Techniques for externally stimulating the release and diffusion of drugs from polymeric drug delivery systems enable finer tuned control of drug delivery for selective treatments. For example, the external stimulus can be triggered when the delivery system is placed close to the targeted delivery site or reaches an appropriate environment. For some time ultrasound has been investigated as a useful external stimulus for enhancing mass-transport of drugs (1). Most of the previous reports of ultrasound-stimulated drug delivery investigate an acoustic field generated by an ultrasonic horn located remotely from the drug-containing polymer (2), or indeed have considered the use of ultrasound to enhance transport across barriers such as the skin (3).
A low-frequency ultrasonic horn is a common device but with a complex operating mode. The most common usage of such ultrasonic horns is to generate 20–30 kHz continuous-wave fields into water or an aqueous solution that has undergone no special treatment (such as degassing or deionizing) under atmospheric pressure (4). For such conditions the threshold acoustic pressure to generate inertial cavitation is around 100–120 kPa (zero-to-peak amplitude). As the driving pressure exceeds those values, a relatively broad range of initial bubble sizes (from microns to tens of microns radius) will generate inertial cavitation (4). The direct sound field that is generated from the immersed tip of the ultrasonic horn has an amplitude that decreases rapidly with distance from the tip. Even in the absence of cavitation, the overall sound field in the liquid itself is a combination of this direct field and that generated by the reverberation and reflections from walls of the vessel which contain the liquid (5).
Ultrasonic horns immersed in aqueous fluids have been reported to stimulate, remotely, drug-release from polymers. For a recently published example, the application of ultrasound using an immersed transducer held remotely at a distance of 1.5 cm enhanced release of fluorescein from poly(lactic acid-co-glycolic acid), PLGA, implants that were injected into tissue-mimicking hydrogel phantoms (acrylamide, 2kPa elastic modulus) (6). However, in that example, ultrasound at 3MHz and only at an intensity of 2.2 W/cm2, but not at 0.7 W/cm2, was capable to enhance release of fluorescein with a concomitant induced degradation of the PLGA. Ultrasound of 43kHz was applied remotely at a distance of 3.7cm to cellulose hydrogels, and this was reported to enhance the release of a drug (mimosa) but with the ultrasound power from 5 to 30 W (7).
The mechanism for such ultrasound-stimulated release, particularly for transport of molecules across the skin barrier, is usually attributed to inertial cavitation since the intensities of the ultrasound at which maximum enhanced release occurred were about 14 W/cm2 for 20 kHz and 17 W/cm2 for 40 kHz (3), which are above the cavitation threshold power. Such intensities also have the potential for damage to surrounding materials and tissues, since it has been reported that the intensities at which maximum enhancement of skin permeability was induced by low-frequency ultrasound (14 W/cm2 for 20 kHz and 17 W/cm2 for 40 kHz) also caused pitting damage of aluminum foil due to the cavitation (3). Indeed, intensities of low-frequency ultrasound (23 kHz) above the cavitation threshold degraded aqueous polyvinyl alcohol (PVA) polymer (8).
This paper reports the use of low-frequency ultrasound to influence transport in porous hydrogels with a transducer attached in direct contact with the hydrogel. This is a different configuration than for ultrasound-generating devices utilized previously for enhancing transport of molecules. The advantages of the system reported in this manuscript are that
An apparatus was designed to provide a direct contact between a low-frequency ultrasound transducer and a hydrogel that contained a tracer nanoparticle (
The diffusion of the tracer gold nanoparticles was quantified by measuring the change in transmittance of the “receiver-gel” that was firmly contacted to the “tracer-gel”. The change in transmittance was measured by recording the output voltage of a photodiode (marked as 7 in
The principle of the
Therapeutic molecules were loaded into the hydrogel (“tracer-gel”) that was attached to the ultrasound transducer. The therapeutic drugs replaced the gold nanoparticles that were used for the
The principle was to prepare a dried film of lipid, rehydrate the lipid in a buffer containing the drug to be packaged and then produce homogeneous liposomes of 1μm diameter using an extrusion procedure. The result of that procedure was a 1ml volume of a concentrated solution of liposomes of 1μm diameter that were loaded with the drug. This was achieved using the following procedure. A quantity of 40μg of L-α-lecithin (Sigma, P-5638) was added to a glass tube and 400ml of chloroform was added to the glass tube in order to achieve a lipid concentration of 100mg/ml and vortex mixed to ensure the lipid was completely dissolved. The lipid/chloroform mixture was dried under a stream of N2 gas to remove the volatile chloroform. The glass tube was left under vacuum for at least 12 hours at room temperature in order to remove all traces of chloroform. Then a 1ml solution of the therapeutic drug, at the required concentration, was addd to the glass tube and vortex mixed for 5 minutes to ensure complete solution of the lipid. The tube was left to stand at room temperature for at least 15 minutes, then vortex mixed again for 5 minutes. This mixture was transferred to a syringe for extrusion through a 1μm polycarbonate filter 19 times in order to produce a homogeneous solution of liposomes. The liposome mixture was transferred to an Eppendorf tube and centrifuged at 9,000rpm for 30 minutes in order to pellet the liposomes, the supernatant was removed and the pellet resuspended in a clear 150mM NaCl (unbuffered) solution. This final solution ws incorporated into the agarose “tracer-gel”.
