Abstract:
An ultrasound transducer for use in intra-vascular ultrasound (IVUS) imaging systems including a single crystal composite (SCC) layer is provided. The transducer has a front electrode on a side of the SCC layer; and a back electrode on the opposite side of the SCC layer. The SCC layer may have a dish shape including pillars made of a single crystal piezo-electric material embedded in a polymer matrix. Also provided is an ultrasound transducer as above, with the back electrode split into two electrodes electrically decoupled from one another. A method of forming an ultrasound transducer as above is also provided. An IVUS imaging system is provided, including an ultrasound transducer rotationally disposed within an elongate member; an actuator; and a control system controlling activation of the ultrasound transducer to facilitate imaging.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     The present application claims priority to and the benefit of U.S. Provisional Patent Application No. 61/745,425, filed Dec. 21, 2012, which is hereby incorporated by reference in its entirety. 
    
    
     TECHNICAL FIELD 
     The present disclosure relates generally to intravascular ultrasound (IVUS) imaging inside the living body and, in particular, to an IVUS imaging catheter that relies on a mechanically-scanned ultrasound transducer, including embodiments where the transducer includes a single crystal composite material. 
     BACKGROUND 
     Intravascular ultrasound (IVUS) imaging is widely used in interventional cardiology as a diagnostic tool for a diseased vessel, such as an artery, within the human body to determine the need for treatment, to guide the intervention, and/or to assess its effectiveness. IVUS imaging uses ultrasound echoes to create an image of the vessel of interest. The ultrasound waves pass easily through most tissues and blood, but they are partially reflected from discontinuities arising from tissue structures (such as the various layers of the vessel wall), red blood cells, and other features of interest. The IVUS imaging system, which is connected to the IVUS catheter by way of a patient interface module (PIM), processes the received ultrasound echoes to produce a cross-sectional image of the vessel where the catheter is placed. 
     In a typical rotational IVUS catheter, a single ultrasound transducer element is located at the tip of a flexible driveshaft that spins inside a plastic sheath inserted into the vessel of interest. The transducer element is oriented such that the ultrasound beam propagates generally perpendicular to the axis of the catheter. A fluid-filled sheath protects the vessel tissue from the spinning transducer and driveshaft while permitting ultrasound signals to freely propagate from the transducer into the tissue and back. As the driveshaft rotates (typically at 30 revolutions per second), the transducer is periodically excited with a high voltage pulse to emit a short burst of ultrasound. The same transducer then listens for the returning echoes reflected from various tissue structures, and the IVUS imaging system assembles a two dimensional display of the vessel cross-section from a sequence of these pulse/acquisition cycles occurring during a single revolution of the transducer. 
     In the typical rotational IVUS catheter, the ultrasound transducer is a piezoelectric ceramic element with low electrical impedance capable of directly driving an electrical cable connecting the transducer to the imaging system hardware. In this case, a single pair of electrical leads (or coaxial cable) can be used to carry the transmit pulse from the system to the transducer and to carry the received echo signals from the transducer back to the imaging system by way of a patient interface module (“PIM”) where the echo signals can be assembled into an image. An important complication in this electrical interface is how to transport the electrical signal across a rotating mechanical junction. Since the catheter driveshaft and transducer are spinning (in order to scan a cross-section of the artery) and the imaging system hardware is stationary, there must be an electromechanical interface where the electrical signal traverses the rotating junction. In rotational IVUS imaging systems, this problem can be solved by a variety of different approaches, including the use of rotary transformers, slip rings, rotary capacitors, etc. 
     While existing IVUS catheters deliver useful diagnostic information, there is a need for enhanced image quality to provide more valuable insight into the vessel condition. For further improvement in image quality in rotational IVUS, it is desirable to use a transducer with broader bandwidth and to incorporate focusing into the transducer. A piezoelectric micro-machined ultrasound transducer (PMUT) fabricated using a polymer piezoelectric material offers greater than 100% bandwidth for optimum resolution in the radial direction, and a spherically-focused aperture for optimum azimuthal and elevation resolution. While this polymer PMUT technology offers many advantages, the electrical impedance of the PMUT is too high to efficiently drive the electrical cable connecting the transducer to the IVUS imaging system by way of the PIM. Furthermore, the transmit efficiency of polymer piezoelectric material is much lower compared to that of the traditional lead-zirconate-titanate (PZT) ceramic piezoelectric. Therefore, the signal-to-noise ratio of a PMUT will be compromised unless the deficiency in acoustic output can be compensated for by improved transmit electronics and/or other signal processing advances. 
     Current approaches to form a focused ultrasound beam include the use of an acoustic lens using conventional PZT transducers. For example, a rubber lens with an acoustic velocity of 1.0 mm/μsec has been used for elevation focus in phased array ultrasound systems. These approaches pose complex fabrication problems and the difficulty of removing imaging artifacts in the resulting signal. 
     Accordingly, there remains a need for improved devices, systems, and methods for implementing focused piezoelectric micro-machined ultrasonic transducers within an intravascular ultrasound system. 
     SUMMARY 
     According to some embodiments, an ultrasound transducer for use in intra-vascular ultrasound (IVUS) imaging systems is provided that includes a single crystal composite (SCC) layer; a front electrode on a side of the SCC layer; and a back electrode on the opposite side of the SCC layer. In some embodiments, the SCC layer includes pillars made of a single crystal piezo-electric material. The pillars are embedded in a polymer matrix in some instances. The SCC layer has a dish shape, defined by a concave surface and opposing convex surface, in some embodiments. The back electrode is split into two electrodes electrically decoupled from one another in some implementations. 
     A method of forming an ultrasound transducer for use in IVUS imaging systems in some embodiments includes etching a single crystal; forming a polymer layer on the etched single crystal to form a single crystal composite (SCC) having a first thickness; placing a first electrode on a first side of the SCC; forming the SCC to a second thickness; placing a second electrode on a second side of the SCC; and placing the SCC on a molded tip. 
