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Author Manuscript
Lab Chip. Author manuscript; available in PMC 2008 February 23.
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Published in final edited form as:
Lab Chip. 2007 April ; 7(4): 463–468.
On-chip generation of microbubbles as a practical technology for
manufacturing contrast agents for ultrasonic imaging
Kanaka Hettiarachchia, Esra Talub, Marjorie L. Longob, Paul A. Daytonc, and Abraham P.
Leea,d
aDepartment of Biomedical Engineering, University of California at Irvine, Irvine, CA 92697, USA
bDepartment of Chemical Engineering and Materials Science, University of California at Davis, Davis, CA
95616, USA
cDepartment of Biomedical Engineering, University of California at Davis, Davis, CA 95616, USA
dDepartment of Mechanical and Aerospace Engineering, University of California at Irvine, Irvine, CA 92697,
USA. E-mail: aplee@uci.edu
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Abstract
This paper presents a new manufacturing method to generate monodisperse microbubble contrast
agents with polydispersity index (σ) values of <2% through microfluidic flow-focusing. Micronsized lipid shell-based perfluorocarbon (PFC) gas microbubbles for use as ultrasound contrast agents
were produced using this method. The poly(dimethylsiloxane) (PDMS)-based devices feature
expanding nozzle geometry with a 7 m orifice width, and are robust enough for consistent production
of microbubbles with runtimes lasting several hours. With high-speed imaging, we characterized
relationships between channel geometry, liquid flow rate Q, and gas pressure P in controlling bubble
sizes. By a simple optimization of the channel geometry and Q and P, bubbles with a mean diameter
of <5 m can be obtained, ideal for various ultrasonic imaging applications. This method
demonstrates the potential of microfluidics as an efficient means for custom-designing ultrasound
contrast agents with precise size distributions, different gas compositions and new shell materials
for stabilization, and for future targeted imaging and therapeutic applications.
Introduction
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Ultrasound contrast agents are encapsulated microbubbles with diameters of the order of 1 to
10 m. The use of contrast agents in ultrasonic imaging has been shown to improve the accuracy
of detecting functional abnormalities, and provides the potential for early detection and
characterization of disease.1,2 Owing to the density and compressibility of their gas core, these
stabilized microbubbles are substantially more echogenic than the interfaces between different
types of tissue, and therefore improve the sensitivity and specificity of 2-D and 3-D ultrasound
imaging by increasing the reflection of sound waves. They have a proven clinical utility,
particularly as a diagnostic tool in cardiology,3 radiology,4 and oncology.5
The size of a microbubble contrast agent affects its ability to cross the pulmonary
microcirculation as well as the degree of its reflectivity of ultrasound. They must be below 7
m in diameter to safely pass through the microvessels of the lungs without causing obstruction,
but the ultrasound scattering efficiency of a microbubble is a function of the sixth power of its
radius, meaning that smaller microbubbles also have poor reflectivity.6 Thus the optimum
microbubble size is between 2 and 5 m in diameter.
In-vivo contrast agents have enabled only small increases in the ultrasound signal at target sites.
7,8 Although methods to improve sensitivity are currently underway, such as delivering a
higher payload of contrast agents to the target site and adding site-specific adhesion molecules
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to the shell, little has been done towards optimization of the size of contrast agents.
Conventional methods used to produce lipid-encapsulated microbubble suspensions rely on
simple agitation to entrain a portion of the bulk gas phase into the bulk aqueous phase. The
random nature of this homogenization process results in a highly polydisperse size distribution.
9 The current leading FDA-approved ultrasound contrast agent DEFINITY® (Bristol-Myers
Squibb Medical Imaging) has microspheres that average in size from 1.1 to 3.3 m, but the
maximum bubble diameter can be as large as 20 m. Since the resonance frequency of a
microbubble depends on its diameter, an aliquot of current contrast agents has a wide range of
resonance frequencies. Limitations in bandwidth of ultrasound transducers reduce the
sensitivity of the imaging system to a small portion of the contrast agent population.
Microfluidic systems are ideal for biomedical (bio-MEMS) applications due to their small size
and batch manufacturability, providing a versatile platform for rapidly performing complex
syntheses, measurements, and analysis.10-17 The potential exists for the use of microfluidics
for highly specialized applications in large markets. For example, the ability to form micronsized bubbles or drops is important in many pharmaceutical and food processing applications,
from processes to actual products.18 Microfluidic techniques have been shown to generate
highly controlled droplet dispersions19-24 and microbubbles.24-27
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Several groups have recently utilized hydrodynamic flow-focusing methods for the massgeneration of microbubbles, but to date no group has demonstrated microbubble production in
the diameter range required for use as ultrasound contrast agents, or the feasibility of producing
lipid shell-based perfluorocarbon gas microbubbles. In this work, we demonstrate a practical
microfluidic manufacturing technique for the generation of monodisperse microbubbles with
parameters optimized for use as ultrasonic contrast agents that does not rely on the randomness
of mechanical agitation.
