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WO2024197304A2 - Système d'endoprothèse à fonction avancée pour l'administration chronique de médicament à la demande, la commande sonogénétique de cellules et la modulation acoustique de la fonction cellulaire - Google Patents

Système d'endoprothèse à fonction avancée pour l'administration chronique de médicament à la demande, la commande sonogénétique de cellules et la modulation acoustique de la fonction cellulaire Download PDF

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Publication number
WO2024197304A2
WO2024197304A2 PCT/US2024/021334 US2024021334W WO2024197304A2 WO 2024197304 A2 WO2024197304 A2 WO 2024197304A2 US 2024021334 W US2024021334 W US 2024021334W WO 2024197304 A2 WO2024197304 A2 WO 2024197304A2
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WO
WIPO (PCT)
Prior art keywords
transducer
microfluidic
support structure
stent
ultrasound
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Pending
Application number
PCT/US2024/021334
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English (en)
Other versions
WO2024197304A3 (fr
Inventor
Jose Morales
Haotian Lu
Huachen CUI
Xiaoyu Zheng
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)
University of California Berkeley
University of California San Diego UCSD
Original Assignee
University of California Berkeley
University of California San Diego UCSD
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Publication of WO2024197304A2 publication Critical patent/WO2024197304A2/fr
Publication of WO2024197304A3 publication Critical patent/WO2024197304A3/fr
Anticipated expiration legal-status Critical
Pending legal-status Critical Current

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Classifications

    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
    • A61F2/00—Filters implantable into blood vessels; Prostheses, i.e. artificial substitutes or replacements for parts of the body; Appliances for connecting them with the body; Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
    • A61F2/82—Devices providing patency to, or preventing collapsing of, tubular structures of the body, e.g. stents
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
    • A61F2250/00—Special features of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof
    • A61F2250/0001—Means for transferring electromagnetic energy to implants
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61F—FILTERS IMPLANTABLE INTO BLOOD VESSELS; PROSTHESES; DEVICES PROVIDING PATENCY TO, OR PREVENTING COLLAPSING OF, TUBULAR STRUCTURES OF THE BODY, e.g. STENTS; ORTHOPAEDIC, NURSING OR CONTRACEPTIVE DEVICES; FOMENTATION; TREATMENT OR PROTECTION OF EYES OR EARS; BANDAGES, DRESSINGS OR ABSORBENT PADS; FIRST-AID KITS
    • A61F2250/00—Special features of prostheses classified in groups A61F2/00 - A61F2/26 or A61F2/82 or A61F9/00 or A61F11/00 or subgroups thereof
    • A61F2250/0058—Additional features; Implant or prostheses properties not otherwise provided for
    • A61F2250/0067—Means for introducing or releasing pharmaceutical products into the body
    • A61F2250/0068—Means for introducing or releasing pharmaceutical products into the body the pharmaceutical product being in a reservoir

Definitions

  • the global implantable drug delivery devices market size was estimated at USD 16.6 billion in 2016 and is anticipated to exhibit a lucrative CAGR of 7.5% over the coming years.
  • a rise in the incidence of chronic neurological conditions, such as neurodegenerative, neurooncological, and neuroinflammatory diseases, is creating a need for efficient medication delivery.
  • the blood brain barrier significantly reduces the concentration of cytoactive drugs that penetrate into the central nervous system to affect its action on target cells.
  • drugs like Nusinersen (and other such gene therapies) that are used to treat spinal muscular atrophy and other such neurogenetic diseases require discomforting lumbar and/or cervical punctures in patients who often have skeletal deformities and poorly tolerate this procedure without significant discomfort.
  • Neuro-oncological, neurogenerative (Alzheimer's disease, Parkinson's Disease, etc.), and the wide variety of difficult to control neurologic diseases require frequent and poorly penetrating drugs to be systemically administered.
  • the surgical insertion of electrodes for deep brain and /or cortical neuromodulation often results in damage to normal brain parenchyma, as well as progressive scarring/gliosis at the tissue interface directly contacting the neuroanatomical target leading to waning efficacy over time.
  • the volume of tissue activated by the electrically conducting electrode is constrained by the Euclidean distance to the neuroanatomical target of interest.
  • the spatiotemporal expression or control of intrinsic or exogenous genes can be performed with targeted acoustic stimulation of cells transfected with select Transient Receptor Potential channels, which also have the potential to enhance precision treatment of neurologic diseases and disorders by activating, silencing, or modulating gene function.
  • a minimally invasive, chronically implantable solution harnessing ultrasound opens up vast possibilities for the efficacious treatment of neurological disease.
  • a stent system including a support structure having first and second ends having one or more members defining interior and peripheral regions for the structure, a lattice interfaced with the support structure, at least one transducer attached to the lattice, one or more leads electrically connected to the at least one transducer, and a power source connected to the one or more lead wires.
  • the at least one transducer is at least partially formed in one or more shapes selected from the group consisting of: non-planar, curved, cylindrical, concentrical circles, concentric circles with struts, hemispherical, annular, ring, half-ring, spherical, rectangular, prism, pyramidal, helical and polyhedral.
  • the at least one transducer is arranged in one or more patterns or configurations selected from: radial, radially outward, radially inward, focused inward, focused outward, circumferential, peripheral, helical, double-helical, triplicate, staggered, eccentric, biased, symmetrical, asymmetrical, or grouped.
  • the at least one transducer has length of 5 pm to 50 mm, a width or diameter of 5 pm to 50 mm, and a height of 5 pm to 50 mm. In some embodiments, the at least one transducer has a curvature ranging from 0 degrees to 180 degrees, a focal depth ranging between 0.1 mm and 25 mm, or an arc length ranging from 0.1 pm to 10 mm.
  • the support structure has a length of 10 mm to 60 mm, and a width or diameter of 0.5 mm to 20 mm. In some embodiments, the one or more members of the support structure have a width or diameter between 0.01 mm and 2 mm. In some embodiments, the openings of the lattice range between 0.1 mm and 5 mm.
  • the one or more members of the support structure are formed of one or more materials selected from the group consisting of: elastic material, superelastic material, sputtered superelastic material, resilient material, nonconductive metal or alloy, poorly conductive metal or alloy, platinum, magnesium, stainless steel, nickel-titanium alloy/nitinol, nitinol thin fdm (i.e., NiTi thin fdm), cobalt-nickel alloys, tantalum alloys, cobaltchromium alloys, platinum-tungsten metal, cobalt alloy, platinum-iridium alloy, cobalt, chromium, conductive metal, gold, copper, silver, titanium, titanium alloys, or any combination thereof.
  • the lattice is formed of one or more materials selected from the group consisting of: elastic material, superelastic material, sputtered superelastic material, resilient material, nonconductive metal or alloy, poorly conductive metal or alloy, platinum, magnesium, stainless steel, nickel-titanium alloy/nitinol, nitinol thin film (i.e., NiTi thin film), cobalt-nickel alloys, tantalum alloys, cobalt-chromium alloys, platinum-tungsten metal, cobalt alloy, platinum-iridium alloy, cobalt, chromium, conductive metal, gold, copper, silver, titanium, titanium alloys, or any combination thereof.
  • the lead wires are formed of one or more materials selected from the group consisting of: conductive metal, conductive metal allow, gold, copper, aluminum, silver, insulating material, biocompatible material, polymer material, synthetic material, plastic, or any combination thereof.
  • At least one of the support structure, lattice, at least one transducer and one or more lead wires are at least partially coated with one or more coatings selected from the group consisting of: biocompatible coating, conductive coating, nonconductive coating, insulating coating, oxide-film, titanium-nitride-oxide, organic polymer, polyurethane, polyimide, bioactive/covalently bonded coating, covalently-linked heparin, polymerized allylamine, poly(dimethylsiloxane), phosphorylcholine, 2-methylacryloyloxyethyl phosphorylcholine, poly(2 -methoxyethyl acrylate), enzyme-loaded inducible nitric acid oxide, synthase oxygenase on polyethylenimine, chromium, platinum, gold, silver, stealth coating, trimethylolpropane triacrylate (TMPTA), or any combination thereof.
  • one or more coatings selected from the group consisting of: biocompatible coating, conductive coating,
  • the at least one transducer is a piezoelectric device including a bipolar ceramic stack, at least one electrode, a back mass, a front mass, and at least one biocompatible coating covering at least a portion of the at least one transducer, wherein the transducer is coated in trimethylolpropane triacrylate (TMPTA).
  • TMPTA trimethylolpropane triacrylate
  • the system further includes a computing system communicatively connected to the power source, having a processor and a non-transitory computer-readable medium with instructions stored thereon, which when executed by a processor, perform steps including providing power at one or more voltages in one or more frequencies from the power source to the at least one transducer for at least one duration.
  • the one or voltages is between 0.1 V to 500 V
  • the one or more frequencies are between 0.1 Hz and 10000 MHz
  • the at least one duration is between 0.1 ps and 1 hour.
  • the system further includes at least one microfluidic chamber positioned in the interior region of the support structure fluidly connected to at least one microfluidic reservoir positioned outside of the support structure with at least one conduit having at least one lumen.
  • the at least one microfluid chamber is at least partially formed in one or more shapes selected from: cylinder, disk, toroid, annulus, capsule, pill, cube, sphere, dome, arc, spiral, helix, double helix, or any combination thereof.
  • the at least one microfluid chamber has a volume ranging between 1 pl and 1000 pl or a diameter between 0.1 mm and 1 mm, the at least one microfluid reservoir has a volume ranging between 100 pl and 2 ml.
  • the at least one conduit has a length ranging between 1 mm and 200 cm.
  • the at least one microfluidic chamber includes first and second microfluidic chambers, and the at least one microfluidic chamber includes first and second microfluidic chambers.
  • the first microfluidic chamber and first microfluidic reservoir are at least partially fluidly filled with a solution including microbubbles
  • the second microfluid chamber and second microfluidic reservoir are at least partially fluidly filled with a solution including at least one therapeutic.
  • the at least one therapeutic is selected from the group consisting of: gene therapy, Nusinersen, Levodopa, anti-seizure medication, DNA or RNA based biologic/viral-vector gene replacement therapies, CRISPR, Nusinersen for SMA, RGX-111, RGX-1 12, condensed tuberin, AP0E2, hTERT, GBA, SOD-1 , C9orf72, HTT, antibody-based biologic therapies, Lecanemab, small molecules, proteins, or any combination thereof.
  • the system further includes at least one port in each microfluidic reservoir configured to receive one or more fluids or solutions.
  • the system further includes an implantable housing at least partially enclosing the power source, the first microfluidic reservoir, and the second microfluidic reservoir.
  • the system further includes at least one tubular conduit having at least one lumen extending from the housing to a position on the first end of the support structure.
  • aspects of the present invention relate to a method for spatiotemporal control of intrinsic or exogenous genes, including providing any disclosed stent system, implanting the stent system near or in a target site of a subject, providing power at one or more voltages in one or more frequencies from the power source to the at least one transducer for at least one duration.
  • the one or voltages is between 0.1 V to 500 V
  • the one or more frequencies are between 0.1 Hz and 10000 MHz
  • the at least one duration is between 0.1 ps and 1 hour.
  • aspects of the present invention relate to a drug delivery method including providing any disclosed stent system, implanting the stent system in or near a target site of a subject, filling the at least one microfluidic chamber at least partially with a solution including at least one therapeutic and microbubbles, producing microbubble cavitation via the at least one transducer and delivering the at least one therapeutic to the target site.
  • aspects of the present invention relate to a method of fabricating a stent system including fabricating a stent, fabricating at least one PZT transducer, fabricating at least one microfluidic chamber, fixing the at least one PZT transducer to the stent, disposing the at least one microfluidic chamber within the stent.
  • the step of fabricating the PZT transducer includes 3D printing a PZT base layer on a polymer support layer, debonding the PZT base layer from the polymer support layer, sintering the PTZ base layer in a PTZ powder bed with heat to form a PTZ element, submerging the PTZ element into an isolation liquid, polarizing the dipoles within the sintered PTZ element, applying an electric field to the PTZ element and cooling the PTZ element.
  • Fig. 1A and Fig. IB depict an exemplary advanced function stent having one or more transducers according to aspects of the present invention.
  • Fig. 1 A depicts a side view of an exemplary advanced function stent.
  • Fig. IB depicts a perspective view of an exemplary advanced function stent.
  • Fig. 2 depicts an exemplary Intravascular Ultrasound transducer design (IVUS).
  • IVUS Intravascular Ultrasound transducer design
  • Fig. 3 depicts an exemplary Advanced Function Stent System according to aspects of the present invention.
  • Fig. 4 depicts an illustrative computer architecture for a computer for practicing the various embodiments of the invention.
  • Fig. 5A through Fig. 5C depict an exemplary Intravascular Ultrasound (IVUS) transducer design and an exemplary 3D printed Piezoelectric transducer according to aspects of the present invention.
  • Fig. 5A depicts an exemplary manufacturing approach comprising additive manufacturing (AM) and a liquid phase sintering.
  • Fig. 5B depicts an exemplary fully 3D printed IVUS transducer with focusing features.
