03:47:29PM S3194 # MHz 180mm YNHH. MSc Thesis. 3D Transesophageal Echocardiography using a fast-rotating transducer. Kyriakos Nathanail

Transcription

1 03:47:29PM S3194 # MHz 180mm YNHH MSc Thesis 3D Transesophageal Echocardiography using a fast-rotating transducer Kyriakos Nathanail

2

3 3D Transesophageal Echocardiography using a fast-rotating transducer THESIS submitted in partial fulfillment of the requirements for the degree of MASTER OF SCIENCE in BIOMEDICAL ENGINEERING by Kyriakos Nathanail born in Athens, Greece Biomedical Engineering Department of Electrical Engineering Faculty of Electrical Engineering, Mathematics and Computer Science Delft University of Technology

4

5 3D Transesophageal Echocardiography using a fast-rotating transducer by Kyriakos Nathanail Erasmus Medical Center Laboratory: Thoraxcenter Biomedical Engineering Delft University of Technology Laboratory: Biomedical Instrumentation Committee Members: Advisor: Co-advisor: Chairperson: Member: Member: Assist. Prof. Johan G. Bosch, Exp Echo, Erasmus MC Prof. Nico de Jong, Exp Echo, Erasmus MC Prof. Paddy J. French, EI, TU Delft Prof. Gerard C. M. Meijer, EI, TU Delft Assist. Prof. Georgi N. Gaydadjiev, CE, TU Delft i

6

7 Contents ACKNOWLEDGEMENTS... V FOREWORD...VII INTRODUCTION ULTRASOUND IMAGING BASICS ECHOCARDIOGRAPHY THREE-DIMENSIONAL ECHOCARDIOGRAPHY MOTIVATION FOR RAPID 3D TEE... 9 MATERIALS AND METHODS THE VINGMED SYSTEM FIVE MULTIPLANE TEE PROBE PROTOTYPES OF THE FAST ROTATING PROBE FOR 3D TEE First prototype Second prototype Third prototype CONSIDERATIONS ON DATA MORPHOLOGY Example of actual distribution Conclusions on data distribution RECONSTRUCTION SOFTWARE PHANTOMS FOR LABORATORY EXPERIMENTS VISUALIZATION SOFTWARE IN-VITRO IMAGING THREE-DIMENSIONAL PHANTOMS Coffee-cup phantom Dolphin phantom Balloon phantom in 3D Rectangular block phantom FOUR-DIMENSIONAL PHANTOMS Rotating dolphin phantom Balloon 4D phantom setup EXPERIMENTS WITH PHANTOMS Issue of additional delay Beam order switching Effect of the number of frames used for reconstruction Shifting of the elevation angles Experiments with the 4D setups CONCLUSIONS ON THE IMAGING OF PHANTOMS IN-VIVO IMAGING IN-VIVO ANIMAL EXPERIMENTS CLINICAL TRIAL SHORT DISCUSSION ON IN-VIVO RESULTS CONCLUSIONS CONCLUDING REMARKS SUGGESTIONS FOR FUTURE WORK EPILOGUE BIBLIOGRAPHY iii

8

9 Acknowledgements The outcome of this thesis has been the fruit of collaboration among a certain number of people, on a scientific level. Furthermore, the journey to this result would probably have been dull, pointless and even unbearable without the mental, spiritual, and, in times, nutritional support from a number of people. I am reluctant to place here a lengthy list with all these names because I have it reserved for other purposes. However, I need to mention some people who have actively been involved in this project, in multiple ways. I wish to thank my advisors Assistant Prof. Johan G. Bosch and Prof. Nico de Jong for their guidance and on-demand feedback, throughout my time in the Experimental Echocardiography group, as well as Prof. Paddy J. French for giving me the opportunity to spend this last year at the Erasmus MC. Further, I wish to thank for their patience Marijn van Stralen (while discussing all my questions and remarks on his work) and the entire Exp Echo group. I need to acknowledge the hard work of the people at Oldelft Ultrasound B.V. and especially the close cooperation of Christian Prins and Franc van den Adel. The technical and medical support provided by Robert Beurskens, Gerard van Burken, Jan Honkoop, Wim van Alphen, Geert Springeling, Cees Pakvis, Stefan Krabbendam, Wim Vletter, Jackie Vletter-McGhie, Dr. Folkert ten Cate, Dr. Marcel Geleijnse and Dr. Michelle Michels is gratefully acknowledged. Finally, I wish to thank Aikaterini Brisimi and Ana Laura Rodriguez-Santos for their assistance with the sound realization of this thesis text. The date below might have been a couple of weeks later if it was not for their support. I am grateful to Christos Strydis for proofreading this text while on Easter vacation and for providing useful input for the outline used herein. Kyriakos Nathanail Delft, The Netherlands May 5 th, 2008 v

10

11 Foreword Echocardiography is a diagnostic modality which is widely used in cardiology, using ultrasound to image the living heart and assess its physiological function. Threedimensional echocardiography surpasses the visual limitations of conventional echocardiography, offering a three-dimensional perspective of the heart. Traditional techniques for acquiring three-dimensional echocardiographic images have been successfully applied to the transesophageal approach to echocardiography. The important clinical contribution of three-dimensional transesophageal echocardiography (3D TEE), as well as the unique diagnostic opportunities it offers, have been acknowledged. However, the applied techniques are based on a slow procedure, cumbersome for the patient and present certain pitfalls. Recent technological advancements have allowed more flexible approaches to 3D TEE. Such approaches are, nonetheless, based on an expensive solution with inherent limitations, available only with specialized, high-end imaging systems. The present thesis is a description of the author s work on the realization of threedimensional transesophageal echocardiography using a TEE probe with a fast-rotating transducer array. It describes the entire endeavor of visualizing the heart and its inherent structures in three dimensions by implementing a novel approach in transesophageal echocardiographic imaging. The approach itself and subsequent research to implement and optimize it are portrayed. The author s work was carried out in its entirety at the Biomedical Engineering department of the Thoraxcenter, Erasmus Medical Center, Rotterdam, the Netherlands, within the Experimental Echocardiography group (ExpEcho) in close cooperation with Oldelft Ultrasound B.V (Oldelft), Delft, the Netherlands. The organizational structure of the thesis attempts to introduce the reader to background knowledge necessary for understanding the subject, present the technical aspects involved and discuss the results and conclusions of the work. The first chapter establishes the background for concepts and terminology related to the thesis subject. In the second chapter, the details on resources and devices used are illustrated, connecting them with the work involved. Further, the design concepts are justified according to specific requirements of the approach. Chapters 3 and 4 describe the associated experimental work. The outcome of that work and its implications about the success of the initial concept are portrayed. The images in these two chapters which correspond to avi-files located in the accompanying CD-ROM are marked with the tag Video available. Finally, chapter 5 summarizes the results of the work, presenting certain conclusions extracted from that work. Suggestions regarding the future direction of research on the topic are provided. vii

12

13 Chapter 1 Introduction A basic introduction is required to delve into the project details and work therein. This introductory information comprises background details on the project, terminology used and the motivation for embarking on such a project. Projected objectives and justification for such expectations are laid down. In this context, the current chapter gives a brief description on ultrasound imaging of the heart. A preface on basic aspects of ultrasound imaging is given. The different concepts involved in recording images of the heart by use of ultrasound are presented in a way so as to clarify the motivation for the project. Finally, the definition of results originally envisioned is ordained. 1.1 Ultrasound imaging basics Ultrasound, as a term, refers to mechanical pressure waves with frequencies well beyond the audible range of 20 Hz to 20 KHz. [1] Depending on the application, frequencies of 1 to 50 MHz are used. Devices that use ultrasound for various applications employ piezoelectric elements, widely known as transducers. Transducers can be singleelement or arrays, the latter consisting of a number of elements arranged in a specific layout. Piezoelectric elements can emit (electrical signals converted to mechanical pressure) and receive (voltage generated by application of mechanical pressure) ultrasonic waves. [2] Emitted waves travel at different speeds through different media, depending on the elastic and inertial properties of the media. [3] The device carrying the ultrasonic transducer module or array is known as ultrasound probe. Medical imaging with ultrasound When ultrasound propagates through biological tissue, it interacts with different tissue types and is partially reflected, scattered or absorbed by tissue. [2] The reflected or back-scattered waves, known as echoes, are received by the transducer. Information can be obtained from the intensity and originating location of the received echoes. Images of the area of tissue through which ultrasound has propagated can be formed using this information, as explained later. 1

14 CHAPTER 1 2 While several diagnostic modes exist, the one widely applied and also of direct interest for the comprehension of the principles used in this project is two-dimensional tissue imaging. Two-dimensional (2D) images of cross sections of the biological structures (organs) imaged with ultrasound can be obtained. The elements of the transducer array transmit an ultrasound pulse in the form of a beam. The beam is fired under different angles, scanning over a 2D sector of an imaginary semicircular disk. The direction of the scanning of the beam is referred to as the elevation direction and the angles at which it is fired are known as elevation angles (sign: θ ). The transducer array receives the echoes of the transmitted ultrasound beams which have been reflected by the scanned structures. [20] An ultrasound imaging system processes the information of the echoes received from each transmitted beam and constructs respective scan lines of the 2D sector image. Each line consists of a number of samples. Samples are grey-scale representations of the echoes received in the form of RF signals along the propagation direction of the transmitted beam. The direction of the beam is also referred to as depth (sign: r ). One can also refer to a specified distance in the propagation direction as a depth level. The grey-scale values are based on the intensity levels of reflected echoes corresponding to the propagation time of the ultrasound pulse from the transducer surface to the reflector. The collection of constructed lines forms a 2D sector-image of the targeted structures. The 2D semicircular disk sector is defined by the angle of its arc. The angle at which the first line of the image-frame is received is called angle of origin. The spacing between adjacent image lines is called angular spacing. Figure 1: 2D sector with image (scan) lines. Each line consists of a number of samples in the depth direction. The angle of origin of the 2D sector is shown, in the case that scanning of the sector is performed from left to right. ([4]) Images are refreshed at a certain frame-rate, on the screen of the imaging system to visualize potential motion of the scanned biological structures. The maximum depth level determines the repetition frequency of fired pulses (PRF). The number of lines per 2D plane determines the angle of the conical 2D sector. Both have a direct impact on the number of frames per second (fps) that can be constructed and visualized. Some special

15 3 INTRODUCTION techniques for firing and receiving ultrasound pulses are applied to achieve an increased frame rate of the generated 3D sequences. [4] Bones, gas masses and foreign objects inside the body can cause very strong reflections, prohibiting further propagation of ultrasound. They are, therefore, considered obstacles because no information about structures behind them on the path of ultrasonic waves can be obtained. [2] 1.2 Echocardiography The term echocardiography is derived from the Greek words echo [ηχώ = echo, sound], cardia [καρδία = heart] and graphy [γραφή = writing, recording]. It refers to the specific application of ultrasound imaging to recording images of the heart and its structures as well as their motion through the cardiac cycles. After being introduced in the mid-1950s, it has constantly and rapidly developed, featuring different modes such as M- mode, 2D Echo and Doppler imaging. [5] Echocardiography is of large clinical significance in cardiology as a diagnostic modality to assess the physiological function of the heart and potential disorders. [1], [5] In order to comprehend and justify certain points of the present thesis, it is required to discuss two-dimensional echocardiography to some extent, omitting other echocardiographic modes. 2D ultrasound imaging of the heart requires that specific areas of the body are used for placing the ultrasound probe. These areas are known as imaging windows since they present an ultrasonically transparent pathway from the transducer to the heart; a passage-way for propagation of ultrasound beams, unhindered by any form of obstacles such as bones or gas-filled cavities. Since the heart is situated behind the thorax, these windows necessarily lie in between ribs (namely 3 rd, 4 th and 5 th intercostals) and more specifically on the left-hand side of the thorax (3 cm from the sternal border). At this level, the heart is not covered by the left lung and the distance from the chest wall to the furthest point of the heart is approximately 12 cm. [5] Ultrasound imaging from the thorax is known as precordial or transthoracic echocardiography (TTE). Transesophageal echocardiography Transesophageal echocardiography (TEE) makes use of the esophagus as an imaging window in the immediate vicinity of the heart. The technique employs a flexible endoscope, with an ultrasound transducer array positioned at its tip. The esophagus wall lies directly posterior to the heart (Fig. 2a, b), allowing for cardiac ultrasound imaging without interference from the ribs or the lungs. What is more, the small distance between the transducer array and the heart s structures to be imaged allows for use of higherfrequency ultrasound pulses and thus increased resolution. [7] This is crucial for imaging structures such as the mitral valve and assessing its potential dysfunction or looking for other malfunctions. Further, the technique is useful for imaging the thoracic aorta but also

16 CHAPTER 1 4 providing short-axis view of the heart by progressing the probe further in the stomach (transgastric view). The position is useful for intra-operative monitoring of global and regional myocardial function. [7] a) b) Figure 2: a) Patient at supine position with a TEE probe through the esophagus. ( b) Artistic interpretation of TEE imaging (Yale Medicine, While transthoracic echocardiography fails to consistently obtain high-quality imaging of the heart and aorta in all patients, TEE overcomes related limitations and offers valuable additional data. [6] It can further be used with obese or elderly patients on whom TTE may have unsatisfactory results. [2] 1.2 Three-dimensional echocardiography Visualization of the heart in three spatial dimensions offers a perspective of the heart very much closer to the real anatomical structure. Correct spatial information of the structures provided by 3D images overcomes the need for conceptualization of the heart s model from 2D cross-sectional images. [8] The possibility of visually removing structures to view inner parts of the heart offers surgeons information on structure morphology and functionality. The value of such views becomes profound if one takes into account the non-physiological state of the heart on which the surgeon operates. [8], [9] A realistic view of the cardiac valves can be obtained, useful for the diagnosis of certain congenital abnormalities and heart valve dysfunctions (e.g. bicuspid aortic valve, mitral valve prolapse, vegetations etc.) Initial efforts on 3D echocardiography (3D Echo) provided, for example, a more in-depth understanding of the mitral valve s saddle shape and led to redefining the diagnostic criteria for mitral valve prolapse. [10] Direct evaluation of cardiac chamber volumes, without using geometric approximations is another of the multiple advantages of 3D Echo. [11] Assessment of the effectiveness of surgical interventions and the functionality of prosthetic valves can benefit from 3D visualization of respective structures.

