FIELD OF THE INVENTION

This invention relates to medical diagnostic ultrasonic imaging and, in particular, to ultrasound probes which use capacitive micromachined ultrasonic transducers (CMUTs).

BACKGROUND OF THE INVENTION

The ultrasonic transducers used for medical imaging have numerous characteristics which lead to the production of high quality diagnostic images. Among these are broad bandwidth and high sensitivity to low level acoustic signals at ultrasonic frequencies. Conventionally the piezoelectric materials which possess these characteristics and thus have been used for ultrasonic transducers have been made of PZT and PVDF materials, with PZT being the most preferred. However the ceramic PZT materials require manufacturing processes including dicing, matching layer bonding, fillers, electroplating and interconnections which are distinctly different and complex and require extensive handling, all of which can result in transducer stack unit yields which are less than desired.

Furthermore, this manufacturing complexity increases the cost of the final transducer probe. As ultrasound system mainframes have become smaller and dominated by field

programmable gate arrays (FPGAs) and software for much of the signal processing functionality, the cost of system mainframes has dropped with the size of the systems.

Ultrasound systems are now available in inexpensive portable, desktop and handheld form. As a result, the cost of the transducer probe is an ever-increasing percentage of the overall cost of the system, an increase which has been accelerated by the advent of higher element- count arrays used for 3D imaging. Accordingly it is desirable to be able to manufacture transducer arrays with improved yields and at lower cost to facilitate the need for low-cost ultrasound systems.

Recent developments have led to the prospect that medical ultrasound transducers can be manufactured by semiconductor processes. Desirably these processes should be the same ones used to produce the circuitry needed by an ultrasound probe such as a CMOS process. These developments have produced micromachined ultrasonic transducers or MUTs. The individual MUT cells can have round, rectangular, hexagonal, or other peripheral shapes. MUTs have been fabricated in two design approaches, one using a semiconductor layer with piezoelectric properties (PMUTs) and another using a diaphragm and substrate with electrode plates that exhibit a capacitive effect (CMUTs). The CMUT transducers are tiny diaphragm-like devices with electrodes that convert the sound vibration of a received ultrasound signal into a modulated capacitance. For transmission the capacitive charge applied to the electrodes is modulated to vibrate the diaphragm of the device and thereby transmit a sound wave. Since these devices are manufactured by semiconductor processes the devices generally have dimensions in the 10-200 micron range, but can range up to device diameters of 300-500 microns. Many such individual CMUTs can be connected together and operated in unison as a single transducer element. For example, four to sixteen CMUTs can be coupled together to function in unison as a single transducer element. A typical 2D transducer array currently will have 2000-10,000 piezoelectric transducer elements. When fabricated as a CMUT array, upwards of 50,000 CMUT cells will be used. Surprisingly, early results have indicated that the yields on semiconductor fab CMUT arrays of this size should be markedly improved over the yields for PZT arrays of several thousand transducer elements.

CMUTs were initially produced to operate in what is now known as an "uncollapsed" mode. Referring to FIGURE 1, a typical uncollapsed CMUT transducer cell 10 is shown in cross-section. The CMUT transducer cell 10 is fabricated along with a plurality of similar adjacent cells on a substrate 12 such as silicon. A diaphragm or membrane 14 which may be made of silicon nitride is supported above the substrate by an insulating support 16 which may be made of silicon oxide or silicon nitride. The cavity 18 between the membrane and the substrate may be air or gas-filled or wholly or partially evacuated. A conductive film or layer 20 such as gold forms an electrode on the diaphragm, and a similar film or layer 22 forms an electrode on the substrate. These two electrodes, separated by the dielectric cavity 18, form a capacitance. When an acoustic signal causes the membrane 14 to vibrate the variation in the capacitance can be detected, thereby transducing the acoustic wave into a corresponding electrical signal. Conversely, an a.c. signal applied to the electrodes 20,22 will modulate the capacitance, causing the membrane to move and thereby transmit an acoustic signal.

