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Non-invasive imaging: tissue characterisation
CARDIOLOGY
Non-invasive imaging: tissue characterisation
US (reflectivity, attenuation, back-scattering and density as contrast production, and post-processing to enhance visual perception), analyzing the received US echoes; magnetic resonance imaging (MRI e the use of proton Larmor frequency resonance to identify tissue properties), using sequential radiofrequency bursts to produce evidence; Other (infrared oximetry and nuclear imaging e positron emission tomography, single positron emission computed tomography, etc.). We will focus here on traditional US and will not describe other techniques, such as MRI or tissue Doppler imaging (to study the motility of cardiac tissue, Doppler signals are reflected from the myocardium instead of from the red blood cells), which have their own particular method of extracting information in terms of diagnostic images. When used for imaging, US allows a visualisation of soft tissues that cannot easily be obtained using other techniques. Tissue characterisation makes use of US properties: tissues are composed of cells or groups of cells that act as complex boundaries to the propagating wave. The wave propagates throughout these complex structures, and reflected and transmitted waves are generated at each interface, depending on the local density, compressibility and absorption of the tissue. The cells (or groups thereof) can be called ‘scatterers’ since they scatter acoustic energy in their environment. The back-scattered field, travelling back towards the transducer, is used to generate the US image, making use of the coherency of the signals (the US signals being well correlated, as opposed to random signals). A grainy appearance of the US-derived tissue signature it described as ‘speckle’; this is produced by the constructive and destructive interference of the scattered signals from structures smaller than the US wavelength. This interference gives rise to bright and dark echoes (for constructive and destructive interference respectively).1 Speckle is not associated with a particular structure in the tissue, so the phenomenon places a limit on the maximum spatial resolution attainable. In the past, several authors have tried different speckle cancellation techniques in order to increase the image quality of traditional diagnostic US. Apart from being a nuisance, however, speckle offers one important characteristic: despite the fact that it is appropriately described only in statistical terms, due to the variable position of the scatterers, speckle is not a random signal. In fact, the coherency of speckle leads to the preservation of its characteristics when the sourceereceiver distance varies. Hence, motion estimation techniques that can determine, for example, the blood flow or elastic properties of tissues are made possible within the framework of so-called ‘speckle tracking’. In cardiology, tissue characterisation is needed for: angioplasty, to identify re-stenosis (follow-up requiring accurate in vivo diagnostic capabilities to isolate degeneration); infarcted tissue (imaged by using kinetics, contrast echo, perfusion, etc.); plaque characterisation (referring to atlases of specific recognisable kinds of molecules in plaque); thrombus (specific recognisable kinds of molecules in the bloodstream); tissue degeneration (specific recognisable kinds of molecules in overgrowth).
Mauro Grigioni Giuseppe D’Avenio
Abstract In the past 30 years, ultrasound has become a very powerful imaging modality, due mainly to its unique temporal resolution, low cost, nonionising characteristics and portability. Understanding the physical details of the acoustic waveetissue interaction is necessary to realise the full potential of ultrasound techniques, which have steadily improved in terms of quality and range of applications in recent years. In this paper, we give an overview of the fundamentals of diagnostic ultrasound and a brief summary of its applications and methods aimed at tissue characterisation. Besides more conventional techniques such as A-mode, B-mode and M-mode, recent successful advances such as harmonic imaging, three-dimensional visualisation, elasticity imaging and the use of contrast agents are discussed. With these technological improvements, ultrasound techniques will be able to provide high-quality diagnostic tools in everyday clinical practice.
Keywords contrast echocardiography; harmonic imaging; scattering; tissue characterisation; ultrasound
In the past 30 years, ultrasound (US) has become a very powerful imaging modality, mainly due to its favourable temporal resolution, low cost, non-ionising characteristics and portability. Recently, unique features such as harmonic imaging, threedimensional visualisation, transducer fabrication techniques, elasticity imaging and the possibility of using contrast agents have contributed to the improvement in quality and range of applications of diagnostic US images. A short overview of the fundamentals of diagnostic US is presented here, along with a brief summary of methods aimed at tissue characterisation. Tissue characterisation is the study of the properties of human tissue for diagnostic purposes. The main requirements to be able to perform such a study are the integration of methods (the use of US echoes), the skills of the operators and clinicians using a specific technique, and knowledge (years of clinical and technique-related experience to assist the interpretation of results). This integration generally allows clinicians to build a protocol that optimally addresses the study of each specific problem under investigation, exploiting the characteristics of each diagnostic technology. Several techniques are used to investigate tissue by image contrast in order to establish a clinical diagnosis:
Mauro Grigioni is Research Director in the Department of Technology and Health, Istituto Superiore di Sanita` (ISS) Rome, Italy. Giuseppe D’Avenio is Researcher in the Department of Technology and Health, ISS, Rome, Italy.
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Ó 2009 Published by Elsevier Ltd.