17 5 INTRODUCTION Basics on 3D Echo techniques The principles of the most widely used 3D Echo techniques throughout the history of echocardiography should be briefly explained. One can generally describe the procedure as the sequential acquisition of 2D echocardiographic images from discrete, adjacent locations. The position of the 2D images in 3D space is registered and temporal information regarding the point of image data acquisition within the cardiac cycle is simultaneously recorded. This last aspect of gathering image data for different positions in time is mainly the reason for referring sometimes to 3D echocardiography as four-dimensional (4D) the fourth dimension being time or cardiac phase. Further, besides electrocardiogram (ECG)-gating of the image acquisition, respiration gating is sometimes used to minimize motion artifacts caused from breathing. The acquired 2D image data is processed along with information on the position of the 2D images to align them in space and time. Interpolation is performed to fill the gaps between images adjacent in space and/or time and redistribute the data in a regular image grid. A 3D volume is derived from the processed data which visualizes the imaged structures in three-dimensions. [10], [11], [12] Different techniques are used for spatially varying the imaging plane to scan the entire volume of interest by free-hand motion of the transducer array. The transducer s position relative to some reference point can be recorded, thus obtaining positional data of the acquired 2D images. [8], [10], [13] The probe is translated over a linear range (Fig. 3a) and the acquired images are positioned according to the obtained data. Another technique is based on the assumption that the patient and the housing of the transducer remain in a relatively fixed position. [15] In this perspective, the 2D images can be acquired by computer-controlled tilting of the transducer array in specified arc angles for acquiring multiple 2D images (fan-like acquisition, Fig. 3b) or by rotating the transducer array and obtaining images at predetermined intervals (Fig. 3c). [15] A frame or an entire cardiac cycle per discrete recorded spatial position is acquired. a) b) c) Figure 3: Different techniques to scan the entire volume of interest and acquire 2D frames for 3D echocardiography; a) Linear translation of the probe, b) Tilting of probe for fan-like acquisition. (a, b: [13]), c) Sequential rotation of the transducer array at predetermined intervals to acquire multiple 2D frames for 3D reconstruction. ([14])

18 CHAPTER 1 6 The problems regarding these approaches are primarily related to the lengthy duration of the acquisition procedure. ECG and respiratory gating theoretically ensure that all acquired image data is from the same discrete point of the cardiac cycle. [12], [14], [15] This means that the reconstructed images depict a specific point in time of the cyclic motion of heart structures. However, certain artifacts can be caused by involuntary movement of the patient. It is reported and also constitutes common knowledge that it is difficult to stay still for a period of minutes, during an examination. [14] Fast rotation for 3D TTE Echocardiography Techniques to overcome the problem have been proposed and implemented. [12], [15], [16] One of these is the fast, continuous rotation of the transducer array of a TTE probe. Such a probe has earlier been developed by the ExpEcho group in cooperation with Oldelft. Multiple 2D images from many different angles of the imaged volume are recorded in a minimum amount of time in this way. The rotation angle is derived from the continuous rotation speed. Roughly described, the rotation speed is obtained with high precision by correlating the recorded frames with an arbitrary reference frame and extracting a periodic signal whose periodicity is related to the speed. [12] Since this is a continuous rotation at a constant speed, the relative position of each frame in space can be extrapolated and 3D cone volumes reconstructed. Besides the technique for data acquisition it is important to understand the morphology of the acquired image data. Contrary to gated, sequential acquisition, the image plane is curved in space. This is due to the scanning of the beam through the 2D sector while the transducer is rotating. The axis of rotation is at the center of the imaging plane. We assume that the first line acquired is on the left side of the sector as seen in figure 4a and rotation is counter-clockwise. Each subsequent line acquired as the beam scans through the sector has an increasing rotational angle since the transducer array has rotated further in the counter-clockwise direction. Between the first and last elevation beam of each frame the rotational angle varies more than a) b) Figure 4: Morphology of acquired image data; a) A curved frame in the spherical and Cartesian coordinate systems. The frame is divided by the rotation axis in a left and a right half. b) The samples of different beams at a certain depth level are presented. White dots indicate beams from the left half of each frame and black dots indicate those from the right half. (a, b: [12])

19 7 INTRODUCTION An additional point concerning image data morphology is that due to the fast rotational motion of the transducer, the volume is oversampled around the axis of rotation compared to points close to the cone border. Figure 4b depicts the line samples, at a certain depth level, of different frames acquired during rotation. The area close to the axis of rotation is densely populated by image samples. Data becomes sparse near the borders of the 3D volume. Techniques for 3D transesophageal echocardiography The concept of rotating the transducer array to obtain 2D images (also known as slices) of the heart has also been used to introduce the transesophageal approach to 3D echocardiography. Latter designs of TEE transducers featured a modality for rotating the transducer array within the tip of the probe (Fig. 5) to view different 2D slices of the heart at the desired esophageal levels. The modality was used for 3D TEE Echo to acquire ECG- and respiration-gated 2D images of the heart from different angles, at increments of 3 0 to 6 0 degrees. [17] The acquisition process is, on average, 3 minutes but can be extended to 10 minutes with certain patients having an irregular heart rate. [14] Figure 5: Rotating transducer array of TEE probe. The transducer array of elements is represented as a horizontal black line based on a circular disk support. The disk, and thus the array on it, can rotate clockwise and anti-clockwise. Rotation is limited to a certain range of degrees as explained in chapter 2, under section 2.2, where the multiplane TEE probe is described. The fact that the TEE-acquired 2D images are of superior quality compared to the TTE approach is a definite advantage. [7] Nonetheless, ECG and respiration-gating may result in decreased image quality in specific situations. The lengthier the procedure, the more prone to motion artifacts it becomes. Experience has proven that due to the semiinvasive character of the examination the patient cannot be expected to hold perfectly still for long periods of time.

20 CHAPTER 1 8 a) b) Figure 6: Three-dimensional images of the mitral valve orifice reconstructed from ECG- and respiration-gated TEE-acquired 2D images. (a:[17], b: [9]) A technique which addresses 3D echocardiography in a more flexible way compared to the ones already mentioned is the use of electronic beam steering in two dimensions. A matrix array transducer is used to image the heart either precordially (matrix TTE) or through the esophagus (matrix TEE). The transducer is made of a 2D grid of piezoelectric elements. Elements are fired non-simultaneously to steer the ultrasound beam in two perpendicular directions, obtaining a pyramidal shape of image data of the imaged volume. Data is processed online and visualized on-screen. For angles of the pyramidal 3D volume up to 30 0 by 50 0, the 3D volume of a cardiac cycle can directly be processed and presented online without any ECG gating, thus the name real-time, associated with the technique. Normally, such a volume is insufficient to image the left ventricle. Then, ECG-gating is employed for acquiring four 3D sectors of about 23 0 by 90 0 during four consecutive heart beats. The four sectors are then combined, side-to-side, to visualize a pyramidal 90 0 by 90 0, so-called full volume data set. [18] 3D imaging with a matrix TEE is flexible, fast and powerful. However, such a probe is extremely expensive and can only be used with specialized ultrasound imaging systems. Also, the 2D image quality of the probe is compromised. Figure 7: Transthoracic echocardiography (TTE) with the matrix array transducer. The pyramidal sector shown is equivalent to the one scanned with matrix-tee. Scanning of sectors with wider angles is possible integrating four sectors, acquired during four consecutive heart cycles. ([18])

21 9 INTRODUCTION 1.4 Motivation for rapid 3D TEE Even though the matrix transducer is considered the state of the art for 3D Echo imaging, it fails to maintain optimal quality in 2D imaging. The aforementioned department of the Erasmus Medical Center in Rotterdam has previous experience with the fast rotating TTE probe. [12] Therefore, the next logical step was to assess the possibility and advantages of multiplane TEE for rapid 3D echocardiography. The demands, though, for rapid 3D TEE are definitely different from those for TTE. Probably the most profound difference is the constraint of the transducer array used in TEE probes to rotate through a limited range of rotation. Continuous rotation at constant speeds is not an option and therefore, left-right rotation with varying speeds is needed. Required spatial information about each recorded image line cannot be directly extrapolated from speed and has to be measured. The concept had to be further investigated to design a suitable probe for acquisition and proper reconstruction of 3D data sets from the acquired data. Purpose of work and projected goals The idea of rapid 3D TEE using a fast-rotating transducer needed to be explored. The actual purpose of the work described in this thesis was to assess the feasibility of generating 3D sequences of the human heart s cycle with the fast-rotating probe. As soon as the usefulness of the concept was proven with imaging of 3D objects and results showed the potential of rapid 3D TEE, implementation of the concept for 3D imaging of the human heart was expected to follow. We had to combine the work of Oldelft on designing and developing the probe, the reconstruction software developed by a colleague and the expertise available in the department on the different topics involved, so as to reach the reported aim. In that process, we had to suggest and in cases implement improvements and necessary modifications for the designed probe which would lead to the desired goal. In order to offer such suggestions we designed and conducted various experiments with ultrasound imaging of motionless and moving objects. The improved understanding of the entire procedure, stemming from the experimental results, led to pinpointing specific problems and issues thereof, and potential solutions to overcome them. Clinical trials were conducted thereafter, showing the feasibility of the concept and offering valuable experience in the clinical setting. What is more, significant issues concerning clinical application of the probe came to the foreground which could only be hypothesized before. Finally, the outcome of the clinical trials verifies not only the feasibility but also the clinical value of the realization of the concept. Discussing the combined results from phantom and clinical trials, we have concluded on some additional improvements. Such

22 CHAPTER 1 10 improvements will be implemented in a further commercial version of the probe and software.

23 Chapter 2 Materials and methods This chapter introduces the reader to the equipment and resources used throughout the project. As mentioned, the project features the work of two parties. Oldelft specializes in the development, manufacturing and service of ultrasound transducers. [32] The ExpEcho group is part of the Biomedical Engineering department of the Thorax Center at the Erasmus Medical Center in Rotterdam, the Netherlands. The group investigates the physical and technical aspects of ultrasound and its application for cardiac diagnosis. [20] The author worked for the project at the ExpEcho group. Software developed in the group was used for data processing. For the visualization of the results commercial software was used. Some of these various resources hardware and software required little or no additional work on our behalf to be utilized. Still, the majority of them had to be adjusted, designed from the beginning or redesigned and adapted to fit the needs of the project. The involved parties had discrete roles in the project. Franc van den Adel and Christian Prins from Oldelft Ultrasound B.V. were responsible for the design and implementation of the custom Fast Rotating Transesophageal Echocardiography (FRTEE) ultrasound probe. The company developed the mechanics, electronics and software for additional control of the probe and for acquisition of rotational angle data, required by this variant of 3D Echo applications. [17] The tasks of the author working for the Experimental Echocardiography group involved the utilization of the designed FRTEE probe for acquisition of 4D image datasets. 4D refers to the motion of the imaged 3D object in time. The group had already developed software for reconstruction of three dimensional images and four dimensional datasets. Marijn van Stralen originally developed the software for a TTE probe featuring continuous fast rotation of the transducer. He also adapted it for the needs of bidirectional transducer rotation used for 3D TEE acquisition. Later, the author adjusted and optimized it to accommodate the designed FRTEE probes. Further, we had to cope with pitfalls arising from both experimental laboratory trials and the actual clinical application of the probes. Modifications were proposed to 11