The CMUT is inherently a quadratic device so that the acoustic signal is normally the harmonic of the applied signal, that is, the acoustic signal will be at twice the frequency of the applied electrical signal frequency. To prevent this quadratic behavior a bias voltage is applied to the two electrodes which causes the diaphragm to be attracted to the substrate by the resulting coulombic force. This is shown schematically in FIGURE 2, where a source of DC bias voltage VB is applied to a terminal 24 and is coupled to the membrane electrode 20 by a path which poses a high impedance Z to a.c. signals such as an inductive impedance. A.c. signals are capacitively coupled to and from the membrane electrode from a signal terminal 26. The positive charge on the membrane 14 causes the membrane to distend as it is attracted to the negative charge on the substrate 12. The CMUT cell only weakly exhibits the quadratic behavior when operated continuously in this biased state. In another realization of the CMUT driving scheme the DC bias voltage coupling and the a.c. signal extraction from the CMUT cell can be arranged from two separate electrodes of the CMUT cell. It is also possible to apply V _B and extract a.c. signal from the second electrode 22, which is coupled to the substrate 12.

It has been found that the CMUT is most sensitive to received ultrasonic signal vibrations when the membrane is distended so that the two oppositely charged plates of the capacitive device are as close together as possible. A close proximity of the two plates will cause a greater coupling between acoustic and electrical signal energy by the CMUT. Thus it is desirable to increase the bias voltage V _B until the dielectric spacing 32 between the membrane 14 and substrate 12 is as small as can be maintained under operating signal conditions. In constructed embodiments this spacing has been on the order of one micron or less. If the applied bias voltage is too great, however, the membrane can contact the substrate, short-circuiting the device as the two plates of the device are stuck together by VanderWals forces. This sticking can occur when the CMUT cell is overdriven, and can vary from one device to another with the same bias voltage VB due to manufacturing tolerance variations. While permanent sticking can be reduced by embedding the device electrodes in an electrical isolation layer (e.g., silicon nitride), the nonlinearity of operation between collapsed and uncollapsed states is an inherent disadvantage when trying to operate an uncollapsed CMUT in a range of maximal sensitivity.

Ultrasonic imaging can be implemented over a wide range of frequencies. Higher frequencies are preferred, as higher frequency received signals will produce images of better resolution. However there is a countervailing consideration, which is that ultrasonic transmission is subject to depth-dependent attenuation, with higher frequencies being more greatly attenuation by passage through tissue than lower frequencies. Thus, ophthalmic ultrasound may be performed at frequencies approaching 20 MHz because the objects being imaged are only millimeters away from the probe, whereas abdominal imaging will require low frequencies of 2-3 MHz to image tissue and organs which are many centimeters deep in the body. A number of approaches have been employed to try to make the best of this situation, using high frequencies when possible for best resolution while using lower frequencies when deep penetration is required. For instance, tracking filters have been employed which filter received ultrasound signals at ever lower frequencies as signals are received from increasing depths in the body as described in US Pat. 4,016,750 (Green), thereby improving the signal-to-noise ratio of ultrasound signals from deeper depths. Images can be formed which are a blend of high harmonic frequency signals used at shallow depths for better resolution and fundamental lower frequency signals used at greater depths as described in US Pat. 6,283,919 (Roundhill et al.)

Two other imaging techniques are commonly used in ultrasound to deal with depth-dependent attenuation during the reception of echo signals over a range of depths. One is dynamic apodization, by which the receive aperture is progressively increased as echo signals are received from progressively greater depths. One benefit of dynamic apodization is that an increasing number of transducer elements are used to receive the echo signals as they become increasingly attenuated, thereby offsetting at least some of the attenuative effect. Another is that it provides an ever-increasing number of transducer elements for signal focusing, facilitating the second technique which is dynamic focusing. A larger aperture enables the received ultrasound to be focused to a smaller spot size, thereby enabling improved resolution at a time when higher frequency ultrasound is decreasing.