  • Fig. 5C depicts an exemplary application for the disclosed advanced function stent system displaying ultrasound-enabled localization cavitation in a blood vessel phantom.
  • Fig. 6A through Fig. 6G depict an exemplary 3D printed Piezoelectric transducer according to aspects of the present invention. Shown is an exemplary miniaturized ultrasound transducer design and 3D printing of a highly loaded Lead Zirconate Titanate (PZT) composite and liquid phase sintering.
  • Fig. 6A shows an exemplary miniaturized ultrasound transducer with curved PZT element configured to produce an acoustic beam with a focal line sized appropriately to at least partially surround a vein according to aspects of the present invention.
  • Fig 6B shows an exemplary schematic of the liquid sintering resin components.
  • Fig. 6C shows an exemplary novel 3D printing system for liquid phase sintering (EPS) piezoelectric composites.
  • EPS liquid phase sintering
  • FIG. 6D shows an example debonding process to bum off the supportive polymer.
  • Fig. 6E shows an example liquid sealing process to reduce the lead loss during high-temperature sintering.
  • Fig. 6F shows an example LPS process to form dense PZT sample.
  • Fig. 6G shows exemplary generated acoustic pressure curves of the 3D printed 9.75-Mhz miniaturized ultrasound transducers, indicating that the values exceeded many medical thresholds.
  • Fig. 7A through Fig. 7J depict exemplary transducer designs, a PZT free-form fabrication and piezoelectric performance of the as-fabricated piezoelectric elements (i.e, transducers).
  • Fig. 7A shows exemplary transducer designs comprising micro-curved stave elements with different curvatures.
  • Fig. 7B shows exemplary transducer designs comprising hemisphere elements typically found in a general ultrasound transducer for medical imaging and nondestructive testing.
  • Fig. 7C shows an exemplary transducer design comprising helical elements for generating spiral acoustic fields for ultrasound manipulation.
  • Fig. 7D shows exemplary transducer designs comprising cylindrical transducer elements with multiple concentric annular layers.
  • Fig. 7A through Fig. 7J depict exemplary transducer designs, a PZT free-form fabrication and piezoelectric performance of the as-fabricated piezoelectric elements (i.e, transducers).
  • Fig. 7A shows exemplary transducer
  • FIG. 7E shows an exemplary transduce design comprising lattice sensor elements for low-frequency ( ⁇ 100 kHz) ultrasound.
  • Fig. 7F shows an exemplary transducer design comprising a force sensor element with ultrahigh sensitivity.
  • Fig. 7G shows a scanning electron microscope (SEM) image comparison between samples sintered by conventional sintering and optimized sintering methods.
  • Fig. 7H is a plot showing the results for sound speed measurements against the sample thickness.
  • Fig. 71 is a plot showing the results for polarization-electric field (P-E) loop of an exemplary 3D printed PZT transducer. The remnant polarization P r is 32.4 pC/cm 2 , which is 84% of the value of pristine materials.
  • Fig. 7J is a plot showing the results for d33 and coupling factor kt benchmarked with the disclosed 3D printed piezoelectric materials.
  • Fig. 8A through Fig. 8D show an exemplary transducer design and acoustic energy output performance of the as-fabricated miniaturized ultrasound transducer.
  • Fig. 8A is a plot showing tunable attenuation coefficient and acoustic impedance for the exemplary of Fig. 8C.
  • Fig. 8B is a plot showing the acoustic pressure map vs. curvature and arc length for the 3D printed miniaturized ultrasound transducer of Fig. 8C, showing that the generated acoustic pressure already exceeds many medical thresholds.
  • Fig. 8C depicts an exemplary 3D printed miniaturized ultrasound transducer according to aspects of the present invention.
  • Fig. 8D shows an image of an exemplary 3D printed miniaturized ultrasound transducer with a focused micro- PZT element according to aspects of the present invention.
  • Fig. 9A through Fig. 91 show an exemplary miniaturized ultrasound transducer and test results.
  • Fig. 9A shows a schematic of an experiment with a disclosed 3D printed transducer capturing the cavitation signal in a microbubble suspension with an enlarged view of microbubble fragmentation.
  • Fig. 9B is a plot showing the captured signals (Input voltage: 15 Vpp, 40 Vpp, 90 Vpp and 180 Vpp) in time domain.
  • Fig. 9C, Fig. 9D and Fig. 9E are plots showing frequency spectrums of the received signals driven by different input voltages. The subharmonic and broadband are identified as the generation of stable cavitation and inertial cavitation, respectively.
  • Fig. 9F, Fig. 9G, Fig. 9H, and Fig. 91 are a series of images showing microbubble behavior during insonation, showing the processes of oscillation, growth, and collapse of a microbubble.
  • Fig. 10A through Fig. 10H show an exemplary miniaturized ultrasound transducer and induced localized cavitation in a blood vessel phantom.
  • Fig. 10A depicts the transducer positioned at a target site or region of interest.
  • Fig. 10B depicts injecting a microbubble suspension or solution into or near the target site or region of interest.
  • Fig. 10C depicts Ultrasound-induced microbubble opening via the exemplary 3D printed miniaturized ultrasound transducer in the 3D printed blood vessel phantom.
  • Fig. 10D is a plot showing microbubble opening vs. time, indicating that the opening rate is related to the input voltage.
  • Fig. 10A depicts the transducer positioned at a target site or region of interest.
  • Fig. 10B depicts injecting a microbubble suspension or solution into or near the target site or region of interest.
  • Fig. 10C depicts Ultrasound-induced microbubble opening via the exemplary 3D printed miniatur
  • FIG. 10E depicts the transducer trajectory and microbubble on the inner wall of the phantom blood vessel.
  • Fig. 10F depicts microbubble fragmentation as a result of the path of the transducer.
  • Fig. 10G depicts microbubble fragmentation as a result of the path of the transducer.
  • Fig. 10H depicts a localized drug delivery testing wherein the microbubbles can be controlled to open only at the focal area of the transducer.
  • Fig. 11A shows an exemplary two-step debonding process.
  • Fig. 1 IB is a plot showing the debonding time-temperature curve of the process of Fig. 11A.
  • Fig. 12A is a plot showing peak sintering temperature optimized by d.33 constant.
  • Fig. 12B is a plot showing breakdown electrical field of the exemplary samples.
  • Fig. 12C is a plot showing an X-Ray diffraction analysis (XRD) of the sintered PZT sample.
  • XRD X-Ray diffraction analysis
  • Fig. 13 A depicts an exemplary polarization setup or transducer design according to aspects of the present invention.
  • Fig. 13B is a diagram depicting dipole alignment along the applied electrical field.
  • Fig. 13C is a plot showing polarization time and electrical field profile for the transducer design of Fig. 13 A.
  • Fig. 14A depict exemplary transducer material samples prior to sintering.
  • Fig. 14B depict exemplary transducer material samples after sintering, showing their dimension shrinkages were identical after the sintering process.
  • Fig. 14C is a plot showing average linear shrinkage vs probability density (i.e., structural integrity measurement and results) of samples after sintering (p is the mean value while G is the standard deviation).
  • Fig. 14D depicts an exemplary curved transducer material sample with focusing feature prior to sintering.
  • Fig. 14E depicts an exemplary curved transducer material sample with focusing feature after sintering, showing the curvature was well maintained during sintering.
  • Fig. 14A depict exemplary transducer material samples prior to sintering.
  • Fig. 14B depict exemplary transducer material samples after sintering, showing their dimension shrinkages were identical after the sintering process.
  • Fig. 14C is a plot showing average linear
  • FIG. 14F is a plot showing probability distribution for curvature (p is the mean value while G is the standard deviation)
  • Fig. 14G depicts an exemplary 3D printed honeycomb lattice transducer (e g., lattice sample).
  • Fig. 14H is an XCT image of the unit cell of the lattice sample.
  • Fig. 141 is a plot showing probability distribution for strut thickness (p is the mean value while G is the standard deviation).
  • Fig. 15A is a diagram of an experimental set up for sound speed and acoustic attenuation measurement of ultrasound through a sample or material.
  • Fig. 15B depicts an enlarged view of the material to be tested p : acoustic pressure)
  • Fig. 15C are plots depicting the received signals with and without sample in between transducers.
  • Fig. 16A and Fig. 16B show Strain-Electric (S-E) field loop measurement of a commercial sample vs an exemplary 3D printed sample of the present invention.
  • Fig 16A is a diagram showing the experimental the S-E loop testing system.
  • Fig 16B is a plot showing the S- E curves of the commercial PZT sample and the exemplary 3D printed sample.
  • Fig. 17A through Fig. 17F depict various exemplary 3D printed transducers of different geometries and their related energy output capability.
  • Fig. 17A depicts an exemplary transducer design and a plot showing the energy output.
  • Fig. 17B is a diagram depicting a schematic of the acoustic pressure measurement setup.
  • Fig. 17C is a plot of x-y coordinate vs intensity for the transducer of Fig. 17A.
  • Fig. 17D depicts an exemplary transducer design and a plot showing the energy output.
  • Fig. 17E depicts an exemplary transducer design and a plot showing the energy output.
  • Fig. 17F depicts an exemplary transducer design and a plot showing the energy output.
  • Fig. 17A through Fig. 17F depict various exemplary 3D printed transducers of different geometries and their related energy output capability.
  • Fig. 17A depicts an exemplary transducer design and a plot showing the energy output
  • FIG. 18A depicts exemplary transducer designs (one transducer comprising a disk element and an another comprising an element with concentric annular layers).
  • Fig. 18B depicts the disk element transducer design.
  • Fig. 18C depicts the transducer design comprising an element with concentric annular layers.
  • Fig. 18D is a plot showing peak-to-peak acoustic pressure generated by the transducers.
  • Fig. 18E and Fig. 18F show example finite element analysis (FEA) models to predict the acoustic pressure and the acoustic fields generated by the transducers.
  • FEA finite element analysis
  • Fig. 19 is a plot showing impedance curves of the fabricated PZT sample and standard PZT sample.
  • Fig. 20 is a plot showing pulse-echo signal and its frequency spectrum for a 3D printed miniatured ultrasound transducer.
  • Fig. 21 is a diagram with images depicting dye diffusion speed and uniformity calculation based on the pixel value of grayscale images of a disclosed transducer near or in a target site or region of interest.
  • Fig. 22A depicts a dye dispersion experiment set up in water.
  • Fig. 22B depicts a 0.1 ml methylene blue drop is added to the media of the water experiment.
  • Fig. 22C depicts where 9.75-MHz burst waves were input into the transducer, and the dye was stirred by the emitted acoustic wave in the water experiment.
  • Fig. 22D depicts where after 15 seconds of ultrasound exposure, the dye partially dispersed in the water.
  • Fig. 22E depicts a timeline across the experiments of Fig. 22A and Fig. 22F.
  • Fig. 22F depicts a dye dispersion experiment set up in a microbubble suspension.
  • Fig. 22A depicts a 0.1 ml methylene blue drop is added to the media of the water experiment.
  • Fig. 22C depicts where 9.75-MHz burst waves were input into the transducer, and the dye was stirred by the emitted acoustic wave in
  • FIG. 22G depicts a 0.1 ml methylene blue drop is added to the media of the microbubble experiment.
  • Fig. 22H depicts where 9.75-MHz burst waves were input into the transducer, and the dye was stirred by the emitted acoustic wave in the microbubble experiment.
  • Fig. 221 depicts where after 15 seconds of ultrasound exposure, the dye fully dispersed in the microbubble suspension.
  • Fig. 23 is a plot showing a digital image correlation curve of microbubble concentration in the blood vessel phantom vs. image mean pixel value.
  • Fig. 24A is a series of images depicting a progression of microbubble collapse and fragmentation.
  • Fig. 24B is a series of images depicting a progression of microbubble collapse and fragmentation.
  • Fig. 25A is a series of images depicting a progression of dye dispersion from microbubble opening.
  • Fig. 25B is a series of images depicting a progression of dye dispersion from microbubble opening.
  • Fig. 25C is a series of images depicting a progression of dye dispersion from microbubble opening
  • Fig. 26A is a series of images depicting a progression of microbubble fragmentation.
  • Fig. 26B is a series of images depicting a progression of microbubble fragmentation.
  • Fig. 27A is a series of images depicting a progression of microbubble fragmentation.
  • Fig. 27B is a series of images depicting a progression of microbubble fragmentation.
  • patient refers to any animal amenable to the systems, devices, and methods described herein.
  • patient, subject or individual may be a mammal, and in some instances, a human.
  • ranges throughout this disclosure, various aspects of the invention can be presented in a range format. It should be understood that the description in range format is merely for convenience and brevity and should not be construed as an inflexible limitation on the scope of the invention. Accordingly, the description of a range should be considered to have specifically disclosed all the possible subranges as well as individual numerical values within that range. For example, description of a range such as from 1 to 6 should be considered to have specifically disclosed subranges such as from 1 to 3, from 1 to 4, from 1 to 5, from 2 to 4, from 2 to 6, from 3 to 6 etc., as well as individual numbers within that range, for example, 1, 2, 2.7, 3, 4, 5, 5.3, and 6. This applies regardless of the breadth of the range.