24 CHAPTER 2 12 Oldelft along with adjustments and solutions that would render the devices suitable and useful for the clinical environment. This second chapter of the thesis describes the objects and devices employed for the conducted experiments. It further discusses related modifications, applied settings and parameters in the course of this work. Its goal is to familiarize the reader with those devices and their working parameters in order to allow further explanation of the concept on which this project was based. The description starts with the Vingmed System FiVe, an ultrasound imaging system which is the main-frame used for acquiring, storing and visualizing 2D image data. A short description of the multiplane TEE probe follows, acting as a reference for comparison with the three consecutive designs of ultrasound probes used for this project. Section 2.4 offers insight on the software used for reconstruction of 3D images from recorded data sets. The final sections for this current chapter describe the visualization software used as well as objects (phantoms) constructed and used during the experiments of the project. Phantoms are structures and setups mimicking the reflective and locomotory properties of tissue. 2.1 The Vingmed System FiVe Vingmed System FiVe, (General Electric/Vingmed, Horten, Norway) is a commercial echocardiographic system 1. 2D sector images are displayed on the screen of the monitor of System FiVe during each acquisition and refreshed at the respective framerate. Significant information regarding the internal operation of the system was based on knowledge available within the department, experimentation with the system and communication with Johan Kirkhorn from GE Vingmed Ultrasound. Simultaneous acquisition of the electrocardiogram (ECG) signals along with the ultrasound data is possible. In effect, recorded frames can directly be assigned to the different phases of the heart cycle. ECG is recorded in the clip file containing the image data. In conventional ultrasound imaging, ECG is mainly used to give timing information for locating images during the cardiac cycle. [19] Further, it is used to trigger acquisition of a frame (usually at the R-peak of the electrocardiogram). In the current project ECG has been used to retrospectively sort the acquired ultrasound lines belonging to the same phase of the heart cycle. System FiVe has a user panel with various controls. It features a Freeze / Unfreeze button which, when pressed, starts image acquisition in the selected imaging mode and stores the images in a ring buffer. Pressing the Freeze button again, stops the acquisition and a maximal number of the last 5650 frames can be reviewed and stored. Acquisition parameters (frequency of RF signals, ultrasound sector angle, maximum imaging depth, focusing depth, frame-rate etc.) can be adjusted via the user panel. These parameters are inter-related. For example, minimizing the angle of the ultrasound 2D sec- 1 An ultrasound imaging system used for echocardiography

25 13 MATERIALS AND METHODS tor decreases the number of scan lines of the 2D sector image, thus increasing the framerate of image acquisition. The pulse repetition frequency (PRF) of the transmitted pulses varies with imaging depth within the range of around 2.5 KHz to 8.3 KHz. Different acquisition modes are also available (e.g. ECG-triggered, M-Mode, Doppler imaging etc.). a) b) Figure 8: The Vingmed System FiVe echocardiographic system. Front view (a) with the user-controls and side view open (b) showing the signal transmission, acquisition and processing boards. The system stores images or clip files, constructed from the received ultrasound data, on its internal hard drive. A clip file contains a sequence of 2D polar frames (θ, r domain), each frame consisting of adjacent 2D sector scan lines as shown in figure9a. The image data in the form of rectangular frames is known as non-scanconverted data. Besides the image data, the file contains information about the angle of origin and the angular spacing between adjacent lines. With this information, a program can depict the non-scanconverted data as a sector image, similar to figure 9b, which will offer a view of a cross-sectional cut of the structure imaged with the ultrasound transducer. The order in time of beams transmitted ( beam order or theta order ) to acquire image lines can also be recorded by using certain commands via a telnet client connected to the system. Except for the user-controllable parameters described above there are also internal acquisition parameters which cannot be accessed by the user control panel. These can only be adjusted by altering the system and probe files which are loaded when a probe is connected to the system. Such system files are accessed over telnet protocol via a UTP connection between a computer and the system. Three of the internal parameters which have been of significance for our work are the multi-line acquisition (MLA), the MLA stitching algorithm and the interleaving of ultrasound beams.

26 CHAPTER 2 14 a) b) Figure 9: Scan-conversion of the stored rectangular frame into a geometrically correct image (2D sector); a) The information is represented in the (θ,r) domain. b) Pixel values are positioned in Cartesian (x,y) coordinates. The annotations of θ and r on this image serve as reference to the original dimensions. (adapted from [4]) At MLA mode, the transducer array transmits a broad beam and the received signals from all transducer elements are processed in parallel by two or more signal processing circuits known as beamformers (parallel beam-forming). [12], [21] Each beamformer time-delays the element signals differently to generate different receive beams, thus acquiring two or more lines per transmit pulse and increasing the frame rate accordingly. [20] While the effect of the MLA mode is desirable, the stitching algorithm interpolates between acquired lines giving image lines with data averaged from received beams. In our case, significant temporal artifacts 2 can be introduced by MLA stitching, since the beams may be spatially close to each other but bear information from discrete points in time and cannot be arbitrarily fused. At an early stage during our experiments, the issue was identified and the stitching algorithm was switched off. Finally, the interleaving of ultrasound beams refers to the acquisition of ultrasound beams in an interleaved manner, transmitting for example the first, then the third, fifth and seventh beam, then the sixth, fourth and second and so on. After having transmitted one beam, before transmitting again in the same direction, the scanner needs to verify that the previous pulse has traveled to a large depth, significantly attenuated so that it does not affect the current pulse being transmitted. The attenuation level is specified by the AL (attenuation level) parameter in the probe file with the engineering parameters, used by System FiVe. While AL would be expected to largely affect modes such as Doppler and contrast echocardiography, where multiple pulses are fired in the same direction, it might also affect regular 2D scanning if the beam density is high. For our experiments AL was set to 36 db, implying that a new pulse should be transmitted in the same direction only after the previous pulse in that direction has been attenuated 36 db. A second parameter, SL, which refers to the attenuation of the previous pulse in the lateral (sidelobe) direction, was not adjusted to prevent getting unwanted signals in the sidelobe direction. The two chosen parameter values ensured that the beam order would be linearly increasing and 2 Artifacts caused by misplacement of image information in time

27 15 MATERIALS AND METHODS defined only by the total number of lateral samples. This helped in promptly defining the theta order of the acquired lines from the recorded information. 2.2 Multiplane TEE probe The multiplane transesophageal echocardiographic probe manufactured by Oldelft is a flexible endoscope which has a transducer at its tip. The transducer is a 64-element phased array with a pitch of mm between individual elements. It has a central frequency of 5 MHz and bandwidth of above 80%. The shaft of the transducer is 11.0 mm thick. The transducer array can rotate clock- and anti-clockwise in a maximum range of approximately of rotation. This rotation is restricted by the flex-print wires used to collect the RF signals from the elements of the transducer. (Fig. 10) Figure 10: The tip of the multiplane TEE probe and the projected image of the inside of the tip. The flex wires collecting the RF signals from the array as well as the support on which the array is fixed are seen. The arrows indicate the clockwise and anti-clockwise rotation of the array. (adapted from photos provided by Oldelft Ultrasound B.V.) A flexible tube encompasses the steering wires and the RF signal cables. The last part of the tube, bearing the tip where the transducer is located, is rigid. For the motorized commercial version of the probe, the handle of the probe houses the motor which steers the array to the left and to the right by use of two buttons on the side of the handle. Rotation is translated to the transducer array through a wires/pulley mechanism. An angle sensor inside the handle reads out a rough estimate of the transducer angle of rotation to be displayed on-screen of the imaging system. At the top of the handle a large knob allows for right/left flexion of the rigid part of the probe while a smaller wheel underneath the large one enables flexion in the anterioposterior direction. The control knob and rotation buttons offer the clinician a variety of choices for imaging planes, each one presenting

28 CHAPTER 2 16 diagnostically important cross-sections of the heart such as the four-chamber view or the short-axis view. Figure 11: The standard multiplane TEE probe by Oldelft Ultrasound. This version does not feature a motor to rotate the transducer array. The large knob at the top of the handle controls the wires/pulley mechanism, instead. ([30].) The probe tip is equipped with a thermistor functioning as a temperature sensor. Ultrasound imaging systems monitor the temperature at the surface of the transducer. Should the value exceed 41 o, the transducer is switched off to avoid heating tissue up. The signal transmission cables terminate at a 260-pin CANNON ZIF (zero insertion force) connector which can be mounted onto the ultrasound imaging system. The connector features an ID EPROM which contains the identification information of the probe. The system corresponds the identification information with the probe-specific system parameters using a look-up table. In case of a faulty or wrong ID EPROM at the connector, the system will either not recognize the probe or load operational parameters of a different probe to use with the one mounted. This may result in abnormal use of the probe and thus potentially unrealistic and distorted images. 2.3 Prototypes of the fast rotating probe for 3D TEE Three prototypes were developed for the needs of this project. The multiplane TEE probe has been used as a starting point for further development of all prototypes. The second prototype significantly differs from the first one in terms of engineering as well as philosophy of approach. The third one is mainly an adjusted/upgraded version of the second prototype with added functionality, based on the requirements advocated by the ExpEcho group.

29 17 MATERIALS AND METHODS Figure 12: Diagram of proposed FRTEE concept. Acquisition of the image data with simultaneous ECG recording begins. Acquisition is synchronized with initiation of the motor. The motor rotates the mechanism of the probe bi-directionally at a frequency of 5Hz. The angle sensor measures the angle of rotation. Angle values are stored synchronously with acquisition of image data. Angle data and image data are downloaded to the laptop where the reconstruction software will produce the 4D reconstructed image data sets. The diagram presented in figure 12 shows the concept for rapid 3D TEE as it was originally discussed. The addition of the SYNC block and the feedback of the angle sensor for motor control were actually implemented only after the first prototype was designed First prototype The design of the first prototype was based on general brainstorming among members from the ExpEcho group and Oldelft Ultrasound. Discussions aimed to determine what would be needed to initially attest and confirm that fast-rotating 3D TEE is possible. Previous fast rotating TTE designs featured a constant rotational speed of 4-8 Hz. The limitation of the TEE probe s rotation to the range of led to a proposed setup where the transducer array would rotate fast in alternating directions. The angle of the transducer array should be measured throughout acquisition, in order to be able to position accurately the acquired data in 3D space. As can be seen in figure 13 a bulky construction assists the mounting of a motor onto the handle of the probe. The motor rotates a wheel in one direction. An eccentric rod, however, translates the unidirectional rotation of the motor and wheel to a bidirectional motion of the knob on the handle. The translation is similar to the translation of reciprocal motion to wheel rotation in the steam engines of old trains. The function of the large knob is modified to directly rotate the transducer in both directions. In this way, the motor rotates the transducer array clockwise and counterclockwise, allowing it to acquire

30 CHAPTER 2 18 slices of the heart throughout the full range. The element array rotates through a range of roughly but the axis of rotation is the center of the 2D frame. Every rotation of the transducer covers two halves, each of 180 0, for the left and right parts of the acquired 2D images. Figure 13: First prototype; mechanism for transducer array rotation and angle measurement. A thin, cylindrical, diametrically magnetized magnet is fixed over the central rotation axis of the knob at the handle. A programmable magnetic encoder by Austriamicrosystems AG, Schloss Premstätten, Austria is positioned over the magnet to read out the angles of the different positions of the knob s rotation. The magnetic encoder integrated circuit (IC) uses spinning current Hall technology for sensing the magnetic field distribution across the surface of the chip. [22] The IC is used in the serial data transmission mode with 10 bits of the transmitted 16-bit word reserved for the angular position. A control box is also developed by Oldelft, along with the probe, for controlling the motor, storing the angle data and delivering it to a computer so as to be used for the reconstruction of the 3D images. The control box requires a 15 Volt power supply. It has one 9-pin connector for supplying power to the motor, controlling the magnetic sensor data acquisition (latch and clock inputs) and reading its serial output (acquired data). Furthermore, two BNC inputs are labeled Start and Frame. The first input triggers the beginning of rotation and angle acquisition while the second one is simply recorded along with the angle data (as high or low level, 1 or 0 ). The initial idea was that triggering angle acquisition per acquired frame would be feasible and adequate to synchronize angle values with image data. Control of the motor is realized by a programmable intelligent computer (PIC) [24] and a field-effect transistor (FET) which controls the power supply of the motor. A