SUMMARY OF THE INVENTION

An aspect of the present invention relates to a CMUT transducer cell array comprising: a substrate; and a plurality of CMUT cells formed on the substrate, each comprising a first electrode coupled to the substrate; a movable membrane formed in spaced relationship to the first electrode; and a second electrode (28) coupled to the membrane; the CMUT array further comprising a plurality of groups of the CMUT cells, each group disposed at a different distance from a central region of an aperture of the array in an elevation dimension; and a source of DC bias voltages coupled to provide the respective first electrodes and second electrodes of at least some of the CMUT cells with a bias voltage, and controlled to apply different bias voltage levels to the groups of CMUT cells of the array, wherein the bias level applied to each group is set to progressively reduce with an increased distance of the group from the central region. Further, it is an object of the present invention to further provide a CMUT transducer cell with good sensitivity but which is immune to the membrane sticking problem.

It is a further object of the present invention to provide a CMUT transducer array which deals with the problem of depth-dependent higher frequency attenuation.

It is a further object of the present invention to provide a CMUT transducer array which enhances the benefits of dynamic apodization and dynamic focusing.

It is a further object of the present invention to provide a CMUT transducer array which provides improved performance over a significant depth-of-field.

There is provided the ultrasonic transducer CMUT cell array, which can operate in the "collapsed" mode. In the collapsed mode the sticking problem is avoided because the membrane is brought to be continually in contact with the floor of the cavity of the CMUT cell. Hysteresis is avoided by use of a range of operation which does not switch between uncollapsed and collapsed states and continually operates in the collapsed mode. The bias voltage conventionally needed to maintain the membrane in the collapsed mode is used to maintain the membrane in the collapsed state, but with a voltage selected to optimize the frequency band of the cells of the active aperture for the current depth of ultrasonic echo reception. In an illustrated implementation, CMUT cells in elevationally different regions of the array are controlled to exhibit receive frequency bands appropriate for their period of use in dynamic apodization and dynamic focusing.

BRIEF DESCRIPTION OF THE DRAWINGS

In the drawings:

FIGURE 1 is a cross-sectional view of a typical CMUT transducer cell.

FIGURE 2 is a schematic illustration of the electrical properties of a typical

CMUT cell.

FIGURE 3 is a cross-sectional view of a CMUT cell when operated in the uncollapsed state.

FIGURE 4 illustrates a CMUT cell constructed in accordance with the principles of the present invention and operated to be biased into a collapsed state.

FIGURE 5 illustrates a small aperture transducer element or array and its focal characteristic.

FIGURE 6 illustrates a large aperture transducer element or array and its focal characteristic. FIGURE 7 illustrates the different focal characteristics of a transducer array with differently operating elements in the elevation dimension.

FIGURE 8 illustrates a CMUT array constructed and operated in accordance with the principles of the present invention with bias voltage frequency control of elevationally different CMUT cells.

FIGURE 9 illustrates another CMUT array constructed and operated in accordance with the principles of the present invention with bias voltage frequency control of elevationally different CMUT cells.

FIGURE 10 is a block diagram of an ultrasonic imaging system suitable for use with a frequency-controlled CMUT cell array of the present invention.

DETAILED DESCRIPTION OF THE EMBODIMENTS

With reference to FIGURE 3, a schematic cross-section of a CMUT element or cell 5 is depicted. CMUT cell 5 includes a substrate layer 12, an electrode 22, a membrane layer 14, and a membrane electrode ring 28. In this example, the CMUT cell has a circular shape and electrode 22 is circularly configured and embedded in the substrate layer 12. In addition, the membrane layer 14 is fixed relative to the top face of the substrate layer 12 and configured/dimensioned so as to define a spherical or cylindrical cavity 18 between the membrane layer 14 and the substrate layer 12. As previously mentioned, the cell and its cavity 18 may define alternative geometries. For example, cavity 18 could define a rectangular and/or square cross-section, a hexagonal cross-section, an elliptical cross-section, or an irregular cross-section.

