Ultrasonics 57 (2015) 144–152

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MRI compatible head phantom for ultrasound surgery Georgios Menikou a, Tetiana Dadakova c, Matt Pavlina c, Michael Bock c, Christakis Damianou b,⇑[email protected] a

City University, London, UK Cyprus University of Technology, Limassol, Cyprus c University Medical Center Freiburg, Radiology – Medical Physics, Freiburg, Germany b

a r t i c l e

i n f o

Article history: Received 19 June 2014 Received in revised form 29 September 2014 Accepted 9 November 2014 Available online 20 November 2014 Keywords: MRI Ultrasound Brain

a b s t r a c t Objective: Develop a magnetic resonance imaging (MRI) compatible head phantom with acoustic attenuation closely matched to the human attenuation, and suitable for testing focused ultrasound surgery protocols. Materials and methods: Images from an adult brain CT scan were used to segment the skull bone from adjacent cerebral tissue. The segmented model was manufactured in a 3-D printer using (Acrylonitrile Butadiene Styrene) ABS plastic. The cerebral tissue was mimicked by an agar–evaporated milk–silica gel (2% w/v–25% v/v–1.2% w/v) which was molded inside a skull model. Results: The measured attenuation of the ABS skull was 16 dB/cm MHz. The estimated attenuation coefficient of the gel replicating brain tissue was 0.6 dB/cm MHz. The estimated agar–silica gel’s T1 and T2 relaxation times in a 1.5 Tesla magnetic field were 852 ms and 66 ms respectively. The effectiveness of the skull to reduce ultrasonic heating was demonstrated using MRI thermometry. Conclusion: Due to growing interest in using MRI guided focused ultrasound (MRgFUS) for treating brain cancer and its application in sonothrombolysis, the proposed head phantom can be utilized as a very useful tool for evaluating ultrasonic protocols, thus minimizing the need for animal models and cadavers. Ó 2014 Elsevier B.V. All rights reserved.

1. Introduction The thermal effects of delivering large amounts of acoustic energy through a focused ultrasound configuration in biological tissue (in vitro and in vivo) was first observed in 1942 [1]. Later on extensive animal studies [2] with focused ultrasound (FUS), demonstrated the reversibility of induced neurological dysfunction below certain temperature threshold and recognized the important technical and safety issues needed to overcome before employing the FUS in a clinical setting. The absence of imaging modalities capable of tissue temperature monitoring, delayed the development of FUS into a controlled, precise and safe tissue ablating therapy. The introduction of ultrasound and magnetic resonance imaging (MRI) with their fast technological advancement made image-guided focused ultrasound therapy a reality. The unsurpassed soft tissue contrast and its temperature sensitivity set MRI as the optimal imaging modality [3–5]. Currently a commercially integrated MRI guided FUS (MRgFUS) device is approved by the Food and Drugs Administration (FDA) for treating uterine fibroids [6–8] and pain palliation of bone metastases [9]. Clinical trials for treating prostate cancer [10] and breast ⇑ Corresponding author. http://dx.doi.org/10.1016/j.ultras.2014.11.004 0041-624X/Ó 2014 Elsevier B.V. All rights reserved.

cancer [11,12] are in progress. Another important technology is the application of FUS for prostate cancer ablation which utilizes ultrasonic imaging [13,14]. Additionally, Philips Healthcare, Netherland [15] has produced Sonalleve which is an MRI guided robotic system which received CE mark for clinical use in 2009 for the treatment of fibroids. Recently a major breakthrough of MRgFUS was the announcement of the positive results of a pilot clinical trial for treating essential tremor [16]. This study shows that if thermal lesions are produced in a specific location of the Thalamus, then the symptoms of the essential tremor are significantly reduced. Finally, lately there is increasing interested for using MRgFUS for pain palliation of patients with bone cancer [17]. The ‘holy grail’ for researchers and manufacturers in the MRgFUS field is of course the application and optimization of the technology in treating brain disorders [18]. The potential of performing noninvasive transcranial surgery to treat disorders of the central nervous system has many advantages and numerous attractive applications. Noninvasive thermocoagulation of brain tumors using focused ultrasound offers an alternative option to traditional surgery, radiotherapy and chemotherapy. Early attempts involved transcranial in vivo animal studies prepared with craniotomy that tested the feasibility of detecting temperature elevations with MRI at sub ablative temperatures [19]. With the introduction of