31 19 MATERIALS AND METHODS RabbitCore RCM3700 C-programmable core module [23] is included in the box. The module features the Rabbit3000 microprocessor which is running a preloaded program. The microprocessor is embedded on a network card for interfacing via a UTP cable with the host PC. The program controls the PIC for motor start/stop. It waits for a trigger signal ( Start ) which signifies the start of image acquisition. Then, the motor is started and storing of the magnetic sensor s angle measurements on the module s RAM is initiated. Sampling of the angle measurements is performed at 2.5 KHz. Simultaneously, the logic state of the second trigger input ( Frame ) is monitored and recorded for the purpose of synchronizing angle and image acquisition. A program written in the Delphi programming language is provided by Oldelft and runs on the host PC. It communicates with the microprocessor, acknowledges the new angle data storage before acquisition is triggered and later retrieves the stored data and presents it in a Microsoft Excel-compatible format. The Excel file consists of three columns of data, the first one being a fixed column where time is represented in increments of 0.4 milliseconds (2.5 KHz sampling). This time axis will be used to correlate angle values with image data. The second column has the measured angle values per time increment. The third one presents a 1, for all time increments from the beginning of the acquisition of angle values that the Frame input line remains at logic high. It, then, switches to 0 at the time increment of the next angle value sample, after the input line has switched to logic low. Conducted work Ideally, a distinct angle value per acquired image line should be recorded. The original idea was that a frame signal would be available or derived from the System FiVe which triggers a pulse for every frame acquired. We located a small coaxial connector on the boards of System FiVe which gives a pulse per transmitted ultrasound beam thus, per image line pair, due to MLA during image acquisition. We deliberated an approach where this line signal is used as the Start input signal for the control box. Related work is described in a separate internship report. [25] As soon as the Unfreeze button on the user panel is pressed, storing of angle values is initiated. The concept seemed viable but no guarantee can be given for the accurate timing between angle- and image-data acquisitions. The sampling of angle values is expected to begin simultaneously with image data acquisition but this may well not be the case. Therefore, the first pulse of the line signal further triggers a delay circuit which instantly (in the order of nanoseconds) gives a logic high output. After an accurately specified period of time, the fixed delay time, it returns to logic low. This second signal is used as the Frame input. In the case that angle acquisition, and thus recording of the Frame input, starts later than the image acquisition, the Frame input in the third column of the angle-data file will not be 1 for the entire predefined fixed delay time. One can accurately measure the delay time between the initiation of the image acquisition and that of the angle measurement by calculating the difference between the fixed delay time and the time

32 CHAPTER 2 20 that the Frame input seems to have switched to zero in the Excel file. The calculated delay has a maximum accuracy of 0.4ms, limited by the 2.5 KHz sampling. A relatively accurate correlation between lines and corresponding angles can be obtained. Limitations of the first prototype Experimenting with the first prototype, 3D images were reconstructed from 2D images acquired along with angle values corresponding to specific time-points. Values of the recorded angles were interpolated as explained later in the description of the reconstruction code. These preliminary results were not satisfactory but were considered promising in terms of showing identifiable views of the imaged object. They also gave further insight as to potential problems and issues that had to be addressed. Further, there was no motor control so as to adjust the motor s speed in order to obtain an approximation of a triangle. The triangle refers to the waveform that consecutive angle-values ideally create. A steep, rising curve for rotation in one direction and a steep dropping curve for the opposite direction are desired. At the two extremes as few values as possible are desired, as if transition from one direction to the other was instantaneous. This would give the most regular distribution of angle beams in rotation angle. The angle waveform for the prototype was almost sinusoidal. The most important issue with the first prototype was hysteresis in the angle measurement. Measurement is performed at the axis of rotation of the knob rotating the transducer through the wires/pulley mechanism and not directly on the axis of rotation of the transducer array. Such translation of rotational motion by wires may exhibit hysteresis, especially with a bent tube. In fact, hysteresis had been measured by Oldelft suggesting a value of ±10 0 for a straight tube. The findings following the experiments, especially in flexion angles resembling clinical use showed that the value could be significantly bigger, varying for different tube flexions. Angle measurement should essentially be performed right at the tip of the probe, at the axis of rotation of the transducer array Second prototype The second prototype used for the project is superior to its predecessor. The motor is incorporated in the handle (Fig. 15b) alleviating the need for the bulky construction around it. The knob and small wheel underneath are both functional for probe flexion in all four directions. Most significantly though, a magnetic sensor is integrated at the tip of the probe for measurement of rotation of the transducer array. The design features a gear mechanism, seen in figure 14, attached to the pulley which rotates the transducer array. The gear translates the rotation of the transducer to a 281 o range of gear rotation. A magnet is fixed over the gear. The shaft encapsulating the transducer array is slightly thicker (12.25 mm) than that of a standard multiplane TEE probe to accommodate the magnet and the sensor. Cables for providing power supply to both the motor and the

33 21 MATERIALS AND METHODS magnetic sensor and those for communicating with the sensor and transferring angle data, go along the length of the standard cables, inside the tube coating. They exit from the connector s casing to a 9-pin D connector which is connected to the control box. Figure 14: Magnetic sensor for angle measurement integrated at the tip of the probe. The gear for translating the transducer rotation to a rotation is situated left from the pulley and in between the wires for pulley rotation. The magnetic sensor is behind the gear in that photo. The wires connected to the sensor are visible. (Photo courtesy of Oldelft Ultrasound B.V.) The magnetic sensor, an ic-ma angular Hall sensor from ic-haus GmbH, Bodenheim, Germany, is functioning based on the same physical principle as the IC used in the previous design for angle measurements. [26] An analog-to-digital converter with a resolution of 8 bits is used to convert the variation of the magnetic flux passing through two of the four Hall sensors (diagonally facing each other) in a digital signal. A resistor mode of operation, exploiting the taps of an integrated resistive divider, is utilized for reading out the values of the angles. The value of the absolute angular position acts as a wiper and selects one of the 256 taps on the resistor chain. In principle, this can give a resolution of ~ The measurement, though, is performed over the magnet which is fixed on the gear. In this way, a resolution of ~0.9 0 is achieved for measuring the translated rotation of the transducer array. The mechanical tolerances of the design of the structure are minimal and the specific design should allow a hysteresis-free angular readout. The control box is also modified for the requirements of the new probe. Since the motor is incorporated inside the handle housing, it is capable of rotating in both directions. For that reason, a FET driver, consisting of two pairs of FETs, is implemented. The angle-signal is low-pass filtered, analog-to-digital converted (ADC) and fed to the PIC that controls the FETs to provide position feedback of the motor. This allows the motor to maintain a specific rotation angle pattern over time, approximating a triangular pattern. A second low-pass filter (LPF) and ADC give a digital angle output which is stored in the RAM of the microprocessor module.

34 CHAPTER 2 22 a) b) Figure 15: a) The second prototype is functional for clinical use. b) The motor is integrated in the handle of the probe. (Photo courtesy of Oldelft Ultrasound B.V.) Limitations of the second prototype Initial results of 3D reconstructed images from data acquired using the second prototype FRTEE probe were considerably improved. Certain concerns, though, came to the foreground. The most profound issue was that of a form of interference, manifesting on the 2D images acquired by System FiVe as a moving dot pattern. The white dots appeared as soon as the motor was initialized and throughout the entire period of its rotation. The clock signal of the PIC together with the power supply of the motor and the angle sensor were considered responsible for the interference. At lower gain settings interference was not significantly affecting the quality of the reconstructed images and the assessment of the probe s 3D function. During a later upgrade of the probe the topic could be revisited. Furthermore, resulting 3D images showed angular errors. We speculated that these errors occurred due to delay between the angle signal and the acquisition of image data. The low-pass filtering at a cut-off frequency of 50Hz caused a phase-shift and a deformation of the angle signal. It was considered therefore a source of delay. We proposed to remove low-pass filtering of the angle signal and acquisition of discrete angle values per transmitted beam to ensure a delay-free angle signal. Two final issues about the functionality of the probe were related to the actual operating procedure of the control box. Firstly, the delay imposed by the electronics of the control box was variable. Sometimes the delay pulse coming from the delay circuit was not long enough to be recorded together with the acquisition of the angle data. In such case, synchronization of angle-values with image data was impossible which led to repeating the experiment. In a clinical situation it is unacceptable to have a patient wait longer for a repetition. Secondly, it became obvious that the direct and sole coupling of motor rotation and angle acquisition to the imaging acquisition inflicts a practical handicap. The opera-

35 23 MATERIALS AND METHODS tor cannot unfreeze the image to view the imaging plane prior to acquisition without literally powering off the control box to avoid simultaneous motor rotation. Furthermore, the fact that the transducer rotates only in the fast-rotating mode confines the examiner s perception of the imaged area. The effect was largely underlined by examiners during in-vivo experiments with pigs described in chapter 4. Examiners indicated that it should be possible to rotate the transducer before beginning image acquisition for 3D reconstruction. This helps to orientate themselves and assess the range of structures that will be imaged Third prototype The third and final prototype designed by Oldelft to be used for the project addresses all previously noted problems and has consequently led to accurate measurements, resulting in satisfactory three-dimensional image datasets. At the same time, it features increased functionality for the user, taking into account its operation in the clinical setting. The prototype is to date the closest design to a customer product. Adjustments, modifications and upgrades The modifications of the third prototype are small-scale compared to the redesigning and adaptations between the first and second prototypes. A major modification, nonetheless, concerns the allocation of hardware components. The prototype features a printed circuit board (PCB) inside its handle, where the control components of the motor are located. Therefore, the PIC and driving FETs are integrated on the handle of the probe as seen in figure 16. Such integration was considered essential to isolate the path of control pulses running from the PCB and driving FETs to the motor from signal wires, alleviating interference. The idea is based on previous experience with probe designs and indeed proved fruitful. The power supply wires were placed outside the covering of the RF signal wires. The images acquired with the new probe were practically interferencefree. Figure 16: Motor control electronics on a PCB, integrated in the handle of the probe. (Photo courtesy of Oldelft Ultrasound B.V.)

36 CHAPTER 2 24 To tackle the problem of inaccuracy imposed by delays between acquired angle values and image data, our original proposal to record one angle value per transmitted ultrasound beam was implemented. The network-card-embedded microprocessor uses the line signal from System FiVe as an input interrupt. One angle value is stored on every interrupt pulse; therefore the angles measured correspond directly to the acquired image lines. The acquisition of angle values is, thus, essentially synchronized with the image data acquisition. It is also sampled at the pulse repetition frequency of System FiVe, solving any under-sampling issues. In addition, the low-pass filter mentioned earlier has been removed. Oldelft Ultrasound further enhanced the functionality of the Delphi-based program. The program does not use an internal clock for defining the acquisition time. An additional input parameter of total number of beams for which angle values should be acquired allows the user to determine the total running time of acquisition of angle values. The program does not have direct information about the pulse repetition frequency used by the imaging system. In the case of a real PRF value of 6717,17 Hz, the user can input a number of lines, running the entire angle-data acquisition for 10 seconds. For the majority of experiments with the third prototype lines were used. The second added feature allows for selection between an Excel file and Tab-delimited text file as an output for the angle values. The latter facilitates the case of PRFs higher than 6553,6 Hz since MS Excel has a limit of rows being stored in one file. As for the improvement of the probe s clinical functionality, the handle of the third prototype features two buttons for slower, left-right rotation of the transducer array. The fast-rotating mode of the probe is decoupled from the unfreeze mode of System FiVe. Angle acquisition along with fast rotation starts only after both the Start button of the Delphi program and the Unfreeze button of System FiVe are pressed. Clinicians can first rotate the imaging plane of the transducer in both directions to identify the structures of interest. They can then Freeze the system, press the Start button of the program and then start the acquisition procedure by pressing the Unfreeze button on the imaging system. In this way, the first angle data acquired corresponds to the first line recorded in the stored clip file. Finally, the transducer array was changed because there were compatibility issues with System FiVe and the thermistor used with the previous transducer array. This allowed for clinical use of the probe. The transducer array used in the third prototype is a narrowband, 40% bandwidth transducer array. 2.4 Considerations on data morphology The morphology of the acquired image data is one of the factors that define the complexity of the reconstruction procedure. Understanding the distribution of the acquired image data in space allows for observations which can lead to optimizing both acquisition and reconstruction. In order to understand the characteristics of such morphol-

37 25 MATERIALS AND METHODS ogy, we describe an acquisition of image and angle data at a certain depth of the 3D volume. One can then refer to the samples acquired at a specific depth over the entire scanned 3D space. Physical data resolution The physical data resolution is determined by the ultrasound beam width in 3D. This is calculated from the aperture of the transducer [4] and is 1.6mm at focus depth for the transducer we are using. This is assumed to hold both for the lateral direction in-plane and perpendicular to the scanning plane. This sets the standard for the division of samples over the scanning volume: a higher density of scan-lines will not supply additional data, and a lower sampling will results in data loss or aliasing. For example, considering the circle at the border of the 3D volume elevation angle of 36,9 o, at focus depth the maximum transducer rotation angle between two consecutive samples on that circle should be about 2.2 o to achieve the 1.6mm resolution. Ideal and achievable data distribution Since the number of acquired ultrasound beams per second is limited, we want to distribute them evenly to achieve a uniform resolution and highest volume rate. Ideally, the image data should be distributed evenly and regularly over a c-plane, i.e. the density of samples per square millimeter would be constant over the plane ( ideal distribution). In our setup, with the rotating scan plane and a uniform distribution of the beams over elevation angle per frame (determined by the ultrasound machine), this cannot be achieved. The best achievable ( expected ) distribution for this case would thus be a uniform distribution over rotation and elevation angle, which will give a high 2D sample density in the center of the cone and a low density near the borders. In practice, the ( actual ) distribution over rotation angle is both non-uniform (because of variation in rotation speed, deviation from the ideal triangular rotation angle pattern) and irregular (because of lack of correlation to cardiac phase). In this section, we quantify the differences between the ideal, expected and actual distributions Example of actual distribution A typical 9-second acquisition is considered, with transducer rotation at 5Hz and an acquisition of image frames at a rate of ~108 fps each frame is produced from information of 62 transmit pulses. The reconstructed 4D space will be re-sampled in time at specific points within the cardiac cycle, to reconstruct 3D volumes at specific time-points of one heart beat. For every cardiac cycle of the entire acquired data, only the samples which are around the same cardiac cycle time-point will be used to reconstruct each one of the 3D volumes. Therefore, it is interesting to consider the morphology of the acquired data around one of these time-points, for every cardiac cycle, as the rest of the data will not be used for this specific 3D volume. Further, for a heart rate of 80 beats per minute, a cardiac cycle lasts 0,75 seconds. Every cardiac cycle of this 12-beat data acquisition is divided into 16 parts of 47 ms each. In our case, we will have 3780 samples per phase.