The bottom electrode 22 is typically insulated on its cavity-facing surface with an additional layer (not pictured). A preferred insulating layer is an oxide-nitride-oxide (ONO) dielectric layer formed above the substrate electrode and below the membrane electrode. The ONO-dielectric layer advantageously reduces charge accumulation on the electrodes which leads to device instability and drift and reduction in acoustic output pressure. The fabrication of ONO-dielectric layers on a CMUT is discussed in detail in European patent application no. 08305553.3 by Klootwijk et al., filed September 16, 2008 and entitled "Capacitive micromachined ultrasound transducer." Use of the ONO-dielectric layer is desirable with collapsed CMUTs, which are more susceptible to charge retention than are uncollapsed devices. The disclosed components may be fabricated from CMOS compatible materials, e.g., Al, Ti, nitrides (e.g., silicon nitride), oxides (various grades), tetra ethyl oxysilane (TEOS), poly-silicon and the like. In a CMOS fab, for example, the oxide and nitride layers may be formed by chemical vapor deposition and the metallization (electrode) layer put down by a sputtering process. Suitable CMOS processes are LPCVD, ALD and PECVD, the latter having a relatively low operating temperature of less than 400°C.

Exemplary techniques for producing the disclosed cavity 18 involve defining the cavity in an initial portion of the membrane layer 14 before adding a top face of the membrane layer 14. Other fabrication details may be found in US Pat. 6,328,697 (Fraser). In the exemplary embodiment depicted in FIGURE 3, the diameter of the cylindrical cavity 18 is larger than the diameter of the circularly configured electrode plate 22. Electrode ring 28 may have the same outer diameter as the circularly configured electrode plate 22, although such conformance is not required. Thus, in an exemplary embodiment of the present invention, the electrode ring 28 is fixed relative to the top face of the membrane layer 14 so as to align with the electrode plate 22 below.

FIGURE 4 shows the CMUT cell of FIGURE 3 when biased to a collapsed state, in which the membrane 14 is in contact with the floor of the cavity 18. This is accomplished by applying a DC bias voltage to the two electrodes as indicated by voltage V _B applied to the electrode ring 28 and a reference potential (ground) applied to the substrate electrode 22. While the electrode ring 28 could also be formed as a continuous disk without the hole in the center, FIGURE 4 illustrates why this is not necessary. When the membrane 14 is biased to its collapsed state as shown in this drawing, the center of the membrane is in contact with the floor of the cavity 18. As such, the center of the membrane 14 does not move during operation of the CMUT. Rather, it is the peripheral area of the membrane 14 which moves, that which is above the remaining open void of the cavity 18 and below the ring electrode. By forming the membrane electrode 28 as a ring, the charge of the upper plate of the capacitance of the device is located above the area of the CMUT which exhibits the motion and capacitive variation when the CMUT is operating as a transducer. Thus, the coupling coefficient of the CMUT transducer element is improved.

The membrane 14 may be brought to its collapsed state in contact with the floor of the cavity 18 as indicated at 36 by applying the necessary bias voltage, which is typically in the range of 50-100 volts. As the voltage is increased, the capacitance of the CMUT cell is monitored with a capacitance meter. A sudden change in the capacitance indicates that the membrane has collapsed to the floor of the cavity. The membrane can be biased downward until it just touches the floor of the cavity as indicated at 36, or can be biased further downward to increased collapse beyond that of minimal contact. Another way to bring the membrane 14 to its collapsed state is to apply pressure to the top of the membrane. When the cavity is formed in a partial or complete vacuum, it has been found that the application of atmospheric pressure of 1 Bar is sufficient to precollapse the membrane 14 to contact with the floor of the cavity 18. It is also possible to use a combination of pressure differential and bias voltage to controllably precollapse the membrane 14, which is effective with smaller devices that may have a high atmospheric collapse pressure (e.g., 10 Bar.) Once the membrane has been collapsed, it can be maintained in that state during operation by the bias voltage VB or by physical means such as by forming a retention member such as a lens cast on top of the collapsed membrane which retains the membrane in its collapsed state.

