Download Ultrasound-triggered image-guided therapy

Investigations and research
Ultrasound-triggered image-guided
therapy
C. Moonen
UMR 5231 CNRS/University Victor Segalen, Bordeaux, France.
N. de Jong
Erasmus Medical Centre, Department of Biomedical Engineering, Rotterdam, the Netherlands.
O. Steinbach
M. Böhmer
S. Langereis
Philips Research Europe, Biomolecular Engineering, High Tech Campus, Eindhoven, the Netherlands.
The local delivery of therapeutic molecules
mediated by ultrasound is a novel approach to
addressing unmet clinical needs by providing a
minimally-invasive platform for targeted delivery
of pharmaceuticals [1]. It enables high
concentrations of drugs to be deposited under
image guidance to specific locations in the body.
It can, therefore, be a preferred treatment option
for non-systemic diseases. This produces more
constant and controlled drug concentration
profiles with favorable pharmacokinetics.
E Ultrasound mediated
delivery enables high
concentrations of drugs
to be deposited to specific
locations in the body.
With the advancement in imaging techniques,
local drug delivery can now be applied in the
human body with an unprecedented spatial and
temporal resolution. Apart from delivering drugs
to the desired location, another advantage of
using ultrasound is the increase of cellular
uptake through an effect called sonoporation,
which gives the method an additional advantage
over needle- or catheter-based procedures.
Ultrasound for therapy is preferably highly
focused, in order to expose a well-defined area,
while steering of the ultrasound beam should be
flexible and fast to allow for the controlled
exposure of a region of interest with a potentially
complex shape.
Two types of ultrasound-mediated delivery
systems can be distinguished: pressure-mediated
delivery and temperature-mediated delivery.
E Targeted delivery
produces more constant
and controlled drug
concentration profiles with
favorable pharmacokinetics.
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MEDICAMUNDI 54/1 2010
Pressure-mediated delivery is performed using
microbubbles, currently clinically applied as
ultrasound contrast agents, that react to short
ultrasound pulses by oscillation or bubble
destruction [2]. As this is a fast process and the
microbubbles and their destruction can be
conveniently imaged using ultrasound, ultrasound
imaging is the preferred modality for guiding
the drug delivery.
In the case of temperature-mediated delivery, in
which ultrasound is applied for a sufficient
amount of time to establish a local temperature
increase, magnetic resonance imaging (MRI) is
the most appropriate modality for image guidance,
as it offers the possibility of therapy planning as
well as monitoring of the local temperature
distribution during the drug delivery procedure
[3]. The first application of high intensity focused
ultrasound (HIFU) in combination with MRI
was in the ablation of uterine fibroids, which
gives an excellent starting point for temperature
monitoring during HIFU treatment.
In this overview we present recent developments
in ultrasound-triggered, image-guided drug
delivery with a focus on pressure- and
temperature-sensitive drug delivery vehicles.
For temperature-induced delivery we present
material systems, where drug and imaging agent
are both encapsulated in temperature-sensitive
carriers that allow the monitoring of
temperature-sensitive delivery vehicles using MRI
and disintegrate upon temperature increase.
For pressure-sensitive microbubble agents we
discuss the options of drug delivery with either
co-injected drugs or drugs incorporated into the
microbubble carrier, and focus on the induced
permeability increase of cells. Finally an outlook
on potential applications is given, ranging from
the improvement in small-molecule drug delivery
to gene-silencing using small interference RNA as
a new therapeutic drug format.
Temperature-triggered, MRIguided drug delivery
During ultrasound energy deposition, the actual
tissue temperature is difficult to predict.
Physiological processes may modify local heat
conduction and energy absorption. Blood flow
may increase during temperature increase and
thus change heat conduction. Unlike ablation
treatments, the exposure to an elevated
temperature for drug delivery using ultrasound
should remain well below the damage threshold.
In order to improve the therapeutic efficiency
and the safety of the intervention, real-time
mapping of temperature and thermal dose with
MRI appear to offer the best strategy to optimize
such interventions [4].