38 CHAPTER 2 26 We consider a c-plane at a depth of 55mm. Max elevation angle is 36.9 degrees, so cone area corresponds to mm 2. Ideal density is therefore 0.7 samples /mm 2. In other words, an area of 1.43 mm 2 corresponds to every sample. Considering the beam width of 1.6mm, beam area is about 2 mm 2 so a uniform density of 0.5 samples /mm 2 would still be sufficient. Through this perspective, we consider the sample distribution both in terms of mm 2 / sample (to compare to the ideal distribution) and in terms of samples per elevation/rotation cell (to compare to expected distribution). We divide the elevation into 31 bins of 1.21 degrees (1 per image line in the half image) and rotation into 120 bins of 3 degrees. This amounts to 3720 elevation/rotation bins and an expected density of about 1.02 samples per bin or 0.027% of the total number of samples. b) Figure 17: a) Sample distribution of acquired data in 3D space, at a 55mm depth level for one cardiac phase. Dense sampling is observed around the 0o rotation angle due to the non-ideal rotation of the transducer. Additionally, the central part (small elevation angles) is oversampled due to the scanning pattern of the transducer. b) Percentage of samples per rotation and elevation bins. The flat area at a level of % shows the ideal distribution of samples over the entire cone area. Figure 17a shows the distribution of samples (used for the reconstruction of one 3D volume) from a slice of the scanned 3D space, at the depth of 55mm. Certain areas show significant gaps between samples. The center of the 3D volume is densely sampled compared to the outer regions where the samples are further apart from each other. The areas around the 0 0 (or ) and rotation angles are densely sampled as well.

39 27 MATERIALS AND METHODS b) Figure 18: a) Effect of irregular transducer rotation: Sample density is higher at the extremes of the rotation, compared to the rest of the rotation. The green line shows the case of an ideal (uniform) distribution. b) Area around each sample for each elevation angle. Yellow level is the minimum area covered per beam (2 mm 2 ). Green level is the area of 1.43 mm 2, in the ideal case that samples are uniformly distributed. The effect can be clearly seen on figure 18a where the distribution of samples per rotation bin is shown. The irregular rotation of the transducer the motor decelerates to switch direction and accelerates again prevents an even distribution of samples over rotation angles. In the case of a uniform distribution the samples would be evenly distributed over the 120 bins, resulting in a fixed value of 0,8333 % of samples per bin. Excluding the peaks, which amount for about 31% of the data, a mean of is calculated. The area around a sample is shown to be smaller, close to the axis of rotation, at the smaller elevation angles (Fig 18b). This means that the gaps among samples are significantly bigger that desired at larger elevation angles. The yellow-delineated level shows the area covered per beam (actual) while the green one shows the ideal coverage. On the other hand, the distribution of samples over the different elevation angles is almost uniform as can be seen in figure 19a. The value per rotation angle, however, can deviate significantly from the mean value. A coefficient of variation (S.D. / mean) of almost 100% is found. A ripple is observed because for each of the different cardiac cycles which the data for this 3D volume come from, the scanning beam is at an arbitrary point. It is interesting to observe the distance of the samples from their nearest neighbors as opposed to the distance of one sample from another in the rotational direction. Figure 19b shows that for the entire 2D slice of samples at the specified depth, the majority of samples are, at worst, almost 0.04mm apart from their nearest neighbor. Only a few of the samples have a distance close to 2mm from their nearest neighbor. It should be noted that these neighbors are found in the elevation direction (see fig. 17a) while the neighbors in the rotational direction may be much further away. On the other hand, if we consider two

40 CHAPTER 2 28 samples at the border of the 2D slice, adjacent in the rotational direction, we obtain a maximum distance of almost 9.5mm a) b) Figure 19: a) Percentage of samples per elevation angle for one rotational sector of 3 0. The peaks at the extremes of the rotation are not taken into account for the calculated S.D. b) Distance of a sample to its nearest neighbor, considering the entire 2D slice Conclusions on data distribution From the above experiment, we can draw a number of important conclusions: Rotation speed variation accounts for a loss of ~31% of the expected density in 95% of the rotation angles and volume. Irregularity of distribution amounts to a coefficient of variation of almost 100%. An adapted distribution of pulses over elevation angles might result in a uniform spatial distribution up to 37 degrees. That density level is now only achieved for elevation angles of 11 degrees and lower as shown in figure 18b. Although the distances to the nearest neighbor are below 2 mm, the holes at the border of the cone can be up to 5 times larger than the desired distance in the rotational direction.

41 29 MATERIALS AND METHODS 2.5 Reconstruction software The reconstruction software has been a tool in its entirety rather than the outcome of our work. It was developed on the C++ programming language. It makes use of an open-source image processing and visualization toolkit (National Library of Medicine Insight Segmentation and Registration Toolkit). The original software was used for 3D and 4D reconstruction of datasets from the fast-rotating precordial probe [12] also developed by the Experimental Echocardiography group. Marijn van Stralen, the original developer of the software, initially adapted the existing code to fit the original demands of the FRTEE project. Nevertheless, we further adapted the software to solve specific misinterpretations which became obvious during experimentation. Our work on the reconstruction software also involved its adaptation to accommodate the three different probes and experimental setups used, along with respective parameters. Initial adaptation concerned the alternating rotation of the transducer and handling the positional information of the acquired image lines in the rotational direction. Further, the software was adapted to handle image data from the two rotation directions separately. Image data acquired during the anticlockwise rotation of the array, tagged as Upgoing, comes from increasing rotational angle values, from 0 0 to Image data acquired during clockwise rotation, tagged as Down-going, is assigned decreasing angle values, from back to 0 0. A new input parameter about the frames used for reconstruction gives the option to choose between frames from the up-going direction, the down-going one or from both directions. The feature proved useful for gaining insight on the problems of the reconstructed sets especially regarding the angle signal delay and the initial hysteresis problems. Input of the software The software uses the clip file (containing the recorded ECG, 2D image data in polar format, and information on depicting this data as a 2D Cartesian sector image) as its primary input. The order of transmitted beams to acquire image lines is also needed for positioning the acquired image lines at the proper rotation angles. In the example of interleaved beam firing, the beam order is crucial for precise positioning of all image lines. A text file containing the beam order ( theta order file) is also used as input for the software. An index file is the third input file of the software, containing information about the frame-rate and the total number of acquired frames to be used for the reconstruction. Since acquisition of image data does not stop together with acquisition of angle data, frames recorded after angle data acquisition has stopped need to be discarded. A parameter file can be used as the final user-adjustable input file, defining the values of certain parameters of the reconstruction. Such parameters which affect the positioning of the image data include switching of the direction of rotation, flipping of the elevation angle and switching of the beam order. The first two parameters mainly concern the visualization of the reconstructed volume. Beam order switching was implemented

42 CHAPTER 2 30 because it was not completely certain whether the first transmitted beam (thus, the first image line) is actually fired (acquired) on the left or right side of the 2D sector. Reconstruction process Since the heart s motion is practically periodic for a certain period of time, a full data acquisition of multiple heart cycles can be combined into one composed heart cycle. This can be done by defining the individual heart cycles and resampling the data at distinct time-points between consecutive heart cycles. These time-points are termed cardiac phases. The cardiac phase is the relative position of data between two consecutive R- peaks. The program reads the ECG data from the clip file and performs peak detection on the ECG signal to define the recorded cardiac cycles. The user has the option to manually mark the ECG peaks, correcting for any sort of irregularities or mismatch that automated detection induces. Each individual line of each frame is assigned its corresponding rotational and elevational angle values. The position of lines in time is defined by the number of the frame they belong to, their serial beam order number and the frame-rate of the acquisition. This time-stamp of each line is converted to a respective cardiac phase, depending on its position between two consecutive R-peaks. Three-dimensional interpolation of the sparse, irregularly distributed data is performed by applying 3D normalized convolution in the dimensions of elevation θ, rotation φ and time τ (thus a 3D space) for every depth level of the image data. The depth dimension is densely sampled and is therefore left out of the interpolation task. 3D normalized convolution is a complex method for interpolation [33] but a simplified description is attempted. A certainty space is generated corresponding to the spatial distribution of the samples in this 3D space. A maximum certainty value of 1 is applied for the points of the data space where a sample value exists. The rest of the points are given the value of 0. For each point to interpolate in the 3D space, a Gaussian distance-weighted sum of all available data in the neighborhood is taken. The same procedure is applied on the certainty space. The interpolated certainty space is used to normalize the image-data interpolated 3D space and obtain the resulting, dense, interpolated data set. The extent to which an individual image sample will blend out and redistribute its value in its environment (dimensions of θ, φ and τ) is controlled by the variance of the Gaussian filter used. This precise parameter of the interpolation method can be termed as smoothing. Large smoothing is required especially at locations with few samples to define the situation in the space among them. Further, the program samples at distinct points in the time dimension to obtain image data from certain phases of the cardiac cycle. For each defined cardiac phase, the resulting interpolated (θ, φ) planes for all depths are then converted into Cartesian voxel sets. The total number of cardiac phases

43 31 MATERIALS AND METHODS used defines the temporal resolution of the reconstructed 3D volume sequence in time. Increasing the number of phases used provides more discrete time-points in the period of the heart s motion, thus more temporal information about this motion. Fundamental work performed on the reconstruction code The code was cross-checked during experimentation with 3D and 4D phantoms for errors in the reconstruction procedure, based on the resulting images and problems that they showed. We expanded the parameter file to assist in the automation of the reconstruction process. Parts of the code were adapted to support the definition of the angle data used. This was essential work for switching between the probe designs. The second probe design, for example, used angles which belonged, in principle within the -90 o +90 o range of values, while the latest probe design gives an output of values between ~10 o and 190 o. Finally, the way the rotation angles were assigned to the individual lines of acquired frames had to be adjusted. The first two prototypes featured a sampling rate of angle values lower than the beam transmission frequency. Linear interpolation between consecutive angle values was used to calculate the angle of beams acquired at a certain time point. We adapted the code for the latest prototype, where recorded angle values directly corresponded to the transmitted beam. Some further points which required additional modification of the code on our behalf are described in the next chapter along with justification for such changes. 2.6 Phantoms for laboratory experiments To assess the functionality of the aforementioned probe prototypes in acquiring useful data for three-dimensional reconstruction we crafted specific structures mimicking ultrasound response properties of tissue. The structures, known as phantoms, were of different shapes resembling coffee cups, rectangular blocks and a toy-dolphin. Phantom preparation and the motivation underlying their shapes are further described in the next chapter. Resulting images of the phantoms are discussed and assist in further comprehending and justifying the choices. Using these phantoms we initially validated the operation of the software for reconstructing 2D images and angle data into 3D datasets. Nevertheless, in the clinical application the imaged structures are moving periodically following the heart s beating rate. Therefore, before proceeding with in-vivo experiments, the probes should actually be tested with moving phantoms or fourdimensional phantoms. A 4D phantom is an object (i.e. a 3D phantom) which performs a repeatable periodic motion in time. We constructed a 4D phantom setup for the requirements of the project, using as an imaging object a 3D toy-dolphin phantom. It was used to obtain preliminary results in

44 CHAPTER 2 32 four dimensions (periodic motion of the 3D object in time). The dolphin phantom was fixated at the tip of an electric screwdriver and could rotate at a constant speed. Further, a Björk-Shiley valve phantom was redesigned by Cees Pakvis (ExpEcho) and modified by Wim van Alphen and Geert Springeling from the Erasmus MC Precision Mechanics Workshop using a water tank and a balloon as the imaged object. The existing pump was utilized to inflate and deflate the balloon with water. Both setups are described at a later stage. 2.7 Visualization software Three programs were used for 3D visualization purposes in this project. We used MeVisLab by MeVis Research GmbH through the entire project for visualization of 3D images of motionless objects and browsing through 2D slices of the 3D image data sets. For visualizing 4D datasets, we used Qlab by Philips Medical Systems and Echoview and Image Arena by Tomtec (TomTec Imaging Systems GmbH). 4D data visualization was done in cooperation with clinicians of the hosting medical center.