One characteristic of a transducer that impacts image quality is ultrasound beam-width. Ideally the ultrasound beam is narrow, allowing good resolution of small targets. Beam-width and resolution are a function of the aperture size, which affects the focal spot size b and hence the resolution. The beam and focal characteristics of a small aperture transducer 50 is illustrated in FIGURE 5. The beam profile 52 is seen to converge on a focus spot size b at a distance D from the transducer or array of aperture size W. In general, the focal spot size b is related to the focal distance D, the aperture size W, the frequency of the ultrasonic wave f with a wavelength λ in accordance with:

b ~ λϋ/W = F

Larger apertures have improved resolution, but at a cost of reduced depth-of-field (the range of depth D over which a narrow beam is achieved) as shown by the larger element or array 54 with its narrower focal spot size b as depicted in FIGURE 6. Thus, conventional single element or one-dimensional transducer arrays trade off resolution to achieve good image quality over a wide range of depth.

To improve this trade off transducers and arrays can be constructed of multiple elements or element groups as illustrated in FIGURE 7. The transducer depicted in this drawing has three elements or groups of elements, 60 in the center, 62 elevationally offset or surrounding the center element or group of elements 60, and an elevationally outer element or group of elements 64. In the near field, element or group 60 is actuated, producing a near field focus b ₁ . At the mid-range the next ring or group of elements 62 is actuated, producing a mid-range focus b ₂ . In the far field the outer ring or group 64 is actuated, producing a far field focus b ₃ . This transducer can be operated in a zone focus mode in which the three elements or group of elements are actuated sequentially: transmit and receive in the near field with element or group 60, then transmit and receive in the mid-range with element or group 62, then transmit and receive in the far field with element or group 64. After signals have been acquired from all three ranges, a complete image is put together using the echoes acquired from all three ranges. This, of course, requires three transmit-receive cycles which reduces real time frame rate by one-third.

An alternative mode of operation is to transmit with one or several elements or group of elements, then dynamically vary the aperture and receive focus as echoes are received over the full depth of field. For example, elements or groups 60 and 62 can be connected together by closing switch 66 and transmitting a beam focused at the mid-range focal point b ₂ . During receive, the active aperture and focal point are progressively changed by opening switch 66 to initially receive echoes from the near field with element or group 60, with a focal point at b ₁ . After the near field echoes have been received with 60, switch 66 is closed to receive mid-range echoes with both the center and next element or group in elevation 62. The active aperture is now larger, being the combination of both 60 and 62, the latter having a mid-range focal point b ₂ . After reception of the mid-range echoes switch 68 is also closed to receive with the largest aperture, the combination of 60, 62 and 64, receiving far field echoes with the elevationally outermost element or group with a far field focal point b ₃ . A group of transducers (elements) such as 60, 62 or 64 within the array, which are activated to form a current aperture, are referred as an active aperture herein. A combination of all active apertures, which are activated to form a single ultrasound image, is referred as an aperture herein.

In a previous embodiment the focal points of array are illustrated to be located along a line orthogonal to a central region (surface area) of the array. It shall be understood to the skilled person that, if the array is arranged to perform an electronic beam steering, the focal points of the array can lie along directions which cross the surface area of the array at different to orthogonal angles. In addition a central region of the active apertures can be different from the central region the array.

But the transducer remains susceptible to the phenomenon of depth-dependent attenuation. Echoes from increasing depths contain progressively less signal and

progressively more noise at the higher frequencies. This is because the transducer as described has no frequency selectivity, no ability to shape the passband for the lower frequency echo signals received by the elevationally outer elements at increasing depths. One approach to addressing this problem is described in US Pat. 5,976,091 (Hanafy). This patent describes a transducer element in which the transducer crystal and its matching layers and lens have a variable thickness, tapering from thin at the center to thicker at the edges in the elevation dimension. The outer elevation portion of the crystal resonates at the lowest frequency, and is focused, in both transmission and reception, at the deepest part of the image where the low frequencies also provide better penetration. The central portion is thinner, resonating at higher frequencies and exhibits a shallower focus. While this approach provides a desired frequency selectivity, it is extremely difficult and costly to fabricate properly tapered transducer layers and bond them together. Furthermore, in a multi -element transducer array, specialized switches are required to switch in each element at its proper phase of the transmit-receive cycle. Accordingly, an improved approach to providing frequency selectivity is desirable.

