administered agents has been used to

increase local delivery of a magnetic resonance contrast agent [34], a high molecular weight

fluorescein isothyocyanate (FITC)-dextran [27], doxorubicin (DoxilTM) [35] and

ThermodoxTM [36]. The primary radiation force may also be used to direct delivery vehicles,

circulating in the blood pool, near a vessel wall, where they will move at a reduced velocity

compared to the vehicles in the centre of the vessel [37]. In addition, the secondary radiation

force cause individual particles to attract each other [38], resulting in an even larger

microbubble concentration near the vessel wall and thus in an enhanced receptor-ligand

contact which may be beneficial for targeted drug delivery [39,40].

2.5. References

[1] Dalecki D. Mechanical bioeffects of ultrasound. Annu Rev Biomed Eng 2004;6:229-

248.

[2] Miller MW, Miller DL, Brayman AA. A review of in vitro bioeffects of inertial

ultrasonic cavitation from a mechanistic perspective. Ultrasound Med Biol

1996;22(9):1131-1154.

[3] Hueter TF. Messung der Ultraschallabsorption in tierschen Geweben und ihre

Abhangigkeit von der Frequenz. Naturwiss 1948;35(9).

[4] Duck FA. Physical properties of tissue; A comprehensive reference book. London:

Academic Press; 1990.

[5] Lele PP. Induction of deep, local hyperthermia by ultrasound and electromagnetic

fields: problems and choices. Radiat Environ Biophys 1980;17(3):205-217.

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[6] Apfel RE. Acoustic cavitation: a possible consequence of biomedical uses of

ultrasound. Br J Cancer Suppl 1982:140-146.

[7] Leighton TG. The acoustic bubble; 1994.

[8] Hill CR. Ultrasonic Exposure Thresholds for Changes in Cells and Tissues. The

Journal of the Acoustical Society of America 1972;52(2B):667-672.

[9] Apfel RE, Holland CK. Gauging the likelihood of cavitation from short-pulse, low-

duty cycle diagnostic ultrasound. Ultrasound Med Biol 1991;17(2):179-185.

[10] Rahim A, Taylor SL, Bush NL, ter Haar GR, Bamber JC, Porter CD. Physical

parameters affecting ultrasound/microbubble-mediated gene delivery efficiency in

vitro. Ultrasound Med Biol 2006;32(8):1269-1279.

[11] Zarnitsyn VG, Prausnitz MR. Physical parameters influencing optimization of

ultrasound-mediated DNA transfection. Ultrasound Med Biol 2004;30(4):527-538.

[12] Boyle RW, Taylor GB, Froman DK. Cavitation in the track of an ultrasound beam.

Trans Roy Soc Canada Section III 1929;23:187-201.

[13] Harvey EN, Barnes DK, McElroy WD, Whiteley AH, Pease DC, Cooper KW. Bubble

formation in animals. I. Physical factors. J Cell Comp Physiol 1944;24:1-22.

[14] Minnaert M. On musical air-bubbles and sounds of running water. Phil Mag

1933;16:235-248.

[15] Nyborg WL. Basic physics of low frequency therapeutic ultrasound. Boston: Kluwer

Academic; 1996.

[16] Marmottant P, Hilgenfeldt S. Controlled vesicle deformation and lysis by single

oscillating bubbles. Nature 2003;423(6936):153-156.

[17] van Wamel A, Kooiman K, Harteveld M, Emmer M, ten Cate FJ, Versluis M, de Jong

N. Vibrating microbubbles poking individual cells: Drug transfer into cells via

sonoporation. Journal of Controlled Release 2006;112(2):149-155.

[18] Flynn HG, Church CC. Transient pulsations of small gas bubbles in water. J Acoust

Soc Am 1988;84:1863-1876.

[19] Hilgenfeldt S, Grossmann S, Lohse D. A simple explanation of light emission in

sonoluminescence. Nature 1999;398(6726):402-405.

[20] Suslick KS. Ultrasound: its chemical, physical and biological effects. London: VCH

Publishers; 1989.

[21] Blake JR, Gibson DC. Cavitation bubbles near boundaries. Ann Rev Fluid Mech

1987;19:99-123.

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[22] Chen WS, Matula TJ, Brayman AA, Crum LA. A comparison of the fragmentation

thresholds and inertial cavitation doses of different ultrasound contrast agents. J

Acoust Soc Am 2003;113(1):643-651.

[23] Shi WT, Forsberg F, Tornes A, Ostensen J, Goldberg BB. Destruction of contrast

microbubbles and the association with inertial cavitation. Ultrasound Med Biol

2000;26(6):1009-1019.

[24] Lokhandwalla M, McAteer JA, Williams JC, Sturtevant B. Mechanical hemolysis in

shock wave lithotripsy (SWL): II. In vitro cell lysis due to shear. Phys Med Biol

2001;46:1245-1264.

[25] Sundaram J, Mellein BR, Mitragotri S. An experimental and theoretical analysis of

ultrasound-induced permeabilization of cell membranes. Biophys J 2003;84(5):3087-

3101.

[26] Dittmar KM, Xie J, Hunter F, Trimble C, Bur M, Frenkel V, Li KC. Pulsed high-

intensity focused ultrasound enhances systemic administration of naked DNA in

squamous cell carcinoma model: initial experience. Radiology 2005;235(2):541-546.

[27] Yuh EL, Shulman SG, Mehta SA, Xie J, Chen L, Frenkel V, Bednarski MD, Li KC.

Delivery of systemic chemotherapeutic agent to tumors by using focused ultrasound:

study in a murine model. Radiology 2005;234(2):431-437.

[28] Rooney JA, Nyborg WL. Acoustic Radiation Pressure in a Traveling Plane Wave.

American Journal of Physics 1972;40(12):1825-1830.

[29] Bjerknes VFK. Fields of Force. New York: Columbia University Press; 1906.

[30] Mesiwala AH, Farrell L, Wenzel HJ, Silbergeld DL, Crum LA, Winn HR, Mourad

PD. High-intensity focused ultrasound selectively disrupts the blood-brain barrier in

vivo. Ultrasound Med Biol 2002;28(3):389-400.

[31] Seidl M, Steinbach P, Worle K, Hofstadter F. Induction of stress fibres and

intercellular gaps in human vascular endothelium by shock-waves. Ultrasonics

1994;32(5):397-400.

[32] Frenkel V, Kimmel E, Iger Y. Ultrasound-facilitated transport of silver chloride

(AgCl) particles in fish skin. J Control Release 2000;68(2):251-261.

[33] Frenkel V, Kimmel E, Iger Y. Ultrasound-induced intercellular space widening in fish

epidermis. Ultrasound Med Biol 2000;26(3):473-480.

[34] Bednarski MD, Lee JW, Callstrom MR, Li KC. In vivo target-specific delivery of

macromolecular agents with MR-guided focused ultrasound. Radiology

1997;204(1):263-268.

24

[35] Frenkel V, Etherington A, Greene M, Quijano J, Xie J, Hunter F, Dromi S, Li KC.

Delivery of liposomal doxorubicin (Doxil) in a breast cancer tumor model:

investigation of potential enhancement by pulsed-high intensity focused ultrasound

exposure. Acad Radiol 2006;13(4):469-479.

[36] Dromi S, Quijano J, Xie J, Frenkel V, wood B, Li K. Pulsed-high intesity focused

ultrasound (HIFU) enhanced delivery of doxorubicin using heat sensitive liposome

(Thermodox). 2005; Chicago.

[37] Dayton P, Klibanov A, Brandenburger G, Ferrara K. Acoustic radiation force in vivo:

a mechanism to assist targeting of microbubbles. Ultrasound Med Biol

1999;25(8):1195-1201.

[38] Dayton PA, Morgan KE, Klibanov AL, Brandenburger G, Nightingale KR, Ferrara

KW. A preliminary evaluation of the effects of primary and secondaryradiation forces

on acoustic contrast agents. Ultrasonics, Ferroelectrics and Frequency Control, IEEE

Transactions on 1997;44(6):1264-1277.

[39] Borden MA, Sarantos MR, Stieger SM, Simon SI, Ferrara KW, Dayton PA.

Ultrasound radiation force modulates ligand availability on targeted contrast agents.

Mol Imaging 2006;5(3):139-147.

[40] Lum AF, Borden MA, Dayton PA, Kruse DE, Simon SI, Ferrara KW. Ultrasound

radiation force enables targeted deposition of model drug carriers loaded on

microbubbles. J Control Release 2006;111(1-2):128-134.

25

Chapter 3. MRI

3.1. Introduction

Magnetic resonance imaging (MRI) is based on the phenomenon of nuclear magnetic

resonance (NMR), which involves the measurement of signals induced in a receiver coil by

atomic nuclei in response to radio waves. The physical principles of NMR are based on

quantum mechanics, but most phenomena concerning MRI can be explained with classical

mechanics.

3.2. Quantum-mechanical description

Depending on the number of neutrons and protons of which it is composed, a nucleus can

have a net spin, i.e. its spin quantum number, I, is non-zero. Think of the spin of this nucleus

as a magnetic moment vector, causing the proton to behave like a tiny magnet with a north

and south pole. Placed in an external magnetic field, this magnetic moment tends to align

itself in the direction of the field. In quantum mechanical terms, the magnetic moment can

align in only 2I + 1 ways with an external magnetic field; the magnetic moment can only

occupy discrete states that have different energy levels. For hydrogen (1H), a nucleus with

spin ½, only two spin states exist. The energy difference between the lower and higher level,

∆E, depends on the gyromagnetic ratio (γ) of the nucleus and the strength of the local

magnetic field (B0):

πγ2

0hBE =∆ 3-1

where h is Plancks constant. A spin can undergo a transition between the two energy states by

the absorption or emission of a photon of frequency ν0, such that

0000

22B

hhBE γυυ

ππγ

=→==∆ 3-2

Expressing the frequency in angular terms gives the Larmor equation:

00 Bγω = 3-3

where ω0 is the Larmor frequency.

26

When a group of spins is placed in a magnetic field, each spin aligns in one of the two

possible orientations. For spin ½ nuclei, like hydrogen, the relative population of the two spin

states is determined by the Boltzmann distribution:

kTEen

n /∆−+

−

= 3-4

where k is the Boltzmann constant, T is the temperature in Kelvin and ∆E is the energy

difference between the two states, given by equation 3-1 . NMR experiments involve the

disturbance of the Boltzmann distribution by absorption and emission of quanta with energy

∆E. However, it is very cumbersome to describe NMR experiments in a quantum mechanical

way, a classic mechanical approach is more convenient.

3.3. Classic-mechanical description

Living tissue consists of 60% to 80% of water in which macro-molecules are suspended. In

both, water and macro-molecules, hydrogen spins represent the largest group of MR-

observable nuclei. In most aspect of MR experiments the group of nuclei under observation

behaves like a large ensemble of spin ½ nuclei and can be described by classical mechanics.

Herein the net magnetization of the ensemble of spin ½ nuclei is represented by a

magnetization vector (M). If the magnetization vector is placed in a static magnetic field B0,

M will experience a torque. The motion of M is described by the Bloch equation [1]:

( )BMdt

dM ×= γ 3-5

This equation describes the precession of M around B0. The angular frequency of the

precession is identical to the Larmor frequency derived in the quantum mechanical

description above (equation 3-3), showing how the classical and quantum mechanical pictures

coincide. In the presence of a large main magnetic field (B0) as well as magnetic field (B1),

applied perpendicular to the B0-field and oscillating at ω0, the magnetization vector will

precess simultaneously about B0 at ω0 and B1 at ω1. This means M will spiral down from

longitudinal plane into the transversal plane when viewed in the laboratory frame of reference

(Figure 3.1a). In a new frame of reference, the rotating frame, which rotates about z-axis at

ω0, the magnetization vector is rotated about the x-axis at an angular frequency of ω1 (Figure

3.1b).

27

3.4. Relaxation and signal detection

The most common way to carry out an NMR experiment is to apply a short burst of resonant

radio frequency (RF) pulse (B1) to rotate the magnetization vector from the longitudinal plane

into the transverse plane. Once in the transverse plane the magnetization can be detected as it

precesses about z-axis, and this is what gives rise to the NMR signal.

The application of a resonant RF pulse disturbs the spin system. The thermal equilibrium state

of the system will be restored by a process known as spin-lattice relaxation. This involves

exchange of energy between the spin system and its surroundings and the rate at which the

equilibrium is restored is characterised by the spin-lattice or longitudinal relaxation time, T1.

In addition to the T1-relaxation effect, the NMR signal will decrease due to a second and

faster relaxation process, spin-spin relaxation. In this process spins exchange energy amongst

them, resulting in no net change in the population of the energy levels, however it leads to a

lost of phase coherence. The rate of lost of phase coherence is characterised by the spin-spin

or transversal relaxation time, T2. Microscopic and macroscopic field inhomogeneities also

contribute to the lost of phase coherence. This will lead to signal decay with a time constant

that is shorter than T2, which is generally referred to as T2*. The relaxation times T1, T2 and

T2* are very important in imaging, as they have the greatest effect in determining image

contrast.

Figure 3.1 Development of the magnetization vector after a 90° RF-pulse excitation in the

laboratory (a) and rotating frame (b) of reference.

28

3.5. Image acquisition

The principle behind all magnetic resonance imaging is the Larmor equation, which shows

that the resonance frequency ω of a spin is proportional to the magnetic field, B. This means

that a measurement of precession frequency of the magnetization gives information on the

field experienced by that group of spins. By manipulating the spatial variation of the

magnetization field in a known way, this frequency information will yield spatial information.

This idea was first proposed by Lauterbur [2] and Mansfield [3] and lies at the origin of MR

imaging. In practice this is done by superimposing a series of linear magnetic field gradients

in three perpendicular directions onto the main magnetic field. The purpose of the magnetic

field gradients are slice selection and position encoding within the selected slice, known as

frequency and phase encoding. By loop-wise repetition of frequency and phase encoding in a

selected slice, a 2-D frequency map can be obtained. To reconstruct the real image, which is

the spin density distribution, a 2-D Fourier transform is applied, which results in a complex

image. The magnitude and phase image are calculated from the real and imaginary part of the

complex image Figure 3.2). Notice that MR images are not pure proton density images but

represent a weighted proton density that depends on T1, T2, possibly T2* and the acquisition

parameters (repetition time, echo time, flip angle).

Figure 3.2 Example of magnitude (a) and phase (b) MR image.

29

3.6. MR thermometry

A number of parameters that play a role in a MRI experiment, such as the spin-lattice

relaxation time T1 [4], the molecular diffusion coefficient D [5] and the water proton

resonance frequency (PRF) [6,7], are temperature dependent. All MR thermometry in this

study is based on the PRF and therefore will be explained in more detail below. The reader is

referred to recent reviews on MR thermometry by Quesson et al. [8] and Rieke et al. [9] for

more information on other MR thermometry methods.

The temperature dependence of the PRF was first observed by Hindman [10] while studying

the intermolecular forces and hydrogen bond formation between water molecules. MR-

temperature monitoring on the basis of the PRF was first proposed by Ishihara et al. [11] and

De Poorter et al. [12].

The resonance frequency of a nucleus in a molecule is determined by the local magnetic field

Bnuc it experiences, which is a function of the main magnetic field B0 and the chemical shift

σ(T):

0))(1()( BTTBnuc σ+= 3-6

The chemical shift field (in ppm) is the sum of temperature-indepedent contributions σ0 and a

temperature-dependent contribution σT(T):

)()( 0 TT Tσσσ += 3-7

The chemical shift field can be calculated from the phase information in gradient echo

images:

TEBTT ⋅⋅⋅=Φ 0)()( σγ 3-8

where Φ is the image phase, γ is the gyromagnetic ratio and TE is the echo time. In order to

measure only temperature-dependent changes in chemical shift, the term σ0 has to be

cancelled out, which is typically accomplished by subtraction of the field distribution

measured at a given reference temperature T0 (before heating) from the field distribution

measured at temperature T (during heating), resulting in:

TEB

TTTTT

⋅⋅⋅Φ−Φ

=−=∆0

00

)()(

γα 3-9

30

where α is the temperature-dependent water chemical shift in ppm/ºC, this process is also

illustrated in Figure 3.3. A (temperature dependent) contrast change can be observed in the

phase image at the location of the heating (arrow). From the preceding explanation it follows

that the PRF based method measures the temperature difference with respect to reference

temperature. Therefore it is very important to have a reliable measure of this reference

temperature and avoid non-temperature related phase changes. Reference temperature

measurements can be done by absolute MR temperature measurements [13], temperature

sensitive contrast agents [14,15] or by introducing a small thermometer. During all

experiments in this study animal temperature was measured with a rectal probe, as an

indicator of basal tissue temperature [16]. In principle, any gradient-echo method can be used

for PRF-based MR thermometry. However, gradient echo methods can not recover phase

losses from magnetic field inhomogeneities, magnetic susceptibilities and water-fat

incoherences, which introduce non-temperature related phase changes and lead to

thermometry errors. Spin echo sequences can not be used since the temperature-induced phase

contribution will be cancelled by the refocusing pulse.

Figure 3.3 Calculation of temperature map using the phase information of gradient echo

images. Subtraction of the phase image before heating (Φ0) from the phase image during

heating (Φ), multiplied with the PRF constant leads to a temperature map. The arrow

indicates the location of heating.

The main advantages of the PRF based method are its near independence of tissue

composition and its high spatial and temporal resolution, which allows for continuous

temperature monitoring. The most prevalent problem for temperature monitoring with the

PRF based method is motion, because it introduces non-temperature related phase changes.

Motion artefacts during the MR thermometry sequence can be divided in two categories,

intra-scan and inter-scan motion, based on the time scale of the motion as compared to the

31

image acquisition time. Intra-scan motion is caused by the movement of an object during MR

image acquisition, yielding a low quality image with typical ghosting and blurring. These

motion artefacts can be reduced by accelerating the image acquisition. However, trade-offs

between acquisition time and SNR and temperature uncertainty have to be considered. Inter-

scan motion occurs due to motion or displacement of an object between the acquisition of

consecutive images. Examples of inter-scan motion are accidental patient motion or periodic

respiratory and cardiac motion. Reduction of inter-scan motion artefacts is achieved by

motion restraining, synchronization of acquisition with motion [17] or image processing

techniques such as the atlas-based motion correction proposed by Denis de Senneville et al.

[18].

3.7. Automatic control of temperature

Real-time temperature mapping during the hyperthermic procedure allows the development of

automatic feedback coupling of the heating device, ultrasound in this case. This technique is

known as MR guided HIFU [19,20]. In this method online monitoring of temperature

distribution is performed by continuous acquisition of gradient echo images. Temperature

maps are calculated online (using the phase information of GE images) to dynamically

visualize the tissue temperature and to adjust continuously the HIFU power to force the

temperature at the focal point to follow a predefined value. Figure 3.4 shows an example of

this automatic regulation process. The temperature evolution versus time is displayed with in

black the target temperature and in red the measured temperature at the focal point. Notice

that the temperature is displayed as temperature increase with regard to a reference

temperature. MR guided HIFU allows the temperature to be adjusted with a precision in the

range of 1º C [21].

32

Figure 3.4 Typical time course of the temperature evolution during a heating experiment (a),

with the target temperature (in black) and the measured temperature at the focal point (in

red).

3.8. References

[1] Bloch F. Nuclear induction. Phys Rev 1946;70:460.

[2] Lauterbur PC. Image formation by induced local interactions: examples employing

nuclear magnetic resonance. Nature 1973;242:190-191.

[3] Mansfield P, Grannell PK. NMR 'diffraction' in solids? J Phys C: Solid State Phys

1973;6:L422-L426.

[4] Parker DL, Smith V, Sheldon P, Crooks LE, Fussell L. Temperature distribution

measurements in two-dimensional NMR imaging. Med Phys 1983;10(3):321-325.

[5] Le Bihan D, Delannoy J, Levin RL. Temperature mapping with MR imaging of

molecular diffusion: application to hyperthermia. Radiology 1989;171(3):853-857.

[6] De Poorter J. Noninvasive MRI thermometry with the proton resonance frequency

method: study of susceptibility effects. Magn Reson Med 1995;34(3):359-367.

[7] Ishihara Y, Calderon A, Watanabe H, Okamoto K, Suzuki Y, Kuroda K, Suzuki Y. A

precise and fast temperature mapping using water proton chemical shift. Magn Reson

Med 1995;34(6):814-823.

[8] Quesson B, de Zwart JA, Moonen CT. Magnetic resonance temperature imaging for

guidance of thermotherapy. J Magn Reson Imaging 2000;12(4):525-533.

33

[9] Rieke V, Butts Pauly K. MR thermometry. J Magn Reson Imaging 2008;27(2):376-

390.

[10] Hindman J. Proton resonance shift of water in the gas and liquid states. J Chem Phys

1966;44:4582–4592.

[11] Ishihara Y, Calderon A, Watanabe H, Mori K, Okamoto K, Suzuki Y, Sato K, Kuroda

K, Nakagawa S, Tsutsumi S. A precise and fast temperature mapping method using

water proton chemical shift.; 1992; Berlin. p 4803.

[12] De Poorter J, De Wagter C, De Deene Y, Thomsen C, Stahlberg F, Achten E.

Noninvasive MRI thermometry with the proton resonance frequency (PRF) method: in

vivo results in human muscle. Magn Reson Med 1995;33(1):74-81.

[13] Kuroda K, Suzuki Y, Ishihara Y, Okamoto K, Suzuki Y. Temperature mapping using

water proton chemical shift obtained with 3D-MRSI: feasibility in vivo. Magn Reson

Med 1996;35(1):20-29.

[14] Fossheim SL, Il'yasov KA, Hennig J, Bjornerud A. Thermosensitive paramagnetic

liposomes for temperature control during MR imaging-guided hyperthermia: in vitro

feasibility studies. Acad Radiol 2000;7(12):1107-1115.

[15] Pakin SK, Hekmatyar SK, Hopewell P, Babsky A, Bansal N. Non-invasive

temperature imaging with thulium 1,4,7,10-tetraazacyclododecane-1,4,7,10-

tetramethyl-1,4,7,10-tetraacetic acid (TmDOTMA-). NMR Biomed 2006;19(1):116-

124.

[16] Flanagan SW, Ryan AJ, Gisolfi CV, Moseley PL. Tissue-specific HSP70 response in

animals undergoing heat stress. Am J Physiol 1995;268(1 Pt 2):R28-32.

[17] Morikawa S, Inubushi T, Kurumi Y, Naka S, Sato K, Demura K, Tani T, Haque HA.

Feasibility of respiratory triggering for MR-guided microwave ablation of liver tumors

under general anesthesia. Cardiovasc Intervent Radiol 2004;27(4):370-373.

[18] Denis de Senneville B, Quesson B, Desbarats P, Salomir R, Palussiere J, Moonen

CTW. Atlas-based motion correction for on-line MR temperature mapping.; 2004;

Singapore. p 2571–2574.

[19] Salomir R, Vimeux FC, de Zwart JA, Grenier N, Moonen CT. Hyperthermia by MR-

guided focused ultrasound: accurate temperature control based on fast MRI and a

physical model of local energy deposition and heat conduction. Magn Reson Med

2000;43(3):342-347.

34

[20] Vimeux FC, De Zwart JA, Palussiere J, Fawaz R, Delalande C, Canioni P, Grenier N,

Moonen CT. Real-time control of focused ultrasound heating based on rapid MR

thermometry. Invest Radiol 1999;34(3):190-193.

[21] Mougenot C, Salomir R, Palussiere J, Grenier N, Moonen CT. Automatic spatial and

temporal temperature control for MR-guided focused ultrasound using fast 3D MR

thermometry and multispiral trajectory of the focal point. Magn Reson Med

2004;52(5):1005-1015.

35

Chapter 4. Optical imaging

4.1. Introduction

Next to the established imaging techniques such as nuclear medicine imaging and magnetic

resonance imaging, optical imaging starts playing a growing role in the characterization and

measurement of biological processes at the cellular and molecular level. Monitoring of

biological processes in vivo with fluorescence or bioluminescence imaging is hampered by the

absorption and scattering of light. Labeling the molecules and cells of interest with optical

contrast agent (in the near infrared (NIR) region) can improve the quality of the acquired

optical images. Optical labeling techniques can be broadly divided into three categories: small

organic dyes, quantum dots and reporter genes. A large variety of organic dyes are

commercially available and their application in in vivo imaging studies is growing. When

conjugated covalently to targeting molecules, sensitive (near infrared) contrast agents are

created that can be used for enhanced detection of early cancer [1,2], drug target assessment

[3] and imaging of apoptosis [4]. Quantum dots are semiconductor crystals and provide a new

class of biomarkers that could overcome the limitations of organic dyes such as high

photobleaching and low quantum yield. For the moment most applications of quantum dots

are still in vitro or ex vivo, but the number of in vivo applications is expanding [5].

Both organic dyes and quantum dots are examples of so-called direct imaging techniques

whereby an optical probe that specifically localizes on an intended target reports on the

location and concentration of this target. In contrast, to study gene expression [6-9], gene

regulation [10] and protein-protein interactions [11,12] often an indirect imaging strategy is

used. The most common practice is the introduction of a transgene in the cell. The transgene

encodes for a fluorescent or bioluminescent protein, which acts as an intrinsically produced

reporter probe. Transcription and translation of the gene leads to the production of the

fluorescent or bioluminescent protein, which can then be detected with optical imaging

methods. Therefore, gene expression and regulation is imaged indirectly by visualizing and

quantifying the presence of fluorescent or enzymatic activity of the bioluminescent protein in

tissues. Reporter genes allow also for sensitive, quantitative, real-time spatiotemporal analysis

of the dynamics of neoplastic cell growth and metastasis and their response to therapeutic

intervention [13-16].

36

4.2. Origin of light

In optics, the term light refers to electromagnetic radiation with wavelengths of 300 nm (ultra

violet) through 1400 nm (infrared). Light is composed of elementary particles called photons

and can exhibit properties of both waves and particles. Which description fits best depends on

the interaction and technology used. There are many different light sources, but the

underlying process producing the light is the same. An electron absorbs energy and moves to

a higher orbit. When the electron falls back down to a lower energy state, a packet of energy

(i.e. photon) is released and light emission is observed. Electrons may be excited in a number

of different ways, e.g. by exothermic (bio)chemical reactions (chemi- and bioluminescence)

or by absorption of light (fluorescence). Below fluorescence and bioluminescence are

explained in more detail.

4.2.1. Fluorescence

Fluorescence is the property of some atoms and molecules (known as fluorophores) to absorb

light at a particular wavelength and to subsequently emit light of longer wavelength after a

brief interval, this is illustrated by the simple electronic-state diagram (Jablonski diagram)

shown in Figure 4.1. The fluorescence process is governed by three important events: (1)

excitation, (2) internal conversion and vibrational relaxation and (3) emission. Because the

energy associated with fluorescence emission transitions is typically less than that of

absorption, the resulting emitted photons have less energy and are shifted to longer

wavelengths. This phenomenon is generally known as Stokes shift. The primary origin of the

Stokes shift is the rapid decay of excited electrons to the lowest vibrational energy level of the

first excited state by internal conversion and vibrational relaxation. The entire fluorescence

process described above is cyclical. Therefore, a single fluorophore can generate many

thousands of detectable photons, which is fundamental to the high sensitivity of fluorescence

detection techniques.

Fluorophores can be divided into two broad classes, termed endogenous and exogenous.