CARDIOLOGY
In addition, dynamic viewing can be used to obtain contrast pictures to identify particular structures whose acoustic impedance cannot easily be used for imaging. To understand these concepts we must refer to the basic concepts of the propagation of radiation within the tissues.
difficulty of transferring energy from one point of the medium to its surroundings, since viscosity and friction lead to energy dissipation and heat production. Scattering also causes an attenuation of US, by diffusion of the beam energy. Finally, the interactions at boundaries between media are a complex phenomenon. Generally, the classical optics view of reflection and refraction can be used, as follows: when the ultrasonic beam enters the medium with normal incidence, part of the energy is reflected by the interface, and the remainder continues its travel without being deviated, just as a ray of light does at an interface. The reflected energy (the echo) depends on the characteristic impedances of the two media (the more different the latter two quantities are, the higher will be the energy of the returned echo). Characteristic impedance (Z ) is defined as the ratio of the instantaneous acoustic pressure to the velocity of the particles. The difference in characteristic impedance gives rise to important effects: for example, structures that lie behind air-filled organs (e.g. the lung) do not return measurable echoes, since the tissueeair interface does not allow the US signal to reach the target structure. Conversely, a tissue mass (e.g. a tumour) with a characteristic impedance similar to that of surrounding tissue will cause a very small amplitude signal at the receiver. If the angle of incidence of the US beam is not at 90 degrees with the interface, the ultrasonic beam that enters into and is transmitted by the second medium will be deviated from its original path (refraction) and travel at a certain angle (given by Snell’s law) with respect to the perpendicular to the interface. The difference in angles will depend on the ratio of the velocities of US in the two media. Ultrasonic beams are not always narrow, and reflecting surfaces are not always smooth as previously hypothesised. Reflection from a rough surface is called diffuse, and this results in considerable scattering of the reflected beam. All the properties of travelling US could thus have an effect on the interpretation of images taken with equipment of clinical interest.
Basic physical concepts of the propagation of ultrasound waves US is a form of mechanical energy that consists of high-frequency vibrations (the typical frequency range being of the order of MHz). Low-intensity US can pass through living tissues without altering their function. As a non-ionising technique, it is safer for the patient than, for example, X-rays. US is generated by electrically inducing a deformation in a solid, usually by a piezoelectric effect. High-energy US can produce heating and cavitation, thereby altering cell function (e.g. a high local pressure gradient causing the development of gas bubbles). As a beam of ultrasonic energy traverses a medium, the constituent particles of the latter are put in oscillation. The distance between the points of maximum (or minimum) amplitude is known as the wavelength (l). The frequency ( f ) of the US is expressed in hertz (Hz), or number of cycles per second, and the period T (¼ 1/f ) is the time required to complete one cycle. The parameters l and f are related to the propagation velocity (c) by the equation c ¼ f l. The propagation velocity c is dependent on the medium and its temperature.2 Table 1 presents velocity and attenuation coefficient data for various media. The velocity of propagation is determined by the delay between the movement of adjacent particles. This delay depends on the elasticity modulus (E ) and the density (r) of the medium. The velocity of propagation is thus given by c ¼ OE/r. Since the velocity depends on the medium, the wavelength will also depend on the medium, as well as on the frequency. For example, the propagation velocity in water is 1480 m/s and the wavelength for 1 MHz US is 1.48 mm. The velocity for other media can be found in Table 1. Another relevant characteristic is the attenuation of US intensity, which decreases with distance of travel of the US beam through a uniform medium. This loss of amplitude can result from three processes: beam divergence, absorption and scattering. In the first of these effects, the beam energy spreads out over a larger area as the beam progresses through the medium and the energy per unit area decreases. Absorption gives a measure of the
Modes of image formation Pulse-echo display modes There are two main types of pulse-echo display, the A-scan and the B-scan (A and B standing for amplitude and brightness respectively) methods. Originally, the A-scan method plotted amplitude against time, and the B-scan method plotted range against azimuth (i.e. the angle of the measurement region with respect to the transducer axis).
Velocity and attenuation coefficient data Tissue
Velocity (m/s)
Air Lung Soft tissues (except muscle) Bone Water
330 650e1160 1460e1615
Attenuation coefficient at 1 MHz (dB/cm) 10 40 0.3e1.5
2700e4100 1480
3e10 0.002
A-scan method In ultrasonics, the A-scan method plots the amplitude of echo against distance (which is proportional to time, since the velocity of US in the medium is known). Since the image is a grey-scale picture, the amplitude of the signal is displayed, after extraction of the envelope of the radiofrequency signal. The resulting signal is called a detected A-scan, A-line or A-mode scan. Another fashion of displaying the A-scan is as a function of time, especially in cases where tissue motion needs to be monitored (e.g. with the ventricular wall, to assess contractility). In this case, only one A-scan from a particular tissue structure is
Table 1
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CARDIOLOGY
displayed and is periodically recorded at high frequency, generating a depthetime display. This image modality is called M-mode (for motion) scanning. A typical application of the M-mode display is in examining the motion of the heart valve leaflets, or in assessing vascular distensibility.
