CMUT Arrays for Medical Imaging

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Microelectronics Journal 37 (2006) 770–777
www.elsevier.com/locate/mejo

Capacitive micromachined ultrasonic transducer (CMUT)
arrays for medical imaging
Alessandro Caronti a,*, G. Caliano a, R. Carotenuto a,c, A. Savoia a,
M. Pappalardo a, E. Cianci b, V. Foglietti b
a

Dip. di Ingegneria Elettronica, Universita` Roma Tre, Via della Vasca Navale 84, 00146 Roma, Italy
b
Istituto di Fotonica e Nanotecnologie IFN-CNR, Via Cineto Romano 42, 00156 Roma, Italy
c
Dip. I.M.E.T., Universita` degli Studi ‘Mediterranea’ di Reggio Calabria, 89069 Reggio Calabria, Italy
Received 6 August 2005; received in revised form 18 October 2005; accepted 24 October 2005
Available online 13 December 2005

Abstract
Capacitive micromachined ultrasonic transducers (CMUTs) bring the fabrication technology of standard integrated circuits into the field of
ultrasound medical imaging. This unique property, combined with the inherent advantages of CMUTs in terms of increased bandwidth and
suitability for new imaging modalities and high frequency applications, have indicated these devices as new generation arrays for acoustic
imaging. The advances in microfabrication have made possible to fabricate, in few years, silicon-based electrostatic transducers competing in
performance with the piezoelectric transducers. This paper summarizes the fabrication, design, modeling, and characterization of 1D CMUT
linear arrays for medical imaging, established in our laboratories during the past 3 years. Although the viability of our CMUT technology for
applications in diagnostic echographic imaging is demonstrated, the whole process from silicon die to final probe is not fully mature yet for
successful practical applications.
q 2005 Elsevier Ltd. All rights reserved.
Keywords: Capacitive ultrasonic transducers (CMUTs); Medical imaging; Micromachining; MEMS

1. Introduction
Since their first appearance in the mid 1990s [1], capacitive
micromachined ultrasonic transducers (CMUTs) have rapidly
emerged as an alternative to conventional piezoelectric
transducers, especially in the field of medical imaging [2–4].
The basic element of a CMUT is a capacitor cell with a fixed
electrode (backplate) and a free electrode (membrane). The
principle of operation is the well-known electrostatic transduction mechanism. If an alternating voltage is applied between
the membrane and the backplate, the modulation of the
electrostatic force results in membrane vibration with
generation of ultrasounds. Conversely, when the membrane is
subjected to an incident ultrasonic wave, the capacitance
change can be detected as a current or voltage signal. A DC
bias voltage must be used in reception for signal detection, and
it is required in transmission for linear operation. In addition,
* Corresponding author. Tel.: C39 06 55177081; fax: C39 06 5579078.
E-mail address: [email protected] (A. Caronti).

0026-2692/$ - see front matter q 2005 Elsevier Ltd. All rights reserved.
doi:10.1016/j.mejo.2005.10.012

both the transmit and receive sensitivities increase with
increasing the bias voltage.
Although the idea of generating acoustic waves by the
electrostatic attraction force between the plates of a condenser
is very old, recent advances in the microfabrication technology
have made possible to fabricate electrostatic transducers
consisting of a large number of membranes with precisely
controlled geometrical and mechanical properties. For the
operation in the megahertz range, as needed by the echographic
applications, the lateral dimensions of the membranes are on
the order of tens of microns, and the thickness is about 1–2 mm.
Thanks to the surface micromachining, the electrode separation
can be made very small, in the sub-micron range, which
enables high electric fields inside the gap, required to achieve
significant electrostatic force and transduction efficiency.
The main advantages of CMUTs compared to piezoelectrics
are the better acoustic matching to the propagation medium,
resulting in wider immersion bandwidth and improved image
resolution, the ease of fabrication, the ability to be integrated
with electronic circuits on the same wafer, and the expected
reduction of production costs. There is also a great potential for
real-time 3D imaging, through the realization of 2D CMUT

A. Caronti et al. / Microelectronics Journal 37 (2006) 770–777

771

Fig. 1. Basic steps of a CMUT fabrication process.

