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Chủ Nhật, 31 tháng 3, 2013

HIFU in ONCOLOGY


Abstract

High-intensity focused ultrasound (HIFU) is a rapidly maturing technology with diverse clinical applications. In the field of oncology, the use of HIFU to non-invasively cause tissue necrosis in a defined target, a technique known as focused ultrasound surgery (FUS), has considerable potential for tumour ablation. In this article, we outline the development and underlying principles of HIFU, overview the limitations and commercially available equipment for FUS, then summarise some of the recent technological advances and experimental clinical trials that we predict will have a positive impact on extending the role of FUS in cancer therapy.

Focused ultrasound surgery (FUS), using high-intensity focused ultrasound (HIFU) technology in combination with modern imaging methods, has the potential to ablate internal tumour target tissue with great precision, giving it all the benefits of minimally invasive surgery [1]. Damage to adjacent or intervening tissues may be minimised with careful image-based treatment planning and the tumour target may be visualised during treatment. As it does not involve ionising radiation, it is low risk and repeat treatments are possible. The non-invasiveness of FUS reduces toxicity compared with other ablation techniques and adjacent blood vessels may be less vulnerable to damage compared with surgical risks [2,3]. FUS therefore holds great promise as a single or part of a multimodal approach for cancer treatment, especially for patients with cancers unsuitable for other established therapeutic options.

We describe how recent technical developments in HIFU equipment design, electronic control, ablation focusing and target imaging have made rapid advances that are overcoming previous limitations of HIFU for destroying target tumour tissue, especially in shortening FUS treatment times. Together with ongoing worldwide trials exploring oncology applications, this is strengthening confidence in FUS and broadening its scope. As a result, we believe that it is evolving into an increasingly more useful alternative or complementary treatment option and have continued expectation that FUS will be successfully integrated into routine future clinical practice.

PRINCIPLES OF FUS

HIFU transducers are made from piezoelectric materials that oscillate upon application of an alternating voltage, resulting in the generation of ultrasound waves in the receiving medium. They are capable of handling relatively high levels of power and focus the resulting ultrasound beam to a small “cigar”-shaped volume, typically of a few cubic millimetres. Focusing can be achieved geometrically, either by using a curved (spherical section) transducer or by using a plane transducer and a curved lens (Figure 1a). In devices that use an array of small transducers, beam focusing may also be achieved by electronic control (Figure 1b). Modern transducers can create acoustic intensities in a target tissue of ∼100–10 000  W cm−2 and peak compression pressures of up to 30 MPa. In comparison, diagnostic ultrasound transducers deliver intensities of ∼0.0001–0.1000 W cm−2 and a compression of 0.001–0.003 MPa [2].

 

Figure 1.

Diagram illustrating focusing principles of high-intensity focused ultrasound in single (a) and array (b) transducers. Reproduced with permission from Pioneer Bioscience Publishing Company, from Khokhlova and Hwang [16].

Rapid elevation of the local tissue temperature is the main causative mechanism of tissue destruction. Coagulative necrosis occurs as a high amount of acoustic energy is deposited in a short period of time—a function of both tissue temperature and exposure time [37]. This thermal effect was the preferred mode of targeted ablation in early clinical applications of HIFU as it was most predictable and understood [4]. Mechanical tissue effects also occur at very high ultrasound intensities [3,810]. Cavitation, i.e. bubble formation, occurs as microscopic gas bodies are drawn out of solution because of alternating rarefaction and compression and local temperature elevations. A low-pressure acoustic field results in stable cavitation, where microbubbles oscillate. In turn, fluid movement leads to the production of shear forces that cause cell membrane disruption and resulting cell damage—a phenomenon known as microstreaming. With high acoustic pressures, vibration-induced changes in microbubble volume result in inertial cavitation, i.e. violent bubble collapse. If this happens near the cell membrane, destruction of the cell may occur [813]. Radiation forces are also created in tissues owing to the absorption and reflection of the ultrasound wave energy [8]. These cause additional destructive bioeffects, including cell membrane deformation, microstreaming and organelle rotation [8,14]. Mechanical destructive effects have been increasingly exploited as HIFU understanding, experience and technological developments have advanced. Harnessing mechanical bioeffects can result in larger treatment volumes, and hence shorter treatment times, as well as achieving very sharply demarcated precise lesions. This latter effect forms the basis of “histotripsy”—a development of HIFU tissue ablation, which uses short pulses of very high-intensity ultrasound to specifically induce mechanical bioeffects for tissue destruction [3,15].

KEY LIMITATIONS OF HIFU FOR FUS

Since ultrasound is reflected at interfaces between soft tissues and air–gas and is rapidly attenuated in bone, the presence of lung, ribs or gaseous bowel in front of the FUS target region can be problematic. Sonication through the cranium is particularly challenging owing to high attenuation and variable thickness and density of the skull. In addition, non-uniform soft tissues cause the ultrasound beam to propagate variably. Therefore, an appropriate “acoustic window” may be required for an ultrasound beam to propagate through the body to the target volume, restricting the application of FUS to specific patients/tumours. Beam scattering and diffraction may also occur. Unwanted high-energy deposition to tissues in the ultrasound pathway, resulting from energy reflected from acoustically resistant media, such as air, bowel gas or bone, to tissues with strong acoustic absorbance, such as skin, muscle or the gastrointestinal tract, can lead to complications like skin burns or serious side effects like bowel perforation owing to thermal injury [3,16]. Beam scattering, diffraction and reflection therefore need to be prevented or carefully accounted for in planning and during delivery, and acoustic coupling of the transducer to the skin throughout treatment is necessary to avoid skin burns.