frequencies provide very small wavelengths and excellent theoretical resolution, as US wavelength, which determines the maximum resolution, is inversely proportional to frequency. At 30 MHz, the wavelength is approximately 50 mm, which allows an axial resolution of 100 mm. Lateral resolution is about 250 mm. IVUS catheters have a maximum diameter of 2.6e3 Fr (0.89e1 mm). A motorised pull-back of the transducer (0.25e 1 mm/s, usually 0.5 mm/s) enables two-dimensinal tomographic scans of large areas of vessels to be made (Figure 1). By increasing imaging frequencies, we can obtain an improved qualitative assessment of, for example, specific tissue such as atherosclerotic plaques. To provide a practical classification of obtainable information, we refer to the main properties of the imaged atherosclerotic plaques to identify the tissues as follows:
B-scan method The received A-scans can be spatially combined after acquisition (using, for example, a transducer put in motion while acquiring echoes, or suitably combined transducer arrays), enabling a two-dimensional scan of the considered region, called brightness or B-mode, to be obtained; this is the most widely used diagnostic US mode. One of the biggest advantages of US scanning is real-time scanning (up to 100 Hz frame rate), which is achieved due to the limited depth of scanning in most tissues and the usually high speed of sound. Resolution With pulse-echo techniques, there is a limit to distinguishing between two closely spaced structures. This spatial limit is about one wavelength in the particular medium. Since the velocity of US in soft tissue is typically 1500 m/s, the resolution in millimetres can be expressed as a function of f, bearing in mind that l ¼ c/f (so that, at 1 MHz, the resolution will be about 1.5 mm). Intravascular ultrasound In cardiology, intravascular US (IVUS) systems are the in vivo US application with the highest resolution. The applicability of IVUSguided diagnosis is based on the following properties: real-time high-resolution imaging and the two-dimensional tomographic assessments of vessels. In addition, longitudinal and threedimensional computer-assisted reconstruction can be obtained by building three-dimensional models of vessel structures in the case of, for example, coronaries or large vessels. IVUS also allows an assessment of the total vessel lumen and plaque dimension in vivo. The transducers have US frequencies ranging between 20 and 50 MHz (usually 30 MHz). Thus high
soft, low echogenicity; fibrous, high echogenicity; calcified, high echogenicity with acoustic shadowing/ reverberations. Calcific plaque is the simplest tissue type to identify as it yields a bright signal reflection. Calcification is seen in 60e80% of target lesions using IVUS, compared with 30e40% by angiography. It must be emphasised, however, that 180 degrees of vascular circumference must be calcified before it can be visualised by angiography. Fibrous plaques have less echogenicity than calcium, but more than muscle or fat tissue. In general, the brightness of fibrous tissue is similar to that of adventitia. No reverberations are present in this case. Fatty plaque is radiolucent and has a soft grey-scale appearance on IVUS. Shadowing from a heavily fibrotic plaque can be mistaken for lipid. Specific endovascular entities such as thrombus can also be detected by imaging, even though this is one of the most difficult tissue types to identify by IVUS. It produces a sparkling pattern on real-time IVUS imaging. Identification of thrombus after stenting may sometimes be vital.
Intravascular ultrasound showing a section of a vessel (left-hand panel) and a longitudinal scan via pull-back of the transducer (right-hand panel). Figure 1
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Finally, an application relating to the in vivo evaluation of medical devices can be seen by using IVUS to confirm the patency of a stented vessel.
back-scatter from blood, allowing tissue imaging without contrast agents in such cases. Imaging using the harmonic approach (with or without contrast agents) is generally known as harmonic imaging. Due to the much smaller influence of the transmitted pulse on the received spectrum, harmonic imaging offers high image clarity and crisper edge definition.
Post-processing Intravascular IVUS image of a coronary artery allows a delineation of the lumen by the following sequential steps (see Figure 1 above): the green curve is the coarse-grained input to the snake algorithm,3 and the red solid curve is the border after several iterations of the algorithm. Active contours (such as snakes) are widely used to locate object boundaries. In this instance, a new external force (gradient vector flow) for active contours has been used, solving issues about initialisation and poor convergence of boundary concavities.
Back-scattering To understand what the signal is in the case of back-scatter, we must remember that tissue is made up of cells and groups of cells, which act as interfaces to propagating US waves. Depending on the local density, etc., groups of cells back-scatter US. The scatter field is called speckle (grainy) and is used to generate images by means of constructive and destructive interference (signatures of the tissue). Integration gives a measure of the changing properties of pathological or damaged tissue. At the Bambino Gesu` Children’s Hospital, it was recently found that, based on the cyclic variation of integrated back-scatter, it was possible to distinguish between the myocardial physical properties of patients undergoing the Kawashima procedure and those of controls, despite normal left ventricular function, and independently of demonstrable coronary abnormalities.4 The anomalies may have been related to cellular damage induced by occult myocardial ischaemia or may have represented abnormalities of small myocardial vessels that could be observed as a result of this technique.