  • aspects of the present invention relate to advanced function stent systems and methods thereof for producing or delivering one or more therapies to or near a target site on a subject (e.g., acoustic stimulation therapy, sonogenetic control of cells, acoustic modulation of cell function, targeted drug delivery, chronic on-demand drug delivery, etc).
  • a target site on a subject e.g., acoustic stimulation therapy, sonogenetic control of cells, acoustic modulation of cell function, targeted drug delivery, chronic on-demand drug delivery, etc.
  • stent system 100 comprises a support structure 101 (in some examples referred to as a stent scaffold or stent structure) having a first end 113 and second end 115 comprising one or more members 102 defining an interior or inner region 114 and an outer or peripheral region 116, a lattice 103 (in some examples referred to as a lattice structure) interfaced with support structure 101, at least one transducer 104 attached to lattice 103, one or more leads 105 electrically connected to the at least one transducer 104, and a power source (not shown) connected to the one or more leads 105.
  • a support structure 101 in some examples referred to as a stent scaffold or stent structure
  • lattice 103 in some examples referred to as a lattice structure
  • leads 105 electrically connected to the at least one transducer 104
  • a power source not shown
  • stent system 100 may be positioned near, in, within, or in proximity to at least one target site or region of interest of a subject using any known method or device including but not limited to: catheters, cannulas, guide wires, push wires, pusher wires, or the like, or any combination thereof.
  • Support structure 101 comprises one or more members 102 configured, patterned, and/or arranged to form a scaffold such that lattice 103 may interface with the members and form a stent.
  • support structure 101 may comprise any number of members.
  • the one or more members 102 comprise 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19 or 20 members.
  • the one or more members 102 and/or lattice 103 comprise one or more wires interfaced together to form a stent.
  • the one or more wires of support structure 101 and/or the one or more wires of lattice 103 are interfaced, latticed, braided, woven, meshed, laced, interwoven, intertwined, interlaced, reticulated, wired, crossed, crisscrossed, wound, or the like, or any combination thereof, in order to form the stent.
  • the support structure 101 and/or lattice 103 comprise or are formed from rigid, flexible, compressible, or elastic materials. It should be appreciated that portions or all of support structure 101 and/or lattice 103 may have rigid, flexible, compressible and/or elastic portions, regions or areas.
  • lattice 103 partially encloses the members of support structure 101, or fully or partially encloses the members.
  • Support Structure 101, lattice 103 and/or leads 105 may be fixedly and/or releasably attached with any known method including attachments, weldments, bondings, glues, brackets, windings, wire bonding, wire wrapping, wire twisting, thermosonic binding, solders, or the like, or combinations thereof.
  • portions or all of support structure 101 and/or lattice 103 may be formed as a single unit or component, or may be formed separately, or with one or more components or structures each.
  • support structure 101 and/or lattice 103 may be configured or formed to have at least an undeployed state and a deployed state for the delivery and deployment of the stent system.
  • stent system 100 comprises a first state or configuration, and a second state or configuration, wherein the first state is a compressed delivery state against the spring bias of the support structure and/or lattice 103, and the second state is a relaxed, expanded delivery state or configuration.
  • support structure 101 and/or lattice 103 comprise a spring bias that allows them to return to their expanded, relaxed state.
  • support structure 101 and/or lattice 103 comprise one or more expanding sections, or the entirety of each expands, or portions of one or both expand.
  • portions or all of stent system 100 can be compressed into a delivery device, such as a catheter, and later expand upon deployment when implanted.
  • the support structure may also comprise one or more scaffolds or lattice type structures in addition to the one or more members 102, or in replacement of the one or more members 102.
  • the one or more members 102 of the support structure 101 are formed of one or more materials selected from the group consisting of: elastic material, superelastic material, sputtered superelastic material, resilient material, rigid materials, flexible materials, nonconductive metal or alloy, poorly conductive metal or alloy, platinum, magnesium, stainless steel, nickel -titanium alloy/nitinol, nitinol thin fdm (i.e., NiTi thin film), cobalt-nickel alloys, tantalum alloys, cobalt-chromium alloys, platinum-tungsten metal, cobalt alloy, platinumiridium alloy, cobalt, chromium, conductive metal, gold, copper, silver, titanium, and titanium alloys.
  • portions or all of support structure 101 are any of braided, laser cut, and/or magnetron sputtered thin film.
  • the support structure 101 has a length of 15 mm to 60 mm, and a width or diameter of 1.5 mm to 20 mm, or a length between 0.1 mm and 100 mm, and a diameter between 0.001 mm and 200 mm.
  • the one or more members 102 of the support structure have a width or diameter between 0.01 mm and 2 mm, or between 0.001 and 5 mm.
  • support structure 101 has a length of 1 mm to 35 mm and a diameter of 0.1 mm to 5 mm.
  • the lattice has repeating subunits or components with openings, wherein each subunit or component comprises one or more connected members, and each member has a diameter or width between 0.01 mm and 2 mm. In some embodiments, the openings of the lattice range between 0.1 mm and 5 mm.
  • the lattice 103 is formed of one or more materials selected from the group consisting of: elastic material, superelastic material, sputtered superelastic material, resilient material, rigid materials, flexible materials, nonconductive metal or alloy, poorly conductive metal or alloy, platinum, magnesium, stainless steel, nickel-titanium alloy/nitinol, nitinol thin film (i.e., NiTi thin film), cobalt-nickel alloys, tantalum alloys, cobaltchromium alloys, platinum-tungsten metal, cobalt alloy, platinum-iridium alloy, cobalt, chromium, conductive metal, gold, copper, silver, titanium, and titanium alloys.
  • portions or all of lattice 103 are any of braided, laser cut, and/or magnetron sputtered thin film.
  • aspects of the present invention relate to the configuration, positioning, arrangement or patterning of one or more transducers on stent system 100.
  • the at least one transducer 104 may be configured on stent system 100 to direct, focus or beam energy over one or more focal areas or regions.
  • the at least one transducers 104 produces a focused beam, or a broad beam, or a hybrid beam.
  • the at least one transducer 104 is positioned at least partially in inner region 114.
  • the at least one transducer 104 is positioned at least partially in peripheral region 116.
  • the at least one transducer spans or is positioned across inner region 114 and peripheral region 116.
  • the at least one transducer 104 is arranged, configured, or positioned in one or more patterns or configurations selected from: radial, radially outward, radially inward, focused inward, focused outward, circumferential, peripheral, helical, double-helical, triplicate, staggered, eccentric, biased, symmetrical, asymmetrical, and grouped.
  • system 100 comprises a plurality of transducers (i.e., at least 2 transducer 104) arranged radially along peripheral region 116 of support structure 101.
  • the plurality of transducers are configured and/or attached eccentrically in a serial linear arrangement along the length of support structure 102.
  • the plurality of transducers are configured, arranged and/or attached rotationally along the length of support structure 101. In some embodiments, plurality of transducers are arranged and/or attached in a corkscrew arrangement along the length of support structure 101. Although example configurations are provided, it should be appreciated that the plurality of transducers may be arranged in any configuration along the length of support structure 101 in order to focus energy directed towards the inside of the stent or the outside of the stent. In some embodiments, the arrangement of the at least one transducer 104 confers tunable acoustic focusing properties for selective targeting of cells or tissue (e.g., neuronal populations) of interest.
  • cells or tissue e.g., neuronal populations
  • stent system 100 and/or the at least one transducer 104 is configured to selectively target one or more areas or regions in our surrounding the system.
  • stent system 100 may be configured or formed to selectively target within the inner region 114, or outwardly from the peripheral region 116, or both, either separately or simultaneously, or on a program or schedule.
  • stent system 100 may be positioned near, in, within, or in proximity to at least one target site or region of interest of a subject using any known method or device including but not limited to: catheters, cannula, guide wires, push wires, pusher wires, or the like, or any combination thereof.
  • the pusher wire fixedly attaches, releasably attaches, or connects to one end of support structure 102.
  • stent system 100 is moved along at least one trajectory to treat the at least one target site or region of interest.
  • a disclosed pusher wire length is 120 cm to 200 cm and a diameter of 0.035 mm to 0.80 mm.
  • the at least one transducer 104 is electrically connected to one or more leads 105 with any known method or material known by one of ordinary level of skill in the art.
  • the one or more leads 105 comprise one or more electrically conductive wires.
  • the one or more lead wires may comprise positive and negative lead wires.
  • leads 105 further comprise one or more insulators.
  • transducers 104 may be electrically connected to the one or more lead wires in any configuration known by one of ordinary level of skill in the art (e.g., series connection, parallel connection).
  • the one or more leads 105 may comprise any terminating connectors or plugs known in the art.
  • the one or more leads 105 are interfaced with any of support structure 101, one or more members 102, and/or lattice 103.
  • the one or more leads 105, the one or more members 102, and/or lattice 103 are interfaced, latticed, braided, woven, meshed, laced, interwoven, intertwined, interlaced, reticulated, wired, crossed, crisscrossed, wound, or the like, or any combination thereof.
  • one or more leads 105 partially encloses the one or more members 102 of support structure 101, or the repeating components of lattice 103, or fully encloses the members and components.
  • the one or more leads 105 are formed of one or more materials selected from the group consisting of: conductive metal, conductive metal allow, gold, copper, aluminum, silver, insulating material, polymer, or plastics, or any combination thereof.
  • at least one of the support structure 101, one or more members 102, lattice 103, and one or more leads 105 are at least partially coated with one or more coatings selected from the group consisting of: biocompatible coating, conductive coating, non-conductive coating, insulating coating, oxide-fdm, titanium-nitride-oxide, organic polymer, polyurethane, polyimide, bioactive/covalently bonded coating, covalently-linked heparin, polymerized allylamine, poly(dimethylsiloxane), phosphorylcholine, 2-methylacryloyloxyethyl phosphoryl choline, poly(2 -methoxyethyl acrylate), enzyme-loaded inducible nitric acid oxide, synthediditric acid
  • transducer 104 according to aspects of the present invention.
  • stent system 100 and/or transducer 104 further comprises at least one piezoelectric device or stack.
  • transducer 104 comprises a housing 140, a backing layer 142, a matching layer 144, a PZT element 146 and at least one electrode 148.
  • transducer 104 comprises a bipolar ceramic stack, electrode(s), a back mass and a front mass.
  • the at least one transducer 104 is a piezoelectric device comprising a bipolar ceramic stack, at least one electrode, a back mass, a front mass, and at least one biocompatible coating covering at least a portion of the at least one transducer 104.
  • the at least one transducer 104 is at least partially formed in one or more shapes selected from the group consisting of: non-planar, curved, cylindrical, concentrical circles, concentric circles with struts, hemispherical, annular, ring, half-ring, spherical, rectangular, prism, pyramidal, helical and polyhedral.
  • the at least one transducer 104 has length of 5 pm to 50 mm, or 15 pm to 25 mm a width or diameter of 5 pm to 50 mmm or 15 pm to 25 mm, and a height of 5 pm to 25 mm, or 15 pm to 25 mm.
  • the at least one transducer 104 has a curvature ranging from 0 degrees to 180 degrees, or 20 degrees to 160 degrees, or 40 degrees to 140 degrees, or 60 degrees to 120 degrees, or 80 degrees to 115 degrees, or 90 to 110 degrees, or 105 degrees to 115 degrees, or about 110 degrees, or an arc length ranging from 0.1 pm to 10 mm. It should be appreciated that the size of the transducer and the degree of curvature or arc length may create one or more focal distance lengths for the transducer.
  • the at least one transducer 104 has one or more focal lengths, such as lengths of about 0.1 gm, 0.2 gm, 0.3 gm, 0.4 gm, 0.5 gm, 0.6 gm, 0.7 gm, 0.8 gm, 0.9 gm or between 0.1 gm and 200 gm, or between 0.1 gm and 2 mm, or about 0.1 mm, 0.2 mm, 0.3 mm, 0.4 mm, 0.5 mm, 0.6 mm, 0.7 mm, 0.8 mm, 0.9 mm, 1.0 mm, or between 0.1 mm and 5 mm, or about 1 mm, or about 1.5 mm, or between 1 mm and 100 mm, or between 1 mm and 50 mm, or between 1 mm and 10 mm. It should be appreciated that any known transducer material, geometry or size may be used or incorporated into the at least one transducer 104 in order to tune or configure the at least one transducer
  • the at least one transducer 104 is at least partially coated with one or more coatings selected from the group consisting of: biocompatible coating, conductive coating, non-conductive coating, insulating coating, oxide-film, titanium-ni tri deoxide, organic polymer, polyurethane, polyimide, bioactive/covalently bonded coating, covalently-linked heparin, polymerized allylamine, poly(dimethylsiloxane), phosphorylcholine, 2-methylacryloyloxyethyl phosphorylcholine, poly(2-methoxyethyl acrylate), enzyme-loaded inducible nitric acid oxide, synthase oxygenase on polyethylenimine, chromium, platinum, gold, silver, trimethylolpropane triacrylate (TMPTA), parylene, anti-moisture coating and/or any anticorrosion coating.
