WO2021252897A2 - Systèmes d'administration d'agents thérapeutiques et leurs procédés de formation et leurs utilisations - Google Patents

Systèmes d'administration d'agents thérapeutiques et leurs procédés de formation et leurs utilisations Download PDF

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WO2021252897A2
WO2021252897A2 PCT/US2021/037010 US2021037010W WO2021252897A2 WO 2021252897 A2 WO2021252897 A2 WO 2021252897A2 US 2021037010 W US2021037010 W US 2021037010W WO 2021252897 A2 WO2021252897 A2 WO 2021252897A2
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lidocaine
hydrogel
nanoparticle
release
therapeutic agent
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PCT/US2021/037010
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WO2021252897A3 (fr
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Jessica Anne Yu ROVE
Dae Won Park
Adam J. ROCKER
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The Regents Of The University Of Colorado, A Body Corporate
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Publication of WO2021252897A2 publication Critical patent/WO2021252897A2/fr
Publication of WO2021252897A3 publication Critical patent/WO2021252897A3/fr
Priority to US18/075,627 priority Critical patent/US20230110354A1/en

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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61K—PREPARATIONS FOR MEDICAL, DENTAL OR TOILETRY PURPOSES
    • A61K9/00—Medicinal preparations characterised by special physical form
    • A61K9/48—Preparations in capsules, e.g. of gelatin, of chocolate
    • A61K9/50—Microcapsules having a gas, liquid or semi-solid filling; Solid microparticles or pellets surrounded by a distinct coating layer, e.g. coated microspheres, coated drug crystals
    • A61K9/51—Nanocapsules; Nanoparticles
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61K—PREPARATIONS FOR MEDICAL, DENTAL OR TOILETRY PURPOSES
    • A61K9/00—Medicinal preparations characterised by special physical form
    • A61K9/10—Dispersions; Emulsions
    • A61K9/107—Emulsions ; Emulsion preconcentrates; Micelles
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61K—PREPARATIONS FOR MEDICAL, DENTAL OR TOILETRY PURPOSES
    • A61K38/00—Medicinal preparations containing peptides

Definitions

  • This invention relates to therapeutic agent delivery systems. Specifically, the invention provides a novel biodegradable micelle-containing hydrogel system and methods of formation and use thereof.
  • hydrogels and nanoparticles have been widely used for controlled drug release in patients, however one drawback includes the release of the drug too quickly.
  • Many nanoparticle-hydrogel delivery systems have been developed, however not all delivery systems can be used to deliver all drugs or treat all diseases. Each delivery system is unique with respect to the composition of the hydrogel, crosslinking of the hydrogel, composition and type of nanoparticle to use with the hydrogel, types of therapeutic agents that can be encapsulated in the particular nanoparticle, stability of the composition, etc.
  • the inventors have developed a novel therapeutic agent delivery system which uses novel nanoparticles embedded within a crosslinked hydrogel to provide sustained release of a therapeutic agent.
  • the use of the novel delivery system can be varied depending on the type of therapeutic agent used.
  • the inventors have expanded the use of the delivery system to deliver growth factors to aid in angiogenesis after myocardial infarction.
  • the delivery system can be coated onto a catheter, such as a chest tube, to sustainably deliver an anesthetic agent.
  • Coronary artery disease is one of the leading causes of death in the United States, faulted for approximately 1 in 7 deaths with 660,000 new coronary attacks each year.
  • CAD Coronary artery disease
  • Ml myocardial infarction
  • the current standard of care for Ml aims for early reperfusion of the occluded vessels to prevent further cell death using surgical or pharmacological agents.
  • biomedical approaches to restore blood supply to ischemic myocardium, via highly localized angiogenesis may present a faster and less invasive Ml treatment option, with the essential benefit of inducing cardiac tissue regeneration.
  • SRTG injectable, sulfonated reverse thermal gel
  • PSHU poly(serinol hexamethylene urea)
  • PNIPAM poly(N- isopropylacrylamide)
  • SPSHU-PNIPAM or SRTG poly(N- isopropylacrylamide)
  • the PSHU backbone is highly functionalizable and has been shown to exhibit a favorable microenvironment for neuronal 6 - 9 and cardiac 10 - 12 tissue engineering applications.
  • the negatively charged sulfonate groups control the release of positively charged proteins by utilizing their electrostatic binding interaction.
  • VEGF vascular endothelial growth factor
  • PDGF Platelet-derived growth factor
  • Interleukin-10 is a potent anti-inflammatory cytokine with broad immunoregulatory activity, and it can suppress many proinflammatory mediators involved in chronic HF. 22 ’ 23 It has been shown that IL-10 treatment in acute Ml rodent models significantly improved left ventricular (LV) function, reduced infarct size, and attenuated infarct wall thinning.
  • IL-10 has a short half-life of less than 5 h in vivo, and it normally requires repeated injections to reach therapeutic efficacy, potentially causing undesirable side effects with a high treatment cost.
  • the SRTG delivery system may provide a localized and sustained release of IL-10 through electrostatic interactions, which could reduce issues associated with high therapeutic doses.
  • Chest tubes are a necessary part of performing intrathoracic operations with the vital function of draining blood and preventing tamponade, evacuating air and preventing pneumothoraces. With over 500 open heart cases performed a year at University of Colorado Hospital (UCH), and an average of 2-3 drains per case, over 1 ,000 chest drains are placed annually for cardiac operations alone. Chest tubes account for a large proportion of pain stimulation after a cardiac or thoracic operation. In studies examining pain and number of chest tubes, fewer chest tubes reduces pain and opioid use after thoracotomy.
  • UCH University of Colorado Hospital
  • the inventors have developed a novel therapeutic agent delivery system that can be used to deliver one or more therapeutic agents to a patient.
  • the novel delivery system is comprised of at least one nanoparticle at least partially encapsulating at least one therapeutic agent.
  • the at least one nanoparticle is embedded within a crosslinked hydrogel.
  • Administration of the delivery system to a patient allows for the therapeutic agent to be sustainably released in the patient.
  • the therapeutic agents that can be used with the novel delivery system include anesthetics, antimicrobials, and growth factors. Depending on the type of therapeutic agent administered, this novel delivery system can be used to treat pain, infection, or even promote angiogenesis and reduce inflammation after myocardial infarction as described below.
  • a sequential protein delivery system comprising of SRTG encapsulating growth factor-loaded polymeric micelle nanoparticles (NPs) has been developed.
  • NPs growth factor-loaded polymeric micelle nanoparticles
  • VEGF and IL-10 were mixed into the aqueous polymer solution before gelling; these proteins should be released according to their binding affinity for sulfonate groups.
  • Novel polymeric micelles were incorporated to provide the sequential and sustained release of PDGF. It was hypothesized that this system will deliver VEGF first, followed by IL-10, and finally PDGF, which may promote optimal cardiac repair after Ml.
  • ischemic myocardium may be protected from necrosis.
  • the subsequent delivery of IL- 10 may repress proinflammatory mediators, which protects the heart from excessive inflammatory injury, while still allowing the important mechanism of clearing dead cells from the infarcted area.
  • the final delivery of PDGF from the encapsulated micelles may stabilize new vessels with pericytes and reduce the incidence of their regression.
  • This specific combination of proteins for therapeutic delivery, targeting the different stages of pathological remodeling following Ml, has not been previously utilized in a potential application for cardiac repair.
  • the inventors evaluated the efficacy of this SRTG micelle delivery system for the spatiotemporal and sequential delivery of VEGF, IL-10, and PDGF for the promotion of mature angiogenic vessels and reducing inflammation.
  • an anesthetic therapeutic agent such as lidocaine
  • a hydrogel may optionally also contain an anesthetic, such as lidocaine.
  • This delivery system may then be administered to the patient by various means including subcutaneous injection or topical application to relieve pain with sustained release of the lidocaine.
  • the novel delivery system can be stably coated onto a catheter, such as a chest tube, to relieve pain associated with placement of a chest tube as described herein.
  • a system for delivery of a therapeutic agent comprising: at least one nanoparticle wherein the at least one nanoparticle is a micelle formed of a polymer comprising a hydrophobic segment and at least one hydrophilic segment; a therapeutic agent encapsulated within the at least one nanoparticle; and a hydrogel wherein the at least one nanoparticle is embedded within the hydrogel.
  • the system allows for sustained release of the therapeutic agent over an extended period of time.
  • the system further comprises an additional amount of the therapeutic agent being embedded within the hydrogel itself to allow for immediate release of the therapeutic agent in addition to the sustained release of the therapeutic agent from the nanoparticles.
  • the hydrophobic segment may be poly(serinol hexamethylene urea) (PSHU) and the at least one hydrophilic segment may be polyethylene glycol (PEG), polycaprolactone (PCL), poly(glycolic acid) (PGA), poly(lactic acid) (PLA), or poly(lactic acid-co-glycolic acid) (PLGA).
  • the hydrogel may be gelatin crosslinked with glutaraldehyde.
  • the therapeutic agent may be selected from anesthetics, antimicrobials, or growth factors.
  • the therapeutic agent is an anesthetic selected from the group consisting of lidocaine, marcaine, bupivacaine, prilocaine, mepivacaine, etidocaine, ropivacaine, and levobupivacaine.
  • the therapeutic agent is at least one growth factor selected from the group consisting of vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF) and interleukin-10 (IL-10).
  • VEGF vascular endothelial growth factor
  • PDGF platelet-derived growth factor
  • IL-10 interleukin-10
  • a method of reducing pain associated with placement of a catheter in a patient in need thereof comprising the steps of: applying an anesthetic-eluting drug delivery system to a surface of a catheter to produce a coated catheter and inserting the coated catheter into the patient.
  • the system allows for sustained release of the anesthetic over an extended period of time to relieve the pain associated with the placement of the catheter.
  • the drug delivery system coated onto the catheter comprises: at least one nanoparticle wherein the at least one nanoparticle is a micelle formed of a polymer comprising a hydrophobic segment, such as poly(serinol hexamethylene urea) (PSHU), and at least one hydrophilic segment, such as polyethylene glycol (PEG), polycaprolactone (PCL), poly(glycolic acid) (PGA), poly(lactic acid) (PLA), or poly(lactic acid-co-glycolic acid) (PLGA); an anesthetic encapsulated within the at least one nanoparticle; and a hydrogel wherein the at least one nanoparticle is embedded within the hydrogel which may be gelatin crosslinked with glutaraldehyde.
  • the drug delivery system may further comprise an amount of the anesthetic being embedded in the hydrogel itself to allow for immediate release in addition to the sustained release of the anesthetic from the nanoparticles.
  • the anesthetic may be selected from the group consisting of lidocaine, marcaine, bupivacaine, prilocaine, mepivacaine, etidocaine, ropivacaine, and levobupivacaine.
  • a method of promoting angiogenesis in patient after myocardial infarction comprising: administering a therapeutic agent delivery system to the patient.
  • the therapeutic agent delivery system comprises: at least one nanoparticle; a therapeutically effective amount of growth factors selected from vascular endothelial growth factor (VEGF), platelet-derived growth factor (PDGF), interleukin-10 (IL-10), or combinations thereof wherein at least one of the growth factors is at least partially encapsulated within the at least one nanoparticle; and a hydrogel wherein the at least one nanoparticle is embedded within the hydrogel.
  • the growth factors not encapsulated within the nanoparticle are embedded in the hydrogel.
  • the delivery system allows for both immediate and sustained release of the combination of growth factors.
