Patent Application
20250144387
This invention relates to drug delivery devices, methods and systems.
Transdermal drug delivery (TDD) is a prominent method for administering drugs through the skin that has been continuously explored for almost 50 years. This method provides a painless, less intrusive (sometimes even non-invasive) approach to administering drugs for motion sickness, smoking cessation, dementia, type 1 diabetes, and several other ailments. Currently developed TDD systems utilize MNs and/or nanoencapsulated drugs, which need time (typically 1˜6 hours) for a controlled, sustained release to fully ensue. This makes these systems unsuitable for rapid release requirements, which are a necessity for placating the effects of many ailments. A need, therefore, exits for a system that benefits from the painless delivery of TDD while providing rapid release of the drug into the user. In this context, with this invention, such a system is explored through the combination of (1) microneedles (MNs), (2) nanoencapsulated drug, and (3) micropumps.
A transdermal drug delivery (TDD) device, method and/or system has been devised, integrating a Shape Memory Alloy (SMA) activated micropump, microneedles (MNs), and nanoencapsulation for accelerated drug release, marking a significant advancement in the field. Embodiments of the invention leverage the compactness of a 3D-printed, SMA-triggered micropump, the inventors have achieved enhanced precision in drug delivery. The inventors' introduction of polymeric nanoparticles offers a revolutionary leap in drug deposition, outperforming conventional methods.
Through ex vivo skin studies, the inventors have showcased a marked elevation in the deposition efficiency of these polymeric nanoparticles versus traditional drug mediums.
The low power consumption, swift responsiveness, and compatibility with wearable integration signals a transformative potential for transdermal insulin administration.
1. Integration of nanoencapsulation for enhanced drug deposition: The integration of nanoencapsulated drugs with this rapid-delivery system allows for efficient skin absorption. Polymeric nanoparticles, which are generally used to prolong drug retention within tissues, here serve a dual role: 1. they enable deep penetration into skin layers, and 2. they provide an initial burst of drug deposition when combined with the micropump's mechanical action. This synergistic integration of nanoparticles into the micropump for both immediate and sustained release in a single system is innovative and not evident from prior art approaches that consider these functions separately.
2. Unique role of the SMA spring for triggering rather than pumping: In this system, the SMA spring is only used as a trigger mechanism rather than as a direct actuator for pumping. This reduces the complexity and energy consumption of the mechanism compared to that of a traditional SMA-based micropump where the spring could also unfavorably heat the drug (as it is wrapped around it during pumping). The SMA spring's contraction triggers a controlled release mechanism, which then allows a pre-loaded spring to drive the fluid through the hollow microneedles. This separation of roles is not evident in existing designs, where SMA is often employed as the primary actuator for fluid movement.
3. Compact design: The invention achieves this functionality within a compact and energy-efficient design that is suitable for wearable applications. The use of a small SMA-triggered mechanism allows the entire system to be miniaturized without compromising its power or speed of delivery. This level of miniaturization, combined with a high-speed release capability, is particularly advantageous for portable, patient-friendly devices.
4. Precision control of drug delivery speed: The actuation speed of the membrane is controlled by the interaction between the pre-loaded spring, the SMA-triggered release, and the damping foam. This provides a design space for precise control over the pressure applied for drug deposition.
The present invention can further be characterized as a transdermal drug delivery method where a transdermal drug delivery system is used. This system has (i) a triggering subsystem with a Shape Memory Alloy (SMA) spring attached and above an iron trigger, (ii) a drug delivery subsystem situated below the triggering subsystem, where the drug delivery subsystem has a preloaded pumping spring situated above a magnet, and as such the preloaded pumping spring is positioned in between the iron trigger and the magnet, and (iii) a microneedle subsystem with a pump chamber containing a drug (e.g. a nanoencapsulated drug), an elastic membrane lining an inside of the pump chamber, and an array of hollow microneedles at a bottom of the pump chamber. A current to the SMA spring is controlled that results in the SMA spring to heat and shrink to an original length of the SMA spring. This heating and shrinking separates the SMA spring and attached iron trigger from the magnet resulting in extension of the preloaded pumping spring. The extension of the preloaded pumping spring results in pushing the drug in the pump chamber through the elastic membrane and then through the array of hollow microneedles. In case the array of hollow microneedles is in contact with a skin portion, the drug is pushed into the skin portion via the array of hollow microneedles.