Endogenous fluorophores, such as amino acids, structural proteins, enzymes, vitamins and

lipids, are those that occur naturally. Examples of exogenous fluorophores are cyanine dyes,

photosensitizers and molecular markers such as green fluorescent protein (GFP). These are

synthetic dyes or modified biochemicals that are added to a specimen to produce fluorescence

with specific spectral properties. When choosing the best fluorophore for a certain application

several parameters are used to describe and compare different fluorophores: extinction

37

coefficient, quantum yield, excitation and emission wavelength and photostability. The molar

extinction coefficient (ranging from 5000 to 200,000 cm-1 mol-1) is a direct measure of a dye's

ability to absorb light and at the same time determines the amount of light a molecule can

generate via fluorescence emission. The fluorescence quantum yield is a measure of the

efficiency with which the excited molecule is able to convert absorbed light to emitted light. It

is defined as the fraction of absorbed photons that are converted to fluorescence emission.

Typical quantum yields for commonly used fluorophores range from 0.05 to 1.0. For in vivo

imaging of fluorophores it is useful to have a probe that has its excitation and emission

wavelengths in the range of 600 to 900 nm (this is explained in more detail in 4.3).

Photostability is the ability of a dye to undergo repeated cycles of excitation and emission

without undergoing chemical modifications while being in the excited state. Chemical

modification of the excited state dye, referred to as "photobleaching," is an important factor

that limits fluorescence detection under high-intensity illumination.

Figure 4.1 Jablonski diagram illustrating the processes responsible for the fluorescence of

fluorophores. The absorption of a photon of energy hνEX by a fluorophore creates an excited

electronic singlet state, S1’ (event 1). Internal conversion and vibrational relaxation cause the

partial dissipation of the energy in S1’, yielding a relaxed singlet excited state S1 (event 2).

Finally, the fluorophore returns to its ground state (S0) by emitting a photon of energy hνEM

(event 3).

4.2.2. Bioluminescence

Bioluminescence is the production and emission of light as a result of a biochemical reaction

during which chemical energy is converted into light energy. In nature, bioluminescence is

found in many species, particularly in the marine environment, and each species has an

38

independently evolved biochemical system for light production [17]. However, for the in vivo

use of bioluminescence as biomarker the firefly luciferase system is by far the most

commonly used [18,19]. The firefly (Photinus pyralis) luciferase is an enzyme and catalyzes

the oxidation of the substrate luciferin in the presence of the cofactors adenosine triphosphate

(ATP), Mg2+ and O2. The oxidation of luciferin forms oxyluciferin in its electronically excited

state. Oxyluciferin returns to ground state while emitting broadband light in the green to

yellow region (500-700 nm), with a peak emission wavelength at 560 nm (Figure 4.2) [19].

The firefly luciferase system has a couple of properties that make it very useful as reporter

gene, such as its high sensitivity and tight coupling of protein synthesis with enzyme activity.

Furthermore, luciferase is a monomer that does not require any post-translational

modifications; it is available as a mature enzyme directly upon translation from its mRNA.

Hence, the luciferase assay provides a nearly instantaneous measure of total reporter

expression in the cell [20], if not limited by deficient co-factors. The use of firefly luciferase

requires exogenous administration of luciferin (enzyme substrate) to the animal before

imaging. Contag et al. showed that luciferin diffuses within minutes throughout all tissues

after i.v. and i.p. administration and rapidly enters into cells due to its small size and its

zwitter ionic nature [21]. Therefore, the exogenous administration of luciferin is not a limiting

factor for bioluminescence imaging. Notice that the measured luciferase activity changes with

the time after luciferin injection, as well as with dose. Therefore, experiments have to be

performed under constant conditions in order to make a quantitative comparison between the

results possible.

39

Figure 4.2 Firefly (Photinus pyralis, (b)) luciferase catalyzes the light-producing reaction of

luciferin to oxy-luciferin in the presence of adenosine triphosphate (ATP), Mg2+ and O2 (c).

The emitting light is in the range of 500-700 nm, with a peak emission wavelength at 560 nm

(a).

4.2.3. Fluorescence vs. bioluminescence

In the field of in vivo reporter gene technologies, bioluminescence (BLI) has several

advantages over fluorescence (FL). Foremost is the inherent low background of BLI markers

as compared with FL reporters, leading to higher sensitivity. FL suffers from autofluorescence

of tissue, resulting from the use of high power external excitation light source, causing an

increased background signal. Caceres et al. showed a 5-25-fold greater sensitivity in vitro for

luciferase transfected cells than for green fluorescent protein (GFP) transfected cells [22]. The

influence of autofluorescence can be reduced by working in the red to near infrared (NIR)

wavelength range [23], as is shown in Figure 4.3. Additionally, the need to use an external

excitation light source in FL imaging may reduce even further the low signal-to-noise ratio by

photobleaching. Hypothetically BLI signals are more easily quantified than FL signals,

because the signal level observed at a given depth of the animal is directly proportional to the

number of cells. In contrast, in the case of FL, the signal is related to both the number of cells

and the intensity of excitation light, which is difficult to quantify. However, in practice

quantification for both is difficult due to complicating factors such as scattering of light. FL

40

has also a couple advantages over BLI. First of all, it is much brighter and therefore poses

fewer restrictions on the imaging system (see also 4.4). Furthermore, FL can also be used with

a large variety of exogenous contrast agents and it has already found applications in clinical

use [24]. Another interesting evolution in the field of fluorescence is the development of

fluorescence tomography. This technique allows for the 3-D reconstruction of the fluorophore

distribution and improves quantification accuracy [25].

Figure 4.3 Wavelength-dependent autofluorescence during in vivo fluorescence imaging.

Tissue autofluorescence was imaged immediately after sacrifice using three different

excitation/emission filter sets: (a) white light, no filters; (b) blue/green (460-500 nm/505-560

nm); (c) green/red (525-555 nm/590-650 nm); and (d) NIR (725-775 nm/790-830 nm).

Arrows mark the location of the gallbladder (GB), small intestine (SI) and bladder (BI).

Image from [23].

4.3. Interaction light-tissue

If light is sent into tissue, different processes can occur. Most light enters the tissue, but a

small part can be reflected off the tissue surface (depending on the angle of incidence and the

refractive index). Inside the tissue, the light can be absorbed or scattered [26-28].

Light absorption in tissue is strongly wavelength dependent as is illustrated in Figure 4.4,

since different absorbing chromophores, absorb in different wavelength regions. At

wavelengths in the ultraviolet and blue region, oxy- and deoxyhemoglobin, other proteins and

41

amino acids absorb strongly. When working in the red to near-infrared (NIR) wavelength

there is less absorption and thus a deeper penetration of the light into the tissue. However, at a

wavelength of 900 nm and above, water becomes a strong absorber. Thus the optimal optical

window for light penetration tissue is: 600-900 nm (Figure 4.4). The differences in

attenuation relative to wavelength have a couple of important implications when performing

in vivo optical imaging. First, when using reporter molecules that have a part of their emission

spectrum below 600 nm, such as firefly luciferase and GFP, their emitted light will be largely

absorbed by tissue. Figure 4.5 shows that only 30% of the emission spectrum of firefly

luciferase is above 600 nm, which is probably the region of the emission spectrum detected

when used as an in vivo reporter. Notice that in in vivo FL applications not only the emitted

light is absorbed, but also the excitation light. Secondly, the emission spectrum of light

depends on the depth of the light source in the tissue. Therefore, light emitted at the surface of

an animal will contain more of the blue spectrum than light originating from deep within the

animal. Both phenomena complicate quantification of the emitted light in fluorescence and

bioluminescence imaging.

Whereas absorption depends on chromophores in the tissue, scattering is caused by tissue

structures in the order of magnitude of the light wavelength, such as elastin and collagen (i.d.

Rayleigh scattering) and mitochondria, cell nucleus, Golgi apparatus (i.d. Mie scattering).

Scattering decreases monotonically with increasing wavelength. However the ratio of

scattering to absorption coefficient increases with the wavelength. Thus, working in the

optical window (600-900 nm), a significant amount of light can escape the tissue, but the

emitted light is highly diffuse. This makes precise localization and quantification of the

emitted light difficult.

42

Figure 4.4 Interaction of light with tissue. The absorption coefficient of light in tissue is

dependent on wavelength and results from absorbers such as hemoglobins, lipids and water.

From [29].

4.4. Light measurement

The collection and measurement of the emitted photons is done by an image sensor.

Photomultiplier tubes are mostly used in the case of confocal microscopy. Charged coupled

device (CCD) cameras are the most commonly used sensors. CCDs are silicon-based

integrated circuits consisting of a dense matrix of photodiodes that operate by converting light

into defined electric charges. Electrons generated by the interaction of photons with silicon

atoms are stored in a potential well and can subsequently be transferred across the chip

through registers and output to an amplifier for storage as an electric image on a computer.

Because CCDs are used in very low light circumstances (in case of BLI), a very high

efficiency and sensitivity is needed. The quantum efficiency of a CCD, which is the ratio

between the number of photons absorbed and the number of electrons created, can reach

values up to 90% in the visual spectrum. The sensitivity of CCDs is optimized by reducing all

noise sources, such as dark current and read-out noise.

43

Figure 4.5 Wavelength dependent emission spectrum during in vivo bioluminescence imaging.

Bioluminescence images were taken of the same mouse with luciferase transfected tumor cells

implanted on its back with different emission filters: (a) no filter; (b) band pass filter (535-

580 nm); and (c) long pass filter (from 610) nm. 60% of the detected light is emitted in the

region at wavelengths longer than 610 nm. (From J.L. Coll, )

44

4.5. References

[1] Becker A, Hessenius C, Licha K, Ebert B, Sukowski U, Semmler W, Wiedenmann B,

Grotzinger C. Receptor-targeted optical imaging of tumors with near-infrared

fluorescent ligands. Nat Biotechnol 2001;19(4):327-331.

[2] Weissleder R, Tung CH, Mahmood U, Bogdanov A, Jr. In vivo imaging of tumors

with protease-activated near-infrared fluorescent probes. Nat Biotechnol

1999;17(4):375-378.

[3] Bremer C, Tung CH, Weissleder R. In vivo molecular target assessment of matrix

metalloproteinase inhibition. Nat Med 2001;7(6):743-748.

[4] Petrovsky A, Schellenberger E, Josephson L, Weissleder R, Bogdanov A, Jr. Near-

infrared fluorescent imaging of tumor apoptosis. Cancer Res 2003;63(8):1936-1942.

[5] Michalet X, Pinaud FF, Bentolila LA, Tsay JM, Doose S, Li JJ, Sundaresan G, Wu

AM, Gambhir SS, Weiss S. Quantum dots for live cells, in vivo imaging, and

diagnostics. Science 2005;307(5709):538-544.

[6] Bhaumik S, Gambhir SS. Optical imaging of Renilla luciferase reporter gene

expression in living mice. Proc Natl Acad Sci U S A 2002;99(1):377-382.

[7] Wu JC, Sundaresan G, Iyer M, Gambhir SS. Noninvasive optical imaging of firefly

luciferase reporter gene expression in skeletal muscles of living mice. Mol Ther

2001;4(4):297-306.

[8] Yang M, Baranov E, Moossa AR, Penman S, Hoffman RM. Visualizing gene

expression by whole-body fluorescence imaging. Proc Natl Acad Sci U S A

2000;97(22):12278-12282.

[9] van Roessel P, Brand AH. Imaging into the future: visualizing gene expression and

protein interactions with fluorescent proteins. Nat Cell Biol 2002;4(1):E15-20.

[10] Ichikawa T, Hogemann D, Saeki Y, Tyminski E, Terada K, Weissleder R, Chiocca

EA, Basilion JP. MRI of transgene expression: correlation to therapeutic gene

expression. Neoplasia 2002;4(6):523-530.

[11] Massoud TF, Paulmurugan R, De A, Ray P, Gambhir SS. Reporter gene imaging of

protein-protein interactions in living subjects. Curr Opin Biotechnol 2007;18(1):31-37.

[12] Ray P, Pimenta H, Paulmurugan R, Berger F, Phelps ME, Iyer M, Gambhir SS.

Noninvasive quantitative imaging of protein-protein interactions in living subjects.

Proc Natl Acad Sci U S A 2002;99(5):3105-3110.

45

[13] Edinger M, Sweeney TJ, Tucker AA, Olomu AB, Negrin RS, Contag CH.

Noninvasive assessment of tumor cell proliferation in animal models. Neoplasia

1999;1(4):303-310.

[14] Sweeney TJ, Mailander V, Tucker AA, Olomu AB, Zhang W, Cao Y, Negrin RS,

Contag CH. Visualizing the kinetics of tumor-cell clearance in living animals. Proc

Natl Acad Sci U S A 1999;96(21):12044-12049.

[15] Chishima T, Yang M, Miyagi Y, Li L, Tan Y, Baranov E, Shimada H, Moossa AR,

Penman S, Hoffman RM. Governing step of metastasis visualized in vitro. Proc Natl

Acad Sci U S A 1997;94(21):11573-11576.

[16] Yang M, Baranov E, Wang JW, Jiang P, Wang X, Sun FX, Bouvet M, Moossa AR,

Penman S, Hoffman RM. Direct external imaging of nascent cancer, tumor

progression, angiogenesis, and metastasis on internal organs in the fluorescent

orthotopic model. Proc Natl Acad Sci U S A 2002;99(6):3824-3829.

[17] Hastings JW. Chemistries and colors of bioluminescent reactions: a review. Gene

1996;173(1 Spec No):5-11.

[18] de Wet JR, Wood KV, DeLuca M, Helinski DR, Subramani S. Firefly luciferase gene:

structure and expression in mammalian cells. Mol Cell Biol 1987;7(2):725-737.

[19] de Wet JR, Wood KV, Helinski DR, DeLuca M. Cloning of firefly luciferase cDNA

and the expression of active luciferase in Escherichia coli. Proc Natl Acad Sci U S A

1985;82(23):7870-7873.

[20] Wood KV. The chemistry of bioluminescent reporter assays. Promega notes 1998:14-

21.

[21] Contag CH, Spilman SD, Contag PR, Oshiro M, Eames B, Dennery P, Stevenson DK,

Benaron DA. Visualizing gene expression in living mammals using a bioluminescent

reporter. Photochem Photobiol 1997;66(4):523-531.

[22] Caceres G, Zhu XY, Jiao JA, Zankina R, Aller A, Andreotti P. Imaging of luciferase

and GFP-transfected human tumours in nude mice. Luminescence 2003;18(4):218-

223.

[23] Frangioni JV. In vivo near-infrared fluorescence imaging. Curr Opin Chem Biol

2003;7(5):626-634.

[24] Ito S, Muguruma N, Kimura T, Yano H, Imoto Y, Okamoto K, Kaji M, Sano S, Nagao

Y. Principle and clinical usefulness of the infrared fluorescence endoscopy. J Med

Invest 2006;53(1-2):1-8.

46

[25] Ntziachristos V. Fluorescence molecular imaging. Annu Rev Biomed Eng 2006;8:1-

33.

[26] Cheong WF, Prahl SA, Welch AJ. A review of the optical properties of biological

tissues. IEEE J Quantum Electron 1990;26:2166-2185.

[27] Rice BW, Cable MD, Nelson MB. In vivo imaging of light-emitting probes. J Biomed

Opt 2001;6(4):432-440.

[28] Tuchin V. Tissue optics: light scattering methods and instruments for medical

diagnosis: International society for optical engineering; 2000.

[29] Shah K, Weissleder R. Molecular optical imaging: applications leading to the

development of present day therapeutics. NeuroRx 2005;2(2):215-225.

47

Part II. Spatio-temporal control of gene

activation

48

49

Chapter 5. Regulatable gene expression systems

5.1. Introduction

The objective of gene therapy is to express a therapeutic gene in the region where therapy is

required and for the duration necessary to achieve a therapeutic effect and to minimize

systemic toxicity. Although a lot of progress has been made in developing vectors for targeted

delivery of genes, it looks quite unlikely that delivery vectors emerge that are specific for a

certain tissue, organs or regions in need of therapy (except perhaps those based on cells with

“natural” homing capabilities such as immune cells and stem cells). Improvements in gene

delivery and spatio-temporal control of gene expression are important requirements in order

to introduce gene therapy in the clinical environment. Several approaches for controlling gene

expression have been proposed in literature [1]. Tissue-specific or disease-specific promoters

can provide spatial control of gene expression [2,3], whereas small molecule-dependent gene

switches, such as tetracycline[4,5], mifepristone [6,7] and rapamycin [8,9], can give temporal

control. Alternatively, physical stimuli such as ionizing radiation [10] and heat [11] offer the

possibility to activate gene expression in deep tissue with excellent spatial definition and

allow temporal control of the start of gene activation. The time course of activation and

subsequent deactivation of the promoters, which respond to these physical stimuli, may be

viewed as a basal form of temporal control. However, this form of temporal control is an

intrinsic property of the promoter and can not be manipulated readily to generate a desired

temporal expression profile that would result in an optimal therapy. Deliberate control of both

spatial and temporal regulation may be obtained via the use of two- or three component

systems (Figure 5.1) comprising (i) a small molecule dependent transactivator whose

expression is placed under the dual control of an inducible promoter and a transactivator-

responsive promoter and (ii) a transactivator-responsive promoter to which a transgene of

interest is linked [12]. With these multi-component systems not only spatial and temporal

control of gene expression is achieved, but also unintentional activation can be prevented.

50

Figure 5.1 Heat-activated and small molecule ligand-dependent three component gene switch

comprising two copies of a transactivator gene, of which one is controlled by an hsp promoter

and the other by a transactivator-responsive promoter (trp), and a transactivator-responsive

promoter to which a transgene of interest is linked. The arrow pointing to the right indicates

transactivator turnover. Adapted from [1].

As ionizing radiation may be a limiting factor when repeated activation is required, heat

appears to be the more suitable approach for controlling local gene expression. The objective

of this part of the thesis is to demonstrate the possibility of spatial and temporal control of

transgene expression using MR guided HIFU in combination with a temperature sensitive Hsp

promoter. Heat shock proteins (Hsps) are part of a family of proteins, whose synthesis is

elevated in response to stress. Combining the Hsp promoter with a reporter gene (e.g.

51

luciferase or green fluorescent protein) allows deliberate activation and kinetic follow-up of

the reporter gene. A general introduction in the biology of heat shock proteins and

applications of the Hsp promoter in gene therapy is described in Chapter 6. The promoter’s

activity was assessed with respect to temperature and duration of hyperthermia in vitro as well

as in vivo, which is presented in Chapter 7 and Chapter 8, respectively. Finally, in Chapter 9

the in vivo local activation of a transgene under control of Hsp promoter using MRgHIFU is

presented.

5.2. References

[1] Vilaboa N, Voellmy R. Regulatable gene expression systems for gene therapy. Curr

Gene Ther 2006;6(4):421-438.

[2] Gorski K, Carneiro M, Schibler U. Tissue-specific in vitro transcription from the

mouse albumin promoter. Cell 1986;47(5):767-776.

[3] Melo LG, Gnecchi M, Pachori AS, Kong D, Wang K, Liu X, Pratt RE, Dzau VJ.

Endothelium-targeted gene and cell-based therapies for cardiovascular disease.

Arterioscler Thromb Vasc Biol 2004;24(10):1761-1774.

[4] Gossen M, Bujard H. Tight control of gene expression in mammalian cells by

tetracycline-responsive promoters. Proc Natl Acad Sci U S A 1992;89(12):5547-5551.

[5] Rendahl KG, Leff SE, Otten GR, Spratt SK, Bohl D, Van Roey M, Donahue BA,

Cohen LK, Mandel RJ, Danos O, Snyder RO. Regulation of gene expression in vivo

following transduction by two separate rAAV vectors. Nat Biotechnol

1998;16(8):757-761.

[6] Wang Y, DeMayo FJ, Tsai SY, O'Malley BW. Ligand-inducible and liver-specific

target gene expression in transgenic mice. Nat Biotechnol 1997;15(3):239-243.

[7] Wang Y, O'Malley BW, Jr., Tsai SY, O'Malley BW. A regulatory system for use in

gene transfer. Proc Natl Acad Sci U S A 1994;91(17):8180-8184.

[8] Rivera VM, Clackson T, Natesan S, Pollock R, Amara JF, Keenan T, Magari SR,

Phillips T, Courage NL, Cerasoli F, Jr., Holt DA, Gilman M. A humanized system for

pharmacologic control of gene expression. Nat Med 1996;2(9):1028-1032.

[9] Ye X, Rivera VM, Zoltick P, Cerasoli F, Jr., Schnell MA, Gao G, Hughes JV, Gilman

M, Wilson JM. Regulated delivery of therapeutic proteins after in vivo somatic cell

gene transfer. Science 1999;283(5398):88-91.

52

[10] Hallahan DE, Mauceri HJ, Seung LP, Dunphy EJ, Wayne JD, Hanna NN, Toledano A,

Hellman S, Kufe DW, Weichselbaum RR. Spatial and temporal control of gene

therapy using ionizing radiation. Nat Med 1995;1(8):786-791.

[11] Madio DP, van Gelderen P, DesPres D, Olson AW, de Zwart JA, Fawcett TW,

Holbrook NJ, Mandel M, Moonen CT. On the feasibility of MRI-guided focused

ultrasound for local induction of gene expression. J Magn Reson Imaging

1998;8(1):101-104.

[12] Vilaboa N, Fenna M, Munson J, Roberts SM, Voellmy R. Novel gene switches for

targeted and timed expression of proteins of interest. Mol Ther 2005;12(2):290-298.

53

Chapter 6. Heat shock proteins

6.1. Introduction

Heat shock proteins (Hsp) are present in both prokaryotic and eukaryotic cells. Their highly

conserved primary structure (60-78% similarities in eukaryotes) suggests that they play a

crucial role in cellular processes [1]. Hsps are constitutively expressed in cells under normal

conditions for which they function as molecular chaperons [2]. They play a critical role in

normal protein homeostasis to assist in protein folding [3,4], the assembly and disassembly of

protein complexes [5], inhibition of improper protein aggregation [6,7] and to direct newly

formed proteins to target organelles for final packaging, degradation or repair. In response to

stress some forms of Hsp can be up regulated and they assist in refolding and repair of

denaturized proteins as well as facilitating synthesis of new proteins to repair damage [8]. The

induction of Hsp production is initiated by stressful conditions such as hyperthermia [9],

ischemia [10], hypoxia [11], depletion of ATP [12], free radicals [13] and various viruses.

The stress response is evoked primarily in response to the presence of damaged molecules

[14]. The proposed mechanism of stress-induced increase in Hsps is illustrated in Figure 6.1.

Under normal conditions heat shock factors (HSF) are bound to Hsps and are inactive. Under

stress conditions, such as heat shock, HSFs are separated from the Hsps. Protein kinase or

other serine/threonine kinases phosphorylate the HSFs, which cause them to form trimers in

the cytosol [15]. The trimers enter the nucleus and bind to the heat-shock elements (HSE)

located on the promoter region of the Hsp genes, and become further phosphorylated by HSF

kinases. Hsp mRNA is transcribed, transported from the nucleus to the cytoplasm and

translated into Hsp proteins. The newly synthesized Hsps bind to HSFs to prevent further

synthesis of Hsps. The above described pathway and associated feedback control mechanism

for the expression and regulation of Hsps is just a general description of this in vivo process.

In the context of this thesis, a lot of details are left out for clarity or are still unknown such as

how the basal level of Hsps is maintained in non-stress conditions or how the rate of Hsp

synthesis is upgraded in proportion to the nature and magnitude of the stress loading.

54

Figure 6.1 Proposed mechanism of stress-induced increase in Hsps. HSFs residing in the

cytosol are normally bound by Hsp and are inactive. Under stress, such as heat shock, HSFs

separate from Hsp, are phosphorylated by protein kinases such as PKC, and form trimers in

cytosol that enter the nucleus to bind HSEs in the promoter region of Hsp gene. HSF is

phosphorylated further, and Hsp mRNA is transcribed and leaves the nucleus for cytosol. In

cytosol, new Hsp is synthesized. HSF returns to the cytosol and is bound once again by HSF.

From [1].

6.2. Hsp promoters in gene therapy

Hsp promoters, particularly Hsp70 promoters, have been quite often used for gene therapy

strategies because they are both heat-inducible and efficient. Notice that not all Hsp70

promoters are inducible, though in the remainder of this thesis we will use Hsp70

promoter/protein to indicate the inducible form of Hsp70 promoter/protein if not stated

otherwise. Further, Hsp70’s almost universally presence in cells from bacteria to humans

makes it possible that virtually any Hsp70 promoter from any eukaryotic organism may be

used in any eukaryotic host cell [16,17]. This property is quite convenient for gene therapy

because strategies can be achieved with Hsp promoters from different species and tested in

different cellular and animal models. Although, the Hsp70 promoter looks very promising for

gene therapy, we have to keep in mind that some limitations do exist such as uncontrolled

55

activation of Hsp70 promoter, thermotolerance and the need for temporal and spatial control

of activation. As mentioned before, the Hsp70 promoter can be stimulated by a variety of

stresses of both environmental and physiological origins. The construction of a minimal

Hsp70 promoter, containing only three elements, solved largely the problem of uncontrolled

activation, because this promoter is believed to respond almost exclusively to heat [18,19].

Additional control could be obtained by rendering the Hsp promoter activation dependent on

an exogenous agent. The phenomenon of thermotolerance is discussed in more detail in

Chapter 8 and the temporal and spatial control of gene activation is the main subject of

Chapter 9.

6.3. References

[1] Kiang JG, Tsokos GC. Heat shock protein 70 kDa: molecular biology, biochemistry,

and physiology. Pharmacol Ther 1998;80(2):183-201.

[2] Freeman BC, Michels A, Song J, Kampinga HH, Morimoto RI. Analysis of molecular

chaperone activities using in vitro and in vivo approaches. Methods Mol Biol

2000;99:393-419.

[3] Saibil H. Molecular chaperones: containers and surfaces for folding, stabilising or

unfolding proteins. Curr Opin Struct Biol 2000;10(2):251-258.

[4] Zimmerman SB, Minton AP. Macromolecular crowding: biochemical, biophysical,

and physiological consequences. Annu Rev Biophys Biomol Struct 1993;22:27-65.

[5] Schroder M, Kaufman RJ. The mammalian unfolded protein response. Annu Rev

Biochem 2005;74:739-789.

[6] Ellis RJ. Molecular chaperones: avoiding the crowd. Curr Biol 1997;7(9):R531-533.

[7] Ohtsuka K, Hata M. Molecular chaperone function of mammalian Hsp70 and Hsp40--

a review. Int J Hyperthermia 2000;16(3):231-245.

[8] Martin J, Horwich AL, Hartl FU. Prevention of protein denaturation under heat stress

by the chaperonin Hsp60. Science 1992;258(5084):995-998.

[9] Ostberg JR, Kaplan KC, Repasky EA. Induction of stress proteins in a panel of mouse

tissues by fever-range whole body hyperthermia. Int J Hyperthermia 2002;18(6):552-

562.

[10] Richard V, Kaeffer N, Thuillez C. Delayed protection of the ischemic heart--from

pathophysiology to therapeutic applications. Fundam Clin Pharmacol 1996;10(5):409-

415.

56

[11] Patel B, Khaliq A, Jarvis-Evans J, Boulton M, Arrol S, Mackness M, McLeod D.

Hypoxia induces HSP 70 gene expression in human hepatoma (HEP G2) cells.

Biochem Mol Biol Int 1995;36(4):907-912.

[12] Kabakov AE, Gabai VL. Heat-shock proteins maintain the viability of ATP-deprived

cells: what is the mechanism? Trends Cell Biol 1994;4(6):193-196.

[13] Kukreja RC, Kontos MC, Loesser KE, Batra SK, Qian YZ, Gbur CJ, Jr., Naseem SA,

Jesse RL, Hess ML. Oxidant stress increases heat shock protein 70 mRNA in isolated

perfused rat heart. Am J Physiol 1994;267(6 Pt 2):H2213-2219.

[14] Kultz D. Molecular and evolutionary basis of the cellular stress response. Annu Rev

Physiol 2005;67:225-257.