  • the PDGF is at least partially encapsulated within the at least one nanoparticle.
  • the combination of growth factors are released sequentially with the VEGF released first, the IL- 10 released second and the PDGF released last.
  • FIG. 1 A-B are a series of images depicting synthesis of the micelles.
  • A Synthesis of PEG-PSHU- PEG by Step 1 : synthesis of PSHU; Step 2: conjugation of PEG to terminal isocyanate groups on PSHU.
  • B 1 H NMR (500 MHz, DMSO-d6) spectrum of PEG-PSHU-PEG.
  • FIG. 2A-E are a series of images of the SRTG and the micelle NPs.
  • A-D SEM images of the SRTG and micelle NPs.
  • E Size distribution of micelles from DLS measurements.
  • FIG. 3 is a graph depicting Cumulative release profile of VEGF, IL-10, and PDGF from the delivery system in vitro. A burst release was observed for VEGF and IL-10 in the first 24 h, with a sustained release following over time. PDGF displayed minimal burst release and slowly released over 28 days, demonstrating the sequential release of all three factors from the SRTG micelle system. Error bars represent standard deviation.
  • FIG. 4A-E are a series of images depicting immunohistochemical assessment of vascularization by vessel formation, endothelial cell count, and functional vascular endothelial cell count.
  • A Representative images show co-staining of CD31 (green), VWF (red), and DAPI (blue) 21 days after subcutaneous injections. Functional vascular cells were characteristic of CD31 + and VWF+. The scale bar represents 100 pm.
  • Quantification of immunohistochemical assessment of vascularization by B) endothelial cell count, (C) functional vascular endothelial cell count, and (D) ratio of functional vascular endothelial cells to total endothelial cells.
  • E Represents the legend for the graphs. Error bars represent standard deviation and * indicates p ⁇ 0.05.
  • FIG. 5A-D are a series of images depicting immunohistochemical assessment of mature vascularization by pericyte cell count.
  • A Representative images show co-staining of CD31 (green), a-SMA (red), and DAPI (blue) 21 days after subcutaneous injections. Pericytes were characteristic of a-SMA+ associated with CD31 +. The scale bar represents 100 pm.
  • Quantification of immunohistochemical assessment of vascularization by B) pericyte cell count and (C) ratio of pericytes to endothelial cells.
  • D Represents the legend for the graphs. Error bars represent standard deviation and * indicates p ⁇ 0.05.
  • FIG. 6A-C are a series of images depicting Immunohistochemical assessment of inflammatory response by macrophage cell count.
  • A Representative images show staining of CD68 (red) and DAPI (blue) 7 and 21 days after injection. Scale bar represents 50 pm.
  • B Quantification of immunohistochemical assessment of inflammatory response by macrophage cell count.
  • C Represents the legend for the graphs. Error bars represent standard deviation and * indicates p ⁇ 0.05.
  • FIG. 8 is an image depicting the delivery system using growth factors and micelles in sulfonated reverse thermal gel (SRTG).
  • SRTG sulfonated reverse thermal gel
  • FIG. 9 is an SEM image of the lidocaine loaded micelles.
  • FIG. 10 is a graph depicting in vitro release profiles of different nanoparticle (NP) formulations loaded with 100 mg of lidocaine.
  • the release profiles from the various NP formulations show the differences in lidocaine release rates are dependent on the amount of NP loaded into the hydrogels.
  • the data display the sustained release of the lidocaine over 10 days for the 20 mg and 10 mg NP groups.
  • the data for last day of each group shows last timepoint where lidocaine could be detected in the release solution.
  • FIG. 11 is a graph depicting in vitro release profiles of different nanoparticle formulations loaded with varied amounts of lidocaine.
  • the release profiles from the various NP formulations show the differences in lidocaine release rates are dependent on the amount of NP and lidocaine loaded into the hydrogels.
  • the data display the sustained release of the lidocaine over 11 days for the 50 mg lidocaine groups.
  • the data for last day of each group shows last timepoint where lidocaine could be detected in the release solution.
  • FIG. 12 is a graph depicting in vitro release profiles of the nanoparticles loaded with varied amounts of lidocaine and embedded in gelatin hydrogels. The data display the sustained release of the lidocaine over eleven days. Error bars represent standard deviation.
  • FIG. 13A-B are a series of images depicting subcutaneous implantation mouse model of catheter tubing with and without gelatin-nanoparticle hydrogel coating. Representative photographs of harvested tissue showing fibrous capsule formation around tubing with and without hydrogel coating 14 days after implantation procedure.
  • the uncoated tubing sample (A) shows similar fibrous tissue formation compared to coated tubing sample (B). This biological tissue response is due to a foreign body reaction from the implanted tubing.
  • FIG. 14 is a graph depicting the lidocaine release profile from the gelatin-lidocaine hydrogel comparing gel mold to coated tubing. Error bars represent standard deviation.
  • FIG. 15 is a graph depicting in vitro release profiles of lidocaine from different gelatin hydrogel formulations.
  • the release profiles from the various gelatin formulations show the differences in lidocaine release rates are dependent on the concentration of gelatin and glutaraldehyde (GA).
  • the data display the sustained release of the lidocaine over 10 days.
  • Concentrations, amounts, solubilities, and other numerical data may be expressed or presented herein in a range format. It is to be understood that such a range format is used merely for convenience and brevity and thus should be interpreted flexibly to include not only the numerical values explicitly recited as the limits of the range, but also to include all the individual numerical values or sub-ranges encompassed within that range as if each numerical value and sub-range is explicitly recited. As an illustration, a numerical range of “about 1 to about 5” should be interpreted to include not only the explicitly recited values of about 1 to about 5, but also include the individual values and sub-ranges within the indicated range.
  • a nanoparticle includes a plurality of nanoparticles, including mixtures thereof.
  • the term “comprising” is intended to mean that the products, compositions and methods include the referenced components or steps, but not excluding others. “Consisting essentially of” when used to define products, compositions and methods, shall mean excluding other components or steps of any essential significance. “Consisting of” shall mean excluding more than trace elements of other components or steps.
  • “Patient” is used to describe an animal, preferably a human, to whom treatment is administered, including prophylactic treatment with the compositions of the present invention.
  • “Pharmaceutically acceptable carrier” means any of the standard pharmaceutically acceptable carriers.
  • the pharmaceutically acceptable carrier can include diluents, adjuvants, and vehicles, as well as implant carriers, and inert, non-toxic solid or liquid fillers, diluents, or encapsulating material that does not react with the active ingredients of the invention. Examples include, but are not limited to, phosphate buffered saline, physiological saline, water, and emulsions, such as oil/water emulsions.
  • the carrier can be a solvent or dispersing medium containing, for example, ethanol, polyol (for example, glycerol, propylene glycol, liquid polyethylene glycol, and the like), suitable mixtures thereof, and vegetable oils.
  • Formulations are described in a number of sources that are well known and readily available to those skilled in the art. For example, Remington’s Pharmaceutical Sciences (Martin EW [1995] Easton Pennsylvania, Mack Publishing Company, 19 th ed.) describes formulations which can be used in connection with the subject invention. In some embodiments, phosphate buffered saline is used as the pharmaceutically acceptable carrier.
  • compositions of the present invention are delivered to the patient for treatment or prevention purposes.
  • the composition can be delivered via subcutaneous injection or topically in dosage unit formulations containing conventional nontoxic pharmaceutically acceptable carriers, adjuvants, and vehicles as desired. Administration may occur once or multiple times.
  • sustained release refers to a composition comprising a therapeutically effective amount of a composition, when administered to a patient, continuously releases a stream of one or more therapeutic agents over a predetermined time period at a level sufficient to achieve a desired effect, such as preventing or treating infections; preventing or treating pain; or treating myocardial infarction by improving cardiac function, promoting angiogenesis or reducing inflammation, throughout the predetermined time period.
  • Reference to a continuous release stream is intended to encompass release that occurs as the result of biodegradation of the composition, or component thereof, or as the result of metabolic transformation or dissolution of the added nutrients or other desired agents.
  • “Therapeutic agent” as used herein is defined as a substance, component or agent that has measurable specified or selective physiological activity when administered to an individual in a therapeutically effective amount.
  • therapeutic agents include growth factors, anesthetics and antimicrobials such as antibiotics, antivirals, antifungals, antiprotozoals, and antiparasitics. At least one therapeutic agent is used in the compositions of the present invention. In some embodiments, multiple therapeutic agents are used and are released in a sequential manner.
  • a “therapeutically effective amount” as used herein is defined as concentrations or amounts of components which are sufficient to effect beneficial or desired clinical results, including, but not limited to, any one or more of treating symptoms of myocardial infarction, infection or pain; or preventing myocardial infarction, infection or pain.
  • Prevention refers to any of: halting the effects of myocardial infarction, infection or pain; reducing the effects of myocardial infarction, infection or pain; reducing the incidence of myocardial infarction, infection or pain; reducing the development of myocardial infarction, infection or pain; delaying the onset of symptoms of myocardial infarction, infection or pain; increasing the time to onset of symptoms of myocardial infarction, infection or pain; and reducing the risk of development of myocardial infarction, infection or pain.
  • Treatment refers to any of the alleviation, amelioration, elimination and/or stabilization of a symptom, as well as delay in progression of a symptom of a particular disorder.
  • treatment may include any one or more of the following: amelioration and/or elimination of one or more symptoms associated with myocardial infarction, infection or pain; reduction of one or more symptoms of myocardial infarction, infection or pain; stabilization of symptoms of myocardial infarction, infection or pain; and delay in progression of one or more symptoms of myocardial infarction, infection or pain.
  • “Infection” as used herein refers to the invasion of one or more microorganisms such as bacteria, viruses, fungi, yeast or parasites in the body of a patient in which they are not normally present.
  • Antimicrobial refers to natural or synthetic compositions capable of killing or inhibiting the growth of microorganisms including, but not limited to, bacteria, fungi, viruses, protozoa and parasites. Antimicrobials used herein include antibiotics, antivirals, antifungals, antiprotozoals, and antiparasitics.
  • Anesthetics refers to a natural or synthetic composition capable of producing a local, regional or general loss of sensation. Anesthetics are generally used to induce an insensitivity to pain. As used herein, the term “anesthetic” is used to refer to local anesthetics. Exemplary anesthetics that may be used herein include, but are not limited to, lidocaine, marcaine, bupivacaine, prilocaine, mepivacaine, etidocaine, ropivacaine, levobupivacaine.
  • Growth factors refers to a secreted biologically active molecule capable of affecting the growth of cells, promote or inhibit mitosis, or affect cellular differentiation.
  • growth factors include angiogenic an immunoregulatory proteins.
  • growth factors examples include, but are not limited to, vascular endothelial growth factor (VEGF); platelet-derived growth factor (PDGF); epidermal growth factor (EGF); granulocyte-macrophage colony-stimulating factor (GM-CSF); insulin-like growth factor 1 (IGF-1 ) and 2 (IGF-2); transforming growth factor (TGFa or TGFp); fibroblast growth factor (FGF); cytokines such as interleukins (IL-1 through IL-18) and interferons (IFN-a, IFN-b, IFN-y).
  • VEGF vascular endothelial growth factor
  • PDGF platelet-derived growth factor
  • EGF epidermal growth factor
  • GM-CSF granulocyte-macrophage colony-stimulating factor
  • IGF-1 insulin-like growth factor 1
  • IGF-2 insulin-like growth factor 1
  • TGFa or TGFp transforming growth factor
  • FGF fibroblast growth factor
  • cytokines
  • Polymer refers to a relatively high molecular weight organic compound, natural or synthetic, whose structure can be represented by a repeated small unit, the monomer. Synthetic polymers are typically formed by addition or condensation polymerization of monomers. Some combinations of polymers form the nanoparticles while other polymers or combinations thereof form the hydrogel. A polymer comprised of two or more different monomers is a copolymer.