  • one or more coatings selected from the group consisting of: biocompatible coating, conductive coating, non-conductive coating, insulating coating
  • transducer 104 comprises Lead Zirconate Titanate (PZT). In some embodiments, the transducer further comprises at least one biocompatible coating. In some embodiments, the biocompatible coating comprises trimethylolpropane triacrylate (TMPTA). In some embodiments, transducer 104 comprises a matching layer 144. In some embodiments, the matching layer 144 is a thin layer of material located between the transducer's piezoelectric crystal and the target to be treated. The matching layer 144 functions to improve acoustic energy transfer between the transducer and the target medium by reducing the impedance mismatch between these two materials. In some embodiments, the matching layer 144 comprises a mixture of carbon fiber (CF) and RIGID resin (Formlab, USA).
  • CF carbon fiber
  • RIGID resin Formlab, USA
  • RIGID resin is a mixture comprising urethane dimethacrylate and isobomyl methacrylate.
  • transducer 104 comprises a backing layer 142 and/or backing material.
  • the backing layer 142 of an ultrasound transducer is a layer of material located behind the piezoelectric crystal. The function of the backing layer 142 is to absorb and dampen the sound waves generated by the crystal, which helps to reduce the ringing or reverberation that can occur after each pulse of sound.
  • the backing layer 144 comprises an Fe magnetic particle-loaded FLEXIBLE resin (Formlab, USA). FLEXIBLE is a mixture comprising acrylated oligomers and acrylated monomers.
  • transducer 104 comprises a parylene coating.
  • transducer 104 comprises any FDA compliant coating, any anti-moisture coating and/or any anti-corrosion coating.
  • aspects of the present invention relate to providing power at one or more voltages in one or more frequencies from the power source to the at least one transducer 104 for at least one duration.
  • transducers, ultrasound transducers and pzeiolectric devices are disclosed herein that operate based on well-known parameters. This includes, but is not limited to, applying electrical power to the transducer in order to produce acoustic energy from the transducer.
  • a controller 108 and/or computer 400 may comprise a power source capable of generating the electrical power, frequencies, signals and other parameters required to operate said transducers in order to perform targeted acoustic stimulation of one or more cells, one or more areas of interest, regions of interest, or target sites on a subject.
  • Adjustable parameters for powering the disclosed transducers include, but are not limited to, voltage (e.g., Vpp), current, power, signal, frequency, duration, bursts, cycles, duty cycles, or the like, and any combination thereof.
  • Vpp voltage
  • the one or voltages is between 0.1 V to 500 V
  • the one or more frequencies are between 0.1 Hz and 10000 MHz
  • the at least one duration is between 0.1 ps and 1 hour
  • duty cycle between 0% and 100%.
  • the at least one transducer 104 is provided with 9.75-MHz burst waves with 50 cycles (20 ms burst period) with a 10% duty cycle.
  • the 9.75-MHz burst waves were set up to have 10% duty cycle.
  • the arrangement and powering of the transducers confers tunable acoustic focusing properties for selective targeting of neuronal populations of interest, and thusly may be arranged or powered in any configuration known by one of ordinary level of skill in the art for achieving the selective targeting of the one or more cells.
  • the performance of transducer 104 may be enhanced by adjoining it with one or more application specific integrated circuits (ASICs).
  • ASICs application specific integrated circuits
  • system 100 comprises at least one microfluidic chamber.
  • Each microfluid chamber comprises one or more side walls having inside and outside surfaces and enclosing an interior region wherein a fluid or solution may be stored securely while the stent system 100 is positioned and/or configured.
  • system 100 further comprises a first microfluidic reservoir 130 external to the stent, connected to and in fluid communication with a first microfluidic chamber 110.
  • system 100 further comprises a second microfluidic reservoir 132 connected to and in fluid communication with second microfluidic chamber 112.
  • first microfluidic reservoir 130 and second microfluidic reservoir 132 are configured to receive drug, solutions, fluids, and/or microbubbles delivered percutaneously via needle and syringe or other methods as would be known by one of ordinary skill in the art.
  • first microfluidic reservoir 130 comprises a first port 131 in fluid communication with first microfluidic chamber 110 and is configured to receive drug, fluid, solution, microbubbles, therapeutics etc delivered via needle and syringe.
  • second microfluidic reservoir 132 comprises a port 133 in fluid communication with second microfluidic chamber 112, and configured to receive fluid, solution, microbubbles, drug, therapeutics, etc via needle and syringe.
  • a single reservoir feeds all of the chambers connected in series, or parallel.
  • the size, location, shape etc of the at least microfluidic chamber may be impacted or controlled by the size, location and shape of the at least one transducer 104. For example, if the at least one transducer is directing energy towards inner region 114, then logically the at least one microfluidic chamber should be positioned in or near the focal area, focal region, or at or near the focal depth of the at least one transducer 104.
  • the at least one microfluidic chamber is disposed or positioned within the inner region 114 of support structure 101. In some embodiments, the at least one microfluidic chamber spans at least a portion of inner region 114 and peripheral region 116. In some embodiments, the at least one microfluidic chamber is deformable. In some embodiments, the at least one microfluidic chamber is porous. In some embodiments, the at least one microfluidic chamber is configured to release drug and/or microbubbles. In some embodiments, the at least one microfluidic chamber is configured to release drug and/or microbubbles when ultrasound is applied to the at least one microfluidic chamber.
  • stent system 100 further comprises least one microfluidic reservoir positioned outside of the support structure 101 with at least one conduit 106 comprising at least one lumen fluidly connecting to the at least one microfluidic chamber.
  • stent system 100 comprising at least one microfluid reservoir connected to at least one microfluidic reserervoir with at least one conduit.
  • stent system 100 comprises a support structure 101 having a first end 113, a second end 115, a length and diameter, and wherein at least one transducer 104 is connected or attached thereto.
  • support structure 101 interfaces with a cylindrical lattice 103 and forms an inner region 114, peripheral region 116.
  • support structure 101 and/or lattice 103 may form or be covered with an outer surface 118 and/or end caps, or end membranes.
  • system 100 further comprises a power source (not shown) electrically connected to at least one transducer 104.
  • system 100 further comprises a controller 108 connected to the power source and/or the at least one transducer 104 or one or more leads 105 (not shown).
  • the at least one conduit 106 is a flexible hypotube with or more lumens connected to a position on support structure 101, and to a position on controller 108.
  • system 100 further comprises at least one microfluidic chamber disposed within support structure 101, and in fluid communication with one or more of the lumens of conduit 106.
  • the at least one microfluidic chamber comprises a first microfluidic chamber 110 and second microfluidic chamber 112, and the at least one microfluidic reservoir comprises a first microfluidic reservoir 130 and second microfluidic reservoir 132.
  • first microfluidic chamber 110 is a microbubble chamber
  • the second microfluidic chamber 112 is a drug chamber.
  • each microfluidic reservoir comprises at least one port configured to receive one or more fluids or solutions.
  • a first microfluidic reservoir 130 comprises a port 131 and second microfluidic reservoir 132 comprises a port 133.
  • the flow of fluid or solution may travel from the one or more microfluidic reservoir through the at least one conduit to the one or more microfluidic chambers using any known method or device. In one example, that this may be achieved by modulating the flow or pressure of the one or more reservoir using any known method or device (e.g., a pump, a syringe fluidly connected to the port, a plunger, other mechanical means).
  • system 100 further comprises an implantable housing 108 at least partially enclosing the power source (not shown), the first microfluidic reservoir 130, and the second microfluidic reservoir 132.
  • the at least one conduit 106 comprises a tubular conduit having at least one lumen extending from the housing 108 to a position on the first end 113 of the support structure 101.
  • the housing 108 is configured to be implanted subcutaneously in the thorax and support structure 101 and lattice 103 is configured to be placed intravascularly.
  • the housing may be at least partially transdermal, subdermal, or subcutaneous, and the stent may be positioned within any structure including vascular, organs and tissue.
  • the stent is implanted in a superior sagittal sinus of a subject, wherein a portion of the system 100 percutaneously exits to a connector block for power, programming and fluid introduction (e.g., housing 108).
  • the microfluidic chambers may be formed in one or more shapes, or at least partially in one or more shapes, and be positioned or configured to any portion or region of stent system 100, support structure 101 and/or lattice 103.
  • the at least one microfluidic chamber is at least partially formed in one or more shapes selected from: cylinder, disk, toroid, annulus, capsule, pill, cube, sphere, dome, arc, spiral, helix, double helix.
  • first microfluidic chamber 110 is formed as a capsule shape and positioned in inner region 114
  • second microfluidic chamber 112 is formed as an annulus or ring configured or formed in or on peripheral region 116. Examples are provided with 1 or 2 microfluidic chambers, but there may be any number of microfluidic chambers, positioned in any configurations in or around the support structure 101 and/or lattice 103.
  • the at least one microfluid chamber comprises sidewalls formed of a permeable membrane, semi-permeable membrane, porous membrane, membrane, a hydrogel, a gel, a biomaterial, a biocompatible material, a bioresorbable material, a biodegradable material, a non- degradable material, a synthetic material, a polymer.
  • the at least one conduit is formed of a substantially non-porous material in order to direct fluid to the at least one microfluidic chamber, wherein the solution with drug may diffuse through the walls of the chamber.
  • the at least microfluidic chamber is configured to lyse, burst, hemorrhage, bust, or pop the microfluidic chamber upon insonation and microbubble cavitation in order to disperse the contained fluid, and is formed of appropriate materials and thicknesses to achieve such.
  • the at least one microfluidic chamber has a diameter ranging between about 0.100 mm to about 5 mm, or about 0.5 mm to 2 mm.
  • the diameter range would be about 0.100 mm to about 0.35 mm.
  • the at least one microfluid chamber is formed at least partially in one or more shapes selected from cylinder, a toroid, a capsule, a pill, a cube, a sphere, a dome, an arc, a spiral, a double helix, irregular, non-planar, curved, banana.
  • the at least one microfluid chamber has a volume ranging between 50 pl and 1000 pl, or 1 pl to 100 mb, or a diameter between 0.1 mm and 1 mm, the at least one microfluid reservoir has a volume ranging between 100 pl and 10 ml.
  • aspects of the present invention relate to fluids, solutions, suspensions or the like that at least partially fluidly fill the disclosed microfluidic chambers and reservoirs.
  • the at least one microfluidic chamber and/or microfluidic reservoir are at least partially fluidly filled with one or more fluids, solutions, suspensions, drug suspensions, drug solutions, microbubble solutions, microbubble suspensions, drug-microbubble solutions, solution comprised drug-encapsulated microbubbles, or the like, or any combination thereof.
  • first microfluidic chamber 110 and first microfluidic reservoir 130 are at least partially fluidly filled with a solution comprising microbubbles
  • second microfluidic chamber 112 and second microfluidic reservoir 132 are at least partially fluidly filled with a solution comprising at least one therapeutic.
  • aspects of the present invention relate to delivering at least one therapeutic to a region of interest, an area of interest, or a target site on a subject.
  • the at least one microfluidic chamber may be pre-filled with a solution prior to placement of the stent, or in another embodiment, the at least one microfluidic reservoir is filled, and the solution travels along one or more lumens of the at least one conduit to the at least one microfluidic chamber.
  • the at least one therapeutic is selected from the group consisting of: gene therapy, Nusinersen, Levodopa, anti-seizure medication, DNA or RNA based biol ogic/viral -vector gene replacement therapies, CRISPR, Nusinersen for SMA, RGX-111, RGX-112, condensed tuberin, AP0E2, hTERT, GBA, SOD-1, C9orf72, HTT, antibody-based biologic therapies, Lecanemab, small molecules and proteins.
  • outer surface 118 comprises any of a membrane, a porous membrane, a hydrogel, a gel, a biomaterial, a biocompatible material, a synthetic material, a polymer.
  • outer surface 118 comprises one or more coatings comprising at least one of oxide-film (e.g., titanium-nitride-oxide and the like), organic polymers (e.g., polyurethane, polyimide, and the like), or other bioactive/covalently bonded coatings (e.g., covalently-linked heparin, polymerized allylamine, poly(dimethylsiloxane), phosphorylcholine, 2-methylacryloyloxyethyl phosphorylcholine, poly(2-methoxyethyl acrylate), enzyme-loaded inducible nitric acid oxide synthase oxygenase on polyethylenimine, and the like), biocompatible coating, stealth coating, or any combination thereof.
  • oxide-film e.g., titanium-nitride-oxide and the like
  • organic polymers e.g., polyurethane, polyimide, and the like
  • bioactive/covalently bonded coatings e.g., co
  • conduit 106 connects along the length of support structure 101. In some embodiments, conduit 106 connects to the first end 113, or the second end 115 of support structure 101. In some embodiments, conduit 106 connects to axially to first end 113 of support structure 101. In some embodiments, conduit 106 connects to second end 115 wherein conduit 106 is centered within the stent, or from the side of the stent. In another example, conduit 106 connects the first end 113 of support structure 101 eccentrically.
  • controller 108 is an electrical power source for providing power to the at least one transducer 104.
  • controller 108 comprises a housing and is configured to be implanted subcutaneously.