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phased array systems, animal brain models in ex vivo human skulls were monitored with MR thermometry and focal lesions were correlated with histology [20]. With the phased array technology the skull aberrations can be corrected by using the newly developed technique of time reversal mirror with amplitude correction which was evaluated through excised human skull [21] and in sheep skulls [22]. Transcranial MRgFUS through intact skulls of primates demonstrated temperature distribution in critical structures like skin, skull and brain surface while thermally ablating brain tissue [23–25]. Recently, in a phase I clinical trial patients with recurrent glioblastoma were treated with transcranial MRgFUS and the ultrasound focus was confirmed with MR temperature imaging [26]. Thermal coagulation of the tumor was not reached due to power limitation at the time of study, but extrapolation of the available data proved that ablation was possible with some device modifications while avoiding overheating the brain surface. Several studies have demonstrated that MRgFUS can be the future of functional neurosurgery. Essential tremor [16,27] and Parkinson’s patients [28–30] have been treated safely and effectively with MRgFUS thalamotomy. The improvement in some of the patients was immediate and tremor symptoms were suppressed even after the first year follow up. Patients with chronic neuropathic pain have also been treated and benefitted with considerable pain relief [31]. Cadaveric models have been used to explore the feasibility of treating trigeminal neuropathic pain along with an in vitro gel phantom encased in a human skull fitted with multiple thermocouples to examine the temperature changes in skull base while targeting the trigeminal nerve region [32]. MRgFUS has also been tested in the temporal disruption of the blood brain barrier (BBB) in animal models [33–34]. In vivo small animal and non-human primates’ studies showed that the disruption was possible at a range of frequencies allowing deep brain penetration and focus formation with minor adverse effects to cerebral tissue (minor extravasation) and without functional deficits [35,36]. Focused ultrasound has been applied in an attempt to treat efficiently Alzheimer’s neurodegenerative disease in mouse models [37,38] by temporarily disrupting the BBB and allowing large molecules of diagnostic and therapeutic agents to extravasate in the brain parenchyma. The introduction of ultrasound microbubbles agent during low intensity sonication proved in histological assessments to reduce the number of subjects experiencing petechiae compared to subjects treated at higher peak negative pressure fields [37]. The opening of the barrier was monitored with MRI and was confirmed in vivo with gadolinium-based, contrastenhanced MR sequences. Sonothrombolysis is another application of therapeutic ultrasound that receives interest. In vitro studies of human blood clots treated with pulsed FUS in combination with a thrombolytic agent known as recombinant tissue plasminogen activator (rtPA) proved to enhance thrombolysis rate compared to rtPA treatment alone [39]. Similarly in another in vitro study, enhanced lytic treatment efficacy for both tPA and liposomes loaded tPA with the parallel application of a 120 kHz unfocused ultrasound has been demonstrated [40]. The possibility of destructing transcranially an in vitro human clot using a hemispheric phased array FUS transducer in the absence of a thrombolytic agent, has been demonstrated [41]. In an in vivo rabbit aorta [42] and rabbit ear marginal vein [43] model studies, clots were treated with pulsed ultrasound with the synergy of rtPA whereas in other similar in vivo studies safe temperature limits and efficacy of treatment protocols was investigated [44,45]. Currently clinical trial using sonothrombolysis are underway such as the CLOTBUST (Combined Lysis of Thrombus in Brain ischemia using transcranial Ultrasound