  • the polymers used to form the nanoparticles exist in a triblock copolymer having an A-B-A structure with polymer A being selected from the group including, but not limited to, polyethylene glycol (PEG), polycaprolactone (PCL), poly(glycolic acid) (PGA), poly(lactic acid) (PLA), poly(lactic acid-co-glycolic acid) (PLGA).
  • Polymer B may be selected from the group including, but not limited to, poly(serinol hexamethylene urea) (PSHU).
  • Exemplary triblock copolymers used to form the nanoparticles include, but are not limited to, PEG-PSHU-PEG, PCL- PSHU-PCL.
  • the polymer used to form the hydrogel exists as a single polymer including, but not limited to, gelatin, collagen, chitosan, cellulose.
  • the diblock copolymer Polyethylene glycol diglycidyl ether) succinic acid (PEGSA) may be used.
  • PEGSA needs to be crosslinked. Tyramine may be conjugated to PEGSA for the crosslinking reaction. Hydrogen peroxide or horseradish peroxidase may then be used to crosslink the ring groups on tyramine.
  • the polymers used to form the hydrogel exist as diblock copolymers having a C-D structure with polymer C being selected from the group including, but not limited to, poly(serinol hexamethylene urea) (PSHU), sulfonated poly(serinol hexamethylene urea) (SPSHU).
  • Polymer D may be selected from the group including, but not limited to, poly(A/-isopropylacrylamide) (PNIPAM).
  • PNIPAM poly(A/-isopropylacrylamide
  • Exemplary diblock copolymers used to form the hydrogel include PSHU-PNIPAM, SPSHU-PNIPAM.
  • Nanoparticle refers to a particle or structure which is biocompatible with and sufficiently resistant to chemical and/or physical destruction by the environment of such use so that a sufficient number of the nanoparticles remain substantially intact after delivery to the site of application or treatment and whose size is in the nanometer range.
  • a nanoparticle typically ranges between about 1 nm to about 1000 nm, preferably between about 50 nm and about 500 nm, more preferably between about 50 nm and about 350 nm, more preferably between about 100 nm and about 250 nm.
  • nanoparticle includes, but is not limited to, micelles, polymeric nanoparticles, and lipid-based nanoparticles such as liposomes and niosomes.
  • a micelle is used as the preferred nanoparticle type.
  • Crosslinking refers to chemically joining two or more molecules by a covalent bond.
  • exemplary crosslinking agents include, but are not limited to, carbodiimides such as 1 -Ethyl- 3-(3-dimethylaminopropyl) carbodiimide (EDC) and dicyclohexylcarbodiimide (DCC); carbonyldiimidazole (CDI); genipin (GP) b-glycerophosphate and glutaraldehyde; and the use of enzymes, including but not limited to transglutaminase, tyrosinases and horseradish peroxidases.
  • Crosslinking agents can be introduced into the materials through physical crosslinking methods as well, including, but are not limited to, dehydrothermal and ultraviolet radiation treatment.
  • Hydrogel refers to a three-dimensional (3D) crosslinked network of biocompatible hydrophilic polymers having an elastic structure and a water content or at least 10%.
  • the polymers may be naturally or synthetically derived.
  • Hydrogels have the capacity to reversibly swell in water.
  • “Gelatin” as used herein refers to a water-soluble polymer obtained from acid, alkaline or enzymatic hydrolysis of collagen. Gelatin derived from acid treatment is referred to as type A while gelatin derived from alkaline treatment is referred to as type B. In some embodiments, the gelatin may be derived from porcine skin, bovine skin, bone, poultry, or fish. In other embodiments, the gelatin may be a recombinant gelatin.
  • Catheter refers to a hollow tube formed from a biocompatible material for insertion into body cavities to drain fluids, inject fluids or distend body passages.
  • the hollow tube may be flexible or rigid depending on the use.
  • Drain is used interchangeably herein with “catheter”.
  • the term “catheter” includes, but is not limited to, pleural and mediastinal drains, as well as various drain types used in other anatomical spaces such as Blake drains, pigtail catheters, or Foley catheters.
  • Exemplary pleural and mediastinal drains include, but are not limited to, chest tubes, pleural catheters, intercostal drains, thoracic catheters, and thoracostomy tubes.
  • the present disclosure provides drug delivery systems, components thereof, methods of use, methods of formation of the systems and components thereof as well as devices using such disclosed drug delivery systems.
  • the following non-limiting examples illustrate exemplary systems and components thereof in accordance with various embodiments of the disclosure. The examples are merely illustrative and are not intended to limit the disclosure in any way.
  • NBOC serinol (0.2873 g, 1 .5 mmol) and urea (0.09 g, 1 .5 mmol) were weighed out and lyophilized at -45°C for 48 h.
  • the reactants were dissolved in 1.5 mL of anhydrous DMF in a 25 mL round- bottom flask at 90°C under gentle stirring and a nitrogen atmosphere.
  • HDI (0.482 mL, 3 mmol) was added dropwise, and the polymerization was carried out for 5 days. After the specified time, an excess of polyethylene glycol) (PEG 1000, 4 mmol) was dehydrated and added to the reaction. The PEGylation reaction was carried out for 24 h at 90°C.
  • the resulting product polyethylene glycol)-block-poly(serinol hexamethylene urea)-block-poly(ethylene glycol) (PEGPSHU-PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50°C and 10 mbar vacuum. Subsequently, the polymer was dissolved in Milli-Q water and dialyzed (MWCO: 3.5 kDa) against 1 L of Milli-Q water for 72 h at room temperature. Then, the product was lyophilized at -45°C for 48h to yield a white flaky material.
  • PEGPSHU-PEG polyethylene glycol)-block-poly(serinol hexamethylene urea)-block-poly(ethylene glycol)
  • Growth factor-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PEG-PSHU-PEG polymer and growth factor were dissolved in 1 mL of DMSO at 3 wt % (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 mL of ultrapure water partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min.
  • the removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant, and then resuspending the micelles in ultrapure water. This DMSO extraction procedure was carried out three times. The resulting micelles were either used immediately or stored at -20°C for later use.
  • Novel micelle NPs were created from a PEG-PSHU-PEG polymer.
  • the PSHU backbone provides the hydrophobic core of the micelles, while the terminal PEG chains provide the hydrophilic interactions on the exterior shell.
  • PEG adds an additional benefit as this provides a “stealth” effect to the NPs for an extended half-life in circulation.
  • the synthesis of PEG-PSHU-PEG from N-BOC serinol, urea, HDI, and PEG was confirmed using 1 H NMR (500 MHz, DMSO-d6).
  • the PSHU backbone structure was previously confirmed by 1 H NMR, 34 and the successful conjugation of PEG to PSHU was confirmed by the chemical shift at d 3.51 ppm (peak I, -0CH2CH20-), which identifies the protons on the PEG repeating unit of the polymer backbone.
  • 35 Scanning electron microscopy (SEM) was used to characterize the morphologies of the micelles.
  • the micelle NPs were uniformly spherical ( Figure 2B).
  • DLS is a technique that can be used to determine the size distribution profile of different particles, or polymers, in solution.
  • the micelle NPs’ size was determined using a Zetasizer Nano ZS. Three separate batches of micelles were made using the same protocol, and the particle size was measured ( Figure 2E).
  • the DLS data are plotted using an intensity-weighted distribution. Micelles had a mean diameter of 216.5 nm with a relative standard deviation (RSD) of 8.62%. As shown with a low RSD value, the micelles are composed of a monodispersed population of particles and batch to batch variability appears to be low.
  • RSD relative standard deviation
  • Lidocaine-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PEG-PSHU-PEG polymer and lidocaine (10 mg polymer to 100 mg drug) were dissolved in 1 mL of DMSO at 3 wt % (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 mL of ultrapure water partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min.
  • the removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant, and then resuspending the micelles in ultrapure water. This DMSO extraction procedure was carried out three times. The resulting micelles were either used immediately or stored at -20°C for later use.
  • lidocaine was used in an embodiment, other anesthetic agents or antimicrobial agents are contemplated for use with the nanoparticles described herein.
  • Micelle size was determined using a Zetasizer Nano ZS.
  • N-Boc-serinol (1 .149 g, 6 mmol) and urea (0.36 g, 6 mmol) were weighed out and lyophilized for 48 hours at -45 °C and 0.045 mbar.
  • the reactants were dissolved in 10 mL of anhydrous DMF in a 25 mL round bottom flask at 90°C under gentle stirring and a nitrogen atmosphere.
  • Hexamethylene diisocyanate (HDI) (1 .928 mL, 12 mmol) was added dropwise, and the polymerization was carried out for 3 days.
  • N-Boc-PSHU-N- Boc (0.5 g) was dissolved in 5 ml anhydrous DMF under a nitrogen atmosphere and stir.
  • Polycaprolactone (PCL, 0.5 g) was added dropwise and a catalytic amount of stannous octoate (0.0025g) was added to the flask, and the polymerization was carried out for 3 days at 70°C.
  • the resulting product, polycaprolactone-block-poly(serinol hexamethylene urea)-block- polycaprolactone (PCL-PSFIU-PCL) was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50°C and 10 mbar vacuum.
  • PCL-PSFIU- PCL can be deprotected by hydrogenation to provide free amine groups for additional chemical modification.
  • micelles can be formed with or without this deprotection step.
  • PCL-PSHU-PCL (0.1 g) was dissolved in DCM and TFA (9 ml each). Deprotection occurred by hydrogenation at room temperature for 15 min providing free amine groups.
  • Deprotected PSHU-PCL (PCL-dPSHU-PCL) was recovered by precipitation in diethyl ether and rotoevaporation. Further purification of dPSHU involved dissolving in TFE, precipitation in diethyl ether, and rotoevaporation. The product was lyophilized for 48h.
  • Lidocaine-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PCL-PSHU-PCL polymer and lidocaine (10 mg polymer to 100 mg drug) were dissolved in 1 mL DMSO at 1 wt% (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 mL of purified water (milliQ) partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min. Removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant and then re-suspending the micelles in purified water. This DMSO extraction procedure was carried out 3 times. The resulting micelles were used immediately.
  • aqueous solution containing gelatin was prepared in PBS (5% w/v) and allowed to dissolve overnight in a 37 S C water bath.
  • lidocaine-HCI was dissolved in the gelatin solution (200 mg lidocaine/1 ml gelatin solution).
  • Glutaraldehyde was then added to the gelatin-lidocaine mixture (1 % w/v) to crosslink the gelatin into a hydrogel. While 1 % w/v of crosslinker was used, a range of between about 0.5% to about 4% w/v is contemplated.
  • NBOC serinol was synthesized as described previously. 8 Briefly, serinol and di-tert-butyl dicarbonate were dissolved in ethanol at 4°C. The solution was heated at 37°C for 1 h, roto- evaporated, and dissolved in an equal volume mixture of ethyl acetate and hexane at 60°C. Additional hexane was added to form crystalline structures, and the precipitate was filtered to remove solvent, yielding N-BOC serinol as a crystalline white product.