arrays with a large number of elements [5], harmonic imaging
applications, and high-frequency applications such as intravascular imaging (IVUS) [6].
The aim of this paper is to review the fabrication
technology, modeling, design and characterization of CMUT
arrays for medical imaging, as developed in our laboratories for
some 3 years.
2. Fabrication process of CMUT arrays
Capacitive micromachined ultrasonic transducers consist
of an array of metallized micro-membranes suspended over a
substrate. CMUTs are commonly fabricated by means of the
surface micromachining technology, using standard integrated circuits techniques. Several processes have been
reported in the literature to fabricate CMUTs, using different
materials and thin-film deposition techniques [7,8]; integrations with CMOS electronics have been also presented
[9,10].
Fig. 1 shows the basic fabrication steps of a process using
PECVD silicon nitride as a membrane structural layer,
evaporated chromium as a sacrificial layer, and sputtered
aluminum for the metallization [11]. The device is fabricated
onto a silicon wafer covered with thick thermal silicon dioxide

Fig. 2. Optical microscope image of the CMUT membranes with top electrodes
and sealed etching holes.

on both sides. After aluminum sputtering and bottom electrode
patterning, a thin-layer of silicon nitride is deposited by rfPECVD (Fig. 1(a) and (b)). A chromium layer, acting as
sacrificial layer, is evaporated and patterned into islands to
define the cavities under the membranes (Fig. 1(c)). The
excellent etching selectivity of chromium against silicon
nitride allows a good control over the cell lateral dimensions
and gap height. The first silicon nitride membrane layer is
deposited at 350 8C using silane, ammonia, and nitrogen
diluted in helium as reactant gases (Fig. 1(d)). The tensile stress
of the film is controlled by varying the silane to ammonia flow
ratio. An aluminium layer is then sputtered on top of the
membranes and patterned to define the top electrodes and
interconnections between adjacent cells (Fig. 1(e)). After
a second silicon nitride deposition, the membranes are released
by wet etching of the sacrificial layer through the etching holes
defined around the perimeter of the membranes (Fig. 1(f) and
(g)). Finally, the etchant holes are sealed by a third silicon

Fig. 3. A portion of a 64-element 1D CMUT array (top-view).

772

A. Caronti et al. / Microelectronics Journal 37 (2006) 770–777

Fig. 4. Axisymmetric model of a CMUT cell with a membrane diameter of 40 mm.

nitride deposition (Fig. 1(h)), in order to avoid water filling the
cavities during immersed operation that degrades the device
performance and eventually leads to a failure.

an effort was made to build custom simplified models, which
are nevertheless able to accurately simulate the various
aspects of CMUT operation.
4. Static operation: FEM modeling of the single cell

The overall performance of the CMUT array in immersion,
in terms of center frequency, bandwidth and sensitivity, are
strongly dependent on the design of the single cells, and their
arrangement within each element as well.
Our fabricated devices consist of circular membranes
(Fig. 2). At the operating frequencies, the membrane dimension
is typically lower than a wavelength in water. In order to design
1D CMUT linear arrays for medical imaging, the diameter of
the membranes and their number in each element must satisfy
the geometrical requirement imposed by the element-toelement distance (element pitch), so as to avoid grating lobes
occurrence in the angular response of the array [12].
Conventional piezolectric arrays typically use a ‘dice and
fill’ technique, where the lateral dimension of the elements is
defined by mechanical dicing and a polymer is infiltrated and
cured into the kerfs. Using this approach, the fabrication of
ultrasonic arrays operating above 20 MHz is very challenging
with piezo-composites [13]. On the other hand, the CMUT
technology is easily applicable to fine-scale arrays.
A realized design for a 7-MHz CMUT linear array is based
on 40-mm diameter membranes, where each element includes
4!288Z1152 membranes arranged with the configuration
shown in Fig. 3.
We employed finite element modeling (FEM) using the
commercial software ANSYS (ANSYS Inc., Canonsburg,
PA, USA) for the analysis, design and optimization of both
CMUT single cells and array elements. Because the
computation time of a FEM simulation can be very long,
Table 1
Mechanical and electrical properties of the materials used in FEM simulations