Compared with HIFU transducer focal volume, clinically relevant tissue target volumes may be very large. This means that the HIFU focus may have to be moved within the target volume to achieve sufficient tissue ablation, either by shooting the beam continuously while moving the transducer or by interrupting the beam and moving the HIFU focus. When combined with the need to frequently verify the location of the focus within the body by means of imaging and ensure unwanted energy deposition to avoid side effects, excessively long treatment times can result.

ADVANCES IN TRANSDUCER DESIGN AND BEAM FOCUS THAT COUNTER LIMITATIONS

HIFU transducers need to be optimally designed for specific clinical applications. The development of piezoactive materials with specific acoustic properties, e.g. lead zirconate titanate-type ceramics and composites of piezoactive elements, that are capable of being driven at high power and can be tailored for the specific clinical application, has been an important step. For curved transducers, the radius of curvature, which determines the distance at which the focal volume is located, and transducer diameter, which determines the surface area, are important parameters. In the case of arrays, the size and number of individual elements required to achieve appropriate acoustic power, their spatial distribution and their relationship to operating frequency are evaluated. Often elements are placed on a curved surface to achieve some geometric focusing [17,18]. The advantage of arrays is that the electrical signals applied to each element can be varied [19]. Using multichannel electronics, the acoustic fields produced by individual elements can be used coherently to produce a single focus that can be adjusted in size, shape and position and manoeuvred through a clinically relevant volume, or several foci can be created simultaneously. This increases the overall volume that can be ablated and achieves faster treatment times. An important factor in transducer array design is the compromise between performance, which favours a large number of elements, and cost and complexity, which favours a small number of elements. Many specific designs for arrays have now been reported. For example, ablation of large deep-seated tissue volumes has been reported with a 256-element phased array [20]; high-power beam steering through human skull was demonstrated with a 200-element sparse phased array [21]; high-power acoustic fields were achieved with an intracavitary 57-element aperiodic array device designed for prostate treatment [22]; and an endorectal transducer with 1000 elements for high-resolution treatment of prostate conditions has been clinically approved [23]. One drawback associated with arrays is that of “grating lobes” caused by sound energy spreading out from the transducer in undesired directions, which occurs when the element spacing is greater than a half wavelength. Several methods to minimise this have now been reported [21,22,24,25], including a patented random array design (Figure 2), which further reduces the time taken to deliver therapy and avoids delivering significant acoustic energy to non-targeted tissues, even when multiple simultaneous foci located off axis are produced [26].

 

Figure 2.

The spherical surface of a patented array transducer with randomly distributed elements that allows multiple simultaneous foci and minimises off-target energy delivery. Reproduced with permission from IOP Publishing Ltd, from Hand et al [26].

Design and testing of a HIFU system with flexible and controllable multifocus pattern ability is another important advance. Using a 256-element spherical section phased array system capable of producing “fit-to-shape” multifocus patterns, e.g. X, S, C, square and Q shapes, simulation and phantom experiments showed that treatment volumes could be up to 6.6 times greater in one sonication. Further, by using three-dimensional (3D) focus steering, it was feasible for other subarrays to operate if some of the elements were blocked by ribs, providing the device with the ability to avoid obstacles [27]. Advanced phased array systems with up to 20 000 elements, allowing 3D multiple foci sonication and rapid beam steering, are currently under further technological research and development.

The new transducer design has also allowed increased exploitation of mechanical effects to enhance ablation. Controlled use of cavitation can induce larger target lesions—thus achieving reduced treatment times—and research in this area is ongoing. A new approach using an endocavitary plane transducer showed that cavitation effects were induced beyond a threshold dose of acoustic intensity in ex vivo studies. Further, when the cavitation effect was combined with the thermal effect, it was possible to necrose cylindrical target volumes up to 31 cm3 in 4 min [28].

In transcranial HIFU, where skull ultrasound wave refraction can cause severe beam degradation, there have been recent important developments to improve focusing, including validation of an in vitro 3D CT adaptive correction method. A specifically designed 300-element spherical array therapeutic transducer was used in conjunction with CT scan acquisitions to deduce acoustic properties of the skull. Precise beam refocusing was achieved through ex vivo human and monkey skulls with a positioning error < .7mm. A later development by the same group, which used MR acoustic radiation force imaging for energy-based adaptive focusing in the human cadaver head, showed greater enhancement of transcranial ultrasound beam focusing [30], paving the way for in vivo human trans-skull FUS.
COMMERCIAL FUS DEVICES IN CLINICAL USE

There are currently two commercially available intracavitary FUS clinical devices: the Ablatherm® (EDAP TMS, Lyon, France), which was jointly developed with the French Institute of Medical Research in the early 1990s, and Sonablate® 500 (SonaCare Medical, Charloltte, NC, previously Focus Surgery Inc., Indianapolis, IN), which was developed in the USA in 1994 (Figure 3a,c). Both use a single moveable probe, are guided by ultrasound imaging, and have been employed in trials for treating prostate cancer [31]. They also have potential application in other pelvic malignancies. Ablatherm has a robotically controlled treatment probe with dual ultrasound transducers. Sonablate 500 has a single transducer and uses a split beam technology that increases the size of the focal zone and allows near simultaneous treatment and imaging. It is more operator dependent but has the practical advantage of being fully portable.

 


Figure 3.