Numbering and labels in figures:
Transdermal drug delivery (TDD) systems can take many forms, but they all serve the same general function. A TDD system transports a drug from outside the body to the blood vessels inside the skin. The skin is a complex multilayered organ which two main components: the epidermis and the dermis. The epidermis, typically 75 to 150 μm in thickness for humans, serves as a layer of protection to the dermis, which is a much thicker (1 to 4 mm) layer that contains capillary blood vessels, sweat ducts, hair follicles, and pain receptors.
One of the most popular and actively researched forms of TDD systems is the microneedle drug delivery system. This is primarily due to the relative ease of manufacturing and customization of such needles with technologies that have seen a drastic improvement in precision over the past decade like additive manufacturing. Microneedle systems pierce through the epidermis layer to deliver the drug into the dermis layer containing the blood vessels without reaching the pain receptors, providing a painless drug delivery process. 3D printing has become an increasingly popular fabrication method for MN systems. Stereolithography (SLA) has been extensively utilized to fabricate a multitude of MN designs because of the versatility, rapid prototyping capability, and relatively remarkable precision made possible by the recent integration of 4K LCD screens in SLA printers.
MN systems can generally be characterized by their design configuration. Solid MNs, for example, are designed to either have a dry drug coat that is released in the skin after piercing the epidermis layer or to have the drug subsequently released in the skin after the MNs have breached the epidermis. Dissolving MNs fully dissolve into the skin, releasing the drug in the process. Hollow MNs, however, allow the drug to reach the dermis through the interior of the MNs. The present invention work focuses on hollow MNs because of the ease of integrating this type of MN design with a controlled or rapid release device (e.g. micropumps) for automated drug delivery.
The integration of micropumps with hollow MNs is an idea that has proven successful in recent studies. This integration provides an element of control over the system which can potentially be combined with machine learning tools to automatically estimate the patient's need for the right dosage. The challenge, however, partly lies in realizing a small enough and easy to manufacture micropump design that can be seamlessly integrated with the MN array. Piezoelectrically actuated, shape memory alloy (SMA) actuated, electromagnetically actuated, and several more actuating mechanisms have been developed over the past two decades to create small footprint micropump designs. 3D printing techniques such as SLA and fused deposition modeling (FDM) have also been utilized to fabricate complicated micropump designs with little difficulty.
MNs have also been combined with the use of nanoparticles to enhance the drug's absorption rate into the skin. Polymeric nanoparticles were found to significantly enhance the deposition into the skin leading to localized sustained release into the skin layers of the loaded drugs. This could be owed to the solid nature of these carriers compared to other types of lipidic nanosystems, which tend usually to disintegrate and diffuse through the skin layers. The use of biodegradable polymers such as PLGA (poly lactic-co-glycolic acid) enhances the biocompatibility of the particles with minimal risks of skin irritation or immune reactions.
In this invention, addressing shortcomings in the art, a novel idea is presented by combining three key components: MNs, SMA triggered micropump, and drug nanoencapsulation. This serves to provide a short-term release system that benefits from TDD's painless delivery method while maintaining a small footprint in a single integrated unit. The resulting system is mostly fabricated using FDM and SLA 3D printed components to benefit from the technologies' versatility and rapid prototyping capability. Finite element analysis methods has been employed to optimize the system's MN arrangement and actuation speed for maximum drug penetration into the skin. A modal analysis study has been then performed to simulate the system's robustness under typical operating conditions (e.g. vibrations from driving a car). The micropump's actual actuation speed was characterized via high-speed camera imaging at 9200 frames per second. The MNs and formulated nanoparticles was characterized by dynamic light scattering and scanning electron microscopy (SEM) for size, zeta potential and morphology validation. The system was finally used to perform an ex vivo skin deposition test using rat skin to evaluate the efficacy of the developed framework.