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and Systemic Recombinant Tissue-Type Plasminogen Activator rt-PA) [46], the interventional application of ultrasound in treating stroke patients [47], and the transcranial low-frequency ultrasound-mediated thrombolysis in brain ischemia (TRUMBI) [48]. The (TRUMBI) trial that uses non-focused, low-frequency (300 kHz) ultrasound showed an increased hemorrhage risk as compared to tPA alone. A recent simulation study [49] demonstrated that the low frequency used in this trial possibly resulted to unwanted standing waves with pressure that exceeds the threshold of inertial cavitation. This not only could possibly lead to increased hemorrhage risk, but also to elevated temperatures in the brain. Finally a randomized, placebo-controlled study with continuous transcranial 2 MHz Doppler ultrasound and perflutren-lipid microspheres was performed [50]. This paper describes the development of a human skull head phantom suitable for testing brain tumor ablation and cerebral sonothrombolysis applications. Blood brain barrier studies involve perfusion through reversibly permeable capillaries that are extremely difficult to simulate reliably in a phantom model. Similarly for testing the safety and efficacy of MRgFUS in neurosurgical applications, neurological assessment and feedback from treated patients is necessary. Currently studies in transcranial applications involve either purely cadaveric [32] or small animal [23] models, animal models with ex vivo human skulls [20] or combine ex vivo human skulls with gel brain mimicking phantoms [32]. Different research groups have used agarose based gel [23], polyvinyl alcohol cryogel PVA-C [51,52] and polyacrylamide hydrogel [53,54] to fabricate brain tissue mimicking phantoms. Human tissue mimicking gel phantoms suitable for HIFU applications and research are commercially available by Onda Corporation [55]. The proposed phantom design consisted of a skull and a brain section. The most important design requirement for the phantom was that the materials used should match as close as possible the geometry and acoustic attenuation coefficient of the replicated biological tissue. The target attenuation coefficients were extracted from relevant literature. Fry and Barger [56] presented measurements made in the frequency range of 0.25–6 MHz in a series of fresh and subsequently formalin immersed human skulls. The measured insertion loss for a typical adult parietal or temporal skull bone varied from 13 to 24 dB/cm at 1 MHz. Another study showed the collected data for attenuation coefficient of human skull obtained from measurements on seven fresh human calvaria [57]. The estimated attenuation was 2.94(±0.66), 20.06(±1,30), 29.01(±3.65) dB/cm for the frequencies of 270, 836, and 1402 kHz, respectively. In another study by Pinton et al. [58], the experimentally measured attenuation across 8 mm thick skull samples was 13.3 ± 0.97 dB/cm at 1 MHz. A study of the acoustic properties of mammalian brain tissue reported that at 1 MHz the brain absorption and attenuation coefficients measured were 0.2 dB/cm and 0.6 dB/cm respectively [59]. The ratio of absorption to attenuation coefficient was the largest when compared to other types of mammalian tissue like heart, kidney, liver and tendon. Similarly the International Commission on Radiation Units and Measurements [60] reported a 0.6 dB/cm MHz for the acoustic attenuation coefficient of brain tissue. The geometry of the replicated skull was considered as a matter of great importance in order to quantify reliably the defocusing effects from the varying bone thickness of a human skull. Since lesion localization and temperature will be validated in future studies with MRI, the phantom was made out of MR compatible materials and the cerebral tissue phantom possessed T1, T2 relaxation times close to human brain. The phantom is expected to reduce significantly the need of using animals for testing in transcranial MRgFUS. Unlike animal studies, using the same phantom model for testing under perfectly controlled conditions is expected to produce results of increased accuracy and reproducibility.

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2. Materials and methods 2.1. Plastic skull design High resolution Computed Tomography head scan images of an anonymized adult male patient were randomly chosen from a hospital’s database (Limassol General Hospital, Cyprus). The selected set consisted of 219 slices acquired with a CT scanner (Toshiba Aquilion 16, Toshiba Europe GmbH) in axial mode with the following scanning parameters: 120 kV, 322 mA s, 24 cm FOV, 16  0.5 mm collimation, 0.73 mm effective slice thickness, and a 512  512 matrix. Images were processed by an interactive segmentation software (Materialise Mimics 10.0, Leuven, Belgium). Fig. 1 demonstrates how the skull bone was isolated, by manually selecting an appropriate range of indicated densities (200–3000 Hounsfield Units). Unwanted scattered fragments produced during segmentation were removed using a minimum diameter filter in a mesh editing software (MeshLab, Visual Computing Lab, ISTI – CNR, Pisa Italy). Following segmentation a surface rendered model of a human skull was exported to stereo-lithography (STL) format as shown in Fig. 2. The model covered only the supraorbital region. The dimensions of the final skull model were 150  170  83 mm. The STL model was imported in the processing application of a Stratasys 3D printer (FDM400, Eden Prairie, Minnesota, USA). The STL file was used to build a 3D geometry of the segmented skull along with the calculations of the build and support materials. The machine used a common thermoplastic material known as Acrylonitrile Butadiene Styrene (ABS). This raw material satisfied the prerequisite of building an MR compatible phantom since ABS is a non-magnetic material. 2.2. Brain tissue substitute material measurements Agar gel was selected as the base of the brain phantom. Agar is known for having a high melting temperature above 85 °C [61], making it ideal to be used at ablative temperatures without losing its integrity. The agar powder used (Himedia Laboratories, Mumbai, India) was of bacteriological grade type. We have selected to use agar in order to produce a gel of intermediate stiffness and