Examples of high-intensity focused ultrasound devices currently in Western clinical use or research. (a) Ablatherm® (EDAP TMS, Lyon, France), (b) ExAblate® OR (InSightec, Haifa, Israel), (c) Sonablate® 500 (SonaCare Medical, Charlotte, NC, previously Focus Surgery Inc., Indianapolis, IN), (d) Sonalleve MR-HIFU (Philips Healthcare, Guildford, UK), (e) ExAblate Neuro (InSightec). Figures are reproduced with permission from the manufacturers.

Extracorporeal FUS devices offer a longer focal length than intracavitary devices and are more versatile overall. MRI-guided FUS (MRgFUS) extracorporeal machines include the ExAblate® system (InSightec, Haifa, Israel), which uses real-time thermometry MRI guidance (Figure 3b) and is currently used worldwide to treat uterine fibroids and in Europe to treat breast cancer, adenomyosis and bone metastasis, and is being investigated in clinical trials for other uses. The Sonalleve MR-HIFU is an alternative system (Philips Healthcare, Guildford, UK), which combines an extracorporeal HIFU system and MR coil elements integrated into a patient table compatible with Philips MRI platforms (Figure 3d). A novel electronic concentric circle beam path method is used to increase ablation volumes. Sonalleve has largely been used to treat uterine fibroids in countries other than the UK but is now under investigation for oncology applications. Extracorporeal ultrasound-guided FUS (USgFUS) machines are more popular in Asia. The Model JC focused ultrasound system (Haifu Technology Co. Ltd, Chongquing, China) originated in China. It can be operated using a choice of transducers with varying focal length and has been used to treat several cancer types including liver and renal cancer [10]. Alternative USgFUS machines include the HIFU-2001 (Sumo Corporation Ltd, Kowloon, Hong Kong) machine, which has been used since 2001 to treat cancer patients in China, Hong Kong and Korea, the HIFUNIT-9000 tumour therapy system (Shanghai Aishen Technology, Shanghai, China) and the FEP-BYTM system (Yuande Biomedical Engineering Lim. Co., Beijing, China). Extracorporeal devices specifically designed for transcranial FUS include the ExAblate Neuro hemispheric phased array HIFU system, which is currently used only for neurosurgery research purposes in brain disorders (Figure 3e) [32].

CLINICAL APPLICATIONS OF HIFU IN ONCOLOGY

Prostate cancer

During the last decade, many trials have assessed intracavitary FUS as a non-invasive alternative to prostatectomy and radiotherapy for localised prostate cancer. The UK National Institute for Health and Clinical Excellence initially supported the use of intracavitary HIFU ablation in the management of prostate cancer [33]. Although it is currently clinically used in other parts of the world, in the UK its use in the National Health Service has been recommended to be confined to clinical trials [34,35] and it is presently under interventional procedure consultation [36]. However, because of its organ-sparing and tumour control ability, retreatment potential, recent technical advances in delivery and imaging and recent promising trial results, HIFU is strengthening as a viable alternative treatment for tumour control—particularly for patients in whom localised cancer control with minimal morbidity or effective salvage are priorities [37]. For example, a review in 2009 on salvage HIFU following recurrent disease after radiotherapy reported biochemical disease-free rates, negative biopsy rates and complication rates similar to other salvage methods [38]. Similar results were reported in a 2011 study of 19 males treated with HIFU for locally recurrent prostate cancer following radical prostatectomy when good cancer control with acceptable morbidity was shown [39]. Both studies indicated better outcome for males with pre-treatment lower risk disease. A study of HIFU as salvage therapy in 22 Tokyo patients in 2011 also reported a good biochemical disease-free rate at 5 years of 52% [40]. The use of HIFU for focal salvage therapy following radiotherapy failure was also recently indicated to reduce the harms of whole-gland salvage therapies [41]. Moreover, recent encouraging results from a trial of HIFU as primary treatment in localised prostate cancer showed no histological evidence of cancer in 30 of 39 males biopsied at 6 months and a low rate of treatment-related genito-urinary side effects [42]. Non-invasive MRgFUS has also been used for prostate cancer ablation and has the advantage of improved targeting and real-time temperature monitoring, but only a few studies have been conducted with human patients [43].

Rectal tumours

Following surgery for rectal tumours, locally recurrent disease is a major concern that is often accompanied by severe pain and incapacitating complications. There is therefore an unmet clinical need for new treatments, especially for patients with residual or progressing disease in whom all current therapies have failed.

Recently, we reported the first case exploring the feasibility of intracavitary HIFU as a therapeutic option for tumour ablation in advanced rectal cancer. The patient had originally undergone surgical resection but developed recurrent local and liver metastatic disease with debilitating symptoms and was not fit for any conventional adjuvant options. Using the Sonablate 500 HIFU device, adjusted to deliver about 50% of the intensity per pulse used for prostate cancer treatment, the exophytic part of the tumour was targeted over 29 min. Symptoms improved within 24 h, there were no complications and repeat MRI at Day 7 showed tumour necrosis of the targeted area. Furthermore, the patient’s overall physical condition improved to the extent that palliative radiotherapy became possible [44]. A UK Phase I/II trial has since been initiated to further investigate the feasibility and efficacy of transrectal HIFU in patients with locally advanced rectal cancer (09/H0808/43).

Liver tumours

Surgical resection or transplantation has been the gold standard treatment for both primary and metastatic liver tumours. Since the first successful HIFU liver ablation in a male in 1993 [45], extracorporeal HIFU approaches have been investigated and developed, concentrating on patients with unresectable hepatocellular carcinoma (HCC) or in whom comorbidity prevents surgery. Particular challenges include beam propagation through the ribs, respiratory movement of the liver and long ablation times owing to large tumour size and small focal volume [46, 47]. The high prevalence of HCC in China has driven HIFU technology to overcome the associated challenges, with emerging encouraging results.