[15] Kroeger PE, Sarge KD, Morimoto RI. Mouse heat shock transcription factors 1 and 2

prefer a trimeric binding site but interact differently with the HSP70 heat shock

element. Mol Cell Biol 1993;13(6):3370-3383.

[16] Pelham HR. A regulatory upstream promoter element in the Drosophila hsp 70 heat-

shock gene. Cell 1982;30(2):517-528.

[17] Voellmy R, Rungger D. Transcription of a Drosophila heat shock gene is heat-induced

in Xenopus oocytes. Proc Natl Acad Sci U S A 1982;79(6):1776-1780.

[18] Leung TK, Rajendran MY, Monfries C, Hall C, Lim L. The human heat-shock protein

family. Expression of a novel heat-inducible HSP70 (HSP70B') and isolation of its

cDNA and genomic DNA. Biochem J 1990;267(1):125-132.

[19] Smith RC, Machluf M, Bromley P, Atala A, Walsh K. Spatial and temporal control of

transgene expression through ultrasound-mediated induction of the heat shock protein

70B promoter in vivo. Hum Gene Ther 2002;13(6):697-706.

57

Chapter 7. In vitro characterization of Hsp70

promoter

7.1. Introduction

From the general introduction of Hsp proteins in Chapter 6 it follows that the promoter region

of the Hsp70 gene functions as a temperature dependent switch for turning on the production

of the Hsp70 protein. This temperature sensitive gene activation is an interesting

characteristic for controlling therapeutic gene expression with local heating. The first results

of heat-activated gene therapy for treating human cancers have already been published. In

these studies a herpes simplex virus thymidine kinase (HSV-tk) was placed under control of

Hsp promoter for treating gastric cancer [1] and breast cancer [2,3] with a combined therapy

of hyperthermia and ganciclovir. The thymidine kinase phosphorylates the nontoxic prodrug

ganciclovir, which then becomes phosphorylated by endogenous kinases to ganciclovir-

triphosphate, causing DNA-damage related apoptosis.

There are a couple of characteristics that make the Hsp70 promoter very suitable for gene

therapy in cancer treatment as described above, but also for treating genetic disorders such as

cystic fibrosis [4]. Several in vitro studies showed that the inducible human Hsp70 promoter

(Hsp70B) has a low basal activity and a high heat-induced expression amplification [5,6]. In

addition, its magnitude of induction depends on temperature as well as duration of the

induced hyperthermia [7,8]. Similar Hsp70 promoter characteristics were reported in vivo in a

number of organs, such as skin [8], muscle [9], prostate [10] and liver [11] that were

genetically modified using either a virus or plasmid to express a reporter gene under the

control of the human Hsp70 promoter.

In general, the local concentration of therapeutic gene determines the efficacy of the therapy.

When using therapeutic genes, such as HSV-tk, under control of a Hsp promoter for gene

therapy, their local concentration depends on the activity of the Hsp promoter and the efficacy

of gene delivery. The objective of this chapter is to assess the promoter’s activity with respect

to the temperature and duration of hyperthermia. The characterization of the exogenous

mouse Hsp70 promoter’s activity with respect to temperature and duration of hyperthermia

was performed on homogeneous bone marrow cell suspensions obtained from femoral bone

of NLF-1 mice. NLF-1 mice contain a transgene that allows firefly luciferase expression

58

under control of the heat shock protein 70 promoter (Hspa1b). Beckham et al. showed that the

bioluminescent light emission (i.e. luciferase activity) is a reliable method for quantifying

Hsp70 promoter’s activity by correlating the results of an ELISA Hsp70 protein concentration

determination and photon counts after heating [12].

In clinical applications of gene therapy, using hyperthermia in combination with a heat

sensitive promoter, normalization of time-temperature data is very important for obtaining

reproducible results and efficient therapeutic effects since both depend on temporal (duration

of heat up and cool down phase) and spatial (tissue inhomogeneities) variations of

temperature. A nowadays general accepted method for normalizing time-at-temperature data

was introduced by Sapareto and Dewey in 1984 [13]. They proposed a thermodynamic

approach that leads to calculating a thermal iso-effect dose (TID), which is the heating

duration at some reference temperature, e.g. 43º C, required for an observed biological effect

induced by a specified time at a specified temperature. Their method was based on Arrhenius

analyses of cell killing during exposure to heat at different temperatures and for different

exposure times [14,15]. The Arrhenius integral describes the measure of thermal damage as

function of temperature and time of exposure:

( ) τdeAtC

Ct

t

o

RT

Ea

∫−

=

=Ω 0ln)( 7-1

where Ω is the tissue damage, defined as the natural logarithm of a ratio of the concentration

of native (undamaged) tissue before heating (C0) to the concentration of native tissue after

heating (C(t)), A is the frequency factor (s-1), Ea is the activation energy (J·mol-1), T is the

temperature of exposure (K) and R is the universal gas constant (R = 8,32 J·mol-1·K-1).

The relationship between thermal damage and temperature is usually illustrated by an

Arrhenius plot in which the logarithm of the thermal damage is plotted as function of the

inverse absolute temperature [16]. The activation energy, Ea, is determined from the slope of

the Arrhenius plot. Several studies obtained a straight line over the temperature range from

43,5 to 57º C with an activation energy of about 140 kcal·mol-1 [17,18], indicating that for an

iso-effect at various temperatures, a decrease of 1º C requires an increase of the hyperthermic

exposure time by a factor K. Dewey et al. showed that this factor K can be calculated from the

activation energy expressed in joules per mol as follows [19]. Iso-effect at different

temperature implies:

21 Ω=Ω 7-2

59

with Ω1 tissue damage at temperature T and with Ω2 tissue damage at temperature T-1. Using

the Arrhenius integral from equation 7-1 and assuming constant temperature this leads to:

( ) KteAteA TR

E

RT

E aa

⋅⋅⋅=⋅⋅ −−−

1 7-3

With K as factor that indicates the required increase in exposure time to obtain iso-effect

when the temperature is lowered with 1º C. Equation 7-3 can be re-written as follows:

)1()ln(

−+−=

TR

E

RT

EK aa 7-4

Leading to the following expression for K:

)1()ln(

−=

TRT

EK a 7-5

For an activation energy of 586 kJ/mol (≈ 140 kcal/mol) this leads to a factor K of 2. With this

Arrhenius method time-temperature data can be normalized by a simple equation [19]:

( )2112

TTKtt −×= 7-6

Where t1 is the time at temperature T1, t2 is the time at temperature T2 and K is a constant

value with a value around 2. When the temperature varies during a heating period, the time-

temperature relationship in equation 7-6 can be used to obtain an equivalent time at 43º C for

each time interval at a given temperature and the equivalent times at 43º C for all intervals

during the heating period can be summed to give the total equivalent time at 43º C for the

entire heating period. This leads to the thermal iso-effect dose (TID) as proposed by Sapareto

and Dewey in 1984:

Equivalent minutes at 43º C = EM43 = τdKt

Tt∫−

0

43 7-7

Sapareto and Dewey also found that K is about 2 for temperatures > 43º C and about 4-6 for

temperatures between 39 and 43º C. The change in slope of the Arrhenius plot (i.e. different

K-factors) is generally thought to be related to development of thermotolerance during heating

[20].

It should be noted that in most studies the Arrhenius formulation was used for analyzing

temperature induced, biophysical tissue damage (i.e. using direct measurements of protein

denaturation or necrosis). Therefore, Ω has been interpreted as a measure of thermal damage

with a value of Ω = 1 that is usually taken to coincide with the threshold of observable

damage. However, the objective of this study is to investigate if luciferase activity (i.e. Hsp70

promoter activity) also follows the Arrhenius relationship without causing thermal tissue

60

damage. Although several studies showed that Hsp70 promoter activity could be modulated

by changing the thermal dose, they never showed the exact relationship between promoter

activity and thermal dose. The hypothesis in this study is that Hsp promoter activity follows

an Arrhenius relationship. The performed Arrhenius analysis was similar to the analyses

based on tissue damage, though the maximum luciferase activity (i.e. maximum light

emission) observed after heating was chosen as the arbitrary endpoint corresponding to Ω = 1.

7.2. Materials & methods

Bone marrow cell extraction

The mouse Hsp70 promoter was characterized in vitro using bone marrow cells of NLF-1

mice. Animals were first anesthetized (2 % isoflurane in air) and then rapidly decapitated.

Bone marrow cells were flushed out of the femoral bone shafts using RPMI-1640 medium

(Invitrogen) supplemented with 2 % Fetal Calf Serum (FCS, Invitrogen). Bone marrow

mononuclear cells were isolated on ficoll by centrifugation (400 g, 20 min) and washed 3

times in RPMI-1640 + 10 % FCS (300 g, 5 min). Cells were counted and seeded (7.0·106

cells·ml-1) in RPMI (at ambient temperature) supplemented with 10 % FCS and mIL-3 (5

ng·ml-1). By using bone marrow cells from NLF-1 mice the similar promoter is used for in

vitro (this chapter) as well as in vivo (Chapter 8) characterization of the Hsp promoter’s

activity.

Heating

Immediately after cell extraction the cells were heated at different temperatures and for

different durations using a thermal cycler (T gradient, Biometra, Germany). Fifty µl of cell

suspension (3.5·105 cells) was loaded in thin-walled plastic tubes, assuring an even

distribution of the heat in the sample. The pre-programmed heating profile of the thermal

cycler consisted of an adaptation phase of 10 min at 37° C, a heating phase with different

temperature-time combinations and a recuperation phase of 10 min at 37° C. Finally the cell

samples were opened and incubated at 37° C in a 5 % CO2 incubator. Using a thermal cycler

machine for heating cell samples provides fast active heating (4° C/s) and cooling (3° C/s)

assuring a tight temporal control of the sample at elevated temperatures and allows testing of

several heating conditions at the same time. However, the use of a thermal cycler for heating

is only possible with cells in suspension and not with adherent cells.

61

Two series of heating experiments were carried out in this study. In the first series of

experiments, the influence of different heating protocols on the cell viability and the time

course of luciferase activity were investigated. Therefore, cell samples were heated for 16 min

at 43º C, 8 min at 44º C and 4 min at 45º C and analyzed for their cell viability and light

emission during a period of 7 hours with 1 hour intervals. These experiments were repeated

three times, each time with bone marrow cells originating from different NLF-1 mice. In the

second series of experiments, the luciferase activity with respect to temperature and duration

of hyperthermia was measured. For this reason light emission was measured 5 hours after

heating cell samples at six different durations (1, 2, 4, 8, 16 and 32 min) and at four different

temperatures (42, 43, 44 and 45° C, ).

Cell viability

Cell viability was measured using trypan blue exclusion dye. The reactivity of trypan blue is

based on the fact that the chromophore is negatively charged and does not interact with the

cell unless the membrane is damaged. Therefore, all the cells which exclude the dye are

viable. 25 µl of cell suspensions was gently mixed with 50 µl trypan blue and 25 µl PBS in an

appropriate tube. 25 µl of stained cell was placed in a hemocytometer and the number of

viable (unstained) cells was counted.

Luciferase light measurement

The luciferase activity was measured 5 hours post heating using the luciferase assay system

(Promega) and a luminometer apparatus (Berthold, Germany). Forty-eight µl of cell

suspension was lysed with 12 µl CCLR lysis buffer (5x). Fifty µl of luciferase assay reagent

was mixed with 5 µl cell lysate in luminometer tubes and mixed by flushing 2-3 times. The

tubes were placed in the luminometer and 10 s reading was started 10 s after mixing the

reagent with the cell lysate.

Arrhenius analysis

To investigate Hsp70 promoter activity as function of time and temperature, six different

durations (1, 2, 4, 8, 16 and 32 minutes) of thermal cycler experiments were performed, each

having four samples at different temperatures (42, 43, 44 and 45° C). To find the activation

energy (Ea) and frequency factor (A) values in equation 7-1 that best describe the

experimental values, a least square fit was performed. This least square fit allowed to find a

minimal difference between the experimental values of Ω and the theoretical values of Ω,

which were calculated by varying the activation energy between 0 and 5·106 J·mol-1 with a

62

step size of 1000 J·mol-1 and varying the natural logarithm of the frequency factor between

100 and 900 with a step size of 4.

A second, more classical method, performed to calculate the activation energy and frequency

rate, is based on the linearization of equation 7-1 by projecting the different exposure times on

one point, or in other words assuming a constant temperature exposure. This results into the

following linear Arrhenius relationship:

RT

EAt a+−=− )ln()ln()ln( ω 7-8

When ln(t)-ln(ω) is plotted versus 1/T, a linear fit can be applied to the data. The slope and y-

intercept of the linear fit were obtained to calculate Ea and A.

7.3. Results

Cell viability was determined, using trypan blue, at multiple time points after applying

different heating protocols (i.e. 16 min at 43º C, 8 min at 44º C and 4 min at 45º C). For all

three protocols cell viability remained constant during the first 5 hours after heating and

decreased significantly after 6 to 7 hours, as shown in Figure 7.1. The severe drop in cell

viability after 6 hours and the recovery of the cell viability one hour later for the case of 16

min heating at 43º C can not be explained by the author. In case of the 16 min heating at 43º C

the cell viability was above 80 % and for the other two cases the cell viability remained

between 70 and 80 % during the first 5 hours after heating.

63

0

20

40

60

80

100

120

1 2 3 4 5 6 7time after heating (h)

live

cells

(%

)16 min @ 43º C

8 min @ 44º C

4 min @ 45º C

Figure 7.1 Cell viability following 16 min heating at 43° C, 8 min heating at 44° C and 4 min

heating at 45° C measured every hour for a period of 7 hours after heating, with an interval

of 1 hour.

Figure 7.2 shows the relative light units (RLU) measured in a luminometer at different time

points after applying the heating protocols as described above. In Figure 7.2a the time course

of luciferase activity, expressed in RLU values, is shown. The peak in light emission is found

3-5 hours after heating. The light emission levels for the three different heating protocols are

comparable during the period of 7 hours. In Figure 7.2b the time course of luciferase activity,

expressed in RLU per viable cell, is shown. This figure also shows a transient luciferase

activity, with a maximum light expression measured 3-5 hours after heating, ignoring the

outer layer at 6 hours after heating for 16 min at 43º C. There is no significant difference in

light emission levels observed between the three different heating protocols.

64

Figure 7.2 Time course of luciferase activity from bone marrow cells from NLF-1 mice

submitted in vitro to 3 different heating protocols. Luciferase activity is expressed by RLU

measured with a luminometer (a) and normalized to the number of viable cells (b).

Figure 7.3 shows the heat-mediated induction of the Hsp promoter with respect to different

temperatures (42, 43, 44 and 45° C) and durations (1, 2, 4, 8, 16 and 32 minutes) of

hyperthermia indicated by luciferase activity. It is shown that expression of the reporter gene

can be modulated by the heating parameters. At 42° C, a weak expression is observed for

heating durations longer than 16 minutes only. At 43° C and 44° C, higher levels of

expression are observed with maximal intensities observed for 32 minutes and 16 minutes

heating, respectively. For these heating protocols a 52/53 fold increase in light emission was

measured compared to the non-heated control sample (i.e. 0 minutes heating). The levels of

emitted light remain comparable for these two temperatures upon reduction of relative

exposure times (i.e. 16 minutes at 43° C / 8 minutes at 44° C). At 45° C, the maximum light

intensity (obtained at 8 minutes) was lower (39 fold increase compared to control) than for

43° C and 44° C and decreased to zero with increasing exposure time. Light emission

decreased also to zero when increasing exposure time from 16 to 32 minutes at 44° C. The

decreased light intensities may be explained by the decreased viability at higher temperatures.

65

0

1000

2000

3000

4000

5000

6000

7000

0 1 2 4 8 16 32

heating duration (min)

RL

U42º C

43º C

44º C

45º C

Figure 7.3 Luciferase activity as an indicator of Hsp70 promoter activity in bone marrow

cells from NLF-1 mice, 5 hours after in vitro heating at different temperatures (42° C, 43° C,

44° C and 45° C) and for different durations (0, 1, 2, 4, 8, 16 and 32 min). Luciferase activity

was measured as emitted light using an in vitro enzymatic assay.

In order to investigate whether the luciferase activity (and thus the promoter activity) follows

an Arrhenius relationship the activation energy (Ea) and the frequency factor (A) described in

equation 7-1 were obtained by performing a least square fit between the experimental data

(i.e. Figure 7.3) and the theoretical data (χ2 = 6,0) . The resulting activation energy and

frequency factor are 630,000 J·mol-1 and 1.7·10104 s-1, respectively. An alternative analysis of

the data is shown in Figure 7.4. In this case the relationship between the luciferase activity

and temperature is illustrated by a ‘classical’ Arrhenius plot in which the logarithm of the

luciferase activity is plotted as function of the inverse absolute temperature [16]. The

activation energy (Ea) is determined from the slope of the Arrhenius plot and the frequency

factor (A) is determined from the y-intercept of the slope. The correlation coefficient of the

linear fit equaled 0.9986 indicating that experimental values followed an Arrhenius

relationship. The resulting activation energy and frequency factor of this method are 647,577

J·mol-1 and 2.3·10103 s-1, respectively. For both methods the experimental data points where

cell death reduces luciferase activity and therefore impedes quantitative analysis of Hsp70

promoter activity, are not taken into account (i.e. 32 min at 44° C and 16 and 32 min at 45°

C).

66

Figure 7.4 ‘Classical’ Arrhenius plot for constant-temperature in vitro water bath

experiments performed at different temperature (42, 43, 44 and 45° C). The error bars

represent the standard deviation of the averages for each temperature. The solid line is a

linear fit with error weighting performed in order to determine the activation energy and

frequency factor.

Equation 7-5 was used to calculate the factor K that indicates by which factor the

hyperthermic exposure time has to be increased to obtain an iso-effect at various

temperatures, when the temperature is decreased by 1° C. The resulting K factors for the least

square fit and the classical analysis are 2.11 and 2.16, respectively.

Figure 7.5 shows a simulation of iso-luciferase activity levels based on the values for

activation energy and frequency factor calculated from the least square fit. As explained in

materials & methods the combination of time and temperature corresponding to the highest

luciferase activity in Figure 7.3 (16 minutes at 44° C) coincides with the iso-level of 1.0 in

Figure 7.5.

67

Figure 7.5 Iso-activation plot calculated from the activation energy (630,000 J/mol) and

frequency factor (e240 s-1) values obtained with a least square fit of the experimental and

theoretical data. The activation was normalized to the experimental data point that resulted in

the highest luciferase activity, which is 960 seconds at 44° C.

7.4. Discussion

The in vitro characterization shows that the thermo-inducible Hsp70 mouse promoter

possesses the same features as the thermo-inducible Hsp70 human promoter already described

in the literature [5,7]. The Hsp70 promoter driving the luciferase expression has low basal

activity and can attain high heat-induced activity, up to 53 fold compared to basal activity

(Figure 7.3). Similar results of high heat-induced expression levels coupled with low levels of

basal expression were reported in other studies [3,21]. Furthermore, it was shown that the

magnitude of luciferase activity can be modulated by temperature and duration of

hyperthermia. At temperatures higher than 42° C the promoter activity has the tendency to be

dose-dependent. However, at prolonged exposure at 44° C (for 32 minutes) and 45° C (for 16

and 32 minutes) luciferase activity decreased. This may be due to the increased cell death at

prolonged exposure at 44° C and 45° C, which was also observed in cell viability assays

performed (see Figure 7.1). An alternative explanation for the decreased light emission may

be the denaturation by heat of proteins and organelles such as ribosomes and other

68

intracellular structures, necessary for the production of the luciferase protein. Also the

malfunction of mitochondria causing a lack of ATP and impairing the cell to carry out the

light producing reaction may cause reduced light emission levels.

The induction of the luciferase gene is transient, as shown in Figure 7.2, with maximal

expression levels 3-5 hours after heat shock. Previous studies of other groups reported similar

post-heating times for maximal luciferase activity [22,23]. An analysis of luciferase activity

later than 7 hours after heat shock was not performed, because the results would be biased by

significant cell death due to unfavorable survival conditions of the cells in the PCR tubes.

Therefore, it is unknown when the induced activity returns to basal levels. However, other

studies found comparable times for maximal expression, e.g. 4-8 hours, and reported a return

to background level after 48 hours [8]. The observed transient induction of luciferase is

consistent with the current model of Hsp protein production in response to stress and

discussed in Chapter 6. The endogenous Hsp proteins having repaired the denatured proteins

and being present in unbound form is causing a down regulation of Hsp protein production.

The magnitude and duration of heat-induced luciferase activity is comparable for the three

hyperthermia protocols. 16 minutes at 43° C, 8 minutes at 44° C and 4 minutes at 45° C result

all three in the same thermal iso-effect dose (TID) assuming that the K-factor in the Arrhenius

relationship equals 2, as was suggested by Sapareto et al. [13]. Several studies showed that

magnitude and duration of heat-induced luciferase activity is proportional to that of the

inducing stress (i.e. thermal dose), explaining the observed results [24,25].

In its classical approach the Arrhenius relationship predicts that the tissue damage (i.e. protein

denaturation) is linearly proportional to the time of exposure at a given constant temperature

and exponentially proportional to the temperature at a given time of exposure. In this study

we showed that Hsp70 promoter activation also follows the Arrhenius relationship over the

temperature range of 42 to 45° C. We found, using two different analysis methods, that the

activation energy values are comparable with respective values of 630,000 J·mol-1 and

647,577 J·mol-1. These activation energy values correspond with K factors of 2.11 and 2.16

respectively. Thus, for obtaining similar expression levels of therapeutic transgene at various

temperatures, a decrease of 1° C requires 2.1 fold increase in hyperthermic exposure time.

Different values for the activation energy were found by Beckham et al. (1.74·106 J·mol-1)

[12], Rylander et al. (2.4·105 J·mol-1) [26] and Diaz et al. (4.5·105 J·mol-1) [27]. This variation

may be explained by different biological endpoints chosen as ‘damage’ indicator. This study

69

and Beckham et al. used Hsp70 promoter activity, reflected by light emission, as arbitrary

endpoint. In contrast, Rylander et al. and Diaz et al. used cell viability as arbitrary endpoint.

Furthermore, all four studies used different cell lines for their research.

The activation energy found in our study is very similar to the activation energy for thermal

denaturation of proteins, which is in the range of 586,000 J·mol-1 [18,19]. Therefore, we

hypothesize that protein denaturation is the underlying phenomenon for the activation of the

Hsp70 promoter as well as tissue damage. Upon a certain threshold of thermal stress Hsps are

produced to protect the tissue from thermal damage. Crossing this thermal stress threshold

results in tissue damage and Hsps become a predictor of tissue damage. Hsp production is a

precursor of tissue damage (depending on the level of stress) and therefore both (Hsp

production and tissue damage) have the same mechanism at their origin (i.e. protein

denaturation), which results in similar activation energies.

Concrete examples of therapy related applications of the Arrhenius analysis are demonstrated

with the help of the iso-activation plot in Figure 7.5. The iso-activation plot can be used to

compare the results of treatments performed with different temperatures and/or durations, but

also for treatment planning. An important observation for the planning phase is that it is

preferable to change exposure times instead of temperature for changing from one iso-level to

another, particularly at exposure times shorter than 300 seconds. This can be explained by the

distance between two iso-levels in temperature, which is higher at longer exposure times, in

combination with the fact that control of temperature is more difficult and less precise than

the control of exposure time, certainly in vivo.

7.5. Conclusions

In this study the Hsp70 mouse promoter activity was assessed with respect to temperature and

duration of hyperthermia in homogeneous bone marrow cell suspensions from NLF-1 mice,

containing the firefly luciferase gene under control of this heat sensitive promoter. It was

shown that the Hsp70 promoter has a low basal activity and a high heat induced activity.

After induction, luciferase activity is transient with a maximal level of expression 3-5 hours

after induction. We demonstrated that the level of transgene expression can be modulated by

changing the duration or temperature of heating. Furthermore, we found that the Hsp70

promoter activity follows an Arrhenius relationship. The modulatable and predictable nature

of the Hsp70 promoter activity makes it very suitable for controlling gene expression.

70

7.6. Reference

[1] Isomoto H, Ohtsuru A, Braiden V, Iwamatsu M, Miki F, Kawashita Y, Mizuta Y,

Kaneda Y, Kohno S, Yamashita S. Heat-directed suicide gene therapy mediated by

heat shock protein promoter for gastric cancer. Oncol Rep 2006;15(3):629-635.

[2] Brade AM, Szmitko P, Ngo D, Liu FF, Klamut HJ. Heat-directed suicide gene therapy

for breast cancer. Cancer Gene Ther 2003;10(4):294-301.

[3] Braiden V, Ohtsuru A, Kawashita Y, Miki F, Sawada T, Ito M, Cao Y, Kaneda Y,

Koji T, Yamashita S. Eradication of breast cancer xenografts by hyperthermic suicide

gene therapy under the control of the heat shock protein promoter. Hum Gene Ther

2000;11(18):2453-2463.

[4] Tate S, Elborn S. Progress towards gene therapy for cystic fibrosis. Expert Opin Drug

Deliv 2005;2(2):269-280.

[5] Gerner EW, Hersh EM, Pennington M, Tsang TC, Harris D, Vasanwala F, Brailey J.

Heat-inducible vectors for use in gene therapy. Int J Hyperthermia 2000;16(2):171-

181.

[6] Huang Q, Hu JK, Lohr F, Zhang L, Braun R, Lanzen J, Little JB, Dewhirst MW, Li

CY. Heat-induced gene expression as a novel targeted cancer gene therapy strategy.

Cancer Res 2000;60(13):3435-3439.

[7] Borrelli MJ, Schoenherr DM, Wong A, Bernock LJ, Corry PM. Heat-activated

transgene expression from adenovirus vectors infected into human prostate cancer

cells. Cancer Res 2001;61(3):1113-1121.

[8] Smith RC, Machluf M, Bromley P, Atala A, Walsh K. Spatial and temporal control of

transgene expression through ultrasound-mediated induction of the heat shock protein

70B promoter in vivo. Hum Gene Ther 2002;13(6):697-706.

[9] Xu L, Zhao Y, Zhang Q, Li Y, Xu Y. Regulation of transgene expression in muscles

by ultrasound-mediated hyperthermia. Gene Ther 2004;11(11):894-900.

[10] Silcox CE, Smith RC, King R, McDannold N, Bromley P, Walsh K, Hynynen K.

MRI-guided ultrasonic heating allows spatial control of exogenous luciferase in canine

prostate. Ultrasound Med Biol 2005;31(7):965-970.

[11] Plathow C, Lohr F, Divkovic G, Rademaker G, Farhan N, Peschke P, Zuna I, Debus J,

Claussen CD, Kauczor HU, Li CY, Jenne J, Huber P. Focal gene induction in the liver

of rats by a heat-inducible promoter using focused ultrasound hyperthermia:

preliminary results. Invest Radiol 2005;40(11):729-735.

71

[12] Beckham JT, Mackanos MA, Crooke C, Takahashi T, O'Connell-Rodwell C, Contag

CH, Jansen ED. Assessment of cellular response to thermal laser injury through

bioluminescence imaging of heat shock protein 70. Photochem Photobiol

2004;79(1):76-85.

[13] Sapareto SA, Dewey WC. Thermal dose determination in cancer therapy. Int J Radiat

Oncol Biol Phys 1984;10(6):787-800.

[14] Henriques FC. Studies of thermal injury V. The predictability and the significance of

thermally induced rate processes leading to irreversible epidermal injury. Archives of

Pathology 1947;43:489-502.

[15] Roizin-Towle L, Pirro JP. The response of human and rodent cells to hyperthermia. Int

J Radiat Oncol Biol Phys 1991;20(4):751-756.

[16] Westra A, Dewey WC. Variation in sensitivity to heat shock during the cell-cycle of

Chinese hamster cells in vitro. Int J Radiat Biol Relat Stud Phys Chem Med

1971;19(5):467-477.