In accordance with the principles of the present invention, the frequency response of a plurality of groups in the CMUT cell array is controlled by controllably adjusting the DC bias applied to the membranes of the cells within each group. FIGURE 8 illustrates a CMUT cell array 72 of this design, with each small circle in the drawing representing an individual CMUT cell or element. The Az bracket indicates the azimuth dimension of the array and the Ele bracket indicates the elevation dimension. Each column of the array can be operated as a single functional transducer element, with scanning performed by transmitting and receiving by each column of cells sequentially across the array. In a constructed embodiment, due to the small size of the CMUT cells, a number of columns would be connected together to operate in unison as one functional transducer element. For example, all of the CMUT cells depicted in FIGURE 8 could be connected to operate in unison as a single functional transducer element.

The CMUT array illustrated in FIGURE 8 may use many of the same components as conventional PZT or PVDF transducers, such as a lens, backing, structural support, cable, and connector. CMUTs do, however, need an additional component, a controllable high voltage power supply V _B to provide a DC bias to operate properly. The requirements for this supply are quite modest since the CMUTs draw very low current. In accordance with the principles of the present invention, multiple bias voltage levels are provided to control the frequency selectivity of the various groups of cells. In the example of FIGURE 8, the bias voltage V _B is applied directly to a first group of more darkly shaded central group of CMUT cells indicated at 74. The full bias voltage V _B pulls the central region of the membrane of each CMUT cell down to create a larger contact area of the membrane 14 with the center of the cell (see FIGURE 4), which causes the non-contact region of the membrane to be smaller and thereby exhibit a higher frequency response; the first group of cells biased by the full bias voltage and disposed in the central region of the array will thus exhibit a greater response (sensitivity) to higher frequencies. The central region of the CMUT array is then operated at relatively higher frequencies for better performance when scanning the near field of the image field. The closure of the single switch 70 causes the actuation of a second group of CMUT cells, the elevationally outer elements indicated at 76 and disposed at a periphery of the array, with a lower voltage dropped across resistor 78. The lower bias voltage applied to the second group of outer elements 76 causes these cells to have a lesser contact area at the center of the collapsed membranes and a greater non-contacting peripheral area (see FIGURE 4) and hence greater sensitivity to lower frequencies. The second group of the outer elements 76 thus exhibits better low frequency performance when scanning the far field of the image area. The lower frequency response of the lower biased outer elements 76 of the active aperture is illustrated by the passband 80, while the higher frequency response of the more highly biased central elements 74 of the active aperture is illustrated by the passband 82. When the aperture is operated as shown with the single switch 70, the opening of the switch causes the operation of the central elements alone to produce a beam profile shown at 84. When the switch 70 is closed, the operation of the full aperture of all elements produces a more deeper focused beam profile as shown at 86.

In an aspect of the invention the source of DC bias voltages is controlled to bias the membranes of the cells to a collapsed state not in all CMUT groups of the plurality. The sensitivity variation corresponding to a respective active aperture can be also achieved, when the source of DC bias voltages is controlled to bias the membranes of the cells to the collapsed state for at least one CMUT group from the plurality of groups in the array.