     An IVUS imaging system according to some embodiments may include an ultrasound emitter and receiver rotationally disposed within an elongate member; an actuator coupled to the ultrasound emitter, the actuator moving the ultrasound emitter through at least a portion of an arc; and a control system controlling the emission of a sequence of pulses from the ultrasound emitter and receiving from the receiver ultrasound echo data associated with the pulses, the control system processing the ultrasound echo data to generate a cross-sectional image of the vessel. In some embodiments the ultrasound emitter and receiver comprises an ultrasound transducer including a single crystal composite (SCC) layer; a front electrode; and a back electrode. In some embodiments the SCC layer includes pillars made of a single crystal piezo-electric material. The pillars are embedded in a polymer matrix in some instances. The SCC layer has a dish shape, with opposing concave and convex surfaces, in some embodiments. 
     These and other embodiments of the present disclosure will be described in further detail below with reference to the following drawings. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG. 1  is a schematic view of an imaging system according to an embodiment of the present disclosure. 
         FIG. 2  is a diagrammatic, partial cutaway perspective view of an imaging device according to an embodiment of the present disclosure. 
         FIG. 3  shows a partial view of an ultrasound transducer according to an embodiment of the present disclosure. 
         FIG. 4  shows a partial cross-sectional side view of a distal portion of an imaging device according embodiment of the present disclosure. 
         FIG. 5A  shows a partial cross-sectional axial view of the distal portion of the imaging device of  FIG. 4  along section line A-A′. 
         FIG. 5B  shows a partial cross-sectional axial view of the distal portion of the imaging device of  FIG. 4  along section line B-B′. 
         FIG. 6A  shows a partial plan view of a single crystal composite according to an embodiment of the present disclosure. 
         FIG. 6B  shows a partial plan view of a single crystal composite according to another embodiment of the present disclosure. 
         FIG. 6C  shows a partial plan view of a single crystal composite according to yet another embodiment of the present disclosure. 
         FIG. 7A  shows a partial side view of an ultrasound transducer according to an embodiment of the present disclosure. 
         FIG. 7B  shows a partial plan view of a distal portion of an imaging device incorporating the ultrasound transducer of  FIG. 7A  according to an embodiment of the present disclosure. 
         FIG. 7C  shows a partial plan view of the ultrasound transducer of  FIG. 7A  according to an embodiment of the present disclosure. 
         FIGS. 8A-F  show a series of partial cross-sectional side views of fabrication stages for an ultrasound transducer according to some embodiments of the present disclosure. 
         FIG. 9  shows a flow chart for a method of forming an ultrasound transducer according to some embodiments of the present disclosure. 
     
    
    
     In the figures, elements having the same reference number have the same or similar functions and/or features. 
     DETAILED DESCRIPTION 
     For the purposes of promoting an understanding of the principles of the present disclosure, reference will now be made to the embodiments illustrated in the drawings, and specific language will be used to describe the same. It is nevertheless understood that no limitation to the scope of the disclosure is intended. Any alterations and further modifications to the described devices, systems, and methods, and any further application of the principles of the present disclosure are fully contemplated and included within the present disclosure as would normally occur to one skilled in the art to which the disclosure relates. In particular, it is fully contemplated that the features, components, and/or steps described with respect to one embodiment may be combined with the features, components, and/or steps described with respect to other embodiments of the present disclosure. For the sake of brevity, however, the numerous iterations of these combinations will not be described separately. 
     Embodiments disclosed herein are for an apparatus and a method of fabrication of the apparatus, the apparatus including a focused transducer to be used in a rotational IVUS catheter. A transducer as disclosed herein provides a broad bandwidth of ultrasound signals having focused beam propagation. Such an ultrasound beam provides a high three-dimensional (3D) resolution for ultra-sound imaging, including depth, lateral and elevation dimensions. In some embodiments, an IVUS catheter of the present disclosure provides a wide bandwidth, focused ultrasound beam without increasing the number of electrical connections to a circuit rotating together with the transducer. An ultrasound transducer according to embodiments disclosed herein may include a single crystal composite material that provides a wide bandwidth, focused beam. The single crystal composite material is shaped into an element having a curvature designed to provide a focused beam (e.g., defining a concave emitting surface for the ultrasound transducer) in some instances. 
       FIG. 1  shows an IVUS imaging system  100  according to an embodiment of the present disclosure. In some embodiments of the present disclosure, the IVUS imaging system  100  is a rotational IVUS imaging system. In that regard, the main components of the rotational IVUS imaging system are a rotational IVUS catheter  102 , a patient interface module (PIM)  104 , an IVUS console or processing system  106 , and a monitor  108  to display the IVUS images generated by the IVUS console  106 . Catheter  102  includes an ultrasound transducer  150  in some embodiments. PIM  104  implements the appropriate interface specifications to support catheter  102 . According to some embodiments, PIM  104  generates a sequence of transmit pulse signals and control waveforms to regulate the operation of ultrasound transducer  150 . PIM  104  may also receive a response signal form transducer  150  through the same pair of lines. 
     Ultrasound transducer  150  transmits ultrasound signals towards the vessel tissue based on the trigger signals received from PIM  104 . Ultrasound transducer  150  also converts echo signals received from the vessel tissue into electrical signals that are communicated to PIM  104 . PIM  104  also supplies high- and low-voltage DC power supplies to the rotational IVUS catheter  102 . In some embodiments, PIM  104  delivers a DC voltage to transducer  150  across a rotational interface. Options for delivering DC power across a rotating interface include the use of slip-rings, rotary transformers, and/or the implementation of the active spinner technology. 
       FIG. 2  shows a diagrammatic, partial cutaway perspective view of catheter  102 , according to an embodiment of the present disclosure.  FIG. 2  shows additional detail regarding rotational IVUS catheter  102 . Rotational catheter  102  includes an imaging core  110  and an outer catheter/sheath assembly  112 . Imaging core  110  includes a flexible drive shaft that is terminated at the proximal end by a rotational interface  114  providing electrical and mechanical coupling to PIM  104  (cf.  FIG. 1 ). The distal end of the flexible drive shaft of the imaging core  110  is coupled to a transducer housing  116  containing ultrasound transducer  150  and associated circuitry. 