Contrast echo Despite the fact that diagnostic US is an older imaging modality than MRI or positron emission tomography, it is nevertheless continuing to expand as a field, diversifying in its applications. Much progress has been brought about in recent years by faster computer processors, contrast agents, the utilisation of nonlinear wave propagation, signal-processing techniques and complex transducer architecture. During the interaction of the wave with the tissues, non-linear effects also occur besides the linear wave propagation discussed earlier, especially at higher US intensities. As a result, non-linear waves can be generated; these depend on acoustic pressure, the characteristics of the medium and the depth of the beam path inside the body. One of the main problems with the standard use of US arises from high attenuation in some tissues, most prominent in the imaging of small vessels and in velocity measurements. In order to overcome this limitation, contrast agents are routinely employed for blood f1ow measurements. Contrast agents are typically microspheres of encapsulated gas or liquid coated by a shell (e.g. albumin). This gas or liquid usually creates a high impedance mismatch, so the resulting back-scatter by the contrast agents is much higher than that of the blood particles. An alternative method of generating higher back-scatter due to the increased impedance mismatch is based on the harmonics generated by the bubble’s interaction with the ultrasonic wave. This interaction results in a vibration of the latter at a particular resonance frequency, as well as at multiples thereof. The second harmonic is particularly important as it can show information usually shadowed by imaging at the fundamental frequency. Indeed, the filtering of undesired echoes from stationary media surrounding the region of interest results in a weakening of the overall signal at the fundamental frequency. Therefore, since high-order harmonics are caused by moving scatterers, motion characteristics can be obtained from the higher harmonic part of the echoes, after high-pass filtering and cancellation of the fundamental frequency spectrum. Besides their application in blood motion measurements, contrast agents can be used for tissue imaging as well: after injection into the bloodstream, the contrast agents can enter and remain in the tissues, offering high-value diagnostic information. However, contrast agents are not always needed for imaging tissues at higher harmonics, especially since back-scatter from the tissues can be up to two orders of magnitude higher than
PAEDIATRICS AND CHILD HEALTH 19:S2
Elastography Another rapidly evolving field is elasticity imaging. Owing to the significant differences between the mechanical properties of several tissue components, the information contained in the coherent scattering (speckle) of US following a mechanical stimulus is sufficient to characterise these tissues. As an example, the range of elastic modulus (relating deformation to imposed stress) for several different normal and pathological human breast tissues in principle allows the detection of tumoural tissues (which are usually stiffer than normal tissues), which are traditionally assessed, in qualitative terms, by palpation.5 Coherent echoes can be tracked while (or after) the tissue in question undergoes motion and/or deformation caused by the mechanical stimulus. This provides the basis for measuring the elasticity of the tissues, as elasticity images (or elastograms). The external application of deformation to the tissue can be provided as a quasi-static compression, a monochromatic lowfrequency vibration (e.g. in sonoelasticity imaging) or a transient shear wave (transient elastography, which allows a determination of the viscoelastic properties of the tissues). Elasticity imaging techniques are achieving increasing success as a tissue imaging modality. Even larger diffusion is warranted by the fact that freehand scanning will be a viable option in the near future.6
Conclusion The wide application of such different US techniques highlights the relevance of a technology that, despite its rather poor resolution compared with other techniques, meets the needs of patients with regard to routine examinations in the general healthcare system, before proceeding, if necessary, to more indepth investigations. In several clinical fields, these US
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2 Feigenbaum H. Echocardiography. 5th edn. St Louis: Mosby-Yearbook; 1994. 3 Xu C, Prince JL. Snakes, shapes, and gradient vector flow. IEEE Trans Image Process 1998; 7: 359e69. 4 Leonardi B, Giglio V, Pasceri V, de Zorzi A, Sanders SP. Myocardial ultrasound tissue characterisation in children after Kawasaki disease. Circulation 2007; 116: II_512. 5 Moradi M, Mousavi P, Abolmaesumi P. Computer-aided diagnosis of prostate cancer with emphasis on ultrasound-based approaches: a review. Ultrasound Med Biol 2007; 33: 1010e28. 6 Havre RF, Elde E, Gilja OH, et al. Freehand real-time elastography: impact of scanning parameters on image quality and in vitro intra- and interobserver validations. Ultrasound Med Biol 2008; 34: 1638e50.
techniques, assisted by new strategies and clinical knowledge, will in future allow a more specific diagnostic use of these repeatable and easy-to-use scanning methods.
Conflict of interest None of the authors of this paper has a financial or personal relationship with other people or organisations that could inappropriately influence or bias the content of the paper. A
REFERENCES 1 Burckhardt C. Speckle in US b-mode scans. IEEE Trans Sonics Ultrason 1978; 25: 1e6.