FIGURE 9 illustrates a CMUT array 72 with an aperture having three different groups corresponding to different regions of frequency selectivity, a central group of elements 74 having the highest frequency selectivity (sensitivity), an elevationally outer group 94 having a lower frequency selectivity and actuated by the closure of switch 90, and an elevationally outermost group 96 having the lowest frequency selectivity and actuated by the closure of switches 90 and 92. Thus, the array of the present example comprises three groups of the CMUT cells, wherein each group is disposed at a different distance from the center of the array and operates at a different frequency being lower the farther the distance is from the central region. The full bias voltage V _B is applied to the central group of cells 74, a lower bias voltage dropped across resistor 78 is applied to the next outer group of cells 94, and the lowest bias voltage dropped across resistor 79 is applied to the outermost group of cells 96. When switches 90 and 92 are closed in succession during echo signal reception, the CMUT array of FIGURE 9 will exhibit the same performance as the multi-element array of FIGURE 7, but with progressively lower frequency selectivity in correspondence with the effects of depth-dependent frequency attenuation at progressively deeper regions of the image field.

FIGURE 10 illustrates in block diagram form an ultrasonic diagnostic imaging system suitable for use with a frequency-controlled CMUT array of the present invention. The CMUT array 100 is located on the tip of a catheter or ultrasound probe 100', together with an integrated circuit microbeamformer 112. The CMUT array 100 is a two-dimensional array of CMUT transducer elements capable of scanning a 2D plane or scanning in three dimensions for 3D imaging. The microbeamformer 112 controls the transmission and reception of signals by the CMUT array cells. Microbeamformers are capable of at least partial beamforming of the signals received by groups or "patches" of transducer elements as described in US Pats. 5,997,479 (Savord et al.), 6,013,032 (Savord), and 6,623,432 (Powers et al.) The microbeamformer is coupled by the catheter or probe cable to a transmit/receive (T/R) switch 116 which switches between transmission and reception and protects the main system beamformer 120 from high energy transmit signals when a microbeamformer is not used and the transducer array is operated directly by the main system beamformer. The transmission of ultrasonic beams from the CMUT transducer array 100 under control of the microbeamformer 112 is directed by a transducer controller 118 coupled to the T/R switch and the main system beamformer 120, which receives input from the user's operation of the user interface or control panel 38. The transducer controller is further adapted to define the plurality of groups in the CMUT cell array to be activated for a given transmission event. Another function controlled by the transducer controller is the direction in which beams are steered. Beams may be steered straight ahead from (orthogonal to) the transducer array, or at different angles for a wider field of view. The transducer controller 118 also controls the DC biases applied to the CMUT cells by DC bias control 104 in accordance with the present invention, which bias the cell membranes 14 to a collapsed state for operation of the CMUTs in the collapsed mode, and bias elevationally different groups of cells at different bias voltage levels for different frequency selectivity, as described previously.

The partially beamformed signals produced by the microbeamformer 112 on receive are coupled to a main beamformer 120 where partially beamformed signals from individual patches of transducer elements are combined into a fully beamformed signal. For example, the main beamformer 120 may have 128 channels, each of which receives a partially beamformed signal from a patch of dozens or hundreds of CMUT transducer cells. In this way the signals received by thousands of transducer elements of a CMUT transducer array can contribute efficiently to a single beamformed signal.

The beamformed signals are coupled to a signal processor 122. The signal processor 122 can process the received echo signals in various ways, such as bandpass filtering, decimation, I and Q component separation, and harmonic signal separation which acts to separate linear and nonlinear signals so as to enable the identification of nonlinear echo signals returned from tissue and microbubbles. The signal processor may also perform additional signal enhancement such as speckle reduction, signal compounding, and noise elimination. The bandpass filter in the signal processor can be a tracking filter, with its passband sliding from a higher frequency band to a lower frequency band as echo signals are received from increasing depths, thereby rejecting the noise at higher frequencies from greater depths where these frequencies are devoid of anatomical information.

The processed signals are coupled to a B mode processor 126 and a Doppler processor 128. The B mode processor 126 employs amplitude detection for the imaging of structures in the body such as the tissue of organs and vessels in the body. B mode images of structure of the body may be formed in either the harmonic mode or the fundamental mode or a combination of both as described in US Pat. 6,283,919 (Roundhill et al.) and US Pat.