A large randomised study in China in 2005 using the Model JC Haifu system in patients with stage IVA HCC reported median survival time to be significantly longer in patients who received combined HIFU and transcatheter arterial chemoembolisation (TACE) therapy (11.3 months vs 4 months; p=0.004) [48]. A 2011 Chinese study of unresectable HCC showed slightly longer median survival of 12 months after combined HIFU+TACE treatment. 45% of patients achieved complete ablation, with ablation response reported as a significant prognostic factor [49]. For HIFU treatment alone, a report in 2011 of 49 patients from a Hong Kong cancer centre who received single HIFU treatment for unresectable HCC concluded that HIFU was an effective treatment modality with a high effectiveness rate and favourable survival outcome: complete tumour ablation was reported in 80% and local tumour control was 67% at 24 months [50]. However, serious complications have recently been reported in a minority of HCC patients, including rib fractures, diaphragmatic rupture, biliary obstruction, pleural effusion, pneumothorax and fistula formation [51]. These have arisen from unwanted thermal damage, indicating the need for caution and improved targeting of beam energy to lower risk.

Renal tumours

Many malignant renal lesions are small, so a non-invasive nephron-sparing therapeutic method is attractive. Initial studies, which used either multiple elements in a concave disc or the Storz investigational HIFU prototype device (Storz Medical, Schaffhausen, Switzerland), showed skin burns and problems with tissue ablation, inhibiting clinical use [45,52]. More contemporary extracorporeal and laparoscopic HIFU systems have produced smaller but better defined lesions and thus better results: however, they remain as investigative procedures, requiring improvements in order to compete with other ablative techniques [53]. A preliminary trial in patients with advanced renal cancers was carried out in 2003 using the Model JC Haifu device. A decrease in both flank pain (90%) and haematuria (89%) were reported with no adverse events [54]. A later study using the same device reported stable lesions in two-thirds of patients with minimal morbidity [55]. A Phase I study of laparoscopic HIFU in 2008 showed feasibility and demonstrated that this more invasive method helped to resolve the limitations caused by bowel, rib cage and abdominal wall obstruction and respiratory motion, although technological and methodological refinements were necessary to improve targeted ablation [56]. Feasibility, good tumour ablation and low morbidity with laparascopic HIFU was also shown in 2011 [57]. Methods in development, such as photoacoustic real-time monitoring [58] and respiration-induced movement correlation modelling [59], or the application of MR image guidance to monitor temperature changes for optimal heat deposition and safety [60] may improve future non-invasive renal FUS.

Pancreatic tumours

Most patients with pancreatic cancer present with inoperable disease, such that palliative treatment for local tumour control and pain relief are the main aims of treatment for which HIFU may have significant benefits. The long treatment times previously required owing to large target volumes are being addressed by the development of new multi-array devices as well as methodology harnessing mechanical tissue effects to enhance tissue ablation. Increased clinical experience is further enabling its development [16]. Early clinical studies in China supported HIFU as a primary therapy for pain relief, with no major adverse events reported [6163]. Recent studies have confirmed pain palliation and have also indicated efficacy. A report in 2009 of all stage unresectable patients in Peking, China, treated with an FEP-BY device showed pain improvement in 80.6% of patients, an overall median survival of 8.6 months and no complications [64]. A Phase II trial in 2010 of concurrent gemcitabine and HIFU in locally advanced pancreatic cancer using a HIFUNIT-9000 system also showed promising activity, with 78% pain relief rate, 43% response rate and a median survival rate of 12.6 months [65]. In a 2011 report of mixed stage inoperable patients treated with HIFU alone, an 87.5% pain relief rate, no complications and an 8-month median survival was shown [66]. In a recent European study in 2010, all six patients with tumours in difficult to treat locations showed pain relief and full tumour ablation, with one experiencing a serious complication [67]. A minority incidence of serious complications, including third-degree burns and fistula formation, has been separately reported [51].

Breast tumours

The breast is suited to HIFU treatment as it offers a soft-tissue acoustic window and can easily be immobilised. In a 2001 feasibility study of MRgFUS of 11 breast fibroadenomas using a custom-made device, 8 lesions indicated complete or partial tissue devascularisation and necrosis [68]. In 2003, MRgFUS using ExAblate as an adjunct to tamoxifen in patients with breast carcinoma reported negative biopsies in 19 of 24 patients at 6 months [69]. A 2007 study of MRgFUS using ExAblate in Japanese females with ductal carcinoma showed only 1 case of recurrence in 21 patients over a median follow-up of 14 months [70]. Similar favourable results have been reported in China using the Model JC USgFUS device, with a 95% 5-year disease-free survival rate [71] in one study and pathology confirming ablation in all cases in another [72]. However, limitations have included the risk of tissue damage to proximate skin, rib and lungs, which are currently being addressed by technological improvements, for example in device design [73], focal aberration correction [74] and novel contrast enhancement agents [75].

Bladder cancer

As ultrasound is commonly used as a first-line imaging method for investigation of urinary tract symptoms, HIFU offers an attractive means to visualise and treat bladder cancers at the same time. Encouraging results were reported in the first study of extracorporeal HIFU in superficial low-grade transitional cell bladder carcinoma, with no recurrence seen in 67% of treated patients [76]. However, the drawbacks of long treatment time and the need for regional anaesthesia require more research. Current interest is largely placed on ultrasound-based combination therapy [77].