In an exemplary embodiment, the inside of the pump's pumping chamber 146 is 9.7 mm in diameter and 4 mm in height to provide a theoretical total pumping volume of ˜295.6 mm3. The pump's outer dimensions measure 18×16×15 mm3. The SMA spring 150 (Kellogg's Research Labs, USA) is a nitinol squared and ground ends compression spring with a wire diameter of 0.25 mm, mean diameter of 3 mm, 2 turns, free length of 5 mm, and solid length of 2 mm. The pumping spring 145 is also a squared and ground ends compression spring with a free length of 10 mm, mean diameter of 3 mm, solid length of 2 mm, and a spring constant of 72.5 N/m (section infra the details behind the selection of this pumping spring 145). The iron trigger 152 is a cylinder with a diameter and height of 3 mm and 4 mm, respectively. The magnet 156 utilized is an N52 grade cylindrical neodymium magnet 8 mm in diameter and 2 mm in height.
The pump's top casing 142 is fitted into the pumping chamber 146 with pins that pass through the membrane 154. The pumping spring 145 is fitted from one end through a guide in the top pump casing 142 that holds it firmly in place and the other end is fixed to the magnet 156 via an adhesive (Akfix 705). The pumping spring 145 is also fitted through the inner diameter of a cylindrical hollow damping memory foam 144 (Royal Foam, Egypt) with an outer diameter of 8 mm and height of 10 mm. The SMA spring 150 is similarly fitted in the top pump casing 142 and fixed to the iron trigger 152 with the same adhesive.
Hollow MNs 158 were designed with an approach similar to that of Yeung (C. Yeung et al., “A 3D-printed microfluidic-enabled hollow microneedle architecture for transdermal drug delivery,” Biomicrofluidics, vol. 13, no. 6, November 2019, doi: 10.1063/1.5127778).
The pump's triggering mechanism is shown in
The pumping speed is primarily determined by the stiffness k and the damping coefficient c of the pumping spring and the foam damper, respectively. Assuming that the magnetic attraction between the iron trigger and the magnet is negligible after separation compared to the force of the spring (because of the quick actuation of the SMA spring producing a large enough separation), the resulting system after triggering would be a typical spring-mass-damper system with no external force. In 1D, this can be written as
Where m is the effective mass of the spring and the magnet, c is the damping coefficient, and k is the stiffness of the system. This means that the peak time (time where the system reaches its maximum amplitude) can be found from solving Eq. 1 as
is the natural frequency and
is the damping ratio of the system. Following Hou (P. Hou, F. Zheng, C. D. Corpstein, L. Xing, and T. Li, “Multiphysics Modeling and 502 Simulation of Subcutaneous Injection and Absorption of Biotherapeutics: Sensitivity 503 Analysis,” Pharm Res, vol. 38, no. 6, pp. 1011-130 Jun. 2021, doi: 10.1007/s11095-504 021-03062-4), the actuation speed for hand driven skin injections is optimal at 0.9 mm/s and 56 kPa. A lower pressure and higher speed, however, are expected due to the nature of the actuation technique of the system, which motivates an optimization for maximum pressure and minimum actuation speed. In the design, the pressure and the actuation speed are related in 1D by
Where p is the pump's actuation pressure, x is the total deflection of the spring after actuation, and r is the radius of the pumping chamber. It can be noted that the total deflection and the geometry of the magnet (and its mass) are constrained to the dimensions outlined infra to maintain a small footprint for the micropump. This constrains the iron trigger's maximum diameter, which suggests that the stiffness of the pumping spring has an upper limit. This is because the pumping force (and in turn the pressure generated) cannot exceed the magnetic attraction force between the magnet and the iron trigger at equilibrium. To bypass the computational expense and difficulty of simulating that magnetic attraction force, the magnetic separation force was experimentally measured to be 0.8 N with a strain gauge (PASCO PS-3202, USA) for an iron trigger with diameter d=3 mm. The resulting value for the stiffness can then be used to select an appropriate spring, and the resulting peak time can be experimentally realized by trying different damping foam thicknesses and checking the peak time with a high-speed camera.
The pumping chamber 146 along with the integrated MNs 158 was printed using an SLA 3D printer. Table 1 summarizes the parameters used for printing this section. The print time for this section was 46 minutes.
The pump's top casing 142 was printed in fused deposition modeling (FDM) on a Creality CR-10 Smart. Table 2 summarizes the printing parameters for that section.