6,458,083 (Jago et al.) The Doppler processor 128 processes temporally distinct signals from tissue movement and blood flow for the detection of the motion of substances such as the flow of blood cells in the image field. The Doppler processor typically includes a wall filter with parameters which may be set to pass and/or reject echoes returned from selected types of materials in the body. For instance, the wall filter can be set to have a passband characteristic which passes signal of relatively low amplitude from higher velocity materials while rejecting relatively strong signals from lower or zero velocity material. This passband characteristic will pass signals from flowing blood while rejecting signals from nearby stationary or slowing moving objects such as the wall of the heart. An inverse characteristic would pass signals from moving tissue of the heart while rejecting blood flow signals for what is referred to as tissue Doppler imaging, detecting and depicting the motion of tissue. The Doppler processor receives and processes a sequence of temporally discrete echo signals from different points in an image field, the sequence of echoes from a particular point referred to as an ensemble. An ensemble of echoes received in rapid succession over a relatively short interval can be used to estimate the Doppler shift frequency of flowing blood, with the correspondence of the Doppler frequency to velocity indicating the blood flow velocity. An ensemble of echoes received over a longer period of time is used to estimate the velocity of slower flowing blood or slowly moving tissue.

The structural and motion signals produced by the B mode and Doppler processors are coupled to a scan converter 132 and a multiplanar reformatter 144. The scan converter arranges the echo signals in the spatial relationship from which they were received into a desired image format. For instance, the scan converter may arrange the echo signal into a two dimensional (2D) sector-shaped format, or a pyramidal three dimensional (3D) image. The scan converter can overlay a B mode structural image with colors corresponding to motion at points in the image field corresponding with their Doppler-estimated velocities to produce a color Doppler image which depicts the motion of tissue and blood flow in the image field. The multiplanar reformatter will convert echoes which are received from points in a common plane in a volumetric region of the body into an ultrasonic image of that plane, as described in US Pat. 6,443,896 (Detmer). A volume renderer 142 converts the echo signals of a 3D data set into a projected 3D image as viewed from a given reference point as described in US Pat. 6,530,885 (Entrekin et al.) The 2D or 3D images are coupled from the scan converter 32, multiplanar reformatter 44, and volume renderer 142 to an image processor 130 for further enhancement, buffering and temporary storage for display on an image display 40. In addition to being used for imaging, the blood flow velocity values produced by the Doppler processor 128 are coupled to a flow quantification processor 134. The flow quantification processor produces measure of different flow conditions such as the volume rate of blood flow. The flow quantification processor may receive input from the user control panel 38, such as the point in the anatomy of an image where a measurement is to be made. Output data from the flow quantification processor is coupled to a graphics processor 136 for the reproduction of measurement values with the image on the display 40. The graphics processor 136 can also generate graphic overlays for display with the ultrasound images. These graphic overlays can contain standard identifying information such as patient name, date and time of the image, imaging parameters, and the like. For these purposes the graphics processor receives input from the user interface 38, such as a typed patient name. The user interface is also coupled to the transmit controller 18 to control the generation of ultrasound signals from the transducer array 100 and hence the images produced by the transducer array and the ultrasound system. The user interface is also coupled to the multiplanar reformatter 144 for selection and control of a display of multiple multiplanar reformatted (MPR) images which may be used to perform quantified measures in the image field of the MPR images. Other variations of the present invention will readily occur to those skilled in the art. For example, a high bias voltage can be applied to all of the CMUT cells of the active aperture during near field operation of the array and a lower bias voltage can be applied to all of the CMUT cells of the active aperture during far field operation of the array. This variation causes all of the CMUT cells of the active aperture of the array to exhibit high frequency selectivity during near field operation, and to exhibit lower frequency selectivity during far field operation. Another variation is to continuously apply different bias voltages to different cells of the array, and use their selection (or non-selection) in the active aperture (e.g., their use or omission from transmission and/or reception) to govern the frequency selectivity characteristics of the operation of the array.