  • PNIPAM was synthesized as described previously. 30 In short, NIPAM and 4,4-azobis(4- cyanovaleric acid) were dissolved in methanol and heated at 68°C for 3h. PNIPAM was recovered by precipitation in ultrapure water at 60°C, purified via dialysis (MWCO 12-14 kDa), and lyophilized, yielding a white product.
  • SRTG SPSHU-PNIPAM
  • 31 N-BOC serinol, urea, and HDI were dissolved in DMF and heated at 90 °C for 7 days.
  • PSHU was recovered by precipitation in diethyl ether and rotoevaporation, yielding the polyurea as a white powder.
  • PSHU was dissolved in DCM and TFA. Deprotection occurred by hydrogenation at room temperature for 45 min providing, free amine groups. Deprotected PSHU (dPSHU) was recovered by precipitation in diethyl ether and rotoevaporation.
  • dPSHU Further purification of dPSHU involved dissolving in 2,2,2- trifluoroethanol, precipitation in diethyl ether, and rotoevaporation.
  • an equivalent mass of PNIPAM was conjugated to dPSHU.
  • PNIPAM, EDC, and NHS were dissolved in DMF and activated for 24 h.
  • dPSHU was dissolved in DMF, added to the activated PNIPAM solution, and reacted for 24 h.
  • PSHU-PNIPAM was recovered by precipitation in diethyl ether, rotoevaporation, purified via dialysis (MWCO 12-14 kDa), and lyophilized. Afterward, the remaining free amine groups were sulfonated.
  • SEM Scanning Electron Microscopy
  • SRTG samples 5%, phosphate-buffered saline (PBS)
  • PBS phosphate-buffered saline
  • the inventors demonstrate the efficacy of the SRTG micelle polymer delivery system for the controlled release of proteins to promote therapeutic angiogenesis and reduce inflammation.
  • PEG-PSHU-PEG polymeric micelles were synthesized and characterized for chemical structure and morphology.
  • the PDGF-loaded micelles were encapsulated into the SRTG and examined for sequential and sustained release of the growth factor.
  • VEGF and IL-10 were encapsulated within the thermal gel system and their spatiotemporal release was evaluated.
  • An in vitro release test revealed that the delivery system is sequentially releasing all three factors and at a sustained rate for 28 days.
  • NIPAm N-lsopropylacrylamide
  • DMF dimethyl sulfoxide
  • ABTS 2,2'-azino-bis(3-ethylbenzothiazoline-6-sulfonic acid)
  • Goat anti-Rabbit IgG (H+L) Secondary Antibody Alexa Fluor 594, Goat anti-Rat IgG (H+L) Secondary Antibody Alexa Fluor 488, Rabbit anti-Goat IgG (H+L) Secondary Antibody Alexa Fluor 594, CD31 Antibody (Rat lgG2a), and smooth muscle actin antibody (a-SMA, rabbit IgG) were purchased from Thermo Fisher Scientific (Waltham, MA).
  • Anti-Von Willebrand factor antibody (VWF, sheep IgG) and anti-CD68 antibody (rabbit IgG) were purchased from Abeam (Cambridge, U.K.).
  • Human PDGF-BB Standard ABTS ELISA development kit, Murine IL-10 Standard ABTS ELISA development kit, Murine VEGF Standard ABTS ELISA development kit, Recombinant Human and Murine PDGF-BB, Recombinant Murine IL-10, and Recombinant Murine VEGF165 were purchased from Peprotech (Rocky Hill, NJ). 10% formalin was purchased from JT Baker (Phillipsburg, NJ). Sucrose (RNASE and DNASE free) was purchased from VWR Life Science (Radnor, PA). An optimal cutting temperature (OCT) compound was purchased from Sakura (Torrance, CA). Alexa Fluor 594 (goat anti-rabbit IgG) was purchased from Life Technologies (Carlsbad, CA).
  • DAPI Flouromount-G was purchased from Electron Microscope Sciences (Hartfield, PA). Spectra/Por dialysis membranes (molecular weight cut-off (MWCO): 3500-5000 and 12 000-14 000 Da) were purchased from Spectrum Laboratories (Rancho Dominguez, CA).
  • Proton nuclear magnetic resonance (1 H NMR) was performed on a Varian Inova 500 NMR spectrometer, and samples were run in DMSO-d6 at room temperature. Polymer morphology was imaged using a JEOL (Peabody, MA) JSAM-6010LA analytical scanning electron microscope. Nanoparticle size was measured using a Zetasizer Nano ZS (Malvern Instruments Ltd, Worcestershire, U.K.) ⁇ ELISA color development was monitored with a BioTek Synergy 2 Multi- Mode Reader at 405 nm with wavelength correction set at 650 nm. Confocal images were taken using the Zeiss LSM 780.
  • NBOC serinol was synthesized as described previously. 8 Briefly, serinol and di-tert-butyl dicarbonate were dissolved in ethanol at 4 °C. The solution was heated at 37 °C for 1 h, rotoevaporated, and dissolved in an equal volume mixture of ethyl acetate and hexane at 60°C. Additional hexane was added to form crystalline structures, and the precipitate was filtered to remove solvent, yielding N-BOC serinol as a crystalline white product.
  • PNIPAM was synthesized as described previously. 30 In short, NIPAM and 4,4-azobis(4- cyanovaleric acid) were dissolved in methanol and heated at 68°C for 3 h. PNIPAM was recovered by precipitation in ultrapure water at 60°C, purified via dialysis (MWCO 12-14 kDa), and lyophilized, yielding a white product.
  • SRTG SPSHU-PNIPAM
  • 31 N-BOC serinol, urea, and HDI were dissolved in DMF and heated at 90°C for 7 days.
  • PSHU was recovered by precipitation in diethyl ether and rotoevaporation, yielding the polyurea as a white powder.
  • PSHU was dissolved in DCM and TFA. Deprotection occurred by hydrogenation at room temperature for 45 min providing, free amine groups. Deprotected PSHU (dPSHU) was recovered by precipitation in diethyl ether and rotoevaporation.
  • dPSHU Further purification of dPSHU involved dissolving in 2,2,2- trifluoroethanol, precipitation in diethyl ether, and rotoevaporation.
  • an equivalent mass of PNIPAM was conjugated to dPSHU.
  • PNIPAM, EDC, and NHS were dissolved in DMF and activated for 24 h.
  • dPSHU was dissolved in DMF, added to the activated PNIPAM solution, and reacted for 24 h.
  • PSHU-PNIPAM was recovered by precipitation in diethyl ether, rotoevaporation, purified via dialysis (MWCO 12-14 kDa), and lyophilized. Afterward, the remaining free amine groups were sulfonated.
  • the PEGylation reaction was carried out for 24 h at 90°C.
  • the resulting product polyethylene glycol)-blockpoly(serinol hexamethylene urea)-block-poly(ethylene glycol) (PEGPSHU-PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50 °C and 10 mbar vacuum. Subsequently, the polymer was dissolved in Milli-Q water and dialyzed (MWCO: 3.5 kDa) against 1 L of Milli-Q water for 72 h at room temperature. Then, the product was lyophilized at -45°C for 48 h to yield a white flaky material.
  • Growth factor-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PEG-PSHU-PEG polymer and growth factor were dissolved in 1 ml_ of DMSO at 3 wt % (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 ml_ of ultrapure water partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min.
  • the removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant, and then resuspending the micelles in ultrapure water. This DMSO extraction procedure was carried out three times. The resulting micelles were either used immediately or stored at -20°C for later use.
  • SRTG samples (5%, phosphate-buffered saline (PBS)), with and without micelles, were gelled at 37°C for 15 min. The hydrogels were snap-frozen in liquid nitrogen. The frozen samples were quickly broken to expose the structure within the scaffold and lyophilized. Samples were sputtercoated with gold and palladium and examined by SEM. Dynamic Light Scattering (DLS). The micelle’s size was determined using a Zetasizer Nano ZS. Three separate batches of micelles were made using the same protocol, and the particle size was measured.
  • DLS Dynamic Light Scattering
  • Proton nuclear magnetic resonance (1 H NMR) was completed with a Varian Inova 500 NMR spectrometer. 1 H NMR was used to confirm the structure of PEG-PSHU-PEG.
  • the PEG-PSHU- PEG samples to be analyzed were dissolved in 600 pl_ of DMSO-d6. Protein Release Test.
  • SRTG 5 wt % polymeric solutions (SRTG + VEGF + NPs + PDGF + IL-10) were created using PBS with 0.2% BSA with samples run in triplicates. Solutions were mixed and left to dissolve overnight at 4°C. 500 ng of each of IL-10, the VEGF, and the PDGF was added to the gels at different times.
  • the micelle NPs loaded with the PDGF were added 30 min before gelling, and the VEGF and IL- 10 were added 10 min before gelation.
  • 50 pL gels were formed in 2 mL vials and placed in a 37 °C incubator for 5 min to promote gelation and encapsulation of proteins and micelles.
  • 1 mL of warm (37°C) release solution (1 c PBS with 0.2% BSA) was added to the vials. After 5 min, 1 mL of release solution was removed from each sample for the day 0 time point, and 1 mL of fresh warm release solution was added. Samples were then taken every 24 h and immediately placed in a -80°C freezer for further analysis.
  • Enzyme-linked immunosorbent assay was used to quantify the concentration of proteins in the release solution samples following the manufacturer’s instructions (Peprotech, Rocky Hill, NJ). ELISA color development was monitored with a BioTek Synergy 2 Multi-Mode Reader at 405 nm with wavelength correction set at 650 nm. The results of three replicates were averaged. The hydrogels were initially allowed to stabilize for 5 min at which a sample was taken to determine the loading efficiency of the proteins in the hydrogels (day 0).
  • mice (The Jackson Laboratory) weighing 24-28 g were maintained on a light/dark (14 h light, 10 h dark) cycle with access to food and water ad libitum. The mice were anesthetized using continuous isoflurane and oxygen inhalation. Initial induction was at 5% isoflurane in oxygen and then maintained at 2% isoflurane in oxygen.
  • mice A total of 36 adult male C57BL/6 mice were used for the study, 3 mice per each injection group (saline, SRTG, SRTG + VEGF, SRTG + VEGF + PDGF, SRTG + VEGF + NPs + PDGF, or SRTG + VEGF + NPs + PDGF + IL-10) for 2 time points (7, 21 days). The mice were allowed 7 days to acclimate prior to injections. Proteins were loaded as described in the protein release test section.
  • Saline 60 pL of saline
  • SRTG 1%, 60 pL of saline
  • SRTG + VEGF 250 ng, 60 pL of saline
  • SRTG + VEGF + PDGF 1 %, 250 ng of VEGF and PDGF each, 60 pL of saline
  • SRTG + VEGF + NPs + PDGF 1%, 250 ng of VEGF and PDGF each, 3 mg of NPs, 60 pL of saline
  • SRTG + VEGF + NPs + PDGF + IL-10 1%, 250 ng of VEGF, PDGF, and IL-10 each, 3 mg of NPs, 60 pL of saline
  • mice were euthanized by carbon dioxide and cervical dislocation 7 or 21 days after injection. After removing the hair around the injection site with depilatory cream, the subcutaneous tissue (2 cm x 2 cm) was harvested at the site of injection. The site was clearly visible as a result of shaving the backs of the mice at the time of the initial injection.
  • the subcutaneous tissue was fixed in formalin (10%, PBS) overnight, cryoprotected with sucrose (30%, PBS) for 1 day, embedded in optimal cutting temperature (OCT) compound, and frozen at -80 °C. The subcutaneous tissue was sectioned transversely with a thickness of 5 pm.