A particular advantage of MRI for guiding
thermal procedures is that MRI not only allows
for temperature mapping but can also be used
for target definition. Although temperature
effects on MR measured parameters (resonance
frequency, relaxation, diffusion) had been well
described previously, the first report of
temperature mapping by MRI appeared in 1983
[5]. This method was based on the longitudinal
relaxation time (T1). T1 based methods have
been used in the clinic and preclinical
applications, in particular at low magnetic field
[6, 7]. Since then, alternative MR temperature
imaging methods have been proposed.
The most widely used MR temperature mapping
is based on the temperature dependence of the
water proton resonance frequency (PRF) [8-11].
The excellent linearity of the temperature
dependency of the proton resonance frequency
and its near-independence with respect to tissue
type make the PRF-based method the method
of choice, in particular at mid to high field
strength. The principles of MR thermometry
have been described in previous publications
[3, 12]. As an indication, a precision of about
1°C can be reached with the PRF method at 1.5T
for a spatial resolution of 2 mm, and a temporal
resolution of 1s.
In future applications, the incorporation of MRI
contrast agents into temperature-sensitive drug
vehicles will not only offer the possibility of
guiding and controlling the temperature in the
lesion, but can also be used to visualize and
quantify the drug delivery process. However,
MRI contrast agents have an effect on the
relaxation phenomena, which are also
temperature-dependent, and therefore complicate
the quantification and reliability for monitoring
the drug release. Some compromise must be
made between sensitivity to co-release of such
contrast agents and sensitivity to temperature
change. Promising contrast agents to allow for
temperature mapping and simultaneous following
of the release are 1H CEST (Chemical Exchange
Saturation Transfer) [13] and 19F MR contrast
agents.
Temperature-sensitive liposomes
A drug delivery technology that has advanced
to clinical trials for MRI-guided, temperatureinduced drug delivery is based on temperaturesensitive liposomes. Liposomes consist of a lipid
bilayer encapsulating an aqueous interior
containing the active molecules. The
1
G
Figure 1. Temperature-sensitive liposomes containing co-encapsulated hydrophilic drugs and MRI
contrast agents that are released at the phase transition temperature (Tm) of the lipid bilayer.
2
G
Figure 2. Temperature-sensitive liposomal CEST and 19F MR contrast agents. Upon approaching
the Tm of the lipid membrane the CEST effect vanishes, while the 19F NMR signal appears.
temperature-sensitive liposomes release the
encapsulated molecules at their phase transition
temperature (Tm), where the lipid membrane
displays a transition from a gel to a liquid
crystalline phase.
Studies with doxorubicin-loaded temperaturesensitive liposomes in combination with an
externally applied regional temperature increase
clearly showed the improved efficacy of
temperature-induced drug delivery [14, 15].
Instead of relying on liposomal accumulation in
the tumor, mild hyperthermia was applied during
the first hour after injection of the temperaturesensitive liposomal formulation of doxorubicin.
As a result, doxorubicin was rapidly released in
the microvasculature of the tumor and
subsequently taken up by the tumor cells.
The co-encapsulation of hydrophilic drugs and
MRI contrast agents is shown diagramatically
in Figure 1. Temperature-sensitive liposomal
1
H CEST and 19F MR contrast agents have been
explored as a potential carrier system for MRI
guided drug delivery in combination with local
MEDICAMUNDI 54/1 2010
41
hyperthermia induced by focused ultrasound
[16] (Figure 2).
In their lumen, these temperature-sensitive
liposomes contain both a chemical shift agent
(for 1H Chemical Exchange Saturation Transfer
(CEST) detection) as well as a highly fluorinated
compound (for 19F detection). The encapsulation
of a chemical shift agent in the lumen of the
temperature-sensitive liposome provides a pool
of water protons in their aqueous core that can
be selectively saturated using a narrow radio
frequency pulse, and subsequently the saturation
is transferred to the water of the bulk phase.
E Pressure-mediated delivery
implies an extension of
the current application of
microbubbles.
E Microbubbles consist of
gas encapsulated in a
phospholipid, protein
(albumin) or polymer shell.
At temperatures below the Tm of the
temperature-sensitive liposomal contrast agent,
the chemical shift agent remains intraliposomal and the CEST effect is obtained,
whereas the intensity of the simultaneously
measured 19F NMR signal is negligible. Upon
reaching the Tm, the CEST effect disappears,
due to the release of the chemical shift agent.