Bone tumours

The first successful targeting of bone lesions using HIFU in animal models, causing necrosis of osteocytes, was reported in 2001 [78]. A key potential advantage for primary bone tumours is limb sparing. A recent study using the Model JC Haifu device showed that USgFUS was feasible and effective in primary bone malignancy. Complete tumour ablation was seen in 69 of 80 patients. Further, for patients whose tumours were completely ablated with HIFU and who completed systemic chemotherapy, the 5-year survival rate was greater than reported for other treatments [79]. Encouraging results have also been achieved for pain control of bone metastasis. MRgFUS for pain palliation in patients for whom other treatments were ineffective or not feasible showed HIFU was a safe and effective treatment option; 72% of patients reported significant pain improvement and a 67% reduction in opioid usage was recorded [80]. Supported by clinical studies, the ExAblate MRgFUS system received the European CE mark and US Food and Drug Administration approval for palliative treatment of bone metastasis in 2007 and 2012, respectively. The first UK trial testing of the Sonalleve MRgFUS system for bone metastasis is currently under way.

Brain tumours

There is great interest and potential use of HIFU in brain tumours. Enhancement of drug delivery across the blood–brain barrier (BBB) is a key active area of research, enabled by targeting BBB disruption [81,82]. However, for the development of effective and highly focused transcranial HIFU tissue ablation, physical problems caused by the skull have created significant technical hurdles (see section “Advances in transducer design and beam focus that counter limitations”). A recent pilot study in three glioblastoma patients using transcranial ExAblate showed focal heating was achieved, but greater device power was required to produce focal coagulative necrosis [83].

CONCLUSION

Non-invasive techniques that utilise HIFU to ablate tumours will enable improvements in future healthcare provision as patient morbidity can be minimised while potentially saving costs. The limitations of HIFU that have delayed its potential use in clinical practice are being overcome through advances in technology and design, ongoing research is enabling improvements and reducing risk, and experimental clinical trials for various types of tumours are showing considerable promise: for some tumour types, e.g. prostate and pancreatic cancer, randomised controlled trials are now required to compare FUS with standard treatments. Clinical applications of FUS are thus continuing to expand and improve and we predict that its benefits along with its increasingly clinically relevant fast treatment times will rapidly result in its adoption as a routine part of multimodal therapy for many cancers.

Thứ Tư, 27 tháng 3, 2013

ULTRASOUND FIRST FORUM PROCEEDINGS

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Selection of US Transducer in Practice


Transducer Characteristics for Imaging

This section discusses the criteria for linking which properties of ultrasound imaging transducers and their formats need to be identified for various clinical applications.3 What follows is primarily applicable to clinically used imaging transducers that operate in the frequency range of 1 to 20 MHz. Transducers operating above this frequency are used for special applications such as intravascular imaging (see Figure 4, F and G) or preclinical imaging of small animals but are also included in the discussion wherever possible.

Acoustic Windows

How well is the type of transducer suited to the “acoustic window” or location where it makes contact with the body to visualize the organs or tissues of interest? Standard acoustic windows provide an unobstructed view of an organ or region; many, by convention, have specific names, such as “transabdominal” or “parasternal long-axis,” so that images can be compared and described consistently.3 Typical windows are located in or on the following general regions of the body: head, chest, abdomen, pelvis, limbs, vessels, and various orifices of the body. Transducers can be associated with certain regions through Latin prefixes: “trans,” through or across; “intra,” into or inside; and “endo,” within, etc. An example is transthoracic, a category that includes transducers that image through the chest. Transcranial describes a transducer that images the head through the skull.

As already mentioned, for the transthoracic window, the phased array would be the most appropriate if the imaging task requires the transducer to be placed between the ribs; it is designed to fit into intercostal spaces and maximize the scanned area (format 2 of Figure 3). For most contact surfaces that are relatively flat and/or slightly deformable (eg, ones used for small parts or vascular imaging), the most general and frequently used transducer type is the linear array designed to make contact with flat surfaces, with the footprint decreasing in size with increasing frequency. Here rectangular and trapezoidal formats (1 and 4 of Figure 3) provide appropriate viewing areas. With abdominal imaging, to increase the viewing area with minimal increases of the contact area, convex arrays (Figure 4C) produce format 3 (shown in Figure 3) and are designed to make surface contact in deformable soft areas of the body.

Specialized Transducers

Specialized transducers are designed to operate inside the body. These include transesophageal probes that are phased arrays suitable for manual manipulation within the esophagus, format 2 and transducer type B in Figure 4. A number of other specialty probes have also been developed for interventional or surgical use such as laparoscopic arrays and intracardiac arrays. These probes can be either linear or phased arrays, depending on the application and access windows. Several endo probes: such as endovaginal, endorectal, and endocavity (type D shapes), are functionally like end-fire phased arrays (format 2 and Figure 4B) or convex arrays (format 3 and Figure 4C) at the end of a small-diameter cylindrically shaped handle to fit in orifices and yet maximize the FOVs. Another example is the intravascular ultrasound transducer (Figure 4H), which is inserted in veins to produce an image plane, format 8 or scanned volume, format 9.

Resolution and Penetration

The selected scan depth allows viewing over the range of interest. Factors involved in imaging capability include the size of the active aperture (occult to the user, but typically a low F number [F#; focal depth/active aperture width] value of 1–2 is used), the transmit focal depth, and time-gain control settings available. Penetration is the minimum scan depth at which electronic noise is visible, despite optimization of available controls (usually at the deepest transmit focal setting and maximum gain), and electronic noise stays at a fixed depth even when the array is moved laterally. Penetration is primarily determined by the center frequency of the transducer: the higher the frequency, the shallower the penetration because the absorption of the ultrasound wave traveling through tissue increases with frequency.