Background
Microbubble production methods
Conventional mechanical agitation techniques produce sufficient force to introduce gas from
the ambient environment into the liquid solution, resulting in the formation of stabilized
microbubbles. The rate of reciprocation and motion is important in determining the amount
and size of the microbubbles formed.28 A popular mechanical shaker used by hospitals to
create a leading ultrasound contrast agent has a shaking frequency of approximately 4500
oscillations per minute.
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The flows of fluids in microfluidic systems are usually characterized by low Reynolds numbers
(laminar flow), dominated by viscous stresses and pressure gradients. To create a condition
within passive microfluidic devices to generate microbubbles, flow-focusing is utilized to force
a central stream of gas and two side sheath flows of a liquid mixture through a narrow orifice
into a second chamber held at ambient pressure. The focusing effect of the surrounding flow
of liquid creates a microjet which breaks at the orifice into microbubbles.
Expanding the nozzle geometry generates monodisperse microbubbles, focusing the bubble
break-off location to one single point located at the orifice (Fig. 1). The narrowest point incurs
the highest shear force, and the subsequent nozzle expansion generates a velocity gradient in
the flow direction that allows the head of the gas thread to break continuously at the orifice,
which provides uniform control of bubble sizes.29
For liquids with moderate to small surface tension and an aqueous viscosity range, Garstecki
et al.30 summarized a scaling law relation [eqn (1)] for the diameter of the bubbles d produced
b
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by fluidic systems having the relation (Qg/Ql < 1), where Qg and Ql are the gas and liquid flow
rates, and D is the orifice diameter:
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(
)
db ∕ D ∝ Qg ∕ Ql 0.4
(1)
In microfluidic flow-focusing systems, the bubble size primarily scales with the liquid and gas
flow rates, with the continuous-phase surface tension having a negligible effect.30
Ideal contrast agents
Microbubble contrast agents with longer survival times are composed of higher molecular
weight gases and more rigid shell materials.28,31-34
Gas composition is a major factor in determining the length of time a microbubble lasts in the
circulation. The diffusivity of a gas is described by the Ostwald partition coefficient L =
Cwater/Cgas which is equal to the ratio of the amount concentrations C in the liquid and in the
gas. Nitrogen has a high Ostwald coefficient (L = 14 480 at 35 °C) and therefore a higher water
solubility compared with a PFC gas such as n-C3F8 (L = 530 at 35 °C). An adapted model by
Kabalnov et al.35 [eqn (2)] describes the rate of bubble shrinkage by the dissolution of gas in
the bloodstream, where D is the gas diffusivity in water, L is the Ostwald partition coefficient,
Patm is atmospheric pressure, P* is an excess pressure term impacted by the blood pressure
and gas metabolism, γ is the interfacial tension, r is the bubble radius, and t is time:
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(
dr ∕ dt = − DL P ∗ + 2γ
(2)
The Laplace pressure ΔP = 2γ/r is the main mechanism responsible for the disappearance of
a bubble. The use of high molecular weight gases such as perfluorocarbons reduces diffusion
out of the microbubble core and enhances bubble stability and circulation lifetime by
counterbalancing the Laplace and blood pressures.31
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A stabilized shell for the microbubble is ideally composed of multiple components serving
different functions (Fig. 2).34 An amphiphilic biocompatible phospholipid shell in the form
of a monolayer is ideal, where the water-insoluble hydrocarbon tails are in contact with the gas
core, leaving the charged phosphate head groups to interact with a polar aqueous environment.
The primary lipid shell component, a saturated diacyl phosphatidyl choline such as DPPC, is
effectively neutral and lowers the interfacial tension, adding rigidity and reducing gas escape.