[17] Borrelli MJ, Thompson LL, Cain CA, Dewey WC. Time-temperature analysis of cell

killing of BHK cells heated at temperatures in the range of 43.5 degrees C to 57.0

degrees C. Int J Radiat Oncol Biol Phys 1990;19(2):389-399.

[18] Dewey WC, Freeman ML, Raaphorst GP, Clark EP, Wong RSL, Highfield DP, Spiro

IJ, Tomasovic SP, Denman DL, Coss RA. Cell Biology of hyperthermia and radiation.

Withers RMaHR, editor. New York: Raven Press; 1980. 589–621 p.

[19] Dewey WC, Hopwood LE, Sapareto SA, Gerweck LE. Cellular responses to

combinations of hyperthermia and radiation. Radiology 1977;123(2):463-474.

[20] Dewhirst MW. Thermal dosimetry. Seegenschmiedt MH, Bolomey J-C, Fessenden P,

Vernon CC, Brady LW, Heilmann H-P, editors. Berlin: Springer; 1995. 123-136 p.

[21] Brade AM, Ngo D, Szmitko P, Li PX, Liu FF, Klamut HJ. Heat-directed gene

targeting of adenoviral vectors to tumor cells. Cancer Gene Ther 2000;7(12):1566-

1574.

[22] Arai Y, Kubo T, Kobayashi K, Ikeda T, Takahashi K, Takigawa M, Imanishi J,

Hirasawa Y. Control of delivered gene expression in chondrocytes using heat shock

protein 70B promoter. J Rheumatol 1999;26(8):1769-1774.

[23] Blackburn RV, Galoforo SS, Corry PM, Lee YJ. Adenoviral-mediated transfer of a

heat-inducible double suicide gene into prostate carcinoma cells. Cancer Res

1998;58(7):1358-1362.

72

[24] Landry J, Bernier D, Chretien P, Nicole LM, Tanguay RM, Marceau N. Synthesis and

degradation of heat shock proteins during development and decay of thermotolerance.

Cancer Res 1982;42(6):2457-2461.

[25] Li GC, Werb Z. Correlation between synthesis of heat shock proteins and

development of thermotolerance in Chinese hamster fibroblasts. Proc Natl Acad Sci U

S A 1982;79(10):3218-3222.

[26] Rylander MN, Diller KR, Wang S, Aggarwal SJ. Correlation of HSP70 expression and

cell viability following thermal stimulation of bovine aortic endothelial cells. J

Biomech Eng 2005;127(5):751-757.

[27] Diaz SH, Nelson JS, Wong BJ. Rate process analysis of thermal damage in cartilage.

Phys Med Biol 2003;48(1):19-29.

73

Chapter 8. In vivo characterization of the Hsp70

promoter

8.1. Introduction

In Chapter 7 the influence of different heat shock protocols on Hsp70 promoter activity was

investigated in vitro. It was concluded that the promoter activity was transient and could be

modulated by changing temperature and/or duration of the heat shock. It was also

demonstrated that in the range between 42 and 45° C the promoter activity followed an

Arrhenius relationship, making comparison between heating protocols with different time-

temperature history possible. Furthermore, it was shown that for iso-effect at various

temperatures, a decrease of 1° C requires the hyperthermic exposure to be increased with a

factor of about 2.

In this chapter a transgenic mouse model is used to assess the kinetics of the Hsp70 promoter

activity with respect to temperature and duration of hyperthermia in vivo. The in vivo mouse

model gives also the possibility to investigate the time course of the promoter activity after

multiple sequential heatings. This is a situation that may resemble to a clinical situation,

where multiple sequential heatings might be necessary, due to the transient character of the

promoter, in order to obtain a prolonged treatment. Therefore, we investigated the effects of a

second and third heat shock at different time points after the first heat shock. In this case a

phenomenon called acquired thermotolerance may play an important role in manipulating the

magnitude of transgene activity.

Acquired thermotolerance is defined as the ability of a cell or organism to become resistant to

heat stress after a prior sub-lethal heat exposure [1-3]. The phenomenon of acquired

thermotolerance is transient in nature and depends primarily on the severity of the initial heat

stress. In general, the greater the initial heat dose, the greater the magnitude and duration of

thermotolerance. Although, there are several observations that suggest that Hsps are directly

involved in thermotolerance [4-6], the precise mechanism is not well understood. For

example, there are studies that found a correlation between the kinetics of thermotolerance

induction and decay versus changes in Hsp induction and degradation, but no causal

relationship has been established [7,8]. It is also not the objective of this study to enlighten the

phenomenon of thermotolerance: we are only interested in the behavior of transgene

74

expression after multiple activations. However we are aware that thermotolerance can result

in ‘unexpected’ expression levels.

8.2. Material & methods

Animals

In a first preliminary study female (n = 9) homozygote transgenic mice (NLF-1) in the age of

23-25 weeks were used [9]. In the succeeding studies only male (n = 54) NLF-1 mice in the

age of 15-22 weeks were used, to exclude variations due to hormonal period. One week

before being included in the protocol, each individual mouse was tested regarding to their

phenotype. Bioluminescence images (BLI) were acquired to verify the presence of low light

emission at the position of the feet and tail after luciferin injection. In contrast to testing

method used in Chapter 9, the mice were not heated: only the basal luciferase activity was

measured. One day before performing the heating experiments the posterior part of the mouse

was shaved with a clipper and depilatory cream. The absence of animal hair assured optimal

contact between the water and the skin during heating and less absorption and scattering of

light during bioluminescence imaging.

Heating

Heating was done in a water bath with automatic temperature regulation, preventing

fluctuations in temperature during the experiment. Water temperature was measured with a

calibrated thermometer (Luxtron, California, USA). Before and during heating, the animals

were sedated with isoflurane (2 % in air). For applying the heat shock, their left leg was put in

the water while the rest of their body was lying on isolation material to prevent overheating of

the mouse.

For the preliminary study 9 female mice were heated in groups of 3 at 43° C for duration of 4,

8 and 16 minutes, respectively. The same heating protocol was repeated on the same mice one

week later.

To investigate more extensively the influence of temperature and duration of hyperthermia on

the luciferase activity, 48 male mice were heated, in groups of 4, at different constant

temperatures (43, 44, 45 and 46° C) for three different exposure times (ranging between 30

seconds and 32 minutes, depending on the chosen temperature). This resulted in 12 different

combinations of temperature and exposure time. The same heating protocol was repeated on

the same mice one and two weeks later for three different heating protocols.

75

Control studies were performed on 6 mice. In a first series of experiments mice (n = 3) were

heated 16 minutes at 36.5° C during 3 weeks with 1-week interval. In a second series of

experiments mice (n = 3) were heated a first time 8 minutes at 36.5° C. The second and third

heating were performed at 43° C (8 minutes) 1 and 2 weeks after the first heating,

respectively.

All heatings were performed on the left hind leg of the mice and were followed by kinetic

measurement of luciferase activity with BLI as described below.

Bioluminescence image acquisition and analysis

To establish the time course of luciferase activity after applying different heating protocols,

bioluminescence imaging (BLI) was continued at two hours intervals for 10 hours.

Bioluminescence images of mice were acquired using a NightOWL LB 981 system equipped

with a NC 100 CCD camera (Berthold Technologies, Germany). Mice were injected intra-

peritoneally with D-luciferin (Promega, 2.9 mg in 100µL sterile phosphate buffer). Two

minutes later, mice were sedated with isoflurane (2 % in air) and bioluminescence images (2

minutes integration period, 2×2 binning) were taken 7 and 10 minutes after the luciferin

injection in prone and supine positions, respectively. A low light imaging standard (Glowell,

LUX biotechnology, UK) was placed next to the animal during each image acquisition to

provide a constant reference for the resulting images.

Analysis of the BLI was done manually by placing a small region of interest (ROI) at the

level of the knee of the heated leg, see Figure 8.1. This region was chosen because it does not

suffer from hyper-intense light spots due to small skin wounds. Within this ROI, the mean

light intensity (in photons s-1 mm-2) was measured.

Arrhenius An Arrhenius analysis was performed to assess the in vivo promoter activity with respect to

different temperatures (43, 44, 45 and 46° C) and different durations (varying between 30

seconds and 16 minutes) of hyperthermia. The light emission values were measured at

different time points (2, 4, 6 and 8 hours) after heating. As described in Chapter 7 a least

square fit was performed between the experimental and theoretical values to obtain the

activation energy (Ea) and frequency factor (A) that describe best the relationship between

different heating protocols. For a detailed description of the performed least square fit the

reader is referred to section 7.1. Also the second method, based on the linearization of

76

Arrhenius equation, was performed to calculate the activation energy and frequency rate. For

both analyses the data of all 12 different heating protocols were used.

8.3. Results

Figure 8.1 shows an example of a bioluminescence image taken four hours after heating the

left leg of a mouse in a water bath. There is only significant light coming from the heated leg.

The position of the ROI, used for quantification of the bioluminescence signal, is also

indicated. When the same ROI is placed on the knee of the non-heated leg it gives the basal

level of light emission, which is around 8000 photons s-1 mm-2.

Figure 8.1 Bioluminescence image taken 4 hours after heating the left leg of NLF-1 mouse in

a water bath. The colors show the light intensity measured with an optical camera in the

range of 200 photons per second (dark blue) to 4000 photons per second (red). The position

of the ROI used for quantification of the bioluminescence signal is also indicated.

Figure 8.2 shows the in vivo time course of luciferase activity after applying a heat shock at

different temperatures (43° C (a), 44° C (b), 45° C (c) and 46° C (d)) and for different

exposure times. The exposure times were changed according to the temperature to prevent

77

tissue damage: the higher the temperature, the shorter the exposure time. The luciferase

activity is transient for each combination of temperature and exposure time and the maximal

light emission was always found 4 hours after heating. For all temperatures tested, an increase

in luciferase activity of about 2 folds is observed when the exposure time is double at each

constant temperature.

Figure 8.2 Kinetic study of luciferase activity during 10 hours after heating a mouse leg with

different heating protocols. (a) 4, 8 and 16 minutes at 43° C, (b) 2, 4 and 8 minutes at 44° C,

(c) 1, 2 and 4 minutes at 45° C and (d) 30 seconds, 1 and 2 minutes at 46° C.

To demonstrate the influence of 1° C temperature increase at constant time of exposure the

data in Figure 8.2 are re-arranged and shown in Figure 8.3. Here the time courses of luciferase

activity after 4 minutes heating at 43, 44 and 45° C are plotted. An increase of 1° C results is

an increase of luciferase activity of about 2. This increase in activity was similar to the

increase of activity after doubling the time of heat exposure, as we saw in Figure 8.2.

78

0,0E+00

2,0E+05

4,0E+05

6,0E+05

8,0E+05

2 4 6 8 10

time after heating (h)

SI (

ph/s

/mm

2)

43° C

44° C

45° C

Figure 8.3 Kinetic study of luciferase activity during 10 hours after heating. For heating the

mouse leg was put 4 minutes in a water bath with a temperature of 43, 44 and 45° C,

respectively.

The measurements of luciferase activity at several time points after heating for different

temperatures and durations of hyperthermia were used for Arrhenius analysis in time. In

Figure 8.4 the logarithm of the luciferase activity 6 hours after heating is plotted as function

of the inverse absolute temperature. The data point corresponding to 46° C has a large

standard deviation and does not coincide with the linear tendency observed for the other three

data points. This may be explained by the short periods of heating (30 seconds, 1 and 2

minutes) at 46° C. These short heating periods may not be sufficient to obtain a homogenous

distribution of the elevated temperature in the leg, which causes a larger uncertainty in the

measured values. Therefore the data points measured at 46° C were not used for the Arrhenius

analysis. The solid line is a linear fit with error weighting performed on three data points,

corresponding to 43, 44 and 45° C. The correlation coefficient of the linear fit equaled 0.9441

indicating that experimental values followed an Arrhenius relationship.The activation energy

(Ea) was determined from the slope of the linear fit and the K-factor is calculated with help of

equation 7-5. The activation energy and K-factor were also calculated by performing a least

square fit between experimental data (without the data corresponding to 46° C) and

79

theoretical data . Figure 8.5 shows the evolution in time of the K-factor for the two different

analysis methods. The mean K-factor equals 1.9 ± 8% for both methods.

Figure 8.4 ‘Classical’ Arrhenius plot for constant-temperature in vivo water bath experiments

6h after heating. The error bars represent the standard deviation of the averages for each

temperature. The solid line is a linear fit with error weighting performed on three data points,

corresponding to 43, 44 and 45° C.

To simulate a clinical protocol where multiple sequential heatings are required for obtaining a

prolonged therapy, multiple similar heatings were performed at a week interval. In a

preliminary study the effect of two repeated heatings was measured for 3 different heating

protocols at the same temperature but with different thermal dose (i.e. 4, 8 and 16 minutes at

43° C). The time course of luciferase activity was measured during 10 hours after the first and

second heating for each protocol (Figure 8.6). The maximal luciferase activity was found 4

hours after heating, independent of the heating protocol and the number of heatings performed

before. However, the light intensity was consistently higher after the second heating,

independent of the heating protocol. The mean increase in signal intensity during 10 hours

was 2.1, 2.6 and 4.1 for 4, 8 and 16 minutes, respectively. The increase in luciferase activity

80

after the second heating appears to be larger when increasing the thermal dose of the heating

protocol.

0,00

0,50

1,00

1,50

2,00

2,50

0 2 4 6 8 10 12

time after heating (h)

K-f

act

or

Least square fit

'classic method'

Figure 8.5 Time course of K-factor for different methods of analysis. The K-factor was

calculated with help of a least square fit using either all data points or only the data points

corresponding to 43, 44 and 45° C. The K-factor was also calculated for both data sets with

linear fit method as demonstrated in figure 4.

Next, the effect of heating three times with an interval of one week was investigated for 3

different heating protocols with equivalent thermal dose (assuming a K-factor of 2) but

different temperature (i.e. 8 minutes at 43° C, 4 minutes at 44° C and 2 minutes at 45° C,

figure 5). The time course of luciferase activity was followed during 10 hours after each

heating and is shown in Figure 8.7. As observed before, the maximal luciferase activity was

found 4 hours after heating, independent of the number of heatings already performed before.

A mean increase of 4.8, 3.3 and 2.3 was found after the second heating for 8 minutes at 43° C,

4 minutes at 44° C and 2 minutes at 45° C, respectively. Apparently there is a smaller increase

in luciferase activity after the second heating at higher temperatures, but equal thermal dose.

In contrast, after the third heating a significant decrease in luciferase activity is observed

81

compared to the second heating. The luciferase activity returns to the same level as after the

first heating.

In the first series of control experiments no change in promoter activity was observed after

multiple heatings at 36,5° C during 16 minutes. The mean light emission levels 4 hours after

heating were 7709, 9521 and 6646 photons·s-1·mm-2 after the first, second and third heating,

respectively. These emission levels are similar to the emission levels of the non-heated leg

(6691 photons·s-1·mm-2).

In the second series of control experiments no increase in light emission was observed in the

heated leg (8 minutes at 36.5° C) compared to the unheated leg. After a second heating,

performed during 8 minutes at 43° C, similar levels of light emission were reached as was

observed for the first heating of 8 minutes at 43° C in Figure 8.2. After a third heating, also

performed during 8 minutes at 43° C, an increase in light emission was observed compared to

the emission levels after the second heating. This observation corresponds with the increase in

light intensity after the second heating shown in Figure 8.7.

82

Figure 8.6 Time course of luciferase activity after a first and a second similar heating in the

same animal for 4 (a), 8 (b) or 16 minutes (c) at 43° C.

83

Figure 8.7 Time course of luciferase activity after a first, second and third similar heating in

the same animal. Three different heating protocols were investigated: (a) 8 minutes at 43° C,

(b) 4 minutes at 44° C and (c) 2 minutes at 45° C.

84

8.4. Discussion

There are already a substantial amount of studies that demonstrate that modulation of Hsp70

promoter activity is possible in vitro by changing temperature and duration of hyperthermia

[10,11]. However, an extensive study of in vivo modulation of Hsp70 promoter activity has

not yet been performed. Silcox et al. showed that the canine prostate that received the longest

hyperthermic treatment above 42° C also expressed the highest level of luciferase activity

[12]. However, the prostates were genetically modified using a virus to express luciferase

under control of Hsp70 promoter and the measurement of promoter activity was invasive.

Therefore, a heterogeneous distribution of the reporter gene may be expected and a kinetic

analysis of promoter activity in the same animal is impossible. In this study the in vivo

characterization of Hsp70 promoter was performed with a transgenic mouse model,

expressing the luciferase gene under control of Hsp70 promoter, and water bath heating. The

transgenic mouse model excludes variations in luciferase expression due to heterogeneous

gene delivery and allows multiple sequential measurements of luciferase activity in the same

mouse.

The light emission originating from the non-heated leg of the NLF-1 mouse is very low (8000

photons s-1 mm-2), implying a low basal activity of the Hsp70 promoter. After heating, the

luciferase activity can increase up to 75-fold compared to the basal activity. This increase in

luciferase activity is comparable with the result from the in vitro study (Chapter 7).

Furthermore, it was shown that the magnitude of luciferase activity can be modulated by

changing the temperature and duration of hyperthermia. An about 2-fold increase of luciferase

activity is observed after doubling heating duration at constant temperature and after

increasing temperature with 1° C at constant time of exposure. This factor 2 increase in

luciferase activity was confirmed by performing an Arrhenius analysis. Two different analysis

methods resulted in a K-factor of 1.9, which is comparable with the value found in the in vitro

study (i.e. 2.1). Notice that the Arrhenius analysis is based on the light emission measured

after different heating protocols. The level of light emission depends on the production (i.e.

promoter activity) and elimination of the luciferase protein. The transient character of the

light emission levels show that the relative contribution of both process change in time.

However, the K-factor remains almost constant in time.

After repetition of the same heating protocol at 1-week interval, considerably different

luciferase activity levels were observed. The second heating led to a significant increase

compared to the first heating, whereas after the third heating the luciferase activity returned to

85

a level comparable to that of the first heating. In addition, the increase in luciferase activity

after a second heating appeared to be thermal dose and temperature dependent. To the

author’s knowledge this is the first study that showed an increased luciferase activity after a

second heating. Further research is needed to verify that this phenomenon is not related to our

transgenic mouse model and to analyze the consequences for gene therapy. With the current

knowledge of the behavior of Hsp70 and its role in thermotolerance these observations are

difficult to explain. Studies with different transgenic models showed that Hsp has a protective

function when it is present at elevated concentrations. For example, it was shown that

transgenic mice with an elevated constitutive Hsp70 expression had a higher resistance to

cardiac ischemia [13,14]. Similar observations were made when knock out animals were used

to study the influence of different genes on the heat shock response [15,16]. In contrast, the

animals in our study did not show any luciferase activity just before the second and third

heating, indicating that the Hsp70 promoter activity had returned to basal level. Furthermore,

in literature it was shown that the life time of endogenous Hsp after heat shock is at maximum

3-5 days [7,17]. Therefore, it is very unlikely that the increased luciferase activity observed

after the second heating is due to the increased presence of endogenous Hsp that is produced

after the first heating. Future research should investigate whether the increased luciferase

activity is really due to increased promoter activity or due to a change in the reaction of

luciferine to oxy-luciferin, by looking at the mRNA levels of luciferase and endogenous Hsp

after the first and second heating. Theodorakis et al. suggested that thermotolerant cells,

containing large concentrations of Hsps, limit Hsp70 expression by transcriptional and

pretranslational mechanisms, perhaps to avoid the potential cytotoxic effect of these proteins

[18]. This mechanism could explain the lower luciferase activity observed after the third

heating in our study. Assuming that the increased luciferase activity after the second heating

is due to an increased promoter activity a high concentration of Hsps may be expected. The

complete degradation of this large amount of Hsps may exceed the 3-5 days as mentioned

before and therefore result in an increased level of Hsps after 1 week. In future work this

theory should be verified by measurements of endogenous Hsp levels.

8.5. Conclusion

In this study it was shown that the activity of Hsp70 promoter can be modulated in vivo by

changing temperature and/or duration of hyperthermia. Hsp70 promoter activity follows an

Arrhenius relationship making temperature-time normalization of different heating protocols

possible. Temperature-time normalization allows comparison of achievable therapeutic levels

86

after different heating protocols and could be used during treatment planning. When multiple

sequential heatings are necessary for obtaining the desired therapeutic effect, treatment

planning may become more complicated. The until now unexplained increased luciferase

activity after a second heating compared to the first heating makes estimation of final

therapeutic effect after multiple heatings difficult.

8.6. References

[1] Landry J, Chretien P. Relationship between hyperthermia-induced heat-shock proteins

and thermotolerance in Morris hepatoma cells. Can J Biochem Cell Biol

1983;61(6):428-437.

[2] Li GC, Werb Z. Correlation between synthesis of heat shock proteins and

development of thermotolerance in Chinese hamster fibroblasts. Proc Natl Acad Sci U

S A 1982;79(10):3218-3222.

[3] Mizzen LA, Welch WJ. Characterization of the thermotolerant cell. I. Effects on

protein synthesis activity and the regulation of heat-shock protein 70 expression. J

Cell Biol 1988;106(4):1105-1116.

[4] Feder ME, Cartano NV, Milos L, Krebs RA, Lindquist SL. Effect of engineering

Hsp70 copy number on Hsp70 expression and tolerance of ecologically relevant heat

shock in larvae and pupae of Drosophila melanogaster. J Exp Biol 1996;199(Pt

8):1837-1844.

[5] Johnston RN, Kucey BL. Competitive inhibition of hsp70 gene expression causes

thermosensitivity. Science 1988;242(4885):1551-1554.

[6] Riabowol KT, Mizzen LA, Welch WJ. Heat shock is lethal to fibroblasts

microinjected with antibodies against hsp70. Science 1988;242(4877):433-436.

[7] Landry J, Bernier D, Chretien P, Nicole LM, Tanguay RM, Marceau N. Synthesis and

degradation of heat shock proteins during development and decay of thermotolerance.

Cancer Res 1982;42(6):2457-2461.

[8] Li GC. Elevated levels of 70,000 dalton heat shock protein in transiently

thermotolerant Chinese hamster fibroblasts and in their stable heat resistant variants.

Int J Radiat Oncol Biol Phys 1985;11(1):165-177.

[9] Christians E, Campion E, Thompson EM, Renard JP. Expression of the HSP 70.1

gene, a landmark of early zygotic activity in the mouse embryo, is restricted to the

first burst of transcription. Development 1995;121(1):113-122.

87

[10] Borrelli MJ, Schoenherr DM, Wong A, Bernock LJ, Corry PM. Heat-activated

transgene expression from adenovirus vectors infected into human prostate cancer

cells. Cancer Res 2001;61(3):1113-1121.

[11] Smith RC, Machluf M, Bromley P, Atala A, Walsh K. Spatial and temporal control of

transgene expression through ultrasound-mediated induction of the heat shock protein

70B promoter in vivo. Hum Gene Ther 2002;13(6):697-706.

[12] Silcox CE, Smith RC, King R, McDannold N, Bromley P, Walsh K, Hynynen K.

MRI-guided ultrasonic heating allows spatial control of exogenous luciferase in canine

prostate. Ultrasound Med Biol 2005;31(7):965-970.

[13] Marber MS, Mestril R, Chi SH, Sayen MR, Yellon DM, Dillmann WH.

Overexpression of the rat inducible 70-kD heat stress protein in a transgenic mouse

increases the resistance of the heart to ischemic injury. J Clin Invest 1995;95(4):1446-

1456.

[14] Plumier JC, Ross BM, Currie RW, Angelidis CE, Kazlaris H, Kollias G, Pagoulatos

GN. Transgenic mice expressing the human heat shock protein 70 have improved

post-ischemic myocardial recovery. J Clin Invest 1995;95(4):1854-1860.

[15] Huang L, Mivechi NF, Moskophidis D. Insights into regulation and function of the

major stress-induced hsp70 molecular chaperone in vivo: analysis of mice with

targeted gene disruption of the hsp70.1 or hsp70.3 gene. Mol Cell Biol

2001;21(24):8575-8591.

[16] Xia W, Vilaboa N, Martin JL, Mestril R, Guo Y, Voellmy R. Modulation of tolerance

by mutant heat shock transcription factors. Cell Stress Chaperones 1999;4(1):8-18.

[17] Kregel KC. Heat shock proteins: modifying factors in physiological stress responses

and acquired thermotolerance. J Appl Physiol 2002;92(5):2177-2186.

[18] Theodorakis NG, Drujan D, De Maio A. Thermotolerant cells show an attenuated

expression of Hsp70 after heat shock. J Biol Chem 1999;274(17):12081-12086.

88

89

Chapter 9. Local gene activation using MR guided

HIFU

9.1. Introduction

The objective of gene therapy is to express a therapeutic gene in the region where therapy is

required (i.e. spatial control) and for the duration necessary (i.e. temporal control) to achieve a

therapeutic effect while minimizing systemic toxicity. Several approaches for controlling

gene expression have been presented in Chapter 5. The use of heat in combination with heat-

inducible promoters, such as the Heat Shock Protein (Hsp) promoters, for local activation of

gene expression is a versatile method and has already been successfully employed [1]. In vitro

experiments showed that the human Hsp70 promoter [2,3] as well as the murine Hsp70

promoter (Chapter 7) are suitable for controlling gene activation. Both promoters showed low

basal activity and high heat-induced expression amplification. In addition, their magnitude of

induction depends on temperature as well as on the duration of the induced hyperthermia

[4,5]. Similar promoter characteristics were reported in vivo in a number of organs, such as

skin [5], muscle [6], prostate [7] and liver [8] that were genetically modified using either a

virus or plasmid to express a reporter gene under the control of the Hsp70 promoter. Non-

invasive in vivo local heating, and thus local activation of gene expression, could be achieved

with High Intensity Focused Ultrasound (HIFU) [8-10], but a subtle compromise has to be

found in the heating strategy since excessive tissue heating can induce tissue damage and

necrosis [11]. Local temperature distribution can be monitored with Magnetic Resonance

temperature Imaging (MRI) to evaluate efficacy of heat-induced gene activation technique as

well as safety. Furthermore, recent developments in MRI guided HIFU techniques allow

automatic control of the HIFU energy deposition in order that the local tissue temperature at

the targeted location follows a predefined temperature evolution [12,13]. This fully automated

approach has been developed and tested in vivo for tumor models expressing a fluorescent

protein reporter gene under transcriptional control of Hsp70B promoter [9]. Unfortunately,

due to limitations of currently available fluorescence imaging capabilities in deep tissues, the

kinetics of gene expression is difficult to assess in vivo. In addition, tumor models show a

high variability in the level of gene expression due to the spatial heterogeneity of cancerous

tissue and physiology. As a consequence, a quantitative spatio-temporal correlation between

MR temperature maps and transgene expression as well as an assessment of promoter activity

90

with respect to the temperature and duration of hyperthermia could not be established.

Moreover, the presence of necrosis complicated the evaluation of the non-invasiveness of the

method.

To overcome these limitations, in the present work a healthy transgenic mouse strain was

chosen to compare spatio-temporal control of gene expression with different heating patterns

under reproducible conditions. This animal model, expressing the firefly luciferase reporter

gene under control of a Hsp promoter, allowed in vivo assessment of the spatial extent and

time course of gene expression via bioluminescence imaging. In addition, the potential

influence of the mechanical and thermal stress of HIFU on Hsp promoter activation was

compared. Finally, a systematic evaluation of the damage induced by HIFU hyperthermia in

the targeted tissue was performed to evaluate the safety of this gene activation strategy.