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Non-invasive imaging: tissue characterisation
US (reflectivity, attenuation, back-scattering and density as contrast production, and post-processing to enhance visual perception), analyzing the received US echoes; magnetic resonance imaging (MRI e the use of proton Larmor frequency resonance to identify tissue properties), using sequential radiofrequency bursts to produce evidence; Other (infrared oximetry and nuclear imaging e positron emission tomography, single positron emission computed tomography, etc.). We will focus here on traditional US and will not describe other techniques, such as MRI or tissue Doppler imaging (to study the motility of cardiac tissue, Doppler signals are reflected from the myocardium instead of from the red blood cells), which have their own particular method of extracting information in terms of diagnostic images. When used for imaging, US allows a visualisation of soft tissues that cannot easily be obtained using other techniques. Tissue characterisation makes use of US properties: tissues are composed of cells or groups of cells that act as complex boundaries to the propagating wave. The wave propagates throughout these complex structures, and reflected and transmitted waves are generated at each interface, depending on the local density, compressibility and absorption of the tissue. The cells (or groups thereof) can be called ‘scatterers’ since they scatter acoustic energy in their environment. The back-scattered field, travelling back towards the transducer, is used to generate the US image, making use of the coherency of the signals (the US signals being well correlated, as opposed to random signals). A grainy appearance of the US-derived tissue signature it described as ‘speckle’; this is produced by the constructive and destructive interference of the scattered signals from structures smaller than the US wavelength. This interference gives rise to bright and dark echoes (for constructive and destructive interference respectively).1 Speckle is not associated with a particular structure in the tissue, so the phenomenon places a limit on the maximum spatial resolution attainable. In the past, several authors have tried different speckle cancellation techniques in order to increase the image quality of traditional diagnostic US. Apart from being a nuisance, however, speckle offers one important characteristic: despite the fact that it is appropriately described only in statistical terms, due to the variable position of the scatterers, speckle is not a random signal. In fact, the coherency of speckle leads to the preservation of its characteristics when the sourceereceiver distance varies. Hence, motion estimation techniques that can determine, for example, the blood flow or elastic properties of tissues are made possible within the framework of so-called ‘speckle tracking’. In cardiology, tissue characterisation is needed for: angioplasty, to identify re-stenosis (follow-up requiring accurate in vivo diagnostic capabilities to isolate degeneration); infarcted tissue (imaged by using kinetics, contrast echo, perfusion, etc.); plaque characterisation (referring to atlases of specific recognisable kinds of molecules in plaque); thrombus (specific recognisable kinds of molecules in the bloodstream); tissue degeneration (specific recognisable kinds of molecules in overgrowth).
Mauro Grigioni Giuseppe D’Avenio
Abstract In the past 30 years, ultrasound has become a very powerful imaging modality, due mainly to its unique temporal resolution, low cost, nonionising characteristics and portability. Understanding the physical details of the acoustic waveetissue interaction is necessary to realise the full potential of ultrasound techniques, which have steadily improved in terms of quality and range of applications in recent years. In this paper, we give an overview of the fundamentals of diagnostic ultrasound and a brief summary of its applications and methods aimed at tissue characterisation. Besides more conventional techniques such as A-mode, B-mode and M-mode, recent successful advances such as harmonic imaging, three-dimensional visualisation, elasticity imaging and the use of contrast agents are discussed. With these technological improvements, ultrasound techniques will be able to provide high-quality diagnostic tools in everyday clinical practice.
Keywords contrast echocardiography; harmonic imaging; scattering; tissue characterisation; ultrasound
In the past 30 years, ultrasound (US) has become a very powerful imaging modality, mainly due to its favourable temporal resolution, low cost, non-ionising characteristics and portability. Recently, unique features such as harmonic imaging, threedimensional visualisation, transducer fabrication techniques, elasticity imaging and the possibility of using contrast agents have contributed to the improvement in quality and range of applications of diagnostic US images. A short overview of the fundamentals of diagnostic US is presented here, along with a brief summary of methods aimed at tissue characterisation. Tissue characterisation is the study of the properties of human tissue for diagnostic purposes. The main requirements to be able to perform such a study are the integration of methods (the use of US echoes), the skills of the operators and clinicians using a specific technique, and knowledge (years of clinical and technique-related experience to assist the interpretation of results). This integration generally allows clinicians to build a protocol that optimally addresses the study of each specific problem under investigation, exploiting the characteristics of each diagnostic technology. Several techniques are used to investigate tissue by image contrast in order to establish a clinical diagnosis:
Mauro Grigioni is Research Director in the Department of Technology and Health, Istituto Superiore di Sanita` (ISS) Rome, Italy. Giuseppe D’Avenio is Researcher in the Department of Technology and Health, ISS, Rome, Italy.
PAEDIATRICS AND CHILD HEALTH 19:S2
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Ó 2009 Published by Elsevier Ltd.