     Catheter/sheath assembly  112  includes a hub  118  supporting rotational interface  114  and provides a bearing surface and a fluid seal between rotating and non-rotating elements of catheter  102 . In some embodiments, hub  118  includes a luer lock flush port  120  through which saline is injected to flush out the air and fill the inner lumen of the sheath with an ultrasound-compatible fluid at the time of use of the catheter. Saline also provides a biocompatible lubricant for the rotating driveshaft. In some implementations, hub  118  is coupled to a telescope  122  that includes nested tubular elements and a sliding fluid seal that permits catheter/sheath assembly  112  to be lengthened or shortened. Telescope  122  facilitates axial movement of the transducer housing within an acoustically transparent window  124  at the distal portion of catheter  102 . 
     In some embodiments, window  124  is composed of thin-walled plastic tubing fabricated from material(s) that readily conduct ultrasound waves between the transducer and the vessel tissue with minimal attenuation, reflection, or refraction. A proximal shaft  126  of catheter/sheath assembly  112  bridges the segment between telescope  122  and window  124 . In some embodiments, proximal shaft  126  is composed of a material or composite that provides a lubricious internal lumen and optimum stiffness to catheter  102 . Embodiments of window  124  and proximal shaft  126  in catheter  102  may be as described in detail un US Pat. Application entitled “Intravascular Ultrasound Catheter for Minimizing Image Distortion,” U.S. Patent Application No. 61/746,958 filed on Dec. 28, 2012, now published as U.S. Patent Application Publication No. 2014/0187964 A1 on Jul. 13, 2014, the contents of which are hereby incorporated in their entirety by reference, for all purposes. 
       FIG. 3  shows a partial view of ultrasound transducer  150  according to some embodiments disclosed herein. Transducer  150  includes a single crystal composite material (SCC)  301  having pillars  320  of a single crystal piezo-electric material embedded in a polymer matrix  330 . In some embodiments polymer matrix  330  is formed by epoxy. The epoxy used as filler in polymer matrix  330  provides flexibility to the SCC material forming ultrasound transducer  150 . 
     In some embodiments, an impedance matching layer  310  is included in ultrasound transducer  150 . Impedance matching layer  310  facilitates coupling of the acoustic wave with the medium surrounding ultrasound transducer  150 . Soft polymer matrix  330  reduces the acoustic impedance to SCC  301 , thus providing high efficiency and broad bandwidth to transducer  150  for acoustic coupling. In some embodiments, matching layer  310  may be a quarter-wave matching layer added to SCC  301 , to further improve efficiency and bandwidth of transducer  150 , thus enhancing sensitivity. 
     According to some embodiments disclosed herein, pillars  320  form structures elongated in an axial direction (Y-axis in  FIG. 3 ) having a narrow diameter in cross section (Z-axis in  FIG. 3 ). The cross section of pillars  320  is in a plane of SCC  301  forming ultrasound transducer  150 . Further according to some embodiments, polymer matrix  330  is continuous in the axial direction (Y-axis) and in the plane of SCC  301  forming ultrasound transducer  150  (XZ-plane). The anisotropic nature of SCC  301  confines the electric field E within pillars  320 , which have a high dielectric constant. Fringe fields at the edges of electrodes  151  and  152  are mitigated by polymer matrix  330 . Thus, in some embodiments the performance of transducer  150  is not degraded by fringe fields at electrode boundaries. According to some embodiments, the thickness (or height) of pillars  320  may be 50 μm or less, for 40 MHz center frequency operation. In some embodiments a thickness-to-width aspect ratio of at least 2 or greater may be desirable, resulting in pillars  320  having a diameter of 20 μm or less. 
     Accordingly, SCC  301  can be made using deep reactive ion etching (DRIE) applied to a single crystal material. Etch a matrix pattern using DRIA and fill the etched trenches with epoxy. Then grind away back side and polish front side and have a resulting composite layer. A horizontal resonant frequency (oscillations in the XZ plane in  FIG. 3 ) is so far apart from vertical frequency (oscillations along the Y-axis in  FIG. 3 ) there is little energy expended by horizontal resonance. This makes the transducer more efficient. Wide bandwidth is achieved by efficiently coupling into medium (for example, using a matching layer). A matching layer overcomes acoustic impedance mismatch between transducer material and the transmitting medium. A PZT has impedance of about 30 while that of blood/saline solutions is about 1.6/1.5. A matching layer allows the transition from the PZT material to the transmitting medium more efficient. In some embodiments, the epoxy used to form SCC  301  may be used as an impedance matching layer. The impedance of epoxy is about 3, while impedance of SCC  301  depends on distribution of PZT ceramic pillars within the epoxy matrix. In some embodiments the acoustic impedance of SCC  301  may be approximately 10. Adding a matching layer and/or a backing material to transducer  150  increases the bandwidth. The shape of pillars  320  may lightly impact the device bandwidth and center frequency of operation. Acoustic loss of the epoxy matrix affects the bandwidth of transducer  150 . The epoxy serves to absorb or dissipate sound. Any energy that attempts to stay in the plastic will be absorbed quickly. An added advantage of SCC  301  is the higher electric field density in pillars  320  relative to epoxy matrix  330  due to the higher dielectric constant of the PZT ceramic relative to the epoxy. This increases the coupling efficiency of the transducer. 
     A piezoelectric material typically has a 20:1 acoustic impedance mismatch with blood and saline. A composite material increases the proportion of epoxy and polymers in transducer  150 , reducing acoustic impedance and providing better impedance matching. Bandwidth may be improved by including a backing material overlaid on transducer  150  to absorb acoustic energy, increasing bandwidth at the cost of somewhat reduced signal strength. 