The rabbits used for these experiments were female New Zealand White of 2–2.5kg in weight. The rabbits were anaesthetised, prior to any procedure, using inhalational anaesthesia that comprised a mixture of 2.0 – 3.0% isoflurane with a combination of oxygen and nitrous oxide (in the ratio 2.5:1). The gas mixtures are delivered to the rabbit via a close-fitting nose-cone that was attached to a Stinger anaesthesia station. The concentration of isoflurane was varied within the range specified during the procedures in order to maintain optimal respiration for the rabbits and to maintain a surgical level of anaesthesia. All procedures were acute and the rabbits did not recover from anaesthesia. The animal care was humane and in accord with institutional guidelines procedures that were approved by the local Animal Ethics Authority.
After the depth of anaesthesia was confirmed by a lack of corneal and muscle reflexes, a lid-speculum was gently inserted to retract the eyelids of the rabbit. Where necessary due to a narrow palpebral aperture, the globe was exposed by incisions at the lateral and medial canthi, removing the nictitating membrane and reflecting the eyelids. The “tracer-gel” attached to the ultrasound transducer was placed in contact with the limbal region on the temporal side of the eye. The transducer was activated to deliver the 40 kHz ultrasound for a period of 1-2 minutes. For each eye of 3 rabbits a different therapeutic molecule was tested, either incorporated directly into the agarose “tracer-gel” or encapsulated in liposomes that were then incorporated into the agarose “tracer-gel”. Some eyes were not treated so as to be negative controls for the analyses of the eye tissues. In some eyes the therapeutic molecules were injected directly into the vitreous chamber to provide a positive control for the analyses of the eye tissues. After this period was completed for each eye, the rabbit was euthenased by CO2 asphyxiation, both eyes enucleated and immediately fixed in formalin for the subsequent fluorescent and immunohistochemical analyses.
The enucleated and fixed eyes were then embedded in paraffin and chilled on a cold plate at -5°C for 15-30 minutes. Thin sections (4 μm thickness) were cut using a new disposable knife blade and floated on a water bath at 45-50°C while the paraffin-embedded eye remained on the cold plate. The floated sections were collected on ATS (3-aminopropyltriethoxy-silane) coated slides, which were then dried in an oven at 55°C for 1 hour. The sections were deparaffinized with the following sequence of solvents; xylene (1 minute, for each of two times), absolute alcohol (1 minute, for each of four times) and 70% alcohol (1 minute, for one time).
The presence of bevacizumab or ranibizumab was detected with a published immunohistochemistry reaction (9) that used a sensitive and highly specific anti-human IgG antibody raised in donkey (Jackson ImmunoResearch Laboratories, cat# 709-165-149) to bind to either the bevacizumab or ranibizumab. The antibody was diluted in glycerol according to the manufacturer’s specifications and stored at –20°C until used. Stringent blocking was undertaken using donkey serum to minimise background binding of the antibody to VEGF receptor iso-types. The blocking solution consisted of PBS, donkey serum (1:20 dilution), BSA (0.5%), Triton X-100 (0.1%) and sodium azide (0.05%). The tissue sections were incubated in the blocking solution overnight at 4°C. The next day the antibody was further diluted 1:500 in the blocking solution, from the glycerol storage solution. The tissue sections were then incubated with the antibody overnight at 4°C. The imaging of the stained sections was conducted using either a confocal microscope or a fluorescence microscope.
An antibody-binding reaction was unnecessary to detect the presence of verteporfin in the sections of the rabbit eye since verteporfin emits light in response to excitation at around 488nm. The emission peak is around 650nm but has a broad tail. Thus, although the fluorescence microscope was optimised to detect FITC emissions (around 560nm) the intensity of the emission from verteporfin was sufficient for detection using a digital camera attached to a fluorescence microscope.
This experiment was conducted in 2 parts with gold tracer nanoparticles incorporated in the “tracer-gel”, which were
The measured transmittance in the “receiver-gel” at x=1mm below the interface is shown in
This limitation on how far the gold nanoparticle could diffuse into the gel is illustrated by the data shown in
To support those interpretations of the experiment, the “receiver-gel” was imaged by scanning electron microscopy (SEM) and the particles that were observed in the SEM images were analysed by energy dispersive spectroscopy (EDS).
Several of the particles in
A negative control experiment was conducted to confirm whether ultrasound-induced changes in transmittance could be explained by the transport of potential degradation products of the polymers. In that control experiment, the “tracer-gel” did not contain any gold nanoparticles and the transmittance in the “receiver-gel” was measured at a distance of x=1mm below the interface. The results for this control experiment are shown in
This (negative) control experiment provided supporting evidence that the diffusion of the gold nanoparticles was enhanced by the 40kHz ultrasound and that the gold nanoparticles could only diffuse 1mm into the “receiver-gel”.
The results of these
This study reports a low-frequency ultrasound system in which the transducer is in direct contact with a polymer gel that contains the molecule or nanoparticle to be delivered. The results of the
The
This report demonstrates that low-frequency ultrasound (40kHz) at a low power (0.2 W/cm2) is capable of enhancing the diffusion of gold nanoparticles in an agarose gel (of concentration 0.5% w/v). The enhanced diffusion was stimulated by an ultrasound transducer in direct contact with the agarose gel that contained the gold nanoparticles. Such a direct contact allowed for low power to be generated to stimulate the diffusion, and this power was below the threshold for cavitation. This principle of ultrasound-stimulated diffusion in an agarose gel was demonstrated in a biological application for the delivery of therapeutic molecules to the outer retina of rabbit eyes.