  • controller 108 is a programmable thorax implant.
  • controller 108 comprises a console 126, a battery 128, a first microfluidic reservoir 130, a first port 131, a second microfluidic reservoir 132 and a second port 133.
  • controller 108 controls the release of fluid, solution or drug from the reservoirs.
  • controller 108 comprises a graphical user interface (GUI) that allows a user to control the release of fluid or drug from the at least one reservoir (e g., from the first microfluidic reservoir 130).
  • GUI graphical user interface
  • controller 108 provides a schedule for the release of drug from the first microfluidic reservoir 130 and/or the acoustic stimulation from the at least one transducer 104.
  • system 100 and/or controller 108 comprise at least one sensor measuring at least one signal from the subject, wherein drug, fluid or solution from first microfluidic reservoir 130 is released when at least one threshold value is measured with the sensor (e.g., a temperature sensor, a pressure sensor).
  • system 100 is a closed-loop system wherein feedback from at least one sensor provides control over system 100.
  • system 100 is an open-loop system wherein control over the system is provided by at least one user or a health care professional.
  • controller 108 comprises one or more computing devices (e g., computing device 400 as elaborated upon herein).
  • controller 108 is a subcutaneous implant and is a programmable device that may be battery powered or wirelessly charged (near field radio frequency).
  • controller 108 encases reservoirs for the percutaneously injected bubbles and drugs.
  • controller 108 controls on/off functions and parameters for at least one transducer 102.
  • system 100 further comprises a conduit 106 comprising a hyptotube.
  • conduit 106 comprises a first end 107, a second end 109, and a length therebetween.
  • conduit 106 comprises at least one lumen extending axially therethrough.
  • conduit 106 connects support structure 101 and controller 108.
  • the conduit 106 comprises at least one microfluidic tubular channel forming a lumen fluidly connected to the at least one microfluid chambers, and at least one channel forming a lumen for running a length of an electrical lead (e g., one or more leads 105) or electrical wire connections for the at least one transducer 104.
  • an electrical lead e g., one or more leads 105
  • conduit 106 fluidly connects microfluidic chamber 110 and first reservoir 130. In some embodiments, at least one microfluidic tubular channel fluidly connects second microfluidic chamber 112 and second microfluidic reservoir 132. In some embodiments, conduit 106 is configured to traverse through a blood vessel. In some embodiments, the conduit 106 comprises individual connections between the stent system 100 (e g., microfluidic chambers, transducers , sensors, etc.) and traverses through a vessel and out to a subcutaneous surgical implant on the body of the subject (e.g., the thorax).
  • the stent system 100 e g., microfluidic chambers, transducers , sensors, etc.
  • conduit 106 may be formed or comprise any suitable material, including but not limited to: flexible materials, synthetics, polymers, elastic materials, metal (e.g., stainless steel, nitinol, etc.), or an extruded polymer (e.g., PTFE, polydimethylsiloxane, silicone, etc.) or the like.
  • suitable material including but not limited to: flexible materials, synthetics, polymers, elastic materials, metal (e.g., stainless steel, nitinol, etc.), or an extruded polymer (e.g., PTFE, polydimethylsiloxane, silicone, etc.) or the like.
  • conduit 106 may comprise any material as would be known by one of ordinary level of skill in the art.
  • the at least one conduit 106 has a length ranging between 1 mm and 200 cm, and a diameter ranging between 0.01 mm and 3 mm. In some embodiments, conduit 106 has a length ranging between 1 mm and 1000 mm, and a width or diameter ranging between 0.01 mm and 2 mm. In some embodiments, conduit 106 comprises at least one lumen, wherein the lumen size ranges between about 0.01 mm and about 1 mm, or about 0.1 mm to about 0.35 mm.
  • aspects of the present invention relate to a method of fabricating a stent system comprising fabricating a stent, fabricating at least one PZT transducer, fabricating at least one microfluidic chamber, fixing the at least one PZT transducer to the stent, disposing the at least one microfluidic chamber within the stent.
  • the step of fabricating the PZT transducer comprises 3D printing a PZT base layer on a polymer support layer, debonding the PZT base layer from the polymer support layer, sintering the PTZ base layer in a PTZ powder bed with heat to form a PTZ element, submerging the PTZ element into an isolation liquid, polarizing the dipoles within the sintered PTZ element, applying an electric field to the PTZ element and cooling the PTZ element.
  • an exemplary method of fabricating a stent system comprises the steps of fabricating a stent, fabricating at least one PZT transducer, fabricating at least one microfluidic chamber, fixing the at least one PZT transducer to the stent, disposing the at least one microfluidic chamber within the stent.
  • the step of fabricating the PTZ transducer comprises 3D printing a PZT base layer on a polymer support layer, debonding the PZT base layer from the polymer support layer, sintering the PTZ base layer in a PTZ powder bed with heat to form a PTZ element, submerging the PTZ element into an isolation liquid, polarizing the dipoles within the sintered PTZ element, applying an electric field to the PTZ element; and cooling the PTZ element.
  • aspects of the present invention relate to spatiotemporal expression or control of genes (e.g., intrinsic or exogenous genes) with at least one therapy (e.g., targeted acoustic stimulation) of cells transfected with select Transient Receptor Potential channels using the disclosed stent system 100.
  • stent system 100 enhances precision treatment of neurologic diseases and disorders by activating, silencing, or modulating gene function by activating, silencing, or modulating gene function.
  • aspects of the present invention relate to a method for spatiotemporal expression or control of intrinsic or exogenous genes comprising providing a stent system with transducers (e g., system 100), implanting the stent system near or in a target site of a subject, providing power at one or more voltages in one or more frequencies from the power source to the at least one transducer for at least one duration.
  • the one or voltages is between 0.1 V to 500 V
  • the one or more frequencies are between 0.1 Hz and 10000 MHz
  • the at least one duration is between 0.1 ps and 1 hour.
  • aspects of the present invention relate to a drug delivery method comprising providing a stent system with at least one transducer and at least one microfluidic chamber (e.g, system 100), implanting the stent system in or near a target site of a subject, fdling the at least one microfluidic chamber at least partially with a solution comprising at least one therapeutic and microbubbles, producing microbubble cavitation via the at least one transducer and delivering the at least one therapeutic to the target site.
  • a drug delivery method comprising providing a stent system with at least one transducer and at least one microfluidic chamber (e.g, system 100), implanting the stent system in or near a target site of a subject, fdling the at least one microfluidic chamber at least partially with a solution comprising at least one therapeutic and microbubbles, producing microbubble cavitation via the at least one transducer and delivering the at least one therapeutic to the target site.
  • an exemplary method of use comprises a drug deliver method.
  • an exemplary drug delivery method comprises the steps of providing a system 100 of the present invention, supplying at least one therapeutic and microbubbles via at least one microfluidic chamber, producing microbubble cavitation via the transducer and delivering the at least one therapeutic to a target site.
  • system 100 is implanted in the superior sagittal sinus of a subject, and at least a portion of the system percutaneously exits the subject to a connector block or controller for power, programing and fluid introduction.
  • system 100 is an advanced function stent system that enables transient blood brain barrier disruption and direct drug delivery into the central nervous system.
  • aspects of the present invention relates to an implantation method for an advanced stent system.
  • a 2.1 - 7 French catheter is used to implant at least a portion of system 100 at a target site.
  • a long sheath, a guide catheter, an intermediate catheter, a microcatheter, and/or a microguidewire is used to implant at least one of portion of system 100 at a target site.
  • the target site is reached via percutaneous access via the jugular, brachial, femoral, or subclavian veins or via the femoral, radial, carotid, or brachial arteries to the intracranial vessels (i.e., Superior sagittal sinus, straight sinus, great vein of galen, transverse sinus, sigmoid sinus, inferior petrosal sinus, internal cerebral vein, basal vein, vein of trolard, vein of Labbe, or the internal cerebral artery, middle cerebral artery, or anterior cerebral artery, respectively (or any of its branches with adequate diameters, generally > 1.5 mm) within the intracranial vault.
  • intracranial vessels i.e., Superior sagittal sinus, straight sinus, great vein of galen, transverse sinus, sigmoid sinus, inferior petrosal sinus, internal cerebral vein, basal vein, vein of trolard, vein of Labbe, or the internal cerebral artery, middle cerebral artery, or anterior cerebral artery, respectively (or any of its branches with adequate diameter
  • a method of using system 100 comprises the steps of providing any system 100 of the present invention, implanting stent system 100 into a subject, supplying microbubble and at least one drug to system 100, and insonating the microbubbles.
  • system 100 comprises a computing system (e.g., computer 400) communicatively connected to the transducers of system 100, comprising a processor and a non-transitory computer-readable medium with instructions stored thereon, which when executed by a processor, perform steps comprising applying an electrical current from the power source to the transducer for a predetermined period of time, or one or more durations.
  • a computing system e.g., computer 400
  • a processor e.g., computer 400
  • non-transitory computer-readable medium with instructions stored thereon which when executed by a processor, perform steps comprising applying an electrical current from the power source to the transducer for a predetermined period of time, or one or more durations.
  • system 100 is used for on-demand transient blood brain barrier disruption, which may modulate cellular microenvironments. In some embodiments, system 100 is used adjunctively for enhanced drug delivery with or without microbubbles or some other carrier. In some embodiments, system 100 is used for sonogenetic control of cells. In some embodiments, system 100 is used for acoustic modulation of cell function. In some embodiments, system 100 is used for sonogenetic control of cells transfected with TRP channels. In some embodiments, system 100 is used for acoustic wave mediated neuromodulation of brain circuits.
  • the delivery location is the intracranial cerebral arteries such was the internal carotid artery, basilar artery, vertebral artery, middle cerebral artery, posterior cerebral artery, or anterior cerebral artery.
  • the venous structures may be the implantation site, such as the superior sagittal sinus, the straight sinus, the inferior sagittal sinus, transverse sinus, sigmoid sinus, or superior petrosal sinus, or inferior petrosal sinus.
  • software executing the instructions provided herein may be stored on a non-transitory computer-readable medium, wherein the software performs some or all of the steps of the present invention when executed on a processor.
  • aspects of the invention relate to algorithms executed in computer software.
  • Parts of this invention are described as software running on a computing device. Though software described herein may be disclosed as operating on one particular computing device (e.g. a dedicated server or a workstation), it is understood in the art that software is intrinsically portable and that most software running on a dedicated server may also be run, for the purposes of the present invention, on any of a wide range of devices including desktop or mobile devices, laptops, tablets, smartphones, watches, wearable electronics or other wireless digital/cellular phones, televisions, cloud instances, embedded microcontrollers, thin client devices, or any other suitable computing device known in the art.
  • a dedicated server e.g. a workstation
  • software e.g. a dedicated server or a workstation
  • software is intrinsically portable and that most software running on a dedicated server may also be run, for the purposes of the present invention, on any of a wide range of devices including desktop or mobile devices, laptops, tablets, smartphones, watches, wearable electronics or other wireless digital/cellular phones, televisions, cloud instances, embedded microcontrollers, thin client
  • parts of this invention are described as communicating over a variety of wireless or wired computer networks.
  • the words “network”, “networked”, and “networking” are understood to encompass wired Ethernet, fiber optic connections, wireless connections including any of the various 802.11 standards, cellular WAN infrastructures such as 3G, 4G/LTE, or 5G networks, Bluetooth®, Bluetooth® Low Energy (BLE) or Zigbee® communication links, or any other method by which one electronic device is capable of communicating with another.
  • elements of the networked portion of the invention may be implemented over a Virtual Private Network (VPN).
  • VPN Virtual Private Network
  • FIG. 4 and the following discussion are intended to provide a brief, general description of a suitable computing environment in which the invention may be implemented. While the invention is described above in the general context of program modules that execute in conjunction with an application program that runs on an operating system on a computer, those skilled in the art will recognize that the invention may also be implemented in combination with other program modules.
  • program modules include routines, programs, components, data structures, and other types of structures that perform particular tasks or implement particular abstract data types.
  • program modules may be located in both local and remote memory storage devices.
  • Fig. 4 depicts an illustrative computer architecture for a computer 400 for practicing the various embodiments of the invention.
  • the computer architecture shown in Fig. 4 illustrates a conventional personal computer, including a central processing unit 450 (“CPU”), a system memory 405, including a random access memory 410 (“RAM”) and a read-only memory (“ROM”) 415, and a system bus 435 that couples the system memory 405 to the CPU 450.
  • the computer 400 further includes a storage device 420 for storing an operating system 425, application/program 430, and data.
  • the storage device 420 is connected to the CPU 450 through a storage controller (not shown) connected to the bus 435.
  • the storage device 420 and its associated computer- readable media provide non-volatile storage for the computer 400.
  • computer-readable media can be any available media that can be accessed by the computer 400.
  • Computer-readable media may comprise computer storage media.
  • Computer storage media includes volatile and non-volatile, removable and non-removable media implemented in any method or technology for storage of information such as computer-readable instructions, data structures, program modules or other data.