     SCC  301  provides high efficiency and broad bandwidth for ultrasound generation and sensing, which is desirable in medical applications. According to some embodiments, single crystal piezoelectric materials used in SCC  301  have a high electromechanical coupling coefficient. The electromechanical coupling coefficient of single crystal piezoelectric materials is typically higher than PZT ceramic. Thus, the voltage levels needed for a predetermined volume change is lower for the single crystal materials used in SCC  301 , relative to that of piezo-electric ceramics. This increases the power conversion efficiency of SCC  301  from radiofrequency energy into sound, and from sound into radiofrequency energy. Some embodiments include narrow pillars  320  that remove the lateral constraint on the piezoelectric material that is present in a continuous slab of material. The lateral constraint of a bulk crystal is related to the rigidity of the material, as the pillars embedded in epoxy stretch longitudinally, there is less resistance from the surrounding epoxy material since the epoxy material is less rigid. In such embodiments, low frequency lateral modes (in the XZ-plane in  FIG. 3 ) in the vicinity of the desired ultrasound frequency are suppressed in narrow pillars  320  by the surrounding polymer matrix  330 . Thus, most of the RF electrical energy in SCC  301  is transferred to ‘height’ vibration modes (Y-axis in  FIG. 3 ) in pillars  320 , which couple to the ultrasound waves forming the probe beam. In some embodiments, polymer matrix  330  reduces the acoustic impedance of SCC  301  compared to that of a single crystal material. Indeed, the young modulus of polymer matrix  330  is lower than that of the single crystal  320 , or that of a piezo-electric ceramic. For example, in some embodiments SCC  301  may include 75% in volume of polymer matrix  330 . Such a composite has low acoustic impedance compared to a slab of single crystal piezoelectric or piezo-ceramic material. This low acoustic impedance is better matched to tissue acoustic impedance, therefore providing high efficiency and broad bandwidth to SCC  301 . 
     The dimensions of SCC  301  vary according to the specific application sought. For example, the target ultrasound frequency and bandwidth determine the specific dimensions of SCC  301  in some instances. In some embodiments pillars  320  are about 10 μm in diameter (Z-axis in  FIG. 3 ) with 10 μm deep kerfs (pillar height, Y-axis in  FIG. 3 ). For high frequency IVUS, it may be desirable to have even smaller structures and kerfs in SCC  301 . In some embodiments single crystal materials may be desirable in SCC  301  for high frequency applications because single crystals may be patterned using deep reactive ion etching (DRIE). DRIE techniques may be used to pattern the crystalline substrate with micron accuracy to fabricate SCC  301  materials on a wafer scale. 
     The volume fraction of polymer matrix  330  in SCC  301  may also vary according to the specific application. For example, the volume fraction of polymer matrix  330  determines the impedance of the transducer material which is beneficial to match acoustical impedance of the tissue of interest for the use of the ultrasound beam in some instances. The thickness of the composite crystal is determined by the resonance frequency desired. The thickness of SCC  301  is chosen to obtain a pre-selected center frequency of a transmitted ultrasound signal from transducer  150 . 
       FIG. 4  shows a partial perspective view of transducer housing  116 , including ultrasound transducer  150  according to some embodiments. Ultrasound transducer  150  includes SCC  301  and impedance matching layer  310 . Other details of ultrasound transducer  150  are omitted in  FIG. 4 , for clarity. It is understood that ultrasound transducer  150  in  FIG. 4  may include the same or similar elements as shown in  FIG. 3 . For example, ultrasound transducer  150  in  FIG. 4  may include back electrode  151 , and front electrode  152 . 
     In some embodiments SCC  301  is deformed into a curved shape. For example, SCC  301  is deformed into a dish-shaped structure having a symmetry axis included in a plane that also includes the Z-axis in  FIG. 3 . In some embodiments, the dish-shaped structure may be symmetric about the BD axis, which may be parallel to the Y-axis, or may be forming an angle relative to the Y-axis. This may be desirable for providing a focused ultrasound beam. For example, in some instances SCC  301  is deformed such that the upper surface of SCC  301  as viewed in  FIG. 3  becomes concave. In some implementations, the concave shape of the upper surface of SCC  301  is generally spherical. Further, in some instances the deformation of SCC  301  results in the lower surface of SCC  301  being convex. In some implementations, the convex shape of the lower surface of SCC  301  is generally spherical. In some particular implementations, the concave upper surface and the convex lower surface are both generally spherical with a common center point. 
     Transducer housing  116  is in a distal portion of catheter  102  according to an embodiment of the present disclosure. In particular,  FIG. 4  shows an expanded view of aspects of the distal portion of imaging core  110 . In this exemplary embodiment, imaging core  110  is terminated at its distal tip by housing  116 . Housing  116  may be fabricated from stainless steel or other suitable biocompatible material, and have a bullet-shaped or rounded nose, and an aperture  128  for an ultrasound beam. Thus, ultrasound beam  130  may emerge from housing  116 , through aperture  128 . In some embodiments, flexible driveshaft  132  of imaging core  110  is composed of two or more layers of counter wound stainless steel wires. Flexible driveshaft  132  is welded or otherwise secured to housing  116  such that rotation of flexible driveshaft  132  also imparts rotation to housing  116 . In the illustrated embodiment, an electrical cable  134  delivers the high-voltage transmit pulse and carries the low amplitude echo signal back to PIM  104 . with an optional shield  136  provides electrical power to SCC  301 . Electrical cable  134  extends through an inner lumen of flexible driveshaft  132  to the proximal end of imaging core  110  where it is terminated to the electrical connector portion of the rotational interface  114  (cf.  FIG. 2 ). SCC  301  is mounted onto molded tip  148 . Molded tip  148  may be formed of a polymer material such as epoxy, and serve as an acoustic backing material to absorb acoustic reverberations propagating within housing  116 . Molded tip  148  provides strain relief for electrical cable  134  at the point of soldering to electrodes  151  and  152  in some instances. In some embodiments, a flexible sheet of material is molded into bowl shaped substrate to have a concave shape. 