The pump's cover 140 was also printed using FDM on the same printer, but a flexible polyurethane with a shore hardness of 85 A (NinjaTek, USA) was used to allow the cover to be easily installed or removed. The self-healing inlet plug 148 was molded from Ecoflex 00-10 platinum silicone for its highest elasticity compared to other Ecoflex products. The material is also biocompatible, displays high elongation at break, and can cure relatively quickly at 60° C. in an oven. 0.2 g of each of the two components of EcoFlex 00-10 were thoroughly mixed for 1 min before being placed in a vacuum chamber for 2 min to rid the mixture of trapped air bubbles. The resulting mixture is poured into an FDM 3D printed mold and placed in an oven at 60° C. for 60 min.
The membrane's deflection speed was characterized via a high-speed imaging camera (Phantom VEO 640, Vision Labs). An open version of the pumping chamber was printed with an FDM 3D printer and attached to the top section to provide an unobstructed view of the membrane. This test section was placed on a black tabletop, and two high intensity light sources (Vision Labs, USA) were utilized to provide the camera with sufficient light for high frame rate videos. This setup can be seen in
The captured video was processed on Tracker Motion Capture software. The videos were captured at 9200 frames per second at a resolution of 640×480 pixels. The membrane's velocity and position were extrapolated from the video by setting a calibration stick and an image search for the membrane's location with a frame evolution of 10%.
In an exemplary embodiment, an ibuprofen solution was prepared by accurately weighing the drug (Ibuprofen, Sigma Aldrich, Germany) and dissolving it into phosphate buffered saline pH 7.4 to obtain a concentration of 1 mg/ml. For the polymeric nanoparticles, PLGA was dissolved in ethyl acetate and the drug was added until complete dissolution. The organic solution was then added to an aqueous 1% PVA (Polyvinyl alcohol, Sigma Aldrich, Germany), and the mixture was exposed to ultrasoncation using a probe sonicator for 4 minutes in an ice bath. The resulting emulsion was then directly transferred to a round bottom flask and fixed to a rotary evaporator for the evaporation of the organic solvent resulting in the polymeric nanoparticles suspended in the aqueous PVA solution. The particles were characterized for their particle size and zeta potential after 10 folds dilution using dynamic light scattering using Zetasizer Nano ZN (Malvern Panalytical Ltd, United Kingdom) at a fixed angle of 173° at 25° C. The samples were analyzed in triplicates. The shape and surface morphology of the PLGA nanoparticles was then examined using scanning electron microscopy (Hitachi S 2460N, Tokyo, Japan). One drop of the nanoparticles' dispersion was placed on the surface of an aluminum stub, left to dry overnight, and then sprayed with a gold coat before being examined by the scanning electron microscope.
Franz diffusion cells were used for the study of skin deposition with the receptor compartment filled with phosphate buffered saline (pH7.4) and maintained at 37° C. throughout the experiment duration. Skin was excised from the back of sacrificed Sprague Dawley rats, shaved with electric clippers and then mounted between the donor and receptor compartments of the Franz cell. The pump was filled with either the ibuprofen solution or the nanoparticles and fixed with a 3D printed fixture on top of the rat skin (see
A numerical analysis approach was utilized to simulate different numbers of MN arrays' penetration. The combination of the largest number of MNs and maximum penetration was selected for the design. The following sections outline the methodology employed for these simulations.
The MN array used was optimized for maximum penetration into the skin without wrinkling. Numerical FEM based simulations were conducted to reduce the number of design trials. Two parameters were of specific interest: the number of needles in the array and the spacing between those needles. Five configurations of needle arrays constrained by the geometry of the pump's base were applied to simulate the effect of these parameters on human skin. Table 3 shows the number of needles and spacing for each configuration simulated. The simulations were conducted using the Explicit Dynamics package in Ansys.
The geometry of each configuration has two parts as indicated in
Human skin is a composite multi-layered structure that normally behaves mechanically as a viscoelastic material. In biomechanical testing, skin acts homogenously and can be approximated as a coherent structure. As discussed earlier, the skin has two main layers: epidermis and dermis—with variations in thickness of each layer depending on factors like age, genetics, and physical condition. The designed MNs are required to penetrate the epidermis and part of the dermis layer. Accordingly, the modeled skin was chosen to be isotropic with a thickness of 1 mm.