A negatively charged shell component such as DPPA is added to enhance stabilization by
repulsive forces, preventing direct contact between microbubbles. The third component, an
emulsifier such as a poly(ethylene glycol) (PEG) or PEGylated material bound to the lipid
membrane of the microbubble, prevents coalescence and provides a physical barrier to various
enzymatic agents, adsorption of blood plasma proteins, and prevents phagocytosis by
macrophages of the immune system.36
Materials and methods
Design of the microfluidic channels
Critical channel widths were incorporated in the flow-focusing device [Fig. 3(a)] such as a
narrow 7 m orifice [Fig. 3(b)] and closely spaced filtering channels [Fig. 3(c)] to prevent
clogging due to PDMS debris accumulation. Reduction in the widths of the liquid and gas inlet
channels to 50 and 35 m respectively enabled the use of lower liquid flow rates and gas
delivery pressures to generate <5 m microbubbles.
The gas inlet distance from the orifice region was decreased to reduce gas diffusion. To
minimize bubble contact, the outlet reservoir immediately follows the expansion chamber. The
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punched outlet reservoir hole is of the same size as the punched inlet holes to reduce large
pressure variations that will affect bubble stability.
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Chemicals
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Microbubbles were prepared using specific lipid shell and gas core components similar to those
described for formulation of commercially manufactured ultrasound contrast agents (Table 1).
In one recipe, the lipids DPPC (1,2-dipalmitoyl-sn-glycero-3-phosphocholine, Avanti Polar
Lipids), DPPA (1,2-dipalmitoyl-sn-glycero-3-phosphate) and poly(ethylene glycol) (PEG)
lipid conjugate DPPE-PEG5000 (1,2-dipalmitoyl-sn-glycero-3-phosphoethanolamine-N[methoxypoly(ethylene glycol)-5000], Avanti Polar Lipids) were combined at an 81 : 8 : 10
ratio of molar percentages, dissolved in chloroform (CHCl3) and exposed to nitrogen under
vacuum to create a homogenous mixture. The fluorescent probe DiI-C18 (1,1′-dilinoleyl-3,3,3′,
3′-tetramethylindocarbocyanine perchlorate, Molecular Probes) was added at 1 mol% for
fluorescence microscopy studies. Water, purified using a Millipore system and mixed with
NaCl (sodium chloride, Fisher) to give a 4 mg mL−1 saline solution, was added to the vial
containing the lipid mixture, sonicated at room temperature for 20 min, and combined with a
10% aqueous glycerol/propylene glycol (GPW) mixture. In another recipe, the lipid DSPC
(1,2-distearoyl-sn-glycero-3-phosphocholine, Northern Lipids) and surfactant emulsifier
PEG-40 stearate (polyoxyethylene 40 stearate – Myrj 52, Sigma-Aldrich) were combined at a
9 : 1 molar ratio and prepared in a similar manner as described previously. Ultra pure DI water
containing 2% Tween 20 (polyoxyethylene 20 sorbitan monolaurate, Sigma-Aldrich)
surfactant was used for initial microbubble size experiments.
Microfabrication and assembling
Fabrication of the poly(dimethylsiloxane) (PDMS)-based microfluidic flow-focusing device
followed standard soft lithography techniques.10 First, a 3 inch silicon wafer was spin-coated
with a 25 m layer of a UV-curable epoxy (SU8-25, MicroChem) and exposed to UV-light
through a high resolution 20 000 dpi photomask containing the channel pattern and developed.
The wafer was used to cast a replica in PDMS (Sylgard 184, Dow Corning) consisting of a 10 :
1 prepolymer and curing agent ratio and bonded to clean soda lime glass (Corning) after oxygen
plasma treatment.
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Nitrogen (N2, Airgas) or octofluorocyclobutane (C4F8, Specialty Chemical Products) is
supplied from a pressurized tank via flexible Tygon tubing and delivered into the gas inlet of
the microfluidic chamber using a homemade micro flow meter consisting of a high-accuracy
filled pressure gauge (Cole-Parmer EW-68022-02, Cole-Parmer Instrument Company)
coupled by a three-way pressure gauge tee to a micro-metering needle valve assembly
(Upchurch P-445, Upchurch Scientific). The continuous liquid phase mixture is pumped at a
constant flow rate using a digitally controlled syringe pump (Pico Plus, Harvard Apparatus).
Analysis and imaging of the lipid microbubbles
An inverted Nikon microscope and high-speed camera (Fastcam PCI-10K, Photron Ltd.) is
used to capture still images and record movies of the microbubbles. A file viewer (PFV, Photron
Ltd.) and image analysis program (ImageJ, NIH) are used for data processing and
measurements.
For bubbles in contact with the top and bottom PDMS walls, the volume was approximated as
Vb = πdb2h/4, where db is the channel wall width and channel height h is 25 m. The
polydispersity index σ =δ/davg× 100% was calculated from the average bubble size davg and
standard deviation δ, determined by measuring the sizes of at least 100 microbubbles from
recorded images.