The electrostatic analysis allows the computation of the
membrane deformation caused by an applied bias voltage. An
axisymmetric model of the CMUT cell is shown in Fig. 4; this
includes a silicon nitride membrane with a buried aluminum
electrode, a thin silicon nitride insulation layer over the bottom
electrode, and a vacuum gap between the membrane and the
insulation layer. The bottom electrode is assumed to be
infinitesimally thick, because it does not affect the membrane
displacement and the electrostatic solution. The metal
interconnects linking adjacent cells are neglected by the
axisymmetric model.
The PLANE121 electrostatic elements, and the PLANE82
structural elements, were used to mesh the structure. The
coupled electrostatic-structural analysis was performed using
the ANSYS macro ESSOLV, which automatically iterates
between an electrostatic field solution and a structural solution
until the field and the structure are in equilibrium. The material
parameters used in the simulations are listed in Table 1.
If the CMUT is operated in conventional regime, the best
performance are obtained when the bias voltage is close to the
collapse voltage, which is the voltage at which the electrostatic
Static membrane displacement

0

–50

U [nm]

3. Design of CMUT arrays

–100

–150

Parameters
Young’s
modulus (GPa)
Poisson’s ratio
Density (kg/m3)
Dielectric
constant

SiNx
170

SiO
300

Aluminum

center
average

67.5
–200

0.25
2600
7

0.25
2220
4

0.35
2700


0

50

100

150

200

250

Bias voltage [V]
Fig. 5. Center and average membrane displacement as a function of the bias
voltage.

A. Caronti et al. / Microelectronics Journal 37 (2006) 770–777
0

Static cell capacitance

23

Average pressure [dB re Pin]

C0 [ fF ]

BW–3dB = 135%@ 8.0 MHz

–3

22

21

20

19

–6
–9
–12
–15
–18
pitch = 110% Dmem
pitch = 120% Dmem
pitch = 140% Dmem
pitch = 160% Dmem

–21
–24
–27

18

773

–30
0

50

100

150

200

0

250

2

4

6

8

10 12 14 16 18 20 22 24 26

Frequency [MHz]

Bias voltage [V]

force overcomes the elastic restoring force of the membrane,
and the membrane collapses onto the substrate [14]. The
collapse voltage was calculated as 232 V for the cell of Fig. 4,
with an uncertainty of 1V based on a non-convergence
criterion.
The effects of membrane metallization upon the static
displacement of the diaphragm, its mechanical resonance
frequency, the cell capacitance, and the collapse voltage, for
different sizes of the top electrode, are analyzed in [15]. A plot
of the membrane displacement is shown in Fig. 5 as a function
of the applied bias voltage. The cell capacitance C0 was
computed using the ANSYS macro CMATRIX, relating the
charges on the electrodes with the voltage drop. A plot of the
capacitance versus the bias voltage, before the collapse to take
place, is shown in Fig. 6.
5. Dynamic operation: FEM modeling of the CMUT array
element
The most simple model to analyze the dynamic operation of
an immersion CMUT is the unbounded transducer model,

Fig. 8. FEM simulated transmit pressure on the surface of the unbounded
CMUT with increasing distance between the membranes.

representing infinity of identical membranes, all driven in
phase. The unbounded FEM model for the CMUT element with
the membrane arrangement of Fig. 3 includes two quarters of
membranes in contact with a fluid column with rigid walls, as
shown in Fig. 7. The validity of this approach is stated by the
principle of image sources [18].
The unbounded model can be run fast to provide a first order
approximation of the actual CMUT array element response.
4.0

Average displacement [nm]

Fig. 6. Cell capacitance versus the bias voltage.

piston
unbounded CMUT
CMUT array element

3.0

2.0

1.0

0.0

0

2

4

6

8

10

12

14

16

18

20

22

24

18

20

22

24

Frequency [MHz]

Average pressure [dB re Pin]

0
–3
–6
–9
–12
piston
unbounded CMUT
CMUT array element

–15
–18

0

2

4

6

8

10

12

14

16

Frequency [MHz]
Fig. 7. FEM model of an unbounded transducer with the membrane
arrangement shown in Fig. 3. Absorbing acoustic elements are placed on top
to avoid reflections of outgoing pressure waves.