  • Computer storage media includes, but is not limited to, RAM, ROM, EPROM, EEPROM, flash memory or other solid state memory technology, CD-ROM, DVD, or other optical storage, magnetic cassettes, magnetic tape, magnetic disk storage or other magnetic storage devices, or any other medium which can be used to store the desired information and which can be accessed by the computer.
  • the computer 400 may operate in a networked environment using logical connections to remote computers through a network 440, such as TCP/IP network such as the Internet or an intranet.
  • the computer 400 may connect to the network 440 through a network interface unit 445 connected to the bus 435.
  • the network interface unit 445 may also be utilized to connect to other types of networks and remote computer systems.
  • the computer 400 may also include an input/output controller 455 for receiving and processing input from a number of input/output devices 460, including a keyboard, a mouse, a touchscreen, a camera, a microphone, a controller, a joystick, or other type of input device. Similarly, the input/output controller 455 may provide output to a display screen, a printer, a speaker, or other type of output device.
  • the computer 400 can connect to the input/output device 460 via a wired connection including, but not limited to, fiber optic, Ethernet, or copper wire or wireless means including, but not limited to, Wi-Fi, Bluetooth, Near-Field Communication (NFC), infrared, or other suitable wired or wireless connections.
  • a wired connection including, but not limited to, fiber optic, Ethernet, or copper wire or wireless means including, but not limited to, Wi-Fi, Bluetooth, Near-Field Communication (NFC), infrared, or other suitable wired or wireless connections.
  • a number of program modules and data files may be stored in the storage device 420 and/or RAM 410 of the computer 400, including an operating system 425 suitable for controlling the operation of a networked computer.
  • the storage device 420 and RAM 410 may also store one or more applications/programs 430.
  • the storage device 420 and RAM 410 may store an application/program 430 for providing a variety of functionalities to a user.
  • the application/program 430 may comprise many types of programs such as a word processing application, a spreadsheet application, a desktop publishing application, a database application, a gaming application, internet browsing application, electronic mail application, messaging application, and the like.
  • the application/program 430 comprises a multiple functionality software application for providing word processing functionality, slide presentation functionality, spreadsheet functionality, database functionality and the like.
  • the computer 400 in some embodiments can include a variety of sensors 465 for monitoring the environment surrounding and the environment internal to the computer 400.
  • sensors 465 can include a Global Positioning System (GPS) sensor, a photosensitive sensor, a gyroscope, a magnetometer, thermometer, a proximity sensor, an accelerometer, a microphone, biometric sensor, barometer, humidity sensor, radiation sensor, or any other suitable sensor.
  • GPS Global Positioning System
  • the blood brain barrier is a major impediment to drug therapy into the central nervous system. Pulsed Ultrasound in conjunction with intravenous ultrasound transiently disrupts the blood brain barrier. The skull remains a significant barrier to effective ultrasound energy delivery.
  • a device referred to as the SonoCloud has been developed for neuro-oncologic clinical indications and is currently undergoing clinical trials, but it requires drilling a hole into the skull. The invasiveness of the device limits its generalizability for the treatment of a variety of other neurological disorders.
  • the use of sputtered Niti Thin film as a scaffold for sophisticated microelectronic mechanical systems has been demonstrated to be a versatile method of maintaining the structural integrity of MEMS components embedded on a stent structure.
  • Fig. 5A, Fig. 5B, and Fig. 5C show an exemplary design and additive manufacturing of fully packaged freeform ultrasonic transducers capable of traveling inside millimeter-sized channels and the delivery of ultrasound energy to precise locations within submm 2 and achieve cavitation-enabled drug delivery. This is achieved through a novel integrated printing and sintering technique that achieves dense 3D printable piezoceramic elements with properties matching pristine ceramics.
  • IVUS intravascular ultrasound
  • 3D printing functional materials requires the introduction of active material inclusions and binders, such as UV sensitive monomers or inks as the matrix material where functional feedstock is embedded in. This results in a significant trade-off between processability and functionality.
  • the disclosed 3D printing and liquid phase sintering method results in properties matching that of the fully dense pristine ceramics, making 3D printing a viable path for high-performance functional materials.
  • Second, a design and printing library that enables the selection and 3D assembly of active, backing and matching layers into focusing ultrasonic transducer architectures for a myriad of medium for maximum acoustic energy output.
  • the novel 3D printable transducer technology enables localized energy delivery in the sub-mm domain. The energy density is capable of reaching the threshold of triggering cavitations via bursting microbubbles, paving the way for precision drug delivery, therapeutics and diagnostics.
  • the manufacturing of these architectures is either dependent on conventional machining methods, including etching, dicing and hot pressing [10], or limited to 3D printed composite materials containing piezoelectric nanoparticles and polymer matrices [15],
  • the mechanical stress caused by the machining processes results in grain pullout, reduced strength and depolarization, leading to significant degradation of the piezoelectricity of the manufactured elements [12],
  • the piezoelectric materials with these architectures exhibit piezoelectric coefficients much lower than those of their pristine ceramic counterparts and weak emission pressures when used as ultrasonic transducers.
  • the viscosity of the highly particle-loaded composite colloids makes printing a uniform layer difficult, and the light scattering effect [28] of the particles induces broadening of the printed features, making printing precise features challenging.
  • the as- fabricated composites (“green parts”) are sintered at high temperature to bum off the polymer matrices and regrow the ceramic grains.
  • piezoceramics including sodium potassium niobate (KNN) [29] and barium titanate (BTO) [30] are most commonly used for 3D printing
  • the conventional sintering process induces deformation, cracks, and high porosity in the 3D printed elements, resulting in degraded mechanical and piezoelectric properties.
  • an SLA-based AM method with an optimized post-sintering process for printing dense lead zirconate titanate (PZT) elements and packaged transducers with microscale features and a high piezoelectric response starts with high-resolution projection micro- stereolithography (PpSL) combined with a tape-casting recoating process to ensure accurate control of the features of the green parts.
  • PpSL projection micro- stereolithography
  • a liquid phase sintering method compatible with 3D printed parts was developed and a liquid sealing method to suppress lead evaporation at high temperature, which reduces the porosity and lead loss of 3D printed piezoceramics.
  • the as-fabricated PZT elements reached a piezoelectric strain constant and an electromechanical coupling factor of up to 583 pC/N, corresponding to 92.5% of that of the pristine material, outperforming currently printable piezoelectric materials.
  • printable packaging material pallets were developed that include a backing layer, an impedance matching layer and physical housing with a large range of tailorable impedances and attenuation coefficients to ensure optimal target performance for targeted applications.
  • IVUS intravascular ultrasound
  • microscale focusing features which is capable of generating high and localized acoustic pressure in blood vessels with diameters as low as 2 mm, allowing localized cavitation triggering and intravenous drug delivery.
  • the 3D printable microscale features together with the high piezoelectric properties of the printed piezoceramic materials will make a great leap forward toward new applications of 3D transducers.
  • the viscosity of the highly particle-loaded composite colloids makes printing a uniform layer difficult, and the light scattering effect [28] of the particles induces broadening of the printed features, making printing precise features challenging.
  • the as- fabricated composites (“green parts”) are sintered at high temperature to bum off the polymer matrices and regrow the ceramic grains.
  • piezoceramics including sodium potassium niobate (KNN) [29] and barium titanate (BTO) [30] are most commonly used for 3D printing
  • the conventional sintering process induces deformation, cracks, and high porosity in the 3D printed elements, resulting in degraded mechanical and piezoelectric properties.
  • the process starts with high-resolution projection micro- stereolithography (PpSL) combined with a tape-casting recoating process to ensure accurate control of the features of the green parts.
  • PPS liquid phase sintering
  • LPS liquid phase sintering
  • a liquid sealing method [32] was developed to suppress lead evaporation at high temperature, which reduces the porosity and lead loss of sintered elements.
  • the as-fabricated PZT elements reached a piezoelectric charge constant and an electromechanical coupling factor of up to 583 pC/N, corresponding to 92.5% of that of the pristine material, outperforming currently printable piezoelectric materials.
  • printable packaging material pallets were developed that include a backing layer, an impedance matching layer and physical housing with a large range of tailorable impedances and attenuation coefficients to ensure optimal target performance for targeted applications.
  • a miniaturized ultrasound transducer was printed with microscale focusing features, which is capable of generating high and localized acoustic pressure in blood vessels with diameters as low as 2 mm, allowing localized cavitation triggering, enhanced drug delivery, and ultrasonic modulation of cellular activity.
  • the 3D printable microscale features together with the high piezoelectric properties of the printed piezoceramic materials will make a great leap forward toward new applications of 3D printed transducers.
  • 3D printing of a highly loaded PZT composite and liquid phase sintering The free-form fabrication achievable with 3D printing technology allows one to exploit acoustic pressure and focusing capabilities with ranges and resolutions that are not feasible with conventional piezoelectric transducers comprising flat piezoelectric elements.
  • the miniatured ultrasound transducer is widely used for blood vessel disease diagnosis and treatment.
  • currently existing miniatured ultrasound transducers have limited focusing or acoustic energy emitting capability [33-36] due to the inability to fabricate high-resolution focusing features.
  • the generated acoustic beam of the ultrasound transducers was manipulated by tuning the PZT element curvature. This curved shape leverages the ability to focus emitted acoustic energy along one line to improve the element sensitivity and acoustic beam resolution, as shown in Fig. 6A.
  • the disclosed fabrication method starts with UV-sensitive piezoelectric preparation.
  • the highly particle-loaded slurry comprises 15-45 vol% PZT nanoparticles (APC 855), 5-15 vol% liquid phase sintering (LPS) agent, lead nitrate (PbNO3) and a UV-sensitive resin made from a mixture of monomers (polyethylene glycol diacrylate, PEGDA) and a photoinitiator, as shown in Fig. 6B.
  • APC 855 features a high piezoelectric charge constant (630 pC/N), which ensures the high performance of the printed part.
  • LPS has been widely used in piezo ceramic fabrication [37,38] and has been implemented in powder-based 3D printing techniques of piezoelectric materials [39],
  • incorporating LPS for light-based additive manufacturing is challenging as most sintering additives (lead oxide [38] and CdBiB [40]) introduce additional light absorbance and increases viscosity, thus decrease the slurry printability.
  • lead nitrate was used as the LPS agent due to its low UV light absorption, ensuring the high printability of the slurry.
  • High PZT loading in the slurry could improve the piezoelectric properties but limited the printability.
  • the combination of 35 vol% PZT and 5 vol% lead nitrates was chosen, which was experimentally optimized to guarantee the printability and piezoelectrical properties.
  • the slurry was fed to the resin tray of a custom-made PpSL system for 3D printing.
  • a printing window (2 mm) made of polydimethylsiloxane (PDMS) was utilized during the process of separating the printed part and the printing window, which prevents small features from being damaged.
  • the oxygen permeability of the PDMS window leads to an oxygen inhibition layer between the printing window and resin that increases the distance between the printed part and printing window and reduces the suction force [41],
  • FEP fluorinated ethylene propylene
  • the as-printed piezocomposites were then sintered to grow the PZT grain size and form the dense PZT ceramic with designed geometries.
  • the printed sample was debonded using a two-step debonding process (Fig. 11) to remove the supportive polymer, as shown in Fig. 6D.
  • the two-step debonding process reduced the deformation of the printed elements compared to the currently existing direct debonding process, allowing fabrication of small-scale features.
  • lead nitrate within the elements decomposes into lead oxide at around 470 °C.
  • extra PZT powders were placed on the bottom of an alumina crucible, and the printed element was placed on top of the powder bed.
  • a smaller crucible was used to cover the element and pushed down into the powder bed.
  • the temperature was then increased to 1100 °C in 1.5 hours and held for 3 hours to grow the PZT grains and form a dense ceramic (see Fig. 12A and Fig. 12B).
  • LPS started at an elevated temperature.
  • Lead oxide melted at 888 °C and contracted the PZT particles via surface tension, densifying the PZT architecture, as shown in Fig. 6E.
  • the lead oxide inside the element also serves as a lead source during sintering to prevent lead loss [45],
  • the liquid sealing method was implemented by placing lead oxide along the edge of the small crucible (see Fig. 6F) to further reduce the lead loss at high temperatures.
  • the temperature and electric field profile of the polarization process was optimized (see Fig. 13C). After reaching the peak value, the electric field was maintained during the cooling process of the isolation liquid to prevent depolarization of the printed element.
  • the optimized sintering procedure ensured the formation of a dense ceramic with low porosity and well-maintained lead content, leading to high piezoelectric charge constant and coupling factor.
  • the acoustic pressure generated by the fabricated 9.75-MHz transducer can be effectively increased by using the curvature design of the elements (Fig. 6G).
  • the acoustic pressure is tunable according to the change in the element curvature degree and can exceed some medical thresholds [46-50] at 9.75 MHz, enlarging the range of applications of the fabricated device.
  • Piezoelectric performance of the as-fabricated piezoelectric elements The technique allows the fabrication of high-performance piezoelectric transducers with complex 3D geometries that are not achievable by any conventional fabrication methods for piezoceramics, including hot-pressing, molding, sanding, and dicing. These complex shapes allow for potential custom transducer applications.
  • Fig. 7A shows curved elements with designed curvatures for potential miniatured transducer applications.