Simultaneously, the 19F MRI probe is freed from
the influence of the paramagnetic shift agent,
causing an appearance of a 19F MRI signal.
The combined CEST and 19F MR temperaturesensitive liposomal carrier provides CEST MR
based contrast enhancement, to localize the
liposomes before release, while the 19F MR signal
can potentially be used to quantify the local
release of drugs with MRI.
Pressure-sensitive systems
An ultrasound device suitable for ultrasoundtriggered, ultrasound image-guided therapy
preferably consists of a dual transducer set-up
[17] comprising a focused transducer and an
imaging transducer. The focused transducer is
used to deliver a preset intensity for a precisely
defined time at a well-defined location. Mostly
the frequencies used range between 1-2 MHz,
with an approximate focal zone of a 1-2 mm in
diameter and 6-10 mm in length. The imaging
transducer is then used to obtain a wider view
around the targeted area and, for instance, to
follow the inflow of microbubbles. The
combination allows acoustic parameters such as
frequency, pressure, and pulse length to be set
independently for the therapeutic and imaging
ultrasound transducers.
Pressure-mediated delivery implies an extension
of the current application of microbubbles.
These microbubbles consist of gas encapsulated
in a phospholipid, protein (albumin) or
polymer shell.
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Microbubbles with lipid shells and albumin shells
are already used clinically. They are very
sensitive ultrasound contrast agents, as only
about 109 bubbles have to be injected, leading
to efficient imaging of wall motion abnormalities
and perfusion. However, these microbubbles
disappear quickly from the blood stream.
They are taken up by the liver and spleen and,
especially, because they may contain a fraction
of bubbles above 4-5 micron they may be trapped
in the lung capillaries. The fluorinated gas used
in these microbubbles is very hydrophobic, but
gas exchange does take place in the lungs.
Polymer-shelled microbubbles are easily prepared
with a more narrow size distribution. This avoids
significant accumulation in the lungs. Polymer
shells are more rigid, leading to different
acoustic characteristics from those of lipidshelled agents, governed by shell inhomogeneities
and higher activation thresholds in terms of the
acoustic pressure needed for destruction. If
polymers with a hydrophobic moiety are chosen,
liquid inflow into the capsules can be prevented
and gas exchange in the lungs does not occur for
these types of agents because fluorinated gases
are not necessary [18]. Because of the typical
size of microbubbles, less than 1 micron, liver
and spleen accumulation is very difficult to avoid.
The microbubbles are especially efficient in
permeabilizing the vessel walls when exposed to
ultrasound [19, 20]. This allows for delivery of
material in tissue where normally no, or only
limited, extravasation would take place.
Examples can be found in cardiology, where
local gene delivery has been established in
preclinical models both for marker genes as well
as for therapeutic genes [21]. In addition, in
tumors, an increase of the delivery of model
compounds in the tumor tissue has been
demonstrated [22].
As the increased cell permeability caused by
sonoporation remains for a limited time, having
a high local drug concentration is advantageous.
Several concepts have been developed to attach
drugs to the bubble shell, to build them into
the shell, or to put them in an additional liquid
reservoir in the inside of the microbubble [23].
For incorporation in the shell, a polymeric
microbubble with a relatively thick shell offers
the highest volume for drug incorporation.
However, drugs incorporated in the shell do not
release quickly when the bubble is destroyed, as
they tend to remain associated with the bubble
fragments that are formed [24]. For this reason,
systems in which drugs are present in an
additional liquid reservoir form a more useful
delivery vehicle, because half the volume of the
microbubble or more can be used. High CEST
F
Figure 3. Three still frames of a
Brandaris movie recorded at a frame
rate of 10 Mfs. The bubbles were lying
against the cell and there were
insonified with 10 cycles of a frequency
of 1 MHz. Left: before ultrasound,
Middle: maximum expansion.
Right: maximum compression
3
4
amounts can be incorporated, and when the
bubble breaks the release is instantaneous [25].
However, this method is restricted to
hydrophobic drugs.
A more flexible method is to attach multilayers
[26] or liposomes to the outside of microbubbles
[27-29]. In this case, both lipid and polymeric
microbubbles can be used and both hydrophilic
and hydrophobic drugs can be incorporated.