flexible enough to withstand high intensity ultrasound compression forces without cracking. In order to replicate the acoustic properties of the phantom to brain tissue, a scattering material (Crystalline silica dioxide) was added to control the gel’s scattering coefficient. Crystalline silica dioxide powder (Merck Millipore, Darmstadt, Germany) which is used as a scattering material is insoluble in water and possesses a high melting temperature (1750 °C). Four silica dioxide samples with volume of 100 ml were placed in glazed ceramic containers. Different amounts of silica dioxide powder were added to each sample in order to measure the scattering coefficient of the gel for a preliminary agar concentrations of 2% w/v). The target scattering coefficient was chosen as 0.35 dB/cm MHz based on previously published data for porcine brain at 1 MHz [62]. The scattering coefficient was measured using the methods described by Worthington et al. [63]. The gels were brought slowly to boiling point by running 5–6 consecutive 30 s microwave cycles on full power (800 W) until the agar powder was fully dissolved. The amount of water that evaporated during boiling was replaced. We kept stirring all samples until they started solidifying (35–40 °C) in order to distribute the silica powder evenly within the gel’s matrix. Care was taken to avoid the creation of any bubbles that are known to reflect ultrasound. In order to create a brain phantom with realistic acoustic properties we had to raise the attenuation coefficient to 0.6 dB/cm MHz by loss mechanisms other than scattering. A recipe described by Madsen et al. [64] was tested, where evaporated milk is added to agarose to produce a solid gel of very low scatter tissue mimicking material. In the absence of a scattering material in a homogeneous medium, it is safe to assume that any loss in the acoustic signal is a result of absorption mechanisms. Four 100 ml samples of 10%, 20%, 40% and 50% v/v combined with a 2% w/v agar gel were used to quantify the effect of evaporated milk’s concentration to the gel’s attenuation coefficient. More specifically for each sample 2 g of agarose powder ((Himedia Laboratories, Mumbai, India) was mixed in 75 mL of distilled water. Similarly like in silica experiments, agarose–water mixture was boiled in a microwave oven until agar melted. Then

Fig. 1. Interactive segmentation of skull while discarding surrounding soft tissue.

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Fig. 2. Multi-planar 3D views of the segmented skull model exported in STL format.

25 mL of evaporated milk were filtered through a 540 Whatman filter paper (Whatman International Ltd., Maidstone, UK) using a simple water flow lab Bernoulli vacuum filter. The filtered milk was boiled up to 55 °C. Temperature was monitored using an electronic thermometer. When the temperature of agarose dropped to 55 °C, the warm evaporated milk was added and was well mixed. Each sample was poured in a ceramic glazed container and was left to solidify. Although the recipe by Madsen et al. [64] included a desiccation agent and a fungicide, they were not used for the purposes of this work since the final product was destined for a single use. In order to deduce the attenuation coefficient, the samples were scanned with a pulse-receive 3.6 MHz planar transducer (Piezotechnologies, Indianapolis, USA). The received acoustic signal induced a proportional electric signal across the piezoelectric crystal which was displayed on an oscilloscope. The displayed waveform was made of multiple characteristic peaks each one representing a reflection of the acoustic wave at every interface across its pathway. The first peak with highest amplitude represented the reflection at the transducer-gel interface. The transducer and gel were coupled with ultrasound gel. The second largest peak was the reflection signal from the gel-container bottom interface. The amplitude of this peak was used to calculate the attenuation of the signal from the gel in decibels compared to a reference signal travelling for the same distance in water. It was assumed that the acoustic wave in water was not attenuated. The attenuation coefficient was calculated as the drop of the signal in dB per centimeter of travel in the gel. The coefficient for each sample was then divided by the scanning frequency (3.6 MHz) to obtain the equivalent coefficient in dB/cm MHz. The target was to deduce the evaporated milk concentration for which the gel’s attenuation coefficient was approximately 0.25 dB/cm MHz in order to fulfill the requirement of a total attenuation coefficient of 0.6 dB/cm MHz after the appropriate amount of silica will be added to the final brain phantom.