Fig. 9. Simulated average displacement (top) and surface pressure (bottom) of a
CMUT array element, and its equivalent piston.

A. Caronti et al. / Microelectronics Journal 37 (2006) 770–777

Frequency [MHz]
4

6

8

10

12 14

16 18

22

Measured

4
Hydrophone signal [mV]

20

24

0
–6
–12

2

–18
0
–24
–2

acoustic
interactions

–4

–30

substrate
ringing

FFT amplitude [dB]

2

– 36

VDC = 140V

0.0

0.5

1.0

1.5

– 42
3.0

2.5

2.0

Time [ms]

Frequency [MHz]
0.1

1

2

3

4

5

6

7

8

9

10 11 12 13
0
–10

0.0

–20
–30

–0.1

–40

FFT amplitude [dB]

Fig. 8 shows the average transmit pressure of a device with the
cell configuration of Fig. 4, for increasing values of the centerto-center distance between the membranes (membrane pitch).
As can be seen, the acoustic coupling determines the immersion
bandwidth: the smaller the pitch, the larger the acoustic coupling
and the transmit bandwidth. The dip around 22 MHz is in
between the first and second symmetrical modes of the
membranes, where the average displacement is minimum, and
can be called a mechanical anti-resonance. Unlike piezoelectric
transducers, fractional bandwidths as high as 130% can easily be
obtained with CMUTs, thus improving the axial resolution of
the reconstructed image; this also enables other modern
techniques like tissue harmonic imaging [19].
Acoustic coupling has other important effects, which cannot
be investigated by means of the simple unbounded model.
Because of the negligible mass, acoustic interactions occur
between the membranes of a CMUT with finite dimensions,
such as the long and narrow 1D array elements used in medical
imaging applications (Fig. 3). We have deeply investigated the
mechanism of acoustic interactions between the CMUT
membranes in immersion, through both finite element
modeling and experimental measurements [16–18]. We
developed a reduced 3D FEM model of one element by
using appropriate periodicity conditions, so as to represent an
infinitely long CMUT array element. All the membranes in the

Signal amplitude [V]

774

–50
–0.2
31.0

31.5

32.0

32.5

33.0

33.5

34.0

34.5

–60
35.0

Time [ms]
Fig. 11. Pulse-echo impulse response, and its Fourier transform, as measured in
oil with a voltage amplifier.

element are assumed to be driven in phase. As a result of the
pressure waves propagation, the membranes are subjected to
different acoustic loading, depending on their position in the
element, and dips occur in the average displacement and
transmit pressure, at frequencies where interaction effects are
stronger [18].
The average displacement and pressure responses of a
typical CMUT array element are shown in Fig. 9, where a
comparison is reported with the results of the unbounded model
and the behavior of a piston having the same mechanical
impedance as the CMUT element. The remarkable bandwidth
of the piston is due to the absence of both the higher order
modes and the fluid reactive effects, originating from the
membrane flexural motion.
6. Experimental results and images
We validated our models through optical displacement
measurements, hydrophone measurements, and pulse-echo
measurements. Fig. 10 displays the pressure impulse response,
and its Fourier transform, as measured by a hydrophone on the
acoustic axis of a CMUT array element in water (top); the element
was biased at 140 V and driven with a short pulse through

Normalized transmit pressure
0
FEM

–12
–18

acoustic
interactions

–24
–30

surface average
far-field on axis
2

4

6

8

10

12

FFT amplitude [dB]

–6

– 36

14

16

18

20

22

– 42
24

Frequency [MHz]
Fig. 10. Transmit pressure of a CMUT array element in water: hydrophone
measurement (top), and FEM simulation (bottom).

Fig. 12. Sixty-four-element CMUT linear array mounted on a PCB, with an
epoxy resin protecting the wire-bond connections.