9.2. Material & methods

Animals

Male homozygote transgenic mice (NLF-1) in the age of 17-26 weeks were used in this study

[14]. NLF-1 mice contain a transgene that allows firefly luciferase expression under control of

the mouse heat shock protein 70 promoter (Hspa1b). Each individual mouse was tested with

regard to their phenotype by heating both front legs in a water bath at constant temperature

(43°C) during 5 minutes three weeks prior to the start of the protocol. Bioluminescence

images (BLI) were acquired to verify the presence of light emission in the heated area (see

below for details of the BLI protocol). The posterior part of the mouse was shaved with a

clipper and depilatory cream to improve ultrasound and light propagation prior to HIFU

studies. All procedures were performed according to protocols approved by the French law

and the rules of animal care of the institute.

HIFU device

Ultrasound experiments were performed with an in-house designed single channel focused

ultrasound transducer (built by Imasonic SA, Besancon, France) incorporated in the bed of the

1.5 Tesla Magnetic Resonance Imaging (MRI) system. The transducer had a focal length of

80 mm and external radius aperture of 60 mm. The sinusoidal signal (1.5 MHz) was generated

using a Yokogawa FG110 wave generator (Yokogawa, Tokyo, Japan) and amplified using a

Kalmus KMP 170F amplifier (Kalmus, Bothell, WA). The dimensions of the focal point were

91

1 × 1 × 4 mm3 (at -3 dB). The wave generator was remotely controlled via in-house written

software running under Windows.

Animal preparation for MRI guided HIFU heating

Animals were sedated with isoflurane (2% in air) and received muscle relaxants (100 µl

lidocaine (2.5 mg·ml-1) in 3 different positions intra-muscularly) in the left hind leg to prevent

unwanted muscle contraction during sonication. Injection of 200 µl of the local anesthetic

ropivacaine (0.5 mg·ml-1) was performed subcutaneously in 2 different positions and 500 µl

ketamine (1.67 mg·ml-1) intra peritoneal as an analgesic (t = 0). To maintain the animal body

temperature constant at 35° C, the animal was positioned over a heating plate. At t = 10-15

minutes, the animal was positioned prone into the MRI, above the focused ultrasound

transducer immersed in a thermo regulated (35° C) water bath with 0.1% (w/v) manganese. At

t = 30 minutes, the deposition of energy by ultrasound was started.

MRI thermometry during HIFU sonication

Online monitoring of temperature distribution was performed with MRI thermometry, using

the proton resonance frequency technique [15]. Dynamic MR temperature imaging was

performed on Philips Achieva 1.5 Tesla with a segmented EPI sequence (EPI-factor = 5, TE =

15 ms, TR = 34 ms, flip angle = 160, FOV = 64 × 56 mm2, matrix = 64 × 64, 3 slices),

resulting in continuous and volumetric acquisition of gradient echo images with a resolution

of 1 × 1 × 2 mm3 every 1.6 seconds. A 23 mm surface receiver coil was positioned above the

mouse hind leg to provide sufficient signal to noise ratio on MRI images. MRI temperature

maps were calculated using the phase information in the gradient echo images and displayed

online. Potential drift of the temperature (caused by drift in time of the phase of the MRI

signal not related to temperature) was compensated by subtracting the average temperature in

a reference Region Of Interest (ROI) selected within the muscle outside the heated area. Short

(~ 10 s), low power HIFU was performed to verify the position of the focal point. The

maximal tissue temperature increase resulting from this test did not exceed 4°C.

Heating protocols

HIFU heating experiments were performed with automatic feedback regulation of the tissue

temperature at the HIFU focal point based on dynamic analysis of the MRI temperature

images to adjust the output power of the HIFU system [16]. Three heating conditions were

investigated, all consisting of a ramp time of 1 minute followed by a plateau at 43°C of either

2, 5 or 8 minutes. The required temperature increase to reach 43° C was determined for each

92

animal based on measurement of the rectal temperature using a MRI compatible thermometer

(Luxtron, California, USA). The electric power was registered for each protocol. The

efficiency of the transducer in water was used to calculate acoustical power levels. The

equivalent amount of energy was then applied to a separate group of animals at constant

power (equal to the mean value found in the initial set of experiments with automatic

feedback) in a pulsed manner (20% duty cycle at 1 Hz, increasing the duration by a factor of

5) to investigate the influence of the mechanical stress induced by HIFU on the resulting gene

activation. These experiments were also performed under MR temperature imaging with

identical imaging parameters, but without automatic MRI-guided feedback controlling the

HIFU power.

Bioluminescence image acquisition and analysis

Bioluminescence images of mice were acquired using a NightOWL LB 981 system equipped

with a NC 100 CCD camera (Berthold Technologies, Germany). Mice were injected intra

peritoneally with D-luciferin (Promega, 2.9 mg in 100 µl sterile phosphate buffer). Two

minutes later, mice were sedated with isoflurane (2% in air) and bioluminescence images (2

minutes integration period, 2 × 2 binning) were taken 7 and 10 minutes after the luciferin

injection in prone and supine positions, respectively. A low light emitting standard (Glowell,

LUX biotechnology, UK) was placed next to the animal during each image acquisition to

provide a constant reference for the resulting images. Grey-scale body-surface reference

images were collected for superposition of BLI images on anatomical maps. Pseudocolor

luminescent images representing the spatial distribution of emitted photons were generated

using IDL programming language (ITT). BLI analysis was done semi-automatically by first

placing a small region of interest (ROI) in the region corresponding with the heated region.

Then, a region growing algorithm was used to extend this ROI automatically to find the

region that corresponded to 320 photons·s-1·mm-2. This threshold was chosen to exclude

signal intensity related to the background and basal activity and was determined by analyzing

the histogram of the number of pixels as function of the signal intensity for 10 heated animals.

The mean light intensity was then evaluated within the resulting ROI.

Histology

Mice were sacrificed 24 hours after heating by cranial dislocation. Muscle samples were

obtained from both muscles (gluteal and rectus femoris) by dissection. The non-heated muscle

samples served as negative control for muscle damage. Samples were fixed in 10% neutral-

93

buffered formalin, followed by paraffin embedding and micron-thick sectioning (3µm). Each

paraffin block, containing the whole muscle sample was totally sliced until exhausted and

each slice analyzed. Then hematoxylin and eosin (H & E) staining was performed for routine

qualitative and quantitative examination of tissue morphology of the entire sample.

Histological examination was done by a confirmed pathologist for skeletal muscle in a

blinded manner, with systematic attention to the following parameters: muscle fiber diameters

(degree and distribution of atrophy), alterations in muscle fibers (centralization of

subsarcolemmal nuclei, changes in contour), cell necrosis, inflammatory infiltrate and size of

damage, if any (measured with microscopic lens).

9.3. Results

Precision of automatic temperature control in vivo in the leg muscle of mouse

Figure 9.1a illustrates the precision of the MRI based temperature control at the HIFU focal

point during 2 minutes heating at 43º C in thigh muscle. The standard deviation of the MRI

measured temperature was 0.61 °C at normal body temperature (observed during repetitive

temperature mapping). The average difference between target profile at the plateau and

experimental data was 0.66 °C during heating demonstrating the precision of the automatic

feedback coupling to ensure the predefined temperature evolution. In this experimental set-up

it was not possible to determine the accurancy of the used method. Figure 9.1b displays a

typical MRI thermal map obtained during automatic temperature control. A color coded

image superimposed on a grey-level magnitude MRI image shows the temperature induced by

HIFU heating. The temperature increase is localized at the HIFU focal point and diffuses out

due to heat conduction.

94

Figure 9.1 Typical time course of the temperature evolution during a heating experiment (a),

with the target temperature (in black) and the measured temperature at the focal point (in

red). MRI temperature map (color coded) superimposed on an anatomical MRI image

(grayscale) of the mouse leg (b).

Comparative analysis of MR temperature mapping and bioluminescence imaging

Figure 9.2a and b show MRI temperature maps (5 minutes after the start of the experiment) of

two mice heated with the same protocol, i.e. 8 minutes heating at 43° C in the focal point.

Figure 9.2c and d show bioluminescence images of the same two mice, taken 6 hours after

heating and following i.p. injection of luciferin. Although both mice underwent the same

heating protocol the resulting temperature distribution in the leg clearly appears different,

whereas the temperature evolution in the focal point was identical. This spatial difference is

also observed in the distribution of emitted photons for both mice (Figure 9.2c and d). These

results demonstrate the close correspondence between spatial distribution of temperature and

expression. Since heat conduction is spatially heterogeneous, the temperature distribution can

be expected to be different from animal to animal despite an identical evolution of the

temperature in the focal point.

95

Figure 9.2 MRI temperature maps of two mice heated with the same protocol, i.e. 8 minutes

heating at 43° C at the focal point (a,b). The colors in image a and b show the local

temperature distribution in the mouse leg in the range of 38° C (blue) to 42° C and higher

(red). Bioluminescence images of the same two mice, taken 6 hours after applying the heating

protocol (c,d). The color in images c and d show the light intensity measured with an optical

CCD camera, in the range of 320 photons per second (purple) to 2000 photons per sec and

more (red).

Nature of gene activation stimulus

HIFU exposure results in thermal and non-thermal mechanical stress to the tissue. From

literature it is known that inducible Hsp70 promoters are predominantly sensitive to

temperature but may also be activated by other stimuli like hypoxia [17], ischemia [18] and

mechanical stress [6]. In order to use MRI thermometry as a reliable method to predict local

transgene activation, gene induction should not be affected by the mechanical part of

ultrasound. To evaluate the influence of mechanical stress, a set of experiments was

performed in a separate group of animals, applying the same amount of total acoustical

energy and the same pressure amplitude, but with varying duty cycle. A 20% HIFU duty

96

cycle and 0.7 W acoustical power (comparable to the mean power found in the heating

experiments) resulted in negligible temperature increase. Figure 9.3 compares

bioluminescence images obtained without HIFU application (a), after MRI-monitored pulsed

HIFU experiment with 20% duty cycle (b) and after MRI-controlled HIFU heating (43°C, 2

min) (c). As expected, local light emission was observed at the spot where the leg of the

animal was heated (c, white arrow). The average acoustical energy deposited for this

experiment was 116 J (n = 4). In contrast, no light was detected at the target location

following pulsed HIFU with 125 J energy deposition (b, white arrow). Low intensity light

emission appeared at random locations (~ 160 photons/s) in all cases, attributed to very low

basal activity of the Hsp70 promoter.

Figure 9.3 Bioluminescence images taken before heating (a), 6 hours after depositing 125 J of

energy with pulsed ultrasound (b) and 6 hours after heating at 43° C for 2 minutes (c). The

arrows indicate the location of HIFU application. The color coding is representing the light

intensity measured with the CCD camera in photons per second.

Kinetics of light intensity after heating in vivo

The results of the in vitro experiments (Chapter 7) suggest that the level of expression of the

transgene is transient and depends on the heating amplitude and duration. Using BLI, the

kinetics of light emission was followed for several heating conditions in the leg muscle.

Figure 9.4 shows the evolution of light intensity for 10 mice heated at 43° C for 2, 5 and 8

minutes, respectively. Light was observed in all cases from 4 hours up to 24 hours post

heating. Light intensity remained nearly constant between 4 and 8 hours, and decreased to

basal levels at 24 hours. Weak emission levels persisted at 24 hours for the 8-minute protocols

and for a single 5-minute experiment. Increasing heating time influenced the level of

expression, with little differences between 2 and 5 minutes heating experiments whereas

97

substantially higher (3 to 4 times) expression was detected for the 8-minute heating protocols.

As expected from in vitro experiments, the heat-induced light production in vivo was found to

be transient and its intensity could be modulated by the heating protocol.

Figure 9.4 Time course of gene expression in vivo upon hyperthermia. The leg muscle of

transgenic mice was heated at 43 ° C for 2 (n = 4), 5 (n = 4) or 8 minutes (n = 2),

respectively. The mean light intensity was reported in photons per second per mm2. Light

emitted by luciferase was measured at 4, 6, 8 and 24 hours after heating. Light emission

below the quantification threshold was marked as ND (not detectable) and measurement not

performed are marked as NA (not available).

Histological analysis of heated muscle

The tissue damage resulting from the different HIFU heating protocols was investigated by

histological analysis and is shown in Figure 9.5. No muscular damage was noticed in muscle

tissue with the pulsed HIFU protocol and following 2 minutes heating at 43° C. However, for

5 minutes protocols at the same temperature a modest variability in fiber diameters without

specific pattern and moderate interstitial edema was noted as well as a few necrotic fibers

with centralized nuclei. Few atrophic fibers and no inflammation were observed. The region

with abnormalities did not exceed 3 millimeters in diameter. When increasing the heating

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duration to 8 minutes, the muscular damage was more extensive and corresponded to a region

measuring up to 11 millimeters in diameter. In this case a marked variability was observed in

fiber diameters with clearly necrosed fibers and severe interstitial edema, without fascicular or

perifascicular pattern. Necrotic fibers were accompanied by inflammatory infiltrates

(lymphocytes and polynuclear cells).

Figure 9.5 Haematoxylin-Eosin stained histology sections of excised muscle tissue 24 hours

after heating at 43° C for 5 minutes (a and c) and 8 minutes (b and e). After 5 minutes heating

a modest variability in fiber diameters and interstitial edema was observed compared to the

un-heated muscle (d). The muscular damage was more extensive after 8 minutes heating,

including necrotic fibers and inflammatory infiltrates. Image a and b are taken with 2x

magnification, whereas image c, d and e are taken with 10x magnification.

9.4. Discussion

Spatial and temporal control of transgene expression is an important requirement for gene

therapy. In this chapter a high similarity is shown between the local temperature distribution

in vivo and the region emitting light,, using MRI guided HIFU in a transgenic mouse

expressing luciferase under the control of a heat sensitive promoter,. Control of transgene

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expression was achieved by automatic adjustment of the HIFU power based on continuous

MRI thermometry to force the temperature to follow a predefined temperature evolution. The

good spatial correspondence between increased temperature and gene expression was

demonstrated by comparing MRI temperature maps and bioluminescence images. This

observation is based on qualitative comparison between images of two different imaging

modalities obtained on living animals, and demonstrates the great potential of MRI guided

HIFU to predict and control the region of gene expression. Future research should investigate

the role of iso-temperature levels and thermal dose on the spatial distribution and level of

gene activation. This was not possible with data acquired in this study due to the low SNR of

gradient echo images.

The spatial positioning of the focal point of the focused ultrasound and the quantity of

deposited energy can be completely controlled by the operator. However, the resulting local

temperature distribution and thus the region of gene activation are not completely predictable

and may depend on several parameters such as tissue absorption of ultrasound, and heat

conduction. These parameters can be spatially heterogeneous and may be affected by

temperature increase itself. Thus, the resulting heated area may vary from animal to animal

when using an identical heating protocol as shown in Figure 9.2a and b. However,

temperature maps of each animal clearly correspond with light emitting regions.

From the in vivo results, shown in Figure 9.4, a dose-dependent promoter activity can be

observed, allowing modulation of the gene expression level possible by appropriate choices of

the heating conditions. Furthermore, it was found that the promoter has a low level of basal

expression since light intensity measured in animals not submitted to HIFU heating was very

weak (~ 160 photons/s). In contrast, the level of induced activity increased at least 10-fold,

even with mild activation protocols, i.e. heating for 2 minutes at 43° C (see Figure 9.3c). We

also demonstrated that the influence of mechanical stress of HIFU on gene activation was

negligible, in agreement with the results from Liu et al. [10].

The induction of the luciferase gene was transient with maximal protein activity occurring 6-8

hours after heating. The time course of activation and subsequent de-activation is an intrinsic

property of the promoter, as well as mRNA and protein processing and therefore does not

allow temporal switch-off control by external factors. Since expression returns to near basal

activity at 24 hours following heat-shock, repetitive heating at the same location may allow

for activation during prolonged periods. Alternatively, temporal regulation may be achieved

via the use of two- or three component systems comprising (i) a small molecule dependent

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transactivator whose expression is placed under the dual control of a Hsp70 promoter and a

transactivator-responsive promoter and (ii) a transactivator-responsive promoter to which a

transgene of interest is linked [19].

The time evolution of luciferase expression upon different activation protocols is comparable

to in vitro observations made by other groups [3,5] and our self (Chapter 7). However,

absolute quantification of the resulting light intensity as a marker of protein activity remains

difficult, since the bioluminescence imaging of the reporter gene used here was a 2-D method

(resulting in weighted projection) whereas multi-slice MRI thermometry was used for

guidance of HIFU. Therefore, light coming from deeper inside the muscle will be attenuated

more than light coming from tissue closer to the surface. Furthermore, light coming from

deeper locations will be more red shifted (section 4.3) and therefore, in the case of luciferase,

be less intense. Both phenomena cause a non-linear signal contribution of deeper tissue to the

measured signal with an optical camera. In preliminary studies, the skin of the heated leg was

removed under anesthesia to expose the underlying tissue. Compared to the case with the skin

still in its place we observed a six times lower light emission when it was removed. These

results could indicate a more important heating of the skin surface then of the muscle.

However, MR temperature maps indicate no heating close to the surface. Variations in

thermal sensitivity and the ability to induce Hsp70 after stress of different tissues are two

alternative explications for the observed phenomenon.

The non-invasiveness of the proposed approach was evaluated by a systematic histological

analysis of the heated muscles. For short duration hyperthermia (i.e. 2 minutes at 43°C), no

damage was observed but, as expected, increasing the duration of the hyperthermia resulted in

an increase of induced damage in the leg muscles with substantial alteration of the muscle

appearance after 8-minute hyperthermia. This illustrates the importance of the choice of the

heating conditions in vivo to observe sufficient induction of expression avoiding tissue

damage. This justifies the fundamental importance of a precise, non invasive and quantitative

temperature measurement and control system provided by the combination of MRI

thermometry and HIFU heating.

Although, the above presented results are very encouraging for using MR guided HIFU as

technique for controlling local gene activation, there are several practical limitations that

should receive attention and improvement before translating this technique into the clinic.

Part of the problems encountered is due the small dimensions of the mouse model and will

probably be solved by using a larger biological model (e.g. rat, rabbit, pig and human). The

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PRF based method used for MR thermometry is motion sensitive, making motion correction

unavoidable for obtaining precise temperature maps. To prevent inter-scan motion of the

mouse leg it was injected with muscle relaxant. The anesthesia did indeed prevent all inter-

scan motion, however it hampered sometimes the MR thermometry by creating black holes in

the images (i.e. pixels lacking signal). This was probably caused by the accumulation of

intramuscularly injected anesthesia or small hemorrhage. Another motion related problem is

caused by the radiation force related motion of less rigid tissue, such as fat, and the overlying

skin. Because the motion is instantaneous at the moment of HIFU activation it is impossible

to correct. In larger muscle structures no macroscopic skin motion will be observed due to the

increased muscle mass and tissue rigidity.

The small size of the mouse leg in combination with the relative large size of the focal point

poses problems at the interface of tissue and air. Due to the large difference in acoustic

impedance between tissue and air, ultrasound may be partially reflected (depending on the

angle of incidence) and thereby creating a second hot spot close to the skin, making skin

burns more likely. To prevent this from happening ultrasound gel, which has the about the

same acoustic impedance as tissue, was put on the animal’s leg, on the distal side of the US

transducer. This solution transfers the problem of a possible second hot spot outside the

animal’s leg.

Another problem related to the small dimensions of the mouse leg is the compensation of

magnetic field drift. Even without heating, the temperature in the mouse leg changes; better

said the phase change without heating. This is caused by magnetic field drift of the main

magnetic field and has to be compensated for reliable temperature maps. The most straight

forward method for compensating the magnetic field drift is subtracting the phase changes in

a region of interest with constant temperature from the phase changes in the heated region.

However, this method assumes that the magnetic field changes are similar in both regions in

absence of heating, which is a condition difficult to meet. Using the 23-mm receiver coil, in

order to obtain highest sensitivity, makes it impossible to image an unheated object (e.g. an

agar gel) placed next to the animal with the same high sensitivity. Furthermore, there is a

quick lost in field homogeneity distaly from the leg. Both phenomena make a similar change

in magnetic field drift in the heated leg and the gel very unlikely. Therefore, we have chosen

to place the ROI of the non-heated area also on the animal’s leg. In this case the ROI will be

very close to the heated region and the coil sensitivity and the field homogeneity will be

almost the same. However, there is now a possibility of low heating in the non-heated region.

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The correct positioning of the focal point before starting the heating procedure is a general

occuring problem during treatment planning. In this study we determined the position (i.e. the

pixel) of the focal point by performing a first heating in agar gel. The same pixel should

contain the focal point in a next heating when the position of the slices is unchanged.

However, in vivo the absorption, diffusion and perfusion of heat are very heterogeneous

compared to the agar gel and may result in a physical movement of the focal point. Therefore,

a first short heating in vivo for finding the position of the focal point is inevitable, with

possible Hsp70 promoter activation as a result. Even though, the pixel containing the focal

point is found with an in vivo test, the position of the focal point may change during the actual

heating, causing problems for automatic control of temperature during heating.

Another problem occurs when spin echo images or gradient echo images with different echo

times with respect to the thermometry sequence are used for anatomical verification of the

focal point position. In this case the pixel containing the focal point (found during the first

heating) will not correspond to the same anatomical position as seen in the anatomical SE and

EPI GE images. This is due to the compensation of chemical shift induced distortion in SE

images and different chemical shift induced distortions in the phase direction of EPI GE

images.

9.5. Conclusions

In this study it was demonstrated that there is a high spatial correspondence between the

heated region and the region where transgene product (i.e. light) is found. Further, it was

shown that Hsp70 promoter was mainly induced by heat and not by mechanical stress caused

by ultrasound. The level of promoter activity can be modulated by changing the activation

protocol. Therefore, the combination of MRI guided HIFU heating and transgenes under

control of heat inducible Hsp70 promoter provides a direct, non-invasive, spatial control of

gene expression via local hyperthermia.

9.6. References

[1] Vekris A, Maurange C, Moonen C, Mazurier F, De Verneuil H, Canioni P, Voisin P.

Control of transgene expression using local hyperthermia in combination with a heat-

sensitive promoter. J Gene Med 2000;2(2):89-96.

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[2] Gerner EW, Hersh EM, Pennington M, Tsang TC, Harris D, Vasanwala F, Brailey J.

Heat-inducible vectors for use in gene therapy. Int J Hyperthermia 2000;16(2):171-

181.

[3] Huang Q, Hu JK, Lohr F, Zhang L, Braun R, Lanzen J, Little JB, Dewhirst MW, Li

CY. Heat-induced gene expression as a novel targeted cancer gene therapy strategy.

Cancer Res 2000;60(13):3435-3439.

[4] Borrelli MJ, Schoenherr DM, Wong A, Bernock LJ, Corry PM. Heat-activated

transgene expression from adenovirus vectors infected into human prostate cancer

cells. Cancer Res 2001;61(3):1113-1121.

[5] Smith RC, Machluf M, Bromley P, Atala A, Walsh K. Spatial and temporal control of

transgene expression through ultrasound-mediated induction of the heat shock protein

70B promoter in vivo. Hum Gene Ther 2002;13(6):697-706.

[6] Xu L, Zhao Y, Zhang Q, Li Y, Xu Y. Regulation of transgene expression in muscles

by ultrasound-mediated hyperthermia. Gene Ther 2004;11(11):894-900.

[7] Silcox CE, Smith RC, King R, McDannold N, Bromley P, Walsh K, Hynynen K.

MRI-guided ultrasonic heating allows spatial control of exogenous luciferase in canine

prostate. Ultrasound Med Biol 2005;31(7):965-970.

[8] Plathow C, Lohr F, Divkovic G, Rademaker G, Farhan N, Peschke P, Zuna I, Debus J,

Claussen CD, Kauczor HU, Li CY, Jenne J, Huber P. Focal gene induction in the liver

of rats by a heat-inducible promoter using focused ultrasound hyperthermia:

preliminary results. Invest Radiol 2005;40(11):729-735.

[9] Guilhon E, Voisin P, de Zwart JA, Quesson B, Salomir R, Maurange C, Bouchaud V,

Smirnov P, de Verneuil H, Vekris A, Canioni P, Moonen CT. Spatial and temporal

control of transgene expression in vivo using a heat-sensitive promoter and MRI-

guided focused ultrasound. J Gene Med 2003;5(4):333-342.

[10] Liu Y, Kon T, Li C, Zhong P. High intensity focused ultrasound-induced gene

activation in solid tumors. J Acoust Soc Am 2006;120(1):492-501.

[11] Hundt W, Yuh EL, Steinbach S, Bednarski MD, Guccione S. Comparison of

continuous vs. pulsed focused ultrasound in treated muscle tissue as evaluated by

magnetic resonance imaging, histological analysis, and microarray analysis. Eur

Radiol 2008;18(5):993-1004.

[12] De Zwart JA, Salomir R, Vimeux F, Klaveness J, Moonen CTW. On the feasibility of

local drug delivery using thermo-sensitive liposomes and MR-guided focused

ultrasound. 2000; Denver. p 43.

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[13] Vimeux FC, De Zwart JA, Palussiere J, Fawaz R, Delalande C, Canioni P, Grenier N,

Moonen CT. Real-time control of focused ultrasound heating based on rapid MR

thermometry. Invest Radiol 1999;34(3):190-193.

[14] Christians E, Campion E, Thompson EM, Renard JP. Expression of the HSP 70.1

gene, a landmark of early zygotic activity in the mouse embryo, is restricted to the

first burst of transcription. Development 1995;121(1):113-122.

[15] Ishihara Y, Calderon A, Watanabe H, Okamoto K, Suzuki Y, Kuroda K, Suzuki Y. A

precise and fast temperature mapping using water proton chemical shift. Magn Reson

Med 1995;34(6):814-823.

[16] Salomir R, Vimeux FC, de Zwart JA, Grenier N, Moonen CT. Hyperthermia by MR-

guided focused ultrasound: accurate temperature control based on fast MRI and a

physical model of local energy deposition and heat conduction. Magn Reson Med

2000;43(3):342-347.

[17] Patel B, Khaliq A, Jarvis-Evans J, Boulton M, Arrol S, Mackness M, McLeod D.

Hypoxia induces HSP 70 gene expression in human hepatoma (HEP G2) cells.

Biochem Mol Biol Int 1995;36(4):907-912.

[18] Richard V, Kaeffer N, Thuillez C. Delayed protection of the ischemic heart--from

pathophysiology to therapeutic applications. Fundam Clin Pharmacol 1996;10(5):409-

415.

[19] Vilaboa N, Fenna M, Munson J, Roberts SM, Voellmy R. Novel gene switches for

targeted and timed expression of proteins of interest. Mol Ther 2005;12(2):290-298.

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Part III. Local drug delivery

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Chapter 10. The role of ultrasound and molecular

imaging in local drug delivery

10.1. Introduction

Oncology is probably the domain where innovations in local drug delivery can best improve

the efficacy of therapy and safety of the patients. Despite a large range of new therapeutic

innovations such as immuno and gene therapy and drugs with lower systemic toxicity,

ordinary chemotherapy remains an often used treatment.

For an anticancer drug to kill a high proportion of cancer cells in a solid tumor, it must

overcome several boundaries (Figure 10.1). Abnormalities in both tumor vasculature and

extracellular matrix lead to alterations in transvascular and interstitial transport, respectively,

which affect ultimately the local drug concentration and thus the efficacy of chemotherapeutic

agents. Compared to normal tissue, blood vessels in tumors are very leaky [1] and lack a

lymphatic system [2], which can create increased interstititial fluid pressure [3]. As a result,

the pressure gradient between intra- and extravascular spaces is reduced, hindering transport

of (large) molecules across vessel walls by convection. Furthermore, the structure and

function of the vascular system is disorganized in solid tumors, increasing the mean distance

between tumor cells and blood vessels [4,5]. This can lead to reduced access of cytotoxic

drugs to those distant tumor cells. Once a molecule has crossed the vessel wall into the tumor,

it must travel through the interstitium to the tumor cells and, depending on the site of action of

the drug, either bind to a membrane receptor or cross the cell membrane. Nearly uniform

pressure, binding to extracellular matrix [6,7] and enzymatic destruction are some of the

problems encountered by molecules crossing the interstitium. In addition, size and charge of

molecules may hinder their transport across the cell membrane [8].