  • Other examples include hemisphere elements (see Fig. 7B) for medical imaging and nondestructive testing, cylindrical transducer elements with multiple concentric annular layers fdled with a polymer (see Fig. 7C) for reducing the traverse vibration and enhancing the thickness vibration, helical elements (see Fig. 7D) for generating acoustic beams with vortex motion for ultrasound manipulation [9], and architected piezo sensors (see Fig. 7E, Fig. 7F) for underwater sensing.
  • Fig. 7G shows the SEM image comparison of an element between the conventional sintering method and the optimized sintering method, confirming the low porosity of the as-fabricated elements.
  • the density of the sintered elements reached 7.2g/cm 3 , which is 94.7% of the pristine material (APC 855).
  • the porosity of the elements 5.3%, is calculated from the density of the printed samples.
  • the sound speed within both the 3D printed element and its pristine material counterpart were measured within multiple PZT plates with the same cross section area (5 x 5 mm) and different thicknesses to verify the low porosity of samples with different volumes.
  • the sound traveling time in each element was captured to calculate the sound speed (Fig. 15). As shown in Fig. 7H, the sound speed in the 3D printed samples reaches approximately 93.2% of the speed in the pristine material, which is consistent with the density measurements.
  • the piezoelectric properties (hysteresis loops, piezoelectric charge constant, kt) of the as-fabricated PZT elements were measured to compare with the pristine materials.
  • Fig. 7H plots the hysteresis of the polarization of the PZT element measured by precision multiferroic analyzer (Radiant Technology. Inc, USA), showing a remnant polarization (P r ) as high as 32.4pC/cm 2 .
  • the strain-electric field loops of the 3D printed and the pristine elements were captured (see Fig. 16A and Fig. 16B), showing that the 3D printed sample can generate near- pristine strain under the same electric field.
  • the piezoelectric charge constant, d33 was measured using a d.33 meter (YE2730A, APC Piezo, USA) and reached 583 pC/N, 92.5% of the value of the pristine material.
  • the electromechanical coupling factor in the thickness direction, kt was measured as 0.57 (see Fig. 19).
  • Fig. 7J compares the d33 and kt of the 3D printed PZT with those of the state-of-the-art 3D printed piezoceramics [11, 51-55],
  • the acoustic signal resolution and energy output of ultrasonic transducers highly depend on the packaging due to the mismatch of acoustic properties of the ceramic transducer element and transmission medium. To achieve optimal resolution and acoustic energy output, the acoustic properties of the matching layer and backing layer of the transducer package need to be optimized for target applications. In addition, the microscale features and complex shapes of a 3D printed transducer require conformal packaging to ensure tight contact with and encapsulation with the transducer element.
  • 3D printable composite material pallets were developed with acoustic impedance ranging from 3 Mrayl to 6.7 Mrayl and attenuation coefficient ranging from -4.3 dB/mm to -10 dB/mm, as shown in Fig. 8A.
  • the acoustic properties of the materials can be tuned by changing the ratio of the resin composites.
  • the matching layer composite featuring low attenuation coefficient and a large range of acoustic impedances, was used as the matching layer, while a backing layer composite, featuring a large range of attenuation coefficients, was used as the backing layer.
  • the large tunable range of impedances and attenuation coefficients allows us to fabricate transducers with optimal performances for targeted applications.
  • Fig. 8A demonstrates the optimal matching layer composition when the transducer is used in different tissues, including blood, muscle, fat, and bone.
  • the selection of the backing layer depends on the applications. For acoustic imaging, as an example, the backing layer requires high attenuation to prevent the ringing effect [57], For high acoustic pressure output, a large impedance mismatch between the backing layer and the transducer is required to maximize the bounce-back pressure [6], Moreover, during in vivo operation, the encapsulation material of the transducer is directly in contact with the tissue and requires biocompatibility.
  • TMPTA trimethylolpropane triacrylate
  • the bold curves with a curvature equal to zero denote the flat piezoelectric transducers with limited acoustic pressure outputs, while the surface plots denote the output pressure range enabled by the 3D printed curved elements.
  • the focusing feature enabled the transducer to minimize its focal zone by 78.6 %, as shown in the subfigure in Fig. 8B.
  • the results show that the miniaturized ultrasound transducer with 3D printed focusing features reaches higher pressure and lateral resolution than the reported ultrasound transducers (see Fig. 17A through Fig. 17F).
  • the 3D printed focusing elements reach the FDA defined pressure limit [50] and many other medical threshold [46-49] at 9.75 MHz.
  • the highly localized acoustic pressure may open the door for applications for targeted ultrasonic therapy, such as noncontact cellular modulation, localized drug delivery, and transient blood brain barrier opening in small areas with length scales under 2 mm.
  • Fig. 8C The schematic (Fig. 8C) of the miniaturized ultrasound transducer shows the package component assembly used for experimental verification.
  • the optimal matching layer CF (28wt%) - matching layer composite
  • backing layer Fe (3wt%) - backing layer composite
  • Fig. 8D shows an optical image of the 3D printed miniaturized ultrasound transducer. Two copper cables (300 mm length, 0.5 mm diameter) were used to connect the device electrodes with the power supply and allow the transducer to be placed in veins.
  • the as-fabricated piezo element was designed with a resonant frequency of 9.75 MHz, which is ideally suited for high resolution focusing and high output pressure in the usual range of the medical ultrasound for procedure guidance [59, 60], 9.75-MHz burst waves were used to drive the miniaturized ultrasound transducer.
  • the output pressures of the transducer with different voltage inputs were captured by a commercial hydrophone (HGL-0400, OND) in DI water and are plotted in Fig. 8B (with “X” shaped dots), which well match the simulated results.
  • Underwater cavitation generated by the as-fabricated transducer The highly localized acoustic output of the 3D printed transducer allows us to exploit its applications for localized cavitation.
  • Cavitation refers to the process in which vapor- or gas-filled cavities formation and undergo oscillation or growth and implosion in liquids upon exposure to acoustic radiation [61], It can generally be classified into two types: stable cavitation and inertial cavitation.
  • Stable cavitation [62] is identified by sustained small amplitude oscillations of the cavities at about their equilibrium, which induces microstreaming in the liquid.
  • Inertial cavitation is characterized by the process of cavities growth and collapse, which induces microjets, causing pitting on solid surface [61],
  • inertial cavitation disrupts the structure of the carrier vesicle and triggers the release of drugs; the induced microstreaming and microjet also makes cell membranes and capillaries more permeable for drug transportation from the blood, making the cavitation a powerful tool for localized endogenous drug delivery.
  • triggering cavitation usually requires high acoustic pressure.
  • the cavitation and the triggering ultrasound can cause collateral damage, including tissue burning, hemolysis and blood temperature increase [61-64], Additionally, the treatment of diseases such as cancer, degenerative and inflammatory diseases, or thromboembolic diseases requires a high concentration of certain drugs to be delivered that have toxic side effects on other tissues [24] . Precise control and localize cavitation at target tissue sites is desired to minimize those side effects.
  • Using microbubble as the cavities can effectively induce the cavitation, making it more controllable [20]
  • the localized cavitation induced by the microbubble oscillation or fragmentation requires precise delivery of acoustic pressure in a controlled manner within a confined volume as small as hundreds of microns, which cannot be achieved by conventional miniatured ultrasonic elements without focusing features.
  • the piezo element has a curvature of 120 deg and an arch length of 2.3 mm, yielding a focal depth of 0.9 mm.
  • the center frequency of the fabricated transducer is 9.75 MHz, as shown in the pulse-echo signal frequency spectrum (see Fig. 20).
  • microbubbles (Lumason, BRACCO, USA) were diluted in a 5 ml 0.9% sodium chloride injection and burst by the 3D printed transducer driving by 9.75-MHz burst waves with 50 cycles.
  • the input voltage of the transducer was adjusted from 15 V PP to 180 V PP and captured the ultrasound signal with a cavitation detector (Sonic Concepts Inc., Bothell, WA, USA) in a large water tank (free field), as shown in Fig. 9A and Fig. 9B.
  • the captured signals were transformed (Fast Fourier Transform) into frequency spectrums to verify cavitation (Fig. 9C, Fig. 9D, and Fig. 9E).
  • microbubble fragmentation was captured by an inverted microscope (1X71, Olympus America Inc., Center Valley, PA, USA). As shown in Fig. 9F, Fig. 24A and Fig. 24B, the microbubbles fragmented and collapsed in the focal area when the transducer was excited by 9.75 MHz 180 V PP pulses. The microbubble behavior during insonation was captured when transducer was driving by 50 V PP pulses with 50 cycles (Fig. 9G through Fig. 9J, Fig. 24A and Fig. 24B).
  • microbubbles experienced oscillation, growth, and collapse, indicating both stable cavitation and inertial cavitation can be generated by the 3D printed transducer [62],
  • the microbubble-induced cavitation shows the capability of the miniatured transducer to assist the drug delivery and diffusion (Cavitation-assisted dye diffusion is shown in Fig. 21, Fig. 22, Fig. 25A, Fig. 25B, and Fig. 25C).
  • a 3D printed transparent blood vessel phantom was prepared with multiple internal channels to mimic the in vivo physical environment and inject the transducer and microbubbles inside it.
  • the transducer was focused on the marked “lesion” location.
  • the microbubble suspension (5 mg/ml) was injected into the blood vessel phantom (see Fig. 10B).
  • pulses were applied to the transducer to generate acoustic pressure.
  • a transparent region appearing in the focal area see Fig. 10B, Fig. 26A and Fig.
  • FIG. 10D plots the bubble content as a function of time captured by digital image correlation (see Fig. 23), showing that the bubbles were fragmented within 0.15 s and 0.08 s for 50 V PP and 90 V PP driving voltages. With a driving voltage lower than the cavitation threshold, no microbubble fragmentation was observed.
  • Fig. 10E bubbles were deposited on the internal wall of the vessel phantom (Fig. 10E) and used the transducer to fragment the bubbles. As shown in Fig. 10E through Fig. 10H, Fig. 27A, and Fig. 27B, as the transducer moved, only the bubbles within the focal area (0.2 mm 2 ) were fragmented, showing high-precision cavitation.
  • the effect of microstructures and curvatures enabled by the printing approach was studied on the focusing resolution and power output and optimized the spatial resolution and acoustic pressure of the output acoustic beam. Additionally, the 3D printed packaging components with tunable acoustic properties adapt the PZT elements to different transmission media and applications, diversifying the versatility of the fabricated ultrasound transducers in medical treatments.
  • the reported processing approach achieves the high energy output application of an ultrasound transducer driven by a 3D printed piezo element. Optimization of the microstructural effect, along with tunable matching and backing layer properties, enables high acoustic pressure confined within sub-mm 2 . Beyond the drug delivery, the localized acoustic energy output could enable future applications including intravascular thrombolysis, in situ imaging, neuromodulation, sonogentic control, and oncology applications.
  • the fabrication process started with the preparation of the UV sensitive piezocomposite colloid comprising PZT 855 (APC Piezo, USA) particles, lead nitrate (MilliporeSigma, USA) and the UV sensitive resin.
  • the UV sensitive resin was comprised of 1 wt% to 3 wt% photoinitiator Irg819 (MilliporeSigma, USA) and PEGDA (MilliporeSigma, USA). 15 vol% to 45 vol% of PZT, 5 vol% to 15 vol% lead nitrate was ball milled together with the UV sensitive resin to form the colloid.
  • the solidified layer was attached to the building platform, or the previous layer and the printing stage was lifted from the printing window, preparing for the coating and curing process of the next layer.
  • the piezoelectric composite with designed 3D geometry was fabricated by repetition of the coating and UV curing processes.
  • the printing resolutions in x-y plane and along z-axis were 20 pm and 15 pm, respectively.
  • the two-step debonding process is shown in Fig. 11A.
  • the debonding temperature profile is shown in Fig. 1 IB.
  • the printed sample was heated under an argon environment. The temperature was increased from 0 °C to 600 °C at a rate of 1 °C/min in 10 hours, and held at 600 °C for three hours, during which the supportive polymer (PEGDA) was carbonized and utilized to maintain the shape of the PZT composite structure. Then, the carbonized polymer was burnt out in the air at 600 °C for 3 hours.
  • the two-step debonding process reduced the deformation of the printed elements compared to the direct debonding process [54], allowing the fabrication of small-scale features.
  • the temperature After debonding, the temperature increased to 1100 °C for grain growth and dense ceramic formation.
  • the peak sintering temperature was optimized according to the d33 testing results and sample breakdown electric field. With a peak sintering temperature higher than 1100 °C, the d33 dropped due to lead loss caused by high temperature (see Fig. 12A). When the peak sintering temperature was lower than 1100 °C, the sample would have a low density, resulting in a low breakdown electric field (see Fig. 12A).
  • Polarization The polarization process was conducted in silicone oil to avoid electric breakdown when the polarization field is over 3 V/pm, which is the breakdown field of the air. Both surfaces of the sample were covered by silver electrode plates wired to the high voltage supply, as shown in Fig. 11 A. The piezoelectric sample was polarized under 6.5 V/pm electric field, resulting in dipole alignment in the direction of the applied field (see Fig. 13B). A combination of an electric field profile and a temperature profile was used to polarize the as- fabricated piezoelectric nanocomposite, as shown in Fig. 13C.