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For imaging of the stained lipid membranes, the microbubbles in the form of a foam were
collected using a glass pipette from the outlet reservoir. Aliquots were transferred onto glass
slides and secured with plastic cover slips to reduce fluid flow. The sample was positioned on
the stage of an upright fluorescence microscope (Nikon Eclipse E800) and illuminated by a
mercury lamp using the optical filters for illumination with green light, = 500–550 nm
(TRITC filter). Detection of fluorescence occurred at = 569 nm due to DiI labeling and
bubbles were imaged at 10× and 40× by a color CCD camera (MicroFire 2 MP, Olympus
America).
Results and discussion
Visualization
Channel geometry in addition to the liquid and gas flow rates are used for precise control of
the bubble sizes. The bubble volume Vb depends on the ratio of the gas pressure P and liquid
flow rate Q (Fig. 4). Vb increases with P for a fixed Q. An increase in Q results in a decrease
in Vb.
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An equilibrium point is reached when the liquid and gas phase form an interface of equal
pressure upstream of the orifice. At this point, no bubbles are produced since there is no pressure
drop along the longitudinal axis of the device and the tip of the gas stream does not enter the
orifice. Higher Q values increase the equilibrium point with P, and decreases the working
bubble generation range when varying P. There is a critical regime near the equilibrium point
where the system generates bubbles of different sizes.
Several device geometries at various Q and P are able to produce micron-sized bubbles.
Depending on the flow rates, it is possible to create microbubbles which are substantially
smaller in diameter than the diameter of the exit orifice.
We observed several channel-related effects on bubble size. PDMS is a gas permeable material
with a permeability to nitrogen of 245 × 10−10 cm3(STP) cm cm−2 s−1 cmHg−1, and reducing
the distance between the gas inlet and orifice results in a lower P necessary to produce the same
Vb for a fixed Q. Increasing the gas channel width Wg to match the liquid channel width Wl
decreases Vb at the same Q and P, and increases the uniformity of bubbles produced.
Production rates as high as one bubble per microsecond, or 6 × 107 bubbles per minute, have
been measured, compared with commercial production of ca. 1010 bubbles per mL during a
45 s agitation of one vial of DEFINITY® in a VialMix activation device.
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Stabilization and analysis of the lipid microbubbles
The choice of shell material greatly affects the microbubble dissolution time (Fig. 5). Upon
generation, the microbubbles quickly adjust to the downstream conditions and we observe
some degree of dissolution at the outlet reservoir when using multiple components for the
microbubble shells.
While monodisperse inside the generation chamber, purely Tween 20-coated microbubbles
exist only a few minutes after generation [Fig. 5(a)] due to rapid bubble collapse from the
Laplace pressure. A shell composed entirely of PEG-40 stearate or DPPE-PEG5000 helps
preserve microbubble monodispersity [Fig. 5(b,c)]. The polydispersity index σ of these
microbubbles was calculated to be <2% after generation, but they slowly expand by air influx
and Ostwald ripening.
There was no observed change in the size of lipid-coated microbubbles from minutes [Fig. 5
(d,e)] to hours after generation. Although highly size-stable for more than two weeks, the
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polydispersity of these lipid-coated microbubbles can be >50% when using high liquid flow
rates and gas pressures (Q > 1.0 L s−1, P > 10 psi) to increase production. Increasing Q and
P decreases the distance between exiting bubbles, and these contact interactions cause them to
coalesce due to the lower shell resistance in a high flow velocity environment as in the
expansion chamber. In addition, DPPC and DSPC lipids exist as liposomal particles in the
aqueous continuous phase, and their opening up and spreading as a monolayer at the gas–liquid
interface upstream of the orifice – a dynamic adsorption process – is effected at high rates of
flow. Using lower Q and P (Q < 1.0 L s−1, P < 5 psi) results in highly monodisperse and stable
5 m lipid-coated microbubbles, at the expense of production rate (Fig. 6). Fluorescence
microscopy confirms the existence of the coatings. The constant brightness along the
membrane wall suggests a constant membrane thickness. In contrast, the polydispersity of
DEFINITY®, the benchmark for our study, is >50% when prepared with the standard
mechanical agitation techniques.