The skin material model was approximated to be linear and highly elastic (high Poisson ratio) with failure criteria determined to be 10 percent extra strain above yield strain. An elastic modulus of 65 MPa and a Poisson ratio of 0.48 were used. PLA was selected for the MNs' material with an elastic modulus of 3.5, an ultimate tensile strength of 60 MPa, and 2.5% elongation at break.
All MN configurations successfully penetrated the required thickness with direct material cut or removal. Higher values of stress were only local to the location of the MNs' penetration with no effect on global skin structure. Configuration 5, however, showed continuous wrinkling of the skin between the subsequent MNs because of their smaller spacing. This wrinkling results in an undesirable decrease of the penetration depth of the MNs into the skin. It may also cause skin irritation which will cause discomfort to the patient. Configuration 4 was therefore selected as the optimum MNs array configuration since it provides the highest number of MNs that penetrate the skin with minimal skin wrinkling.
The results obtained for the membrane position and deflection from the high-speed camera photography (
A modal analysis was conducted to ensure that the iron-magnet attraction does not accidentally separate-thereby triggering the micropump-during operation, which can render the device dangerous to use in daily operations like driving a car, running, etc. Linear modal and harmonic vibrational analysis simulations were conducted using the Ansys Mechanical package. The analysis showed a first, second, and third modes at 548.8 Hz, 646.33 Hz, and 795.3 Hz, respectively. This is sufficiently far from the typical car vibration, running, or typical daily activities.
The nanoparticles were successfully prepared and were found to be spherical, uniform negatively charged particles. The particle size is 215+5.8 nm, polydispersity index is 0.03±0.005, and the zeta potential is −18+2.5. An SEM image of the nanoparticles can be seen in
The results obtained for the ex vivo test 5 minutes after the pump actuation are shown in
The dermis layer showed the highest percentage of deposition compared to the stratum corneum and the epidermis. This is expected because the MNs were designed to pierce the stratum corneum and the epidermis. The pointed shape of the MNs and the shape of their base explain the lower percentages deposited in the stratum corneum and the epidermis. The control MNs array without the pumping mechanism also exhibited a similar trend, with a higher, albeit small (2.3%), deposition percentage at the dermis layer in the short time frame. This demonstrates that the hollow MNs utilized are effective in delivering the drug to the dermis layer. Alone, however, these hollow MNs were not sufficient for substantial drug deposition at a short time frame.
The polymeric nanoparticles displayed a higher total deposition percentage than that exhibited by the drug solution. This is remarkable because polymeric nanoparticles, to the best of the inventors' knowledge, were not introduced to the skin through a micropumping system before, and thus this comparison between an untreated model drug and a polymeric nanoparticle version of it is unique. It is documented in the literature that the utilized polymeric nanoparticles excel in the sustained (long term) release and deposition of a drug into the skin relative to an untreated drug solution. This is anticipated because the particles' nanometer size allows them to travel more easily into the skin's inherently porous structure. It can be concluded, however, that this process can be catalyzed, at least initially, through a micropumping mechanism that provides the particles with a sufficient initial velocity and pressure to traverse through the skin layers.
A novel 3D printed, MN integrated micropump combined with polymeric nanoparticles for rapid TDD is successfully realized in this work. The micropumping mechanism utilized a low current SMA spring to trigger its actuation, and the MNs' geometry was optimized using FEA techniques. Polymeric nanoparticles were developed for a model drug and loaded into the micropump for an ex vivo rat skin deposition experiment. The system exhibited a considerable improvement in the total percentage of polymeric nanoparticles deposited in rat skin compared to a traditional drug solution in a short time frame. In just 5 minutes, the system delineated a 21.4% total deposition of the drug nanoparticles into the skin. This can be compared to only 2.3% total deposition in the same time frame for the nanoparticles without the micropumping mechanism. This ascertains the efficacy of this new concept of combining polymeric nanoparticles with a microneedle integrated micropump.
The short-term response, small footprint, demonstrated robustness, and low power consumption of the system can make it potentially very useful for insulin administration. The system can be integrated into a wearable device with a glucose sensor for automated transdermal insulin administration.
This application claims priority from U.S. Provisional Patent Application 63/596,345 filed Nov. 6, 2023, which is incorporated herein by reference.
| Number | Date | Country | |
|---|---|---|---|
| 63596345 | Nov 2023 | US |