     According to some embodiments, molded tip  148  is formed such that an upper surface of the molded tip is concave so that when ultrasound transducer  150  is placed on the concave upper surface, the flexibility of SCC  301  allows ultrasound transducer  150  to acquire a corresponding curved shape. In some instances, a bottom surface of the ultrasound transducer  150  matches the curvature of the upper surface of the molded tip  148 . Accordingly, in some such instances the bottom surface of the ultrasound transducer  150  becomes convex and an opposing upper surface of the ultrasound transducer becomes concave (as shown in  FIGS. 4 and 5B ). The convex shape of the lower surface of the ultrasound transducer  150  may have an apex along the axis of beam direction BD such that a tangent to apex of the surface forms an angle θ with a longitudinal axis of catheter  102  (Z-direction in  FIG. 4 ). In that regard, in some embodiments the concave upper surface of the ultrasound transducer  150  is symmetrical about the axis of beam direction BD such that ultrasound beam  130  emitted from the ultrasound transducer  150  propagates along direction BD into the vessel tissue.  FIG. 4  shows a BD substantially orthogonal to the longitudinal axis of catheter  102  (0˜90°). One of ordinary skill will recognize that angle θ may have values smaller than 90° or larger than 90°, depending on the desired features for ultrasound data processing. In that regard, in some implementations, the ultrasound transducer  150  is mounted such that the ultrasound beam  130  propagates at an oblique angle with respect to the longitudinal axis of the catheter. 
     The curvature adopted by ultrasound transducer  150  according to embodiments as disclosed herein provides focusing for beam  130 . In some embodiments aperture  128  may be about 500 μm in diameter (d), and a focal length, f:3d, may be desired to obtain sufficient resolution and depth of field. Thus, the geometric focus of ultrasound beam  130  may be about 1 mm outside the sheath (1.5 mm from the aperture). For a curved transducer of this geometry, the depth of the dish should be approximately 20 μm. In some embodiments, the wavelength of a center frequency of an ultrasound signal transmitted by transducer  150  is about 40 μm in the transducer material. Accordingly, the diameter of the transducer may fit about a ten, a dozen, or a similar number of wavelengths within its surface. 
     In some embodiments an acoustic lens may be used to provide focusing to beam  130 . To achieve a lens, some embodiments may use silicone or some other polymer that reduces sound speed through the lens material relative to that of the medium. For example, ultrasound waves may travel at a 1.0 mm/μsec velocity in a silicone lens, versus 1.5 mm/μsec medium velocity. This may provide a similar focusing power (f:3d) to the 20 μm deep dish described above with a lens thickness of approximately 60 μm. Such a lens may be formed by surface tension under a microscope, to control thickness. For example a lens may be formed with a glue drop having a concavity provided by surface tension. A material may be as silicon rubber (slow material). But careful with losses. 
     The curved transducer approach as shown in  FIG. 4  facilitates mitigating reflections, reverberation, attenuation, and other diffraction effects resulting from using refractive elements in the path of ultrasound beam  130  in some embodiments. 
       FIG. 5A  shows a partial cross-section view of a transducer housing including electrical leads  134 - 1  and  134 - 2 , according to some embodiments.  FIG. 5A  results from taking a cut away view of  FIG. 4  along line AA′. Electrical leads  134 - 1  and  134 - 2  may be collectively referred to as leads  134  (cf.  FIG. 4 ). Leads  134  may be coupled to bonding pad  506  (lead  134 - 1 ) and to bonding pad  507  (lead  134 - 2 ). Bonding pads  506  and  507  may have electrical contact with either of electrodes  151  and  152  in ultrasound transducer  150 . In some instances, electric leads  134 - 1  and  134 - 2  provide a high and a low voltage signal coupled to SCC  301  through electrodes  151  and  152 . In some embodiments lead  134 - 1  is coupled to back electrode  151  and lead  134 - 2  is coupled to front electrode  152 . Further, according to some embodiments leads  134 - 1  and  134 - 2  are coupled to different portions of back electrode  151 . In such configurations, front electrode  152  may have a floating voltage having a value between the voltages provided by leads  134 - 1  and  134 - 2 . Embodiments having a floating electrode  152  may reduce the connections used inside housing  116 . In particular, embodiments having a floating electrode  152  may enable the use of a continuous index matching layer  310 . 
       FIG. 5B  shows a partial cross-section view of transducer housing  116 , including ultrasound transducer  150 , according to some embodiments.  FIG. 5B  results from taking a cut away view of  FIG. 4  along line BB′.  FIG. 5B  illustrates aperture  128  formed above ultrasound transducer  150  to allow ultrasound beam  130  to pass through, into and from the vessel tissue.  FIG. 5B  also shows window  124 , which is transparent to the ultrasound beam  130  coupling transducer  150  with the vessel tissue (cf.  FIG. 2 ). 
       FIGS. 6A, 6B, and 6C  show partial plan views of single crystal composites  601 A,  601 B, and  601 C, respectively, according to embodiments disclosed herein. Without loss of generality, SCC  601 A, SCC  601 B, and SCC  601 C in  FIGS. 6A, 6B, and 6C  are shown in a plane XZ consistent with Cartesian coordinate axes shown in  FIGS. 1-5B . One of ordinary skill in the art will recognize that an ultrasound transducer fabricated from any one of SCC  601 A,  601 B, and  601 C may have any orientation in 3D space. In particular, as has been discussed above, an ultrasound transducer formed from SCC  601 A, SCC  601 B, and SCC  601 C may have a 3D curvature forming a dish shape having a symmetry axis, BD, as shown in  FIG. 4 . SCC  601 A, SCC  601 B, and SCC  601 C (collectively referred to as SCC  601 ) include pillars  620 A,  620 B, and  620 C, respectively (collectively referred to as pillars  620 ). Pillars  620  in SCC  601  are embedded in polymer matrix  630 . In some embodiments polymer matrix  630  may be as polymer matrix  330 , described in detail with reference to  FIG. 3 , above. Also illustrated in  FIGS. 6A, 6B, and 6C  is a cutout path  650  in the XZ plane. Cutout path  650  may be formed with a laser beam on portions of SCC  601  including polymer matrix  630 . 