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Figure 10.1 A diagram of physiological barriers to drug delivery into tumor cells after

systemic administration. The barriers include microvessel wall, extracellularmatrix, and

plasma membrane of cells. They hinder (1) transvascular, (2) interstitial, and (3)

transmembrane transport of drugs, respectively, as indicated by the corresponding numbers

and the open arrows next to them. Adapted from [9].

In general, there are two strategies for overcoming the transport barriers and thereby improve

the efficacy of cancer treatment. The first strategy is modification of the drug by specific

targeting or changing size (by incorporation into a carrier) to increase accumulation at the

tumor site. For example, paclitaxel-loaded nanoparticles were shown to exhibit greater

efficacy when they were conjugated to transferring (Tf) ligand [10]. The greater efficacy is

due to cellular uptake of Tf-conjugated nanoparticles via Tf receptors instead of non specific

endocytosis [11]. Dreher et al. investigated the relationship between molecular weight of

different macromolecular drug carriers and tumor vascular permeability. They showed that

there is an optimal molecular weight for macromolecular drug carriers that results in the

highest accumulation in the tumor [12]. The second strategy is the modification of the tumor’s

physiology to reduce its resistance to the drug accumulation. The tumor blood flow can be

improved by pretreating a tumor with anti-angiogenic therapy, which leads to the

normalization of the tumor vasculature and as a consequence to a reduced interstitial fluid

pressure [13,14]. Modifications of the tumor extracellular matrix [15] and reduction of the

drug

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interstitial fluid pressure [16,17] might also facilitate the penetration of drugs into the tumor.

Other methods that are being developed for enhancing local drug delivery involve external

sources of energy (ultrasound, magnetic and electric fields) that change the physiology of the

tumor or induce the local drug release from specially designed carriers [18].

10.2. Ultrasound facilitated local drug delivery

A large number of these physical methods, developed to reduce transport barriers as a way to

improve the delivery of cytotoxic agents, are based on therapeutic ultrasound exposure and its

interaction with tissue. In the general introduction of this thesis (section 2.3) the different

interactions of ultrasound with tissue were discussed in detail. Below, we briefly review how

local ultrasonic heating, cavitation and radiation force have been exploited in the field of local

drug delivery [19].

Hyperthermia may be particularly helpful in facilitating drug deposition in tumors. As tissue

temperature rises, one of the first physiological reactions is an increase in tumor blood flow,

causing increased delivery of a drug [20]. Furthermore, an increased tumor microvessel pore

size has been noticed in some studies during heating resulting in an increased extravasation of

drug delivery vehicles from tumor vessels [21,22]. This effect is visualized in Figure 10.2

[23]. In contrast, there is no effect on liposomal extravasation from normal vessels [22].

Therefore the increased extravasation of liposomes due to local heating can be exploited as a

drug delivery mechanism in cancer, particularly because the effect appears to be more

important in tumors. Additionally, hyperthermia can be used as a modality for increasing

liposomal drug delivery to tumors by spatially and temporally controlled release of drug from

the liposome [24]. Using liposomes for encapsulating drugs is an effective way to improve

drug delivery to solid tumors, compared to the drug alone, where systemic toxicity is lowered

and uptake into cells is increased [25]. The potential of using temperature-sensitive liposomes

in combination with local hyperthermia for targeted control of local drug release was first

shown by Weinstein and Yatvin [26,27]. Liposomes remain relatively stable in the circulation

at temperatures well below the phase transition temperature (Tc) of the liposome membrane.

At Tc distinctive structural changes occur in the lipid bilayer resulting in increased membrane

permeability and the accompanying release of the liposomes’ content [28-30]. The elevated

temperature needed for the release of the liposomes’ content may be generated by water bath

[24], catheter [31], radio frequency [28,32] or microwave [33]. However, local hyperthermia

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generated by focused ultrasound has a large potential to become a clinical tool, because of its

high spatial precision and non-invasive nature [30,34,35].

Figure 10.2 Epi-illumination images of tumor tissue (human ovarian carcinoma) implanted in

nude mouse with its dorsal skin flap placed in a window chamber. Extravasation of 100 nm

rhodamine-labeled liposomes from tumor vessels at 60 minutes after injection at different

temperatures are shown. (a) 34°C; (b) 39°C; (c) 40°C; (d) 41°C; (e) 42°C. Minimal

extravasation of liposomes was seen at 34°C throughout the 60-minute experiment. At 42°C,

focal perivascular fluorescent spots developed that increased in size and became more

diffuse. From [23].

As with heat also acoustic cavitation can play multiple roles in local drug delivery. Vibrating

and collapsing microbubbles may cause, as discussed before, increased permeability of the

cell membranes and disrupt the shell of co-delivered carriers thereby releasing the enclosed

drugs, resulting in an increased local drug concentration and an enhanced extravasation of the

drug. This approach, whereby microbubbles only serve as cavitation nuclei, may be

principally successful in the microvasculature. In larger vessels, an important part of the drug

will be released in the centre of the vessel and consequently not reach the extravascular

spaces. In this case, it may be helpful to position the drug delivery system first closer to the

vessel wall, using the non-destructive radiation force (as was describe before section 2.4) and

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then apply high intensity ultrasound pulses for microbubble destruction [36]. Another strategy

of using microbubbles in drug delivery is by loading the microbubbles with the drugs. In this

case the destruction of microbubbles leads to the release of the drug and extravasation of the

drug due to the increased number of pores in the vasculature. Recent examples of this strategy

are the delivery of anti-sense androgen receptor oligodeoxynucleotide in vivo, resulting in

inhibition of prostate tumor growth [37] and the use of a conjugation between microbubbles

and liposomes carrying the payload [38]. The feasibility of both strategies has been proven

mainly by using (reporter) genes as payload [39-41]. The microbubbles loaded with

drugs/genes can be targeted to specific (pathologic) sites using different targeting ligands

incorporated into bioconjugates [42]. Microbubbles are successfully targeted to the

pathophysiologic processes like inflammation [43], angiogenesis [44] and thrombus formation

[45], important in many disease states (e.g. atheroscelerosis, tumors, transplant rejection,

etc.). Although the targeted microbubbles in these applications improve the efficacy of

diagnostic imaging, they may also improve ultrasound facilitated drug delivery by increasing

the local concentration near the vessel wall. Since US can be focused, well-defined regions

can be treated. Targeted delivery vehicles may further enhance this on a microscopic scale.

Furthermore, as the microbubbles are relatively stable and circulate through the whole body,

they may passively accumulate at unwanted sites and become toxic. This toxic effect at

unwanted sites can also be reduced by targeting the drug carrying microbubbles to specific

tissues.

The contribution of the radiation force to drug delivery may not only be the induction of gaps

between endothelial cells, widening intracellular spaces in epithelial tissue and positioning the

carriers closer to the vessel wall. Crowder et al. showed that the acoustic radiation force was

the physical mechanism for the augmentation of lipid delivery from nanoparticles. The

acoustic energy, at non-cavitational level, stimulates increased interaction between the

nanoparticle’s lipid layer and the targeted cell’s plasma membrane and thereby improves the

transport of the drugs [46].

10.3. Imaging of drug delivery

A first (important) step for improving the efficacy of anticancer therapy and minimize

systemic toxicity is the reduction of transport barriers by changing the properties of the drug

(carrier) or/and the tumor’s physiology as described above. Though, maybe as important as

facilitating delivery is monitoring non-invasively the drug’s pharmacokinetics (absorption,

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distribution, metabolism and elimination) and pharmacodynamics (drug effects, tolerance,

altered blood flow, toxicity, etc) with a high sensitivity and spatial resolution. Following the

distribution and the metabolism of a drug in vivo gives insight in the behavior and fate of a

drug in a living system and is essential in drug development and treatment optimization.

There is a large spectrum of non-invasive techniques to monitor drug delivery (Figure 10.3).

On the basis of its physics and chemistry, each imaging technique has certain limitations or

advantages with respect to resolution, sensitivity and contrast generation. Below these

limitations and advantages of each imaging modality are discussed specifically for monitoring

drug delivery.

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Figure 10.3 Multiple imaging modalities are available for small-animal molecular imaging.

Shown are views of typical instruments available and illustrative examples of images that can

be obtained with these modalities. (A) Positron Emission Tomography (B) Computed

Tomography (C) Single Photon Emission CT (D) Optical reflectance fluorescence imaging

(E) Magnetic Resonance Imaging (F) Optical bioluminescence imaging. Many of these

imaging modalities are in routine clinical use, making translation from animal model to bed

side possible.

Local concentrations of drugs for therapy are in the range of µM. Only nuclear medicine

techniques such as gamma-scintigraphy, single-photon emission computer tomography

(SPECT) or positron emission tomography (PET), and optical techniques have the required

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sensitivity to detect micromolar concentrations. However, nuclear medicine methods are

hampered by relatively low spatial resolution (1.5 mm) which, although acceptable in clinical

applications, presents a serious limitation when studying small animals. Furthermore, these

methods lack chemical specificity, being unable to distinguish whether the emitting

radioisotope is the parent drug molecule or a metabolite. Nuclear medicine techniques are also

lacking intrinsic anatomic information and require the use of radioactive tracers.

Magnetic resonance imaging (MRI), on the other hand, provides high spatial resolution (pixel

dimensions of 100 µm and better) and is non-invasive, but MRI requires tissue concentrations

in the millimolar range. Due to the high spatial resolution detailed morphological information

can be obtained in the small animal models predominantly used in pharmacological research.

However, the low sensitivity makes the use of ingenious amplification strategies necessary to

detect the low drug concentrations. The use of MRI for monitoring pharmacokinetics and

dynamics offers other advantages. First, since the method is non-invasive it allows repetitive

measurements in the same animal, which leads to statistical and economical advantages and

decreased used of animals. Secondly, MRI takes advantage of the enormous library of

acquisition sequences that provide different image contrast that leads to a high contrast

resolution, i.e. the ability to distinguish the differences between two arbitrarily similar but not

identical tissues. Besides the larger number of endogenous parameters (proton density,

relaxation times, water diffusion, water exchange rates) that provide optimal contrast for

different soft-tissue structures, MRI also exploits a large range of so-called biomarkers. A MR

biomarker is an anatomic, physiologic, biochemical or molecular parameter detectable with

MRI used to establish the presence or the rate of a process of interest, such as a disease or the

deposition of a drug. Examples of MR biomarkers are perfusion and diffusion studies,

magnetic resonance spectroscopy (MRS) and contrast agents. Perfusion studies, based on the

dynamic tracking of a bolus of contrast agent, allow early evaluation of anti-angiogenesis

therapies [47,48]. Imaging of water diffusion has also shown great potential in early

evaluation of drug response [49-52]. Spectroscopy studies provide an early analysis of

modifications of metabolism, such as that of choline, a marker of tumor metabolism [53]. The

concentration of the drug itself can also be monitored by MRS when the drug contains atoms

with magnetic nuclei in their structure (e.g. misanidazole (2H) [54], 5-fluorouracil (19F) [55]

and ifosfamide (31P) [56]). However, at the moment the MRS signal is only sufficient for 5-

fluorouracil at concentrations used for therapeutic action. MR contrast agents (e.g.

gadolinium, iron oxide particles) can be loaded into or attached to drug delivery vehicles such

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as liposomes and micelles, which allows real-time monitoring of the distribution of the carrier

in vivo [57,58].

Other imaging methods used for monitoring drug pharmacokinetics are ultrasound (US) and

optical imaging. US imaging appears as very suitable method when using microbubbles as

drug delivery vehicles, because of its very high sensitivity. Even a single microbubble may be

detected. However, upon the collapse of the microbubble and release of the drug, this

hyperechogenicity disappears. Therefore US imaging can only be used to follow the delivery

vehicle to the location for release, but the subsequent pathway of the drug can not be

monitored directly. Optical imaging methods (fluorescence and bioluminescence) are gaining

more and more interest as monitoring modalities, essentially because of the high sensitivity

(10-12 moles/L [59]), albeit only for small animals because of absorption of light by tissue

[60]. But also the large choice in probes and the possibility to work at different wavelengths

makes optical imaging a versatile technique. A very sophisticated example of fluorescence

imaging in drug delivery was shown by Bagalkot et al. [61]. They developed a

multifunctional nanoparticle that is capable of sensing the release of therapeutic modality by a

change in the fluorescence of the imaging modality.

The best way to monitor the pharmacokinetics and pharmacodynamics of a drug is to follow

the drug itself. With PET this is possible for all drugs containing in their structure a

radionuclide (e.g. 18F, 11C and 15O). The same is true for drugs that contain in their structure

atoms with magnetic nuclei when using MRS as imaging method. In MRI this is done by

attaching [62] or co-injecting [63] MR contrast agents with the drug. When the contrast is co-

injected with the drug the relation between the drug concentration and the signal distribution

of the contrast agent has to be known. In case the contrast agent is attached to the drug this

relationship is known, but the contrast agent may modulate the pharmacokinetics of the drug

compared to the unbound state. A new approach is loading the drug and the contrast agent in

the same delivery system [64]. Viglianti et al. used MnSO4/doxorubicin loaded liposomes to

monitor in vivo liposome concentration distribution and drug release [31]. In a follow-up

study they confirmed that there was a linear relationship between a change in MRI contrast

and the amount of deposited drugs [65]. Results from this study are shown in Figure 10.4.

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Figure 10.4 Validation of T1-map based doxorubicin (DOX) concentrations released from

temperature-sensitive liposomes (TSL) loaded with MnSO4 by invasive methods in two

independent studies. (A,D) Raw signal intensity map (0-125 a.u. color bar) shows an axial

view of a rat bearing a flank fibrosarcoma (top left) with a central heating catheter at

beginning and at 45 min after DOX injection, respectively. (B,E) T1-intensity map (0-3000 ms

color bar) at beginning and at 45 min after injection, respectively. Note that the regions that

are enhanced in d have reduced T1 intensity in e, indicating contrast/drug presence through

T1 shortening. (C) The calculated DOX concentration (ng/mg) on a pixel-by-pixel basis using

images b and e. (F) An enlarged image of c showing the heterogeneity in drug delivery that

can be imaged and quantified by this MRI technique. (G) The results for the HPLC validated

[DOX] measurements from each animal. (H) The results for the fluorescence validated

[DOX] measurements. i: An overlay of both experiments is displayed, showing the precision

and accuracy of MRI for measuring DOX at lower concentrations. Figure adapted from[65].

To take maximal profit of the different imaging modalities in drug delivery, they should not

only be used for real-time monitoring of drug distribution, but also for treatment planning and

follow-up. However, each analysis has specific requirements that can not be found all in the

same image modality. Therefore, in order to prevent making concessions on essential

parameters (such as spatial resolution, sensitivity, cost, availability etc.) of the analysis,

117

hybrid systems are used more and more. Combinations of two or more different imaging

modalities such as PET/MR [66], PET/CT [67] and MR/fluorescence [68,69] combine the

strengths of the different imaging techniques to fulfil the demanding requirements in local

drug delivery.

10.4. References

[1] Gerlowski LE, Jain RK. Microvascular permeability of normal and neoplastic tissues.

Microvasc Res 1986;31(3):288-305.

[2] Leu AJ, Berk DA, Lymboussaki A, Alitalo K, Jain RK. Absence of functional

lymphatics within a murine sarcoma: a molecular and functional evaluation. Cancer

Res 2000;60(16):4324-4327.

[3] Boucher Y, Baxter LT, Jain RK. Interstitial pressure gradients in tissue-isolated and

subcutaneous tumors: implications for therapy. Cancer Res 1990;50(15):4478-4484.

[4] Jain RK. Delivery of molecular medicine to solid tumors. Science

1996;271(5252):1079-1080.

[5] Jain RK. The next frontier of molecular medicine: delivery of therapeutics. Nat Med

1998;4(6):655-657.

[6] Berk DA, Yuan F, Leunig M, Jain RK. Direct in vivo measurement of targeted

binding in a human tumor xenograft. Proc Natl Acad Sci U S A 1997;94(5):1785-

1790.

[7] Juweid M, Neumann R, Paik C, Perez-Bacete MJ, Sato J, van Osdol W, Weinstein JN.

Micropharmacology of monoclonal antibodies in solid tumors: direct experimental

evidence for a binding site barrier. Cancer Res 1992;52(19):5144-5153.

[8] Schlicher RK, Radhakrishna H, Tolentino TP, Apkarian RP, Zarnitsyn V, Prausnitz

MR. Mechanism of intracellular delivery by acoustic cavitation. Ultrasound Med Biol

2006;32(6):915-924.

[9] Wang Y, Yuan F. Delivery of viral vectors to tumor cells: extracellular transport,

systemic distribution, and strategies for improvement. Ann Biomed Eng

2006;34(1):114-127.

[10] Sahoo SK, Ma W, Labhasetwar V. Efficacy of transferrin-conjugated paclitaxel-

loaded nanoparticles in a murine model of prostate cancer. Int J Cancer

2004;112(2):335-340.

118

[11] Wagner E, Curiel D, Cotten M. Delivery of drugs, proteins and genes into cells using

transferrin as a ligand for receptor-mediated endocytosis. Adv Drug Delivery Rev

1994;14(1):113-135.

[12] Dreher MR, Liu W, Michelich CR, Dewhirst MW, Yuan F, Chilkoti A. Tumor

vascular permeability, accumulation, and penetration of macromolecular drug carriers.

J Natl Cancer Inst 2006;98(5):335-344.

[13] Jain RK. Normalizing tumor vasculature with anti-angiogenic therapy: a new

paradigm for combination therapy. Nat Med 2001;7(9):987-989.

[14] Jain RK. Normalization of tumor vasculature: an emerging concept in antiangiogenic

therapy. Science 2005;307(5706):58-62.

[15] McKee TD, Grandi P, Mok W, Alexandrakis G, Insin N, Zimmer JP, Bawendi MG,

Boucher Y, Breakefield XO, Jain RK. Degradation of fibrillar collagen in a human

melanoma xenograft improves the efficacy of an oncolytic herpes simplex virus

vector. Cancer Res 2006;66(5):2509-2513.

[16] Griffon-Etienne G, Boucher Y, Brekken C, Suit HD, Jain RK. Taxane-induced

apoptosis decompresses blood vessels and lowers interstitial fluid pressure in solid

tumors: clinical implications. Cancer Res 1999;59(15):3776-3782.

[17] Rubin K, Sjoquist M, Gustafsson AM, Isaksson B, Salvessen G, Reed RK. Lowering

of tumoral interstitial fluid pressure by prostaglandin E(1) is paralleled by an increased

uptake of (51)Cr-EDTA. Int J Cancer 2000;86(5):636-643.

[18] Besic E. Physical mechanisms and methods employed in drug delivery to tumors. Acta

Pharm 2007;57(3):249-268.

[19] Tachibana K, Tachibana S. The use of ultrasound for drug delivery. Echocardiography

2001;18(4):323-328.

[20] Karino T, Koga S, Maeta M. Experimental studies of the effects of local hyperthermia

on blood flow, oxygen pressure and pH in tumors. Jpn J Surg 1988;18(3):276-283.

[21] Gaber MH, Wu NZ, Hong K, Huang SK, Dewhirst MW, Papahadjopoulos D.

Thermosensitive liposomes: extravasation and release of contents in tumor

microvascular networks. Int J Radiat Oncol Biol Phys 1996;36(5):1177-1187.

[22] Kong G, Braun RD, Dewhirst MW. Hyperthermia enables tumor-specific nanoparticle

delivery: effect of particle size. Cancer Res 2000;60(16):4440-4445.

[23] Kong G, Braun RD, Dewhirst MW. Characterization of the effect of hyperthermia on

nanoparticle extravasation from tumor vasculature. Cancer Res 2001;61(7):3027-

3032.

119

[24] Kong G, Anyarambhatla G, Petros WP, Braun RD, Colvin OM, Needham D, Dewhirst

MW. Efficacy of liposomes and hyperthermia in a human tumor xenograft model:

importance of triggered drug release. Cancer Res 2000;60(24):6950-6957.

[25] Allen TM. Liposomes. Opportunities in drug delivery. Drugs 1997;54 Suppl 4:8-14.

[26] Weinstein JN, Magin RL, Yatvin MB, Zaharko DS. Liposomes and local

hyperthermia: selective delivery of methotrexate to heated tumors. Science

1979;204(4389):188-191.

[27] Yatvin MB, Weinstein JN, Dennis WH, Blumenthal R. Design of liposomes for

enhanced local release of drugs by hyperthermia. Science 1978;202(4374):1290-1293.

[28] Bos C, Lepetit-Coiffe M, Quesson B, Moonen CT. Simultaneous monitoring of

temperature and T1: methods and preliminary results of application to drug delivery

using thermosensitive liposomes. Magn Reson Med 2005;54(4):1020-1024.

[29] Fossheim SL, Il'yasov KA, Hennig J, Bjornerud A. Thermosensitive paramagnetic

liposomes for temperature control during MR imaging-guided hyperthermia: in vitro

feasibility studies. Acad Radiol 2000;7(12):1107-1115.

[30] Frenkel V, Etherington A, Greene M, Quijano J, Xie J, Hunter F, Dromi S, Li KC.

Delivery of liposomal doxorubicin (Doxil) in a breast cancer tumor model:

investigation of potential enhancement by pulsed-high intensity focused ultrasound

exposure. Acad Radiol 2006;13(4):469-479.

[31] Viglianti BL, Abraham SA, Michelich CR, Yarmolenko PS, MacFall JR, Bally MB,

Dewhirst MW. In vivo monitoring of tissue pharmacokinetics of liposome/drug using

MRI: illustration of targeted delivery. Magn Reson Med 2004;51(6):1153-1162.

[32] Aoki H, Kakinuma K, Morita K, Kato M, Uzuka T, Igor G, Takahashi H, Tanaka R.

Therapeutic efficacy of targeting chemotherapy using local hyperthermia and

thermosensitive liposome: evaluation of drug distribution in a rat glioma model.

International Journal of Hyperthermia 2004;20(6):595-605.

[33] Khoobehi B, Peyman GA, Niesman MR, Oncel M. Hyperthermia and temperature-

sensitive liposomes: selective delivery of drugs into the eye. Jpn J Ophthalmol

1989;33(4):405-412.

[34] Dromi S, Frenkel V, Luk A, Traughber B, Angstadt M, Bur M, Poff J, Jianwu Xie J,

Libutti SK, Li KCP, B.J. W. Pulsed-High Intensity Focused Ultrasound and Low

Temperature Sensitive Liposomes for EnhancedTargeted Drug Delivery and

Antitumor Effect. Clin Cancer Res 2007;13(9):2722-2727.

120

[35] McDannold N, Fossheim SL, Rasmussen H, Martin H, Natalia Vykhodtseva N,

Hynynen K. Heat-activated liposomal MR contrast agent: initial in vivo results in

rabbit liver and kidney. Radiology 2004;230(3):743-752.

[36] Shortencarier MJ, Dayton PA, Bloch SH, Schumann PA, Matsunaga TO, Ferrara KW.

A method for radiation-force localized drug delivery using gas-filled lipospheres.

IEEE Trans Ultrason Ferroelectr Freq Control 2004;51(7):822-831.

[37] Haag P, Frauscher F, Gradl J, Seitz A, Schafer G, Lindner JR, Klibanov AL, Bartsch

G, Klocker H, Eder IE. Microbubble-enhanced ultrasound to deliver an antisense

oligodeoxynucleotide targeting the human androgen receptor into prostate tumours. J

Steroid Biochem Mol Biol 2006;102(1-5):103-113.

[38] Kheirolomoom A, Dayton PA, Lum AF, Little E, Paoli EE, Zheng H, Ferrara KW.

Acoustically-active microbubbles conjugated to liposomes: Characterization of a

proposed drug delivery vehicle. J Control Release 2006.

[39] Miller DL, Pislaru SV, Greenleaf JE. Sonoporation: mechanical DNA delivery by

ultrasonic cavitation. Somat Cell Mol Genet 2002;27(1-6):115-134.

[40] Shohet RV, Chen S, Zhou YT, Wang Z, Meidell RS, Unger RH, Grayburn PA.

Echocardiographic destruction of albumin microbubbles directs gene delivery to the

myocardium. Circulation 2000;101(22):2554-2556.

[41] Vannan M, McCreery T, Li P, Han Z, Unger E, Kuersten B, Nabel E, Rajagopalan S.

Ultrasound-mediated transfection of canine myocardium by intravenous

administration of cationic microbubble-linked plasmid DNA. J Am Soc Echocardiogr

2002;15(3):214-218.

[42] Klibanov AL. Microbubble contrast agents: targeted ultrasound imaging and

ultrasound-assisted drug-delivery applications. Invest Radiol 2006;41(3):354-362.

[43] Lindner JR, Coggins MP, Kaul S, Klibanov AL, Brandenburger GH, Ley K.

Microbubble persistence in the microcirculation during ischemia/reperfusion and

inflammation is caused by integrin- and complement-mediated adherence to activated

leukocytes. Circulation 2000;101(6):668-675.

[44] Ellegala DB, Leong-Poi H, Carpenter JE, Klibanov AL, Kaul S, Shaffrey ME, Sklenar

J, Lindner JR. Imaging tumor angiogenesis with contrast ultrasound and microbubbles

targeted to alpha(v)beta3. Circulation 2003;108(3):336-341.

[45] Schumann PA, Christiansen JP, Quigley RM, McCreery TP, Sweitzer RH, Unger EC,

Lindner JR, Matsunaga TO. Targeted-microbubble binding selectively to GPIIb IIIa

receptors of platelet thrombi. Invest Radiol 2002;37(11):587-593.

121

[46] Crowder KC, Hughes MS, Marsh JN, Barbieri AM, Fuhrhop RW, Lanza GM,

Wickline SA. Sonic activation of molecularly-targeted nanoparticles accelerates

transmembrane lipid delivery to cancer cells through contact-mediated mechanisms:

implications for enhanced local drug delivery. Ultrasound Med Biol

2005;31(12):1693-1700.

[47] Miyazaki K, Collins DJ, Walker-Samuel S, Taylor JN, Padhani AR, Leach MO, Koh

DM. Quantitative mapping of hepatic perfusion index using MR imaging: a potential

reproducible tool for assessing tumour response to treatment with the antiangiogenic

compound BIBF 1120, a potent triple angiokinase inhibitor. Eur Radiol

2008;18(7):1414-1421.

[48] Muruganandham M, Lupu M, Dyke JP, Matei C, Linn M, Packman K, Kolinsky K,

Higgins B, Koutcher JA. Preclinical evaluation of tumor microvascular response to a

novel antiangiogenic/antitumor agent RO0281501 by dynamic contrast-enhanced MRI

at 1.5 T. Mol Cancer Ther 2006;5(8):1950-1957.

[49] Cui Y, Zhang XP, Sun YS, Tang L, Shen L. Apparent diffusion coefficient: potential

imaging biomarker for prediction and early detection of response to chemotherapy in

hepatic metastases. Radiology 2008;248(3):894-900.

[50] Dudeck O, Zeile M, Pink D, Pech M, Tunn PU, Reichardt P, Ludwig WD, Hamm B.

Diffusion-weighted magnetic resonance imaging allows monitoring of anticancer

treatment effects in patients with soft-tissue sarcomas. J Magn Reson Imaging

2008;27(5):1109-1113.