A useful first approximation for estimating a depth of penetration (dp) for a given frequency is dp = 60/f cm-MHz, where f is given in megahertz. Thus, one might expect a 6-cm penetration from a 10-MHz center frequency transducer. As noted earlier, the absorption coefficient (acoustic power loss per unit depth) is a function of frequency and varies from tissue to tissue (values for soft tissues range from 0.6 to 1.0 dB/cm-MHz4). A more general term describing acoustic loss is the attenuation coefficient, which includes additional losses due to scattering and diffusion and hence is always greater than the absorption coefficient. The attenuation coefficient is highly patient and acoustic path dependent.

To optimize image resolution, users and manufacturers have worked on increasing the imaging frequencies for the various examination types. For example, some 30 years ago, people might have imaged the abdomen with a frequency of 2.25 MHz, whereas today the number is more often 3.5 MHz with some obstetric and gynecologic imaging reaching up to 5 MHz.5 Similarly, the last decade has seen a steady increase in breast imaging reaching the low teens in megahertz.

Transducer Properties and Imaging

Other criteria to be included in the above-discussed selection process are the transducer efficiency,2 transducer-system design, system signal-to-noise ratio, and, as already noted, absorption of the tissues being imaged. A major factor is absorption, the compositions and relative positions of different tissue types in the acoustic path. For example, a thick layer of adipose tissue will reduce penetration due to refractive or aberration errors in the acoustic path to the site of interest. Similarly increased amounts of amniotic fluid with fetal imaging enhance penetration and may permit the use of frequencies higher than those ordinarily used at the given scanning site.

The frequency range, or bandwidth,1,2,6 of the transducer will determine whether it can support B-mode imaging at different center frequencies and also operate in Doppler, harmonic, and color flow modes. With Doppler-based imaging modes, we often need to operate with lower frequencies than the B-mode frequency to minimize aliasing. With harmonic imaging, by definition, one uses a receive frequency that is a multiple (usually 2) of the transmitted frequency; hence the need for the wide bandwidth. The bandwidth and focusing properties will also influence image resolution. In clinical practice, it is essential to ensure that the image obtained can discern the smallest possible dimensions in both the lateral and axial directions.

Finally, the number of individual transducer elements is of interest because the number of active elements (with the exception of phased arrays or angularly scanned 2D arrays) determines the lateral extent or width of the image. For phased arrays, an increasing number of elements is associated with improved resolution and penetration depth. For 2D arrays (usually symmetric), the number of elements along the x and y directions determines the extent of the volume for linearly scanned arrays. For a 2D phased array, resolution and penetration increase with a greater number of elements along the x and y directions, but the angular shape or FOV remains the same, independent of the number of active elements used. The focusing in a fixed direction can indirectly affect imaging because focusing is positioned at only one depth and is poorer elsewhere. For 3D imaging, mechanically scanned 2D arrays suffer from the same fixed elevation focal depth limitation encountered in 2D imaging. In contrast, all the elements of fully populated 3D imaging or matrix arrays focus electronically at one point in both azimuth and elevation planes to provide far better resolution.

At the deepest depths, it is the maximum number of available active channels in the system that determines the resolution (along with strength of focusing and system noise). Spatial resolution is generally poorer (typically by a factor of 2) than the temporal resolution along scan lines; in the discussion presented here, resolution refers to spatial resolution, unless noted otherwise. For phased arrays, the number of channels usually corresponds to the maximum number of elements. As a general rule of thumb, because elements are typically on half-wavelength spacing, the more elements, the better the spatial resolution, which is inversely proportional to the active aperture in wavelengths. For example, a 64-element array, 32-wavelength aperture will have maximum spatial resolution 2-fold lower (wider beam) than that of a 128-element, 64-wavelength array. In the case of a linear array, which may have several hundred elements, the number of elements determines the lateral extent of the image, but it is the number of active channels that governs the resolution. For these 1D arrays, the resolution out of the imaging plane (also known as slice thickness) is poor except near the fixed elevation focal length. For 2D arrays, the spatial resolution is inversely proportional to the active apertures that form the sides of the 2D array. Two-dimensional arrays have superior resolution compared to 1D array focusing with fixed elevation focusing because true-point focusing can be achieved simultaneously in azimuth and elevation for 3D imaging.

Another way of looking at resolution is F#. The smaller the F#, the better the resolution.1 A simple estimate of the full (beam) width in millimeters, a common measure of resolution, neglecting absorption, is approximately F# × λ, where λ is wavelength (1.5 mm/μs/f [MHz]). For example, resolution would be 0.3 mm at 5 MHz for an F# = 1. Focal depths are also active aperture dependent. For example, for a 128-element 64-wavelength array, the deepest focal depth achieved at maximum aperture and an F# = 1 is F = F# × L = 64 wavelengths. The actual penetration or useable scan depth would, of course, be deeper than the maximum focal depth.

Matching Transducers to Clinical Applications

Now that transducer types and properties have been related to imaging and acoustic windows, they can contribute to the selection of transducers for specific clinical applications.3 The appropriateness of certain transducers to specific applications evolved historically and through specific tailored designs. The primary considerations are the target region of interest and its extent and the available acoustic windows needed for access.

Abdominal Imaging

When transducer arrays were initially introduced commercially for abdominal imaging (including obstetrics and gynecology) in the 1970s, they were of the linear array type (type A in Figure 4 with image format 1 of Figure 3). In most cases, the contact area with the patient was not a critical issue, and some of these linear arrays were quite long (eg, 8 cm) to cover, say, the third-trimester fetal head.3,5 However, it was soon realized that one could achieve sufficiently large coverage by the use of curved or convex arrays (type C in Figure 4) without incurring the penalty of having to manipulate the rather unwieldy linear array transducers.