     One of ordinary skill will recognize that the portion of the total area of SCC  601 A,  601 B, and  601 C covered by pillars  620 A,  620 B, and  620 C may vary. In some embodiments pillars  620 A,  620 B, and  620 C may cover an area of about 25% of a surface area of SCC layer  601 A,  601 B, and  601 C, respectively. 
     As shown in  FIG. 6A , SCC  601 A includes pillars  620 A having a circular cross-section in the XZ plane. As shown in  FIG. 6B , SCC  601 B includes pillars  620 B having a square cross-section in the XZ plane. As shown in  FIG. 6C , SCC  601 C includes pillars  620 C having puzzle-piece cross-section in the XZ plane. One of ordinary skill would recognize that the particular shape of pillars in SCC  601  in the XZ plane is not limiting. Some embodiments may include pillars having cross-sections in the XZ plane with dog-bone shape, pseudo-random shape, and hexagonal shape. 
     Embodiments such as SCC  601 A,  601 B,  601 C, or similar non-traditional shapes provide improved fill efficiency in the XZ plane, improved adhesion to polymer matrix  630 , greater flexibility, and better suppression of undesired lateral modes (in the XZ plane). Furthermore, SCC  601  provides improved mechanical integrity during the wafer thinning process. Patterning the finished transducer with cutout path  650  is also a valuable benefit. In some embodiments, cutout path  650  may form a circular or elliptical transducer shape. Ultrasound transducers having circular or elliptical shapes offer good performance in terms of side-lobe levels, compared to cutout paths having rectangular or square shapes. 
     The geometrical configuration of pillars  620  shown in  FIG. 6  is not limiting to patterns  620 A,  620 B, or  620 C. One of ordinary skill will recognize that many configurations are possible. In some embodiments the aperture formed by SCC  601  may be apodized by adjusting the density of pillars  620  near the edges of the aperture (close to cutout path  650 ) to further reduce side-lobe levels. Some embodiments include pillars  620  having cross-sections with shapes obtained from Escher style tessellations of XZ-plane. In some embodiments, odd-shaped but uniform pillars  620  are used. 
       FIG. 7A  shows a partial side view of an ultrasound transducer  750  according to some embodiments disclosed herein. Embodiments of split back electrode transducer  750  include a back electrode divided into two equal halves  751 - 1  and  751 - 2 . In some embodiments, halves  751 - 1  and  751 - 2  have a D-shape where the transducer has a circular or elliptical profile. Halves  751 - 1  and  751 - 2  are electrically decoupled from one another, so that each half may be coupled to a different voltage. The front electrode is continuous over the entire front surface of transducer  750  in some instances. In ultrasound transducer  750  the electrode connections to electrical cables  734 - 1  and  734 - 2  are provided from the back side. Thus, the back electrode in ultrasound transducer  750  includes back side portion  751 - 1  connected to cable  734 - 1 , and back side portion  751 - 2  connected to cable  734 - 2 . According to some embodiments, front electrode  752  may float with no direct contact to an outside voltage source, or ground. Ultrasound transducer  750  includes SCC  701 , which may include single crystal pillars embedded in a polymer similar to SCC  301  and SCC  601  as described in detail above (cf.  FIGS. 1, 6A, 6B, and 6C ). 
     Some embodiments of ultrasound transducer  750  with a split back electrode configuration as in  FIG. 7A  include SCC  701  having two halves  701 - 1  and  701 - 2 , poled in opposite directions. For example, a first half SCC  701 - 1  coupled to electrode  751 - 1  may be poled in a first direction, and a second half SCC  701 - 2  coupled to electrode  751 - 2  may be poled in a second direction opposite to the first direction. SCC may support a split polarization without significant artifacts due to the separation between individual pillars provided by the polymer matrix. According to some embodiments, cable  734 - 1  may couple electrode  751 - 1  to a voltage supply at a first voltage. Also, cable  734 - 2  may couple electrode  751 - 2  to a voltage supply at a second voltage, higher than the first voltage. When the two back electrodes are excited with equal and opposite signals, the front electrode remains at virtual ground by symmetry, and each of transducer halves  701 - 1  and  701 - 2  receive equal and opposite electrical excitation. Electric field  761  is opposite in direction to electric field  762 . Likewise, the polarization induced in SCC  701 - 1  by electric field  761  is opposite to the polarization induced in SCC  701 - 2  by electric field  762 . Since SCC  701 - 1  and SCC  701 - 2  are poled in opposite directions, the piezo-electric effect on first half  701 - 1  is the same as the piezo-electric effect on second half  701 - 2 . Thus, an acoustic wave-front including the two halves of split electrode transducer  750  is generated. Accordingly, in some embodiments halves  701 - 1  and  701 - 2  vibrate in phase with one another, providing a full aperture beam. 
     A single crystal composite as disclosed herein is particularly well suited to the split back electrode configuration. Fringe fields at the boundary between the split electrodes  751 - 1  and  751 - 2  are mitigated by polymer matrix  330 . This ensures that poling of halves  701 - 1  and  701 - 2  provides a well-defined orientation near their border. 
     Some embodiments using ultrasound transducer  750  including a split electrode may yield a lower capacitance (higher impedance) device. Indeed, each of the two capacitors formed between electrode  751 - 1 ,  752 , and  751 - 2  has a lower capacitance than a capacitor made of the same SCC  701  material and having the same thickness, but double the area. Furthermore, in the split electrode configuration the two capacitors formed between electrodes  751 - 1 ,  752 , and  751 - 2  are connected in series, thus reducing the net capacitance of SCC  701  as compared to a configuration where back electrodes  751 - 1  and  751 - 2  form a single electrode. Thus, embodiments of SCC  701  having a split back electrode may use a higher excitation voltage to achieve the same ultrasound output as a conventional electrode. Embodiments consistent with the split electrode configuration illustrated in  FIG. 7A  provide desirable manufacture features, since front electrode  752  is floating and may not use a direct connection to a voltage source, or ground. This simplifies the configuration and manufacturing of ultrasound transducer  750  and tip housing  116 . For example, an impedance matching layer such as layer  310  (cf.  FIG. 3 ) may be formed as a continuous layer on top of front electrode  752 . 