[51] Jordan BF, Runquist M, Raghunand N, Baker A, Williams R, Kirkpatrick L, Powis G,

Gillies RJ. Dynamic contrast-enhanced and diffusion MRI show rapid and dramatic

changes in tumor microenvironment in response to inhibition of HIF-1alpha using PX-

478. Neoplasia 2005;7(5):475-485.

[52] Kim H, Morgan DE, Zeng H, Grizzle WE, Warram JM, Stockard CR, Wang D, Zinn

KR. Breast tumor xenografts: diffusion-weighted MR imaging to assess early therapy

with novel apoptosis-inducing anti-DR5 antibody. Radiology 2008;248(3):844-851.

[53] Glunde K, Jacobs MA, Bhujwalla ZM. Choline metabolism in cancer: implications for

diagnosis and therapy. Expert Rev Mol Diagn 2006;6(6):821-829.

[54] Evelhoch JL, McCoy CL, Giri BP. A method for direct in vivo measurement of drug

concentrations from a single 2H NMR spectrum. Magn Reson Med 1989;9(3):402-

410.

122

[55] Brix G, Bellemann ME, Haberkorn U, Gerlach L, Bachert P, Lorenz WJ. Mapping the

biodistribution and catabolism of 5-fluorouracil in tumor-bearing rats by chemical-

shift selective 19F MR imaging. Magn Reson Med 1995;34(3):302-307.

[56] Rodrigues LM, Maxwell RJ, McSheehy PM, Pinkerton CR, Robinson SP, Stubbs M,

Griffiths JR. In vivo detection of ifosfamide by 31P-MRS in rat tumours: increased

uptake and cytotoxicity induced by carbogen breathing in GH3 prolactinomas. Br J

Cancer 1997;75(1):62-68.

[57] Mulder WJ, Strijkers GJ, van Tilborg GA, Griffioen AW, Nicolay K. Lipid-based

nanoparticles for contrast-enhanced MRI and molecular imaging. NMR Biomed

2006;19(1):142-164.

[58] Torchilin VP. Multifunctional nanocarriers. Adv Drug Deliv Rev 2006;58(14):1532-

1555.

[59] Massoud TF, Gambhir SS. Molecular imaging in living subjects: seeing fundamental

biological processes in a new light. Genes Dev 2003;17(5):545-580.

[60] Gumbleton M, Stephens DJ. Coming out of the dark: the evolving role of fluorescence

imaging in drug delivery research. Adv Drug Deliv Rev 2005;57(1):5-15.

[61] Bagalkot V, Zhang L, Levy-Nissenbaum E, Jon S, Kantoff PW, Langer R, Farokhzad

OC. Quantum dot-aptamer conjugates for synchronous cancer imaging, therapy, and

sensing of drug delivery based on bi-fluorescence resonance energy transfer. Nano

Lett 2007;7(10):3065-3070.

[62] Peira E, Marzola P, Podio V, Aime S, Sbarbati A, Gasco MR. In vitro and in vivo

study of solid lipid nanoparticles loaded with superparamagnetic iron oxide. J Drug

Target 2003;11(1):19-24.

[63] Saito R, Bringas JR, McKnight TR, Wendland MF, Mamot C, Drummond DC,

Kirpotin DB, Park JW, Berger MS, Bankiewicz KS. Distribution of liposomes into

brain and rat brain tumor models by convection-enhanced delivery monitored with

magnetic resonance imaging. Cancer Res 2004;64(7):2572-2579.

[64] De Zwart JA, Salomir R, Vimeux F, Klaveness J, Moonen CTW. On the feasibility of

local drug delivery using thermo-sensitive liposomes and MR-guided focused

ultrasound. 2000; Denver. p 43.

[65] Viglianti BL, Ponce AM, Michelich CR, Yu D, Abraham SA, Sanders L, Yarmolenko

PS, Schroeder T, MacFall JR, Barboriak DP, Colvin OM, Bally MB, Dewhirst MW.

Chemodosimetry of in vivo tumor liposomal drug concentration using MRI. Magn

Reson Med 2006;56(5):1011-1018.

123

[66] Pichler BJ, Judenhofer MS, Wehrl HF. PET/MRI hybrid imaging: devices and initial

results. Eur Radiol 2008;18(6):1077-1086.

[67] Chowdhury FU, Scarsbrook AF. The role of hybrid SPECT-CT in oncology: current

and emerging clinical applications. Clin Radiol 2008;63(3):241-251.

[68] Kircher MF, Weissleder R, Josephson L. A dual fluorochrome probe for imaging

proteases. Bioconjug Chem 2004;15(2):242-248.

[69] Mulder WJ, Strijkers GJ, Habets JW, Bleeker EJ, van der Schaft DW, Storm G,

Koning GA, Griffioen AW, Nicolay K. MR molecular imaging and fluorescence

microscopy for identification of activated tumor endothelium using a bimodal lipidic

nanoparticle. Faseb J 2005;19(14):2008-2010.

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125

Chapter 11. MRI monitoring of ultrasound

mediated drug delivery

11.1. Introduction

The pharmacological action of a chemotherapeutic drug, as well as its toxicological effect is

related to its tissue concentration. Therefore, a local increase of therapeutic drug in the region

where therapy is required would increase the efficacy of the therapy with a reduced systemic

toxicity. Modification of the drug (carrier) or the tissue’s physiology may reduce the

resistance to drug accumulation in the pathologic tissue [1]. It has been shown that therapeutic

ultrasound can improve the delivery of genes and drugs into cells [2-6] and can alter vascular

permeability for increased extravasation and hence improved delivery to whole tissues [7,8].

However, interaction of ultrasound with living tissues may also cause irreversible tissue

damage, such as microvessel rupture [9], hemorrhage [10], apoptosis [11], hemolysis [12] and

necrosis [13], which are incompatible with the objective of non destructive drug deposition.

The likelihood of the desired as well as the adverse effects, depends on the interaction of

ultrasound with tissue [14,15].

Cavitation is considered as the most important mechanism to enhance gene/drug delivery.

This effect can be induced in deep tissue non-invasively by ultrasound waves, depending on

the frequency and pressure of the sonication. The threshold of acoustic pressure for inducing

cavitation can be lowered by administration of US contrast agent [16], reducing the possibility

of other physical effects such as tissue heating.

The efficacy of increased delivery by cavitation is influenced by acoustic parameters, tissue

characteristics and the type of US contrast agent, if present. Rahim et al. and Zarnitsyn et al.

have investigated in vitro the influence of several physical sonication parameters and

microbubbles concentration on the efficacy of intracellular gene delivery [17,18]. Hallow et

al. have analyzed the influence of different acoustic energy levels on the delivery of an optical

contrast agent as pseudo-drug in ex vivo arteries [19]. Most in vivo optimization of ultrasound

mediated drug delivery is performed with anticancer cytotoxic agents. The efficacy of the

used parameters is determined from tumor regression after sonication. Larkin et al. showed,

using a tumor model, the influence of acoustic energy intensity, duration of sonication and

duty cycle on the cytotoxicity of bleomycin [20]. Iwanaga et al. showed in a similar

126

experiment that presence of microbubbles increases the therapeutic effect of bleomycin [21].

Bekeredjian et al. have shown an increased capillary permeability in tumor and muscle using

Evans Blue as reporter molecule [22].

Although it was obvious from these studies that ultrasound improved the therapeutic effect,

distinct differences between the different ultrasound protocols were only observable after 3 to

4 days. Despite the growing interest of such non-invasive therapeutic approaches, no real-time

imaging methods have been proposed to monitor the effect of different sonication parameters.

Non-invasive and near real-time monitoring with high sensitivity and spatial resolution of the

pharmacokinetic and dynamic effects during drug deposition is essential for treatment

optimization.

Visualization of the local deposition of a MR contrast agent in the brain was proposed by

Treat et al [23]. In their study, the local disruption of the blood-brain barrier by focalized

pulsed US was visualized immediately after the procedure on T1-weighted images by

injecting a MR contrast agent. The delivered doses of MR contrast agent and of co-

administrered drug (Doxorubicin) were demonstrated to be dependent on the acoustic power

and injected US contrast agent. However, no analysis of the time course of the effect was

investigated.

This study presents a method for monitoring in vivo with MRI the changes of distribution of a

reporter macromolecule in hepatic tissue induced by cavitation. Experiments were performed

on small animals with clinical US imaging and MRI devices. Cavitation effect was enhanced

with the help of a clinically accepted contrast agent (Sonovue) and the reporter

macromolecule was a MRI contrast agent (Vistarem). This macromolecule influences the

longitudinal relaxation time (T1) of the MR signal but does not spontaneously diffuse outside

the vasculature [24]. Therefore, modifying the local distribution of Vistarem in hepatic tissue

by cavitation should result in apparent T1 modifications. Comparison of T1 changes in time of

hepatic tissue in presence or absence of cavitation should thus provide a non-invasive

estimate of the relative influence of sonication characteristics (e.g. pressure, duration,

presence of microbubbles) on the potential extravasation effects.

127

11.2. Materials and methods

Animal

Male Wistar rats (350 to 550g) were anesthetized with isoflurane in air (3% for the induction,

and 2% for the remaining experiment). The abdomen was carefully shaved with depilatory

cream to improve ultrasound propagation in the liver. A 24-gauge cannula (Insyte®, BD) was

inserted into a tail. A continuous slow perfusion of sodium chloride (0.9%) was maintained

until injections of US and MRI contrast agents were performed through this catheter with the

help of a 3 track line tap. This animal procedure was approved by the university committee

for the use and care of animals.

US contrast agent

SonovVue (Bracco) is a microbubble contrast agent of the second generation consisting of a

sulphur hexafluoride (SF6) gas core surrounded by a thin and flexible shell of phospholipids.

Sonovue contains microbubbles of different sizes, ranging between 1 and 10 µm, with a mean

size of 2.5 µm [25]. In humans the maximum blood concentration is reached within 1-2

minutes and then rapidly declines (T1/2 ~ 7 minutes) [26]. There is no such information for

rodents. Sonovue was prepared according to the manufacturer’s instructions; leading to a

concentration of microbubbles in the range of 1,0·109 microbubbles per ml . Sonovue was

injected in a slow bolus at the dose of 0,05ml/100g through the perfusion line.

MRI contrast agent

P792 (Gadomelitol, Vistarem ®, Guerbet) is a macromolecular blood pool MR contrast agent

with rapid clearance (T1/2 ~ 20 minutes) [27]. This contrast agent was selected in the present

study since it remains in the blood for a prolonged time and has a higher relaxivity as

compared to conventional gadolinium based contrast agents (e.g. Dotarem and Magnevist).

Therefore, it may serve as MR reporter agent for monitoring changes of capillary permeability

of the rat liver induced by destruction of Sonovue by ultrasound. This blood pool contrast

agent was administered in a slow bolus (20-30 sec) at the dose of 40 µmol/kg through the

perfusion line.

Ultrasound device

A clinical echograph (Acuson sequoia 512-Siemens) was used at an operating frequency of 2

MHz with the 4C1 transducer. The probe was positioned on the abdomen after application of

a colloid on the skin. Localization of the liver was performed on two dimensional echographic

128

images prior to injection of Sonovue, with a scan depth of 3.5 cm and a constant focus located

at half of this value. The CPS (cadence Contrast Pulse Sequence) mode was used to visualize

the arrival of microbubbles into the liver (about 3 to 5 seconds after injection) and was rapidly

switched to the pulsed mode imaging with color Doppler for potential local destruction of

microbubbles. Destruction of microbubbles within the liver was visualized on the screen of

the echograph. Manipulation of the echograph was systematically performed by the same

operator to minimize inter-operator variation.

Magnetic resonance imaging

Dynamic measurements of T1 values were performed on Philips Achieva 1.5 Tesla by

repeating a fast inversion recovery (IR) sequence (Look-Locker sequence , TR = 23ms, TE =

11 ms, EPI factor = 11, flip angle = 30°, FOV = 96 × 96 mm2 matrix = 96 × 96, NSA = 4)

before and during one hour after intravenous co-administration of the MRI blood pool agent

and microbubbles. About 30 acquisitions at different inversion times were performed to

measure the recovery of the longitudinal magnetization, with respiratory synchronization,

leading to an acquisition time of 2 to 3 minutes, depending on the breathing period.

Image processing and data analysis

The images obtained from each Look-Locker measurement were automatically fitted with in-

house developed software written in IDL language (ITT Corporation). To obtain a parametric

map of the T1 values, the temporal evolution of the magnitude of the longitudinal

magnetization was fitted (Levenberg-Marquardt algorithm) for each pixel with the following

equation:

( ) 10

T

t

z ebaMtM−

⋅−⋅= 11-1

Mo corresponded to the longitudinal magnetization at the equilibrium and Mz corresponded to

the longitudinal magnetization in time. The initial values of the parameters a and b were set to

1 and 2 (ideal inversion recovery), respectively, and these parameters were allowed to vary in

the fitting process between 0,5 and 1,5 for a and between 1,5 and 2,5 for b, to account for

potential local variations of the flip angles related to non uniform B1 values. This automatic

process was repeated for each individual Look-Locker measurement. Then, three regions of

interest (ROI) were manually selected by an experienced radiologist in the left, center and

129

right parts of the liver to analyze the evolution of T1 (mean ± standard deviation) as a function

of time. The T1 values obtained were expressed in percentage of the initial T1 of each animal.

Experimental design

The change in distribution of macromolecular MRI contrast agent due to ultrasound and

microbubbles was monitored with the proposed method. For this purpose, animals (n = 12)

were placed in supine position in the MRI scanner with a 4.7 cm in diameter surface receiver

coil taped on the shaved abdomen. A reference T1-map was measured using the Look-Locker

sequence. Next the macromolecular MRI contrast agent (Vistarem) was injected, followed by

a second injection, of the US contrast agent. In one group of rats (n = 5), these microbubbles

were immediately visualized and destroyed with the clinical echograph located near the table

top of the MRI scanner. For microbubble destruction sonications were performed during 2

minutes at a mechanical index (MI) of 1.5 and a frequency of 2 MHz. In the control group (n

= 7), no ultrasound was applied. Measurements of the hepatic T1 values in time were

performed with the Look-Locker technique (see above for details) during 1 hour. To ensure

that the animal remained in the same position before and after sonication, a home made MRI

compatible animal holder was designed. This apparatus could slide horizontally on the table

top to allow sonications within the liver with the non-MR compatible echograph entered in

the Faraday cage. After the 2 minutes sonication, the echograph was removed and the holder

was translated back to its initial position at the magnet center. The total duration required for

injections, sonication and repositioning the animal was approximately 4 minutes. Body

temperature of the animal was continuously monitored with a rectal optical probe (Luxtron®),

since variations of temperature may influence T1 values.

130

Figure 11.1 Timing diagram of performed experiments. First a reference T1 map was

measured using the Look-locker sequence, followed by i.v. injections of Vistarem and

microbubbles. Next ultrasound was applied during 2 minutes for one group of animals.

Finally, T1 maps were acquired continuously during 1 hour.

11.3. Results

Figure 11.2 displays typical results obtained from the fast inversion-recovery imaging

sequence. Image intensities in transverse orientation are displayed as a function of the

inversion times (Figure 11.2a). Figure 11.2b shows typical results of the evolution of the MR

signal in a single pixel located in the liver. The signal intensity was measured for 29

successive inversion times (black dots). The solid line is the result of the fit of these data with

equation 11-1. Parametric maps of the M0 and T1-values are displayed in Figure 11.2c and d

to illustrate the quality of the results.

131

Figure 11.2 Typical results obtained from the fast inversion-recovery imaging sequence. (a)

Image intensities in transverse orientation are displayed as a function of increasing inversion

times. (b) Typical results of the evolution of the MR signal in a single pixel located in the

liver. The signal intensity was measured for 29 successive inversion times (black dots). The

solid line is the result of the fit of these data with equation 11-1. (c) M0 map with ROI used for

temporal analysis. (d) Typical T1-map.

In order to investigate non-cavitation related variation of T1 values on an anesthetized animal

during 1 hour, a first set of experiments (n = 2) was performed without any injections nor

sonications. Negligible variations of T1-value (3%) in individual rats were observed, though

there were small oscillations in body temperature (< 0,5 °C). The T1-values measured in the

different rats were not identical (385 ± 9 ms and 310 ± 15 ms). Therefore, T1 values measured

at different time point were normalized to the initial value (average of 2 to 4 measurements).

The ROI indicated on the T1-map in Figure 11.2c was used for the temporal analysis of T1-

values. Figure 11.3 compares the evolutions in time of the mean T1-value in this ROI for one

rat with sonication (square) and for one rat without sonication (circle). Monitoring of the T1

values was started approximately 4 minutes after the end of sonication. For both experiments,

T1 decreased after injection of MR contrast agent and sonication. Monitoring of the T1 values

after the sonication with the Look-Locker method shows an immediate and slow recovery of

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the T1 values for both experiments. For both cases, the T1-values did not return to the initial

T1-value 60 minutes after injection of the macromolecular contrast agent.

Figure 11.3 Temporal evolution of mean T1 values for an ultrasound treated rat (square) and

for a control rat (circle)

Figure 11.4 shows the temporal evolution of normalized T1 values at 7, 15 minutes and 60

minutes in the ROI as indicated in Figure 11.2c for the two groups of animals. A clear

difference was observed between the 2 groups, with systematic lower values for the group

with destruction of microbubbles (n = 5) as compared to the control group (n = 7), indicative

of a change of the interaction of the MRI contrast agent with the liver tissue. Seven minutes

after injection of Vistarem the normalized T1 represented 63% (± 6%) of the initial T1 in the

group with microbubbles destruction versus 81% (± 4%) in the control group. This difference

progressively decreased in time but still remained one hour after injection (79% ± 3% vs 90%

5%). A similar behavior was observed for the two ROIs located on the left and right parts of

the liver, respectively. This was expected since the ultrasonic field induced by the probe

(4C1) used for destruction of the Sonovue covered the complete liver of the animals. Body

temperature remained stable during these experiments (mean variation of 1,1 ± 0,7 °C).

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Figure 11.4 Normalized T1 values for ultrasound treated (square) and non-ultrasound treated

control rats (circle) at t = 7, 15 and 60 minutes after Vistarem and microbubble injection.

11.4. Discussion

The proposed method allows for non invasive imaging of the time course of the change in the

interaction between a blood pool MR contrast agent and the surrounding hepatic tissue. For

this purpose, one group of animals was sonicated with a high mechanical index after injection

of MR contrast agent and microbubbles in order to evoke cavitation mediated extravasation of

the MR contrast agent. A control group was also injected with MR contrast agent and

microbubbles, but was not sonicated. After injections and in one case sonication, quantitative

evolution of the longitudinal relaxation time was measured dynamically in vivo in rats with

the help of a fast inversion-recovery sequence.

Measurements performed on the control group showed a slow recovery of the T1 values after

injection of Vistarem, with an elimination time in the range of 1 hour. This value corresponds

to invasive blood measurements reported in the literature [28].

In addition, the difference between T1 values in the two groups of animals remained

significantly different after one hour. With respect to this time scale, the temporal resolution

of the imaging sequence (about 2 minutes) was considered sufficient for reasonable sampling

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of the temporal evolution of the redistribution of the macrogadolinium. The observed

differences in T1 between the two groups clearly demonstrate a transient change of the

interaction between the macrogadolinium and the hepatic tissue related to cavitation.

However, the precise local distribution of the macrogadolinium was not investigated in the

present study, since histological visualization of the selected reporter macromolecule was not

possible. Quantification would have been possible with the help of mass spectroscopy [29].

However, this would require a perfusion of the ex vivo livers to wash out the remaining

Vistarem in the vasculature. In addition, this technique is limited by local sampling of the

tissue and was not compatible with temporal monitoring due to its invasive nature. Recent

developments of multi-modality contrast agents, such as magnetofluorescent nanoparticles

[30], offer possibilities to correlate in vivo MR data with histological or in vivo optical

analysis. Developments of multi-modality probes and optical imaging devices (fluorescence

tomography, Cell~vizio™ (Mauna Kea Technolgies)) dedicated to small animal should help

in analyzing the nature of the influence of cavitation effect for increased local drug delivery.

The analysis of the precise mechanisms responsible for redistribution of the macrogadolinium

between tissue and vasculature was out of the scope of the present work. The selected reporter

macromolecule does not strictly mimic a real drug nor a nano drug carrier. However, the

proposed method may remain applicable in the presence of such molecules/vehicles and

allows for online visualization of the influence of cavitation on the change of interaction

between hepatic tissue and Vistarem.

In the present work, only two different sonication conditions were compared (with and

without cavitation), but no optimization of the ultrasound parameters was performed. In

addition, since the US device was not MR compatible, no direct monitoring of the changes

during sonication could be performed. This limitation could be overcome by the use of

dedicated US imaging device, as recently reported [31]. Further, no spatial differences were

observed between the different parts of the liver due to the size of the US transducer. In the

perspective of local drug delivery, the precise control of the localization of the cavitation

could be performed by the use of a MR compatible focused ultrasound transducer [32]. The

dimensions of the focal beam are typically a few millimeters, depending on the transducer

design and operating frequency. Therefore, the spatial resolution achieved with the Look-

Locker sequence (1 mm2) should allow for direct visualization of the changes of the T1 values

during focused ultrasound sonications. However, these apparatus do not usually include

ultrasonic imaging capabilities and therefore do not allow for a direct visualization of

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microbubbles transient passage in the targeted organ, as in our case. The absence of

monitoring of the T1 changes during sonication was not considered problematic in the present

work, since it lasted less than 4 minutes whereas the time course of the observed phenomenon

was about 1 hour.

The proposed monitoring method may help in optimizing the sonication protocol (amplitude,

pulse repetition frequency, sonication duration and microbubble concentration), as a function

of the targeted organ, since the local quantity of US contrast agent is perfusion dependent.

Although the data shown here were obtained on the liver, this method may be applied to any

soft tissue that can be imaged with MRI.

In the present work, the imaging devices and the injected molecules were clinically approved.

However, Sonovue has received the agreement for clinical usage only as a diagnostic contrast

agent, but not for the purpose of increased cavitational purposes. The proposed method may

help in direct visualization of the enhanced drug delivery by cavitation in humans without

inducing unwanted irreversible damages.

11.5. References

[1] Tredan O, Galmarini CM, Patel K, Tannock IF. Drug resistance and the solid tumor

microenvironment. J Natl Cancer Inst 2007;99(19):1441-1454.

[2] Bao S, Thrall BD, Miller DL. Transfection of a reporter plasmid into cultured cells by

sonoporation in vitro. Ultrasound Med Biol 1997;23(6):953-959.

[3] Greenleaf WJ, Bolander ME, Sarkar G, Goldring MB, Greenleaf JF. Artificial

cavitation nuclei significantly enhance acoustically induced cell transfection.

Ultrasound Med Biol 1998;24(4):587-595.

[4] Lawrie A, Brisken AF, Francis SE, Tayler DI, Chamberlain J, Crossman DC,

Cumberland DC, Newman CM. Ultrasound enhances reporter gene expression after

transfection of vascular cells in vitro. Circulation 1999;99(20):2617-2620.

[5] Lu QL, Liang HD, Partridge T, Blomley MJ. Microbubble ultrasound improves the

efficiency of gene transduction in skeletal muscle in vivo with reduced tissue damage.

Gene Ther 2003;10(5):396-405.

[6] Taniyama Y, Tachibana K, Hiraoka K, Aoki M, Yamamoto S, Matsumoto K,

Nakamura T, Ogihara T, Kaneda Y, Morishita R. Development of safe and efficient

novel nonviral gene transfer using ultrasound: enhancement of transfection efficiency

of naked plasmid DNA in skeletal muscle. Gene Ther 2002;9(6):372-380.

136

[7] Price RJ, Skyba DM, Kaul S, Skalak TC. Delivery of colloidal particles and red blood

cells to tissue through microvessel ruptures created by targeted microbubble

destruction with ultrasound. Circulation 1998;98(13):1264-1267.

[8] Song J, Chappell JC, Qi M, VanGieson EJ, Kaul S, Price RJ. Influence of injection

site, microvascular pressure and ultrasound variables on microbubble-mediated

delivery of microspheres to muscle. J Am Coll Cardiol 2002;39(4):726-731.

[9] Skyba DM, Price RJ, Linka AZ, Skalak TC, Kaul S. Direct in vivo visualization of

intravascular destruction of microbubbles by ultrasound and its local effects on tissue.

Circulation 1998;98(4):290-293.

[10] Dalecki D, Child SZ, Raeman CH, Cox C, Carstensen EL. Ultrasonically induced lung

hemorrhage in young swine. Ultrasound Med Biol 1997;23(5):777-781.

[11] Feril LB, Jr., Kondo T, Zhao QL, Ogawa R, Tachibana K, Kudo N, Fujimoto S,

Nakamura S. Enhancement of ultrasound-induced apoptosis and cell lysis by echo-

contrast agents. Ultrasound Med Biol 2003;29(2):331-337.

[12] Dalecki D, Raeman CH, Child SZ, Cox C, Francis CW, Meltzer RS, Carstensen EL.

Hemolysis in vivo from exposure to pulsed ultrasound. Ultrasound Med Biol

1997;23(2):307-313.

[13] Cline HE, Hynynen K, Watkins RD, Adams WJ, Schenck JF, Ettinger RH, Freund

WR, Vetro JP, Jolesz FA. Focused US system for MR imaging-guided tumor ablation.

Radiology 1995;194(3):731-737.

[14] Deckers R, Rome C, Moonen CT. The role of ultrasound and magnetic resonance in

local drug delivery. J Magn Reson Imaging 2008;27(2):400-409.

[15] Frenkel V. Ultrasound mediated delivery of drugs and genes to solid tumors. Adv

Drug Deliv Rev 2008;60(10):1193-1208.

[16] Taniyama Y, Tachibana K, Hiraoka K, Namba T, Yamasaki K, Hashiya N, Aoki M,

Ogihara T, Yasufumi K, Morishita R. Local delivery of plasmid DNA into rat carotid

artery using ultrasound. Circulation 2002;105(10):1233-1239.

[17] Rahim A, Taylor SL, Bush NL, ter Haar GR, Bamber JC, Porter CD. Physical

parameters affecting ultrasound/microbubble-mediated gene delivery efficiency in

vitro. Ultrasound Med Biol 2006;32(8):1269-1279.

[18] Zarnitsyn VG, Prausnitz MR. Physical parameters influencing optimization of

ultrasound-mediated DNA transfection. Ultrasound Med Biol 2004;30(4):527-538.

137

[19] Hallow DM, Mahajan AD, Prausnitz MR. Ultrasonically targeted delivery into

endothelial and smooth muscle cells in ex vivo arteries. J Control Release

2007;118(3):285-293.

[20] Larkin JO, Casey GD, Tangney M, Cashman J, Collins CG, Soden DM, O'Sullivan

GC. Effective tumor treatment using optimized ultrasound-mediated delivery of

bleomycin. Ultrasound Med Biol 2008;34(3):406-413.

[21] Iwanaga K, Tominaga K, Yamamoto K, Habu M, Maeda H, Akifusa S, Tsujisawa T,

Okinaga T, Fukuda J, Nishihara T. Local delivery system of cytotoxic agents to

tumors by focused sonoporation. Cancer Gene Ther 2007;14(4):354-363.

[22] Bekeredjian R, Kroll RD, Fein E, Tinkov S, Coester C, Winter G, Katus HA, Kulaksiz

H. Ultrasound targeted microbubble destruction increases capillary permeability in

hepatomas. Ultrasound Med Biol 2007;33(10):1592-1598.

[23] Treat LH, McDannold N, Vykhodtseva N, Zhang Y, Tam K, Hynynen K. Targeted

delivery of doxorubicin to the rat brain at therapeutic levels using MRI-guided focused

ultrasound. Int J Cancer 2007;121(4):901-907.

[24] Port M, Corot C, Raynal I, Idee JM, Dencausse A, Lancelot E, Meyer D, Bonnemain

B, Lautrou J. Physicochemical and biological evaluation of P792, a rapid-clearance

blood-pool agent for magnetic resonance imaging. Invest Radiol 2001;36(8):445-454.