  • Sections were fixed in formalin (10%, PBS) for 10 min and washed three times with wash buffer (0.1% Tween 20, PBS) for 5 min each.
  • Permeabilization buffer (0.5% Triton X-100, PBS) was used for 10 min, and the sections were washed three times with a wash buffer for 5 min each.
  • Blocking buffer (0.25% Triton X-100, 2% bovine serum albumin (BSA), 4% bovine y globulins, and PBS) was used for 60 min on the sections at room temperature. All antibodies were diluted in dilution buffer (0.25% Triton X-100, 2% BSA, 4% bovine y globulins, and PBS).
  • the sections were stained with primary antibodies CD31 (1 :50), VWF (1 :50), SMA (1 :250), and/or CD68 (1 :500) overnight at 4 °C and washed three times in wash buffer for 5 min each.
  • the sections were stained with secondary antibodies Alexa Fluor 488 (1 :500) for CD31 and/or the associated Alexa Fluor 594 (1 :500) for VWF, SMA, and CD68 at room temperature.
  • the sections were washed three times in wash buffer for 5 min each and washed three times in ultrapure water for 5 min each.
  • DAPI Fluoromount-G was used to stain nuclei and mount the sections.
  • Novel micelle NPs were created from a PEG-PSHU-PEG polymer, which is based on the same PSHU backbone as SRTG, making the polymer biocompatible and biodegradable. As this polymer is used to fabricate micelles, it must have differing sections of hydrophobic and hydrophilic chains to form these particles.
  • the PSHU backbone provides the hydrophobic core of the micelles, while the terminal PEG chains provide the hydrophilic interactions on the exterior shell, which can ionically bind to growth factors (Figure 1A). PEG adds an additional benefit as this provides a “stealth” effect to the NPs for an extended half-life in circulation.
  • the micelles When the micelles are injected together with SRTG, the micelles can provide long-term and sequential release of angiogenic growth factors, such as PDGF, while the SRTG provides a scaffold for their retention at the injection site.
  • angiogenic growth factors such as PDGF
  • the PSHU backbone structure was previously confirmed by 1 H NMR, 34 and the successful conjugation of PEG to PSHU was confirmed by the chemical shift at d 3.51 ppm (peak I, -0CH2CH20-), which identifies the protons on the PEG repeating unit of the polymer backbone.
  • 35 Scanning electron microscopy (SEM) was used to characterize the morphologies of the SRTG, micelles, and the combined delivery system.
  • the SRTG alone is composed of polymer sheets (Figure 2A), while the micelle NPs were uniformly spherical (Figure 2B).
  • Figure 2C shows the morphology of the SRTG encapsulating the micelle NPs.
  • FIG. 2D is an enlargement of the porous area of the same polymeric gel showing that the micelles are embedded within the gel and their uniform spherical shape has been maintained.
  • the black arrows indicate the micelles in the gel.
  • DLS is a technique that can be used to determine the size distribution profile of different particles, or polymers, in solution.
  • the micelle NPs’ size was determined using a Zetasizer Nano ZS. Three separate batches of micelles were made using the same protocol, and the particle size was measured ( Figure 2E).
  • the DLS data are plotted using an intensity-weighted distribution. Micelles had a mean diameter of 216.5 nm with a relative standard deviation (RSD) of 8.62%. As shown with a low RSD value, the micelles are composed of a monodispersed population of particles and batch to batch variability appears to be low.
  • RSD relative standard deviation
  • VEGF vascular endothelial growth factor
  • VEGF and PDGF Increased Functional Vascular Endothelial Cell Recruitment and Endothelial Cell Differentiation into Vascular Cells
  • the SRTG polymeric delivery system was subcutaneously injected in the middle back of mice.
  • the process of angiogenesis encompasses protein binding to endothelial cell receptors, basement membrane degradation by matrix metalloproteinases, endothelial cell proliferation and migration, subsequent vessel formation and remodeling, and finally vessel stabilization by pericytes. 38 Immunohistochemistry was performed to directly identify functional blood vessels. This staining technique allows for the quantification of the different cell types involved in the process of blood vessel formation.
  • Endothelial cells were stained with a CD31 antibody, as this protein makes up a substantial portion of endothelial cell intercellular junctions.
  • VWF a factor involved in hemostasis, was used to identify functional vascular endothelial cells when endothelial cells were stained with both VWF and CD31 ( Figure 4).
  • TNF-a tumor necrosis factor-a
  • SRTG + VEGF + NPs + PDGF + IL-10 group displayed significantly more endothelial cells than saline, SRTG, and SRTG + VEGF + PDGF after 7 days, and it showed significantly more endothelial cells than saline and SRTG after 21 days ( Figure 4B).
  • the SRTG + VEGF group did not show a statistically significant difference in functional endothelial cells compared to the SRTG and saline groups, which shows that the delivery of PDGF is important to increase the proliferation of vascular endothelial cells (Figure 4C). Furthermore, the ratio of functional vascular endothelial cells to endothelial cells was quantified to determine the number of endothelial cells maturing into vascular cells. Both sequential delivery groups, with the PDGF-loaded micelle NPs, demonstrated a significant increase in functional vascular endothelial cell maturation after 21 days, while the other groups did not (Figure 4D). This highlights the importance of sequentially delivering different angiogenic factors that are involved in distinct stages of the vessel formation process to promote more mature and functional vessels.
  • pericytes The recruitment of pericytes (mural cells or vascular smooth muscle cells) was used to identify stable and mature vasculature. Newly formed vessels can become hemorrhagic and may regress over time if they are not stabilized with recruited pericytes. 14 ’ 20 ’ 41 Immunohistochemistry was implemented to observe the recruitment of pericytes to endothelial cell vessel lumens. Pericytes were stained with a-smooth muscle actin (a-SMA), which is a major component of microfilament bundles responsible for contractile functions. The association of a- SMA+ cells around CD31 + cells was used to identify mature and stable blood vessels at 7 and 21 days after injection (Figure 5). More mature vessels should show an organized circle of endothelial cells surrounded by vascular smooth muscle cells.
  • a-SMA a-smooth muscle actin
  • the SRTG + VEGF + PDGF group did not show any increase in pericyte recruitment between 7 and 21 days, unlike the other delivery systems loaded with growth factors (Figure 5B).
  • This group does not have PDGF encapsulated within NPs, which indicates that VEGF and PDGF are most likely burst releasing simultaneously.
  • early-stage angiogenic factors can have inhibitory effects on late-stage growth factors and vice versa, when presented concurrently. 15 - 17 ’ 20
  • These results for the SRTG + VEGF + PDGF group further confirm the importance of sequentially delivering angiogenic growth factors to develop stable and mature blood vessels.
  • both sequential delivery groups showed a significant increase in the ratio of vascular smooth muscle cells to endothelial cells compared to that of the other three polymer injection groups after 21 days (Figure 5C).
  • the ratio approached the physiological ratio shown in the saline control, as would be expected.
  • the lack of vessel stabilization and maturation observed with the other growth factor delivery groups shows the limitations of VEGF delivery without the sequential release of a growth factor to recruit pericytes.
  • SRTG + VEGF + NPs + PDGF + IL-10 exhibited a significant reduction in macrophages compared to the other four polymer injection groups after 21 days ( Figure 6B). This result demonstrates that IL-10 is significantly reducing the immune response caused by the polymer gel. Additionally, the SRTG + VEGF, SRTG + VEGF + PDGF, and SRTG + VEGF + NPs + PDGF + IL-10 all showed a significant reduction in macrophages from 7 to 21 days ( Figure 6B). Although the saline group showed significantly less macrophages than all polymer delivery groups (Figure 6B), it is expected that the macrophage cell count would continue to decrease over time and ultimately reach normal physiological counts.
  • An injectable protein polymer delivery system that can revascularize heart tissue could dramatically reduce healthcare costs related to CAD, which would also improve the quality of life for millions of people suffering from this disease.
  • the process of treating CAD can involve many different tests and expensive medications, in addition to the almost inevitable open-heart surgery that must be implemented.
  • a standard injectable protein delivery system that can stimulate new blood vessel formation around a blocked artery could allow many people to avoid paying burdensome healthcare bills.
  • Incorporating a supplementary anti-inflammatory protein into this system might also minimize myocardial necrosis and optimize cardiac repair following Ml.
  • the polymer system may help with reducing the number of heart attacks experienced by patients because the delivery system could be injected before a complete occlusion of a coronary artery occurs, further reducing healthcare costs.
  • polymer delivery systems for CAD treatment localized and controlled release of proteins to revascularize damaged heart tissue and reduce the detrimental effects of a prolonged inflammatory response remains inadequate.
  • VEGF vascular endothelial growth factor
  • the SRTG delivery system provided spatiotemporal controlled release of VEGF to answer some of these growth factor delivery requirements. 5 ’ 12
  • novel PDGF-loaded micelle NPs were incorporated, which are embedded within the thermal gel and provide the sequential release of the late-stage angiogenic factor that is needed to stabilize newly formed vessels with vascular smooth muscle cells.
  • the inventors previously demonstrated the release of a drug (triamcinolone acetonide or TA) from polymeric micelles fabricated from a similar triblock copolymer. In that study, the authors examined the release of TA from the micelles alone and from the micelles encapsulated within a similar reverse thermal gel.
  • That release test showed a comparable drug release rate from the encapsulated micelles to the PDGF-loaded micelles shown here, demonstrating that encapsulating the micelles in a reserve thermal gel results in no burst release as seen with the micelles alone.
  • the inventors have not performed a release study with the PDGF-loaded micelles alone, the inventors believe that the release rate would be analogous to the release profile from the previous study. Based on the triblock polymer structure, there is no functional group with covalent bonding to the drug or protein, hence the inventors believe that there will be a zero-order release rate regardless of which micelle is evaluated.
  • the polymer delivery system was evaluated for the controlled and sequential release of three proteins using an in vitro release study.
  • the results demonstrated that the SRTG system sequentially released all three factors: VEGF first, followed by IL-10, and finally PDGF.
  • the inventors believe the different heparin-binding affinities of IL-10 and VEGF provided the different release rates between the two factors, although additional testing is necessary to confirm this.
  • Kinetic studies have demonstrated that some angiogenic factors bind to sulfated scaffolds in order of their specific binding affinities (KA); 37 ’ 48 thus, the inventors believe the proteins are being released with a similar mechanism.
  • the release test results showed that PDGF was being slowly released from the micelles that were embedded within the SRTG, which provided the sequential release of PDGF from the scaffold last. This spatiotemporal release of growth factors from the polymer scaffold is critical to promote stable and mature angiogenic vessel formation.
  • the inventors have provided a modified release profile (Figure 7), which demonstrates the sustained and sequential release rate of the proteins.
  • the graph was created by removing the burst release and starting the release profile for each factor with the data from day 1 , while normalizing the release profile to start at 0. The inventors believe that this may better exemplify the data since it can be difficult to discern the release rates from the full graph. While there are only partial nanogram amounts of each protein releasing after a certain amount of time, the inventors believe that these small amounts are still enough to stimulate a cellular angiogenesis response as demonstrated by the subcutaneous injection study and by evidence from other studies.
  • TGF-p1 very low concentrations of TGF-p1 , such as 10 pg/mL, stimulated the growth rate of NIH3T3 cells in low serum or serum-free medium. 53 With this study taken in conjunction with the animal study, the inventors believe that partial nanogram amounts of growth factor can have a significant cell impact.