     Split back electrode transducer  750  is desirable in embodiments including matching layer  310 . The use of a split back electrode permits matching layer  310  to be formed at the wafer level fabrication of transducer  750  without having a conductive material making contact with front electrode  752 . Thus, fabrication methods according to some embodiments may avoid cutting a hole in matching layer  310  for a front electrode contact. 
       FIG. 7B  shows a partial plan view of ultrasound transducer  750  according to some embodiments disclosed herein.  FIG. 7B  illustrates back electrodes  751 - 1  and  751 - 2 .  FIG. 7B  also illustrates molded tip  148  (cf.  FIG. 4 ). In some embodiments electrodes  751 - 1  and  751 - 2  in the distal area close to the tip of molded tip  748  may include a gold plated diamond grit. Bond pads  761 - 1  and  761 - 2  provide electrical contact to electrodes  751 - 1  and  751 - 2  from electrical cables such as cables  134 - 1  and  134 - 2  (cf.  FIG. 5A ). Such configuration ensures efficient and reliable electrical contact to SCC  701 . Bond pads  761 - 1  and  761 - 2  may be formed of any conductive material, like gold or silver. One of ordinary skill would recognize that the specific material forming bond pads  761 - 1  and  761 - 2  is not limiting and any conductive material or alloy thereof may be used, without limitation. 
     In embodiments using a gold plated diamond grit, SCC  701  is pressed and glued onto molded tip  148 . Thus, protuberances in the diamond grit poke into the electrode plating on the back of the sheet formed by SCC  701 , providing a low resistance electrical connection. Some embodiments may include anisotropic conductive adhesives to provide a reliable electrical connection to SCC  701 . For example, an insulating epoxy-like material filled with gold or silver spheres provides an anisotropic conductive adhesive in some implementations. In such embodiments the density of the conductive spheres is low enough that the material is non-conductive, but when the material is compressed into a thin film between two conductive surfaces, the spheres are squished between the conductors and they bridge the narrow gap to again form a low resistance connection along the compression direction. 
       FIG. 7C  illustrates front electrode  752 , which may be the common electrode for SCC transducer  750 . In some embodiments electrode  752  includes alignment tab  770  to orient the device properly within molded tip  148 . The SCC may include an epoxy matching layer. An acoustic impedance approximately equal to 3 is desirable. 
     According to some embodiments, SCC  701  including electrodes  752 ,  751 - 1 , and  751 - 2  is glued into molded tip  148  forming a dish-shape for providing focused beam  130  (cf.  FIG. 4 ). 
       FIGS. 8A-F  show a partial view of fabrication stages for an SCC  801 , according to some embodiments.  FIG. 8A  illustrates single crystal material  802  formed into a slab of material  801 A, patterned using photolithography and DRIE (or other suitable etching and/or material removal processes) to etch away portions  825  of material. SCC material  802  may be any single crystal, piezo-electric material. For example, some embodiments may use a single crystal including lead magnesium niobate-lead titanate (PMN-PT). Slab  801 A may be formed on a wafer, having a front surface (top of  FIG. 8A ) and a back surface (bottom of  FIG. 8B ). This leads to a slab of material  801 B having isolated pillars or ribs  820 , partially formed through the wafer, as illustrated in  FIG. 8B . In some embodiments a pattern of trenches is etched in the piezo-electric substrate using DRIE to produce vertical walls (Y-direction) and a very precise geometry (XZ plane), typically with 1 μm resolution. After etching, the trenches are filled with a polymer  830  such as epoxy or silicone, as illustrated in  FIG. 8B . 
       FIG. 8C  illustrates the forming of slab  801 C, according to some embodiments. Polymer layer  830  may be on the front side of the ultrasound transducer in slab  801 B, and material  802  may be on the back side of slab  801 B. In some embodiments polymer layer  830  may be polished, ground, or etched to a thickness such that polymer layer  830  and pillars  830  have an edge on the front side of SCC  801 . 
     Thus, slab  801 C includes pillars  820  of a piezo material, isolated from one another on the front side (top of  FIG. 8C ), contained within polymer matrix  830 . The flexibility of slab  801 C is adjustable based on the size of the trenches formed in the DRIE step and the properties of the polymer used in matrix  830 . Further, slabs  801 C may have different geometries obtained by photolithography and DRIE steps, as described above. In some embodiments, the pattern of pillars  820  may be isolated islands separated by large moats. 
       FIG. 8D  illustrates forming of slab  801 D, including a front electrode  852 . Forming slab  801 D may include forming the SCC layer into a desired thickness. To accomplish this, material  802  in the back side of slab  801 C (bottom of  FIG. 8C ) may be polished, ground, or etched to a thickness such that polymer matrix  830  and pillars  820  have an edge on the back side of SCC  801 D. When the substrate is thinned to form a composite sheet having pillars  820  embedded in polymer matrix  830 , individual transducer elements forming an aperture can be selected by tracing a desired outline and removing polymer matrix  830 . In some embodiments, tracing the desired outline of individual elements and removing the polymer may be performed using a laser. The individual transducer elements are then electroplated to form a front electrode  852  in slab  801 D in some instances. Front electrode  852  is formed by electroplating a conductive material on the top portion of slab  801 D in some implementations. In some embodiments front electrode  852  and matching layer  810  are formed while the structure is part of the single wafer. The thickness of the structure may be 50 μm, 40 μm, 30 μm, or less. In some embodiments, the epoxy layer may be ground to form an impedance matching layer having a ¼ wavelength thickness (or approximately 15 μm in epoxy). 