45 Chapter 3 In-vitro imaging The actual goal of the project has been three-dimensional ultrasound imaging of the human heart. Several experiments were conducted in advance, outside the human body, using structures and objects of different shapes. Since fast-rotating TEE is a novel concept, we first needed to verify that the acquired data of imaged objects can produce accurate 3D representations of these objects in terms of size and shape. Motion of objects in time was at this point excluded. The software for reconstructing the acquired data had originally been developed for the fast rotating TTE probe. [12] The continuous rotation of the fast-rotating unit (FRU) as well as the calculation of speed significantly variegates the original version of the software from the one modified for FRTEE. Rotation of the FRTEE is bidirectional and positional information of the acquired data is required. Verification of proper software function was essential to conclude that the approach implemented is correct, presenting sound results. The imaged objects were made out of tissue mimicking material. These tissue mimicking phantoms are placed in tanks filled with water during experiments to facilitate ultrasound propagation. The material is considered to have equivalent acoustic properties to tissue (in terms of scattering, attenuation etc.), allowing for improved quality of ultrasound image representation. The different shapes of phantoms used are described further in this chapter. The most prominent experiments conducted are outlined and results that led to changes in the acquisition and reconstruction procedures are presented. The experiments also allowed us to familiarize with ultrasound imaging, acquisition and perspective of images and structures there-in. We further gained valuable experience to be able to pinpoint certain aspects of interest which might be revised in the procedure of the reconstruction. After obtaining promising results from imaging and reconstruction of 3D objects and implementing improvements, the next step was the imaging of moving objects. These experiments were closer to the physiological application of the device in humans, offering insight on the actual 4D reconstruction procedure and potential problematic points. Section 3.2 discusses two setups of four-dimensional phantoms, objects which perform 33

46 CHAPTER 3 34 some form of periodic motion and produce a time-sequence of 3D reconstructed images. Experiments with the different phantoms are described in section Three-dimensional phantoms Certain phantom structures such as blocks made of tissue-mimicking material (TMM) were readily available for experimentation, already used during preliminary experiments with the first probe prototype. We decided, though, to create new 3D phantoms to be able to visualize different aspects on the 2D and 3D images for further familiarization with the probe and the reconstructed results. Tissue-mimicking material from the components shown in figure 20 was prepared from a standardized recipe [27] and cast in several molds. The 3D images in this section are reconstructed by imaging the respective objects with the transducer array facing the phantom on the front side, the one seen in each enface reconstructed image. The only exception is the dolphin phantom which is imaged with the transducer imaging the dolphin phantom from the belly to its back. Component Weight % (pure components) Glycerol 11,21 Demineralised water 82,95 Benzankonium Chloride 0,47 SiC (400 mesh) 0,53 Al 2 O 3 (0,3µm) 0,88 Al 2 O 3 (3µm) 0,94 Agar 3,02 Sum: 100% Figure 20: Table with the components used to produce tissue mimicking material. [27] Coffee-cup phantom A tall coffee-cup is used in which TMM is cast to give a cylindrical structure which can then be placed in the water tank.

47 35 IN-VITRO IMAGING a) b) Figure 21: The cup (a) used as mold for making the phantom and the 3D reconstructed image (b). The phantom is tall enough to image its upper part, leaving out of the imaging view the lower part. In this way the bottom surface of the tank is not imaged. It had earlier been observed that the tank walls gave large reflection. The phantom is put upsidedown to image the bottom rim detail for additional structural information. It was also later used as support to position other phantoms at a certain level from the bottom surface of the tank Dolphin phantom Following early experiments with large blocks made of TMM and those with the coffee-cup phantom, we understood that phantom shapes should provide visual information of their orientation in space and relatively to the apex of the 2D sector image. In this way, potential spatial problems concerning one or both halves of the plane of acquisition can be examined and spatial position of artifacts affecting large parts of the 3D image can easily be isolated. a) b) Figure 22: Photo of the dolphin phantom (a) and the respective reconstructed 3D image. The fin on the back of the dolphin, the tail, the head and the protruding support cane on the right side of the dolphin were used as landmarks to define orientation on reconstructed images. While looking for objects that would fulfill such shape requirements, a two-piece mold was crafted by J. G. Bosch by using a plastic toy dolphin. We prepared the mold with a layer of grease on both sides. A hole was drilled on the one side, at the right fin of

48 CHAPTER 3 36 the dolphin to provide an orifice for filling the mold with tissue-mimicking material. We placed the plastic body (piston) of a thin syringe through the hole to serve as a vertical handling rod. The needle of the syringe was pierced through the piston of the syringe to provide additional internal support for the dolphin in the direction perpendicular to the piston. The two pieces of the mold were then united with each other. TMM was prepared and poured inside the mold. After the TMM cooled down and congealed the two parts of the mold were separated to release the dolphin phantom. The dolphin phantom was a successful 3D phantom and, as will be described later, was also used for our first 4D experimental setup. Its success refers to its extensive use throughout experiments and the insight that such experiments provided. It is attributed mainly to its shape which is unique and asymmetric in all three dimensions of space meaning, the head is distinguishable from the tail, the belly from the back with the fin and the right from the left side due to the protruding support cane Balloon phantom in 3D The balloon phantom is actually part of the four-dimensional balloon phantom setup. Imaged as a motionless 3D object, it is a regular latex balloon, filled with water, its opening fixed at the small orifice of a larger cylinder which can function as a pump. For the 3D experiments the pump function was merely used to initially inflate the balloon s volume with water to a desired point. The first experiments were done in 2D to assess the possibility and further the quality of visualization of the balloon as an ultrasound phantom. a) b) Figure 23: Balloon phantom (a) and 3D reconstructed image of the balloon (b). A cut through a certain point of the volume of the reconstructed balloon can be observed. This is because of the angle of incidence of the ultrasound beam to the balloon surface at this point

49 37 IN-VITRO IMAGING The 2D images that were obtained show indeed sharp slices of the balloon. An artifact from far echoes from the water tank wall is present at larger depths of the 2D sector image. Due to the low attenuation in the depth direction of the transmitted beam, the signals from these depths return to the next image lines. At a level from the balloon s center, the balloon wall is faintly reflecting the ultrasound beam, as can be observed in figure 23b. This can be expected since at this point the ultrasound beam is almost tangential to the balloon surface, generating a weak scattering signal Rectangular block phantom The block phantom was not cast, as in the cases of the coffee cup and the dolphin phantoms. It was actually cut from a pre-existing, larger block phantom. The shape was cut as rectangular as possible, aiming to image the right angles formed by the block s sides. En-face imaging of the block s sharply defined shape would provide understanding of potential angular errors between the two directions of rotation. a) b) Figure 24: The block phantom and its reconstructed 3D representation 3.2 Four-dimensional phantoms As mentioned in the previous chapter, it was essential to perform some reconstructions of periodically moving objects. In such way, we could assess functionality of the reconstruction software in a situation similar to the physiological one, before conducting in-vivo experiments Rotating dolphin phantom One of the dolphin phantoms crafted for the 3D phantom experiments was attached via a syringe shaft to the tip of an electrical screwdriver (IKEA products). The syringe shaft served as a handling grip for securely positioning the phantom as required. We fastened a stripe of black paper with a vertical white line of 1mm width around the rotat-

50 CHAPTER 3 38 ing tip. A reflective photomicrosensor by Omron [28] is sensing the change of reflective properties between the black and white parts of the paper stripe during every rotation. The circuit that we developed gives an output pulse for every rotation of the screwdriver s tip. The signal is fed to the ECG leads of System FiVe for simultaneous recording along with the image data. In such way, an input signal similar to that of a real ECG of a patient is made available to be used with the reconstruction program. a) b) Figure 25: a) Rotating dolphin 4D phantom setup. b) The tip of the electrical screwdriver. The photomicrosensor over the tip, as seen on the image, emits light by using a built-in LED and senses the light reflected by the stripe of paper. The white line reflects emitted light different compared to the black part of the paper. While the setup is improvised, it provided an opportunity for assessing the fastrotating probe in an experiment with a periodically moving object, immediately before the first in-vivo experiments with pigs. The rotation speed of the phantom when immersed in water is ~2Hz Balloon 4D phantom setup The balloon four-dimensional phantom setup was also used. The motion of the balloon during inflation and deflation is similar to the motion of the heart in terms of the amount of wall displacement. The balloon 4D setup was used to further understand the effect of different parameters on the reconstruction of 4D data sets and discuss potential improvements.

51 39 IN-VITRO IMAGING The balloon setup is a modified setup of a Björk-Shiley valve phantom which was available. The structure mimicking the left ventricular function is removed and a tank is placed instead. An actuator translates its linear motion to a piston which slides inside a Perspex 3 tube, forming a pump. The tank and the tube with the piston communicate through an opening. After they are both filled with degassed water, the balloon is fixated around the orifice of the opening on the side of the tank. Then, the piston moves forward and the balloon is filled with water from the tube and consequently inflates. This is the stationary position of the balloon as shown in figure 26. Figure 26: Balloon 4D phantom setup. The actuator on the right moves the piston in the Perspex cylinder. The cylinder is terminated on the left side at a small opening on which the balloon is fastened. The pumping function of the piston inflates the balloon with water and deflates it, periodically. The linear actuator is controlled via a custom-made controller which is further interfaced to a computer by a National Instruments board. [29] The user interface for manipulating the actuator, and thus the pump function, has previously been developed in Labview. The motion parameters of the actuator are fully controlled by the software, therefore offering increased functionality for mimicking different motion patterns. Most importantly, the controller has an output trigger signal, its rising and dropping slopes set by the Labview software. This trigger signal is used, via a 1000:1 voltage divider to be of the range of electrocardiographic signals, as an input for System FiVe s ECG recording. The 4D balloon setup became available late in the project and was used only with the third probe prototype. During the experiments the pump was set in motion and the balloon was inflated with water and deflated periodically, with a period of 1Hz. This periodic motion was set to be repeated for a number of cycles 3 The commercial name for a form of acrylic glass.

52 CHAPTER Experiments with phantoms In order to further discuss resulting images and conclusions from these experiments, it is imperative to report on the settings used for the reconstruction. The reconstruction software is intended for periodically moving objects. For the majority of the 3D phantom experiments, the cardiac cycle was arbitrarily defined and divided into four phases. Data needs to be distributed over at least four phases for software implementation reasons. Also, this gives results closer to the 4D situation. In most of the reconstructions the first hundred or so frames were discarded. The transducer has not reached its nominal rotational frequency at that point and does not cover the complete rotation. Therefore, the first unstable part of transducer rotation has been left out. Results from a 4D data set would be visualized as a sequence of the resulting phases. For the 3D phantoms there was no motion of the phantoms and thus no difference in time to be viewed. Therefore, only one of the four resulting data sets was visualized after the reconstruction procedure. In some trials, beam order remained as defined by the input theta order file. It was later proven that it should be switched. The effect is explained later in this section. Finally, unless otherwise mentioned, the reconstructions were done using frames from both directions of rotation. An exception is those reconstructions that used frames acquired during rotation in one direction (up-going) to underline specific aspects of results. These results would potentially, be compared to the results from frames acquired during the opposite direction of rotation (down-going) Issue of additional delay During the preliminary reconstructions from image data acquired with the second of the described prototypes, it was concluded that there was some sort of delay between the angle value acquisition and the image data acquisition. The delay was roughly estimated by the resulting images (Fig. 27a, 27b) to be within the range of 5-10 milliseconds. This delay was different from the delay captured by the fixed delay circuit described in section An additional delay value was implemented in the reconstruction code. We later modified the code in order for the additional delay to be accessible as a program parameter rather than a value in the code, cumbersome to adjust. It was unclear where this additional delay was stemming from. The value of 7 milliseconds was found to present satisfactory results. A number of experiments followed to understand the additional delay and find out if a more suitable value existed. System FiVe s internal parameters were checked for any indication of delays related to processing of image data that would explain the observation. To assess the best value for an additional delay, reconstructions of the block phantom were performed using image data from the anti-clockwise rotation of the transducer array and assigning different values for that delay (Fig. 27c, 27d). Using the same values for the additional delay, reconstructions were also performed using image data from the