CARDIOLOGY
In addition, dynamic viewing can be used to obtain contrast pictures to identify particular structures whose acoustic impedance cannot easily be used for imaging. To understand these concepts we must refer to the basic concepts of the propagation of radiation within the tissues.
difficulty of transferring energy from one point of the medium to its surroundings, since viscosity and friction lead to energy dissipation and heat production. Scattering also causes an attenuation of US, by diffusion of the beam energy. Finally, the interactions at boundaries between media are a complex phenomenon. Generally, the classical optics view of reflection and refraction can be used, as follows: when the ultrasonic beam enters the medium with normal incidence, part of the energy is reflected by the interface, and the remainder continues its travel without being deviated, just as a ray of light does at an interface. The reflected energy (the echo) depends on the characteristic impedances of the two media (the more different the latter two quantities are, the higher will be the energy of the returned echo). Characteristic impedance (Z ) is defined as the ratio of the instantaneous acoustic pressure to the velocity of the particles. The difference in characteristic impedance gives rise to important effects: for example, structures that lie behind air-filled organs (e.g. the lung) do not return measurable echoes, since the tissueeair interface does not allow the US signal to reach the target structure. Conversely, a tissue mass (e.g. a tumour) with a characteristic impedance similar to that of surrounding tissue will cause a very small amplitude signal at the receiver. If the angle of incidence of the US beam is not at 90 degrees with the interface, the ultrasonic beam that enters into and is transmitted by the second medium will be deviated from its original path (refraction) and travel at a certain angle (given by Snell’s law) with respect to the perpendicular to the interface. The difference in angles will depend on the ratio of the velocities of US in the two media. Ultrasonic beams are not always narrow, and reflecting surfaces are not always smooth as previously hypothesised. Reflection from a rough surface is called diffuse, and this results in considerable scattering of the reflected beam. All the properties of travelling US could thus have an effect on the interpretation of images taken with equipment of clinical interest.
Basic physical concepts of the propagation of ultrasound waves US is a form of mechanical energy that consists of high-frequency vibrations (the typical frequency range being of the order of MHz). Low-intensity US can pass through living tissues without altering their function. As a non-ionising technique, it is safer for the patient than, for example, X-rays. US is generated by electrically inducing a deformation in a solid, usually by a piezoelectric effect. High-energy US can produce heating and cavitation, thereby altering cell function (e.g. a high local pressure gradient causing the development of gas bubbles). As a beam of ultrasonic energy traverses a medium, the constituent particles of the latter are put in oscillation. The distance between the points of maximum (or minimum) amplitude is known as the wavelength (l). The frequency ( f ) of the US is expressed in hertz (Hz), or number of cycles per second, and the period T (¼ 1/f ) is the time required to complete one cycle. The parameters l and f are related to the propagation velocity (c) by the equation c ¼ f l. The propagation velocity c is dependent on the medium and its temperature.2 Table 1 presents velocity and attenuation coefficient data for various media. The velocity of propagation is determined by the delay between the movement of adjacent particles. This delay depends on the elasticity modulus (E ) and the density (r) of the medium. The velocity of propagation is thus given by c ¼ OE/r. Since the velocity depends on the medium, the wavelength will also depend on the medium, as well as on the frequency. For example, the propagation velocity in water is 1480 m/s and the wavelength for 1 MHz US is 1.48 mm. The velocity for other media can be found in Table 1. Another relevant characteristic is the attenuation of US intensity, which decreases with distance of travel of the US beam through a uniform medium. This loss of amplitude can result from three processes: beam divergence, absorption and scattering. In the first of these effects, the beam energy spreads out over a larger area as the beam progresses through the medium and the energy per unit area decreases. Absorption gives a measure of the
Modes of image formation Pulse-echo display modes There are two main types of pulse-echo display, the A-scan and the B-scan (A and B standing for amplitude and brightness respectively) methods. Originally, the A-scan method plotted amplitude against time, and the B-scan method plotted range against azimuth (i.e. the angle of the measurement region with respect to the transducer axis).
Velocity and attenuation coefficient data Tissue
Velocity (m/s)
Air Lung Soft tissues (except muscle) Bone Water
330 650e1160 1460e1615
Attenuation coefficient at 1 MHz (dB/cm) 10 40 0.3e1.5
2700e4100 1480
3e10 0.002
A-scan method In ultrasonics, the A-scan method plots the amplitude of echo against distance (which is proportional to time, since the velocity of US in the medium is known). Since the image is a grey-scale picture, the amplitude of the signal is displayed, after extraction of the envelope of the radiofrequency signal. The resulting signal is called a detected A-scan, A-line or A-mode scan. Another fashion of displaying the A-scan is as a function of time, especially in cases where tissue motion needs to be monitored (e.g. with the ventricular wall, to assess contractility). In this case, only one A-scan from a particular tissue structure is
Table 1
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displayed and is periodically recorded at high frequency, generating a depthetime display. This image modality is called M-mode (for motion) scanning. A typical application of the M-mode display is in examining the motion of the heart valve leaflets, or in assessing vascular distensibility.