2.3. Gel thermometry with thermocouples Temperature measurements were needed to confirm that the prescribed gel recipe of agar–evaporated milk–silica gel (2% w/v–25% v/v–1.2% w/v) was capable of absorbing acoustic energy. A gel sample (400 ml) was poured in a plastic container. At the bottom of the container an acoustic absorber was positioned to reduce the reflections of the acoustic wave. A pre-calibrated thermocouple was positioned and held tight once the mixture jellified about 4 cm below the surface of the gel. The gel was then positioned into a plastic water tank. The tank was filled with degassed water. The water level covered the gel container. A spherically focused transducer (Piezotechnologies, Indianapolis, USA) was positioned on a holder that attached to the tank’s walls. The holder allowed

accurate positioning of the transducer in the XYZ axes. The transducer was driven by a signal generator at 1 MHz. The transducer diameter was 40 mm and the radius of curvature was 80 mm. The transducer was driven by an amplifier (RFG 1000, JJA instruments, Duvall, WA 98019, USA). The same transducer and amplifier were used for the MR thermometry experiments (Section 2.5). The method followed to identify the location of the thermocouple was to move the transducer in the horizontal plane until the reading of temperature reached a maximum. The second criterion used to match the beam focus with the thermocouple tip was to observe the rate of temperature drop after the ultrasound beam was turned off. Heat dissipation rate through conduction is maximized at focus site since the temperature difference with the surrounding material is also maximum. The sonifications for localization of the thermocouple were done at minimum input power in order to conserve the gel’s integrity. Temperature readings were recorded for 16 W (acoustical power) using the 1 MHz transducer. With a cross sectional area of focus of 0.78 mm2, the spatial peak temporal average intensity was estimated at 2043 W/cm2. A data acquisition board (6251 DAQ, National Instruments, Texas, USA) was used to measure the temperature in the agar/silica phantom. An analogue input of the board was used to capture the temperature. An Omega (M2813-1205, OMEGA Engineering, INC. Stamford, Connecticut, USA) voltage to temperature converter was used to measure temperature using a software written in MatLab (The Mathworks Inc., Natick, MA). A thermocouple (Omega engineering) was placed in the phantom in order to measure temperature elevation at the focal point. The size of the thermocouple was chosen to be 50 lm, so that the interaction with the beam of ultrasound was minimized. The DAQ station sampled temperature measurements for every 1 s.

2.4. T1 and T2 relaxometry The phantom was scanned with the following sequences on an MRI scanner (Signa Excite HD 1.5 T, GE Healthcare, Milwaukee, USA) using the 8-channel CTL coil (GE Healthcare, Milwaukee, USA) in order to quantify the associated relaxation times of the gel. For measuring the spin–lattice relaxation time (T1), an Inversion Recovery Spin Echo (IR-SE) sequence [65] with the following acquisition parameters were used: Repetition Time (TR) = 5000 ms, Echo Time (TE) = 20 ms, Slice Thickness = 5 mm, Number of Excitations (NEX) = 4, Matrix = 320  160 and variable Inversion Time (TI) = 66, 316, 616, 750 ms. In an IR-SE sequence the initial RF pulse inverts the longitudinal magnetization from its equilibrium state by 180°, with magnitude equal to Mo around the z-axis. Immediately after the application of the inverting pulse, the longitudinal magnetization vector (Mz) points in the opposite direction and relaxes back to equilibrium

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at a rate 1/T1. After a time TI, a second 90° excitation pulse is applied and rotates any longitudinal magnetization that managed to relax in the meantime to the transverse plane. The sequence is repeated at time TR after the initial 180°. Assuming that all pulses are perfect and that field homogeneity is maintained across the sample, the longitudinal magnetization can be expressed as a function of inversion time TI with the following equation.

ð1Þ

Since there is no way of measuring the longitudinal magnetization directly from MRI images an indirect method was used. From Eq. (1) it is obvious that Mz becomes zero when TI = N1/ln2 and therefore no magnetization is flipped with the 90° to the transverse plane. Phantom images were acquired at different inversion times (TI) in order to identify the TI setting for which the signal on magnitude images was nullified. The nulling TI was interpolated by fitting the expression of Eq. (1) to the data and the T1 was calculated. The T2 relaxation time was estimated by taking a series of Fast Spin Echo T2 [65] sequences for different effective echo times (18, 36, 63, 81 and 99 ms). The other parameters used were: TR = 2500 ms, Slice thickness = 3 mm, Matrix = 256x256, FOV = 16 cm, NEX = 1, and ETL = 8. The transverse magnetization was measured in arbitrary units and was then normalized using the factors of the exponential fit described in Eq. (2).