       FIG. 8E  illustrates the forming of a back electrode  851  in slab  801 E. Back electrode  851  and front electrode  852  may be as electrodes  151  and  152  described in detail above (cf.  FIG. 3 ). Back electrode  851  may be formed in the same way as front electrode  852  (cf.  FIG. 8D ). One of the advantages of SCC slab  801 E is that it has relatively low acoustic impedance, so it can provide a broad frequency response even without an acoustic matching layer. 
       FIG. 8F  illustrates slab SCC  801  formed by depositing an acoustic impedance matching layer  810  on top of slab  801 E. Acoustic matching layer  810  is included in some embodiments of SCC  801  to match the acoustic impedance of the vessel tissue. Thus, acoustic matching layer  810  may further broaden the frequency response of an ultrasound transducer using SCC  801 . 
     Once a slab of SCC  801  is complete as shown in  FIG. 8E  or  FIG. 8F , it may be installed in a catheter tip as an ultrasound transducer. According to some embodiments, SCC  801  is pressed into molded tip  148  (cf.  FIG. 4 ). Molded tip  148  may include a curved shape to impart a curved shape to SCC  801  and produce a focused acoustic beam  130 . Molded tip  148  may also provide backing impedance to SCC  801  and attachment of the transducer to driveshaft  132  (cf.  FIG. 4 ). 
     According to embodiments of the fabrication method illustrated in  FIGS. 8A-F , the dimensions of an ultrasound transducer may be defined at the wafer level. Thus, the dimensions of a finished ultrasound transducer may be determined during the formation of slab  801 A (e.g., photolithography step) and slab  801 B (e.g., DRIE step). Furthermore, the finished ultrasound transducer may be segmented into smaller transducers of any desired size and shape. The flexibility of DRIE allows the formation of pillars  820  of arbitrary shape, forming arbitrary patterns within polymer matrix  830 . For example, some pillar cross-sections discussed herein are more desirable than traditional square pillars. Having pillars  820  embedded in polymer matrix  830  allows the formation of a round transducer that is cut out using laser ablation. 
     By having flexibility in the layout and pattern design of an ultrasound transducer, fabrication methods for SCC layers as disclosed herein provide a focused ultrasound beam using a simple electrical coupling to the transducer. Some embodiments further include a custom electronic chip, such as a micro-electromechanical system (MEMS), to provide more sophisticated acoustic beam manipulation or modulation. 
       FIG. 9  shows a flow chart for a method  900  of forming an ultrasound transducer according to embodiments disclosed herein. Method  900  will be described below in relation to the steps and structures illustrated in  FIGS. 8A-F . Reference to the steps and structures in  FIG. 8A-F  is used for illustrative purposes only and is not limiting of the embodiments of method  900  consistent with the general concept expressed in  FIG. 9 . One of ordinary skill would recognize that obvious variations to method  900  may be provided, while maintaining the overall concept as described below. 
     Step  910  includes etching a single crystal according to a pattern formed by lithography, such as in slab  801 A (cf.  FIG. 8A ). In some embodiments, step  910  includes a DRIE procedure. Step  920  includes placing a polymer layer on the etched single crystal, to form a slab such as slab  801 B (cf.  FIG. 8B ). In some embodiments step  920  includes filling a pillar pattern resulting from the etching step  910  with polymer, which may be an epoxy. Step  930  includes forming the polymer layer to a thickness, such as in slab  801 C (cf.  FIG. 8C ). Step  930  may include lapping the surface of the wafer to removing excess epoxy, creating a planar surface and exposing the pillars. In step  940  an electrode is placed on the front side of the SCC. 
     Step  950  includes forming an SCC layer to a thickness, as in slab  801 D (cf.  FIG. 8D ). In some embodiments step  950  includes grinding the back portion of the wafer including slab  801 D to release the composite structure from the wafer. Step  960  includes placing a back electrode to form a slab such as slab  801 E (cf.  FIG. 8E ). According to some embodiments, step  960  may include similar procedures as step  940  to place front electrode  852  on slab  801 D. In some embodiments, slab  801 E is formed with a plurality of individual transducer elements, each forming an aperture. Step  960  may include cutting individual transducers from slab  801 E. The cutting process could be made using a laser to cleanly remove epoxy filler  830  surrounding isolated groups of pillars  820 . Thus, the piezoelectric material in pillars  820  may be left intact in step  960 . 
     Step  970  includes placing an impedance matching layer on one electrode. Step  970  may include grinding the matching layer to a desired thickness. 
     Step  980  includes placing the SCC material thus formed on a molded tip, such as molded tip  148 . Once the individual transducer is available, it can be pressed into a micro-molded housing that will become the tip of the flexible driveshaft in a rotational IVUS catheter. The molded housing may include a dish-shaped depression to form the desired aperture deflection. In some embodiments, step  980  is performed once the front and back electrodes are in place (steps  940  and  960 ). Step  980  may also include forming bonding pads to bridge the gap between the electrical leads inside the driveshaft (e.g., a shielded twisted pair) and the split back electrodes of the transducer. Such bonding pads may be as described in detail above in reference to bond pads  761 - 1  and  761 - 2  (cf.  FIG. 7B ). In some embodiments the fabrication process may include a “Cast-In-Can” method to form a transducer on a molded tip. In some embodiments, the transducer is pressed into the micro-molded tip subassembly. In some embodiments the transducer is placed on a molded tip such that acoustic beam  130  is formed in a plane perpendicular to the longitudinal axis of the catheter (XY plane in  FIG. 2 ). According to some embodiments, the transducer is placed on a molded tip such that acoustic beam  130  extends at an oblique angle with respect to the longitudinal axis (Z-axis) of the catheter. 
     Embodiments of the present disclosure described above are exemplary only. One skilled in the art may recognize various alternative embodiments from those specifically disclosed. Those alternative embodiments are also intended to be within the scope of this disclosure. Accordingly, it is appropriate that the appended claims be construed broadly and in a manner consistent with the present disclosure.