frequencies provide very small wavelengths and excellent theoretical resolution, as US wavelength, which determines the maximum resolution, is inversely proportional to frequency. At 30 MHz, the wavelength is approximately 50 mm, which allows an axial resolution of 100 mm. Lateral resolution is about 250 mm. IVUS catheters have a maximum diameter of 2.6e3 Fr (0.89e1 mm). A motorised pull-back of the transducer (0.25e 1 mm/s, usually 0.5 mm/s) enables two-dimensinal tomographic scans of large areas of vessels to be made (Figure 1). By increasing imaging frequencies, we can obtain an improved qualitative assessment of, for example, specific tissue such as atherosclerotic plaques. To provide a practical classification of obtainable information, we refer to the main properties of the imaged atherosclerotic plaques to identify the tissues as follows:
B-scan method The received A-scans can be spatially combined after acquisition (using, for example, a transducer put in motion while acquiring echoes, or suitably combined transducer arrays), enabling a two-dimensional scan of the considered region, called brightness or B-mode, to be obtained; this is the most widely used diagnostic US mode. One of the biggest advantages of US scanning is real-time scanning (up to 100 Hz frame rate), which is achieved due to the limited depth of scanning in most tissues and the usually high speed of sound. Resolution With pulse-echo techniques, there is a limit to distinguishing between two closely spaced structures. This spatial limit is about one wavelength in the particular medium. Since the velocity of US in soft tissue is typically 1500 m/s, the resolution in millimetres can be expressed as a function of f, bearing in mind that l ¼ c/f (so that, at 1 MHz, the resolution will be about 1.5 mm). Intravascular ultrasound In cardiology, intravascular US (IVUS) systems are the in vivo US application with the highest resolution. The applicability of IVUSguided diagnosis is based on the following properties: real-time high-resolution imaging and the two-dimensional tomographic assessments of vessels. In addition, longitudinal and threedimensional computer-assisted reconstruction can be obtained by building three-dimensional models of vessel structures in the case of, for example, coronaries or large vessels. IVUS also allows an assessment of the total vessel lumen and plaque dimension in vivo. The transducers have US frequencies ranging between 20 and 50 MHz (usually 30 MHz). Thus high
soft, low echogenicity; fibrous, high echogenicity; calcified, high echogenicity with acoustic shadowing/ reverberations. Calcific plaque is the simplest tissue type to identify as it yields a bright signal reflection. Calcification is seen in 60e80% of target lesions using IVUS, compared with 30e40% by angiography. It must be emphasised, however, that 180 degrees of vascular circumference must be calcified before it can be visualised by angiography. Fibrous plaques have less echogenicity than calcium, but more than muscle or fat tissue. In general, the brightness of fibrous tissue is similar to that of adventitia. No reverberations are present in this case. Fatty plaque is radiolucent and has a soft grey-scale appearance on IVUS. Shadowing from a heavily fibrotic plaque can be mistaken for lipid. Specific endovascular entities such as thrombus can also be detected by imaging, even though this is one of the most difficult tissue types to identify by IVUS. It produces a sparkling pattern on real-time IVUS imaging. Identification of thrombus after stenting may sometimes be vital.
Intravascular ultrasound showing a section of a vessel (left-hand panel) and a longitudinal scan via pull-back of the transducer (right-hand panel). Figure 1
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Finally, an application relating to the in vivo evaluation of medical devices can be seen by using IVUS to confirm the patency of a stented vessel.
back-scatter from blood, allowing tissue imaging without contrast agents in such cases. Imaging using the harmonic approach (with or without contrast agents) is generally known as harmonic imaging. Due to the much smaller influence of the transmitted pulse on the received spectrum, harmonic imaging offers high image clarity and crisper edge definition.
Post-processing Intravascular IVUS image of a coronary artery allows a delineation of the lumen by the following sequential steps (see Figure 1 above): the green curve is the coarse-grained input to the snake algorithm,3 and the red solid curve is the border after several iterations of the algorithm. Active contours (such as snakes) are widely used to locate object boundaries. In this instance, a new external force (gradient vector flow) for active contours has been used, solving issues about initialisation and poor convergence of boundary concavities.
Back-scattering To understand what the signal is in the case of back-scatter, we must remember that tissue is made up of cells and groups of cells, which act as interfaces to propagating US waves. Depending on the local density, etc., groups of cells back-scatter US. The scatter field is called speckle (grainy) and is used to generate images by means of constructive and destructive interference (signatures of the tissue). Integration gives a measure of the changing properties of pathological or damaged tissue. At the Bambino Gesu` Children’s Hospital, it was recently found that, based on the cyclic variation of integrated back-scatter, it was possible to distinguish between the myocardial physical properties of patients undergoing the Kawashima procedure and those of controls, despite normal left ventricular function, and independently of demonstrable coronary abnormalities.4 The anomalies may have been related to cellular damage induced by occult myocardial ischaemia or may have represented abnormalities of small myocardial vessels that could be observed as a result of this technique.