A. Caronti et al. / Microelectronics Journal 37 (2006) 770–777

Fig. 13. Final assembly of the CMUT array probe with an echographic cable.

the pulser/receiver GE-Panametrics 5800. A comparison with a
FEM simulation of the transmit pressure is shown on the bottom
of the figure The dashed line represents the average surface
pressure, which is then propagated and compensated for

775

diffraction losses (solid line). The effect of water attenuation is
not included. The acoustic interaction effects are predicted by
FEM and can be observed experimentally, resulting in a moderate
degradation of the output pressure. The mechanical antiresonance of the membranes, limiting the immersion bandwidth
of the device below 20 MHz, is also found in a very good
agreement between FEM simulations and experiments. In
addition, silicon substrate ringing modes are observed at
7.5 MHz and harmonics. These modes have been previously
reported by other researchers [20], and can be eliminated through
proper high impedance backing of the CMUT; substrate ringing is
not predicted by FEM because the substrate is not included in the
model.
The most critical components of an ultrasound imaging
system are the transducer and its front-end electronics. In order
to exploit all the advantages of the CMUT technology, the
design of a low-noise wideband electronic system is an
important issue. Custom-designed electronic systems, based
on voltage (non-inverting) and trans-impedance (inverting)

Fig. 14. Echographic images obtained with the 64-element CMUT probe: (top) cyst phantom; (bottom) human carotid.

776

A. Caronti et al. / Microelectronics Journal 37 (2006) 770–777

amplifier configurations, were realized to evaluate the pulseecho performance of CMUT arrays [21]. A pulse-echo impulse
response of a CMUT array element immersed in oil, as measured
with a voltage amplifier, is shown in Fig. 11. The transducer is
provided with a backing on the rear side to eliminate substrate
ringing. The ringing visible in the echo signal is mainly caused
by the acoustic interactions between the membranes, as also
observed in the FFT spectrum around 3.5 MHz.
Our research is mainly focused on the development of CMUT
array probes for medical imaging. For this purpose, the CMUT
array was wire-bonded to a PCB (Fig. 12), the active membrane
area was covered with a thin-layer of silicone rubber for
continuative immersed operation, and the transducer was housed
in a probe containing biasing and decoupling circuits. A picture of
the final probe with the cable connecting to a commercial
echographic system (Esaote Technos) is shown in Fig. 13.
Fig. 14 displays the image of an echographic cyst phantom
immersed in a uniform parenchyma mimicking the human
body (top), and the in vivo image of a human carotid (bottom),
as obtained by the CMUT probe. The image quality is
comparable with that achieved with a commercial piezoelectric
probe having similar features, although the current sensitivity
of our CMUT probes is about 10 dB below [22,23].
7. Conclusion
In this paper, we presented a summary of the fabrication
technology, design, modeling and characterization of capacitive micromachined ultrasonic transducer (CMUT) arrays for
medical imaging, as developed in our laboratories in the past 3
years. We demonstrated, by experimental evidence, the
viability of our technology to realize multi-element probes,
which can be connected to standard commercial echographic
systems to generate images. At present, the images obtained
with state-of-the-art CMUT probes can be comparable or even
superior in quality to those of commercial PZT probes. Thanks
to the wider pulse-echo bandwidth, typically greater than
100%, CMUT images feature improved axial resolution, which
allows for small targets to be resolved. However, further
improvements in both sensitivity and resolution are needed to
fully compete with piezoelectric arrays, especially in the
applications where high depth of penetration is required. This
paper also reports on finite element modeling of CMUT arrays,
that we are using extensively to design and optimize the
performance of these devices. In particular, we have recently
reported the first comprehensive analysis of the acoustic
coupling in CMUTs, through both finite element modeling and
experiments. The results show that interaction effects between
the membranes can be notable in the operational bandwidth of
the transducer, giving rise to a degradation of the transducer
output pressure and efficiency.
The CMUT technology developed in our laboratories is not
fully consolidated yet. Our probes suffer from low sensitivity
and reliability issues. However, we are strongly confident that
significant improvements can be obtained in the near future, by
conducting research in many directions: the materials quality,
fabrication process stability, CMUT design, front-end

electronics, and probe engineering are the fields where we
have been working on to translate CMUT array demonstrators
into successful practical applications.
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