[25] Greis C. Technology overview: SonoVue (Bracco, Milan). Eur Radiol 2004;14 Suppl

8:P11-15.

[26] Morel DR, Schwieger I, Hohn L, Terrettaz J, Llull JB, Cornioley YA, Schneider M.

Human pharmacokinetics and safety evaluation of SonoVue, a new contrast agent for

ultrasound imaging. Invest Radiol 2000;35(1):80-85.

[27] Port M, Corot C, Rousseaux O, Raynal I, Devoldere L, Idee JM, Dencausse A, Le

Greneur S, Simonot C, Meyer D. P792: a rapid clearance blood pool agent for

magnetic resonance imaging: preliminary results. Magma 2001;12(2-3):121-127.

[28] Corot C, Violas X, Robert P, Gagneur G, Port M. Comparison of different types of

blood pool agents (P792, MS325, USPIO) in a rabbit MR angiography-like protocol.

Invest Radiol 2003;38(6):311-319.

[29] Frame EM, Uzgiris EE. Gadolinium determination in tissue samples by inductively

coupled plasma mass spectrometry and inductively coupled plasma atomic emission

spectrometry in evaluation of the action of magnetic resonance imaging contrast

agents. Analyst 1998;123(4):675-679.

138

[30] Sosnovik DE, Nahrendorf M, Deliolanis N, Novikov M, Aikawa E, Josephson L,

Rosenzweig A, Weissleder R, Ntziachristos V. Fluorescence tomography and

magnetic resonance imaging of myocardial macrophage infiltration in infarcted

myocardium in vivo. Circulation 2007;115(11):1384-1391.

[31] Curiel L, Chopra R, Hynynen K. Progress in multimodality imaging: truly

simultaneous ultrasound and magnetic resonance imaging. IEEE Trans Med Imaging

2007;26(12):1740-1746.

[32] McDannold N, Vykhodtseva N, Hynynen K. Effects of acoustic parameters and

ultrasound contrast agent dose on focused-ultrasound induced blood-brain barrier

disruption. Ultrasound Med Biol 2008;34(6):930-937.

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Part IV. Summaries and perspectives

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Summary

This thesis describes the use of molecular imaging technologies for molecular therapy

applications such as local gene activation and local drug delivery. Molecular imaging

techniques can be used for early detection and characterization of diseases, guidance of

therapy and quantification of induced therapeutic effect. This thesis starts with an introduction

of the main techniques used for research in this work. In the second part feasibility studies are

described to investigate the use of MRI guided HIFU in combination with a heat sensitive

promoter to achieve spatial and temporal control of transgene expression. Bioluminescence

imaging allowed for non-invasive monitoring of local transgene activation. Local deposition

of energy with ultrasound may also facilitate the delivery of drugs to areas in need of therapy.

Molecular imaging techniques allow for real-time monitoring of the fate of the drugs. This

research is described in the third part.

Part I: Introduction of the main techniques used

In Chapter 1 the interaction of ultrasound with tissue is described. In Chapter 3 the use of

MRI thermometry for controlling local hyperthermia is described. The different optical

imaging techniques for monitoring local gene activation and drug delivery are explained in

Chapter 4.

Part II : Spatio-temporal control of gene activation

Local hyperthermia in combination with a heat sensitive heat shock protein (Hsp) promoter

allows for spatio-temporal control of transgene expression. The time course and location of

transgene expressiom is assessed with help of an optical reporter gene (firefly luciferase)

placed under control of the Hsp promoter. Hsp promoters, particularly Hsp70 promoters, have

a couple of characteristics that make them very suitable for gene therapy. The in vitro

characterization of the Hsp70 promoter in Chapter 7 showed that the Hsp70 promoter has a

low basal activity and can attain high heat-induced activity, up to 53 fold compared to basal

activity. Furthermore, it was shown that the magnitude of promoter activity can be modulated

by temperature and duration of hyperthermia. Promoter activity was assessed with respect to

different temperatures and durations of hyperthermia. It was demonstrated that the promoter

activity followed an Arrhenius relationship. Temperature increase of 1º C with a constant

exposure time resulted in a 2-fold increase of luciferase activity. A similar increase in

luciferase activity was observed after doubling the exposure time at constant temperature.

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In Chapter 8 a transgenic mouse model is used to evaluate the in vivo kinetics of the Hsp70

promoter activity with respect to temperature and duration of hyperthermia. This transgenic

mouse model, NLF-1, contains a transgene that allows firefly luciferase expression under the

control of the Hsp70 promoter. The maximal luciferase activity was found 4 hours after

heating, independently of the applied heating protocol. It was demonstrated that the in vivo

promoter activity also followed an Arrhenius relationship. The response of the promoter

activity to a temperature increase of 1º C or a doubling of the exposure time was similar in

vivo and in vitro. In the clinical environment multiple sequential heatings might be necessary

in order to obtain a prolonged treatment, due to the transient character of the promoter. The in

vivo mouse model allowed investigating the time course of the promoter activity after

multiple sequential heatings. A second heating resulted in significant increase of luciferase

activity, whereas after a third heating the emission returned to original light emission levels.

The origin of varying luciferase activity after multiple sequential heatings is unknown.

In Chapter 9 MRI guided High Intensity Focused Ultrasound (HIFU) and Bioluminescence

Imaging (BLI) were used in the transgenic NLF-1 mouse to show a high similarity between

the local temperature distribution in vivo and the region emitting light. Control of transgene

expression was achieved by automatic adjustment of the HIFU power based on continuous

MRI thermometry to force the temperature to follow a predefined temperature evolution. The

good spatial correspondence between increased temperature and gene expression was

demonstrated by comparing MRI temperature maps and bioluminescence images. BLI

demonstrated also to be a reliable method for analyzing the kinetics of gene activation. Mild

heating protocols (i.e. 2 minutes at 43°C) lead to significant amplification of gene expression

without inducing tissue damage. However, increasing the duration of the hyperthermia

resulted in an increase of induced damage in the leg muscles. This illustrates the importance

of the choice of the heating conditions in vivo to observe sufficient induction of expression

avoiding tissue damage. Furthermore, this justifies the fundamental importance of a precise,

non-invasive and quantitative temperature measurement and control system provided by the

combination of MRI thermometry and HIFU heating.

Part III: Local drug delivery

In Chapter 10 the role of ultrasound and molecular imaging technologies in local drug

delivery are explained. The interaction of ultrasound with tissue may improve the deposition

of drug (carriers) in the region in need of therapy. Hyperthermia, cavitation and radiation

force are the underlying physical mechanism that create bio-effects favorable for drug

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delivery. The in vivo distribution and metabolism of a drug and its effect on a living system

can be monitored with a large spectrum of non-invasive imaging technologies. Each

technique has specific advantages and limitations for monitoring drug delivery. Therefore, the

use of hybrid imaging systems is promising.

In Chapter 11 a method is presented for monitoring in vivo with MRI the changes of

distribution of a reporter macromolecule in hepatic tissue induced by cavitation. The temporal

and spatial resolution of MRI was sufficient to follow transient changes in T1 values of

hepatic tissue. Injection of macromolecular contrast agent and microbubbles followed by

ultrasound resulted in lower T1 values as compared to the control experiment without

ultrasound. The proposed method may help in direct visualization of the enhanced drug

delivery by cavitation and in optimizing the sonication protocol for delivery.

Key words: MRI, focused ultrasound, local transgene activation, optical imaging, local drug

delivery

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Perspectives

This study showed the opportunities of the non-invasive HIFU technique for controlling, via

local hyperthermia, the location and level of transgene activation. MRI thermometry was

shown to be essential for monitoring the local temperature distribution and controlling HIFU

power output. The next step will be the application of spatio-temporal activation under control

MR guided HIFU of a therapeutic gene in a disease model. The therapeutic gene HSV-tk

would be a good first candidate, because the principle of activating HSV-tk with

hyperthermia, but without HIFU, has already been demonstrated by other groups.

The local deposition of thermal energy with HIFU may not only be useful in gene therapy, but

also in cell-based therapies. In cell-based therapies stem cells are induced to differentiate into

the specific cell type required to repair damaged or destroyed cells or tissues. Stem cells,

directed to differentiate into specific cell types, offer the possibility of a renewable source of

replacement cells and tissues to treat diseases including Parkinson's and Alzheimer's diseases,

spinal cord injury, stroke, burns, heart disease, diabetes, osteoarthritis, and rheumatoid

arthritis. However, the efficacy of stem cell therapies depends on the introduced cells arriving

where they are needed and differentiating into specific cell types needed for replacing or

rejuvenating damaged cells. MRI guided HIFU may play an important role in the spatio-

temporal control of inducing stem cell differentiation and thereby improve the efficacy and

safety of cell-based therapies.

The perspectives of MR guided HIFU in combination with molecular imaging techniques for

local drug delivery seems to be extraordinary. The large choice in drug carriers, targeting

methods and applications of ultrasound allow the development of drug delivery methods for

each individual case, and thus lead to drastically altered pharmacokinetics and

pharmacodistribution. The drug delivery methods that will enter the clinic first may be based

on combinations of existing techniques and formulations already approved for clinical use.

For example doxorubicin filled liposomes (Doxil) that are passively targeted to tumor site

where therapeutic ultrasound improves the release of the drug out of the carrier and reduces

the physiological barriers. Active targeting of acoustic active carriers such as microbubbles to

diseased sites in combination with HIFU may further improve the efficacy of the treatment

and lower the systemic toxicity.

Furthermore, the non-invasive imaging of drug distribution can help to quantify delivery and

allow the prediction of treatment efficacy based on the distribution and quantity of the

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delivered drug. The large library of reporter molecules developed for molecular imaging

applications using different imaging modalities may also be exploited for drug delivery

monitoring. However, such molecular image probes are in general based on covalent binding

to the molecule of interest which is, at present, a strategy not envisioned for drug delivery

(except with regard to PET labelled drugs). Co-injection of image probes with similar in vivo

behaviour as the drug allows for monitoring drug delivery without changing the properties of

the drug. Further, the development of multi-modality imaging methods is promising because

it allows simultaneous monitoring of the drug carrier as well as the drug itself.

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Résumé

Cette thèse décrit l’utilisation des techniques d’imagerie moléculaire pour des thérapies

moléculaires telles que l’activation génique locale ou le dépôt local de médicaments. Les

techniques d’imagerie moléculaire peuvent être utilisées pour une détection et une

caractérisation précoce de maladies, le guidage d’une thérapie, et la quantification des effets

thérapeutiques induits par un traitement. Cette thèse commence avec une introduction des

principales techniques utilisées dans ce travail. Dans la seconde partie de ce manuscrit, une

étude de faisabilité a été menée pour mettre en évidence le rôle des ultrasons focalisés guidés

par IRM combinés à un promoteur thermosensible pour permettre le contrôle spatio-temporel

d’une expression transgène. L’imagerie par bioluminescence permet un contrôle non-invasif

de l’expression transgène. Le dépôt local d’énergie avec les ultrasons peut aussi faciliter la

libération de molécules thérapeutiques dans des régions nécessitant un traitement. Les

techniques d’imagerie moléculaire permettent le monitorage temps réel du suivi du

médicament. Ce point est abordé dans la troisième partie de cette thèse. Ce manuscrit

commence par un état de l’art des principales techniques utilisées dans cette thématique de

recherche.

Partie I: Introduction des principales techniques utilisées

Dans le chapitre 2, la description de l’interaction entre les ultrasons et les tissus est faite. Dans

le chapitre 3, le rôle de la thermométrie par IRM dans le contrôle d’une hyperthermie locale

est explicité. Les différentes techniques d’imagerie optique permettant un monitorage en ligne

d’une activation génique et le monitorage en ligne d’un dépôt local de médicament sont

détaillées dans le chapitre 4.

Partie II: Contrôle spatio-temporel de l’activation génique

La combinaison de l’hyperthermie locale avec un promoteur thermosensible (Hsp) permet un

contrôle spatio-temporel de l’expression transgène. La cinétique et la localisation de

l’expression transgène est faite à l’aide d’un gène optique reporter (généralement la

luciférase) placé sous contrôle d’un promoteur Hsp. Les promoteurs Hsp, particulièrement

Hsp70, ont des caractéristiques appropriées pour la thérapie génique. La caractérisation in

vitro du promoteur Hsp70 dans le chapitre 7 montre que ce promoteur comporte une activité

de base faible et peut atteindre une activité élevée sous l’effet de la chaleur, avec une

amélioration d’un facteur 53 par rapport à l’activité normale. De plus, il a été montré que

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l’amplitude de l’activité du promoteur peut être modulée grâce à la température appliquée et

la durée effective de l’hyperthermie. L’activité du promoteur a été testée pour différentes

températures et durées de chauffage. Il a été montré que l’activité du promoteur suit une

relation d’Arrhenius. Une augmentation de température de 1°C avec un temps d’exposition

constant résulte dans une augmentation d’un facteur 2 de l’activité de la luciférase. Une

amélioration similaire a été observée si l’on double le temps de chauffage pour une

température constante.

Dans le chapitre 8, un modèle de souris transgénique est utilise pour évaluer la cinétique in

vivo de l’activité du promoteur Hsp70 en fonction de la température et de la durée d’une

hyperthermie. Ce modèle de souris transgénique, NLF-1, possède un transgène qui permet

l’expression de la luciférase sous le contrôle du promoteur Hsp70. Le pic maximal d’activité

de la luciférase a été trouvé 4 heures après chauffage, indépendamment du protocole

d’hyperthermie utilisé. Il a été démontré que l’activité du promoteur in vivo suit également

une relation d’Arrhenius. La réponse du promoteur à une augmentation de temperature de 1°C

ou à un temps de chauffage doublé est similaire pour une expérience in vivo et pour une

expérience in vitro. Dans le cas d’applications cliniques, plusieurs séquences de chauffage

peuvent être nécessaires pour obtenir un traitement efficace ceci à cause du caractère

transitoire du promoteur. Le modèle de souris in vivo permet l’étude de l’activité du

promoteur après plusieurs séquences d’hyperthermie. Une seconde hyperthermie résulte dans

une amélioration significative de l’activité de la luciférase, alors qu’après un troisième

chauffage, l’émission lumineuse revient aux niveaux d’émissions lumineuses originaux.

L’origine de cette variation d’activité de la luciférase après plusieurs chauffages consécutifs

est inconnue.

Dans le chapitre 9, les ultrasons focalisés par IRM (MRgHIFU) et l’imagerie par

bioluminescence (BLI) sont utilisés sur une souris transgénique NLF-1 pour montrer la

correspondance entre une élévation locale de température in vivo et une émission locale de

lumière. Le contrôle de l’expression transgène a été effectué par un ajustement automatique

de la puissance des ultrasons basé sur le monitorage en ligne de la température par IRM afin

de forcer la température à suivre une consigne d’évolution prédéfinie. La correspondance

spatiale entre une élévation de température et l’expression du gène a été démontrée en

comparant les cartes de températures avec les images de bioluminescences. Il a aussi été

démontré que les BLI peuvent être un outil fiable pour analyser la cinétique de l’activation

d’un gène. Des protocoles d’hyperthermie légers (i.e. 2 minutes à 43°C) produisent des

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amplifications significatives de l’expression d’un gène sans pour autant induire des

dommages sur les tissus environnant. Malgré tout, l’augmentation de la durée de

l’hyperthermie résulte dans une augmentation des dommages tissulaires induit dans les

muscles. Ce point illustre l’importance du choix des paramètres de chauffage in vivo

permettant l’observation de l’expression du gène en évitant les dommages tissulaires. De plus,

cela justifie le rôle fondamentale de la méthode précise, non-invasive et quantitative de

mesure de température fournit par la combinaison de la thermométrie par IRM et d’une

hyperthermie appliquée à l’aide des ultrasons focalisés.

Partie III: Dépôt local de médicaments

Dans le chapitre 10, le rôle des ultrasons et des techniques d’imagerie moléculaire dans le

dépôt local de médicaments est expliqué. L’interaction des ultrasons avec les tissus peut

améliorer le dépôt de principes thérapeutiques dans une région nécessitant un traitement.

L’hyperthermie, la cavitation et la force rayonnant sont les mécanismes physiques sous-

jacents créant des effets biologiques favorables au dépôt de médicament. La distribution et le

métabolisme in vivo d’un médicament et ses effets sur un système vivant peut être monitorée

par un large spectre de techniques d’imagerie non-invasives. Chaque technique présente des

avantages et des limitations spécifiques dans le monitorage du dépôt de médicament.

L’utilisation de systèmes d’imagerie hybrides est prometteuse.

Dans le chapitre 11 la méthode présentée permet le monitorage in vivo par IRM des

changements de distribution par cavitation d’une macromolécule reporter dans un tissu

hépatique. La résolution spatiale et temporelle de l’IRM est suffisante pour suivre les

changements transitoires des valeurs de T1 du tissu hépatique. L’injection d’un agent de

contraste macromoléculaire et de microbulles suivie d’un traitement ultrasonore résulte en des

valeurs de T1 plus faible comparé à l’expérience de contrôle sans ultrasons. La méthode

proposée peut être utile pour la visualisation directe de l’effet du dépôt de médicament par

cavitation en optimisant le protocole de sonication pour la libération des principes

thérapeutiques.

Mots-clés: IRM, ultrasons focalisés, l’activation locale transgénique, imagerie optique, le

dépôt local de médicaments

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Perspectives

Cette étude montre les opportunités offertes par les techniques non-invasives d’ultrasons

focalisés pour contrôler, via l’hyperthermie locale, le lieu et le niveau d’une activation

transgène. La thermométrie par IRM est un élément essentiel dans le monitorage d’une

distribution locale de température et dans l’asservissement de la puissance de la sonde HIFU.

La prochaine étape sera l’application d’une activation spatio-temporelle contrôlée par

MRgHIFU d’un gène thérapeutique sur un modèle de maladie. Le gène thérapeutique HSV-tk

pourrait être un bon candidat pour débuter étant donné que le principe d’activation du gène

HSV-tk par hyperthermie (mais sans ultrasons focalisés) a déjà été démontré par d’autres

groupes de recherche.

Le dépôt local d’énergie thermique avec les ultrasons focalisés peut non seulement être utile

dans la thérapie génique, mais aussi dans les thérapies cellulaires. Dans les thérapies

cellulaires, les cellules souches sont destinées à se différencier dans les types de cellules

spécifiques requis pour réparer les cellules ou les tissus détruits ou endommagés. Les cellules

souches offrent la possibilité d’avoir une source de cellules et de tissus de remplacement afin

de traiter des maladies comme la maladie de Parkinson, d’Alzheimer, des blessures de la

moelle épinière, des caillots sanguins, des brulures, des maladies du cœur, du diabète, et de

l’arthrite rhumatoïde. L’efficacité des thérapies basées sur les cellules souches repose sur le

fait que les cellules introduites doivent arriver à l’endroit où les cellules endommagées ont

besoin d’être réparées et aussi sur le fait qu’elles doivent se différencier dans les bons types

cellulaires. Les ultrasons focalisés guidés par IRM peuvent jouer un rôle important dans le

contrôle spatio-temporel et le déclenchement de la différenciation et donc améliorer

l’efficacité et la sécurité des thérapies cellulaires.

Les perspectives liées à l’utilisation des ultrasons focalisés guides par IRM combinés avec les

techniques d’imagerie moléculaire semblent être extraordinaire. Le large choix dans les

transporteurs de médicaments, les méthodes de ciblage, et les applications d’ultrasons permet

le développement de méthodes de dépôt de médicaments spécifiques à chaque cas rencontré,

et conduit à changer considérablement les cinétiques et les modes de distributions

pharmacologiques. La méthode de dépôt de médicament qui sera utilisée dans un premier

temps pourra être basée sur la combinaison de techniques existantes et de formulations déjà

approuvées pour les pratiques cliniques. On peut citer par exemple, l’utilisation de liposomes

porteurs de doxorubicine (Doxil) ciblés passivement sur le lieu d’une tumeur et où

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l’utilisation des ultrasons améliore la libération du principe thérapeutique et réduit l’efficacité

des barrières physiologiques permettant ainsi un traitement plus efficace. Un ciblage actif du

lieu à traiter de transporteurs sensibles aux ondes ultrasonores tels que les microbulles,

combiné aux ultrasons focalisés peut améliorer l’efficacité du traitement et réduire la toxicité

systémique.

De plus, l’imagerie non-invasive de la distribution de médicament peut aider à quantifier la

déposition du principe thérapeutique et permet la prédiction de l’efficacité du traitement à

partir de la distribution et de la quantité de médicament délivrée. La grande variété de

molécules reporters développées pour l’imagerie moléculaire utilisant différentes modalités

d’imagerie peut aussi être exploitée pour monitorer le dépôt local de médicament. En

revanche, ces molécules reporters sont en général basées sur un couplage covalent avec la

molécule thérapeutique, ce qui à l’heure actuelle n’est pas la stratégie souhaitée à long terme

pour la déposition locale de médicament (excepté pour les médicaments associés à la TEP).

L’injection simultanée de molécules reporters ayant des comportements in vivo similaires aux

médicaments utilisés permet le suivi de la libération du principe thérapeutique sans en

changer les propriétés. Le développement de l’imagerie multi-modalité est prometteur car il

permet le suivi simultané du transporteur de médicament et du médicament lui-même.

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Part V. Word of thanks and publications

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Word of thanks

Time flies when you are having fun. I still remember the day that I took the train from

Maastricht in the direction of Bordeaux like it was yesterday. Five minutes after leaving the

railway station we crossed the border with Belgium and from that point on the people that

entered the train spoke a language I didn't understand: French. My big adventure had started!

First I would like to thank the members of the jury for reading my thesis in such detail and

making critical comments on it. Mickaël Tanter and Bertand Tavitian, I was honored to have

you as my ‘rapporteurs’ and thank you for coming to Bordeaux just before Christmas. Pierre

Voisin, Alain Brisson, Wilbert Bartels and Chrit Moonen, thank you for taking place in my

jury as ‘examinateur’ and making my jury even more multi-disciplinary.

Chrit, of course I don’t want to thank you only for being in my jury, but mainly for giving me

the opportunity to do my PhD thesis in your lab. I think we have discussed a lot of things

(science, politics, sports, wine, the French system) in a lot of languages (French, English,

Dutch and last but not least Limburgs). I really enjoyed our morning-coffee discussions and I

am happy that we can continue to do this in Utrecht.

During my thesis Chrit (and sometimes even I) had the most beautiful ideas for new research.

Luckily Franck and Bruno you were also in the lab to put these beautiful ideas in a more

down-to-earth perspective and to help me on a day-to-day basis to realize my objectives.

Mario, although we didn’t work on the same research projects, our routes crossed often for

example during the morning-coffees and when my computer didn’t work. Furthermore, I

would like to thank you for giving me a place to stay during the first weeks in Bordeaux and

of course for supporting my IKEA addiction when I found an apartment.

Greg, pour montrer quel bon prof de français tu es, je te remercie en français. C’est toi qui

m’a appris les mots importants pour une conversation de haut niveau : ‘hi

coquine…remorque…femme à lunette, femme à…’. Tu étais mon collègue, mais surtout mon

pote et j’espère qu’on va avoir encore beaucoup de moments où on peut parler de foot, des

filles et de science.

Yasmina, heureusement tu étais là, avec ta touche féminin pour corriger notre instinct animal

et ta compagnie joyeuse au boulot et dans les sortis.

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Nora, Pierre-Yves, Christelle, Anna and Coralie I enjoyed working with you on the different

projects. Good luck with becoming radiologist, your thesis, your post-doc and your

‘concours’.

Josette, heureusement tu n’es pas partie en retraite avant la fin de ma thèse. Je te remercie de

m’avoir expliqué plein choses en biologie, et surtout de ta patience vu mon niveau de français

(au début).

Pierre and Hélène thank you for taking care of all my sweet little mice. Without them and thus

without you there wouldn’t have been a thesis.

Charles, thank you for always having the correct answers to all my ultrasound related

questions. I hope you will continue doing this for me in Utrecht.

Colette, Marie-France, Matthieu, Silke, Xenia, Baudouin, Mathilde, Sebastien, Bixente,

Gwenaelle, Isabelle, Olivier, Marion, Christophe, Hugues, Hervé and Nicolas I enjoyed

working in the same lab. J’espère que vous n’êtes pas trop triste qu’il y a plus un néerlandais

qui tartine ses sandwiches dans la cuisine?

Luis, Iulius, Claire, Cedric, Marianne, Philippe, Vincent, Sander, Omer, Charles and Thibault

you are, just like me, old members of the IMF lab. Be proud of it and good luck in your

career!

En tout cas je ne vais pas oublier mon séjour en France car j’ai emmené un grand souvenir

aux Pays-Bas. Sophie, merci de m’avoir soutenu pendant ma dernière année de thèse et d’être

venu avec moi à Utrecht.

En natuurlijk, pap en mam, dank jullie wel voor jullie onvoorwaardelijke steun. Jullie

luisterend oor en goede adviezen tijdens de vele telefoongesprekken waren een prima

klankbord voor alle frustaties en blijdschappen die gepaard gaan met een promotie. Dat er nu

een mooi boekje ligt en dat ik nu de mensen aan de andere kant van de grens van Maastricht

ook versta is ook zeker aan jullie te danken.

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List of publications

Papers

Frulio N, Trillaud H, Deckers R, Lepreux S, Corot C, Moonen C and Quesson B, MRI

Monitoring of the Inflence of Microbubbles Destruction by Ultrasound for Local Delivery of

a Macromolecular MRI Contrast Agent in the Rat Liver. (Ready for submission)

Deckers R, Quesson B, Arsaut J, Eimer S, Couillaud F and Moonen C, Non-invasive spatio-

temporal control of gene expression . Proc. Natl. Acad. Sci. U S A. 106(4), 1175-1180 (2009)

Deckers R, Rome C, and Moonen C, The role of ultrasound and magnetic resonance in local

drug delivery. Journal of Magnetic Resonance Imaging. 27, 400-409 (2008)

Peer-reviewed abstracts

Deckers R, Couillaud F, Quesson B, Eimer S and Moonen C, Local Control of Transgene

Expression Using MRI Guided HIFU on a Transgenic Mouse. “World Molecular Imaging

Congress, Nice, France”, (2008)

Frulio N, Trillaud H, Deckers R, Lepreux S, Eker O, Corot C, Moonen C and Quesson B,

MRI Monitoring of the Inflence of Microbubbles Destruction by Ultrasound for Local

Delivery of a Macromolecular MRI Contrast Agent in the Rat Liver. “World Molecular

Imaging Congress, Nice, France”, (2008)

Frulio N, Trillaud H, Eker O, Deckers R, Lepreux S, Laurent C, Corot C, Moonen C and

Quesson B, MRI Monitoring of the Influence of US Contrast Agent Destruction for Local

Delivery of a MRI Blood Pool Contrast Agent in the Rat Liver. “ISMRM, 16th scientific

meeting and exhibition, Toronto, Canada”, (2008)

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Deckers R, Quesson B, Couillaud F and Moonen C, Temporal and spatial control of gene

expression in transgenic mice: a combination of MR-guided HIFU and bioluminescence.

“AMI/SMI Joint Molecular Imaging Conference, Providence, Rhode Island, USA”, (2007)

Deckers R, Quesson B, Rome C, Couillaud F and Moonen C, Non-invasive spatial control of

gene activation by local heating with focused ultrasound under MRI temperature guidance.

“ISMRM, 15th scientific meeting and exhibition, Berlin, Germany”, (2007)

Deckers R, Quesson B, Rome C, Couillaud F and Moonen C, Bioluminescence imaging of

local transgenic expression induced by heat in mice. “First International Conference of the

European Society for Molecular Imaging (ESMI), Paris, France”, (2006)