Contrast echo Despite the fact that diagnostic US is an older imaging modality than MRI or positron emission tomography, it is nevertheless continuing to expand as a field, diversifying in its applications. Much progress has been brought about in recent years by faster computer processors, contrast agents, the utilisation of nonlinear wave propagation, signal-processing techniques and complex transducer architecture. During the interaction of the wave with the tissues, non-linear effects also occur besides the linear wave propagation discussed earlier, especially at higher US intensities. As a result, non-linear waves can be generated; these depend on acoustic pressure, the characteristics of the medium and the depth of the beam path inside the body. One of the main problems with the standard use of US arises from high attenuation in some tissues, most prominent in the imaging of small vessels and in velocity measurements. In order to overcome this limitation, contrast agents are routinely employed for blood f1ow measurements. Contrast agents are typically microspheres of encapsulated gas or liquid coated by a shell (e.g. albumin). This gas or liquid usually creates a high impedance mismatch, so the resulting back-scatter by the contrast agents is much higher than that of the blood particles. An alternative method of generating higher back-scatter due to the increased impedance mismatch is based on the harmonics generated by the bubble’s interaction with the ultrasonic wave. This interaction results in a vibration of the latter at a particular resonance frequency, as well as at multiples thereof. The second harmonic is particularly important as it can show information usually shadowed by imaging at the fundamental frequency. Indeed, the filtering of undesired echoes from stationary media surrounding the region of interest results in a weakening of the overall signal at the fundamental frequency. Therefore, since high-order harmonics are caused by moving scatterers, motion characteristics can be obtained from the higher harmonic part of the echoes, after high-pass filtering and cancellation of the fundamental frequency spectrum. Besides their application in blood motion measurements, contrast agents can be used for tissue imaging as well: after injection into the bloodstream, the contrast agents can enter and remain in the tissues, offering high-value diagnostic information. However, contrast agents are not always needed for imaging tissues at higher harmonics, especially since back-scatter from the tissues can be up to two orders of magnitude higher than
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Elastography Another rapidly evolving field is elasticity imaging. Owing to the significant differences between the mechanical properties of several tissue components, the information contained in the coherent scattering (speckle) of US following a mechanical stimulus is sufficient to characterise these tissues. As an example, the range of elastic modulus (relating deformation to imposed stress) for several different normal and pathological human breast tissues in principle allows the detection of tumoural tissues (which are usually stiffer than normal tissues), which are traditionally assessed, in qualitative terms, by palpation.5 Coherent echoes can be tracked while (or after) the tissue in question undergoes motion and/or deformation caused by the mechanical stimulus. This provides the basis for measuring the elasticity of the tissues, as elasticity images (or elastograms). The external application of deformation to the tissue can be provided as a quasi-static compression, a monochromatic lowfrequency vibration (e.g. in sonoelasticity imaging) or a transient shear wave (transient elastography, which allows a determination of the viscoelastic properties of the tissues). Elasticity imaging techniques are achieving increasing success as a tissue imaging modality. Even larger diffusion is warranted by the fact that freehand scanning will be a viable option in the near future.6
Conclusion The wide application of such different US techniques highlights the relevance of a technology that, despite its rather poor resolution compared with other techniques, meets the needs of patients with regard to routine examinations in the general healthcare system, before proceeding, if necessary, to more indepth investigations. In several clinical fields, these US
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2 Feigenbaum H. Echocardiography. 5th edn. St Louis: Mosby-Yearbook; 1994. 3 Xu C, Prince JL. Snakes, shapes, and gradient vector flow. IEEE Trans Image Process 1998; 7: 359e69. 4 Leonardi B, Giglio V, Pasceri V, de Zorzi A, Sanders SP. Myocardial ultrasound tissue characterisation in children after Kawasaki disease. Circulation 2007; 116: II_512. 5 Moradi M, Mousavi P, Abolmaesumi P. Computer-aided diagnosis of prostate cancer with emphasis on ultrasound-based approaches: a review. Ultrasound Med Biol 2007; 33: 1010e28. 6 Havre RF, Elde E, Gilja OH, et al. Freehand real-time elastography: impact of scanning parameters on image quality and in vitro intra- and interobserver validations. Ultrasound Med Biol 2008; 34: 1638e50.
techniques, assisted by new strategies and clinical knowledge, will in future allow a more specific diagnostic use of these repeatable and easy-to-use scanning methods.
Conflict of interest None of the authors of this paper has a financial or personal relationship with other people or organisations that could inappropriately influence or bias the content of the paper. A
REFERENCES 1 Burckhardt C. Speckle in US b-mode scans. IEEE Trans Sonics Ultrason 1978; 25: 1e6.
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