M xy ¼ M o e

TTE 2

ð2Þ

2.5. MRI thermometry Transverse and coronal images were acquired to monitor the ultrasonic beam using Double Echo Segmented Echo planar imaging (EPI). MR thermometry was evaluated in clinical 32-channel 1.5 T MRI system (Magnetom Tim Symphony, Siemens, Germany) using the integrated spine array coils in combination with an anterior 6-channel body array coil for signal reception. In this article only the transverse images are presented. The parameters used for the transverse imaging are: TE = 9 ms, TR = 87 ms, flip angle (FA) = 35°, Bandwidth = 751 Hz/pixel, Slice thickness = 2.2 mm, filed of view (FOV) = 256  256 mm2, Matrix = 125  128. The temperature was estimated using proton resonance frequency (PRF) methods as described by Chung et al. [66]. The acoustical power used during the acquisition of temperature was 10 W for 30 s using the transducer and amplifier indicated in Section 2.3.

Fig. 3. Attenuation coefficient in dB/cm MHz vs. material type (different color, layer thickness and coating method). (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)

3.2. Agar–silica gel attenuation coefficient measurement results Fig. 4A shows scattering coefficient (dB/cm MHz) vs. silica dioxide concentration using 2% agar. Scattering coefficient of 0.35 dB/ cm MHz was achieved using approximately 1.2% of silica dioxide. Since the desired scattering coefficient was established, the next step was to vary the evaporated milk concentration until the desired ultrasonic attenuation is achieved. Fig. 4B shows the measured attenuation coefficient plotted against the associated evaporated milk to agar concentration. Both results were in good agreement with similar measurements found in the literature [64,67]. In our experiments we used a 60–100 lm size range of silica particles whereas the equivalent experiments in the literature [67] were performed with 0.5–50 lm. Larger silica particles have a larger effective scattering area and therefore it was expected and confirmed that the normalized to frequency attenuation coefficient (dB/cm MHz) for the same silica concentration was larger. Since the brain phantom was of homogeneous density, it was considered safe to assume that the attenuation coefficient varied linearly proportional to frequency. The equation of the linearly

A Scaering Coefficient (dB/cm-MHz)

  TI M z ¼ M o 1  2eT1

3. Results

1.0 0.9 0.8 0.7 0.6 0.5 0.4 0.3 0.2 0.1 0.0

y = 0.2827x + 0.0064 R² = 0.998

0

The acoustic attenuation of the skull was measured at different square blocks of ABS having 2.5 mm thickness. Two different colors of ABS from Stratasys were used (Gray and white). For the white color, development with two different layer thickness were used (0.1 mm and 0.25 mm). Finally a third technique to increase attenuation was to use a coating either in just one side of the sheets or in both sides. Fig. 3 shows the attenuation coefficient in dB/cm MHz vs. material type. The attenuation of white ABS was found to be higher than the gray ABS. This is attributed to the fact that white ABS seem to be porous. This property of white was also confirmed by filling containers made of white ABS with water. The water leaked out of the container indicating porous in this type of material. Producing the sheets with thicker layers increases the attenuation coefficient. Finally, using a special coating in both sides increases the attenuation coefficient.

0.5

1

1.5

2

2.5

3

3.5

Silica Concentraon % (w/v)

B Aenuaon Coefficient (dB/cm-MHz)

3.1. ABS plastic attenuation coefficient measurements

1 0.8 0.6 0.4

y = 0,163x + 0,03 R² = 0,9968

0.2 0 0

10

20

30

40

50

60

Evaporated Milk Concentraon % (v/v) Fig. 4. (A) Scattering coefficient (dB/cm MHz) vs. silica dioxide concentration using 2% w/v agar. Scattering coefficient of 0.35 dB/cm MHz was achieved using 1.2% w/v of silica dioxide. (B) Attenuation Coefficient measurement for different evaporated milk concentrations (10%, 20%, 40%, 50% v/v) to a 2% w/v agar, at 3.6 MHz. The extra attenuation coefficient of 0.25 dB/cm MHz (with scattering 0.35 dB/cm MHz) was achieved using 25% v/v of evaporated milk.