53 41 IN-VITRO IMAGING anti-clockwise rotation of the transducer array. The C-planes 4 of the resulting 3D data would rotate in each case according to the increase or decrease of the additional delay. C- planes of reconstructions with the same value of additional delay were compared to find the ones correlating best with each other. The experiment was repeated for specific values of additional delay. a) b) c) d) Figure 27: Images (a) and (b) are C-planes of two reconstructed 3D volumes, I and II, of the rectangular block phantom. Both volumes were reconstructed from the same acquired dataset, using the same value of additional delay. Volume I was reconstructed using image data acquired only during the Down-going rotation of the transducer array. Volume II was reconstructed using the other part of the image data from the Up-going rotation. Theoretically, the block should be positioned at the same orientation in both images. The observed mismatch indicates a mismatch of angle values with corresponding image data. Images (c) and (d) are reconstructed from the same acquired data set, using image data from the anti-clockwise direction. No additional delay is used for the reconstruction of image (c). Image (d) was reconstructed with the implemented 7ms additional delay. Beam order has not been switched for any of the reconstructions. The result was not very straightforward as the images were not very sharp, potentially due to artifacts from using a wrong beam order for the reconstructions. It showed though that best correlation was to be achieved around the 7 milliseconds of additional delay. Experiments stimulated discussion on the beam order influence and its switching in the reconstruction software. In the meantime, the third prototype was being developed according to the specifications we had provided for acquisition of discrete angle values per transmitted beam. We decided to continue reconstruction experiments with the 7ms value and wait for the new probe Beam order switching One of the user-adjustable parameters of the reconstruction software is the switching of the beam order. If switching is activated, then the order of lines per 2D image sector, as defined by the theta order input file, is switched. This means that if the first line is defined as being the one on the left-hand side of the sector, then it switches position with the last line of the sector, on the right-hand side. Subsequent lines are switched in the same way. This may be needed since the software arbitrarily positions the first image line on the left side of the 2D sector, while the transducer may be firing its first beam on the right side. 4 2D slice of the 3D volume dataset at a specific depth level

54 CHAPTER 3 42 a) b) c) d) e) f) g) Figure 28: Physical explanation of the effect of switching the beam order. Image (a) shows a regular 2D frame acquired by a stationary transducer array. The star indicates the first line acquired. If the beam order is switched (b) the effect is the same as a mirroring effect. When looking at the C-planes of a hypothetical 3D volume reconstructed from multiple stationary 2D frames (c, d), the specific frame will seem as if it has been rotated around the axis of rotation. In the case of fast-rotation of the transducer array, acquired 2D frames are curved in space (e), as explained is section 1.3. In such case the switching is not equivalent to of rotation, showing a mirroring effect as the one hypothesized in (f), where the correct curvature of the frame is maintained. It rather changes the curvature of the frame to a reversed one (g), positioning image points at wrong positions in the 3D space. In the case of flat 2D image planes used for the reconstruction of 3D volumes, beam order switching is not critical for the procedure as shown in figure 28a-d. The result of a reversed beam order would be a mirroring effect of the volume. The situation becomes more complex for curved 2D image planes as the ones acquired by the FRTEE. As shown in figure 28e-g, if scanning of the 2D sector starts on the opposite side than the one that the reconstruction software supposes, image data will be distributed in quite a different way. The resulting reconstructed image will show severe artifacts. a) b) Figure 29: Effect of beam order switching on the dolphin phantom. Image (a) is reconstructed without switching the beam order. For the second reconstruction bream order is switched. Both reconstructions used the same acquired dataset as input. To make the effect comprehensible, we used input data from the Up-going rotation only. This may induce additional small-scale artifacts, such as a scraggy surface, due to data sparseness. Such an effect is visible in (b), along the right side of the dolphin, close to the tail.

55 43 IN-VITRO IMAGING The artifact can also be observed in the dolphin phantom reconstructed images above (Fig. 29). Since the effect was not considered during initial experiments, angular artifacts related to it could not be explained and other potential explanations were sought. Following experimentation with elevation angle flipping, beam order and rotational angle switching, it was deduced that the beam order switching parameter should be active for reconstructing the FRTEE-acquired 2D image data. This means that the software should switch the beam order of acquired frames as this is defined in the theta order file to correspond to the real order Effect of the number of frames used for reconstruction With the introduction of the 7 ms additional delay, the quality of 3D reconstructed images was sufficient to assess other parameters of the reconstruction process. One 3D data set of angles (angle file) and images (clip file) acquired from the coffee cup was reconstructed three times, each time using as image data input a different number of acquired frames. The outcome presented in figure 30 serves as an example for the effect of sparse data in 3D reconstruction. The less frames used for reconstruction, the less samples available for reconstruction and the more degraded the reconstructed image is. Reconstructions were performed using 328, 100 and 50 input frames respectively for the three resulting images. The reader should be reminded that these are actually divided among four phases. The actual number of frames from which image samples contribute to the interpolation for reconstructing the 3D images presented are 82, 25 and 12.5 (the decimal point makes sense since samples from discrete frame-lines are actually used). The hole which can be clearly distinguished in the first image and is largely distorted in the last one was intentionally carved on the phantom. It served as an orientation point for some experiments where the phantom was imaged from different sides and orientations. a) b) c) Figure 30: 3D reconstructions of the inverted coffee cup phantom. Reconstruction is done using (a) 82 frames, (b) 25 frames and (c) 12.5 frames. The example shows that there is a minimum number of frames that should be used for accurately reconstructing 3D images of objects. Further, should 16 cardiac phases be

56 CHAPTER 3 44 used for the reconstruction of a sequence of phases of moving structures, an average frame rate of 100 frames per second will produce around 60 frames of acquired image data to be used per phase. However, there are more aspects affecting the quality of the 3D reconstructions. These include the accuracy of the angle measurements both in terms of rotational angle values and their correspondence to image data and the desired outcome of a potential sequence of 3D reconstructions, i.e. increased temporal and/or spatial resolution Shifting of the elevation angles An interesting observation on the reconstructed 3D data sets led to solving a reconstruction problem. As can be seen in figure 31 while the reconstruction is pretty accurate, there is a small bump running from the opening of the balloon longitudinally to the apex of the balloon. The effect can also be observed in reconstructions with the dolphin phantom as well as the block phantom described later. During experimentation with the second probe prototype its contribution as an artifact was minor compared to severe angular inaccuracies and remained unnoticed. Since the latest probe design offered accurate correlation of the angular data with the image data, the bump became a prominent artifact on the reconstructed images. a) b) Figure 31: Artifact of shifted elevation angles. a) The arrows indicate the visual effect of the artifact on the 3D image. b) The same data set was reconstructed, accounting for the artifact in the reconstruction software. Viewing the resulting C-planes figure 32 from the reconstructed data, the right half of the balloon seems to have a larger radius. Closer inspection indicated that the actual thickness of the wall was similar in both halves. The wall on the right half, though, was depicted further from the center of the 3D volume (which is the axis of rotation).

57 45 IN-VITRO IMAGING a) b) c) Figure 32: C-planes of balloon 3D reconstructions. Images (a) and (b) delineated to emphasize the area of interest indicate the problem on 2D slices in the depth direction of the 3D volume. Image (c) shows the slice from the same depth level after the software has been corrected. The effect could be explained if post-processing of the 2D acquired slices on System FiVe shifted the physical elevation angle for each transmitted beam. We speculated that such shifting would be rational if the system requires to define a central transmission beam (e.g. for M-mode) of the entire scan sequence for one 2D scan plane and the number of transmitted beams is even. If such a hypothesis was valid, the data acquired by one half of the ultrasound plane would be depicted a bit further from the center of the plane while the other half would be depicted a bit closer to the center. This would create the effect of tilting of the images in the scan direction, relative to the center of the 2D sector. A total number of 124 received beams (image lines) per 2D image for the 3D balloon phantom imaging was used. This indicates a total (indeed even) number of 62 transmitted beams due to the parallel beam-forming used during acquisition. All acquired image data sets used an even number of transmit beams to scan the 2D plane Therefore, we assumed that the elevation angle would need be corrected in the reconstruction software The 3D images next to the real photos used to introduce the phantoms in sections 3.1 and 3.2 have also been corrected. The results of this correction on the balloon phantom data set as well as on those of the dolphin and block phantoms show that the artifact indeed disappeared. An exception is the coffee-cup phantom where the artifact was not corrected. Further experiments and understanding of the physical value of elevation angles as this is assigned by the imaging system should explain the effect in more detail. The value used for correcting the elevation is equal to half the spacing between two consecutive transmit beams. This basically brings the beam on the left of the central axis, exactly on the axis itself. This value of half the spacing between two consecutive transmit beams is the value that the reconstruction software uses as the absolute spacing between the received beams. The balloon representation was very accurate in the 3D case. The satisfactory 3D reconstructions of the balloon were later used to compare 3D with 4D acquisitions.

58 CHAPTER Experiments with the 4D setups 4D Dolphin phantom setup The dolphin phantom was expected to give some initial result to be compared with imaging of the same phantom in 3D and discuss the effects that motion of the object had on its representation. Experience with 2D and 3D images from previous experiments would assist in comparing the reconstructions. Another advantage was that orientation could be determined in all three dimensions. The fact that rotational displacement of the phantom was 360 degrees per cycle deemed the experiment quite challenging for the reconstruction software, since the acquired image data would be a lot sparser due to the large amount of motion in one cycle. Nonetheless, we had the opportunity to familiarize with the reconstruction of 4D datasets and evaluate functionality of the software and acquisition procedure with such extreme experimental parameters. Figure 33: Time-stamped 2D slices from reconstructed 3D volumes of the rotating dolphin phantom. While 3D volumes showed certain artifacts, the rotating motion of the dolphin can be depicted. The experiment offered the opportunity to familiarize with 4D reconstructions. Video available

59 47 IN-VITRO IMAGING As can be seen in figure 33 from the sequence of 2D slices from the reconstructed 3D data sets for different phases of the rotation, the phantom is indeed shown to be rotating. Nonetheless, the 3D volumes showed large motion artifacts in the rotational direction, in the outer regions. This is due to the large displacement of the phantom in one period and its high rotational frequency. Further, the parameters used for the reconstruction were still under scrutiny to assess how they affect the resulting images and sequence. The large motion artifacts did not allow for further discussion on the results. However, the visual representation of the rotational motion itself as this was visualized in slices from the 3D volumes was promising. 4D balloon phantom setup All experiments with the balloon phantom were conducted using the third design of the FRTEE probe. While imaging the balloon in 4D to discuss initial reconstructions with the new probe, an artifact was observed. Seemingly, the right half of the C-plane slices was out of phase compared to the left half. The artifact should be further investigated as it may be related with the reconstruction software. Figure 34: Time-stamped 3D volumes of the 4D balloon setup. A video sequence is created, including from these 16 volumes. For acquiring image data for this sequence, the transducer was facing the balloon from what is depicted here as its top, towards its bottom. Initial experiments showed an accurate representation of the balloon and its periodic motion. In figure 34, we present 8 volumes captured from the reconstructed 4D se-

60 CHAPTER 3 48 quence of the motion of the balloon. The different phases of the balloon motion are distinguished. We noticed a star-shaped representation of the middle parts of the balloon which was also present, to some smaller extent, at its walls (Fig. 35). While it looked much like an angular artifact, the available satisfactory reconstructions in 3D led to the conclusion that this is a motion artifact. Since the image data sampling in the center of the image is dense, motion can create large artifacts in time. Less motion artifacts are expected if we increase resolution in time by increasing the number of phases from which reconstructed data is generated. Further, temporal smoothing 5 close to the border of the 3D volume would help since image data at these points is sparse. a) b) Figure 35: C-planes of 4D balloon volumes. Especially the right half of the images shows a starshaped rep-representation of the balloon wall. The effect is severe in the center of the volume, at a depth level closer to the transducer array (a). It is still present at further depths (b). Nevertheless, the more phases used for the reconstruction, the sparser the data for every phase. In such case large gaps can be observed at the 3D volumes of individual phases. In such extreme situations of highly sparse data, temporal smoothing will not assist further. 3.4 Conclusions on the imaging of phantoms At this point in the project, time limitations restricted further reconstruction attempts with the data sets from in-vitro experiments to implement proposed improvements and to obtain potentially improved results. 3D images of the corrected 3D reconstructions of phantoms using the data acquired with the latest prototype prove that 3D reconstruction of still objects has been successful. The resulting reconstructed 3D images of motionless phantoms accurately represent the imaged objects. The current 3D images do not present any obvious rotational artifacts and are geometrically correct. However, the imaged 3D objects are bulky in relation to the fine- 5 as it was defined in section 2.4

61 49 IN-VITRO IMAGING detailed heart valves which are imaged in the clinical application. Imaging of such phantoms, by nature does not provide very useful information on potential angular artifacts at the very center of the 3D volume. Visualization of the motion of the 4D balloon phantom was satisfactory. The inflation, deflation and wobbling of the balloon are clearly presented on the generated sequences. The reconstructed volumes showed minor motion artifacts. The issue should further be investigated. The motion of a structure at the center of the volume is imaged multiple times compared to a structure at the border of the volume. In that sense, at the border of the volume, temporal resolution should be lower and higher temporal smoothing should be implemented. Figure 36: 3D image from a 4D sequence of the 4D balloon setup. Video available Quantitative assessment of 3D and 4D reconstruction accuracy of phantoms should be explored in order to provide results which can be compared with other approaches in 3D echocardiography.