  • Hysteresis loops measurement The P-E loop was measured by Precision Multiferroic Analyzer from Radiant technology. Oil drops were applied to the sample to prevent electric breakdown.
  • the S-E loop measurement was performed by Thin Film Piezoelectric Test Bundles (Radiant technologies, Inc, USA). The testing bundle comprises a precision materials analyzer (Precision Multiferroic II, Radiation technologies, Inc) and a laser vibrometer (VibroOne, Polytec, Inc. USA) (Fig. 16). The tested samples had the same dimensions (2.5 x 2 x 0.25 mm). In S-E loop testing, the samples were polarized under 6.5 V/pm electric fields using the same poling method described in the manuscript.
  • Electromechanical coupling factor The electromechanical coupling factor kt was calculated using resonance frequency f r and anti-resonance frequency f a [68] via:
  • Equation 2 f r and f a were extracted from the impedance curve measured using an E4991B impedance analyzer (Key sight Technology, USA).
  • the spectrum of impedance of the as- fabricated elements (2.1x1.9x0.15 mm) is shown in Fig. 19.
  • the resonant frequency and anti- resonant frequency were 8.27 MHz and 9.77 MHz, respectively, leading to an electromechanical coupling factor of 57.2%.
  • Equation 3 where p nega tive is the peak negative pressure and f is the ultrasound wave frequency.
  • Ultrasound setup A function generator (SDG1025, SIGLENT, China) was connected to a radio frequency power amplifier (100A250A, Amplifier Research, USA) to drive the transducer. 9.75-MHz burst waves with 50 cycles (20 ms burst period) were input into the transducer. For localized cavitation experiments in blood vessel phantom, the 9.75-MHz burst waves were set up to have 10% duty cycle.
  • Miniaturized ultrasound transducer pulse-echo signal The pulse-echo profile of the as-fabricated transducer with optimized packaging is shown in the Fig. 20. JSR pulser receiver (DPR300, JSR Ultrasonics, USA) was used to drive the transducer and receive the echo signal. The frequency spectrum indicates that the transducer has the largest output at frequency is 9.75 MHz. The -6 dB bandwidth of the signal is 57%, which is also suitable for imaging application [69],
  • the 3D printable composite material pallets contain two resin composite systems for matching and backing layer of a transducer.
  • the loaded fiber/particle were sonicated and uniformly mixed with the UV curable resin using ball mill.
  • the acoustic properties (acoustic impedance, attenuation coefficient) of the cured composite can be tuned by adjusting the loading of fibers in the composites.
  • the 3D printable composite materials were printed using the same custom PpSL fabrication system as the PZT fabrication.
  • the acoustic simulations were conducted using COMSOL Multiphysics 6.0.
  • the transmission medium was set to water.
  • the boundaries of simulation models were set to perfect matching layer to avoid reflection.
  • the materials’ properties (piezoelectric properties, acoustic impedance and sound speed) in the model were using the tested properties of the samples.
  • Dye diffusion coefficient and microbubble concentration The dye diffusion coefficient and concentration of microbubble in the container were calculated by a selfdeveloped digital image correlation method based on the transparency of the solution.
  • the screenshots of the target location were converted into gray scale image using MATLAB (MathWorks, USA).
  • the mean pixel values of images were used for computing the dye diffusion speed and microbubble concentration change (see Fig. 21, Fig. 22 and Fig. 23).
  • Fabrication resolution of the 3D piezoelectric nanocomposites During the curing process, the resolution of the custom PpSL fabrication system in x-y plane (20x20 pm) is determined by the pixel size of the projector. The resolution along z-axis is determined by the curing depth of colloid. The curing depth of the solid layer is given as [70]:
  • Equation 4 where, a is the resin absorption coefficient, E and E c are the actual and critical exposure, respectively.
  • the layer thickness was set as 15 pm, which is smaller than half of the cure depth to ensure tight bonding between two consecutive layer [71], This optimized cure depth for each layer was modulated by tuning the exposure time [72], A reductive lens was utilized to increase the fabrication resolution of the printed structure [72], The range of achievable 3-3 vibration mode frequency of a 3D printed thin-film sample is determined by the range of thickness that can be achieved, which is around 10.8 pm. This corresponds to the highest achievable frequency on the order of 171 MHz. Higher frequency can be potentially achieved by reducing the minimal layer thickness of the printing apparatus (employing Z-stage with higher resolution).
  • Element shrinkage during sintering and its effects on the structural integrity To evaluate the effects of shrinkage on structural integrity, three types of geometrical parameters were measured before and after sintering, including length change (characterized on flat samples), curvature deviation (measured on concave samples) and strut thickness deviations.
  • 20 rectangular samples (3 x 2.6 x 0.2 mm) were fabricated and the dimension change before and after sintering was measured, as shown in Fig. 14A through Fig. 14B.
  • Fig. 14C shows the deviation distribution of the average linear shrinkage ratio with an average value of 27.73 % and a standard variation of 1.9 %, [50] indicating a uniform shrinkage ratio among samples.
  • FIG. 14E show the representative sample before and after sintering.
  • the average curvature change is within 0.5° (0.51 % of the original curvature) with a standard deviation of 2.2° (see Fig. 14E and Fig. 14F).
  • the resultant acoustic pressure change is within 0.04 Mpa based on Fig. 8D, which was negligible in the experiments.
  • the strut thickness uniformity in 3D architectures was evaluated. X-ray computed tomography was used (CT, Bruker microCT, Model Skyscan 1174) to capture the morphology of a sintered lattice sample containing 84 struts.
  • Fig. 141 shows the deviation distribution of the strut thicknesses. The thickness of the sintered struts is less than 1.58% thicker than the as-designed value (300 pm) with a standard deviation of 6.6 pm, indicating that the structural integrity was well kept after sintering.
  • Equation 5 where, t indicates the thickness of the material to be tested, Vwater is the sound speed in water.
  • the acoustic impedance is calculated by [73]: Z pVmaterial
  • Equation 6 Equation 6 where, p is the density of the material.
  • the transmission ratio, t P between the transmitted wave acoustic pressure and the incident wave acoustic pressure on the interface [74] can be represented by:
  • Equation 7 where the Zi and Zi indicates the acoustic impedance of the two materials.
  • the acoustic pressure p in the material near the interface I and interface II are represented using the transmission ratio t P i2 at interface I and t P 34 at interface II via:
  • the acoustic pressure ratio between top and bottom side within material can be represented using the voltage amplitude V by:
  • Equation 10 where, Ziand Z2 are the acoustic impedance of the medium (water) and the material to be tested. Vi and V4 are the measured voltage generated by ultrasound transducer placed at the top and bottom side of the material. For a given frequency, converting the unit of Equation 10 into (dB/mm) can yield the attenuation coefficient of the material via:
  • Equation 11 Equation 11 where, d is the thickness of the material.
  • Strain - electric field curves of the as-fabricated material For the S-E loop, the strain in the thickness direction was measured (3-3 mode) of the 3D printed material and a standard commercial material against the applied electric fields (S-E curves) using Thin Film Piezoelectric Test Bundles (Radiant technologies, Inc, USA) at room temperature (21.3 °C), as shown in Fig. 16A.
  • the system comprises a precision materials analyzer (Precision Multiferroic II, Radiation Technologies, Inc) and a laser vibrometer (VibroOne, Polytec, Inc. USA).
  • the precision materials analyzer applied voltage long the polarization direction (thickness direction) of the sample and the displacements of the samples were collected by the laser vibrometer and sent back to material analyzer for processing.
  • the tested samples had the same dimensions (2.5 x 2 x 0.25 mm).
  • the S-E loops of the 3D printed sample and commercial sample are plotted and compared in Fig. 16B, showing that the 3D printed sample can generate near-pristine strain under the same electric field.
  • the matching material was formed by mixing carbon fiber (CF) with RIGID resin (Formlab, USA).
  • RIGID resin is the mixture of urethane di methacryl ate and isobomyl Methacrylate.
  • the acoustic impedance of the matching material reached the values from 3 Mrayl to 6.7 Mrayl by tuning the matching layer’s CF loading from 25 wt% to 45 wt%.
  • the CF could increase the composite density and sound speed with a little increment of attenuation coefficient [75],
  • the thickness of the fabricated matching layer was designed as the one-quarter of the wavelengths of the transmitted ultrasound to reduce the acoustic loss [76],
  • the backing material is the Fe magnetic particle-loaded Flexible resin (Formlab, USA). Flexible is the mixture of acrylated oligomers and acrylated monomers.
  • Flexible is the mixture of acrylated oligomers and acrylated monomers.
  • the matching layer is designed based on the shape of the aperture of the transducer.
  • the material selection is determined by the sound propagation medium, as discussed in this disclosure.
  • the thickness of matching layer was controlled to be less than one-fourth of the ultrasound wavelength to minimize the acoustic attenuation [78],
  • the matching layer was 3D printed to fit the size of the sintered piezo element and glue the matching layer with the piezo part using the same photocurable material used for the matching layer.
  • the matching layer was designed as flat membranes and glue it onto the surface that receives sound wave.
  • the lateral resolution of ultrasonic transducers is determined by the geometries and driving frequencies of piezo elements [85],
  • the acoustic intensity field of the transducer was simulated with COMSOL by inputting transducer aperture geometry and working frequency (see Fig. 17A) and verified the result with experiments (see Fig. 17C).
  • the hydrophone HGL-0200, ONDA, USA
  • the resolution is defined by the -6 dB threshold of the acoustic intensity field.
  • the simulated and measured resolutions of our transducer are 0.08 mm and 0.18 mm, respectively.
  • our transducer has the highest lateral resolution owing to its focusing feature and high working frequency.
  • the negative acoustic pressure of the as-fabricated transducers is plotted in Fig. 18D.
  • the disk element can generate the negative acoustic pressure of 0.86 MPa under the voltage of 150 V PP , while the negative acoustic pressure produced by element with concentric layers can be up to 0.99 Mpa.
  • Finite element analysis (COMSOL, USA) was used to show the generated acoustic field of the transducers (Fig. 18E and Fig. 18F).
  • Dye diffusion coefficient and microbubble concentration The transducer achieves cavitation assisted diffusion of molecules.
  • the efficiency was tested in the diffusion of methylene blue, a salt used as a dye and as a medication in aqueous solutions.
  • the experiment was prepared by gently dropping methylene blue into aqueous solution and measuring its diffusion coefficient [19] with and without the assistance of microbubble cavitation at 90 V PP (Fig. 25A, Fig. 25B, and Fig. 25C).
  • the container was divided into six sections. The change of their pixel value standard deviation was used to evaluate the diffusion uniformity (Fig. 21).
  • the ultrasound radiation force-assisted diffusion and cavitation-assisted diffusion can be calculated based on the methylene blue diffusion coefficient in the water [88],
  • methylene blue in both the water and microbubble suspensions started to disperse, as shown in Fig. 22 and Fig. 25A through Fig. 25C.
  • the dispersion in the microbubble suspension rapidly began when the cavitation threshold was reached, and the corresponding standard deviations decreased faster than that in the tap water group.
  • methylene blue in the microbubble suspension was fully dissolved in the solution, and the standard deviation was decreased by 97 % compared to its original value, while the methylene blue in tap water was mainly dispersed in the bottom part with a 21 % drop in the standard deviation.
  • microbubble solution with different concentrations was injected into the transparent blood vessel phantom.
  • the pixel value - microbubble concentration digital imaging correlation (DIC) curve was plotted in Fig. 23.
  • the microbubble fragmentation speeds were achieved by correlating frames of experiment videos at different times and the DIC curve.
  • Frij link M. E., Goertz, D. E., Van Damme, L. C. a., Krams, R. & W. Van Der Steen, A. F. Intravascular ultrasound tissue harmonic imaging in vivo. IEEE Transactions on Ultrasonics, Ferroelectrics, and Frequency Control 53, 1844-1852 (2006).
  • Implantable Biomedical Microsystems eds. Bhunia, S., Majerus, S. J. A. & Sawan, M. 281-312 (William Andrew Publishing, 2015).

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Abstract

Selon certains aspects, un système d'endoprothèse comprend une structure de support possédant de première et seconde extrémités, comprenant un ou plusieurs éléments définissant des régions intérieure et périphérique pour la structure, un réseau interfacé avec la structure de support, au moins un transducteur fixé au réseau, un ou plusieurs conducteurs connectés électriquement au ou aux transducteurs, et une source d'alimentation connectée au ou aux fils conducteurs. Dans d'autres aspects, l'invention concerne un procédé de commande spatio-temporelle ou d'expression de gènes d'une ou de plusieurs cellules.
PCT/US2024/021334 2023-03-23 2024-03-25 Système d'endoprothèse à fonction avancée pour l'administration chronique de médicament à la demande, la commande sonogénétique de cellules et la modulation acoustique de la fonction cellulaire Pending WO2024197304A2 (fr)

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