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3.3. Silica scattering effect The agarose gels with silica were scanned with a diagnostic ultrasound imaging system (Philips HD7 series Ultrasound Systems, Philips and Neusoft Medical Systems Co. Ltd., Shenyang, China) in order to demonstrate that gel attenuated ultrasound. As expected the pure agarose gel was homogeneously hypoechoic due to the lack of scattering (Fig. 5A). On the contrary, the agarose gel doped with silica appeared homogeneous but with an increased echogenicity, which was a result of a reflected acoustic wave reaching the transducer induced by scattering from the silica particles (Fig. 5B). The grey texture of this image demonstrated that gel phantom behaves ultrasonically like a normal soft tissue.

100 90 80

Temperature (⁰C)

fitted data was used to calculate the appropriate silica concentration for creating a cerebral tissue mimicking phantom with attenuation coefficient 0.6 dB/cm. It was concluded that the target recipe was an agar-evaporated milk-silica gel (2% w/v–25% v/v– 1.2% w/v).

70 60 50 40 30 20 10 0

0

50

100

150

200

250

Time (s) Fig. 6. Temperature elevation vs. time at the focus in the agar/silica gel phantom.

severe artifacts were induced in the vicinity of the phantom. Some susceptibility artefacts can be observed on the T1W FSE image. The best quality images, in terms of signal to noise ratio were taken using the spine coil (Milwaukee, USA). These images provided sufficient proof that the phantom contained only MR compatible material.

3.4. Temperature measurements Fig. 6 shows the temperature raised from a baseline value of 37 °C to a plateau of high temperatures within 9 s. The plateau had an average temperature of 81 °C which was sustained for the whole duration that the ultrasound was activated. The temperature showed a high variability during heating because of possible interactions with the thermocouple. When ultrasound was activated heat absorption worked constructively towards raising the temperature while heat dissipation through conduction to the surrounding decreased temperature. When ultrasound was deactivated the only present mechanism was heat loss through conduction which is a slower process since it depends on the temperature difference between the hot focus and the colder surrounding. With time this temperature difference decreased and therefore heat loss rate also decreased. 3.5. Head phantom construction The skull model was filled with an agar-evaporated milk-silica gel (2% w/v–25% v/v–1.2% w/v) to acoustically mimic the brain tissue. The total volume of the brain mimicking gel was 800 ml (Fig. 7). A thin plastic membrane was placed in between the plastic skull and the gel to prevent the presence of air bubbles. 3.5.1. MR imaging of the head phantom The skull-brain phantom was positioned in a water tank and was imaged with the 1.5 T GE MRI scanner with conventional T1W FSE and T2W FSE sequences (Fig. 8A and B). As expected no

3.5.2. Gel phantom relaxometry Magnitude MRI images represent the distribution of transverse magnetization in space and consequently the mean pixel value in a region of interest is proportional to the average transverse magnetization. The spin lattice relaxation time (T1) of the gel phantom was estimated by taking images and measuring the mean pixel value of the gel phantom for different inversion times. The nulling inversion time was interpolated at 565 ms and using Eq. (1) we estimated T1 at 852 ms. Spin–spin relaxation time (T2) according to Eq. (2) is expected to vary exponential with echo time. T2 relaxation time was estimated by measuring the mean pixel value for different echo times and calculating the inverse exponent of the exponential fit and was found to be 66 ms. Fig. 9A shows MRI thermometry map of the beam in a plane parallel to the transducer beam using Double Echo Segmented (EPI) without the plastic skull. The acoustical power used was 30 W for 60 s. Fig. 9B shows MRI thermometry map of the beam in a plane parallel to the transducer beam using Double Echo Segmented (EPI) with the plastic skull. Note that with the skull the temperature is decreased substantially, and the beam losses its focal capabilities.

4. Summary This paper described the design of a head phantom suitable for testing in parametric studies of transcranial MRgFUS. The phantom

Fig. 5. Ultrasound imaging using a 12 MHz transducer of (A) 0% silica to agarose gel and (B) 2.1% w/v silica to agarose gel.

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Skull filled with Agar/Silica gel

ABS skull model

B

A

Fig. 7. (A) ABS prototype of an adult human skull and (B) ABS prototype of an adult human skull filled with an agar–silica gel.

B

A

Fig. 8. (A) T1-W FSE image of head phantom and (B) T2W-FSE image of head phantom.