62

63 Chapter 4 In-vivo imaging The latest of the fast-rotating TEE probe prototypes undoubtedly produces exceptional quality 3D reconstructed images of motionless objects. What is more, the results obtained with the moving phantoms showed potential for imaging living structures. The motion pattern of the dolphin phantom could be distinguished in 2D slices of the reconstructed data set, as discussed before. Nevertheless, the results from that phantom were not satisfactory in terms of 3D image quality. The accuracy of the measurements and the reconstruction parameters were still issues which, in principle, affected the quality of reconstructed image. However, the large displacement per cycle of the dolphin phantom was considered to contribute largely to the deterioration of the 3D image quality. The range of motion of the heart s structures on the other hand, is significantly smaller than the motion encountered in the of rotation of the dolphin phantom. Imaging of the 4D balloon phantom, which shows a motion pattern closer to that of the heart, had produced satisfactory 3D sequences. In order to prove the practical, clinical capability of the probe and the reconstruction procedure of 3D sequences of the beating heart, we should image the heart in-vivo, during its physiological function. The TEE procedure causes significant discomfort to the patient, therefore we did not want to experiment on patients without testing the image quality in a comparable invivo setting. Within this context we first planned an animal study with anesthetized pigs and conducted experiments with assistance from the Catheterization Laboratory (CatLab) team as well as experienced clinicians. The results and experience from the initial pig experiments provided important remarks to use as feedback for the new design of the third prototype and highlighted some issues to be expected during the clinical application. Satisfactory images of the heart structures of the pig were reconstructed using the newest probe prototype and the optimized version of the reconstruction software. All these results and points of interest are presented and discussed in the first section of this chapter. After obtaining satisfactory results from the pig experiments a small clinical study with humans was planned. The second section of the chapter discusses briefly the clinical procedure and some preliminary results, as the study was carried out while this thesis was 51

64 CHAPTER 4 52 being written. The first images of the human heart, reconstructed from 3D data acquired with the fast-rotating TEE probe, are presented and briefly discussed. 4.1 In-vivo animal experiments We imaged the heart of pigs which were to be sacrificed for another clinical study in the Experimental Cardiology department of Erasmus MC. The pigs used for the animal experiments were anesthetized, as mentioned earlier, and put on respiratory support. Since it is quite difficult to perform a transesophageal acquisition on the pig, the Catlab team proposed to remove one rib from the right side of the pig s thorax and create a hole which would allow the TEE probe to be maneuvered to the pericardium (fig. 37). Following the first experiments, it was decided to open the chest of the pig and image directly on the pericardium after having acquired data sets from the side-hole. Both of these imaging windows are not comparable with transesophageal imaging in terms of the technique used. For example, the clinician has direct visual contact with the heart when imaging from an open-chest and can swiftly identify the desired imaging planes. Nonetheless, imaging through the hole on the side of the thorax of the pig can be a creative way of simulating the transesophageal procedure while the open chest imaging can potentially offer improved image quality of the acquired data. a) b) Figure 37: The two sites from which the heart of the pig was imaged. The hole on the side of the chest of the pig (a) provides an alternative to simulate the procedure of TEE imaging and manipulation$n of the probe. Notice the saline-filled cavity (b) of the thorax. Saline solution allowed imaging the heart without touching the pericardium and inducing potential motion artifacts during acquisition. After experimenting on the influence of the different pre-processing imaging settings on the image quality of phantoms we decided to use default imaging settings for the compression, dynamic range and rejection parameters. One of the issues arising during the first experiments was the fact that the examiner could not view the 2D perspective of the imaged structures without simultaneously triggering fast rotation of the transducer. This was due to the coupling of the beginning of rotation to the beginning of image acquisition. The issue was provisionally alleviated by

65 53 IN-VIVO IMAGING turning off the power of the control box for as long as the examiner was looking for the different imaging planes to acquire. Then, image acquisition was stopped momentarily, the control box was powered on and acquisition of the image and angle data started. However, the second prototype did not offer any manual rotation of the transducer array. Therefore it was still difficult for the examiner to establish a good imaging site. Both of these issues are unacceptable for the clinical application with a patient. We accounted for them and they were addressed by the design of the third prototype, as described in chapter 2. The images of the first open-chest experiments showed promising quality but considerable motion artifacts. We discussed that part of the motion artifacts can be originating from the translation of the heart s motion to the probe, when the probe was touching the pericardium. Therefore, we decided to carry out subsequent experiments were with the open chest cavity filled with saline solution and the probe imaging at a small distance from the pericardium rather than directly on it. In a similar perspective, control of respiration seemed to facilitate the procedure. The pig s breathing was stopped during the actual acquisition period of 10 seconds to avoid further artifacts from body and chest motion. Due to these changes, the experimental protocol became more similar to the human rapid 3D TEE procedure. Figure 38: Aortic valve and outflow tract from the heart of the pig. The valve leaflets (arrow) are barely distinguishable from the surrounding blood. We decided to use lower gain settings in further experiments. Video availabl e The 2D slices of the reconstructed 3D volume in figure 38 show that there is very low contrast between tissue and blood in the cavities imaged. This has a consequence of a solid 3D volume where structures such as valve leaflets and vessel walls are not distinguishable from the cavity during their motion. We concluded that the gain settings of the imaging system should be lowered for our next experiments so that blood would appear as black. The lower gain should also minimize the motor interference which was present on the acquired images as described earlier. Later experiments with lower gain and respiration control produced promising results as shown in figure 39.

66 CHAPTER 4 54 a) b) Figure 39: Mitral valve of the anesthetized pig. En face view of the mitral valve during left ventricular systole (a) and diastole (b). Open-chest imaging from the apex of the heart. Experiments on the last animal were conducted using the third prototype of the FRTEE probe. Reconstructions of the acquired data sets were performed after obtaining satisfactory results from the 3D and 4D phantoms with the same probe. The results were satisfactory although the tissue/blood contrast was still not optimal. Small rotational artifacts remained visible especially in the central axis of rotation. This implied that the software process for the 4D volume reconstructions should further be investigated. We then reconstructed 4D sequences using a higher number of phases per cardiac cycle in an attempt to decrease errors caused by temporal smoothing. Simultaneously, the smoothing parameters were adjusted using different values to study and further optimize their effect on the resulting 3D sequence in relation to the number of phases used. We did not reach any specific conclusions about the result of adjusted values for smoothing combined with higher number of phases. Problems with the visualization software prevented us from viewing quality sequences of more than 20 phases. However, resulting sequences produced from 16 three-dimensional volumes per dataset presented very promising results. The sequence presented in figure 40 clearly illustrates the mitral and aortic valves on a 3D slice. The motion of the mitral valve opening and closing is nicely depicted in the center of the sequence. The aortic valve orifice, located a bit in the background, is shown, one of its leaflets closing upon opening of the mitral valve. The reconstructed data sets were visualized with the Echoview visualization software, Tomtec GmbH.

67 55 IN-VIVO IMAGING a) b) c) d) e) f) Figure 40: Sequence showing the mitral valve en face from the atrium. The aortic valve situated at the background of the 3D image is initially open (a). One of its leaflets is shown as it closes (b, c). The mitral valve is then depicted as its leaflets open inwards (d, e, f). Video available The clinicians who assisted in visualizing the specific 4D volume reported that the resulting sequence offers a quite comprehensible view of the valves. The concept had proven to be successful in imaging the structures of the heart, initially on animals. 4.2 Clinical trial The outcome of the last series of animal experiments was promising enough to embark on a clinical study with humans, using the latest prototype of the FRTEE probe. The goals of the study included acquisition of data sets to reconstruct imaged structures of the human heart, evaluation of the results and assessment of the quality of 2D and 3D images. Further, the medical examiner s assessment of the functional use of the probe was sought for providing further feedback on the clinical use of the device. Two volunteers, J. B. and K. N. offered to undergo the transesophageal echocardiography procedure with the FRTEE probe. The procedure itself is similar to the standard TEE procedure [31] but significantly less time-consuming. An exception is that during the 10 seconds of acquisition the subjects were asked to hold their breath to avoid any motion artifacts due to respiration. In the first volunteer no satisfactory 2D or 3D images were obtained. J. B. experienced continuous spasms of the esophagus and stomach, inhibiting a stable contact between the probe and the esophagus. Imaging in K. N. was successful.

68 CHAPTER 4 56 The clinician reported difficulty in getting acquainted with the manual rotation of the imaging plane since it was rotating very fast upon each button click for clockwise and anticlockwise rotation. As a consequence the entire procedure was slightly prolonged. The examiner s task became also complicated by the lack of the on-screen angular value indication which normally depicts the transducer s angle of rotation. Reconstruction of the acquired data was performed creating 3D sequences of the heart cycle divided into 16 phases. For these reconstructions, imaged structures of interest (aortic and mitral valves) were mainly located at the center of the imaging plane, where image data is oversampled. Low temporal smoothing was first used. It was later discussed that in order to obtain higher temporal resolution the reconstructed data should be divided into more phases (e.g. 64). The technical problems we experienced in visualizing quality 4D volumes of 64 cardiac phases using the available visualization software prevented us from evaluating the theoretical improvement expected. Figure 41 shows the 2D slice from a reconstructed 3D data set. The contrast between valve tissue and blood was still not optimal. The same effect can be observed in figure 42 where the motion of one of the mitral valve leaflets is depicted, as the valve is opening and closing. The clinician noted that image quality can vary from patient to patient depending on positioning of probe inside the esophagus and individual anatomical characteristics. Figure 41: Left ventricle and mitral valve from human heart. The image is a 2D slice from a reconstructed 3D data set. The mitral leaflets are distinguishable. Tissue-to-blood contrast can further be improved. Video availabl e Nevertheless, visualization of the 3D data sets on the Tomtec Image Arena gave satisfactory sequences. The 3D frames presented in figure 43 show the aortic valve of one of the two test subjects.

69 57 IN-VIVO IMAGING a) b) c) Figure 42: Mitral leaflet motion. A mitral leaflet of the human heart is indicated with the red arrow (a). Its motion can be followed through images (b) and (c) as the valve opens and closes. Video available A striking remark is that one of the test subjects was found to have a bicuspid aortic valve 6, as observed in the sequence of figure 43 and shown in figure 44. It is interesting to state that the problem was not known beforehand and was observed during evaluation of the 4D sequences together with clinicians. The congenital anomaly was further investigated and confirmed through a standard congenital transthoracic echo examination. The cardiologist reported that there was no malfunction of the valve at present and suggested annual or biannual check-ups for monitoring the condition of the valve. The unexpected finding signified the first diagnostic use of the fast-rotating 3D TEE approach. a) b) c) d) e) f) Figure 43: Aortic valve of human test subject. The sequence shows the valve as it is open during systole (a) and then closes (b, c, d, e, f). Video availabl e 6 An aortic valve with two leaflets instead of three

70 CHAPTER 4 58 Figure 44: The bicuspid aortic valve. The arrow indicates the horizontal orifice formed by the two leaflets. Video available The probe was tested thereafter as part of a regular clinical procedure on patients who would undergo regular transesophageal echocardiography. After one unsuccessful trial where the patient had an uncontrollable gag reflex, the probe was used with one patient who was screened for thrombi in the left atrial appendage in advance of an intervention to treat atrial fibrillation. Data was acquired during the regular 2D procedure. The sequence presented in figure 45 shows the patient s aortic valve as it closes. The tissue/blood contrast is acceptable. Figure 46 shows the normal aortic valve when it is closed. The image may serve as a reference for comparison with the bicuspid valve presented in figure 44. a) b) c) d) e) f) Figure 45: Normal (tricuspid) aortic valve. Closing of the valve is depicted through images (a) (f). Video available

71 59 IN-VIVO IMAGING Figure 46: Tricuspid aortic valve from human heart. 4.3 Short discussion on in-vivo results The results from the animal study presented satisfactory 4D sequences of the mitral and aortic valves of the heart of the anesthetized pig. The reconstructed sequences of the human heart are promising and of confirmed diagnostic value. They further prove the FRTEE concept as a clinical application with significant potential. The 3D frames presented, still show minor reconstruction artifacts which should be investigated further in relation with both the reconstruction procedure and the chosen values for spatial and temporal smoothing. At the time of the writing of this thesis, minor trials with further optimization of the reconstruction software were conducted. Some points from these trials are presented as concluding remarks in the next chapter.