  • PNIPAM-based reverse thermal gels have different properties below and above their respective lower critical solution temperature (LCST, 32 °C for PNIPAM).
  • LCST lower critical solution temperature
  • the polymers are soluble, and the hydrophilic interactions are dominant.
  • higher than the LCST the polymers become progressively hydrophobic and insoluble, creating the formation of a physical gel. This transition above the LCST causes the thermoresponsive block to shrink, which causes the proteins to be expelled until the gel has stabilized. 54
  • the inventors believe that this is why there is a rapid release of the proteins in the initial first day.
  • the previous study determined that the SRTG goes through a rapid, reversible phase transition around 32 °C (LCST) by UV-visible spectroscopy.
  • the primary goal for this embodiment is toward injecting it into the heart to help optimize cardiac repair and limit the damaging effects of inflammation following Ml.
  • the inventors have previously demonstrated that the SRTG scaffold can effectively deliver VEGF to the heart in a Ml model to provide cardioprotective benefits. 12
  • the inventors are investigating the efficacy of this scaffold for delivering various combinations of heparin-binding proteins, such as basic fibroblast growth factor (FGF-2), stromal cell-derived factor 1 a (SDF-1a), transforming growth factor-b ⁇ (TGF- b1), and hepatocyte growth factor (HGF).
  • FGF-2 basic fibroblast growth factor
  • SDF-1a stromal cell-derived factor 1 a
  • TGF- b1 transforming growth factor-b ⁇
  • HGF hepatocyte growth factor
  • NBOC serinol (0.2873 g, 1 .5 mmol) and urea (0.09 g, 1 .5 mmol) were weighed out and lyophilized at -45° C and 0.045 mbar for 48 h.
  • the reactants were dissolved in 1 .5 ml_ of anhydrous DMF in a 25 ml_ round-bottom flask at 90° C under gentle stirring and a nitrogen atmosphere.
  • HDI (0.482 ml_, 3 mmol) was added dropwise, and the polymerization was carried out for 5 days. After the specified time, an excess of polyethylene glycol) (PEG 1000, 4 mmol) was dehydrated and added to the reaction.
  • the PEGylation reaction was carried out for 24 h at 90 °C.
  • the resulting product polyethylene glycol)-block-poly(serinol hexamethylene urea)-block-poly(ethylene glycol) (PEG- PSHU-PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50° C and 10 mbar vacuum. Subsequently, the polymer was dissolved in Milli-Q water and dialyzed (MWCO: 3.5 kDa) against 1 L of Milli-Q water for 72 h at room temperature. Then, the product was lyophilized at -45 °C for 48 h to yield a white flaky material.
  • Lidocaine-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PEG-PSHU-PEG polymer and lidocaine (about 1 -20 mg polymer to about 50-200 mg lidocaine) were dissolved in 1 mL of DMSO at 3 wt% (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 mL of ultrapure water (milliQ or equivalent) to give a 1 :20 organic:aqueous phase ratio partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min.
  • DMSO methyl methacrylate
  • the removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant, and then resuspending the micelles in ultrapure water. This DMSO extraction procedure was carried out three times. The resulting micelles were either used immediately or stored at -20 °C for later use.
  • aqueous solution containing gelatin was prepared in PBS (5% w/v) and allowed to dissolve overnight in a 37 S C water bath.
  • Micelle nanoparticles loaded with lidocaine-HCI 50-200 mg lidocaine, 1 -20 mg polymer) were added to the gelatin solution (0.5 ml gelatin solution).
  • GA was then added to the gelatin-nanoparticles-lidocaine mixture (1% w/v GA) to crosslink the gelatin into a hydrogel.
  • the delivery system can have applicability to reducing or eliminating pain associated with a wound by being administered topically or injected subcutaneously into a patient in need thereof.
  • the dual micelle-hydrogel system allows for controlled sustained release of an anesthetic such as lidocaine.
  • an anesthetic such as lidocaine.
  • a further use for the delivery system can be found in Example 6 in which the lidocaine-loaded nanoparticle gel is coated onto the surface of a chest tube to reduce pain associated with placement of the chest tube.
  • lidocaine as the anesthetic
  • other anesthetics are contemplated for use in the invention.
  • NBOC serinol was synthesized as described previously. 8 Briefly, serinol and di-tert-butyl dicarbonate were dissolved in ethanol at 4 °C. The solution was heated at 37 °C for 1 h, rotoevaporated, and dissolved in an equal volume mixture of ethyl acetate and hexane at 60 °C. Additional hexane was added to form crystalline structures, and the precipitate was filtered to remove solvent, yielding N-BOC serinol as a crystalline white product.
  • PNIPAM was synthesized as described previously. 30 In short, NIPAM and 4,4-azobis(4- cyanovaleric acid) were dissolved in methanol and heated at 68 °C for 3 h. PNIPAM was recovered by precipitation in ultrapure water at 60 °C, purified via dialysis (MWCO 12-14 kDa), and lyophilized, yielding a white product. SRTG (SPSHU-PNIPAM) was synthesized similarly as described previously. 31 N-BOC serinol, urea, and HDI were dissolved in DMF and heated at 90 °C for 7 days.
  • PSHU was recovered by precipitation in diethyl ether and rotoevaporation, yielding the polyurea as a white powder.
  • PSHU was dissolved in DCM and TFA. Deprotection occurred by hydrogenation at room temperature for 45 min providing, free amine groups.
  • Deprotected PSHU (dPSHU) was recovered by precipitation in diethyl ether and rotoevaporation. Further purification of dPSHU involved dissolving in 2,2,2- trifluoroethanol, precipitation in diethyl ether, and rotoevaporation.
  • an equivalent mass of PNIPAM was conjugated to dPSHU.
  • PNIPAM, EDC, and NHS were dissolved in DMF and activated for 24 h.
  • PSHU-PNIPAM was recovered by precipitation in diethyl ether, rotoevaporation, purified via dialysis (MWCO 12-14 kDa), and lyophilized. Afterward, the remaining free amine groups were sulfonated. 1 ,3-Propane sultone, potassium tert-butoxide, and PSHU-PNIPAM were dissolved in DMF and reacted at 60 °C for 3 days.
  • SRTG was recovered by precipitation in diethyl ether, rotoevaporation, purified via dialysis (MWCO 12-14 kDa), and lyophilized, yielding a light- yellow product. Polymer characterization was performed as described previously to confirm material produced. 5 ’ 9 ’ 32
  • NBOC serinol (0.2873 g, 1 .5 mmol) and urea (0.09 g, 1 .5 mmol) were weighed out and lyophilized at -45 °C for 48 h.
  • the reactants were dissolved in 1 .5 ml_ of anhydrous DMF in a 25 ml_ round- bottom flask at 90 °C under gentle stirring and a nitrogen atmosphere.
  • HDI 0.482 ml_, 3 mmol
  • an excess of polyethylene glycol) (PEG 1000, 4 mmol) was dehydrated and added to the reaction.
  • the PEGylation reaction was carried out for 24 h at 90°C.
  • the resulting product polyethylene glycol)-blockpoly(serinol hexamethylene urea)-block-poly(ethylene glycol) (PEGPSHU-PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50 °C and 10 mbar vacuum. Subsequently, the polymer was dissolved in Milli-Q water and dialyzed (MWCO: 3.5 kDa) against 1 L of Milli-Q water for 72 h at room temperature. Then, the product was lyophilized at -45 °C for 48 h to yield a white flaky material.
  • Lidocaine-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PEG-PSHU-PEG polymer and lidocaine were dissolved in 1 ml_ of DMSO at 3 wt % (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 ml_ of ultrapure water partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min.
  • the removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant, and then resuspending the micelles in ultrapure water. This DMSO extraction procedure was carried out three times. The resulting micelles were either used immediately or stored at -20 °C for later use.
  • Example 6 Lidocaine loaded PEG-PSHU-PEG micelles embedded in crosslinked gelatin used to coat chest tube
  • lidocaine loaded micelle nanoparticles which can be embedded within a bio-absorbable polymer hydrogel to form a drug delivery system that can be applied to a catheter, such as a chest tube, to allow sustained release of the lidocaine over a period of between about 3-7 days.
  • Polymer concentration, lidocaine-loading amount and crosslinker concentration can be optimized to control the amount and rate of lidocaine release.
  • Lidocaine-loaded PEG-PSHU-PEG micelles were manufactured as described previously in Example 4 and detailed again below.
  • NBOC serinol (0.2873 g, 1 .5 mmol) and urea (0.09 g, 1 .5 mmol) were weighed out and lyophilized at -45° C and 0.045 mbar for 48 h.
  • the reactants were dissolved in 1 .5 ml_ of anhydrous DMF in a 25 ml_ round-bottom flask at 90° C under gentle stirring and a nitrogen atmosphere.
  • HDI (0.482 ml_, 3 mmol) was added dropwise, and the polymerization was carried out for 5 days. After the specified time, an excess of polyethylene glycol) (PEG 1000, 4 mmol) was dehydrated and added to the reaction.
  • the PEGylation reaction was carried out for 24 h at 90 °C.
  • the resulting product polyethylene glycol)-block-poly(serinol hexamethylene urea)-block-poly(ethylene glycol) (PEG- PSHU-PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50° C and 10 mbar vacuum. Subsequently, the polymer was dissolved in Milli-Q water and dialyzed (MWCO: 3.5 kDa) against 1 L of Milli-Q water for 72 h at room temperature. Then, the product was lyophilized at -45 °C for 48 h to yield a white flaky material.
  • Lidocaine-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PEG-PSHU-PEG polymer and lidocaine (about 1 -20 mg polymer to about 50-200 mg lidocaine) were dissolved in 1 mL of DMSO at 3 wt% (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 mL of ultrapure water (milliQ or equivalent) to give a 1 :20 organic:aqueous phase ratio partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min.
  • DMSO methyl methacrylate
  • the removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant, and then resuspending the micelles in ultrapure water. This DMSO extraction procedure was carried out three times. The resulting micelles were either used immediately or stored at -20 °C for later use.
  • aqueous solution containing gelatin was prepared in PBS (5% w/v) and allowed to dissolve overnight in a 37 S C water bath.
  • Micelle nanoparticles loaded with lidocaine-HCI 50-200 mg lidocaine, 1 -20 mg polymer) were added to the gelatin solution (0.5 ml gelatin solution).
  • GA was then added to the gelatin-nanoparticles-lidocaine mixture (1% w/v GA) to crosslink the gelatin into a hydrogel.
  • the crosslinked hydrogel containing the embedded lidocaine-loaded micelles was applied to the chest tube and dehydrated thus resulting in a thin, stable polymer coating on the tubing.
  • the viscous solution was poured onto the catheter tubing and allowed to fully polymerize for 5 minutes.
  • the tubing was slowly spun around by hand during the polymerization, to ensure a smooth, even coating of the gel. From here, the hydrogel coating on the tubing can be fully dehydrated overnight at room temperature or used immediately for further applications.
  • the fully dehydrated samples result in a thin, rigid polymer coating on the tubing, while the fresh hydrogel samples produce a thick, elastic, jelly-like coating. Both forms of the gelatin coating were analyzed with an ex vivo rib test and a lidocaine release study was performed.
  • a mold can be used, instead of hand-spinning, to apply the hydrogel to the tubing.
  • an aqueous solution containing gelatin was prepared in PBS (5% w/v) and allowed to dissolve overnight in a 37 S C water bath.
  • Lidocaine-HCI was dissolved in the gelatin solution (200 mg lidocaine/1 ml gelatin solution).