A

B

Transducer

Transducer

Fig. 9. MRI thermometry map of the beam in a plane parallel to the transducer beam using Double Echo Segmented (EPI) (A) without the plastic skull and (B) with the plastic skull. The power used was 30 W for 60 s. The transducer diameter was 40 mm, the radius of curvature was 100 mm and the frequency was 1 MHz.

fulfilled the requirement of being compatible in the vicinity of a magnetic field inside an MRI scanner. A skull model was produced in the form of a 3D printed plastic replica following segmentation of the patient’s head CT scan images. The plastic model retained accurately the geometrical characteristics of the patient’s skull which is essential for testing the degree of induced defocusing effects. The attenuation coefficient of the plastic skull model was

matched as close as possible to that of a typical human skull by selecting properly the ABS material. To achieve an attenuation coefficient as close to that of humans (13–24 dB/cm MHz), we have use white ABS material from Stratasys, manufactured with 0.25 mm layers, and using a special coating on both sides of the skull. Using these techniques the attenuation coefficient of the plastic skull was 16 dB/cm MHz.

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The white ABS material includes porous, but since in all the evaluations that we performed (including the attenuation measurement), the skull is immersed in water, then the possibility that bubbles are trapped in these small cavities is minimal. However, the presence of these cavities definitely could increase scattering, and that is why white ABS has larger attenuation than Grey ABS (which is presumably non-porous). Another difference between the 3D printed skull and a real human skull is the bone layers of a human skull where there is a tubercular layer in between two cortical layers which has different acoustic properties and may behave differently in different angles and frequency. This first design of a plastic skull is simplistic and at the moment it is impossible to include all aspects of real skulls. However, its overall behavior regarding attenuation makes it as useful tool for evaluating MRgFUS. Unlike all previous transcranial MRgFUS studies, the proposed design avoided using biological tissue in the form of cadaver or animal models. Unless they are used fresh, cadaveric skulls require formalin fixation in order to be preserved for future study. Formalin fixated human skulls were found to exceed significantly the acoustic attenuation loss of fresh cadaver skulls on average 3 dB over for an average attenuation of 10 dB [56]. In vivo studies usually use small animal models and therefore their results cannot be translated reliably to the clinical setting with humans. The replacement of animal models with the phantom is expected to reduce the running costs of research in MRgFUS. Similarly the attenuation coefficient of the agarose gel was matched to that of a typical brain by varying the concentration of silica in the gel. The set of temperature measurements proved that the prescribed gel recipe not only attenuates the acoustic beam by scattering, but also through absorption. The acoustic absorption is observed as a rise in temperature in the gel. For the power used we managed to exceed 80 °C within seconds starting from the baseline, which lies well above the threshold temperature of thermocoagulation. Although the skull can be used for long term use, the agar solution that mimics brain usually lasts for 8–10 days. The life of the agar phantom can be extended by adding preservative material in the solution. Due to the low cost of purchasing and preparing the agar phantom, at the moment no preservatives were used. If the agar phantom is heated to high temperatures, then due to the liquidation of the gel, the agar phantom has to be disposed. Again this is not a problem due to the low cost of agarose. The measured relaxation times T1, T2 of the prescribed agar–silica gel recipe were within the range of values found in bibliography for a normal brain [68,69]. The measured T1 and T2 relaxation times for the agar/silica phantom were 852 ms and 66 ms respectively. The effectiveness of the skull to reduce ultrasonic heating was demonstrated using MRI thermometry. Clearly, the beam of the single element spherical transducer used is defocused. Moreover, the temperature produced, compared to the situation that the skull is not present, is greatly reduced. A major future study is to use phased array technology using MRI thermometry to demonstrate the refocusing of the beam. Another issue not explored in this preliminary evaluation of the skull phantom was, the heating of the skull near the tissue/skull interface. This can be assess with MRI thermometry, and corrected by circulating cold water in the area of the skull that is exposed to the ultrasonic beam. This phantom can be set as a tool of absolute thermal dosimetry in MRgFUS ablation techniques and use it to calibrate such systems. The skull phantom is literally customizable to each patient skull geometry and therefore it can serve as a treatment planning tool where treatment plan will be validated before treatment application. Using conventional T1 and T2 sequences lesion forming can be monitored and fused with thermometry maps to evaluate the extent of the treatment. The dependence relaxation times to

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