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Studies on optimizing parameters for long-term stability (>one month) after production are
currently underway. In theory, a N2/PFC gas mixture helps to maintain bubble size by setting
an osmotic equilibrium with water-soluble gases, enabling the PFC gas to resist the Laplace
and outer fluid pressures. The issue of creating long-lasting lipid shell-based PFC gas
microbubbles in microfluidic systems poses interesting studies concerning the impact of device
geometry and scale, flow parameters, and the synergy between shell, internal components, and
the surrounding medium on the stabilization of the microbubbles.
Conclusions
In summary, we have demonstrated a microfluidic flow-focusing system for manufacturing
monodisperse microbubble contrast agents in the size range desired for ultrasonic imaging.
The microfluidic device is fully compatible with the pharmaceutical ingredients in existing
contrast agents and can serve as a low-cost alternative for current manufacturing systems.
Eliminating the size disparity issue is the most noteworthy aspect of our device. With a
monodisperse contrast agent population, we expect similar microbubble radial oscillations
from pulses of ultrasound, and thus a smaller variation in the received echoes. This we hope
translates into a better ultrasound image. Acoustic testing of these agents is outside the scope
of the current paper, but will be demonstrated in future work.
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A significant benefit of microfluidic systems is the ability to integrate multiple assays or steps
(i.e. synthesis, purification, analysis, and diagnostics) on a single device. Given the
straightforwardness of the basic flow-focusing design, this approach is easy to implement for
creating a multiplexed microbubble generator that can achieve the level of contrast agent
required for imaging.
Our successful production of lipid shell-based PFC gas microbubbles demonstrates the
potential of microfluidics as an efficient means for custom-designing contrast agents with
different gas composition and new shell materials for stabilization. A nice prospect is the
building of functionalized ‘smart’ microbubbles for targeted imaging and therapeutic
applications such as localized drug delivery.
Acknowledgements
The authors would like to thank Lisen Wang for microfabrication assistance, and funding from UC Discovery Grant
BIO-ELE04-10462 and by the National Institutes of Health through the NIH Roadmap for Medical Research, Grant
# 1 R21 EB005325-01.
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Fig. 1.
3-D rendering of the flow-focusing expansion chamber. The widths of the liquid and gas inlet
channels (Wl and Wg) are 50 and 35 m respectively. The orifice width is 7 m. Channel height
h is 25 m.
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Fig. 2.
Stabilized microbubble concept. The PFC gas core resists the combined Laplace and blood
pressure forces.
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Fig. 3.
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(a) Schematic view of the microfluidic flow-focusing device. All channels have a rectangular
cross section and a height of 25 m. The widths of the liquid and gas inlet channels are {50,
75, 100 m} and {35, 50 m} respectively. The devices feature an expanding nozzle with a
range of orifice widths {7, 10, 15, 20, 25 m}. The outlet channel connects to an open reservoir
for bubble collection. (b) Main functional area with a 7 m orifice and 3 m microbubble
generation. The arrows indicate direction of flow. (c) Magnified images of liquid inlet (left)
and gas inlet (right) filtering channels.
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Fig. 4.
Sequence of high-speed images showing the relationship between N2 gas flow rate and bubble
size upon application of several different liquid flow rates (0.5–2.0 L s−1). The widths of the
liquid and gas inlet channels are 50 and 35 m respectively, and the orifice size is 7 m.
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Fig. 5.
Relationships between microbubble shell materials and dissolution time using the same PFC
gas and flow conditions.
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Fig. 6.
Fluorescence microscopy image of ca. 5 m lipid/PFC microbubbles. Microbubbles of varying
sizes can be formed under controlled conditions.
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Table 1
Lipid microbubble compositions
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Component
Abbreviation
Shell
Shell
DPPC
DPPA
Shell
Shell
DSPC
PEG-40 stearate
Shell
DPPE-PEG5000
Shell
Gas core
Gas core
DiI-C18
N2
C4F8
Name
1,2-Dipalmitoyl-sn-glycero-3phosphocholine
1,2-Dipalmitoyl-sn-glycero-3-phosphate
1,2-Distearoyl-sn-glycero-3phosphocholine
Polyoxyethylene 40 stearate (Myrj 52)
1,2-Dipalmitoyl-sn-glycero-3phosphoethanolamine-N-[methoxypoly
(ethylene glycol)-5000]
1,1′-Dilinoleyl-3,3,3′,3′tetramethylindocarbocyanine perchlorate
Nitrogen
Octafluorocyclobutane
Molecular weight
Category
734
671
Phospholipid
Phospholipid
790
2704
Phospholipid
Surfactant emulsifier
5745
Lipopolymer emulsifier
934
28
200
Lipophilic membrane stain
Osmotic agent
Osmotic agent
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