  • Glutaraldehyde was then added to the gelatin- lidocaine mixture (1% w/v) to crosslink the gelatin into a hydrogel.
  • the tubing was inserted into the mold containing the gelatin-glutaraldehyde solution and the polymerization was allowed to continue until fully gelled (approximately 1 hour).
  • the hydrogel coated tubing was removed from the mold. After polymerization the coated chest tubes are dehydrated to form a stably fixed coating.
  • the introduction of a lyophilizer may control the dehydration step and further aid in eliminating variation in elution that may be related to this step.
  • a nasogastric tube was inserted into the chest tube lumen during the coating and polymerization process.
  • wooden skewers were inserted into the lumen allowing the entire length of the Blake drain to be coated.
  • the coating may be applied in the operating room similar to the application of BioGlue®. Applying the coating in the operating room has the advantage of being able to adjust the dose if there is a concern for toxicity based on weight or hepatic impairment.
  • components can be filtered prior to crosslinking, applied to the tube and the tube subsequently packaged in a sealed bag and sterilized in an autoclave for in vivo application.
  • an exemplary chest tube is coated as described above with the composition.
  • the coated chest tube is then inserted ex vivo between pork spareribs using a technique simulating chest tube insertion in a patient and the durability and stability of the coating is observed.
  • a coated tubing which comprises a dried 1 cm by 0.5 cm piece of gelatin-lidocaine (100 mg lidocaine in 0.5 ml gelatin, 1% GA) was used.
  • the coated tubing was placed in a 7 ml glass scintillation vial for the release study.
  • the release solution medium (1 ml of PBS) was added to the scintillation vials, and the sample was incubated at 37 s C. 1 ml samples were taken daily, and the release solution was replenished. The day 0 sample was taken 5 minutes after adding the release solution to the vials to measure the amount of lidocaine not incorporated into the gels.
  • a UV-Vis spectrophotometer GENESYS 10S UV-Vis
  • lidocaine-polymer 5 mm x 10mm chest tube pieces are coated with equal volumes of lidocaine-polymer. SEM and lidocaine release testing are performed before and after autoclave sterilization.
  • Group 1 No incision or chest tube (2 subjects);
  • Group 2 Bilateral incisions, no chest tube (2 subjects);
  • Group 3 One control tube, one lidocaine-polymer coated tube (11 subjects);
  • Group 4 One Lidocaine-polymer tube only (2 subjects).
  • Data collected include daily weight, food consumption, mobility, analgesia requirements, other morbidity and mortality.
  • 64 Drug levels and inflammatory markers are measured daily out to 7 days.
  • Behavior and in vivo neurophysiology core technicians perform daily, blinded, non-biased assessment of pain.
  • the skin is probed with Von Frey hairs around the tube and response graded on the nociceptive scale. How far from the tube a change in nociception occurs is also assessed.
  • Tissue is prepared for histology. Core services are utilized to characterize scar formation, cytotoxicity and inflammation.
  • PEG-PSHU-PEG and Micelle Fabrication and Characterization Novel micelle NPs were created from a PEG-PSHU-PEG polymer.
  • the PSHU backbone provides the hydrophobic core of the micelles, while the terminal PEG chains provide the hydrophilic interactions on the exterior shell.
  • PEG adds an additional benefit as this provides a “stealth” effect to the NPs for an extended half-life in circulation.
  • Scanning electron microscopy (SEM) was used to characterize the morphologies of the micelles.
  • the micelle NPs were uniformly spherical ( Figure 9).
  • the release profiles from the various NP formulations show the differences in lidocaine release rates are dependent on the amount of NP loaded into the hydrogels.
  • the data display the sustained release of the lidocaine over 10 days for the 20 mg and 10 mg NP groups.
  • the data for last day of each group shows last timepoint where lidocaine could be detected in the release solution.
  • Figure 11 depicts the in vitro profiles of different nanoparticle formulations that have been loaded with varied amounts of lidocaine.
  • the release profiles from the various NP formulations show the differences in lidocaine release rates are dependent on the amount of NP and lidocaine loaded into the hydrogels.
  • the data display the sustained release of the lidocaine over 11 days for the 50 mg lidocaine groups.
  • the data for last day of each group shows last timepoint where lidocaine could be detected in the release solution.
  • the optimal formulations of the gelatin-nanoparticle delivery system were selected for additional examination with release tests using three samples each. Embedding 10 mg of nanoparticles into the hydrogel, while loading 50 or 100 mg of lidocaine in the nanoparticles, displayed the ideal sustained release profile of lidocaine over 11 days. These two formulations were further examined to additional release testing. The release tests show the sustained and linear release of lidocaine from the hydrogel delivery system with either 100 or 50 mg of the drug loaded into the nanoparticles ( Figure 12).
  • Subcutaneous Implantation Mouse Model of Catheter Tubing with and without Gelatin-Nanoparticle Hydrogel Coating Gelatin-nanoparticle hydrogel coated tubing samples were prepared as described previously. Nanoparticles (0.2 mg) were loaded with 1 mg of lidocaine and embedded within gelatin hydrogels for coating procedure using a 0.5 cm by 1 cm piece of catheter tubing. For the mouse model, a 2 cm incision was made in the skin over the right posterior chest and a subcutaneous space was created. The silicone tube was inserted either uncoated or coated depending on whether control or experimental. The entirety of the tube was implanted subcutaneously, and the incision was closed with an external suture closure. The mice were euthanized on postoperative day 14 after the procedure. Tissue from the skin around the tube was harvested and analyzed by gross examination (Figure 13). The images show similar fibrous-tissue capsule formation between the coated and uncoated tubing samples.
  • Lidocaine loaded micelles embedded in crosslinked gelatin hydrogel produce a durable and stable coating to chest tube
  • the coating does not flake off the chest tube after being applied.
  • the coating moves in coordination with the tube thus allowing the tube to remain flexible.
  • the coating remains stably fixed on the chest tube. The stable fixation of the coating allows the surgeon to maintain their techniques for handling and placing the chest tube in the patient.
  • the stable fixation of the coating to the tube prevents obstruction of the chest tube openings or central lumen. Similarly, when using a Blake drain, none of the drain channels are blocked.
  • Coated chest tube is able to stably release lidocaine
  • a post-sterilization lidocaine release profile indicates continued stable elution over a period of about 3-7 days.
  • the maximum subcutaneous dose of lidocaine hydrochloride is about 4.5 mg/kg with a maximum dose at one time not to exceed 300 mg.
  • the targeted lidocaine release for a chest tube or Blake drain is 4.5 mg/kg/day.
  • lidocaine as the anesthetic
  • other anesthetics are contemplated for use in the invention.
  • a chest tube is used in this embodiment, other types of catheters may be similarly coated.
  • Example 7 Lidocaine loaded crosslinked gelatin used to coat chest tube
  • Gelatin-Lidocaine Hydrogel Formulation An aqueous solution containing gelatin was prepared in PBS (5% w/v or 10% w/v) and allowed to dissolve overnight in a 37 s C water bath. Lidocaine-HCI was dissolved in the gelatin solution (200 mg lidocaine/1 ml gelatin solution). Glutaraldehyde (GA) was then added to the gelatin-lidocaine mixture (1 -4% w/v) to crosslink the gelatin into a hydrogel. Before the hydrogel was fully formed, the viscous solution was poured onto the catheter tubing and allowed to fully polymerize for 5 minutes. The tubing was slowly spun around by hand during the polymerization, to ensure a smooth, even coating of the gel.
  • PBS 5% w/v or 10% w/v
  • the hydrogel coating on the tubing can be fully dehydrated overnight at room temperature or used immediately for further applications.
  • the fully dehydrated samples result in a thin, rigid polymer coating on the tubing, while the fresh hydrogel samples produce a thick, elastic, jelly-like coating.
  • Both forms of the gelatin coating were analyzed with an ex vivo rib test and a lidocaine release study was performed as described below.
  • the introduction of a lyophilizer may control the dehydration step and further aid in eliminating variation in elution that may be related to this step.
  • a nasogastric tube was inserted into the chest tube lumen during the coating and polymerization process.
  • wooden skewers were inserted into the lumen allowing the entire length of the Blake drain to be coated.
  • the coating may be applied in the operating room similar to the application of BioGlue®. Applying the coating in the operating room has the advantage of being able to adjust the dose if there is a concern for toxicity based on weight or hepatic impairment.
  • components can be filtered prior to crosslinking, applied to the tube and the tube subsequently packaged in a sealed bag and sterilized in an autoclave for in vivo application.
  • an exemplary chest tube is coated as described above with the composition.
  • the coated chest tube is then inserted ex vivo between pork spareribs using a technique simulating chest tube insertion in a patient and the durability and stability of the coating is observed.
  • Lidocaine release study For the lidocaine release study, two methods were used to assess the potential drug release difference between the hydrated and dehydrated coatings.
  • the first method uses a coated tubing which comprises a dried 1 cm by 0.5 cm piece of gelatin-lidocaine (100 mg lidocaine in 0.5 ml gelatin, 1 % GA). The coated tubing was placed in a 7 ml glass scintillation vial for the release study.
  • the second method uses a gel mold comprising a fresh, hydrated gelatin-lidocaine gel (200 mg lidocaine in 1 ml gelatin, 1 % GA) that was polymerized directly into a 7 ml glass scintillation vial thus forming a hydrogel disc. The hydrogel disc was allowed to stabilize for 30 minutes before examining the lidocaine release.
  • the release solution medium (1 ml of PBS) was added to the scintillation vials, and the sample was incubated at 37 s C. 1 ml samples were taken daily, and the release solution was replenished. The day 0 sample was taken 5 minutes after adding the release solution to the vials to measure the amount of lidocaine not incorporated into the gels.
  • the release profiles from the various gelatin formulations show the differences in lidocaine release rates are dependent on the concentration of gelatin and glutaraldehyde (GA).
  • the data display the sustained release of the lidocaine over 10 days.
  • SEM is performed for the lidocaine- polymer coated chest tube before and after autoclave sterilization.
  • lidocaine-polymer 5 mm x 10mm chest tube pieces are coated with equal volumes of lidocaine-polymer. SEM and lidocaine release testing are performed before and after autoclave sterilization.
  • Group 1 No incision or chest tube (2 subjects);
  • Group 2 Bilateral incisions, no chest tube (2 subjects);
  • Group 3 One control tube, one lidocaine-polymer coated tube (11 subjects);
  • Group 4 One Lidocaine-polymer tube only (2 subjects).
  • Data collected include daily weight, food consumption, mobility, analgesia requirements, other morbidity and mortality.
  • 64 Drug levels and inflammatory markers are measured daily out to 7 days.
  • Behavior and in vivo neurophysiology core technicians perform daily, blinded, non-biased assessment of pain.
  • the skin is probed with Von Frey hairs around the tube and response graded on the nociceptive scale. How far from the tube a change in nociception occurs is also assessed.
  • Tissue is prepared for histology. Core services are utilized to characterize scar formation, cytotoxicity and inflammation.
  • the coating does not flake off the chest tube after being applied.
  • the coating moves in coordination with the tube thus allowing the tube to remain flexible.
  • the coating remains stably fixed on the chest tube. The stable fixation of the coating allows the surgeon to maintain their techniques for handling and placing the chest tube in the patient.
  • the stable fixation of the coating to the tube prevents obstruction of the chest tube openings or central lumen. Similarly, when using a Blake drain, none of the drain channels are blocked.