The convex arrays (Figure 4C) are the tools of choice for most general 2D imaging applications involving the abdomen. The general form factor, related to ergonomic factors and the suitability of the transducer shape and FOV to the application, for abdominal 3D imaging is still evolving. Three key descriptors for these arrays are the footprint (overall aperture size), FOV, and radius of curvature (Figure 1C). The footprint describes the contact area usually in the form of a rectangle, circle or ellipse. Even though for abdominal imaging, access is not usually a concern, when these types of transducers are considered for new applications, window access is of primary importance. The radius of curvature and FOV (expressed in degrees of maximum angular coverage) are related to image extent and coverage. Advanced signal processing has been added to some systems to increase penetration; however, this feature is usually only available on certain probes.

For the mechanical 3D probes, the currently preferred form factor is a mechanically swept convex array (Figure 4G and format 6 in Figure 3); however, fully electronic convex 2D arrays are now becoming available. For these cases, two FOVs are given for the orthogonal scan directions. Alternatively, phased arrays, because of their small footprint and wide sector image format, are also used for abdominal imaging. Finally, 2D or matrix arrays are becoming increasingly prevalent for these applications because of their superior image quality, resolution, and ease of use.

Intercostal Imaging

The primary applications of this imaging grouping are cardiac scanning and examination of the liver from between the ribs. Simply because of the restrictive anatomy and the limited acoustic windows caused by the ribs and the often encroaching lungs, the transducer choice here is limited to phased arrays (Figure 4B). Even in this area, initial attempts were made to use linear arrays; however, these were rapidly dropped due to the rib shadowing and the superiority of phased array transducer format 2. For cardiac applications, the probes tend to have array dimensions on the order of 20 × 14 mm depending on the manufacturer. The patient contact area will be slightly larger. These numbers have evolved over the last 40 years and depend on a number of things, such as the general size of the patient population. Age is another consideration; rib spacing and the depth penetration needed vary as children mature into adults.

For noncardiac intercostal applications, the dimensions of the arrays are somewhat larger. As noted earlier, the existence of these anatomic limitations creates an upper performance limit for spatial resolution since resolution performance is inversely related to the size of the aperture, as explained above. In both cardiac and general intercostal imaging applications, the imaging depth is deep (depending on the patient size, it may be as deep as 24 cm), forcing the use of lower (1–3.5 MHz) frequencies and resulting in some further loss of imaging performance.

There is an interesting aspect of cardiac imaging that has had a profound effect on the nature of the probes. Due to the presence of the ribs and other acoustically hostile tissue in the ray path, echocardiography suffers from imaging artifacts due to reverberant noise. The introduction of harmonic imaging has proven to be highly successful in reducing this noise. As a consequence, the importance of transducer bandwidth has become critical in cardiac transducer design. Today, most cardiac systems transmit at frequencies between 1.5 and 2.0 MHz and, of course, receive signals at frequencies twice that range.

A major development in cardiac imaging was the implementation of fully populated 2D or matrix arrays (type E) containing thousands (typically 50 × 50 or so) of elements. These make possible real-time (4D) depiction of pyramidal volumes (format 7, Figure 3), visualization of arbitrary cut planes, and 4D cardiac imaging and color flow imaging. In addition, true electronic focusing in both the xz and yz planes provides superior resolution in comparison with all other 1D array transducers.

Superficial and Breast Imaging

This category refers to “superficial” imaging of carotids, leg veins, breasts, thyroids, testicles, etc and includes the categories of small parts, musculoskeletal, and peripheral vascular imaging. It is the last bastion of the application of linear arrays (type A), which formed the starting design type for the applications discussed earlier. In this clinical category, access is usually not an issue, and the sizes of the probes themselves can be small (because of the use of high 7- to 15-MHz frequencies and the resultant small element sizes). Musculoskeletal applications for imaging muscles, ligaments, and tendons also use arrays of this type. In the last 10 years, breast imaging has gone to very high frequencies (eg, 14 MHz), while imaging of the peripheral vasculature has remained at lower (about 3–11 MHz) values due to the need to include deeper leg veins and successful Doppler performance. Usually the capability of the array to add trapezoidal imaging (format 4) is a considerable advantage. As in abdominal imaging, 3D imaging with mechanically swept probes or electronic 2D arrays is now available for superficial and breast applications, greatly improving the coverage available and image quality. For applications involving imaging vasculature, some probes have advantages of including modes that enhance flow visualization.

Obstetrics and Gynecology

At the present time, mechanically scanned convex or linear arrays (types G and F) are used widely to provide 3D and 4D imaging of fetuses in vivo (formats 5–7). Matrix or fully populated 2D arrays (type E) are also available for this application (typically format 7).

For gynecology, specialized endo-array probe shapes are used (type D). Typically, the arrays are at the end of the probe (end-fire arrays) and are convex or curved arrays with wide FOVs (format 3); however, phased arrays in an endo-array package (type D) can also be used (format 2). Frequencies used are typically 5 MHz and higher. As in other applications, 2D arrays have been designed for 3D imaging of these cases.

Neonatal and Pediatric

Pediatric transducers tend to have smaller footprints than transducers used for adults, applications A–C, and operate at higher-frequency varieties (≥7 MHz) of those that are used for adults. Depending on the body region, types of transducers similar to those for adults are applicable. Phased arrays (type B) and 3D transducers (types E and G) are appropriate for cardiac imaging. Other arrays that are also useful for these clinical needs include static (2D) and, for 3D, mechanically swept linear arrays and convex arrays.