The delivery of doxorubicin, siRNA and pDNA
has been established.
Ultrafast cameras allow the microbubble
behavior to be studied on the microsecond
timescale, and their interaction with cells to be
observed in detail [30]. A camera developed at
Erasmus Medical Centre and Twente University
uses the rotating mirror principle to project
images onto 128 CCD cameras mounted over
an angle of 70˚. This camera provides recording
speeds up to 25 million frames per second
(25 Mfs) which is well above the ultrasound
frequency, so that microbubble behavior during
an ultrasound pulse can be studied in detail.
It has been found that lipid-shelled microbubbles
can show both linear and strong non-linear
behavior, which is useful for contrast-specific
ultrasound imaging [31, 32]. Polymer-shelled
microbubbles with a thin, usually inhomogeneous
shell show reversible indentations in the shell
before they actually break and release the
encapsulated gas [25]. This indentation results
in signals visible on an ultrasound scanner.
The ultrafast camera has been used to investigate
sonoporation as well. Van Wamel et al. [33]
presented results that demonstrate that
(endothelial) cell membranes follow the oscillations
of adjacent microbubbles. The amplitude of the
oscillations is on the micrometer scale and the
time scale is in microseconds, so the movements
in the cell membrane are very fast, in the order
of m/s. The displacement of the cell caused by
the expansion and compression of the bubbles
around the cell has an amplitude of 2 μm,
indicated by arrows in Figure 3.
G
Figure 4. Fluorescence images of
propidium iodide uptake of the one
cell surrounded with bubbles. The
left frame is before US exposure.
The right frame is the fluorescence
image taken 15 s after US was
turned off. The two circles indicate
the positions of the microbubbles.
This leads to an enhanced uptake of a marker
dye as shown in Figure 4 (right frame). This
shows the propidium iodide uptake recorded
15 seconds after the ultrasound has been switched
off. Only cells with bubbles in close proximity
showed uptake. After three minutes, no propidium
iodide was present in the cell, indicating that
the cell was still viable. This experiment clearly
demonstrates that it is not necessary to induce
large pores into membranes to change the
porosity, and conditions exist at which cell
membrane becomes only temporarily porous.
In recent work, the importance of changes in
calcium channels and the need for ATP in the
membrane because of ultrasound and
microbubbles has been highlighted [34, 35].
MEDICAMUNDI 54/1 2010
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E The future of ultrasoundtriggered drug delivery
will be determined by
the combination of many
factors.
Therefore, not only the opening but also the
closure of cell membranes after ultrasound
treatment is a very relevant item in ultrasoundinduced drug delivery research.
Summary and conclusions
The first application of HIFU/MRI was in the
ablation of uterine fibroids, which gives an
excellent starting point for temperature monitoring
during HIFU treatment. Further clinical trials
are running for delivery of doxorubicin using
temperature-sensitive liposomes (Thermodox™,
Celsion). New options to improve image-guided
drug delivery will need new imaging technology,
material development and an expansion of the
knowledge on mechanisms of uptake of drugs
into cells, especially for high molecular weight
therapeutics. These complex questions are
currently addressed in public-private consortia
such as the European Framework 7 project
“Sonodrugs”.
Specific attention needs to be paid to the choice
of drug and its formulation. Small-molecule drugs,
such as doxorubicin and paclitaxel, are well-
established therapies. By depositing a larger dose
at the region of interest, the efficacy of cancer
therapy can be improved. Novel therapeutic
formats, using DNA-based gene therapy or gene
silencing by therapeutic siRNAs, need specific
mechanisms to enter the cell, currently often by
the use of viral vector. With sonoporation, a
different way of entering the cell is introduced,
avoiding potential adverse reactions and complex
delivery systems such as viruses. Here
ultrasound-mediated delivery becomes a true
enabling application
For chemotherapy, dose reduction and thereby
the increase of the therapeutic window is the
differentiating advantage. While for smallmolecule drugs entry to the cells can be achieved
passively, nucleic acids require an uptake
mechanism. Here, ultrasound can also be used
as the trigger.
The specific combination of ultrasound, delivery
and contrast vehicles, imaging technology and
the medical application will finally determine
the future development of image-guided,
ultrasound-triggered drug delivery K
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