  • Coated chest tube is able to stably release lidocaine
  • a post-sterilization lidocaine release profile indicates continued stable elution over a period of about 3-7 days.
  • the maximum subcutaneous dose of lidocaine hydrochloride is about 4.5 mg/kg with a maximum dose at one time not to exceed 300 mg.
  • the targeted lidocaine release for a chest tube or Blake drain is 4.5 mg/kg/day.
  • lidocaine as the anesthetic
  • other anesthetics are contemplated for use in the invention.
  • a chest tube is used in this embodiment, other types of catheters may be similarly coated.
  • the concept of a drug-eluting drain has significant scientific importance because of its potential use for any drain placed inside the body.
  • the technique of using a bioabsorbable polymer for drug delivery signifies the potential to deliver different types of drugs or combination of drugs, with elution profiles tailored to the application.
  • Drug delivery on a bio-absorbable scaffold is a rapidly evolving field.
  • Example 8 Lidocaine loaded PEG-PSHU-PEG micelles embedded in lidocaine-loaded crosslinked gelatin used to coat chest tube - Prophetic
  • lidocaine loaded micelle nanoparticles which can be embedded within a lidocaine-loaded bio-absorbable polymer hydrogel to form a drug delivery system that can be applied to a catheter, such as a chest tube, to allow for immediate as well as sustained release of the lidocaine.
  • Polymer concentration, lidocaine-loading amount and crosslinker concentration can be optimized to control the amount and rate of lidocaine release.
  • Lidocaine-loaded PEG-PSHU-PEG micelles were manufactured as described previously and detailed again below.
  • NBOC serinol (0.2873 g, 1 .5 mmol) and urea (0.09 g, 1 .5 mmol) were weighed out and lyophilized at -45 °C for 48 h.
  • the reactants were dissolved in 1 .5 ml_ of anhydrous DMF in a 25 ml_ round- bottom flask at 90°C under gentle stirring and a nitrogen atmosphere.
  • HDI 0.482 ml_, 3 mmol
  • PEG 1000 polyethylene glycol
  • 4 mmol polyethylene glycol
  • the resulting product polyethylene glycol)-blockpoly(serinol hexamethylene urea)-block-poly(ethylene glycol) (PEGPSHU-PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50°C and 10 mbar vacuum. Subsequently, the polymer was dissolved in Milli-Q water and dialyzed (MWCO: 3.5 kDa) against 1 L of Milli-Q water for 72 h at room temperature. Then, the product was lyophilized at -45 °C for 48 h to yield a white flaky material.
  • PEGPSHU-PEG polyethylene glycol)-blockpoly(serinol hexamethylene urea)-block-poly(ethylene glycol)
  • Lidocaine-loaded micelles were fabricated by a traditional emulsification-sonication procedure.
  • the PEG-PSHU-PEG polymer and lidocaine were dissolved in 1 mL of DMSO at 3 wt % (polymer/DMSO). This solution was then added dropwise to a beaker containing 20 mL of ultrapure water partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min.
  • the removal of DMSO was carried out by centrifugation at 11 ,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant, and then resuspending the micelles in ultrapure water. This DMSO extraction procedure was carried out three times. The resulting micelles were either used immediately or stored at -20 °C for later use.
  • aqueous solution containing gelatin is prepared in PBS (5% w/v or 10% w/v) and allowed to dissolve overnight in a 37 S C water bath.
  • Lidocaine-HCI is dissolved in the gelatin solution (200 mg lidocaine/1 ml gelatin solution).
  • Micelle nanoparticles loaded with lidocaine-HCI 50-200 mg lidocaine, 1 -20 mg polymer) are added to the gelatin solution (0.5 ml gelatin solution).
  • Glutaraldehyde (GA) is then added to the gelatin-micelle-lidocaine mixture (1 -4% w/v) to crosslink the gelatin into a hydrogel.
  • the crosslinked hydrogel containing the embedded lidocaine-loaded micelles are applied to the chest tube and dehydrated thus resulting in a thin, stable polymer coating on the tubing.
  • a hydrogel containing is prepared as described above. Before the hydrogel is fully formed, the viscous solution is poured onto the catheter tubing and allowed to fully polymerize for 5 minutes. The tubing is slowly spun around by hand during the polymerization, to ensure a smooth, even coating of the gel. From here, the hydrogel coating on the tubing can be fully dehydrated overnight at room temperature or used immediately for further applications. The fully dehydrated samples result in a thin, rigid polymer coating on the tubing, while the fresh hydrogel samples produce a thick, elastic, jelly-like coating. Both forms of the gelatin coating are analyzed with an ex vivo rib test and a lidocaine release study is performed.
  • a mold can be used, instead of hand-spinning, to apply the hydrogel to the tubing.
  • an aqueous solution containing gelatin is prepared in PBS (5% w/v) and allowed to dissolve overnight in a 37 S C water bath.
  • Lidocaine-HCI was dissolved in the gelatin solution (200 mg lidocaine/1 ml gelatin solution).
  • Micelle nanoparticles loaded with lidocaine-HCI 50-200 mg lidocaine, 1 -20 mg polymer) are added to the gelatin solution (0.5 ml gelatin solution).
  • Glutaraldehyde is then added to the gelatin-lidocaine mixture (1% w/v) to crosslink the gelatin into a hydrogel.
  • the tubing Before the hydrogel was fully formed, the tubing is inserted into the mold containing the gelatin-micelle-glutaraldehyde solution and the polymerization was allowed to continue until fully gelled (approximately 1 hour). Subsequently, the hydrogel coated tubing was removed from the mold. After polymerization the coated chest tubes are dehydrated to form a stably fixed coating.
  • the introduction of a lyophilizer may control the dehydration step and further aid in eliminating variation in elution that may be related to this step.
  • a nasogastric tube is inserted into the chest tube lumen during the coating and polymerization process.
  • wooden skewers are inserted into the lumen allowing the entire length of the Blake drain to be coated.
  • the coating may be applied in the operating room similar to the application of BioGlue®. Applying the coating in the operating room has the advantage of being able to adjust the dose if there is a concern for toxicity based on weight or hepatic impairment.
  • components can be filtered prior to crosslinking, applied to the tube and the tube subsequently packaged in a sealed bag and sterilized in an autoclave for in vivo application.
  • an exemplary chest tube is coated as described above with the composition.
  • the coated chest tube is then inserted ex vivo between pork spareribs using a technique simulating chest tube insertion in a patient and the durability and stability of the coating is observed.
  • Lidocaine release is determined using a UV-Vis spectrophotometer (GENESYS 10S UV-Vis) at 280 nm wavelength.
  • the coated chest tube is in a sealed vial with phosphate buffered saline (PBS) and aliquots tested every 24 hours. Testing is performed both before and after autoclave sterilization.
  • PBS phosphate buffered saline
  • SEM is performed for the lidocaine- polymer coated chest tube before and after autoclave sterilization.
  • one lidocaine coated piece of tube is inserted subcutaneously over the posterior rib cage on one side and an uncoated piece of tube in the exact location on the opposite side.
  • Power analysis using the resource equation estimates 22 subjects to detect a difference of 20% (nociceptive pain scale 0 - 10).
  • 11 subjects are needed. Additional controls include lidocaine-specific morbidity (see group 4), Von Frey testing at baseline (see group 1 ) and the incision alone (see group 2).
  • Group 1 No incision or chest tube (2 subjects);
  • Group 2 Bilateral incisions, no chest tube (2 subjects);
  • Group 3 One control tube, one lidocaine-polymer coated tube (11 subjects);
  • Group 4 One Lidocaine-polymer tube only (2 subjects).
  • Data collected include daily weight, food consumption, mobility, analgesia requirements, other morbidity and mortality.
  • 64 Drug levels and inflammatory markers are measured daily out to 7 days.
  • Behavior and in vivo neurophysiology core technicians perform daily, blinded, non-biased assessment of pain.
  • the skin is probed with Von Frey hairs around the tube and response graded on the nociceptive scale. How far from the tube a change in nociception occurs is also assessed.
  • Tissue is prepared for histology. Core services are utilized to characterize scar formation, cytotoxicity and inflammation.
  • Novel micelle NPs were created from a PEG-PSHU-PEG polymer.
  • the PSHU backbone provides the hydrophobic core of the micelles, while the terminal PEG chains provide the hydrophilic interactions on the exterior shell.
  • PEG adds an additional benefit as this provides a “stealth” effect to the NPs for an extended half-life in circulation.
  • Scanning electron microscopy (SEM) was used to characterize the morphologies of the micelles.
  • the micelle NPs were uniformly spherical ( Figure 9).
  • Lidocaine loaded micelles embedded in crosslinked gelatin-lidocaine hydrogel produce a durable and stable coating to chest tube
  • the coating does not flake off the chest tube after being applied.
  • the coating moves in coordination with the tube thus allowing the tube to remain flexible.
  • the coating remains stably fixed on the chest tube.
  • the stable fixation of the coating allows the surgeon to maintain their techniques for handling and placing the chest tube in the patient.
  • the stable fixation of the coating to the tube prevents obstruction of the chest tube openings or central lumen. Similarly, when using a Blake drain, none of the drain channels are blocked.
  • Coated chest tube is able to stably release lidocaine
  • a post-sterilization lidocaine release profile indicates continued stable elution over a period of about 3-7 days.
  • the maximum subcutaneous dose of lidocaine hydrochloride is about 4.5 mg/kg with a maximum dose at one time not to exceed 300 mg.
  • the targeted lidocaine release for a chest tube or Blake drain is 4.5 mg/kg/day.
  • lidocaine as the anesthetic
  • other anesthetics are contemplated for use in the invention.
  • a chest tube is used in this embodiment, other types of catheters may be similarly coated.
  • lidocaine-loaded micelles embedded within the hydrogel can provide a slow, sustained release over time.
  • the general procedure for loading the lidocaine into the hydrogel and micelles can remain the same as above. However, less than the initial amount of lidocaine may be loaded into the hydrogel as more lidocaine is added with the addition of the lidocaine-loaded micelles.
  • the inventors have developed a novel nanoparticle capable of being loaded with a therapeutic agent such as an anesthetic, antimicrobial or growth factor. These nanoparticles can be embedded within a crosslinked hydrogel to form a drug delivery system.
  • the drug delivery system can be administered directly to the patient or alternatively, are capable of being applied to various types of catheters or drains, such as chest tubes.
  • Application of the composition to a catheter or drain allows for the sustained release of the therapeutic agent, such as an anesthetic, to the patient over an extended period of time.

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Abstract

Un nouveau système d'administration d'agent thérapeutique, ainsi que des nouveaux procédés d'utilisation et des nouveaux procédés de formation de celui-ci sont présentés. Le nouveau système d'administration est constitué de nouvelles nanoparticules capables d'encapsuler au moins partiellement un agent thérapeutique tel qu'un anesthésique, un antimicrobien, un facteur de croissance ou une protéine. Les nanoparticules sont incorporées dans un hydrogel réticulé. L'hydrogel peut être administré directement à un patient ou peut être revêtu sur un dispositif tel qu'un cathéter. Le système d'administration permet une libération prolongée de l'agent thérapeutique sur une période de temps prolongée.
PCT/US2021/037010 2020-06-11 2021-06-11 Systèmes d'administration d'agents thérapeutiques et leurs procédés de formation et leurs utilisations WO2021252897A2 (fr)

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US8273366B2 (en) * 2002-06-05 2012-09-25 University Of Florida Research Foundation, Incorporated Ophthalmic drug delivery system
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