Intracavity Probes

Intracavity probes constitute a large group of specialty transducers that are designed to image inside the body cavity. Transesophageal transducers are used to enable imaging of internal organs, especially the heart, from inside the esophagus (see Figure 5). They use higher frequencies (≥5 MHz) and are implemented as phased arrays with manipulators and motors to adjust the orientation of the transducer. Miniature transesophageal 2D arrays offer electronic scanning for 3D and 4D imaging.

Transducers can be highly specialized for viewing usually within body openings or vessels. Intracardiac phased arrays are inserted through a vessel to gain access to the inner chambers of the heart. Surgical specialty probes include laparoscopic arrays inserted through small incisions to image and aid in laparoscopic surgery (similar to endo probes); these are remarkable for their FOV despite small diameters. Intraoperative arrays are specially shaped to be placed on vessels, organs, and regions made accessible during open surgery (see Figure 5). Others in this class are surgical and interventional probes with unique shapes (see Figure 5).

As already noted, the probes that are inserted into the body are designed to fit through small openings and have a wide FOV (90°–150°). These probes include transrectal (or endorectal) for imaging of the pelvic region using the anus for access and the already described endovaginal (also called transvaginal) for imaging the female pelvis and reproductive organs using the vagina as entry for gynecologic and obstetric applications. These endo probes, described earlier, are cylindrical to fit into small orifices and have convex arrays (typically 3–9 MHz) at their ends with large fields of view, biplanes, or mechanically swept convex arrays. Probes for urologic applications include the biplane type.

A unique transducer is the biplane probe, which consists of two orthogonal arrays producing images in planes xz and yz. Typically the arrays are small (8–12 mm) and of the convex type. Each format and transducer would correspond to those of a single–imaging plane transducer such as format 3 of Figure 3 and the convex array of Figure 4C. However, sector or linear array formats are also possible, depending on the transducer construction, so that several combinations can be used in practice. Alternatively, a subset of the imaging capability of a 2D array is the simultaneous presentation of two orthogonal 2D images.

Intravascular transducers are inserted into blood vessels to image the vessel walls for various pathologic conditions (type H and formats 8 and 9). They are most often mechanically rotated single transducers with frequencies greater than 20 MHz and dedicated imaging systems, although there are also tiny (about 2-mm-diameter) arrays designed for this purpose.

Head Probes

Transcranial imaging of the brain and its vasculature is conducted through limited acoustic windows through the skull such as the temples or eyes. Transorbital arrays are high-frequency (typically >20 MHz) ophthalmologic transducers and are used to image the eye or use the eye as an acoustic window. Transcranial probes are usually lower-frequency (1–4 MHz) phased arrays used to image blood vessels within the skull through the temples as windows.

Conclusions

Many ultrasound imaging transducers are designed to operate in certain regions of the body for specific applications. A primary objective of this article is to provide a systematic approach that would aid in matching a transducer to a clinical application, starting with the acoustic window and the region and depth to be imaged. To this end, a checklist for selecting a transducer is given in Table 1.

 

As indicated earlier, the first consideration for imaging a target region or organs is the access: the intended acoustic window. The transducer type must provide access through the selected acoustic window. The transducer type is linked with the image format, and common selections previously discussed include the linear, phased, convex, and 2D arrays. The size or transducer footprint must fit within the window, and in extreme cases in which the transducer window is an orifice, the transducer shape must conform to the available opening. As noted above, in some applications, specialty probes, such as endorectal transducers, are needed that are small enough in diameter (size) and have the elongated shape suitable for entering a body orifice.

Second, the size or FOV and image format are selected to obtain the desired coverage over the region of interest. Here both the scan depth and image width or FOV are important. For linear arrays, the availability of trapezoidal imaging may be necessary for adequate coverage. For 3D or volumetric imaging, the extent of the image may be given as a set of maximum scan angles in orthogonal directions or a FOV and an angle. A somewhat more hidden parameter for 2D imaging for determining coverage for the region of interest is the elevation focal depth that describes the region with the thinnest slice thickness.

Third, the maximum scan depth selected determines the highest achievable frequency through the penetration relation given in the “Resolution and Penetration” section above. For example, if the scan depth is 10 cm, then, as already discussed in the “Resolution and Penetration” section, the frequency from the penetration depth d is equal to 60/d = 60/10 = 6 MHz. This frequency provides an estimate of the best lateral resolution of about 1 wavelength for an F# = 1, or, for this example, the resolution is λ = c/f = 0.25 mm (from the “Transducer Properties and Imaging” section). Exceptions to this rule are systems that use advanced signal processing to enhance sensitivity and increase penetration. In addition, the use of piezoelectric materials such as piezo composites or domain-engineered single crystals can increase sensitivity or, correspondingly, penetration depth.6

Fourth, the coverage of essential diagnostic imaging modes can be determined. From the manufacturer-provided data, the effective bandwidth needed to support different modes of interest may be extracted, or for the system considered, the actual modes of interest may be listed, such as pulsed wave Doppler, multiple imaging frequencies available, or elastographic mode. Transducers with piezoelectric materials such as piezo composites or domain-engineered single crystals can increase bandwidth substantially.6

In conclusion, transducers and image formats have evolved to better suit specific clinical applications. The classification and organization given in this article provide the background for the selection of a transducer for a particular purpose. In addition, the understanding provided can aid in determining transducer characteristics needed for new cases, thereby extending the range of transducer use.